Nanoparticules multifonctionelles pour la résonance magnétique et l ...

tel-00661206, version 1 - 18 Jan 2012 

N° d’ordre :4399 

THÈSE 

EN CO-TUTELLE PRÉSENTÉE A 

L’UNIVERSITÉ BORDEAUX 1 ET 

L’UNIVERSITÉ d’ AVEIRO 

ÉCOLE DOCTORALE DES SCIENCES CHIMIQUES 

Par Sonia Luzia Claro de PINHO 

POUR OBTENIR LE GRADE DE 

DOCTEUR 

SPÉCIALITÉ : Physico-chimie de la matière condensée. 

MULTIFUNCTIONAL NANOPARTICLES FOR MR AND 

FLUORESCENCE IMAGING 

Soutenue le 14 décembre 2011 

Après avis de : 

S. Begin-Colin, University of Strasbourg Rapporteur 

H. D. Burrows, University of Coimbra Rapporteur 

Devant la commission d’examen formée de : 

MME BEGIN-COLIN Sylvie, Professeur des Universités Rapporteur 

M. BURROWS HUGH D., Professeur en établissement étranger Rapporteur 

M. CARLOS Luis, Professeur en établissement étranger Rapporteur 

MME DELVILLE Marie-Hélène, Directeur de recherche CNRS Directeur de thèse 

M. ETOURNEAU Jean, Professeur des Universités Rapporteur 

M. GERALDES Carlos F.G.C., Professeur en établissement étranger Rapporteur 

M. ROCHA Joao, Professeur en établissement étranger Directeur de thèse 

M. VIEIRA Joaquim Manuel, Professeur en établissement étranger Président 

Université Bordeaux 1 

Les Sciences et les Technologies au service de l’Homme et de l’environnement

tel-00661206, version 1 - 18 Jan 2012



Sonia Luzia Claro de 

Pinho 

Universidade de Aveiro 

2011 

Departamento de Química 

Nanopartículas Multifuncionais para Imagens de RM 

e Fluorescência 

Multifunctional Nanoparticles for MR and 

Fluorescence Imaging

tel-00661206, version 1 - 18 Jan 2012



Sonia Luzia Claro de 

Pinho 

Universidade de Aveiro 

2011 

Departamento de Química 

Nanopartículas Multifuncionais para Imagens de RM 

Fluorescência 

Multifunctional Nanoparticles for MR and 

Fluorescence Imaging 

Dissertação apresentada à Universidade de Aveiro para cumprimento dos 

requisitos necessários à obtenção do grau duplo de Doutor em Química pela 

Universidade de Aveiro e de Physico-chimie de la matière condensée pela 

Universidade de Bordéus 1, realizada sob o regime de co-tutela com a orientação 

científica do Doutor João Carlos Matias Celestino Gomes da Rocha, Professor 

Catedrático do Departamento de Química da Universidade Aveiro, do Doutor Luís 

António Ferreira Martins Dias Carlos Professor Catedrático do Departamento de 

Física da Universidade Aveiro e da Doutora Marie-Hélène Delville, Directeur de 

Recherche do Institut de Chimie de la Matiere Condenseé de Bordeaux. 

Apoio financeiro da FCT e do FSE no 

âmbito do III Quadro Comunitário de 

Apoio.


v


À memória da minha avó 

Aos meus pais e irmãos, 

Ao Pedro

tel-00661206, version 1 - 18 Jan 2012


o júri / the jury 

presidente / president Prof. Doutor Joaquim Manuel Vieira 

Professor Catedrático do Departamento de Engenharia Cerâmica e do Vidro da Universidade de 

Aveiro 

vogais / members Prof. Doutor Syvie Begin-Colin 

Professor catedrático do Institut de Physique et de chimie des Matériaux de Strasbourg da 

Universidade de Strasbourg 

Prof. Doutor Hugh Douglas Burrows 

Professor Catedrático da Faculdade de Ciências e Tecnologia da Universidade de Coimbra 

Prof. Doutor Jean Etourneau 

Professor catedrático do Institut de Chimie de la Matiere Condensee de Bordeaux da Universidade 

de Bordeaux 1 

Prof. Doutor Carlos Frederico de Gusmão Campos de Geraldes 

Professor Catedrático da Faculdade de Ciências e Tecnologia da Universidade de Coimbra 

Prof. Doutor João Carlos Matias Celestino Gomes da Rocha 

Professor Catedrático do Departamento de Química da Universidade de Aveiro 

Prof. Doutor Luís António Ferreira Martins Dias Carlos 

Professor Catedrático do Departamento de Física da Universidade de Aveiro 

Doutora Marie-Hélène Delville 

Directeur de Recherche CNRS do Institut de Chimie de la Matiere Condensee de Bordeaux da 

Universidade de Bordeaux 1 .

tel-00661206, version 1 - 18 Jan 2012


agradecimentos / 

acknowledgements 

After almost four years of working on this project it was time to wrap it up in a 

thesis. Through good and less good times, this story came to a ‘happy ending’ 

and I should now thank people for their help, support, friendship or for just 

being around making my life colorful. 

I have to say that this thesis was crafted by several remarkable people. First of 

all, I would like to thank my supervisors/co-supervisors Dr. Marie-Hélène 

Delville, Dr. João Rocha, Dr Luis Carlos and finally Dr. Carlos Geraldes, whose 

importance for this thesis as well as for my personal development is 

inexpressible. They have all shared with me their passion for Science, the most 

important jewel that I will take home. Marie-Hélène, you’ve been an amazing 

teacher and an example for me. You have let me into your world and your 

philosophy of life and sense of humour. Dr. João Rocha and Dr. Luís Carlos 

you have given me this PhD opportunity, you have always respected my 

choices in the research line I have traced and most importantly you’ve 

supported me in these choices. Dr. Carlos Geraldes, you have introduced me 

into the addictive and sophisticated world of NMR. Thank you all for everything! 

Next in my list are all my labmates from ICMCB in Bordeaux and at the 

University of Aveiro, a huge “thank you” for the patience, help, brainstorming 

and enjoyable moments spent together.I will miss those afternoon coffee 

breaks. A special thanks to my good friend Laetitia Etienne, all your dropping 

by ‘just to say hi’ and our great discussions about life. You were always there 

when I needed you, it always felt good to share either my successes or 

troubles, being sure of your sincere support in everything. Of course I cannot 

forget my good friend Dr Patricia Lima for all her support, availability and 

friendship. 

Next in my list are all my labmates from ICMCB in Bordeaux and at the 

University of Aveiro, a huge “thank you” for the patience, help, brainstorming 

and enjoyable moments spent together.I will miss those afternoon coffee 

breaks. A special thanks to my good friend Laetitia Etienne, all your dropping 

by ‘just to say hi’ and our great discussions about life. You were always there 

when I needed you, it always felt good to share either my successes or 

troubles, being sure of your sincere support in everything. Of course I cannot 

forget my good friend Dr Patricia Lima for all her support, availability and 

friendship. 

I cannot forget to thank Dr. Stéphane Mornet for his help in chemistry at the 

early stage of this work and his availability in the lab, as well as, Dr. Pierre 

Voisin and Dr. Emeline Ribot from RMSB for openning up their lab to me. Many 

thanks should also be given to François, Patrick, Laetitia and Virginia from the 

ICMCB, for all their support, friendship and wonderful moments. 

I am also very grateful to Dr. Giovannia Pereira for all her help at the NMR lab. 

She has tought me well. 

I cannot forget to thank Dr. Henrique Faneca, Dr. Sophie Laurent and Inês 

Violante for the precious scientific inputs. Without you I would not be able to 

have obtain such good biological results 

My thanks go also to the friends that were always there, in a paralell world, to 

keep my sanity by giving me pleasant non-scientific moments. They proven 

over and over again that true friends are able to grow old without growing apart. 

Finally, but not less important, I would like to thank my family, specially my 

parents, and my husband Pedro for all their unconditional support, 

understanding and affection.

tel-00661206, version 1 - 18 Jan 2012


palavras-chave 

resumo 

Nanopartículas multifuncionais, core-shell/corona, sílica, Fe2O3, lantanídeos, 

agentes de contraste para IRM, agentes de contraste ópticos, relaxometria, 

RMND, fotoluminescência 

Nos últimos anos, surgiu uma nova geração de nanopartículas (NPs) 

multifuncionais destinada a aplicações biomédicas, mais concretamente para 

uso em técnicas de diagnóstico, reconstrução celular e em diversas aplicações 

terapêuticas. Em relação às suas antecessoras, estas novas nanopartículas 

apresentam uma estrutura mais elaborada, integrando vários componentes 

ativos com diferentes funcionalidades que, em princípio, permitem realizar 

diversas tarefas em simultâneo (como o direcionamento ativo para 

determinadas células ou compartimentos celulares, imagem e libertação de 

fármacos). Estas nanopartículas são designadas, por isso, de nanopartículas 

multifuncionais. 

A presente dissertação relata o desenvolvimento de dois tipos de sondas 

bimodais e as propriedades físico-químicas destas, nomeadamente a sua 

textura, estrutura e relaxometria e dinâmica de 1 H , com o objectivo de avaliar 

o seu potencial como agentes de contraste para Imagem por Ressonância 

Magnética Nuclear (IRM). São, também, apresentados estudos de 

fotoluminescência que permitem avaliar o potencial daquelas sondas para 

serem usadas como agentes de contraste óptico. Estes materiais combinam as 

propriedades dos complexos de lantanídeos trivalentes (Ln 3+ ) e das NPs 

funcionando, assim, como agentes bimodais. 

Foram desenvolvidos os seguintes sistemas fotoluminescentes e com 

contraste T1 em IRM em que os iões Ln 3+ magnetica (Gd 3+ ) e opticamente 

(Eu 3+ , Tb 3+ ) activos se encontram à superfície das NPs de sílica: 

SiO2@APS/DTPA:Gd:Ln e SiO2@APS/Pyd-DTPA:Gd:Ln (Ln = Eu ou Tb). No 

que respeita às propriedades de relaxometria, na presença destas NPs 

observa-se um aumento moderado de r1 e considerável de r2, especialmente a 

campos magnéticos altos (devido aos efeitos de susceptibilidade para r2). Os 

iões Eu 3+ apresentam um único ambiente local de baixa simetria, sendo que a 

emissão de fotoluminescência não é influenciada pela presença simultânea de 

Gd 3+ e Eu 3+ . Verificou-se que a presença de Tb 3+ (em lugar do ião Eu 3+ ) 

aumenta ainda mais o valor r1, diminuindo r2. A internalização das NPs em 

células vivas é rápida e resulta num aumento de intensidade nas imagens 

ponderadas em T1. Foram estudadas as características ópticas de pastilhas de 

células (“cellular pellets”) contendo NPs, tendo-se confirmado o interesse das 

novas sondas propostas enquanto agentes para imagem bimodal. 

Esta dissertação relata, ainda, agentes com contraste em T2 para IRM, que 

consistem em um sistema núcleo-coroa (“core-shell”) em que é ajustável a 

espessura da coroa de sílica envolvendo o núcleo de óxido de ferro. A 

espessura do revestimento de sílica tem um efeito significativo sobre a 

relaxividade r2 e r2* , sendo aqui proposto um modelo para explicar este 

comportamento. A viabilidade celular e a expressão da desidrogenase 

mitocondrial das células da microglia foram também avaliadas.

tel-00661206, version 1 - 18 Jan 2012


keywords 

abstract 

Multifunctional nanoparticles, core-shell/corona, silica, Fe2O3, 

lanthanides, MRI contrast agents, optical contrast agents, relaxometric, 

RMRD, photoluminescence 

In the past few years a new generation of multifunctional nanoparticles 

(NPs) has been proposed for biomedical applications, whose structure is 

more complex than the structure of their predecessor monofunctional 

counterparts. The development of these novel NPs aims at enabling or 

improving the performance in imaging, diagnosis and therapeutic 

applications. The structure of such NPs comprises several components 

exhibiting various functionalities that enable the nanoparticles to perform 

multiple tasks simultaneously, such as active targeting of certain cells or 

compartmentalization, imaging and delivery of active drugs. 

This thesis presents two types of bimodal bio-imaging probes and 

describes their physical and chemical properties, namely their texture, 

structure, and 1 H dynamics and relaxometry, in order to evaluate their 

potential as MRI contrast agents. The photoluminescence properties of 

these probes are studied, aiming at assessing their interest as optical 

contrast agents. These materials combine the properties of the trivalent 

lanthanide (Ln 3+ ) complexes and nanoparticles, offering an excellent 

solution for bimodal imaging. 

The designed T1- type contrast agent are SiO2@APS/DTPA:Gd:Ln or 

SiO2@APS/PMN:Gd:Ln (Ln= Eu or Tb) systems, bearing the active 

magnetic center (Gd 3+ ) and the optically-active ions (Eu 3+ and Tb 3+ ) on 

the surface of silica NPs. Concerning the relaxometry properties, 

moderate r1 increases and significant r2 increases are observed in the 

NPs presence, especially at high magnetic fields, due to susceptibility 

effects on r2. The Eu 3+ ions reside in a single low-symmetry site, and the 

photoluminescence emission is not influenced by the simultaneous 

presence of Gd 3+ and Eu 3+ . The presence of Tb 3+ , rather than Eu 3+ ion, 

further increases r1 but decreases r2. The uptake of these NPs by living 

cells is fast and results in an intensity increase in the T1-weighted MRI 

images. The optical features of the NPs in cellular pellets are also studied 

and confirm the potential of these new nanoprobes as bimodal imaging 

agents. 

This thesis further reports on a T2 contrast agent consisting of core-shell 

NPs with a silica shell surrounding an iron oxide core. The thickness of 

this silica shell has a significant impact on the r2 and r2* relaxivities, and a 

tentative model is proposed to explain this finding. The cell viability and 

the mitochondrial dehydrogenase expression given by the microglial cells 

are also evaluated.

tel-00661206, version 1 - 18 Jan 2012


Mots clés 

resumée 

Nanoparticules multifonctionnelles, coeur-écorce/couronne, silice, Fe2O3, 

lanthanides, agents de contraste IRM, agents de contraste optiques, 

relaxométrie, RMRD, photoluminescence 

Cette thèse décrit une stratégie de synthèse de nouvelles générations 

des nanoparticules (NPs) pour applications biomédicales, visant à une 

amélioration de leurs performances pour l’imagerie, le diagnostic 

thérapeutique. Ces NPs présentent plusieurs fonctionnalités leur 

permettant de réaliser des tâches multiples. Deux types de sondes 

bimodales ont été développés et étudiés afin d'évaluer leur potentiel 

comme agents (1) de contraste en IRM et (2) luminescents. Ces objets 

combinent les propriétés des complexes de lanthanide (Ln 3+ ) et celles 

des NPs de silice ou de type coeur-écorce Fe2O3@SiO2 pour une 

imagerie bimodale. Ces NPs testées sur des cellules vivantes ont permis 

d’illustrer la preuve du concept aussi bien en IRM avec une augmentation 

d'intensité des images et un impact significatif sur les relaxivities r1, r2 et 

r2* qu’en photoluminescence. L’étude du système coeur-écorce a montré 

que l’influence du contrôle fin de l’écorce autour du noyau d'oxyde de fer 

a pu être modélisée.


Contents 

1. Introduction 

1.1 Context of the thesis 3 

1.2. Scope and Objectives of the thesis 8 

1.3. Outline of the thesis 9 

1.4. References 11 

2. Magnetic Resonance Imaging Background Concepts 

2.1 Magnetic Resonance Imaging 17 

2.2. MRI Contrast Agents 22 

2.2.1. Relaxation 24 

2.2.1.1. Spin-Lattice Relaxation 24 

2.2.1.2. Spin-Spin Relaxation 26 

2.2.2. Relaxivity 28 

2.2.2.1. Inner Sphere Relaxivity 30 

2.2.2.2. Second and Outer Relaxivity 35 

2.2.2.3. Paramagnetic Relaxation Parameters 37 

2.2.3. Nuclear Magnetic Resonance Dispersion 39 

2.3. Classification of CAs 44 

2.3.1. Chemical Composition, Magnetic Properties and Effect on the 

MRI Image 

2.3.2. Biodistribution and Applications 48 

2.3.2.1. Non-specific Agents 48 

2.3.2.2. Specific or Targeted Agents 52 

2.3.2.3. Nono-injectable Organ Specific Agents 56 

2.3.2.4. Responsive, Smart or Bioactivated Agents 58 

2.3.2.5. Contrast Agents Based on Other Properties 63 


i 

45


3. Fluorescence Imaging Background Concepts 

3.1 Optical Imaging 80 

3.2. Luminescence 82 

3.3. Tissue Optical Properties 91 

3.4. Optical Imaging Technology 93 

3.5. Optical Contrast Agents 95 


4. Lanthanide-Chelate Grafted Silica Nanoparticles as Bimodal-Imaging 

Contrast Agents 

4.1 Introduction 113 

4.2. Experimental Procedures 116 

4.3. Results and Discussion 121 

4.3.1. Characterization of Nanoparticles 121 

4.3.2. Photoluminescence Properties 131 

4.3.3. Relaxivity Properties 151 

4.3.4. Cell Imaging 159 

4.4. Conclusions 161 


5. Core-Shell Nanoparticles for Bimodal-Imaging Contrast Agents 

5.1 Introduction 172 





5.3.3. Cytotoxicity 199 



6. Final Conclusions and Future Work 

Final Conclusions and Future Work 211 

ii


Nomenclature, symbols and 

acronyms 

B0: The static, homogeneous magnetic field used to polarize spins, 

creating magnetization. This can refer to both the direction and the 

magnitude of the field. The direction of B0 defines the longitudinal axis. 

B1: Magnetic component of a radio-frequency field applied perpendicular 

to the longitudinal axis (B0) to perturb the magnetization in some manner 

(e.g., excitation pulses, inversion pulses, etc). 

BOLD, or blood oxygenation level dependent: The BOLD effect is the 

source of contrast in FMRI. The presence of deoxygenated blood leads to 

signal loss due to a reduction in T2*. 

BPA: Blood pool agents. 

Contrast Agent (CA): A substance that enhances contrast, i.e., shortens 

the relaxation time of water molecules, making them appear ‘brighter’ in 

T1 or T2 weighted MR images. 

Dephasing: In an ensemble of spins, each has a phase angle in the 

transverse plane. The ensemble can either be coherent (have the same 

angle) or incoherent (varying angle). Loss of coherence is referred to as 

"dephasing", which leads to net signal loss. BOLD and diffusion contrast 

are both based on dephasing. 

DNP: Dynamic nuclear polarization. 

Echo time (TE): The time between the excitation of magnetization and the 

acquisition of signal. For long acquisition windows, the TE is usually 

defined as the point at which the acquisition is closest to the center of k- 

space. 

ED: Electric-dipole. 

Emission quantum efficiency (η): the fraction of emission processes in 

which emission of light is involved: 

 

 

 

EXP 

RAD 

. 

RAD 

 

A 

RAD 

A 

A 

NRAD 

Emission quantum yield (ϕ): the ratio of the number of photons emitted 

number of 

emitted 

photons 

number of 

absorved photons 

and the absorbed: 

.



acronyms (cont.) 

Extracellular fluid agents (ECF): are the contrast agents that have been in 

clinical use for the longest period of time in contrast-enhanced MR 

imaging of the liver. 

Frequency encode (FE) direction: The direction along which individual 

lines are acquired in k-space. Note that one may also refer to the 

frequency encode direction in the image (i.e., if the FE direction is along 

kx, one may also refer to the x direction in image space as the FE 

direction). See also phase encode (PE) direction. 

Gradient: Spatially varying magnetic field used to manipulate the 

resonance frequency across an object. MRI scanners incorporate three 

linearly-varying gradient fields across x, y and z. The strength of the 

gradient is controlled by the pulse sequence and can be rapidly 

manipulated (or "switched"). 

Gradient echo, or GRE/GE: Any MRI sequence that detects an un- 

refocused signal (i.e., a non-spin-echo sequence). 

Inner Coordination Sphere. Used to describe any water molecules that 

are directly coordinated to the metal ion of a contrast agent. 

Inversion: Rotation of magnetization from alignment with B0 (along +z) to 

anti-alignment (along -z), caused by a RF pulse with 180º flip angle. 

Inversion time (IT): The time between inversion of magnetization and its 

excitation (often to acquire signal that has been "prepared" with an 

inversion pulse). 

k-space: The conceptual space in which MRI images are acquired. Data 

in k-space provides a map of the amount of structure in the image that 

can be attributed to each spatial frequency. The image and k-space are 

related by the mathematical operation called the Fourier transform (FT). 

ii




Larmor frequency: See resonance frequency. 

LLBs: Lanthanide luminescent bioprobes. 

LMCT: ligand to-metal charge-transfer. 

Longitudinal axis (z): The direction parallel to the main magnetic field (B0), 

which represents the direction along which magnetization is in 

equilibrium. After an RF pulse, the magnetization recovers to this 

equilibrium according to the rate T1. The component of magnetization 

along z cannot be detected. 

MD: Magnetic dipole. 

Magnetization: The net nuclear magnetic moment induced in an object or 

tissue when exposed to an external magnetic field. The magnetization is 

at equilibrium when it is aligned parallel to the external field (along the 

longitudinal axis) and at its maximum magnitude. MRI experiments 

manipulate the magnetization away from this equilibrium, into the 

transverse plane to detect its signal. 

MRI: Magnetic Resonance Imaging, a non-invasive technique using 

strong magnetic fields and radiofrequency pulses to create images of 

internal anatomy. 

NIR: Near-infrared. 

NPs: Nanoparticles. 

NMR: Nuclear Magnetic Resonance, an analytical technique commonly 

employed in chemistry. 

NMRD. Nuclear Magnetic Resonance Dispersion profile. 

OCT: Optical coherence tomography. 

Outer Coordination Sphere: Used to describe any water molecules that 

are not in the inner- or second-coordination spheres. 

iii




PARACEST: Paramagnetic chemical exchange saturation transfer. 

PHIP: Parahydrogen-induced polarization. 

PET: Positron-emission tomography. 

Phase: In MRI, phase most commonly refers to the angle of the 

magnetization in the transverse plane. 

Phase encode (PE) direction: The direction perpendicular to individual 

lines in k-space. Note that one may also refer to the phase encode 

direction in the image (i.e., if the PE direction is along ky, one may also 

refer to the y direction in image space as the PE direction). Image 

artefacts are often most severe along the PE direction. See also 

frequency encode (FE) direction. 

Polarization: The tendency for spins to align in the presence of an 

external magnetic field, creating a net magnetic moment (or, 

magnetization). 

Precession: Gyration of a spinning body, which traces out a cone about 

the axis of precession. In MRI, the magnetization that is out of alignment 

with the longitudinal axis precesses about this axis. 

Relaxation Time: The time it takes nuclear spins to return to their 

equilibrium state after excitation by an RF pulse. 

Relaxivity: The ability of a contrast agent to shorten the relaxation time of 

nearby water protons. The higher the relaxivity, the shorter the relaxation 

time. 

Pulse sequence: The series of RF pulses, gradient field amplitudes and 

acquisition periods applied to acquire a spectrum or an image. Also used 

to refer to sets of sequences with common properties (e.g., the family of 

spin echo pulse sequences include structural and diffusion variants). 

QDs: Quantum dots. 

iv




R1: The expression of time constant T1 as a rate: R1=1/T1. 

R2: The expression of time constant T2 as a rate: R2=1/T2. 

Relaxation: The process by which the magnetization slowly returns to 

equilibrium following the rotation of the magnetization away from the 

longitudinal axis with an RF pulse. Relaxation (decay) in the transverse 

plane has the characteristic time T2; relaxation (recovery) along the 

longitudinal axis has characteristic time T1. 

Refocusing: The rotation of magnetization in the transverse plane by 

180º, usually with the intention of removing the effect of field 

inhomogeneities on signal. 

Repetition time (TR): The time between repeated excitations of a given 

component of the magnetization. For multi-slice sequences that 

sequentially excite different imaging slices, the TR is the time between 

repeated excitations of the same slice. For volume-excite (3D) sequences 

that repeated excite the entire imaging volume, the TR is the time 

between repeated excitations of the volume. 

Resonance frequency (ω0): The frequency at which the magnetization 

can be excited and detected. The frequency varies directly with magnetic 

field strength, and is normally in the radio frequency (RF) range. Also 

called Larmor frequency. 

RF pulse: A brief transmission of energy in the RF range. In MRI, RF 

pulses are used to manipulate the direction of the magnetization (e.g., to 

excite or invert). 

v




Saturation: Removal or reduction of a component of the magnetization 

using excitation into the transverse plane. Most commonly used to in the 

context of "T1 saturation", where repeated excitations of magnetization 

result in incomplete T1 recovery, and therefore signal reduction. Can also 

refer to an excitation followed by a "spoiling" mechanism, which 

deliberately removes signal. 

Second Coordination Sphere. Used to describe any water molecules that 

are coordinated to only the contrast agent ligand, not to the metal ion 

inside a contrast agent. For example, a water molecule that is hydrogen 

bound to a polar group on the ligand. 

Shimming: The process of improving field homogeneity by compensating 

for imbalances in the main magnetic field of an MRI system. 

Accomplished by a combination of constant ("passive") shims and 

controllable ("active") shim coils. 

Spin: The property exhibited by atomic nuclei that contain an odd number 

of protons and/or neutrons. The property spin causes nuclei to behave as 

though they are spinning charges. These nuclei are often referred to 

colloquially as "spins". 

Spin echo, or SE: Any MRI pulse sequence characterized by the use of a 

refocusing pulse to reverse the effect of off-resonance precession. The 

signal is said to form a spin echo when the off-resonance precession of all 

spins is reversed, such that the spins re-align to form a signal peak (or 

"echo"). 

T1: The time constant defining the rate of recovery of magnetization along 

the longitudinal (z) axis following an RF pulse. Also called the spin-lattice 

relaxation time. 

T2: The time constant defining the irreversible loss of magnetization (and 

therefore signal) in the transverse (xy) plane following excitation. Also 

called the spin-spin relaxation time. In tissue, T2 is shorter than T1 (often 

by an order of magnitude). This signal loss cannot be recovered by a spin 

echo. 

vi




T2*: The time constant defining the loss of signal following excitation. Two 

components contribute to T2*. First, some signal loss occurs due to T2 

relaxation (i.e., loss of magnetization). Second, some signal loss is 

caused by variation in precession angles for different spins within a voxel. 

This does not represent loss of magnetization, and this component can be 

recovered wiwth a spin echo. 

T1M and T2M: The longitudinal and transverse proton relaxation 

enhancement experienced by the inner-sphere water molecules, 

respectively. 

TRL: Time-resolved luminescent. 

Transverse plane: The plane orthogonal to the longitudinal axis, denoted 

as the xy plane. It is defined by the direction of the main (B0) field, and the 

plane in which the magnetization is observable (i.e., where signal can be 

detected). 

Voxel: A resolution element in a 3D imaging experiment (coming from the 

term "volume element"); the 3D extension of a pixel (a "picture element"). 

C: the molar concentration of the paramagnetic complex. 

q: the number of water molecules in the inner-coordination sphere directly 

coordinated to the paramagnetic centre. 

ΔωM: the chemical shift difference between the free and the bound water 

molecules. 

γ: the proton gyromagnetic ratio. 

τM: the mean residence lifetime of the coordinated water protons within 

the inner-sphere coordination sphere. 

τR: The rotational correlation time of a contrast agent. This is related to 

how fast the contrast is physically tumbling in solution. 

τRAD: Radiative lifetime. 

vii


viii


1 

1. 

Introduction


Introduction 

1.1. Context of the thesis 3 

1.2. Scope and objectives of the thesis 8 

1.3. Outline of the thesis 9 


2


1.1. CONTEXT OF THE THESIS 

Nanomedicine is defined as the medical application technology in the nanometer 

range (ranging from biomedical imaging to drug delivery and therapeutics. 1 This is a new, 

fast expanding and growing medical field. The overall goals of traditional medicine and 

nanomedicine are the same: early and accurate diagnosis, effective treatments, free from 

side effects and non–invasive evaluation of the efficacy of the treatment. Nanotechnology 

used in nanomedicine brings not only improvements to the existing techniques but also 

provides completely new tools and capabilities. 2,3 

Nanotechnology is the study, understanding and control of matter at dimensions in 

the range of 1 to ~100 nanometres, where unique phenomena enable novel applications. 4 

It is a very broad interdisciplinary research field that involves various areas of science, 

such as chemistry, physics, engineering, biology and medicine, while eroding the 

traditional boundaries between them. 5 

Indeed, there is a radical change in the physical and chemical properties of 

materials as their size is scaled down to small clusters of atoms. When the size decreases 

the surface-area-to-volume ratio increases, since there is a larger percentage of surface 

atoms compared with bulk materials. 6 As nanodevices in the same range of dimension as 

antibodies, membrane receptors, nucleic acids and proteins, among other biomolecules 

the can possess biomimetic features. 7 

Given all these outstanding features nanoparticles are powerful tools for imaging, 

diagnosis and therapy. Over the last few years, a new generation of nanoparticles has 

arisen with improved performances for the same type of applications and with a more 

complex structure compared to the simpler monofunctional nanoparticles. These complex 

nanoparticles comprise different components that can carry out various functions, which 

enable these nanoparticles to perform multiple tasks simultaneously (active targeting a 

given type of cells or compartment, imaging them and delivering an active compound). For 

example, a core particle may be linked to a specific targeting functional group that 

recognises the unique surface signatures of the target cells. The same particle can be 

modified with an imaging functional group to monitor the drug transport process, a 

molecular entity to evaluate the therapeutic efficacy of a drug, a specific cellular 

3


Introduction 

penetration moiety and a therapeutic agent. 7 Therefore, these complex nanoparticles are 

termed multifunctional nanoparticles. 8 A schematic representation of these systems is 

given in Figure 1.1. 

Figure 1.1. Schematic representation of a multifunctional nanoparticle 7 

A large variety of therapeutic entities can be incorporated or attached on the surface 

of the multifunctional nanoparticles while other components can be used for targeting 

and/or imaging. Nanoparticles are prepared with organic polymers (organic nanoparticles) 

and/or inorganic elements (inorganic nanoparticles). 7 Figure 1.2 depicts examples of 

nanomaterials and nanocarrier systems. There are three major categories in which the 

multifunctional nanoparticles may be classified: liposomes and micelles, polymeric 

carriers, and core-shell/corona structures. 

4


Stability, 

Table 1.1. Strategies for constructing multifunctional nanoparticles. 6 

Properties Benefits Functional group Refs 

biocompatibility 

Specific targeting 

Intracellular 

penetration 

Imaging 

Stimulus-sensitive 

drug release 

Maintain drug levels in the blood, 

thereby improving specificity 

Increase efficiency, reduce toxicity 

Modify nanoparticle pharmacokinetics 

and biodistribution, increasing drug 

efficacy 

Report real-time nanoparticle 

biodistribution 

Control bioavailability, reduces toxicity 

Liposomes and micelles 

5 

Polyethylene glycol 

Modified acrylic acid polymers 

Phospholipid micelles 

Polypeptides 

Antibodies 

Peptides 

Aptamers 

Carbohydrate 

Folic acid 

Peptides 

Trans-activating 

transcriptional activator (TAT) 

Ligands 

Transferrin 

Positively charged moieties 

Cationic lipids 

Cationic polymers 

Quantum dots 

Magnetic nanoparticles 

pH-labile 

Photosensitive 

Thermosensitive 

Magnetic sensitive 

Photothermal sensitive 

Redox sensitive 

One of the earliest forms of nanomedicine was the development of liposomes as 

drug delivery vesicles. As their sizes vary from 100 nm up to a few micrometers they can 

[9] 

[10] 

[11] 

[12] 

[13] 

[14] 

[15] 

[16] 

[17] 

[18] 

[19] 

[20] 

[21] 

[22] 

[23] 

[18] 

[24] 

[25] 

[26] 

[27] 

[28]


Introduction 

carry within their phospholipid bilayered membrane different types of vesicles (e.g. small 

molecules to proteins, peptides, DNA, magnetic nanoparticles). 29,30,31 

Liposomes are generally considered non-toxic, biodegradable and non- 

immunogenic, although recently some adverse effects have been described. .32 Although 

the underlying mechanism has not yet been fully elucidated, it appears that the choice of 

liposome characteristics plays an important role and that choosing the appropriate lipids 

and their size helps to overcome these unexpected effects. 33,34 

PEGylation is the process of covalent attachment of polyethylene glycol polymer 

chains to another molecule, normally a drug or therapeutic protein. By employing 

PEGylation the circulation time is increased and by combining the liposomes with a 

therapeutic agent the multifunctionality of the vesicles is increased. 35 

Micelles (lipid vesicles of 20 to 100 nm) present similar features as liposomes (long 

circulating behaviour, high biocompatibility and the possibility of increasing the 

functionality number). They also offer a better method of delivering hydrophobic drugs. In 

this system, the hydrophobic parts form the core to minimize their exposure to aqueous 

surroundings, whereas the hydrophilic blocks form the corona-like shell that stabilizes the 

core through direct contact with water. 36 The hydrophobic core is capable of carrying 

pharmaceuticals (poorly water soluble drugs), with high loading capacity (5–25% weight), 

while its hydrophilic shell provides not only a steric protection for the micelle (with 

increased stability in blood) but also functional groups suitable for further micelle 

modification. 37 

Polymeric carriers (polymeric or dendrimers) 

The polymeric carriers are formed by a hydrophobic core that is capable of carrying 

pharmaceuticals (poorly water soluble drugs) with high loading capacity (5–25% weight) 

and a hydrophilic shell that provides not only a steric protection for the micelle (with 

increased stability in blood) but also functional groups suitable for further micelle 

modification (repeated relative to previous phrase, reword) 38,39 , such as starch, 40 

poly(ethylene glycol) –PEG, 41,42 poly(lactic-coglycolic acid) -PLGA 43,44 ). 

Dendrimers are a relatively novel class of synthetic polymers with highly ordered 

structure. They are highly branched (


density of functional groups on the surface. Their functionalisation possibilities, symmetry 

perfection, diameters in the range of 10 to 100 nm and internal cavities provide many 

potential applications in biochemistry, gene therapy and nanomedicine. Nanoparticles 

coated with dendrimers can alter the charge, functionality, reactivity and enhance their 

both stability and dispersibility. 45,46,47 

Core – shell/corona structures (Inorganic or magnetic nanoparticles, 

quantum dots or carbon nanotubes) 

The core – shell/corona architecture is another attractive alternative approach for 

developing multifunctional nanoparticles. These structures normally comprise optical, 

targeting and delivery functions. Nanoparticles visualisation is normally achieved through 

the core’s functionality. The shell/corona protects the core and acts as the scaffold for 

adding further functions, such as drug delivery (drugs are incorporated and released in a 

controlled manner, either via diffusion or active delivery via pH), temperature modulation, 

active targeting (through surface modification with specific ligands), or even a tool for an 

alternative way of visualisation through the incorporation of a different imaging agent. A 

large variety of combinations of materials that are assembled in core-shell/corona 

architecture can be found in the literature as the core quantum dots 48,49,50 , fluorophores 51 

or magnetic nanoparticles 52,53 are the most researched moieties. The most commonly 

used coating layers are either organic layers (polymethyl methacrylate) (PMMA), 54 

polylactic acid (PLA), 55,56 poly(lactic-co-glycolic acid) (PLGA), 26,43 polyvinyl alcohol 

(PVA) 25,57 ) or inorganic materials as silica (amorphous, 58 mesoporous 59 ), apatite, 60 or have 

a metallic layer of gold 61 or silver. 62 

Extensive in vivo application of nanoparticles will require a more exhaustive 

exploration of the physicochemical and physiological processes occurring in the biological 

environments. The smallest capillaries in the human body are 4-6 μm in diameter. When 

considering possible medical applications (both in-vivo or in-vitro) the nanoparticles have 

to possess some basic properties. Specifically they must have a size of less than 20 nm 

(especially for intraveneous (i.v.) administration), high surface area with a larger platform 

for surface functionalisation, high colloidal stability (with minimum agglomeration and 

aggregation) and the ability to overcome biological barriers. 63 For the particular case of in- 

7


Introduction 

vivo applications the particles should adhere to more strict requirements: non-toxicity, 

non-immunogenicity, long term retention during blood circulation and the ability to reach, 

and pass through, the endothelial capillary membranes without causing an embolism of 

the bigger vessels. 64,65 

The three major applications of nanoparticles in medicine are imaging (diagnosis), 

drug delivery (therapeutics) and cell reconstruction. In this thesis I report the design and 

synthesis of nanoparticles for bi-modal imaging with potential to be further developed in 

the future with the addition of cell-targeting agents and drug delivery molecules. 

1.2. SCOPE AND OBJECTIVES OF THIS THESIS 

Nanoparticles have to meet several specifications in order to be applied in-vivo and 

in-vitro. Important features these particles should have are a small overall size, high 

surface area for surface functionalisation, high colloidal stability and the ability to pass the 

biological barriers. 

To achieve the proposed objective two sets of contrast agents were designed, 

synthesized and characterized: 

T1 contrast agent with an optical Ln 3+ probe 

The nanoparticles we will consider are hybrid inorganic-organic SiO2@APS 

core-corona. Lanthanide chelates are linked onto these particles, for their 

relaxometry (Gd 3+ ) and luminescence (Eu 3+ or Tb 3+ ) properties. Subsequently, 

these nanoparticles are internalised in living cells in the form of cellular pellets 

in order to confirm their potential as new probes for bimodal imaging. 

T2 contrast agent 

These are core-shell nanoparticles, Fe2O3@SiO2. The influence of the silica 

shell on the relaxometry and cytotoxicity properties was studied. 

8


1.3. OUTLINE OF THIS THESIS 

This thesis is organised in six chapters, as follows. 

Chapter 1 provides an overview of different multifunctional nanoparticles for 

applications in medicine. 

Chapter 2 presents a general introduction of magnetic resonance imaging and MRI 

contrast agents, as well as the parameters governing the relaxivity of the latter, presenting 

a summary of the various classes of contrast agents, classified according to their 

applications. 

Chapter 3 gives a brief introduction to optical imaging and its main applications are 

presented. A short overview is provided of the photoluminescence phenomenon, in 

particular of trivalent lanthanides. The optical properties of tissues are succinctly 

described as well as the different techniques available for optical imaging. A summary of 

the various types of optical contrast agents is also presented. 

Chapter 4 describes the synthesis of T1 contrast agents, with an optical Ln 3+ probe 

attached. In order to optimise the nanoparticles emission, two types of chelating agents, 

with and without an aromatic ring, were used. The physical and chemical properties were 

studied by several different techniques. The uptake of these nanoparticles by living RAW 

cells was evaluated using cellular pellets. The optical and magnetic features of these 

probes in the cellular pellet were assessed in order to confirm their potential as bimodal 

imaging agents. 

Chapter 5 reports the synthesis both, the core of the magnetic iron oxide 

nanoparticles, with a well-defined size, and the silica shellI shall show that the silica-shell 

thickness determines the relaxometry properties, and I shall propose a tentative 

theoretical to explain this fact. The effect of these nanoparticles on the viability and the 

mitochondrial dehydrogenase expression of microglial cells were also evaluated. The 

magnetic properties of these nanoparticles and their T1 relaxivities as a function of the 

external magnetic field (Nuclear Magnetic Relaxation Dispersion (NMRD) plots) were also 

studied in detail and rationalized on the basis of current theory. 

the future. 

Finally, chapter 6 presents the concluding remarks of this thesis and an outlook for 

9


Introduction 

1.4. REFERENCES 

1 Bawarski, W. E.; Chidlowsky, E.; Bharali, D. J.; Mousa, S. A., Emerging 

nanopharmaceuticals. Nanomedicine-Nanotechnology Biology and Medicine 2008, 4 (4), 

273-282. 

2 Freitas, R. A., Jr., What is nanomedicine? Nanomedicine : nanotechnology, biology, and 

medicine 2005, 1 (1), 2-9. 

3 Caruthers, S. D.; Wickline, S. A.; Lanza, G. M., Nanotechnological applications in medicine. 

Current Opinion in Biotechnology 2007, 18 (1), 26-30. 

4 National Nanotechnology Initiative. What is nanotechnology? http://www.nano.gov/html/ 

facts/whatIsNano.html Accessed 10 January 2010. 

5 Thompson, M., Nanomedicine - A tremendous research opportunity for analytical chemists. 

Analyst 2004, 129 (8), 671-671. 

6 Moghimi, S. M.; Hunter, A. C.; Murray, J. C., Nanomedicine: current status and future 

prospects. Faseb Journal 2005, 19 (3), 311-330. 

7 Sanvicens, N.; Marco, M. P., Multifunctional nanoparticles - properties and prospects for 

their use in human medicine. Trends in Biotechnology 2008, 26 (8), 425-433. 

8 Liu, Z.; Kiessling, F.; Gaetjens, J., Advanced Nanomaterials in Multimodal Imaging: Design, 

Functionalization, and Biomedical Applications. Journal of Nanomaterials 2010. 

9 Gref, R.; Minamitake, Y.; Peracchia, M. T.; Trubetskoy, V.; Torchilin, V.; Langer, R., 

Biodegradable long-circulating polymeric nanospheres. Science 1994, 263 (5153), 1600- 

1603. 

10 Sathe, T. R.; Agrawal, A.; Nie, S., Mesoporous silica beads embedded with semiconductor 

quantum dots and iron oxide nanocrystals: Dual-function microcarriers for optical encoding 

and magnetic separation. Analytical Chemistry 2006, 78 (16), 5627-5632. 

11 Dubertret, B.; Skourides, P.; Norris, D. J.; Noireaux, V.; Brivanlou, A. H.; Libchaber, A., In 

vivo imaging of quantum dots encapsulated in phospholipid micelles. Science 2002, 298 

(5599), 1759-1762. 

12 Pinaud, F.; King, D.; Moore, H. P.; Weiss, S., Bioactivation and cell targeting of 

semiconductor CdSe/ZnS nanocrystals with phytochelatin-related peptides. Journal of the 

American Chemical Society 2004, 126 (19), 6115-6123. 

13 Lukyanov, A. N.; Elbayoumi, T. A.; Chakilam, A. R.; Torchilin, V. P., Tumor-targeted 

liposomes: doxorubicin-loaded long-circulating liposomes modified with anti-cancer 

antibody. Journal of Controlled Release 2004, 100 (1), 135-144. 

14 Akerman, M. E.; Chan, W. C. W.; Laakkonen, P.; Bhatia, S. N.; Ruoslahti, E., Nanocrystal 

targeting in vivo. Proceedings of the National Academy of Sciences of the United States of 

America 2002, 99 (20), 12617-12621. 

10


15 Farokhzad, O. C.; Cheng, J. J.; Teply, B. A.; Sherifi, I.; Jon, S.; Kantoff, P. W.; Richie, J. P.; 

Langer, R., Targeted nanoparticle-aptamer bioconjugates for cancer chemotherapy in vivo. 

Proceedings of the National Academy of Sciences of the United States of America 2006, 

103 (16), 6315-6320. 

16 Zhu, J.; Xue, J.; Guo, Z.; Zhang, L.; Marchant, R. E., Biomimetic glycoliposomes as 

nanocarriers for targeting P-selectin on activated platelets. Bioconjugate Chemistry 2007, 

18 (5), 1366-1369. 

17 Kukowska-Latallo, J. F.; Candido, K. A.; Cao, Z. Y.; Nigavekar, S. S.; Majoros, I. J.; 

Thomas, T. P.; Balogh, L. P.; Khan, M. K.; Baker, J. R., Nanoparticle targeting of anticancer 

drug improves therapeutic response in animal model of human epithelial cancer. Cancer 

Res 2005, 65 (12), 5317-5324. 

18 Sawant, R. M.; Hurley, J. P.; Salmaso, S.; Kale, A.; Tolcheva, E.; Levchenko, T. S.; 

Torchilin, V. P., "SMART" drug delivery systems: Double-targeted pH-responsive 

pharmaceutical nanocarriers. Bioconjugate Chemistry 2006, 17 (4), 943-949. 

19 Bartlett, D. W.; Su, H.; Hildebrandt, I. J.; Weber, W. A.; Davis, M. E., Impact of tumor- 

specific targeting on the biodistribution and efficacy of siRNA nanoparticles measured by 

multimodality in vivo imaging. Proceedings of the National Academy of Sciences of the 

United States of America 2007, 104 (39), 15549-15554. 

20 Li, W.; Szoka, F. C., Jr., Lipid-based nanoparticles for nucleic acid delivery. Pharmaceutical 

Research 2007, 24 (3), 438-449. 

21 Luten, J.; van Nostruin, C. F.; De Smedt, S. C.; Hennink, W. E., Biodegradable polymers as 

non-viral carriers for plasmid DNA delivery. Journal of Controlled Release 2008, 126 (2), 

97-110. 

22 Derfus, A. M.; Chen, A. A.; Min, D.-H.; Ruoslahti, E.; Bhatia, S. N., Targeted quantum dot 

conjugates for siRNA delivery. Bioconjugate Chemistry 2007, 18 (5), 1391-1396. 

23 Medarova, Z.; Pham, W.; Farrar, C.; Petkova, V.; Moore, A., In vivo imaging of siRNA 

delivery and silencing in tumors. Nature Medicine 2007, 13 (3), 372-377. 

24 Skirtach, A. G.; Javier, A. M.; Kreft, O.; Koehler, K.; Alberola, A. P.; Moehwald, H.; Parak, 

W. J.; Sukhorukov, G. B., Laser-induced release of encapsulated materials inside living 

cells. Angewandte Chemie-International Edition 2006, 45 (28), 4612-4617. 

25 Stover, T. C.; Kim, Y. S.; Lowe, T. L.; Kester, M., Thermoresponsive and biodegradable 

linear-dendritic nanoparticles for targeted and sustained release of a pro-apoptotic drug. 

Biomaterials 2008, 29 (3), 359-369. 

26 Hu, S.-H.; Liu, T.-Y.; Huang, H.-Y.; Liu, D.-M.; Chen, S.-Y., Magnetic-sensitive silica 

nanospheres for controlled drug release. Langmuir 2008, 24 (1), 239-244. 

27 Park, H.; Yang, J.; Seo, S.; Kim, K.; Suh, J.; Kim, D.; Haam, S.; Yoo, K.-H., Multifunctional 

nanoparticles for photothermally controlled drug delivery and magnetic resonance imaging 

enhancement. Small 2008, 4 (2), 192-196. 

11


Introduction 

28 Trewyn, B. G.; Giri, S.; Slowing, I. I.; Lin, V. S. Y., Mesoporous silica nanoparticle based 

controlled release, drug delivery, and biosensor systems. Chemical Communications 2007, 

(31), 3236-3245. 

29 Torchilin, V. P., Recent advances with liposomes as pharmaceutical carriers. Nature 

Reviews Drug Discovery 2005, 4 (2), 145-160. 

30 Koning, G. A.; Krijger, G. C., Targeted multifunctional lipid-based nanocarriers for image- 

guided drug delivery. Anti-Cancer Agents in Medicinal Chemistry 2007, 7 (4), 425-440. 

31 Mezei, M.; Gulasekharam, V., Liposomes - a selective drug delivery system for the topical 

route of administration I. Lotion dosage form Life Sciences 1980, 26 (18), 1473-1477. 

32 Chanan-Khan, A.; Szebeni, J.; Savay, S.; Liebes, L.; Rafique, N. M.; Alving, C. R.; Muggia, 

F. M., Complement activation following first exposure to pegylated liposomal doxorubicin 

(Doxil): possible role in hypersensitivity reactions. Annals of Oncology 2003, 14 (9), 1430- 

1437. 

33 Ishida, T.; Harada, M.; Wang, X. Y.; Ichihara, M.; Irimura, K.; Kiwada, H., Accelerated blood 

clearance of PEGylated liposomes following preceding liposome injection: Effects of lipid 

dose and PEG surface-density and chain length of the first-dose liposomes. Journal of 

Controlled Release 2005, 105 (3), 305-317. 

34 Ishida, T.; Ichikawa, T.; Ichihara, M.; Sadzuka, Y.; Kiwada, H., Effect of the physicochemical 

properties of initially injected liposomes on the clearance of subsequently injected 

PEGylated liposomes in mice. Journal of Controlled Release 2004, 95 (3), 403-412. 

35 Sun, C.; Lee, J. S. H.; Zhang, M., Magnetic nanoparticles in MR imaging and drug delivery. 

Advanced Drug Delivery Reviews 2008, 60 (11), 1252-1265. 

36 Torchilin, V. P., Recent advances with liposomes as pharmaceutical carriers. Nature 

Reviews Drug Discovery 2005, 4 (2), 145-160. 

37 Zhang, L.; Gu, F. X.; Chan, J. M.; Wang, A. Z.; Langer, R. S.; Farokhzad, O. C., 

Nanoparticles in medicine: Therapeutic applications and developments. Clinical 

Pharmacology & Therapeutics 2008, 83 (5), 761-769. 

38 Mikhaylova, M.; Kim, D. K.; Berry, C. C.; Zagorodni, A.; Toprak, M.; Curtis, A. S. G.; 

Muhammed, M., BSA immobilization on amine-functionalized superparamagnetic iron oxide 

nanoparticles. Chemistry of Materials 2004, 16 (12), 2344-2354. 

39 Raynal, I.; Prigent, P.; Peyramaure, S.; Najid, A.; Rebuzzi, C.; Corot, C., Macrophage 

endocytosis of superparamagnetic iron oxide nanoparticles - Mechanisms and comparison 

of Ferumoxides and Ferumoxtran-10. Investigative Radiology 2004, 39 (1), 56-63. 

40 Kim, D. K.; Mikhaylova, M.; Wang, F. H.; Kehr, J.; Bjelke, B.; Zhang, Y.; Tsakalakos, T.; 

Muhammed, M., Starch-coated superparamagnetic nanoparticles as MR contrast agents. 

Chemistry of Materials 2003, 15 (23), 4343-4351. 

12


41 Kim, D. K.; Mikhaylova, M.; Zhang, Y.; Muhammed, M., Protective coating of 

superparamagnetic iron oxide nanoparticles. Chemistry of Materials 2003, 15 (8), 1617- 

1627. 

42 Mikhaylova, M.; Kim, D. K.; Bobrysheva, N.; Osmolowsky, M.; Semenov, V.; Tsakalakos, T.; 

Muhammed, M., Superparamagnetism of magnetite nanoparticles: Dependence on surface 

modification. Langmuir 2004, 20 (6), 2472-2477. 

43 Naik, S.; Carpenter, E. E., Poly(D,L-lactide-co-glycolide) microcomposite containing 

magnetic iron core nanoparticles as a drug carrier. Journal of Applied Physics 2008, 103 

(7). 

44 Cheng, J.; Teply, B. A.; Sherifi, I.; Sung, J.; Luther, G.; Gu, F. X.; Levy-Nissenbaum, E.; 

Radovic-Moreno, A. F.; Langer, R.; Farokhzad, O. C., Formulation of functionalized PLGA- 

PEG nanoparticles for in vivo targeted drug delivery. Biomaterials 2007, 28 (5), 869-876. 

45 Bawarski, W. E.; Chidlowsky, E.; Bharali, D. J.; Mousa, S. A., Emerging 

nanopharmaceuticals. Nanomedicine-Nanotechnology Biology and Medicine 2008, 4 (4), 

273-282. 

46 Lee, C. C.; MacKay, J. A.; Frechet, J. M. J.; Szoka, F. C., Designing dendrimers for 

biological applications. Nature Biotechnology 2005, 23 (12), 1517-1526. 

47 Koo, O. M.; Rubinstein, I.; Onyuksel, H., Role of nanotechnology in targeted drug delivery 

and imaging: a concise review. Nanomedicine : nanotechnology, biology, and medicine 

2005, 1 (3), 193-212. 

48 Darbandi, M.; Thomann, R.; Nann, T., Single quantum dots in silica spheres by 

microemulsion synthesis. Chemistry of Materials 2005, 17 (23), 5720-5725. 

49 Koole, R.; van Schooneveld, M. M.; Hilhorst, J.; Donega, C. d. M.; t Hart, D. C.; van 

Blaaderen, A.; Vanmaekelbergh, D.; Meijerink, A., On the incorporation mechanism of 

hydrophobic quantum dots in silica spheres by a reverse microemulsion method. Chemistry 

of Materials 2008, 20 (7), 2503-2512. 

50 Hu, S.-H.; Liu, D.-M.; Tung, W.-L.; Liao, C.-F.; Chen, S.-Y., Surfactant-Free, Self- 

Assembled PVA-Iron Oxide/Silica Core-Shell Nanocarriers for Highly Sensitive, 

Magnetically Controlled Drug Release and Ultrahigh Cancer Cell Uptake Efficiency. 

Advanced Functional Materials 2008, 18 (19), 2946-2955. 

51 Burns, A.; Ow, H.; Wiesner, U., Fluorescent core-shell silica nanoparticles: towards "Lab on 

a Particle" architectures for nanobiotechnology. Chemical Society Reviews 2006, 35 (11), 

1028-1042. 

52 Heitsch, A. T.; Smith, D. K.; Patel, R. N.; Ress, D.; Korgel, B. A., Multifunctional particles: 

Magnetic nanocrystals and gold nanorods coated with fluorescent dye-doped silica shells. J 

Solid State Chem 2008, 181 (7), 1590-1599. 

13


Introduction 

53 Modak, S.; Karan, S.; Roy, S. K.; Mukherjee, S.; Das, D.; Chakrabarti, P. K., Preparation 

and characterizations of SiO(2)-coated nanoparticles of Mn(0.4)Zn(0.6)Fe(2)O(4). Journal 

of Magnetism and Magnetic Materials 2009, 321 (3), 169-174. 

54 Ninjbadgar, T.; Yamamoto, S.; Fukuda, T., Synthesis and magnetic properties of the 

gamma-Fe2O3/poly-(methyl methacrylate)-core/shell nanoparticles. Solid State Sciences 

2004, 6 (8), 879-885. 

55 Gomez-Lopera, S. A.; Arias, J. L.; Gallardo, V.; Delgado, A. V., Colloidal stability of 

magnetite/poly(lactic acid) core/shell nanoparticles. Langmuir 2006, 22 (6), 2816-2821. 

56 Chen, F.; Gao, Q.; Hong, G.; Ni, J., Synthesis of magnetite core-shell nanoparticles by 

surface-initiated ring-opening polymerization of L-lactide. Journal of Magnetism and 

Magnetic Materials 2008, 320 (13), 1921-1927. 

57 Liu, T.-Y.; Huang, L.-Y.; Hu, S.-H.; Yang, M.-C.; Chen, S.-Y., Core-shell magnetic 

nanoparticles of heparin conjugate as recycling anticoagulants. Journal of Biomedical 

Nanotechnology 2007, 3 (4), 353-359. 

58 Pinho, S. L. C.; Pereira, G. A.; Voisin, P.; Kassem, J.; Bouchaud, V.; Etienne, L.; Peters, J. 

A.; Carlos, L. D.; Mornet, S.; Geraldes, C. F. G. C.; Rocha, J.; Delville, M.-H., Fine Tuning 

of the Relaxometry of gamma-Fe(2)O(3)@SiO(2) Nanoparticles by Tweaking the Silica 

Coating Thickness. Acs Nano 2010, 4 (9), 5339-5349. 

59 Liu, H.-M.; Wu, S.-H.; Lu, C.-W.; Yao, M.; Hsiao, J.-K.; Hung, Y.; Lin, Y.-S.; Mou, C.-Y.; 

Yang, C.-S.; Huang, D.-M.; Chen, Y.-C., Mesoporous silica nanoparticles improve magnetic 

labeling efficiency in human stem cells. Small 2008, 4 (5), 619-626. 

60 Pon-On, W.; Meejoo, S.; Tang, I. M., Substitution of manganese and iron into 

hydroxyapatite: Core/shell nanoparticles. Materials Research Bulletin 2008, 43 (8-9), 2137- 

2144. 

61 Wang, L. Y.; Luo, J.; Fan, Q.; Suzuki, M.; Suzuki, I. S.; Engelhard, M. H.; Lin, Y. H.; Kim, N.; 

Wang, J. Q.; Zhong, C. J., Monodispersed core-shell Fe3O4@Au nanoparticles. Journal of 

Physical Chemistry B 2005, 109 (46), 21593-21601. 

62 Lai, C. H.; Wu, T. F.; Lan, M. D., Synthesis and property of core-shell Ag@Fe3O4 

nanoparticles. IEEE T Magn 2005, 41 (10), 3397-3399. 

63 Kreyling, W. G.; Semmler-Behnke, M.; Moeller, W., Health implications of nanoparticles. 

Journal of Nanoparticle Research 2006, 8 (5), 543-562. 

64 Pathak, P.; Katiyar, V. K., Multifunctional nanoparticles and their role in cancer drug delivery 

– a review. Journal of Nanotechnology Online 2007, 3, 1 - 17. 

65 Tartaj, P.; Morales, M. D.; Veintemillas-Verdaguer, S.; Gonzalez-Carreno, T.; Serna, C. J., 

The preparation of magnetic nanoparticles for applications in biomedicine. Journal of 

Physics D-Applied Physics 2003, 36 (13), R182-R197. 

14


15 

2. 

Magnetic Resonance Imaging 

Background Concepts


Magnetic Resonance Imaging Background Concepts 

2.1. Magnetic Resonance Imaging 17 

2.2. MRI Contrast Agents 22 

2.2.1. Relaxation 24 

2.2.1.1. Spin-Lattice Relaxation 24 

2.2.1.2. Spin-Spin Relaxation 26 

2.2.2. Relaxivity 28 

2.2.2.1. Inner Sphere Relaxivity 30 

2.2.2.2. Second and Outer Relaxivity 35 

2.2.2.3. Paramagnetic Relaxation Parameters 37 

2.2.3. Nuclear Magnetic Resonance Dispersion 39 

2.3. Classification of CAs 44 

2.3.1. Chemical Composition, Magnetic Properties and Effect on 

the MRI Image 

2.3.2. Biodistribution and Applications 48 

2.3.2.1. Non-specific Agents 48 

2.3.2.2. Specific or Targeted Agents 52 

2.3.2.3. Nono-injectable Organ Specific Agents 56 

2.3.2.4. Responsive, Smart or Bioactivated Agents 58 

2.3.2.5. Contrast Agents Based on Other Properties 63 


16 

45


2.1. MAGNETIC RESONANCE IMAGING 

Magnetic Resonance Imaging (MRI) is a non-invasive diagnostic tool providing high- 

resolution ( m 

scale) anatomical images of soft tissue and allowing the quantitative 

assessment of disease pathogenesis by measuring up-regulated biomarkers. MRI is 

based on the principles of Nuclear Magnetic Resonance (NMR), a spectroscopic 

technique applied by scientists to obtain microscopic chemical and physical information 

about molecules. 

Felix Bloch and Edward Purcell independently discovered the phenomenon of 

magnetic resonance in 1946 and for this they were awarded the Nobel Prize in 1952. Until 

the 1970s the NMR technique was used only for chemical and physical analysis. The first 

step towards the discovery of the magnetic resonance was taken in 1971 by Raymond 

Damadian who showed that the nuclear magnetic relaxation times of healthy and 

pathogenic tissues are different. 1 It was only in 1973, however, that imaging moved from 

the single dimension of NMR spectroscopy to the second dimension of spatial orientation 

with the discovery of NMR-zero field gradients by Paul Lauterbur. He published this new 

imaging technique in a short and concise paper in Nature entitled "Image formation by 

induced local interaction; examples employing magnetic resonance". 2 

MRI is an extension of nuclear magnetic resonance spectroscopy used in chemistry. 

The nuclei of certain atoms have an inherent magnetic dipole moment resulting from their 

electrical charge and spin. If these nuclei are placed in a strong, static magnetic field, B 0 , 

they precess around it at the Larmor frequency, I 

determined by their orientation along the axis parallel to 0 

17 

. The energy levels of the spins are 

B (either with or against 0 

There are more spins aligned with the field (parallel - low energy state) than spins aligned 

against the field (anti-parallel- high energy state). Due to this slight excess of parallel 

spins, the net equilibrium magnetization (macroscopic magnetization) is parallel to 

B ). 

B 0 (figure 2.1a). The difference in energy between the two spin states increases as the 

magnetic field strength increases. 

Besides the static 0 B field applied along z, we apply a time varying field 1 

perpendicularly to 0 B and oscillating at 0 , which perturbs the populations of the spin-up 

B ,



and spin-down states and gives phase coherence to the spins. A 90º pulse ( 1 

18 

B along the x 

axis) brings the magnetisation vector to the xy plane (figure 2.2b). In this case, the 

magnetisation precesses about 0 

B at 0 and 1 

B at 1 

. At this point, it is appropriate to 

introduce a new frame of reference for viewing the evolution of the magnetisation vector, 

the rotating frame (x',y',z) rotating about the z-axis at frequency 0 

the magnetisation precesses about the x' axis with frequency 1 

1 

. In the rotating frame 

B 

(figure 2.3c). 

Figure 2.1. a) Precession of the spins in a static the external magnetic field 

and equilibrium magnitisation b) Precession of the net magnitisation in the 

laboratory frame under the influence of the static 0 

B , 

B field and the RF field 1 

c) Net magnitisation viewed in the rotating frame after a on-resonance 90º (x) 

pulse. 3,4 

In MRI the unpaired nuclear spins (mainly from hydrogen atoms in water - 70% to 

90% of most tissues - and organic compounds) align themselves when exposed to a 

magnetic field. A temporary radiofrequency pulse at the Larmor frequency changes the 

alignment of the spins, and their return to thermal equilibrium (relaxation) is detected in a 

coil. Protons from different tissues react differently giving a picture of anatomical 

structures. 5 

c) b) c) b) 

a) 

B 0 

c) 

B 0 

B 1 

, 

,


In order to create a MRI image of a patient it is necessary to create different voxels 

(volume picture elements) of its volume. This is accomplished by applying perpendicular 

magnetic gradients. MRI allows a complete and flexible image orientation; although 

normally in clinical use the reference employed for the magnetic gradient is the principal 

axis of the patient (patient’s images are obtained in the x and y axis while the z axis is the 

head to toe direction). These gradients are small perturbations superimposed on the main 

magnetic field 0 

B , with a typical imaging gradient producing a total field variation of less 

than 1%. Therefore, in the presence of a field gradient, each proton will resonate at a 

unique frequency that depends on its exact position within the gradient field. The spatial 

encoding is obtained by applying magnetic field gradients which encode the position within 

the phase of the signal. 

In the case of one dimension, a linear relationship of phase with respect to position 

can be obtained via collection of data in the presence of a magnetic field gradient. In the 

case of three dimensions (3D), a plane can be defined by "slice selection" where a RF 

pulse of defined bandwidth is applied in the presence of a magnetic field gradient (Figure 

2.2). This procedure is applied in order to reduce spatial encoding to two dimensions (2D). 

Figure 2.2 is an illustration of slice selection during the RF excitation phase using a 

longitudinal magnetic field gradient. The typical medical resolution is about 1 

research models can exceed 1 

3 

m . 6 

Figure 2.2. Illustration of slice selection during the RF excitation phase using 

a longitudinal magnetic field gradient 5 

19 

3 

mm , while



Although the most common sequence is a spin-echo sequence, there are numerous 

pulse sequence types. The spin-echo sequence (Figure 2.3) consists on the application of 

a 90º pulse to the spin system. This 90º pulse rotates the magnetization to the x’y’ plane, 

where the transverse magnetization begins to diphase. Following the 90º pulse a 180º 

rephasing pulse at TE/2 is applied. This pulse rotates the magnetization by 180º about the 

x’ axis, forcing the magnetization to at least partially rephase and generating a signal 

called a spin echo. In terms of time this is an echo time (TE). 

This pulse sequence is repeated at each time interval TR (Repetition Time). With 

each repetition a k-space line acquisition (often refers to the temporary image space, 

usually a matrix, in which data from digitized MR signals are stored during data acquisiton) 

is filled due to a different phase encoding. Three types of gradients are applied: slice 

selection (GS), phase encoding (GP) and frequency encoding (GF). 

One of the advantages of using a spin-echo sequence is that it introduces 2 

20 

T (see 

2.2.1 for definition) dependence to the signal intensity, since some tissues and 

pathologies have similar 1 

T values but different 2 

imaging sequence that can produce images with 2 

T values. It is, thus, of interest to have an 

T dependence (since the 180º 

rephasing pulse compensates for the constant field heterogeneities to obtain an echo that 

is weighted in 2 

T and not in 

weighted proton or spin density and 2 

high signal-to-noise ratio. 

* 

T 2 ). The pulse sequence timing can be adjusted to give 1 

T - 

T -weighted images. This sequence also provides a 

A spin-echo sequence has two essential parameters: repetition time (TR) and echo 

time (TE), in which TR is defined as the time between repetitions of the sequence and TE 

is the time between the 90º pulse and the maximum amplitude in the echo, respectively. 

This sequence has a clear contrast mechanism for a T1 -weighted image, where a short 

TR (450 – 850 ms) and TE (10 – 30 ms) are used. In the case of a T2 -weighted image a 

long TR (>2000 ms) and TE (>60 ms) are used. In the case of 1 H images a long TR with a 

short TE are used.


Figure 2.3. Spin-echo pulse sequence used in MRI. 7 

The signal intensity (SI) is given by Eq. 2.1, where k is a proportionality constant 

dependent on flow, perfusion and diffusion, and ρ is the density of spins in the sample. 

TR/ 

T1 

TE 

/ T 2 

1 e e 

SI k 

(2.1) 

Three types of spin echo sequences are commonly used: standard single echo, 

standard multi-echo, and echo-train spin echo. Standard single-echo sequences are 

generally used to produce 1 

T -weighted images when acquired with relatively short TR and 

TE (less than 700 ms and 30 ms, respectively). A multi-slice loop structure is used with a 

single pair of excitation and refocusing pulses applied per slice loop. The pixel intensity is 

proportional to the number of protons contained within the voxel weighted by the T 1 and 

T 2 relaxation times for the tissues within the voxel. The echo-train spin echo sequence is 

21



similar to the multi-echo sequences, except that each spin echo is acquired with a 

different phase encoding as well as its own TE. The echo-train length (or turbo factor) is 

the number of echoes acquired in each TR period. Here, k-space is filled segmentally with 

one echo from each echo-train filling each segment of k-space and, as such, is an efficient 

sequence. These sequences are typically used to create 2 

22 

T -weighted images. 

Nevertheless, at times insufficient contrast is observed and the administration of 

Contrast Agents (CAs) or Contrast Media is necessary. MRI CAs are able to change the 

1 H relaxation properties of the tissues, leading to MR images with improved contrast. In 

MRI the object or patient is exposed to a powerful magnetic field. The hydrogen protons of 

the water molecules in the object or body are excited by radio-frequency pulses. During 

the time they are “recovering” (or relaxing back) they transmit signals that are recorded 

and compiled into an image. The faster the atomic nuclei return from an excited to an 

unexcited state, the stronger the signal and thus also the contrast. Most approved MRI 

contrast agents are based on the rare earth element gadolinium, adding paramagnetic 

properties to the compound. This element causes the atomic nuclei to relax more quickly 

and thus to transmit stronger signals. 

2.2. MRI CONTRAST AGENTS 

One of the strengths of MRI is the significant amount of intrinsic contrast between 

tissues. This contrast is due to differences in the longitudinal ( 1 

T ) 

T ) and/or transverse ( 2 

proton relaxation times of the tissues under observation accentuated by the chosen TR 

and TE. Pathologic tissue may or may not exhibit significant differences in 1 

T 

T or 2 

relatively to the surrounding normal tissue. For this reason, there may be little signal 

difference between normal and pathologic tissue in spite of the inherent high contrast in 

the images. An example of difference between a 1 

Figure 2.4a and b. 

T and 2 

T -weighted image is given in


aa bb 

cc dd 

Figure 2.4. Brain MRI images with a) w 

of a commercial CA and d) w 

T 1 , b) w 

23 

T 2 , c) w 

T 1 with the administration of a commercial CA 8 

T 1 without the administration 

A MRI Contrast Agent (CA) is a chemical substance introduced to the anatomical or 

functional region being imaged in order to increase the differences between different 

tissues, or between normal and abnormal tissue, by altering the relaxation times. 

Therefore, the aim of using CAs in MRI is to accelerate the relaxation of water proton 

spins with their surrounding. This can be achieved by using paramagnetic substances. 

Examples of these achievements are the experiments of Bloch et al. 9 and Lauterbur et 

al. 10 

The MRI contrast mechanism can be affected by many intrinsic and extrinsic factors 

such as proton density and MRI pulse sequences. Based on their relaxation processes the 

contrast agents can be classified as T 1 or T 2 contrast agents. The commercially approved 

T 1 contrast agents are usually paramagnetic complexes, while the 2 

T contrast agents are 

based on iron oxide nanoparticles, which are the most representative nanoparticulate 

agents.



The most common paramagnetic complexes are the paramagnetic chelates 

containing a lanthanide ion. Currently, gadolinium (III) complexes are by far the most 

widely used CAs in clinical practice because it has seven unpaired electrons making it the 

most paramagnetic (highest spin density) stable metal ion. Gadolinium (III) also has 

another significant feature due to the symmetrical 8 S ground state: its electron spin 

relaxation is relatively slow, which is relevant to its efficiency as an MRI CA. There are a 

number of excellent reviews 11-18 covering the development and properties of first and 

second-generation contrast agents, particularly focusing on gadolinium complexes. 

11,12,13,14,15,16,17,18 

2.2.1. RELAXATION 

Relaxation is a fundamental precess in MR and presents the main mechanism for 

the image contrast. The RF pulse turns the macroscopic magnetization away from the z- 

axis. Once the 1 

magnetic field ( 0 

B -field has been turned off the magnetization rotate around the main 

B ) at the Lamor frequency. Due to relaxation processes, the 

magnetization will eventually return to its equilibrium position along the z-axis and the 

NMR signal will as a result fade away. There are two different relaxation pathways 

possible: the spin-lattice relaxation and the spin-spin relaxation. 

2.2.1.1. SPIN-LATTICE RELAXATION 

The spin-lattice relaxation, or longitudinal relaxation, or 1 

for the return of the z-magnetization component, Z 

24 

T relaxation, is responsible 

M , to its equilibrium value. It provides 

the mechanism by which the protons give up their energy and return to the thermal 

equilibrium. Therefore, if a 90° pulse at the resonant frequency, 0 , is applied to a sample 

its magnetization, 0 

M , will rotate to the x’y’ plane as shown in Figure 2.4-a, and Z 

M will 

be zero. With time, M Z will increase, as the protons release their energy (Figure 2.4-b). 

This return of magnetization to equilibrium follows an exponential growth process, with a 

time constant T 1 describing the rate of growth (equation 2.2):


0 

) / ( 

T1 

1 

e 

M ( ) M 

(2.2) 

where is the time following the RF pulse. After three T 1 time periods, M will have 

returned to 95% of its value prior to the excitation pulse, M 0 . The term spin-lattice refers 

to the fact that the excited proton (“spin”) transfers its energy to its surroundings (“lattice”) 

rather than to another spin, where the energy will no longer contribute to spin excitation. 

This energy transfer to the surroundings has some very important consequences. The key 

to this energy transfer is the presence of some type of molecular motion (e.g., vibration, 

rotation) in the vicinity of the excited proton with an intrinsic frequency, L 

the resonant frequency, 0 

. The closer 0 

is to L 

25 

, that matches 

, the more willingly the motion absorbs 

the energy and the more frequently this energy transfer occurs, allowing the protons to 

return to equilibrium. 

a) a) a) b) b) b) 

Figure 2.4. a) Effect of a 90º RF pulse on the net magnetization, viewed in the 

rotating frame. b) Plot of the relative M Z component as a function of time 

after the RF pulse, know as the 1 

T relaxation curve. 

In general any mechanism which gives rise to fluctuating magnetic fields is a 

possible relaxation mechanism. Therefore all mechanisms contribute to the observable 

relaxation time according equation 2.3. 

1 

 

1 

 

T1 m T1m 

(2.3)



where m 

T 1 is the time constant of mechanism m . The magnitude of the contribution to the 

overall relaxation rate from a single mechanism may differ greatly among different 

chemical groups, the chemical environment and has a dependence regarding parameters 

such as temperature and field strength. 

2.2.1.2. SPIN-SPIN RELAXATION 

Spin–spin relaxation time or transverse relaxation time or the relaxation time 2 

26 

T is 

the time required for the transverse component of M to decay to 37% of its initial value 

via irreversible processes. The net magnetization 0 

M is initially along the z-axis ( 0 

B ). If a 

90° pulse is applied to a sample, M 0 will rotate as shown in Figure 2.5a-1. This causes 

M 0 to rotate entirely into the x’y’ plane, so that the individual spin magnetic moments 

have phase coherence in the transverse plane at the end of the pulse. With time this 

phase coherence disappears and the value of M in the x’y’ plane decreases toward 0. T 2 

or 

* 

T 2 relaxation is the process by which this transverse magnetization is lost. At the end 

of the 90° RF pulse, when the proton spins have absorbed energy and are oriented in the 

transverse plane, all spins precess at the same frequency 0 

and are synchronized at the 

same point or phase of its precessional cycle. Given that a nearby proton of the same type 

will have the same molecular environment, and the same 0 

, it can easily absorb the 

energy that is being released by its neighbour. Therefore the spin–spin relaxation refers to 

this energy transfer from an excited proton to another nearby proton and the absorbed 

energy remains as spin excitation rather than being transferred to the surroundings as in 

T 1 relaxation. Figure 2.5a illustrates a rotating frame slower than 0 

of the net magnetization after a 90º RF pulse, where: 

B , 

1- Net magnetization M is oriented parallel to 0 

and the various steps 

2- following a 90° RF pulse the protons initially precess in phase in the transverse 

plane, 

3- due to inter- and intramolecular interactions the protons begin to precess at 

different frequencies and become asynchronous with each other 

4 and 5- as more time elapses the transverse coherence becomes smaller until, 

6- there is complete randomness of the transverse components and no coherence.


a) a) a) b) b) b) 

Figure 2.5. a) Illustration of the net magnetization M with regards to the x’y’ 

plane within a rotating frame slower than 0 

component as a function of time, known as the 2 

27 

M 

. b) Plot of the relative XY 

T relaxation curve. 

If the protons continue in close proximity and remain at the same 0 

the proton– 

proton energy transfer will occur many times. The local magnetic field of the protons can 

fluctuate due to intermolecular and intramolecular interactions such as vibrations or 

rotations. This fluctuation produces a gradual, irreversible loss of phase coherence of the 

spins as they exchange energy and reduce the magnitude of the transverse magnetization 

as well as reduce the generated signal (Figure 2.5b). With the increase in time this 

transverse coherence completely disappears, only to reform in the longitudinal direction as 

T 1 relaxation occurs. 

There are two main causes for a loss of transverse coherence to M . The first cause 

is the movement of the adjacent spins due to molecular vibrations or rotations 

(responsible for spin–spin relaxation or the true T 2 ). The second cause is due to the 

inhomogeneity of the magnetic field. As the proton spin precesses it experiences a 

fluctuating local magnetic field, causing a change in 0 and a loss in transverse phase 

coherence. 

The proper design of the pulse sequence can eliminate the imaging gradients as a 

source of dephasing. While, the other sources contribute to the total transverse relaxation 

time, 

* 

T 2 equation 2.4:



1 

1 1 1 

 

(2.4) 

* 

T2 T2 

T2M 

T2MS 

where M 

T 2 is the dephasing time due to the main field inhomogeneity and MS 

dephasing time due to the magnetic susceptibility differences. 

28 

T 2 

is the 

The decay of the transverse magnetization following a 90° rf pulse follows an 

exponential process with the time constant of 

M 

XY 

( t) 

where 

* 

* 

T2 rather than just 2 

T equation 2.5: 

( t 

/ T2 

) 

M XY e 

(2.5) 

max 

M XY is the transverse magnetization 

max 

XY 

excitation pulse. For most tissues or liquids, M 

M immediately following the 

T 2 is the major factor in determining 

whereas for tissue with significant iron deposits or air filled cavities, MS 

T determined using the spin-echo experiment). 

( 2 

2.2.2. RELAXIVITY 

T 2 dominates 

Relaxivity is the ability of magnetic compounds in enhancing the relaxation rate of 

the surrounding water proton spins. Therefore, the longitudinal and transverse relaxivity 

values, 1 

r and 2 

r , refer to the amount of increase in 1 T1 and 2 

millimolar concentration of agent (often given as per mM of Gd). 1 

r2 r1 

ratios of 1-2, whereas 2 

19,20,21,22,23 

T agents normally have 1 

* 

T 2 , 

* 

T 2 

1 T , respectively, per 

T agents usually have 

r2 r ratios ~10 or higher. 19-23 

Solomon, Bloembergen and Morgan developed the general theory of solvent nuclear 

relaxation in the presence of paramagnetic substances. 19-23 . The Gd (III) complexes 

induce an increase of both the longitudinal and transverse relaxation rates, 1 T1 and 2 

1 T , 

respectively, of the solvent nuclei (normally water). Diamagnetic and paramagnetic 

relaxation rates are additive and given by equation 2.6:


1 

T 

1 

1 

, i 

i, 

obs Ti 

, d Ti 

, p 

1 

T 

i, 

obs 

i, 

d 

1, 

2 

1 

ri 

i 

T 

where 

Gd, 1, 

2 

1 is the observed solvent relaxation rate, 

T i, 

obs 

29 

T i, 

d 

(2.6) 

(2.7) 

1 the diamagnetic relaxation 

rate, and 1/Ti,p the paramagnetic relaxation rate that depends on the concentration of 

paramagnetic species expressed in millimolar for diluted samples. Relaxivity, ri, is defined 

as the slope of the concentration dependence, equation 2.7., and is normally expressed in 

mM -1 s -1 ; molal (mol/kg) concentrations should be used when dealing with nondilute 

systems. 

Within the Solomon, Bloembergen and Morgan’s (SBM) theory the paramagnetic 

relaxation process is described on the basis of a model that considers ‘inner sphere’ and 

‘outer-sphere’ contributions (equation 2.8). 

1 

T 

1 

1 

, 

is os i 

i, 

p Ti 

, p Ti 

, p 

1 

T 

1 

1 

1 

1, 

2 

3 

1 Gd 1, 

2 

r 

is os 

i 

i 

i, 

obs Ti 

, p Ti 

, p Ti 

, d 

Ti 

, d 

where T i, 

p , 

(2.8) 

(2.9) 

is 

os 

T i, 

p and T i, 

p are, respectively, the total paramagnetic, the inner- and outer- 

sphere paramagnetic contributions to the longitudinal (i=1) or transverse (i=2) NMR 

relaxation times. The inner-sphere contribution is related to the exchange between the 

bound water molecules and bulk water, and the outer sphere-contribution is caused by 

water molecules diffusing near the paramagnetic centre during their translational 

diffusion. 13,24 Often, a third contribution is also taken into account, the ‘second-sphere’, 

that is caused by the presence of mobile protons or water molecules in the second 

coordination sphere of the paramagnetic species (Figure 2.6). 25 The relaxivity ( r i ) is 

defined in equation 2.9 (combination of eqs. 2.14, 2.15 and 2.16), if the second sphere 

contribution is not considered.



Figure 2.6. Schematic representation of the three types of water molecules 

surrounding the metal complex. 

2.2.2.1. INNER SPHERE RELAXIVITY 

The inner-sphere contribution to the overall relaxation rates is described in 

equations 2.10 and 2.11. The parameters involved in this inner-sphere mechanism are i) 

q , the number of water molecules in the inner-coordination sphere directly coordinated to 

the paramagnetic centre, ii) C , the molar concentration of the paramagnetic complex, iii) 

M , the mean residence lifetime of the coordinated water protons within the inner-sphere 

coordination sphere, iv) T1 M and M 

T 2 , the longitudinal and transverse proton times 

enhancement experienced by the proton of the inner-sphere water molecules, 

 

respectively, and v) M , the chemical shift difference between the free and the bound 

water molecules. 

1 

R 

 

qC 

T 

is 

T1 is 

1 

55. 

5 1M 

M 

30 

(2.10)


C 1 

1 

1 

2 

1 is q T2 

M( 

M T2 

M ) 

R2 

 

is 

 

T2 , p 55. 

5 

M 

M 2M 

M 

M 

 

1 

1 

2 2 

T 

 

 

31 

T 

(2.11) 

In the case where the system is in the ‘fast exchange’ regime ( M ) the inner- 

sphere contribution becomes important and is transferred to the bulk water. When the 

coordinated water molecule is in the ‘slow exchange’ regime ( M ), the water 

exchange rate becomes the limiting factor of the relaxivity. 

1 M 

1 M 

T 

The Solomon-Bloembergen-Morgan theory 18,20 provides a foundation to understand 

the basis of iM 

T (Equations 2.12-18). The two components of the T iM term (Equation 

2.12) are dipole-dipole (arising from random fluctuations of the through-space interaction 

of the nuclear dipole with the unpaired electron dipole) and scalar interactions (resulting 

from a through-bond delocalization of the unpaired spin density on the nucleus) that are 

noted by the “DD” and “SC” superscripts respectively. 

1 

T 

iM 

1 

T 

DD 

i 

1 

 

T 

(i = 1,2) (2.12) 

SC 

i 

2 2 2 

1 DD 2 

 

I g BS 

( S 1) 

o 

3 

c1 

7 

c2 

R1 

. 

. 

DD 

6 

2 2 

2 2 

T1 15 rGdH 

4 

1 

I 

c1 

1 

S 

c 

1 SC 2 A e2 

R1 

S( 

S 1). 

SC 

2 2 

T1 3 1 

S 

e 

1 

T 

DD 

2 

1 

T 

2 

2 2 2 

dip 1 

 

 

I g BS 

( S 1) 

o 

3 

1C 

13 

2C 

R2 

. 

4 

6 

2 2 

2 2 c1 

15 rGdH 

4 

1 

I 

c1 

1 

S 

c2 

 

2 

2 

2 

 

 

 

2 

 

 

 

(2.13) 

(2.14) 

(2.15) 

2 

1 A 

e2 

S( 

S 1). 

 

2 2 e (2.16) 

3 1 

S 

e2 

 

SC 

R SC 2 

1 

2



1 

 

ci 

1 

 

ei 

1 

 

 

M 

1 

 

 

M 

1 

 

 

 

R 

1 

T 

ie 

 

1 

T 

ie 

(i=1,2) (2.17) 

(i=1,2) (2.18) 

where I is the proton gyromagnetic ratio, g is the electronic g-factor, B is the 

Bohr magneton, S is the number of unpaired electrons in the paramagnetic metal ion, 

r GdH is the distance between the water protons and the unpaired electrons of the 

paramagnetic metal ion, 0 

is the magnetic permeability of a vacuum, 

32 

A 

 

 

hyperfine coupling constant between the metal electrons and the water protons, I 

are the nuclear and electron Larmor frequencies, respectively ( B 

S 

I , S I , S 

is the 

and 

, where B 

is the magnetic field). For field strengths used in MRI the electronic contributions (the “7” 

term inside the square brackets in equation 2.13) may be conveniently ignored since, 

2 2 

s c2 

1, 

reducing the electronic contributions to an insignificant amount. The nuclear 

contribution (the “3” term also inside the square brackets in equation 2.13) is determined 

by 

2 

I , the proton Larmor frequency, and 1 

, the local correlation time. The relaxation 

enhancement efficiency of the CA depends on how closely matched the proton frequency 

(i.e., the Larmor frequency, I 

 

 

 

 

1 

T 1, 

c 

 

 

 

 

c 

) is to the correlation frequency of the contrast agent 

. The nuclear contributions to relaxation are maximized when 

the Larmor frequency of the protons. 

The local correlation time ( ci 

 

 

 

 

1 

T 1, 

c 

 

 

 

 

approaches 

), defined by equations 2.17 and 2.18, has three 

components, T 1 , e (the electronic relaxation time of the unpaired electrons, M 

) the water 

residency lifetime, and T R – the rotational correlation lifetime. The high-field strengths 

used in MRI again simplify matters, T 1 , e is long enough to reasonably ignore the


contributions from the 

 

 

 

 

1 

T 1, 

c 

 

 

 

 

term. M is the same term from equation 2.10. R is the 

rotational correlation time and is related to the physical tumbling time of the CA in solution. 

Considering equations 2.13 and 2.15, 

motion limit (in the case of small c 

33 

dip 

R1 and 

dipole-dipole mechanism behaviour differs. In this case 

dip 

R2 are roughly equal at the fast 

), while far from this limit (in the case of long c 

dip 

R1 decreases with c 

and 

) the 

is constantly increasing due to the presence of the frequency independent term c 

Regarding the contact contribution, SC, due to its nature the correlation time modulating 

this contribution is not affected by R (equations 2.14 and 2.16) and since 

extremely large, for Gd 3+ complexes, 

the contact contribution ( 

the metal ion 26 . 

dip 

R 2 

4 . 

2 2 

S e is 

SC 

R1 is usually negligible. In the case of R2, however, 

SC 

R2 ) is often the dominant mechanism, mainly from nuclei near 

At high magnetic fields, especially if R 

is much larger than e 

T i, 

, the Curie or 

susceptibility mechanism is another dipolar effect that must be considered (equations 2.19 

and 2.20). 27 

2 2 2 4 4 

1 6 I Bo 

eff 

 

o 

R 

. . 

 

 

 

2 6 

2 2 

T1, 5 4 

( 3kT 

) r 1 

I 

R 

(2.19) 

2 2 2 4 4 

1 1 I Bo 

eff 

 

o 

3 

R 

. . 

4 

 

2 6 R 

2 2 

T 

 

2 5 4 

( 3kT 

) r 1 

I 

R 

, (2.20) 

where 0 

B is the magnetic field strength, T is the absolute temperature, eff 

is the 

effective magnetic moment of the metal ion, k is the Boltzmann constant, r is the 

distance between the nuclear spin and the metal ion. 

The Curie mechanism describes the dipolar interaction between the nuclear spins 

and the magnetic moment generated by the thermally averaged excess of electron



population in the electronic spins levels. The contribution of the Curie mechanism to the 

total longitudinal relaxation is only relevant for slowly rotating macromolecules (long R 

The electronic relaxation rates ( ie 

34 

). 

T 1 ) described by Bloembergen and Morgan20 and 

McLachlan 28 depend on the magnetic field. For Gd 3+ complexes, these are normally 

interpreted in terms of a modulation of the zero field splitting (ZFS) 29 (only once) These 

relaxation rates are described by Equations. 2.21 and 2.22, referred to as the 

Bloembergen-Morgan theory of paramagnetic electron spin relaxation: 

ZFS 

2 

1 

1 4 

 

 

 

 

V 

. 

2 2 

2 2 

T1e 25 

1 

S 

V 1 

4S 

V 

ZFS 

 

4S 

S 1) 

3 

( (2.21) 

 

2 

1 

5 2 

 

 

 

 

 

V 4S( 

S 1) 

3 . 

 

3 

2 2 

2 2 

(2.22) 

T2e 50 

1 

S 

V 1 

S 

V 

where 

2 

is the mean squared ZFS energy and V is the correlation time for the 

modulation of the ZFS interaction. This modulation results from the transient distortions of 

the complex “metal coordination cage”. A transient ZFS of the spin levels can be induced 

by vibration, intramolecular rearrangement and collisions between solvent molecules and 

the metal complexes, allowing the coupling of rotation with spin transitions. The validation 

of equations 2.21 and 2.22 is restricted to certain conditions where 1 

At low magnetic field ( 1 

 

 

is fulfilled. 

0 0. 

B T) the relaxivity of the Gd3+ complexes depends 

mainly on the electronic relaxation (Equation. 2.23). At high magnetic field ( 5 

0 

V 

B 1. 

T) the 

electronic relaxation rate decreases and becomes slower than the rotational rate of the 

complex (Equation. 2.24). 

1 1 1 

 

 

 

, 

 

 

 

 

 

 

 

T1e T2e 

 

R 

1 1 1 

 

 

 

, 

 

 

 

 

 

 

 

T1e T2e 

 

R 

T 

; T1e c1 

; 2e c2 

 

; R c1 

c2 

0 

(2.23) 

(2.24)


that 

A reported temperature and magnetic field dependence EPR study 30 demonstrated 

1 

3+ 

 

 

 

of various aqueous solutions of a series of Gd complexes is described to a 

T1e 

good approximation by the previously developed equations. The major contribution to the 

observed EPR line widths is due to electronic relaxation. 

A more recent description of electron spin relaxation requires EPR measurements 

over a very wide range of temperatures and magnetic fields. 31 This theoretical model 

shows that the electronic relaxation mechanisms at the origin of the EPR line shape arise 

from the combined effects of the modulation of the static crystal field by the random 

Brownian rotation of the complex and of the transient zero-field splitting. 

2.2.2.2. SECOND AND OUTER SPHERE RELAXIVITY 

Water molecules not directly coordinated to the metal ion also experience relaxation 

enhancement in the presence of the CA. These water molecules may be organized into a 

second- and outer-coordination sphere as shown in Figure 2.6. The Solomon- 

Bloembergen-Morgan theory may also be applied to second-sphere water molecules, thus 

SS 

T1 may be modelled from equations 2.10, 2.14 and 2.15. 

The outer-sphere relaxation enhancement may be modelled using theories 

developed by Hwang and Freed 32,33,34 that take into account the electronic relaxation and 

diffusion. Essentially, OS 

T 1 is determined by je 

T , the electronic relaxation time, 

a non-Lorentzian spectral density function, NA is the Avogadro’s number, S 

gyromagnetic ratio, GdH 

35 

J ; is 

i je T 

is electron 

a is the distance of closest approach of the solvent protons to the 

paramagnetic centre, D GdH is the sum of the diffusion coefficients of the water proton and 

of the Gd 3+ complex and GdH is the diffusion correlation time. The symbols not mentioned 

here maintain the meaning given before. 

1 

T 

os 

1 

2 

os 32 

o 

2 2 2 NA 

R1 

. . I S 

S( 

S 1). 

. 3J( I 

; T1e 

) 7J( 

S 

; T2 

e) 

(2.23) 

405 4 

 

a . D 

GdH 

GdH



 

 

 

J( 

 

i ; Tje) 

Re 

 

 

 

1 

ii 

 

 

aGdH GdH 

DGdH 

2 

 

GdH 

 

 

T 

GdH 

je 

 

 

 

 

1 

1 

ii 

4 

 

1 

2 

4 

ii 

9 

 

36 

GdH 

GdH 

 

 

T 

 

 

T 

GdH 

je 

GdH 

je 

 

 

 

 

1 

2 

 

 

1 

ii 

 

 

9 

 

GdH 

 

 

T 

j = 1, 2 ; i = I, S (2.24) 

(2.25) 

In the case of small-sized Gd 3+ complexes the outer sphere contribution is 

responsible for about 50% of the total relaxivity. For macromolecular systems, the outer 

sphere contribution may be considered less important. Complexes of similar shape and 

size have similar diffusion coefficients and outer sphere contribution to the relaxivity. 

The second sphere water molecules should be considered as bound via hydrogen 

bonds to the functional group in the ligand molecule. The second-sphere contribution is 

difficult to evaluate due to the number of second-sphere water molecules and their 

exchange rates, which are unknown. Usually, the second sphere effect is included in the 

outer-sphere contribution. The explanation of unexpected high relaxivity of some Gd 3+ 

complexes, such as [Gd(DOTP)] 7 where no water molecules can be found in the inner- 

sphere, is given by a very strong second-sphere contribution. 35 

It is problematic to separate the inner-, second- and outer-sphere components of 

CA 

T1 . An approach often used to approximate 

GdH 

je 

 

 

 

 

IS 

T1 of a CA is to simply remove 

3 

2 

 

 

 

 

 

 

 

 

CA 

T1 of a 

second sphere component, related CA that has q 0 . In the case of the second CA, 

is zero (since 0 

and 

OS 

T 1 

q ) and the observed relaxation enhancement is attributed solely to 

, therefore providing a reasonable estimate of 

SS 

T1 and 

IS 

T 1 

SS 

T 1 

OS 

T1 for the first CA 

because these parameters are not expected to differ significantly among structurally 

related CAs. Nevertheless, there are noteworthy exceptions to this assumption and this 

approach must be carefully applied. Presently there is no satisfactory experimental 

method to partition the second- and outer sphere components of 

CA 

T1 .


2.2.2.3. PARAMAGNETIC RELAXATION PARAMETERS 

In order to better understand the potential of the different parameters to affect 

37 

CA 

T1 , 

equations 2.6-2.25 are condensed and presented into equation 2.26. Two reasonable 

assumptions are made in simplifying the equations, SC 

7 

1 

 

water proton metal interactions in CA’s and 

1 

T 

CA 

1 

OuterSpher eContribut ion 

 

 

 

1 

kOS 

S( 

S 1) 

C 3 j 

7 j 

 

c2 

2 2 

S 

c2 

T 1 

A 

 

 

1 is omitted because 0 

because 0 

B is sufficiently strong. 

InnerSpher eContribut ion 

SecondSphe reContribu tion 

 

 

 

 

 

qC 

q'C 

 

kIS 

kSS 

1 

1 

 

M 

 

'M 

c 

6 ' 

1 

c1 

S( 

S 1) 

 

 

r' 

S( 

S 1) 

 

2 2 

2 2 

1 I 

 

 

 

c1 

1 I 

' 

 

c1 

 

H 

S 

for 

(2.26) 

The development of more effective contrast agents involves the optimization of 

various parameters governing the relaxivity. The relation between the molecular structure 

and electronic relaxation is still not well established. This task is focused on the 

optimization of three parameters, which are the number of water molecules coordinated to 

the metal ion ( q ), the exchange lifetime ( M ) and the reorientation correlation time ( R ). 

An example of the variation of 

CA 

T1 with respect to these three parameters is given in 

Figure 2.7. Figure 2.7, which represents the inner-sphere proton relaxivity ( 1 

r ) calculated 

at two magnetic fields as a function of M and R for two electronic relaxation time values, 

q = 1 and r = 3.1 Å. Several important conclusions may be drawn from these simulated 

curves: i) the optimal relaxivities are obtained by slowing down the rotation of the complex 

and optimizing the exchange lifetime (for the later, an extreme exchange lifetime has a 

negative influence on the relaxivity); ii) the optimal relaxivity decreases with increasing 

magnetic field strength; iii) when one parameter begins to be optimized the other



parameters become more critical and, therefore a compromise has to be reached; and iv) 

T1 e increases as the magnetic field strength increases and at 0.5 T, e T1 factor, while at 1.5 T it may reach a point where it does not influence 1 

Figure 2.7. Inner-sphere relaxivities calculated as a function of M 

T 1 

values of e 

at 0.5T (~21MHz) and 1.5T (~64MHz).11 

38 

r . 

may be a limiting 

and R 

for 

The relaxation induced by superparamagnetic particles 36 , briefly outlined in the next 

section, is explained by the classical outer-sphere relaxation theory, reformulated by the 

Curie relaxation theory, 24 since the former considers the relaxation rates of water protons 

diffusing near the unpaired electrons responsible for the particle’s magnetization. 37 These 

agents exhibit strong 1 

T relaxation properties and due to susceptibility differences to their 

surroundings, also produce a strongly varying local magnetic field, which enhances T 2 

relaxation. An important result from the outer-sphere theory is that the 1 

r ratio 

increases with increasing particle size and, thus, smaller particles are much better 1 

shortening agents than larger ones. 32,38,39 

2 r 

T -


2.2.3. NUCLEAR MAGNETIC RESONANCE DISPERSION 

The nuclear magnetic relaxation properties of a compound are ideally acquired 

by a magnetic field dependence study. This is performed by measuring the proton 

longitudinal and transverse relaxation over a range of magnetic fields with a Fast-Field- 

Cycling (FFC) spectrometer that switches the magnetic field strength over a range of 

proton Larmor frequencies. The data acquired represents the Nuclear Magnetic 

Resonance Dispersion (NMRD) profile, which can be fitted by eqs 2.10-2.18, eqs 2.21- 

2.22 and eqs 2.23-2.24, in order to obtain the values of the relaxation parameters. The 

underlying complexity is a major drawback to this technique, since there are too many 

influencing parameters inducing possible errors in the fitting of the NMRD profile. For this 

reason, 40 an accurate interpretation of NMRD profiles may only be made by reference to 

independent information from other techniques, such as 17 O NMR, 2 H or 13 C NMR and 

Electron Paramagnetic Resonance (EPR). The 17 O NMR relaxation rates and chemical 

shifts, over a range of magnetic fields and as function of temperature and pressure, allow 

estimates of the number of inner-sphere water molecules ( q ), the rotational correlation 

time ( R 

), the water exchange rate ( M 

), and the longitudinal electronic relaxation rate. 

With 2 H or 13 C NMR it is also possible to determine the rotational correlation time ( R ), 

while the EPR line widths give direct access to transverse electronic relaxation rates. 30 

With these techniques a more reliable determination of the set of parameters 

governing proton relaxivity provide a more stringent test of the relaxation theories applied 

to the techniques and allow a validation of current models for the dynamics in 

paramagnetic solutions. 

Within a NMRD profile, despite possible inaccuracies, it is viable to draw some 

valuable conclusions concerning the relaxation processes: i) at the high magnetic field 

region (10-100 MHz) the inner-sphere relaxation is governed by the reorientational 

correlation ( R 

) time, which is dependent on the molecular weight of the complexes; ii) at 

the low magnetic field, this region is mainly determined by the zero-field electronic 

relaxation time ( s0 

), where 

1 

and is dependent on the symmetry of the 

12 

 

s0 

2 

complex and on the chemical nature of the coordinating groups. 

v 

An example of a typical NMRD profile of low-molecular-weight Gd(III) complexes 

with one inner-sphere water molecule such as [Gd(DTPA-BMA)(H2O)], [Gd(DTPA)(H2O)] 2- 

39



, or Gd(DOTA)(H2O)] - (three clinically approved CAs) have the general forms shown in 

Figure 2.8. 

Figure 2.8. NMRD profile of three commercially-available CAs [Gd(DTPA- 

BMA)(H2O)], [Gd(DTPA)(H2O)] 2- , or Gd(DOTA)(H2O)] - . 14 

The main feature of these profiles is that the relaxivity is limited by fast rotation, 

especially at high frequencies (>10 MHz). As a consequence, the high-field relaxivities of 

these three agents are practically the same, since their sizes and therefore their rotational 

correlation times ( R ) are very similar. The lower value for the water exchange rate (one- 

order of magnitude) determined for [Gd(DTPA-BMA)(H2O)] has no influence on the high- 

field relaxivities, since at these fields the relaxivity is exclusively limited by the rotation of 

the CA. In the low-field region the different relaxivities reflect the extremely slower 

electronic relaxation of the symmetric [Gd(DOTA)(H2O)] - species, as compared to the 

linear chelates. The higher relaxivity is a consequence of its much longer zero-field 

electronic relaxation time ( 0 

s ). The determined s0 

40 

values are 650ps for 

[Gd(DOTA)(H2O)] - 41 and 72 and 81ps for [Gd(DTPA)(H2O)] 2- and [Gd(DTPA-BMA)(H2O)], 

respectively. 41 

In the case of superparamagnetic NPs the relaxation induced by these crystals is 

complicated by another feature. The ferromagnetic crystals are dispersed in a liquid media


to form a colloid suspension. As their size is much smaller than the size of one magnetic 

domain, they are completely magnetized, constituting a nanomagnet made of a fully 

magnetized single domain. In addition to the value of its magnetization, each single mono- 

domain is also characterized by its anisotropy energy. The magnetic energy of each nano- 

magnet depends upon the direction of its magnetization vector with respect to the 

crystallographic directions, increasing with the tilt angle between the magnetization vector 

and the anisotropy directions (or easy axes) which minimize this magnetic energy. The 

difference between the maximum and minimum energy is called the anisotropy energy, Ea, 

which is proportional to the crystal volume (V), Ea = Ka V, where Ka is the anisotropy 

constant. 

In these conditions, the return of the magnetization to equilibrium is determined 

by two different processes. The first one is the Néel relaxation, which is determined by the 

anisotropy energy, and is characterized by a relaxation time constant N, which defines the 

fluctuations that arise from the jumps of the magnetic moment between different easy 

directions causing the return of the magnetization to equilibrium after a perturbation 

(Figure 2.9). The second process is the Brownian relaxation, which characterizes the 

viscous rotation of the entire particle, is characterized by a (Figure 2.9). 36 Therefore, the 

global magnetic relaxation rate of the colloid (1/, where is the global magnetic relaxation 

time) is the sum of the Néel relaxation rate (1/N) and the Brownian relaxation rate (B). 

The Brownian relaxation time is proportional to the crystal volume while the Néel 

relaxation time is an exponential function of the volume. Then, in the case of larger 

particles, B is shorter then N, so the viscous rotation of the particle becomes the 

dominant process determining the global relaxation. In these conditions, the magnetization 

curve is perfectly reversible because the fast magnetic relaxation allows the system to be 

always at thermodynamic equilibrium. This behavior has been named 

“superparamagnetism” by Bean and Livingston. 36 

41



Figure 2.9. Illustration of the two components of the magnetic relaxation of a 

magnetic fluid. 36 

Therefore, the influence of the electron magnetic moment is modulated by the 

Néel relaxation, which depends upon the crystal anisotropy. In the case of large 

superparamagnetic crystals or crystals with a very high anisotropy constant 42 the 

anisotropy energy is larger than the thermal energy, maintaining the direction of the crystal 

magnetic moment very close to that of the anisotropy axes. This characteristic simplifies 

the models, enabling the precession of the electron magnetization. Considering small 

crystals, the anisotropy energy is of the same order of magnitude as the thermal energy, 

therefore it is possible for the magnetic moment to point in a different direction from the 

anisotropy axes, allowing some electron precession. 

In both these cases the explanation of the longitudinal relaxation rate 

dependence with the magnetic field (NMRD profile) is based on the so-called Curie 

relaxation. 27 This relaxation arises from considering separately two contributions to 

relaxation: i) diffusion into the inhomogeneous non-fluctuating magnetic field created by 

the mean crystal moment, aligned onto 0 

B (the accurately termed Curie relaxation) and ii) 

the fluctuations of the electronic magnetic moment or the Néel relaxation. 

The different contributions to proton relaxation, in the simplified model for crystals 

with large anisotropy, are given in Figure 2.10. As shown at low field, the proton 

longitudinal relaxation rate is obtained by introducing into the Freed equations the 

precession prohibition mentioned above (the electron Larmor precession frequency is set 

to zero). 43 Figure 2.10 shows the dispersion of this density spectral function, called Freed 

function, centred around 

1 

I 

. At high field strength the magnetic vector is locked 

 

C 

42


along the 0 

B direction and the Curie relaxation dominates and the corresponding 

relaxation rates are given by Ayant’s model. 44 In the case of intermediate field strength, 

the proton relaxation rates ( 1 

R and 2 

R ) are combinations of the high- and low-field 

strength contributions, weighed by factors depending upon the Langevin function. 45 

Figure 2.10. Different contributions to proton relaxation in the simplified model 

for crystals with large anisotropy. 

In the case of superparamagnetic nanoparticles, the fitting of the NMRD profiles 

by adequate theories provides information on: their average radius ( r ), their specific 

magnetization ( s M ),anisotropy energy ( A 

E ), and Néel relaxation time ( N 

43 

). 46 Figure 2.11 

illustrates a standard NMRD profile of magnetite particles in colloidal suspension. The 

average radius ( r ) may be determined since at high magnetic fields the relaxation rate 

only depends upon D as the inflection point corresponds to the condition I D ~ 1 (see 

Figure 2.11) and given that 

2 

r 

D , the determination of D 

D 

gives the crystal size r . 

Regarding the specific magnetization ( s M ), also at high fields, M s can be obtained from 

the equation: 

M 

1/ 

2 

max ~ 

R 

s 

C 

D 

 

 

, where C is a constant and max 

R is the maximal relaxation 

rate. The absence or presence of dispersion at low fields provides information about the 

magnitude of the anisotropy energy. In the case of crystals characterized by a high



anisotropy energy ( A 

E ) value as compared to the thermal agitation the low field 

dispersion disappears. These conclusions have also been confirmed in previous work with 

cobalt ferrites, 39 which are known to have high anisotropy energy. The relaxation rate at 

very low fields 0 

equal to N 

if D 

R is governed by a “zero magnetic field” correlation time 0 

N . Often this situation is not met; therefore, N 

44 

, which is 

C 

is often reported as 

qualitative information in addition to the crystal size and the specific magnetization. 

Figure 2.11. NMRD profile of magnetite particles in colloidal suspension. 39 

2.3. CLASSIFICATION OF CAs 

According to their various features, the currently available MR CAs may be 

classified in different ways, based on the: i) presence and nature of the metal centre, ii) 

their magnetic properties, iii) effect on the magnetic resonance image (MRI), iv) chemical 

structure and ligands present, and v) biodistribution and applications. A simplification of 

the above classification may be made, since several of these characteristics are closely 

related: on one side the chemical composition (metal and ligands), the magnetic 

properties and the effects on the MRI image and, on the other side, their applications 

resulting from their in vivo bio-distribution.


2.3.1. CHEMICAL COMPOSITION, MAGNETIC PROPERTIES AND EFFECTS 

ON THE MRI IMAGE 

The chemical nature of MRI CAs varies widely. They can be i) small 

mononuclear or polynuclear paramagnetic chelates, ii) metalloporphyins, iii) polymeric or 

macromolecular carriers (covalently or noncovalently labelled with paramagnetic chelates, 

such as dendrimers or proteins), iv) particulate CAs, v) paramagnetic or 

superparamagnetic particles, vi) diamagnetic or paramagnetic chemical exchange 

saturation transfer (PARACEST) polymers or chelates, vii) diamagnetic hyperpolarization 

probes, such as 13 C labelled compounds or ions. 47 

The most common and simplest paramagnetic chelates use Gd 3+ or Mn 2+ (due to 

their properties already described) as metal centres with linear or macrocyclic 

polyaminocarboxylate/phosphonate derivative ligands. In the case of particulate CAs, Gd 3+ 

ions are also used in various forms, for example: bound to amphiphylic chelates in 

paramagnetic micelles, in the bilayer of liposomes, as small chelates in their aqueous 

internal compartment, or bound to porous materials like zeolites. 47 When considering 

paramagnetic or superparamagnetic particles, gadolinium oxide nanoparticles and iron 

oxide particles with different sizes and coatings are considered, respectively. 47 

Other paramagnetic chelates containing different Ln 3+ ions, such as Dy 3+ or Tm 3+ , 

are used as MRI CAs with different variations. Examples of these variations are as 

susceptibility agents, when they are in a compartment, affecting T2 or T2 * relaxation, 48 or 

as PARACEST and LIPOCEST agents, where the paramagnetic shift effect of the ion on 

the proton nuclei of the CA facilitates the irradiation of their shifted resonances and 

consequent saturation transfer by chemical exchange (CEST effect), thus decreasing the 

water proton signal intensity and leading to a negative image contrast. 49 

Less conventional MRI CAs do not contain any metal centre. Examples include 

the oral agents or the CEST agents. 

Recently a new class of MRI CAs has been developed, dynamic nuclear 

polarization (DNP) agents, such as hyperpolarized noble gases ( 3 He, 129 Xe) and 13 C- 

labeled organic compounds or ions like 6 Li + . The probe’s nuclei (with long T1 values) 

exhibit a strong signal enhancement allowing the direct imaging of the probe’s molecular 

distribution. 47 

CAs may also be classified according to their magnetic properties as 

paramagnetic or superparamagnetic agents. Metal ions with one or more unpaired 

45



electrons are paramagnetic and consequently possess a permanent magnetic moment. 

Organic free radicals are also paramagnetic due to their unpaired valence electron. Within 

an aqueous suspension a dipolar magnetic interaction between the electronic magnetic 

moment of the paramagnetic atom and the much smaller magnetic moments of the 

protons of the nearby water molecules is formed. Random fluctuations in this dipolar 

magnetic interaction can be caused by the molecular motions, therefore reducing both the 

longitudinal ( 1 

T ) and the transverse ( 2 

T ) relaxation times of the water protons. As already 

mentioned, Gd 3+ and Mn 2+ are examples of paramagnetic ions used as MR CAs, because 

their physical properties are suitable for efficiently reducing the 1 

relaxation times. 47 

46 

T and 2 

T proton 

Unfortunately, paramagnetic metal ions, like Gd 3+ , can not be used as CAs in 

their ionic form due to their undesirable biodistribution (accumulating in bones, liver or 

spleen) and relatively high toxicity. Therefore, metal ion complexes (chelates) with high 

thermodynamic and kinetic stabilities are required in order to use these paramagnetic 

metal ions in vivo. Small Gd 3+ or Mn 2+ -based paramagnetic chelates are nonspecific CAs 

and have similar 1 

positive contrast 1 

r and 2 

r effects in water. Currently these agents are used mainly for 

T -weighed images. As already mentioned, the relaxivity effects of these 

agents result mostly from both inner- and outer-sphere mechanisms. 

The superparamagnetic agents consist of materials, such as iron oxides, in the 

form of colloids made up of particles (typically 5 – 200nm in diameter) in suspension, 

which are composed of very small crystallites (1 – 10 nm) containing several thousand 

magnetic ions. These agents exhibit a behaviour similar to paramagnetism except that, 

instead of each individual atom being independently influenced by an external magnetic 

field, the magnetic moment of the entire crystallite tends to align with that magnetic field. 47 

Therefore the magnetic moments of the individual ions do not cancel out but are mutually 

aligned, inducing a much higher permanent magnetic moment within the crystallites when 

in the presence of a magnetic field compared to a single molecule of a Gd chelate. 47 

These particles are embedded in a coating such as dextrans (in ferumoxide) or 

siloxanes (in ferumoxsil), which prevents agglomeration. There are three kinds of 

particulate superparamagnetic iron oxides, according to the overall size of the particles: i) 

if they have a diameter dd>50 

nm , they are called small superparamagnetic 

iron oxide (SPIO) particles; and iii) micron-sized particles of iron oxide (MPIO) are large 

particles, with a diameter of several microns. Intravenous administration is only possible


for the former two, while the large particles can only be administered orally to explore the 

gastrointestinal tract, otherwise they would be trapped in the lung alveoli. Yet more 

nomenclatures exist, such as monocrystalline iron oxide particles (MION) and cross-linked 

iron oxides (CLIO). 

developed as 2 

As a consequence of their larger size and magnetic moment SPIOs were initially 

T -agents, producing a dark area on MRI images resulting from their 

negative contrast effect. 50 A new generation of USPIOs with sizes less than 10 nm has 

also been reported to have excellent 1 

T -enhancing properties. 39,51,52,53 As already 

mentioned, the relaxation induced by superparamagnetic particles is explained by the 

classical outer-sphere relaxation theory, reformulated by the Curie relaxation theory. 48 

T 2 or 

The susceptibility agents induce long-range interactions which can dominate the 

* 

T 2 relaxation. This relaxation mechanism, known as susceptibility-induced 

relaxation, 47 is related to the magnetization of the CA, which results from the partial 

alignment of the individual magnetic moments in the direction of the magnetic field. 

An example of this type of agent is when a CA becomes compartmentalized, like 

when superparamagnetic particles are taken up by Kupffer cells, or in a vessel, the 

compartment containing these CA functions as a secondary CA. In this case the water 

protons on the outside of the compartment are affected by the overall magnetization of the 

magnetic bulk material inside that compartment and are, therefore, relaxed by an outer- 

sphere mechanism. 47 The blood oxygen level dependent (BOLD) effect is due to this long- 

range 

T relaxation effect, where the paramagnetic hemoglobin is compartmentalized 

* 

2 T2 

within blood erythrocytes. This phenomenon is the basis of functional MRI 54 and can 

depend on several parameters such as the magnetic moment and local concentration of 

the CA, the dimensions and geometry of the compartment, the diffusion constant of water 

within the compartment and so on. 

As already mentioned, the classification of MR CAs as ‘ 1 

is not always accurate, since any CA that reduces 1 

47 

T also reduces 2 

T agents’ or ‘ 2 

T agents’ 

T . Nevertheless any 

agent that reduces T 2 does not necessarily reduce T 1 , at least at MRI field strengths. 

Therefore the CA functions as a ‘ T 1 agent’ or ‘ 2 

T agent’ depending on the imaging 

sequence used, the magnetic field strength, the size of the CA and how the CA is 

compartmentalized in the tissue. 47



2.3.2. BIODISTRIBUTION AND APPLICATIONS 

The biodistribution of CAs is very important because it is vital to know what 

happens to the CAs and where they go when they are intravenously administered. The 

main distribution sites and excretion pathways for soluble metal complexes are 

summarized in Figure 2.12. 

There are several types of possible classifications and divisions for CAs such as: 

i) non-specific agents, ii) specific or targeted agents, iii) non-injectable organ-specific 

agents, iv) responsive, smart or bio-activated agents and v) CEST and Hyperpolarized 

agents. 

Figure 2.12. Main distribution sites and excretion pathways for intravenously 

administered soluble metal complexes. 47 

2.3.2.1. NON-SPECIFIC AGENTS 

Nonspecific CAs (those that do not interact specifically with any type of cells) 

encompass the extracellular fluid (ECF) agents and the blood pool agents (BPA). The 

48


former are low molecular weight extracellular complexes that equilibrate rapidly between 

the intravascular and interstitial space and are mainly excreted by the kidneys. 55-58 BPA 

have high molecular weight, such as high generation dendrimers, which stay within the 

intravascular space and are slowly excreted via the kidneys and/or the liver. 55,56,57,58 

The ECF agents leak rapidly from the blood into the interstitium with a distribution 

half-life of about 5 minutes and are cleared by the kidney with an elimination half-life of 

about 80 minutes. ECF agents have been extensively used in extra-cranial applications. 

Since these agents rapidly clear out of the blood, the images are typically acquired in the 

early phase following a bolus injection, when used in conjunction with MRA. The most 

common ECF agents are Gd 3+ chelates of linear or macrocyclic polyaminocarboxylate 

ligands and are the main commercially-available MRI CAs. These type of CAs can be 

divided into two groups: neutral and ionic agents (Figure 2.13, Table 2.1). In conclusion, 

Gd 3+ -based ECF agents are typically safe when used in clinically recommended doses, 

and adverse reactions and side effects, such as allergy, are very rare. 47 

Figure 2.13. Structures of commercial ECF contrast agents with intravasculsar 

and extracellular distribution. 36 

49



Table 2.1. Properties of commercial ECF contrast agents with intravascular 

and extracellular distribution. 

Short Name & 

Generic Name 

Gd-DTPA - Gadopentetate 

dimeglumine 

Gd-DOTA – Gadoterate 

meglumine 

Gd-DTPA-BMA - 

Gadodiamine 

Gd-HP-DO3A – Gadoteridol 

Gd-BT-DO3A - Gadobutrol 

Gd-DTPA-BMEA – 

Gadoversetamide 

Gd-BOPTA – Gadobenate 

dimeglumine 

Trade Name 

Magnevist® 

Bayer Schering 

Pharma AG 

Dotarem® 

Guerbet 

Omniscan® GE 

Healthcare 

Prohance® 

Bracco SpA 

Gadovist® 


Pharma AG 

OptiMARK® 

Mallinckrodt 

MultiHance® 

Bracco 

Diagnostics 

Relaxivity (mM -1 s -1 ) 

B =1.0 T (37 ºC) 

0 

r 1 =3.4, 2 

r 1 =3.4, 2 

r 1 =3.9, 2 

r 1 =3.7, 2 

r 1 =3.6, 2 

50 

MRI Enhancement & 

Physiochemical Properties 

r =3.8 positive – charged (ionic) – linear 

r =4.8 

positive – charged (ionic) – 

macrocyclic 

r =4.3 positive – neutral (non-ionic) – linear 

r =4.8 

r =4.1 (at 0.47 T) 

r 1 =3.8, 2 

r 1 =4.6, 2 


macrocyclic 


macrocyclic 

r =4.2 positive – neutral (non-ionic) – linear 

r =6.2 positive – charged (ionic) – linear 

Blood-pool agents (BPA) or intravascular agents are compounds with larger 

‘sizes’ than the previous ECF agents and also have higher 1 

r relaxivities. These two 

characteristics offer many advantages in MR angiography (MRA) when compared to ECF 

agents. 14,59 Because their high molecular weight (>20 kDa) prevents leakage into the 

interstitium, they remain in the intravascular system longer than conventional ECF 

agents. 60 These agents have been primarily developed for MRA and there main properties 

are to have a relatively long vascular half-life and the highest possible 1 

r2 relaxivity must be low enough to avoid excessive signal loss due to 

r -relaxivity. Their 

* 

T 2 T2 

relaxation. 47 

The potential advantage of using a BPA in MRA is the prolonged imaging window 

given that a longer image acquisition time is granted. This translates into a higher image 

resolution and/or signal-to-noise ratio, therefore higher-quality angiograms may potentially 

be attained when compared to gadolinium-based ECF agents. In this respect, vascular 

abnormalities, associated with certain tumours or atherosclerosis, can be more easily 

detected. Additionally, tissue blood volume and perfusion can be measured. 61 The 

disadvantage of this ‘steady-state’ approach is that arteries and veins are equally 

enhanced, thus making their differentiation more challenging.


According to their action mechanism the BPA can be divided into several 

classes: i) the noncovalent binding of low molecular weight Gd 3+ -based complexes to 

human serum albumin –HSA (the most abundant plasma protein) which prevents 

immediate leakage into the intersititial space; ii) systems based on polymers or liposomes 

(increases the size of the CA molecule); and iii) systems based on particles (involves a 

change in the route of elimination). Figure 2.14 and Table 2.2 display several examples of 

these different classes of BPA. 

Figure 2.14. Structures of some HSA-binding and polymeric Gd 3+ complexes, 

as potential or approved blood pool CAs for MRA (simplified structure for 

Gadomer 17). 47 

51



Table 2.2. Properties of commercial BPA contrast agents. 

Short Name & 

Generic Name 

MS-325 – 

Diphenylcyclohexyl 

phosphodiester-Gd-DTPA 

B-22956 or B-22956/1 – 

Gadocoletic acid 

Gadomer-17 or Gd-DTPA-17 

P792 – Gadomelitol 

Trade Name 

Vasovist® 


Pharma AG 

Bracco SpA 1 


Pharma AG 

Vistarem® 

Guerbet 


(37 ºC) 

r 1 =19, 0 

r =27, 0 

r 1 =11.9, 2 

r 1 =42, 2 

Gadofluorine-M _ 1 

AMI-227 – Ferumoxtran-10 

SH U 555 C – Ferucarbotran 

Sinerem® 

Guerbet, 

Combidex® 

AMAG 

Supravist® 


Pharma AG 

52 

B =1.5 T 

B =0.5 T 

r =16.5 0 

r =50 0 

r =137, 0 

r 1 =22.7, 2 

B =0.5 T 

B =0.47 T 

B =1.5 T 

r =53.1 0 

r 1 =14, 0 

B =1.5 T 

B =1.0 T 

2.3.2.2. SPECIFIC OR TARGETED AGENTS 

MRI Enhancement & 


positive – Albumin binding 

molecules – MRA 

vascularisation, capillary 

permeability 

positive – Reversible albumin 

binding – coronary MRA 

positive – Polymeric Gd complex 

– MRA vascularisation and tumor 

differentiation 

positive – Polymeric Gd complex) 

– MRA 

positive – Polymeric Gd complex 

– MRA 

Positive or negative – Coated 

USPIO particles – MRA 

Positive – Coated USPIO 

particles – MRA 

Specific or targeted agents can also be divided into two main groups: those that 

are actively targeted to a molecularly specific site with an appropriate ligand and those 

that are passively directed to a particular type of cell. The first group includes agents that 

target pathologic processes or states, such as inflammation, angiogenesis, apoptosis, 

atherosclerosis and tumour. The cell labelling CAs function through recognition of specific 

molecular markers of those processes at the cell surface (such as cell-specific receptors 

or transport proteins) and accumulate at those molecular sites (usually in the intracellular 

space). 47 Therefore, these cell labelling CAs are essential for MRI molecular imaging. The 

second main group are the organ-specific agents for the liver (hepatobiliary), spleen, 

lymph nodes, bone marrow or brain, based mainly on the agent size and/or chemical 

structure. Nevertheless, all CAs can be considered organ-specific CAs to some extent, as 

they are excreted either by the liver or the kidneys.


Actively Targeted or cell labelling Contrast Agents 

The actively targeted or cell labelling CAs are able to recognize specific 

molecular sites (e.g. cell-specific receptors or transport proteins) at the cellular membrane 

and accumulate at those sites. Researchers involved in the synthesis of MRI contrast 

agents devote special attention to these types of agents. The development of agents 

enables the recognition and imaging of a specific ‘signature’ of a given disease (molecular 

imaging) that simplifies the task of diagnosis and therapy. 62 One of the requirements for 

efficient molecular probes is the development of high affinity ligands and their conjugation 

to contrast agents Figure 2.15A. A major problem is the need to have a local 

concentration of CA of ca. 0.5 mM in order to have 50% enhancement of contrast. It is 

possible to increase the payload of reporter groups delivered at the target site by using 

manyreporters bound to a single carrier, as illustrated in Figure 2.15B. 

A 

B 

Figure 2.15. General structure of a targeted CA for cell labelling a) with a single 

reporter group, b) with a carrier of many reporter groups. 47 

The main applied targeting strategies are cell-surface targeting and receptor 

targeting. In the former, specific epitopes easily available at the cell surface are targeted 

to which the CA stays bound. This strategy has been used together with the pretargeting 

approach in order to facilitate the detection and imaging of tumour cells (as tumour cells 

are known to have abnormally high negative charges on their cell surface). It is clear that 

53



the use of low-molecular-weight targeting CAs that are able to accumulate quickly at 

specific cell surface sites have more advantages with regards to the use of 

macromolecular agents. 47 Another example of cell-surface targeting is the well-known 

non-specific binding of porphyrins to the interstitial space of tumours, e.g. Dy-TPPS 

(TPPS=tetraphenylporphyrin sulfonate). One expanded porphyrin (texaphyrin) complex, 

[Gd(Tex)] 2+ (PCI-120), selectively accumulates in tumours, inducing a prolonged 

enhancement in MRI images and provides the possibility of being used as a radiation 

sensitizer for brain cancer. 63 

The targeting of cell-surface receptors approach can be pursued using labelled 

antibodies or low-molecular-weight targeting complexes. In the first approach, due to the 

slow diffusion of the antibodies, the most accessible targets are those present on the 

endothelial vessels. 47 A typical example is the targeting of the endothelial integrin receptor 

V 

 

3 

, a specific angiogenesis marker whose concentration correlates to the tumour 

grade. An example of an imaging probe containing many reporter groups per carrier is a 

Gd 3+ -containing polymerized liposome. The pretargeting approach was used, where the 

target was bound first to a biotinylated monoclonal antibody against 3 , which is well 

recognized by an avidin moiety present on the liposome surface carrying the Gd 3+ chelate 

reporter groups. 64 The same 3 

V target has also been grafted with lipidic nanoparticles 

containing Gd 3+ chelates. 65 The large molecular size and consequently the slow delivery of 

these systems can be considered the major limitation of this technique. 

A more efficient way to accumulate CAs at the target site is by cell internalization. 

For this process to be completely successful the concentration of the agent inside the cell 

must be higher than at the cell surface. The internalization processes may occur via 

phagocytosis and pinocytosis (or fluid phase endocytosis) mechanisms, which do not 

require a cell receptor, or receptor mediated endocytosis. 47 Gd-DTPA bis-stearylamide 

derivatives are CAs forming insoluble Gd 3+ -containing particles that are biodegraded after 

internalization and become soluble and trapped inside the cell. 66 Gd-HPDO3A is a CA 

used for labelling stem cells via pinocytosis mechanism where the stem cells are 

incubated in a culture medium containing Gd-HPDO3A with a concentration ranging 10– 

50mM. 67 The relaxivity of the CA entrapped in the cell endosomic compartment can be 

seriously limited with the exception of the internalization by electroporation, where they 

are delivered to the cytoplasm. 68 Other cell internalization mechanisms have used 

54 

V


membrane transporters and transmembrane carrier peptides. The latter have proven 

useful for the internalization of a number of substrates like proteins, oligonucleotides and 

plasmid DNA. 

Another interesting development was the synthesis of a bimodal (optical and MR) 

imaging probe consisting of a Gd 3+ /Eu 3+ -DOTA complex, a PNA (peptide nucleic acid) 

sequence and a transmembrane carrier peptide. 47 This system can enter any type of cell, 

however it accumulates only in tumour cells due to the specific binding of the PNA moiety 

to the c-myc mRNA whose production is increased in those cells. 69 

Passively directed - organ-specific agents 

The passively directed or organ specific CAs are passively directed to a 

particular type of cell. Normally, there are organ-specific agents for the liver 

(hepatobiliary), spleen, lymph nodes, bone marrow or brain, and they are generally based 

mainly on the agent size and/or chemical structure. 47 

In general, tissue or organ-specific contrast agents consist of two components: a 

magnetic label capable of altering the signal intensity on MR images and a target-group 

molecule having a characteristic affinity for a specific type of cell or receptor. Some 

suitable residues have been incorporated into either the acetic side arms or the 

diethylenetriamine backbone of Gd-DTPA and Gd-DOTA to obtain the tissue or organ- 

specific contrast agents. This type of CAs provides specific advantages in terms of 

sensitivity of lesion detection and characterization. Examples of these types of agents are 

the hepatobiliary CAs, 70-74 the lymph nodes and bone marrow CAs, 75 the brain CAs 76-79 

and the gastro-intestinal CAs 80,81 that can accumulate in the target sites, increasing 

contrast concentration and can produce stronger signal in the MR images. Table 2.3 

provides a summary of the properties of these CAs. 

70,71,72,73,74 7576,77,78,7980,81 

55



Short Name & 

Generic Name 

Gd-EOB-DTPA – 

Gadoxetic acid 

Gd-BOPTA – 

Gadobenate dimeglumine 

Mn-DPDP – 

mangafodipir 

trisodium 

AMI-25 – 

Ferumoxides (SPIO) 

SH U 555 A – 

Ferucarbotran 

(SPIO) 

Table 2.3. Properties of commercial organ-specific contrast agents. 

Trade Name 

Primovist 

(formerly 

Eovist®) Bayer 

Schering Pharma 

AG 

Multihance® 

Bracco SpA 

Teslascan® GE 

Healthvare, 

Endorem 


Resovist® 

/Cliavist® Bayer 


AG 


(37 ºC) 

r 1 =5.3, 2 

r 1 =4.6, 2 

r 1 =2.3, 2 

r 1 =40.0, 2 

r 1 =25.4, 2 

r =6.1 0 

r =6.2 0 

r =4.0 0 

r =160 0 

r =151 0 

Gadofluorine-M _ r 1 =137, 0 

AMI-227 – 

Ferumoxtran-10 

EPI-2104 R – 

Gd-DTPA 

mesoporphyrin 

(gadophrin) 

Sinerem® 


Combidex® 

AMAG 

EPIX 

Pharmaceuticals, 

Inc 

Gadophrin Bayer 


AG 

r 1 =22.7, 2 

56 

B =0.47 T 

B =1.0 T 

B =1.0 T 

B =0.47 T 

B =0.47 T 

B =1.5 T 

r =53.1 0 

- 

- 

B =1.0 T 

MR Enhancement & 


positive – small charged linear complex 

(ionic) – Hepatobiliary liver lesions 

positive – small charged linear complex 

(ionic) – 

Intravascular/Extracellular/Hepatobiliary 

Neuro/whole body, liver lesions 

positive – small charged – 

Pancreatic/Adrenal/Hepatobiliary 

liver lesions 

negative – dextran-coated SPIO 

particles – RES-directed liver lesions 

negative – dextran-coated SPIO 

particles – RES-directed liver lesions 

lymph nodes 

positive – Polymeric Gd complex – 

lymph nodes 

Positive or negative – Coated USPIO 

particles – lymph nodes 

Positive – gadolinium-based small 

peptide – lymph nodes - visualization of 

blood clots 

Positive - myocardium and necrosis 

targeted 

2.3.2.3. NON INJECTABLE ORGAN-SPECIFIC AGENTS 

When ingested, non-injectable organ-specific agents (oral agents) change the 

signal intensity at the stomach and the intestine relatively to adjacent abdominal tissues. 

Depending on their magnetic properties they may change the contrast of the gastro- 

intestinal (GI) tract through various mechanisms, thus allowing MR cholangiography 

(Table 2.4). 

Diamagnetic agents, such as fatty emulsions, fill the GI tract with materials with 

short T 1, 

enhancing its signal and producing positive contrast relatively to adjacent 

tissues. Within the same level, diamagnetic agents can also generate negative contrast 

either by decreasing the T 2 of GI water protons, such as in the presence of Ba 2+ , Al 3+ , 

Si 4+ -containing suspensions, or by decreasing the proton density through the use of


perfluorinated compounds like perfluorooctylbromide (PFOB). 47 Paramagnetic agents, 

such as MnCl2 solutions, Magnevist® [Gd-(DTPA)] or Gadolite® Gd 3+ -containing zeolite Y 

particle suspensions, 80 ferric ammonium citrate (FAC) solutions, are oral positive CAs that 

decrease the 1 

T of GI water protons. Large superparamagnetic particles for oral uptake, 81 

such as Abdoscan®, composed of monodisperse polymer particles of 3 mm diameter 

coated with crystals of iron oxide, or Lumirem®, a silicone-coated superparamagnetic iron 

oxide suspension, belong to the group of negative oral CAs that decrease proton 

57 

T . 

* 

2 T2 

Their main purpose is to distinguish the loops of the bowel from other abdominal 

structures. When ingested, they flow through and darken the stomach and the small 

intestine in 30–45 minutes with a clear identification of the intestinal loops, and improving 

the visualization of adjacent abdominal tissues such as the pancreas. 47 

Table 2.4. Properties of commercial non-injectable organ-specific contrast 

agents. 

Short Name & 

Generic Name 

Gd-DTPA – 

gadopentetate 

dimeglumine 

Ferric amonium 

citrate Geritol 

MnCl2 – manganese 

chloride 

Gd-zeolyte Y 

particles – Gadolite 

60 gastrointestinal 

AMI-121 - 

ferumoxsil (USAN) 

OMP – Ferristene 

(USAN) oral 

magnetic particles 

PFOB - perfluoro- 

octylbromide 

barium sulfate 

suspensions, clays 

mineral particles 

Trade Name Relaxivity 

Magnevist enteral® 

Bayer Schering Pharma 

AG 

Ferriseltz ® Otsuka 

Pharmaceutical 

LumenHance® ImaRx 

Pharmaceutical Corp. 

Bracco Spa 

Gadolite®Pharmacyclics 1 

Lumirem® / Guerbet 

Gastromark ® AMAG 

Pharmaceuticals 

Abdoscan® GE 

Healthcare 

Imagent-GI® Alliance 

Pharmaceutical 

MR Enhancement & 


T1 reduction Positive - Paramagnetic 

- positive – Paramagnetic 

T1 reduction Positive - Paramagnetic 

T reduction Positive - Paramagnetic 

* 

T2 enhanced 

* 

T2 enhanced 

Proton density reduction, 

signal void 

Various Mixtures Diamagnetic, - 1 

Negative – Superparamagnetic 

Negative – Superparamagnetic 

Negative – Diamagnetic 

T -short Negative – Diamagnetic



2.3.2.4. RESPONSIVE, SMART OR BIOACTIVATED AGENTS 

The term ‘responsive’ refers to paramagnetic systems that are sensitive to a 

given biochemical or physiologic parameter that characterizes their microenvironment. 

Typical parameters to which these systems should be responsive are the pH, 

temperature, oxygen pressure, enzymatic activity, redox potential and concentration of a 

specific ion. The use of these agents induces a type of imaging, Molecular Imaging, which 

aims to noninvasively visualize the expression and function of bioactive molecules that 

often represent specific molecular signatures in disease processes. 48 So far, only very 

few of these CA have progressed to in vivo testing. The in vivo non-invasive detection of 

abnormalities in pH or temperature, in the oxygen pressure, in enzymatic activities or in 

the concentration of metal ions and radicals may serve in the future as an important 

diagnostic tool of the underlying diseases. Information on the redox status, an important 

factor governing tumour aggressiveness, can also help determine the adapted tumour 

treatment. Many of these abnormalities are observable in the extracellular media, which 

largely facilitates the chemical design of the imaging probes that have no need for 

intracellular CA delivery. 

pH-sensitive agents 

The potential use of pH-sensitive probes is vast, though so far pH mapping of 

tissues has been primarily intended to facilitate cancer detection and assess the tumour 

status. The pH on the surface of tumours is ~0.4 units lower than that of normal tissue, but 

in some cases it can be as low as 6.0. 82 The pH-sensitive probes may also indicate 

neuronal activity because they induce a slight acidification of the extracellular medium (pH 

7.2–7.4). 83 One could also imagine the use of pH-responsive probes to determine if the 

brain environment is suitable for a drug that functions in a pH-sensitive environment. 

The main requirements for a system to be pH-sensitive are that either the 

dynamics or structural properties determining its relaxivity are pH-dependent. The pH 

dependence of the relaxivity can reflect changes in the hydration number of the metal 

chelates and the presence of protonatable groups on the ligands can influence these 

36,84 ,85,86,87,88,89 

changes. 

58


Designing of pH-sensing probes has become an intensive field in MR imaging 

contrast agent research. 49,84-89 Some of the pH probes are useful for in vivo application. An 

example is the combination of the pH-sensitive amido-phosphonate derivative of Gd- 

DOTA and a pH-insensitive analogue, which were used in a dual injection to image renal 

pH in mice. 90,91 The assumption was made that both compounds have comparable 

pharmacokinetics; hence the concentration of the former can be inferred from the 

concentration of the latter. This mixture has been also used to obtain extracellular pH MR 

imaging maps in a rat glioma model, with improved spatial resolution compared with 

spectroscopic methods. 92 Differences in the order of 1 pH unit could be detected; the 

absolute pH values have been calculated by using a calibration method. 

The key factor inducing the effectiveness of an agent is the difference between 

the relaxivity of the "on" state compared with that of the "off" state. Recently the amplitude 

of the relaxivity response to pH variation of this low-molecular-weight probe has been 

largely improved (doubled in some cases) by conjugating it to a macromolecular 

dendrimeric scaffold. 93 Improving the relaxivity response to pH by increasing the molecular 

weight may also negatively impact the effectiveness of such agents. As already 

mentioned, large molecules, such as dendrimers, remain in the vasculature longer than 

discrete agents, which are better able to diffuse into all extracellular space. Large 

molecules also tend to slowly clear from the body, resulting in increased liver uptake, 

extending their retention time in the body. Further studies into the in vivo behaviour of 

dendrimer-based MR imaging contrast media will be required to establish whether this 

approach, which is successful for increasing the relaxivity response, will yield agents that 

can actually be applied in vivo. 

Temperature-sensitive agents 

Most Ln 3+ chelates have temperature-dependent NMR properties, such as their 

1 H NMR chemical shifts. These shifts can be monitored and used to determine 

temperature variation. For this reason some of these Ln 3+ chelates are considered to be 

good temperature probes. 94,95 

The encapsulation of Gd 3+ chelates into liposomes is also another example of a 

temperature-dependent probe. 96 The membrane transition from gel to liquid crystal occurs 

59



at a specific temperature. At this precise temperature changes in the permeability of the 

membrane occurs, therefore the mobility of the water molecules through the membrane 

changes and consequently the relaxivity also changes. 47 

Enzymatic activity agents 

Among all responsive agents, enzyme targeting represents a specific 

advantage, which is particularly valid for MR imaging detection given its low sensitivity. A 

small concentration of the enzyme can catalytically convert a relatively high amount of the 

enzyme-responsive magnetic probe, which increases the detection for the enzyme 

compared with other biomolecules. In addition, the remarkable specificity of enzymatic 

reactions, which can be observed as changes in the MR imaging properties, can be 

undoubtedly attributed to the targeted enzyme. The presence of certain enzymatic 

reactions indicates the cellular state and can provide the signature of a given pathology. 

Consequently the real-time non-invasive in vivo detection of specific enzymatic activities 

would have invaluable diagnostic impact. 

As the enzymatic activity has been established in the processes of tumour 

formation, growth and metastasis, it is vital to monitor it. Molecular Biology studies have 

defined enzymatic steps of the apoptotic response to anticancer therapies in vitro and in 

vivo. Utilising this approach the detection of gene markers (such as β-galactosidase) could 

be another important field of application. 47 

The first enzymatically responsive potential MR imaging contrast agent was a 

Gd-DOTA–derivative bearing a galactopyranose residue that avoids water 

coordination. 97,98 This sugar moiety is a substrate for the enzyme β-galactosidase. Its 

enzymatic cleavage by β-galactosidase opens the access of water to the first coordination 

sphere of Gd 3+ , resulting in an enhancement of the relaxivity, thus irreversibly activating 

the agent. This agent has been successfully used in vivo to detect by MRI β - 

galactosidase mRNA expression in living Xenopus laevis embryos. 47 

Among other examples, Nivorozhkin et al 99 reported the enzymatic 

transformation of a prodrug Gd 3+ complex with poor affinity to HSA and low relaxivity to a 

60


species with improved HSA affinity and enhanced relaxivity. 36 Mazooz et al 100 described a 

Gd-DTPA peptide acting as a transglutaminase substrate, which was used to monitor 

transglutaminase activity. 47 Anelli et al 101 functionalized Gd(DTPA) 2– with sulfonamide, 

which is known as a specific carbonic anhydrase inhibitor. The agent reacts with carbonic 

anhydrase and thus targets enzyme-rich tissues. 47 Shiftan et al 102 reported MR imaging 

visualization of hyaluronidase in ovarian carcinoma, related to the aggressiveness of 

ovarian cancer metastasis. 47 Chen et al 103 visualized plaque rupture in atherosclerosis 

with a Gd-DOTA–serotonin derivative, which polymerizes in the presence of neutrophil 

myeloperoxidase, resulting in a remarkable relaxivity increase. Another concept for 

enzyme detection is based on the self-immolative mechanism. 47 Duimstra et al 104 reported 

a Gd 3+ complex with a self-immolative moiety, designed for the detection of β- 

glucuronidase. 47 

Redox potential sensitive agents 

Hypoxia is an important factor governing tumour aggressiveness, as hypoxic 

tissue is more resistant to conventional therapeutics. The methodology of imaging tumour 

redox status would allow the noninvasive application of this potential biomarker of tumour 

sensitivity to existing and novel chemotherapies, as well as radiation therapy. The 

possibility of using such methods could also extend to other pathologies, such as 

cardiovascular disease, since free radical formation is associated with damaging effects 

on the coronary microcirculation during recovery from myocardial infarction. 105 Hypoxia is 

mostly detected by imaging techniques including positron-emission tomography (PET) and 

blood oxygen level-dependent (BOLD) MR imaging. 

Reports on redox-sensitive MR imaging contrast agents have been rather 

scarce. The simplest design of a these agents is based on metal complexes whose metal 

ion can be reduced or oxidized depending on the biologic environment and these two 

oxidation forms have different relaxation properties. As a result, the two redox states 

influence the proton relaxation of the surrounding protons to a different extent, resulting in 

different image intensities. 

The partial oxygen pressure (pO2) is also an important parameter in the 

metabolic processes of the cells and its variation is related with certain pathologies. 47 The 

usual systems used as pO2 probes are based on the redox equilibrium of paramagnetic 

ions. In these systems the relaxivity depends on the oxidation state of the metal ion and 

61



consequently on the oxygen pressure. It has been reported that the adducts formed 

between tpps complexes of Mn (III) and Mn (II) and poly-β-cyclodextrin have considerably 

different relaxivities depending on the redox state of the metal, itself determined by the 

partial oxygen pressure of the solution. This technique can quantify the oxygen 

concentration in the surrounding environment. 106 

More recently, DOTA-based complexes of Gd bearing a thiol moiety were 

synthesized, and they form reversible covalent linkages with HSA, which contains a 

reactive thiol at cysteine-34. This redox-sensitive reversible binding of Gd complexes to 

plasma albumin was exploited for imaging the tissue-redox state. 107 

Metal ion and radical agents 

It is known that the presence of metal ions can induce changes in the structure 

of the paramagnetic complexes, consequently changing their relaxivities. 47 Considering 

the Ca 2+ -sensitive MR imaging probes two different approaches have been attempted. 

The first approach uses Gd-complexes with a 1 

while the second approach uses a 2 

62 

T response on interaction with Ca 2+ ions, 

T agent based on the Ca 2+ -related aggregation of 

superparamagnetic iron nanoparticles and calmodulin. 108 Although these strategies have 

several limitations, the main drawback of the second approach is the relatively long time 

course of the Ca 2+ -dependent aggregation (a few seconds) that prevents the sensing of 

fast Ca 2+ -concentration changes. 

The design of all Gd 3+ -based Ca 2+ ion–sensitive probes reported so far 

involved changes in the coordination sphere of the Gd 3+ ion following coordination of Ca 2+ . 

These probes integrate 2 coordinating units that selectively chelate Gd 3+ and Ca 2+ ions. In 

the absence of the sensed Ca 2+ ion, one or more of the donor groups of the Ca 2+ chelating 

centre, are weakly coordinated to the Gd 3+ ion. Concerning the interaction with Ca 2+ , this 

donor group switches from Gd 3+ to Ca 2+ coordination, consequently liberating one 

coordination position on the Gd 3+ ion. This free coordination position is immediately 

occupied by a water molecule increasing the hydration number and the relaxivity. The 

work of Li et al. 109 is an example of this type of agents utilizing this approach. 

Zn 2+ is the second most abundant transition metal ion in the body and its 

highest concentrations occur in the brain. Hanaoka et al. 110 used a ligand of DTPA with 

N,N,N’,N’-tetrakis (2-pyridylmethyl)ethylenediamine (TREN) as a zinc specific chelator. In 

the absence of zinc, water is bound to the gadolinium ion. In the presence of zinc, the


carboxylic acid and pyridine moieties coordinate to zinc thus restricting the access of 

water to the Gd 3+ , thereby decreasing the hydration number and respective relaxivity in 

the presence of zinc. 47 

Iron is the most abundant transition metal in the body and in the brain. Biologic 

iron is most commonly found in the +2 (ferrous) and +3 (ferric) oxidation states. Aime et 

al. 111 synthesized an iron-sensitive contrast agent by functionalizing DTPA with salicylate 

moieties. In the presence of Fe 3+ the Gd-DTPA–salicylate complexes bind to the iron ions 

via the salicylate functional groups. The relaxivity increases as this binding yields an 

increase in R 

. Recently another high-molecular weight tetrametallic supramolecular 

complex [(Ln-DTPA- phen)3Fe] - (Ln=Gd, Eu, La) has been obtained upon self-assembly 

around one of the three iron(II) ions 1,10-phenantroline- based molecules substituted in 5’- 

position with the polyaminocarboxylate diethylenetriamine- N,N,N’,N’,N’-pentaacetate, 

DTPA- phen 4- . 112 

There are reports of radical responsive CAs, consisting of Gd 3+ chelates 

containing a free thiol group (Gd-HASH-DO3A) conjugated through a disulfide bond 

formed with SH-activated phospholipid molecules incorporated in a liposome. The long 

reorientational motion of the supramolecular adduct ensures a 1 

63 

r relaxivity much larger 

than that of the free complex. 47 The disulfide bonds represent a radical sensitive moiety 

and a large decrease in the relaxivity is observed upon their cleavage. 113 

2.3.2.5. CONTRAST AGENTS BASED ON OTHER PROPERTIES 

New classes of MRI CAs have recently been developed, which do not fit into the 

classification mentioned above. These CAs are based on their NMR properties and can be 

grouped into two families, chemical exchange saturation transfer agents (CEST) 114 and 

hyperpolarized agents. 

Chemical exchange saturation transfer agents (CEST) 

Basically a CEST agent is a molecule possessing exchangeable protons (-NH, - 

OH, etc.) that resonate at a chemical shift different from that of the bulk water signal,



measurable when their exchange with the bulk water protons is slow on the NMR 

timescale. This occurs when the difference in frequency between those chemical 

environments ( 

) is higher than the exchange rate of the process ( ex 

64 

k ). When this 

condition is fulfilled the resonance of the CEST mobile protons may be selectively 

saturated using a specific radio frequency 1 

B . These protons will then transfer into the 

bulk water pool and lead to a reduction of its equilibrium magnetization, resulting in a 

decrease of its signal intensity. Consequently this water saturation process is caused by 

chemical exchange (see Figure 2.16). CEST agents can be used to switch the image 

contrast ‘on’ and ‘off’ just by changing the irradiation parameters. 47 

a) a) 

b) b) 

kk kk ex ex ex ex 

CEST CEST CEST CEST 

Agent Agent Agent Agent 

water water water water molecule molecule molecule molecule 

mobile mobile mobile mobile proton proton proton proton of of of of CEST CEST CEST CEST agent agent agent agent 

Figure 2.16. a) CEST agent illustration b) Schematics of the chemical 

exchange saturation transfer (CEST) method using molecular separation of 

encoding and detection for significant signal amplification. The resonance of 

the detection molecule at high concentration (in this case, water) is observed 

after off-resonance saturation (left spectrum) and after on-resonance 

saturation (right spectrum) of a highly diluted CEST agent.


Several agents contain exchangeable protons and consequently can generate 

CEST contrast. 115 Their signals are often very close to the bulk water signal, and broader 

than 2 ppm, due to the magnetic field inhomogeneity of many tissues. 116 Within these 

conditions it is difficult to distinguish contrast due to the CEST effect and the direct 

saturation of bulk water. Large 

values improve the specificity of the CEST and, as 

increases with the magnetic field strength, the overall relationship between 

k ex will be a function of the field strength of the MR experiment. 47 

Since the attainable saturation transfer (ST) value is directly related to ex 

expected that paramagnetic complexes will display large 

65 

and 

k , it is 

values for the exchanging 

proton resonance and thus may improve the efficacy of the CEST agents. These agents 

are called PARACEST agents and consist of particular Ln 3+ complexes with a coordinated 

water molecule undergoing extremely slow exchange with the bulk water and with very 

large 

values. A good ST effect was reported by Zhang et al. 117 by irradiating the 

metal-bound water protons of Eu 3+ chelates resonating at 50 ppm downfield from the bulk 

water. The very same effect can be obtained with slow exchanging amide protons of Ln 3+ 

complexes of DOTA derivatives. 118 Recent reports show that paramagnetic Ln 3+ 

complexes of tetraamide derivatives of DOTA have ST properties which are markedly 

dependent on pH and lactate concentration, making them responsive CAs. 119,120,121 

The main advantage of CEST agents relatively to the traditional MRI CAs is that 

the generation of contrast only occurs when the RF irradiation frequency is set to the 

same frequency as the absorption frequency of the mobile protons. 49 For this reason it is 

not necessary to register an image before the administration of the CEST agent, as the 

image visualization of CEST agents results from the comparison of the on and off 

resonance MRI scan. Co-administration of different CEST agents is also possible as the 

difference within the resonance frequencies of their mobile protons is large enough to 

avoid the overlapping of the respective CEST resonances. The detection of their 

biodistribution is possible and observable within the same image. 122 

The most critical disadvantage of CEST agents is their low sensitivity. 

Theoretically, the ST process is dominated by several parameters, among which ex 

k and 

the number of mobile protons available are particularly relevant. 47 In the case of small 

sized CEST agents, containing less than 10 mobile protons per molecule, such as amino 

acids, heterocyclic compounds, sugars or paramagnetic chelates, they have a detection 

limit in the range of mM. 114,119,123,124 Several approaches to solve this problem were 

114,125-127 ,126,127,128 

suggested.



Hyperpolarized agents and molecular imaging using MRI. 

Undoubtedly MRI provides incomparable soft tissue contrast, however its low 

sensitivity has limited the clinical use to the imaging of water protons. A radically new 

magnetic resonance imaging (MRI) technique, using hyperpolarized 3 He, 129 Xe, 13 C, 15 N 

and 6 Li is being developed to produce high-contrast images of important body tissues that 

have resisted conventional MRI techniques. A totally different approach for increasing the 

polarization of spins is to create an artificial, non-equilibrium distribution of nuclear spins 

called the hyperpolarized state 49 . A hyperpolarized state is defined as a state in which the 

nuclear spin populations are altered with respect to the equilibrium value described by the 

Boltzmann equation. Since the signal intensity is proportional to the spin population’s 

difference, hyperpolarization leads to an increase in the NMR signal intensity by a factor 

as high as 10 5 . 48 Hyperpolarization can improve the detection of 3 He and 129 Xe by up to a 

hundred thousand times. This technology is showing dramatic results for diagnostic 

imaging of the lungs, brain, and other parts of the body. 49 A wide range of organic 

substances containing 13 C has been hyperpolarized by either parahydrogen-induced 

polarization (PHIP) 129 or by dynamic nuclear polarization (DNP). 130 The potential 

applications of hyperpolarized 13 C imaging include vascular imaging, perfusion imaging 131 , 

catheter tracking 132 and visualization and metabolic/molecular imaging. 133 

66



1 Damadian, R., Tumor Detection by Nuclear Magnetic Resonance. Science 1971, 171 

(3976), 1151-1153. 

2 Lauterbur, P. C., Image Formation by Induced Local Interactions: Examples Employing 

Nuclear Magnetic Resonance. Nature 1973, 242 (5394), 190-191. 

3 http://users.fmrib.ox.ac.uk/~stuart/thesis/chapter_2/image270.gif 

4 http://www.cardiff.ac.uk/biosi/researchsites/emric/basics.html 

5 Berry, C. C.; Curtis, A. S. G., Functionalisation of magnetic nanoparticles for applications in 

biomedicine. Journal of Physics D-Applied Physics 2003, 36 (13), R198-R206. 

6 Mansson S, Bjørnerud A., Physical principles of medical imaging by nuclear magnetic 

resonance. In The Chemistry of Contrast Agents in Medical Magnetic Resonance Imaging, 

Merbach AE, Toth E (eds). Wiley: Chichester, 2001; 1–44. 

7 Alexander, H., An introduction to the basics of Magnetic Resonance. Siemens AG Medical 

Solutions Magnetic Resonance 2003, 1-238. 

8 http://upload.wikimedia.org/wikipedia/commons/thumb/8/85/Bluthirnschranke_nach_ 

Infarkt_nativ_und_KM.png/300px Bluthirnschranke_nach_Infarkt_nativ_und_KM.png 

9 Bloch, F.; Hansen, W. W.; Packard, M., The Nuclear Induction Experiment Physical Review 

1946, 70 (7-8), 474-485. 

10 Lauterbur, P. C., Mendoca-Dias, M. H., and Rudin, A. M. in Frotiers of Biological Energetics, 

Dutton, P. L., Leigh, L. S., and Scarpa, A (Eds), Academic Press, New York, 1978, p.752. 

11 Lauffer, R. B., Paramagnetic Metal-complexes as Water Proton Relaxation Agents for NMR 

Imaging - Theory and Design. Chemical Reviews 1987, 87 (5), 901-927. 

12 Aime, S.; Botta, M.; Fasano, M.; Terreno, E., Lanthanide(III) chelates for NMR biomedical 

applications. Chemical Society Reviews 1998, 27 (1), 19-29. 

13 Aime, S.; Botta, M.; Terreno, E., Gd(III)-based contrast agents for MRI. Advances in 

Inorganic Chemistry - Including Bioinorganic Studies, Vol 57 2005, 57, 173-237. 

14 Caravan, P.; Ellison, J. J.; McMurry, T. J.; Lauffer, R. B., Gadolinium(III) chelates as MRI 

contrast agents: Structure, dynamics, and applications. Chemical Reviews 1999, 99 (9), 

2293-2352. 

15 Aime, S.; Botta, M.; Fasano, M.; Terreno, E., Prototropic and water-exchange processes in 

aqueous solutions of Gd(III) chelates. Accounts of Chemical Research 1999, 32 (11), 941- 

949. 

16 Toth, E.; Burai, L.; Merbach, A. E., Similarities and differences between the isoelectronic 

Gd-III and Eu-II complexes with regard to MRI contrast agent applications. Coordination 

Chemistry Reviews 2001, 216, 363-382. 

17 Jacques, V.; Desreux, J. F., New classes of MRI contrast agents. In Contrast Agents I: 

Magnetic Resonance Imaging, Krause, W., Ed. 2002; Vol. 221, pp 123-164. 

67



18 Mansson, S.; Bjørnerud, A., Physical principles of medical imaging by nuclear magnetic 

resonance. In The Chemistry of Contrast Agents in Medical Magnetic Resonance Imaging, 

Merbach AE, Toth E (eds). Wiley: Chichester 2001. 

19 Bloembergen, N.; Purcell, E. M.; Pound, R. V., Relaxation Effects in Nuclear Magnetic 

Resonance Absorption. Physical Review 1948, 73 (7), 679-712. 

20 Solomon, I., Relaxation Processes in a System of 2 Spins. Physical Review 1955, 99 (2), 

559-565. 

21 Bloembergen, N.; Morgan, L. O., Proton Relaxation Times in Paramagnetic Solutions. 

Effects of Electron Spin Relaxation. Journal of Chemical Physics 1961, 34 (3), 842-850. 

22 Bloembergen, N., Proton Relaxation Times in Paramagnetic Solutions. Journal of Chemical 

Physics 1957, 27 (2), 595-596. 

23 Connick, R. E. and Fiat, D. Oxygen−17 Nuclear Magnetic Resonance Study of the 

Hydration Shell of Nickelous Ion, J. Chem. Phys., 1966, 44, 4103 – 4107. 

24 Toth, E.; Helm, L.; Merbach, A. E. Contrast Agents I Magnetic Resonance Imaging: New 

Classes of MRI Contrast Agents: Relaxivity of MRI Contrast Agents, Top.Curr.Chem., 2002, 

221, 61 - 101. 

25 Botta, M., Second coordination sphere water molecules and relaxivity of gadolinium(III) 

complexes: Implications for MRI contrast agents. European Journal of Inorganic Chemistry 

2000, (3), 399-407. 

26 Peters, J. A.; Huskens, J.; Raber, D. J., Lanthanide induced shifts and relaxation rate 

enhancements. Progress in Nuclear Magnetic Resonance Spectroscopy 1996, 28, 283-350. 

27 Gueron, M., Nuclear relaxation in macromolecules by paramagnetic ions: a novel 

mechanism. Journal of Magnetic Resonance 1975, 19 (1), 58-66. 

28 McLachlan, A. D., Line Widths of Electron Resonance Spectra in Solution. Proceedings of 

the Royal Society of London Series a-Mathematical and Physical Sciences 1964, 280 

(1380), 271-288. 

29 Bertini, I.; Luchinat, C.; Aime, S., NMR of paramagnetic substances. Coordination 

Chemistry Reviews 1996, 150, 1-28. 

30 Powell, D. H.; Merbach, A. E.; Gonzalez, G.; Brücker, E.; Micskei, K.; Ottaviani, M. F.; 

Köhler, K.; Zelewsky, A. V.; Grinberg, O. Y.; Lebedev, Y. S., Magnetic-Field-Dependent 

Electronic Relaxation of Gd 3+ in Aqueous Solutions of the Complexes [Gd(H2O)8] 3+ , 

[Gd(propane-1,3-diamine-N,N,N′,N′-tetraacetate)(H2O)2] − , and [Gd(N,N′-bis[(N- 

methylcarbamoyl)methyl]-3-azapentane-1,5-diamine-3,N,N′-triacetate)(H2O)] of interest in 

magnetic-resonance imaging, Helv.Chim.Acta, 1993, 76, 2129 - 2146. 

31 Rast, S.; Borel, A.; Helm, L.; Belorizky, E.; Fries, P. H.; Merbach, A. E., EPR spectroscopy 

of MRI-related Gd(III) complexes: Simultaneous analysis of multiple frequency and 

temperature spectra, including static and transient crystal field effects. Journal of the 

American Chemical Society 2001, 123 (11), 2637-2644. 

68


32 Hwang, L. P.; Freed, J. H., Generalized Einstein relations for rotational and translational 

diffusion of molecules including spin. Journal of Chemical Physics 1975, 63 (1), 118-130. 

33 Freed, J. H., Dynamic effects of pair correlation functions on spin relaxation by translational 

diffusion in liquids. II. Finite jumps and independent T1 processes. Journal of Chemical 

Physics 1978, 68 (9), 4034-4037. 

34 Koenig, S. H.; Brown, R. D., Field-cycling relaxometry of protein solutions and tissue: 

Implications for MRI. Progress in Nuclear Magnetic Resonance Spectroscopy 1990, 22, 

487-567. 

35 Avecilla, F.; Peters, J. A.; Geraldes, C., X-ray crystal structure of a sodium salt of 

Gd(DOTP) (5-): Implications for its second-sphere relaxivity and the Na-23 NMR hyperfine 

shift effects of Tm(DOTP) (5-). European Journal of Inorganic Chemistry 2003, (23), 4179- 

4186. 

36 Laurent, S.; Forge, D.; Port, M.; Roch, A.; Robic, C.; Elst, L. V.; Muller, R. N., Magnetic iron 

oxide nanoparticles: Synthesis, stabilization, vectorization, physicochemical 

characterizations, and biological applications. Chemical Reviews 2008, 108 (6), 2064-2110. 

37 Gillis, P.; Koenig, S. H., Transverse relaxation of solvent protons induced by magnetized 

spheres: application to ferritin, erythrocytes and magnetite. Magnetic Resonance in 

Medicine 1987, 5 (4), 323-345. 

38 Koenig, S. H.; Kellar, K. E., Theory of 1/T1 and 1/T2 NMRD profiles of solutions of magnetic 

nanoparticles. Magnetic Resonance in Medicine 1995, 34 (2), 227-233. 

39 Roch, A.; Muller, R. N.; Gillis, P., Theory of proton relaxation induced by superparamagnetic 

particles. Journal of Chemical Physics 1999, 110 (11), 5403-5411. 

40 Koenig, S. H., Novel Derivation of Solomon-Blombergen-Morgan Equations - Application to 

Solvent Relaxation by Mn2+ Protein Complexes. Journal of Magnetic Resonance 1978, 31 

(1), 1-10. 

41 Geraldes, C.; Sherry, A. D.; Lazar, I.; Miseta, A.; Bogner, P.; Berenyi, E.; Sumegi, B.; Kiefer, 

G. E.; McMillan, K.; Maton, F.; Muller, R. N., Relaxometry, animal biodistribution, and 

magnetic resonance imaging studies of some new gadolinium (III) macrocyclic phosphinate 

and phosphonate monoester complexes. Magnetic Resonance in Medicine 1993, 30 (6), 

696-703. 

42 Roch, A.; Muller, R. N.; Gillis, P., Water relaxation by SPM particles: Neglecting the 

magnetic anisotropy? A caveat. Journal of Magnetic Resonance Imaging 2001, 14 (1), 94- 

96. 

43 Roch, A.; Muller, R. N., Longitudinal Relaxation of Water Protons in Colloidal Suspensions 

of Superparamagnetic Crystals, Proc. 11th Annu. Meet. Soc. Magn. Reson. Med., 1992, 11, 

1447. 

69



44 Ayant, Y.; Belorizly, E.; Alizon, J.; Gallice, Calculation of spectral density resulting from 

random translational movement with relaxation by magnetic dipolar interaction in liquids, J. 

Phys A, 1975, 36, 991-1004. 

45 Roch, A.; Gillis, P.; Ouakssim, A.; Muller, R. N., Proton magnetic relaxation in 

superparamagnetic aqueous colloids: a new tool for the investigation of ferrite crystal 

anisotropy. Journal of Magnetism and Magnetic Materials 1999, 201, 77-79. 

46 Muller, R. N.; Vander Elst, L.; Roch, A.; Peters, J. A.; Csajbok, E.; Gillis, P.; Gossuin, Y., 

Relaxation by metal-containing nanosystems. Advances in Inorganic Chemistry - Including 

Bioinorganic Studies, Vol 57 2005, 57, 239-292. 

47 Grealdes C. F. G. C., Laurent S., Classification and Basic properties of contrast agents for 

magnetic resonance imaging. Contrast Media Mol. Imaging, 2008, 3, 1 – 23. 

48 Yablonskiy, D. A.; Haacke, E. M., Theory of NMR signal behavior in magnetically 

inhomogeneous tissues. The static dephasing regime. Magnetic Resonance in Medicine 

1994, 32 (6), 749-763. 

49 Aime, S.; Crich, S. G.; Gianolio, E.; Giovenzana, G. B.; Tei, L.; Terreno, E., High sensitivity 

lanthanide(III) based probes for MR-medical imaging. Coordination Chemistry Reviews 

2006, 250 (11-12), 1562-1579. 

50 Vogl, T. J.; Hammerstingl, R.; Schwarz, W.; Mack, M. G.; Muller, P. K.; Pegios, W.; Keck, 

H.; EiblEibesfeldt, A.; Hoelzl, J.; Woessmer, B.; Bergman, C.; Felix, R., Superparamagnetic 

iron oxide-enhanced versus gadolinium-enhanced MR imaging for differential diagnosis of 

focal liver lesions. Radiology 1996, 198 (3), 881-887. 

51 Koenig, S. H.; Kellar, K. E., Theory of 1/T1 and 1/T2 NMRD profiles of solutions of magnetic 

nanoparticles. Magnetic Resonance in Medicine 1995, 34 (2), 227-233. 

52 Kellar, K. E.; Fujii, D. K.; Gunther, W. H. H.; Briley-Saebo, K.; Bjornerud, A.; Spiller, M.; 

Koenig, S. H., NC100150 injection, a preparation of optimized iron oxide nanoparticles for 

positive-contrast MR angiography. Journal of Magnetic Resonance Imaging 2000, 11 (5), 

488-494. 

53 Ahlström K. H., Johansson L. O., Rodenburg J. B., Ragnarsson A. S., Akeson P, Börseth 

A., Pulmonary MR angiography with ultrasmall superparamagnetic iron oxide particles as a 

blood pool agent and a navigator echo for respiratory gating: pilot study., Radiology, 1999; 

211, 865 - 869. 

54 Prasad P. V., Edelman R. R., Epstein F.H., Non-invasive evaluation of intra-renal 

oxygenation using BOLD MRI. Circulation, 1996; 94, 3271 - 3275. 

55 Van Wagoner M., Worah D., Gadodiamide injection. Non-invasive evaluation of intra-renal 

oxygenation using BOLD MRI. Invest. Radiol., 1993, 28, S44 - S48. 

56 Oksendal, A. N.; Hals, P. A., Biodistribution and toxicity of MR imaging contrast media. 

Jmri-Journal of Magnetic Resonance Imaging 1993, 3 (1), 157-165. 

70


57 Bellin, M. F.; Vasile, M.; Morel-Precetti, S., Currently used non-specific extracellular MR 

contrast media. European Radiology 2003, 13 (12), 2688-2698. 

58 Bellin, M.-F., MR contrast agents, the old and the new. European Journal of Radiology 

2006, 60 (3), 314-323. 

59 Saeed, M.; Wendland, M. F.; Higgins, C. B., Blood pool MR contrast agents for 

cardiovascular imaging. Journal of Magnetic Resonance Imaging 2000, 12 (6), 890-898. 

60 Knopp, M. V.; von Tengg-Kobligk, H.; Floemer, F.; Schoenberg, S. O., Contrast agents for 

MRA: Future directions. Journal of Magnetic Resonance Imaging 1999, 10 (3), 314-316. 

61 Daldrup-Link, H. E.; Brasch, R. C., Macromolecular contrast agents for MR mammography: 

current status. European Radiology 2003, 13 (2), 354-365. 

62 Bogdanov A. A., Lewin M., Weissleder R., Approaches and agents for imaging the vascular 

system, Adv. Drug Deliv. Rev., 1999, 37, 279 - 293. 

63 Young S. W., Quing F., Harriman A., Sessler J. L., Dow W. C., Mody T. D., Hemmi G. W., 

Hao Y., Miller R. A., Gadolinium(III) texaphyrin: A tumor selective radiation sensitizer that is 

detectable by MRI. Proc. Natl Acad. Sci. USA. 1996, 93, 6610 - 6615. 

64 Sipkins, D. A.; Cheresh, D. A.; Kazemi, M. R.; Nevin, L. M.; Bednarski, M. D.; Li, K. C. P., 

Detection of tumor angiogenesis in vivo by alpha(v)beta(3)-targeted magnetic resonance 

imaging. Nature Medicine 1998, 4 (5), 623-626. 

65 Winter, P. M.; Caruthers, S. D.; Kassner, A.; Harris, T. D.; Chinen, L. K.; Allen, J. S.; Lacy, 

E. K.; Zhang, H. Y.; Robertson, J. D.; Wickline, S. A.; Lanza, G. M., Molecular Imaging of 

angiogenesis in nascent vx-2 rabbit tumors using a novel alpha(v)beta(3)-targeted 

nanoparticle and 1.5 tesla magnetic resonance imaging. Cancer Res 2003, 63 (18), 5838- 

5843. 

66 Aime, S.; Cabella, C.; Colombatto, S.; Crich, S. G.; Gianolio, E.; Maggioni, F., Insights into 

the use of paramagnetic Gd(III) complexes in MR-molecular imaging investigations. Journal 

of Magnetic Resonance Imaging 2002, 16 (4), 394-406. 

67 Crich, S. G.; Biancone, L.; Cantaluppi, V.; Esposito, D. D. G.; Russo, S.; Camussi, G.; 

Aime, S., Improved route for the visualization of stem cells labeled with a Gd-/Eu-chelate as 

dual (MRI and fluorescence) agent. Magnetic Resonance in Medicine 2004, 51 (5), 938- 

944. 

68 Terreno, E.; Crich, S. G.; Belfiore, S.; Biancone, L.; Cabella, C.; Esposito, G.; Manazza, A. 

D.; Aime, S., Effect of the intracellular localization of a Gd-based imaging probe on the 

relaxation enhancement of water protons. Magnetic Resonance in Medicine 2006, 55 (3), 

491-497. 

69 Heckl S., Pipkorn R., Waldeck W., Spring H., Jenne J., von der Lieth C. W., Corban- 

Wilhelm H., Debus J., Braun K., Intracellular visualization of prostate cancer using magnetic 

resonance imaging. Cancer Res., 2003, 63, 4766 - 4772. 

71



70 Schmitt-Willich H., Brehm M., Evers C. L. J., Michl G., Muller-Fahrnow A., Petrov O., 

Platzek J., Raduchel B., Sulzle D., Synthesis and physicochemical characterization of a 

new gadolinium chelate: the liver-specific magnetic resonance imaging contrast agent Gd– 

EOB–DTPA, Inorg. Chem., 1999, 38, 1134 - 1144. 

71 Schuhmann-Giampieri G., Schmitt-Willich H., Press W. R., Negishi C., Weinmann H. J., 

Speck U., Preclinical evaluation of Gd–EOB–DTPA as a contrast agent in MR imaging of 

the hepatobiliary system. Radiology, 1992, 183, 59 - 64. 

72 Lim K., O., Stark D. D., Leese P. T., Pfefferbaum A., Rocklage S. M., Quay SC. 

Hepatobiliary MR imaging: first human experience with MnDPDP, Radiology, 1991, 178, 79 

- 82. 

73 Reimer P., Rummeny E. J., Daldrup H. E., Balzer T., Tombach B., Berns T., Peters P. E., 

Clinical results with Resovist: a phase 2 clinical trial. Radiology, 1995, 195, 489 - 496. 

74 Hamm B., Thoeni R. F., Gould R. G., Bernardino M. E., Luning M., Saini S., Mahfouz A. E., 

Taupitz M., Wolf K. J., Focal liver lesions: characterization with nonenhanced and dynamic 

contrast material- enhanced MR imaging, Radiology, 1994, 190, 417 - 423. 

75 Misselwitz B, Platzek J, Radüchel B, Oellinger JJ, Weinmann H-J. Gadofluorine 8: initial 

experience with a new contrast medium for interstitial MR lymphography. MAGMA, 1999, 8, 

1352 - 8661. 

76 Silva A .C., Lee J. H., Aoki I., Koretsky A. P., Manganese-enhanced magnetic resonance 

imaging (MEMRI): methodological and practical considerations. NMR Biomed., 2004, 17, 

532 - 543. 

77 Watanabe T., Tammer R., Boretius S., Frahm J., Michaelis T., Chromium( VI) as a novel 

MRI contrast agent for cerebral white matter-preliminary results in mouse brain in vivo, 

Magn. Reson. Med., 2006, 56, 1 - 6. 

78 Poduslo J. F., Wengenack T. M., Curran G. L., Wisniewski T., Sigurdsson E.M., Macura S. 

I., Borowski B. J., Jack C. R., Molecular targeting of Alzheimer’s amyloid plaques for 

contrast-enhanced magnetic resonance imaging. Neurobiol. Dis., 2002, 11, 315 - 329. 

79 Wadghiri Y. Z., Sigurdsson E. M., Sadowski M., Elliott J. I., Li Y. S., Scholtzova H., Tang C. 

Y., Aguinaldo G., Pappolla M., Duff K., Wisniewski T., Turnbull D. H., Detection of 

Alzheimer’s amyloid in transgenic mice using magnetic resonance microimaging, Magn. 

Reson. Med., 2003, 50, 293 - 302. 

80 Rubin D. L., Falk K. L., Sperling M. J., Ross M., Saini S., Rothman B., Shellock F., Zerhouni 

E., Stark D., Outwater E. K., Schmiedl U., Kirby L. C., Chezmar J., Coates T., Chang M., 

Silverman J. M., Rofsky N., Burnett K., Engel J., Young S. W., A multicenter clinical trial of 

Gadolite Oral Suspension as a contrast agent for MRI. J. Magn. Reson. Imag., 1997, 7, 865 

- 872. 

72


81 Lonnemark M., Hemmingsson A., Carlsten J., Ericsson A., Holtz E., Klaveness J., 

Superparamagnetic particles as an MRI contrast agent for the gastrointestinal tract, Acta 

Radiol., 1988, 29, 599 - 602. 

82 Gillies R. J., Raghunand N., Karczmar G. S., Bhujwalla Z. M., MRI of the tumor 

microenvironment. J. Magn. Reson. Imaging, 2002, 16, 430 - 450. 

83 Chesler M., Regulation and modulation of pH in the brain. Physiol Rev., 2003, 83, 1183 – 

1221. 

84 Kalman F. K., Woods M., Caravan P., Jurek P., Spiller M, Tircso G., Kiraly R., Brücher E., 

Sherry A. D., Potentiometric and relaxometric properties of a gadolinium-based MRI 

contrast agent for sensing tissue pH. Inorg. Chem., 2007; 46, 5260 - 5270. 

85 Mikawa M, Miwa N, Brautigam M, Akaike T., Maruyama A., A pH-sensitive contrast agent 

for functional magnetic resonance imaging (MRI), Chem Lett., 1998, 7, 693 – 694 

86 Aime S., Crich S. G., Botta M., Giovenzana G., Palmisano G., Sisti M., A macromolecular 

Gd(III) complex as pH-responsive relaxometric probe for MRI applications, Chem Commun., 

1999, 1577 - 1578. 

87 Lowe M. P., Parker D., Reany O., Aime S., Botta M., Castellano G., Gianolio E., Pagliarin 

R., pH-dependent modulation of relaxivity and luminescence in macrocyclic gadolinium and 

europium complexes based on reversible intramolecular sulfonamide ligation. J. Am. Chem. 

Soc., 2001, 123, 7601 - 7609. 

88 Hovland R., Glogard C., Aasen A. J., Klaveness J., Gadolinium DO3A derivatives mimicking 

phospholipids; preparation and in vitro evaluation as pH responsive MRI contrast agents. J. 

Chem. Soc. Perkin Trans 2, 2001, 6, 929 - 933. 

89 Toth E., Bolskar R. D., Borel A., González G., Helm L., Merbach A. E., Sitharaman B., 

Wilson L. J., Water-soluble gadofullerenes: toward high-relaxivity, pH-responsive MRI 

contrast agents. J. Am. Chem. Soc. 2005, 127, 799 - 805. 

90 Zhang S., Wu K., Sherry A. D., A novel pH-sensitive MRI contrast agent, Angew Chem Int 

Ed Engl., 1999, 38, 3192 - 3194. 

91 Raghunand N., Howison C., Sherry A. D., Zhang S., Gillies R. J., Renal and systemic pH 

imaging by contrast-enhanced MRI. Magn Reson Med., 2003, 49, 249 - 257. 

92 Garcia-Martin, M. L.; Martinez, G. V.; Raghunand, N.; Sherry, A. D.; Zhang, S. R.; Gillies, R. 

J., High resolution pH(e) imaging of rat glioma using pH-dependent relaxivity. Magnetic 

Resonance in Medicine 2006, 55 (2), 309-315. 

93 Ali, M. M.; Woods, M.; Caravan, P.; Opina, A. C. L.; Spiller, M.; Fettinger, J. C.; Sherry, A. 

D., Synthesis and relaxometric studies of a dendrimer-based pH-responsive MRI contrast 

agent. Chemistry-a European Journal 2008, 14 (24), 7250-7258. 

94 Aime, S.; Botta, M.; Fasano, M.; Terreno, E.; Kinchesh, P.; Calabi, L.; Paleari, L., A new 

ytterbium chelate as contrast agent in chemical shift imaging and temperature sensitive 

probe for MR spectroscopy. Magnetic Resonance in Medicine 1996, 35 (5), 648-651. 

73



95 Zuo, C. S.; Mahmood, A.; Sherry, A. D., TmDOTA(-): A sensitive probe for MR thermometry 

in vivo. Journal of Magnetic Resonance 2001, 151 (1), 101-106. 

96 Fossheim, S. L.; Il'yasov, K. A.; Hennig, J.; Bjornerud, A., Thermosensitive paramagnetic 

liposomes for temperature control during MR imaging-guided hyperthermia: In vitro 

feasibility studies. Acad Radiol 2000, 7 (12), 1107-1115. 

97 Louie, A. Y.; Huber, M. M.; Ahrens, E. T.; Rothbacher, U.; Moats, R.; Jacobs, R. E.; Fraser, 

S. E.; Meade, T. J., In vivo visualization of gene expression using magnetic resonance 

imaging. Nature Biotechnology 2000, 18 (3), 321-325. 

Louie, A. Y.; Huber, M. M.; Ahrens, E. T.; Rothbacher, U.; Moats, R.; Jacobs, R. E.; Fraser, S. E.; 

Meade, T. J., In vivo visualization of gene expression using magnetic resonance imaging. 

Nature Biotechnology 2000, 18 (3), 321-325. 

98 Moats, R. A.; Fraser, S. E.; Meade, T. J., A ''smart'' magnetic resonance imaging agent that 

reports on specific enzymatic activity. Angewandte Chemie-International Edition in English 

1997, 36 (7), 726-728. 

99 Nivorozhkin A. L., Kolodziej A. F., Caravan P., Greenfield M. T., Lauffer R. B., McMurry T. 

J., Enzyme-activated Gd(III) magnetic resonance imaging contrast agents with a prominent 

receptor-induced magnetization enhancement. Angew Chem Int Ed Engl, 2001, 40, 2903 – 

2906. 

100 Lauffer RB, McMurry TJ, Dunham SO, et al. Epix Medical Inc. Bioactivated diagnostic 

imaging contrast agents., Patent publication number WO9736619. 10 9, 1997. 

101 Anelli, P. L.; Bertini, I.; Fragai, M.; Lattuada, L.; Luchinat, C.; Parigi, G., Sulfonamide- 

functionalized gadolinium DTPA complexes as possible contrast agents for MRI: A 

relaxometric investigation. European Journal of Inorganic Chemistry 2000, (4), 625-630. 

102 Shiftan L, Israely T, Cohen M, Frydman V., Dafni H., Stern R., Neeman M., Magnetic 

resonance imaging visualization of hyaluronidase in ovarian carcinoma. Cancer Res., 2005, 

65, 10316 – 10323 

103 Chen, J. W.; Pham, W.; Weissleder, R.; Bogdanov, A., Human myeloperoxidase: A potential 

target for molecular MR imaging in atherosclerosis. Magnetic Resonance in Medicine 2004, 

52 (5), 1021-1028. 

104 Duimstra, J. A.; Femia, F. J.; Meade, T. J., A gadolinium chelate for detection of beta- 

glucuronidase: A self-immolative approach. Journal of the American Chemical Society 

2005, 127 (37), 12847-12855. 

105 Matsumoto, H.; Inoue, N.; Takaoka, H.; Hata, K.; Shinke, T.; Yoshikawa, R.; Masai, H.; 

Watanabe, S.; Ozawa, T.; Yokoyama, M., Depletion of antioxidants is associated with no- 

reflow phenomenon in acute myocardial infarction. Clinical Cardiology 2004, 27 (8), 466- 

470. 

74


106 Aime, S.; Botta, M.; Gianolio, E.; Terreno, E., A p(O-2)-responsive MRI contrast agent 

based on the redox switch of manganese(II/III) - Porphyrin complexes. Angewandte 

Chemie-International Edition 2000, 39 (4), 747-750. 

107 Raghunand, N.; Jagadish, B.; Trouard, T. P.; Galons, J. P.; Gillies, R. J.; Mash, E. A., 

Redox-sensitive contrast agents for MRI based on reversible binding of thiols to serum 

albumin. Magnetic Resonance in Medicine 2006, 55 (6), 1272-1280. 

108 Atanasijevic T, Shusteff M, Fam P, Jasanoff A., Calcium-sensitive MRI contrast agents 

based on superparamagnetic iron oxide nanoparticles and calmodulin, Proc Natl Acad Sci U 

S A, 2006, 103, 14707 - 14712. 

109 Li, W. H.; Fraser, S. E.; Meade, T. J., A calcium-sensitive magnetic resonance imaging 

contrast agent. Journal of the American Chemical Society 1999, 121 (6), 1413-1414. 

110 Hanaoka, K.; Kikuchi, K.; Urano, Y.; Narazaki, M.; Yokawa, T.; Sakamoto, S.; Yamaguchi, 

K.; Nagano, T., Design and synthesis of a novel magnetic resonance imaging contrast 

agent for selective sensing of zinc ion. Chemistry & Biology 2002, 9 (9), 1027-1032. 

111 Aime, S.; Botta, M.; Fasano, M.; Terreno, E., Paramagnetic Gd(III)–Fe(III) heterobimetallic 

complexes of DTPA-bis-salicylamide. Spectrochimica Acta Part a-Molecular and 

Biomolecular Spectroscopy 1993, 49 (9), 1315-1322. 

112 Parac-Vogt, T. N.; Elst, L. V.; Kimpe, K.; Laurent, S.; Burtea, C.; Chen, F.; Van Deun, R.; Ni, 

Y.; Muller, R. N.; Binnemans, K., Pharmacokinetic and in vivo evaluation of a self- 

assembled gadolinium(III)-iron(II) contrast agent with high relaxivity. Contrast Media & 

Molecular Imaging 2006, 1 (6), 267-278. 

113 Glogard, C.; Stensrud, G.; Aime, S., Novel radical-responsive MRI contrast agent based on 

paramagnetic liposomes. Magnetic Resonance in Chemistry 2003, 41 (8), 585-588. 

114 Ward, K. M.; Aletras, A. H.; Balaban, R. S., A new class of contrast agents for MRI based 

on proton chemical exchange dependent saturation transfer (CEST). Journal of Magnetic 

Resonance 2000, 143 (1), 79-87. 

115 Zhou, J. Y.; Wilson, D. A.; Sun, P. Z.; Klaus, J. A.; van Zijl, P. C. M., Quantitative description 

of proton exchange processes between water and endogenous and exogenous agents for 

WEX, CEST, and APT experiments. Magnetic Resonance in Medicine 2004, 51 (5), 945- 

952. 

116 Guivel-Scharen, V.; Sinnwell, T.; Wolff, S. D.; Balaban, R. S., Detection of proton chemical 

exchange between metabolites and water in biological tissues. Journal of Magnetic 

Resonance 1998, 133 (1), 36-45. 

117 Zhang, S. R.; Winter, P.; Wu, K. C.; Sherry, A. D., A novel europium(III)-based MRI contrast 

agent. Journal of the American Chemical Society 2001, 123 (7), 1517-1518. 

118 Zhang, S. R.; Merritt, M.; Woessner, D. E.; Lenkinski, R. E.; Sherry, A. D., PARACEST 

agents: Modulating MRI contrast via water proton exchange. Accounts of Chemical 

Research 2003, 36 (10), 783-790. 

75



119 Aime, S.; Delli Castelli, D.; Terreno, E., Novel pH-reporter MRI contrast agents. 

Angewandte Chemie-International Edition 2002, 41 (22), 4334-4336. 

120 Aime, S.; Barge, A.; Castelli, D. D.; Fedeli, F.; Mortillaro, A.; Nielsen, F. U.; Terreno, E., 

Paramagnetic lanthanide(III) complexes as pH-sensitive chemical exchange saturation 

transfer (CEST) contrast agents for MRI applications. Magnetic Resonance in Medicine 

2002, 47 (4), 639-648. 

121 Zhang, S. R.; Michaudet, L.; Burgess, S.; Sherry, A. D., The amide protons of an 

ytterbium(III) dota tetraamide complex act as efficient antennae for transfer of magnetization 

to bulk water. Angewandte Chemie-International Edition 2002, 41 (11), 1919-1921. 

122 Aime S., Carrera C., Castelli D. D., Crich S. G., Terreno E., Tunable imaging of cells labeled 

with MRI-PARACEST agents. Angew. Chem. Int. Edn Engl., 2005; 44, 1813 - 1815. 

123 Aime, S.; Barge, A.; Castelli, D. D.; Fedeli, F.; Mortillaro, A.; Nielsen, F. U.; Terreno, E., 

Paramagnetic lanthanide(III) complexes as pH-sensitive chemical exchange saturation 

transfer (CEST) contrast agents for MRI applications. Magnetic Resonance in Medicine 

2002, 47 (4), 639-648. 

124 Terreno, E.; Castelli, D. D.; Cravotto, G.; Milone, L.; Aime, S., Ln(III)-DOTAMGIY 

complexes: A versatile series to assess the determinants of the efficacy of paramagnetic 

chemical exchange saturation transfer agents for magnetic resonance imaging applications. 

Investigative Radiology 2004, 39 (4), 235-243. 

125 Goffeney, N.; Bulte, J. W. M.; Duyn, J.; Bryant, L. H.; van Zijl, P. C. M., Sensitive NMR 

detection of cationic-polymer-based gene delivery systems using saturation transfer via 

proton exchange. Journal of the American Chemical Society 2001, 123 (35), 8628-8629. 

126 Snoussi, K.; Bulte, J. W. M.; Gueron, M.; van Zijl, P. C. M., Sensitive CEST agents based 

on nucleic acid imino proton exchange: Detection of poly(rU) and of a dendrimer-poly(rU) 

model for nucleic acid delivery and pharmacology. Magnetic Resonance in Medicine 2003, 

49 (6), 998-1005. 

127 Aime, S.; Castelli, D. D.; Terreno, E., Highly sensitive MRI chemical exchange saturation 

transfer agents using liposomes. Angewandte Chemie-International Edition 2005, 44 (34), 

5513-5515. 

128 Winter, P. M.; Cai, K.; Chen, J.; Adair, C. R.; Kiefer, G. E.; Athey, P. S.; Gaffney, P. J.; Buff, 

C. E.; Robertson, J. D.; Caruthers, S. D.; Wickline, S. A.; Lanza, G. M., Targeted 

PARACEST nanoparticle contrast agent for the detection of fibrin. Magnetic Resonance in 

Medicine 2006, 56 (6), 1384-1388. 

129 Bowers, C. R.; Weitekamp, D. P., Para-Hydrogen and synthesis allow dramatically 

enhanced nuclear alignment. Journal of the American Chemical Society 1987, 109 (18), 

5541-5542. 

130 Abragam A., The Principles of Nuclear Magnetism. Clarendon Press: Oxford, 1961. 

76


131 Johansson, E.; Olsson, L. E.; Mansson, S.; Petersson, J. S.; Golman, K.; Stahlberg, F.; 

Wirestam, R., Perfusion assessment with bolus differentiation: A technique applicable to 

hyperpolarized tracers. Magnetic Resonance in Medicine 2004, 52 (5), 1043-1051. 

132 Magnusson, P.; Johansson, E.; Mansson, S.; Petersson, J. S.; Chai, C.-M.; Hansson, G.; 

Axelsson, O.; Golman, K., Passive catheter tracking during interventional MRI using 

hyperpolarized C-13. Magnetic Resonance in Medicine 2007, 57 (6), 1140-1147. 

133 Mansson, S.; Johansson, E.; Magnusson, P.; Chai, C. M.; Hansson, G.; Petersson, J. S.; 

Stahlberg, F.; Golman, K., C-13 imaging - a new diagnostic platform. European Radiology 

2006, 16 (1), 57-67. 

77


Fluorescence Imaging Background Concepts 

78 

3. 

Fluorescence Imaging 

Background Concepts



3.1. Optical Imaging 80 

3.2. Luminescence 82 

3.3. Tissue Optical Properties 91 

3.4. Optical Imaging Technology 93 

3.5. Optical Contrast Agents 95 


79



3.1. OPTICAL IMAGING 

Unlike any other imaging modality, optical imaging derives from the fact that it is 

possible to combine conventional display of tissue volumes using direct-optical, 

transillumination or tomographic techniques, with the capability of gaining information on 

molecular properties and function due to the high instrumental sensitivity for optical 

signals. Taking into account the various principles of optical imaging methods but not 

disregarding the limitation of penetration depth and spatial resolution in thick tissues, a 

number of potential applications are envisioned for clinical diagnostics. Over the past few 

years optical imaging techniques have joined the available methods for the assessment of 

tissue anatomy, physiology, metabolic and molecular function. This technology is 

attracting a lot of interest due to the fact that fluorescent dyes can be detected at low 

concentrations and non-ionizing, harmless radiation can be applied repeatedly to the 

patient. This technology has the additional advantage that it is less expensive (compared 

to other imaging techniques), small in size and, therefore, easily at hand to solve clinical 

problems. Table 3.1 summarizes the range of applications of optical imaging. The two 

leading imaging technologies are optical coherence tomography (OCT) and diffusion 

imaging. 

Within this perspective the design of contrast agents for optical in vivo imaging of 

diseased tissues has gained a remarkable vitality. 1,2 Light is one of the most convenient 

vectors for transmitting signals that can easily reach regions of a complex molecular 

structure that are not accessible to other molecular messengers. When appropriate 

wavelengths are used, penetration depth may be substantial and light can reach regions 

of complex molecular structure which are not accessible to other molecular probes. 

Sophisticated contrast agents have been synthesized and characterized for their 

capability to monitor disease-specific anatomic, physiological and molecular parameters 

through their optical signals. 

80



Table 3.1. Major application areas of optical imaging. 

In vivo optical imaging In vitro diagnostics 

Non invasive clinical imaging 

Tissue imaging 

Microvascular 

Ophthalmology 

Dermatology 

Brain 

Breast 

Preclinical imaging 

Pharmacokinetic 

Drug-efficacy testing 

Patient monitoring 

Intraoperative 

Biosensor implants 

Pulse oximetry 

Optical imaging 

Cancer-cell detection 

81 

Genomics 

Proteomics 

Microarray readers 

Biosensors 

Pathogen detection 

Point-of-care diagnostics 

Semiconductor quantum dots (QDs) (e.g. CdSe nanocrystals with 2-10 nm 

diameter) and their bioconjugates are potential canditates as they are highly luminescent, 

tunable in the entire visible range and display a superior photostability compared to 

organic luminophores. This type of nanoparticles (NP) have been exploited for both in 

vitro analyses and in vivo imaging. Although their has been some reports regarding QDs 

with near-infrared (NIR) emission and with low toxicity Therefore, their introduction into the 

biological and medical media has been still slow. 3,4,5 Alternately trivalent lanthanide ions, 

Ln 3+ , present another option to organic luminescent stains in view of their singular 

properties. They enable easy spectral and time discrimination of their emission bands 

which span both the visible and NIR ranges. 

The very first staining of biological cells with lanthanides dates back to 1969 

by Scaff et al. 6 Bacterial smears (Escherichia coli cell walls) were treated with 

aqueous ethanolic solutions of europium chelate of 4, 4, 4-trifluoro-1-(2-thienyl)



1,3-butanedione (thenoyltrifluoroacetonate, TTA) and bright red spots appeared under 

mercury lamp illumination. Only in the mid-1970s that further attention was accorded to 

luminescent lanthanide bioprobes when Finnish researchers in Turku proposed Eu 3+ , 

Sm 3+ , Tb 3+ , and Dy 3+ polyaminocarboxylates and β-diketonates as luminescent sensors in 

time-resolved luminescent (TRL) immunoassays. 7,8 This technological push projected 

broader interest and subsequent developments. Examples of these developments can be 

found in homogeneous TRL assays, 9 optimization of bioconjugation methods for 

lanthanide luminescent chelates 10 and time-resolved luminescence microscopy (TRLM) 11 

that resulted in applications of lanthanide luminescent bioprobes (LLBs) 12,13 in many fields 

of biology, biotechnology and medicine, including tissue 14,15 and cell imaging, 16,17 analyte 

sensing 18 and monitoring drug delivery. 19 

In order to design optical contrast agents it is necessary to first understand the 

concept of luminescence, the optical properties of tissues as well as comprehend the 

existing technology. 

3.2. LUMINESCENCE 

The term “Luminescence” was defined in 1888, by the German physicist Eilhardt 

Wiedemann, as the light emission characteristics not conditioned by an increase in 

temperature. 20 Within this thesis IUPAC rules 21 regarding molecular luminescence 

spectroscopy will be used, where the term ‘‘fluorescence’’ is used for processes which 

occur without change in spin. These are typically 0 

or 2 

82 

S S (singlet to singlet ground state) 

2 2 

F5 2 

F7 

(Yb 3+ ) transitions and ‘‘phosphorescence’’ for transitions implying a change 

in spin, most commonly 0 

1 

T1 S (triplet to singlet ground state) or J 

F 

5 7 

3+ 

D0 (Eu ) 

transitions. The energy type used in the excitation of a luminescent material can be used 

to classify luminescent processes, as photoluminescence, when the excitation is by 

electromagnetic radiation; cathodoluminescence, when a beam of electrons is used; 

electroluminescence, when excitation is effected by an electric voltage; 

triboluminescence, when mechanical energy is the source of excitation; and 

chemiluminescence, when energy of a chemical reaction is employed. 22 

Not all materials exhibit the luminescence phenomenon even though their molecules 

can be excited to a higher state. This is due to the existence of a non-radiative pathway in



the return to the ground state (GS). This transition from a vibration energy level of an 

excited state to a vibration energy level of the GS is called internal conversion. The 

population of an excited state and the energy difference between the GS and excited 

states are among the important factors that affect the luminescence of a material. Figure 

3.1 illustrates some of the typical processes involved in luminescence. The emission 

quantum yield indicates the efficiency of the luminescent material and will be defined in 

the next section. 23 

Figure 3.1. Partial energy-level diagram for a photoluminescent system. 24 

LANTHANIDE LUMINESCENCE 

The complex optical properties of the trivalent lanthanide ions, Ln 3+ , derives from 

special features of the electronic [Xe] 

83 

N 

4 configuration (N=0-14)). The third ionization 

f 

state is characterized (with the exception of lutetium, cerium and gadolinium) by the



removal of two electrons from the 6s orbital and one form the 4f orbital. Lanthanides (Ce– 

Lu) are unique among the elements (with the exception of the actinides) in markedly 

resembling each other in their chemical properties, particularly regarding oxidation states, 

due to their electronic configuration. 25,26 

These configurations generate a large variety of energy levels. Figure 3.2 displays 

partial energy diagrams for the lanthanide aquo ions. 21 Eu 3+ , Gd 3+ , and Tb 3+ are the best 

ions with respect to the energy gap requirement (as larger is that energy gap higher is the 

probability of radiative transitions) as they present a larger band gaps when compared to 

other trivalent lanthanides with E 

5 7 F 

4 

0 

=12 300 

84 

5 7 

D0 F6 

, 32 200 

6 8 

P7 2 

S7 

2 and 14 800 

D , cm -1 , respectively. Gd 3+ emits in the UV and it is not very useful as a 

luminescent probe since its luminescence interferes with either the emission or absorption 

processes in the organic part of the complex molecules. 23 The Russel-Saunders spin-orbit 

coupling scheme * is normally used to characterize these energy levels, that is based on 

three quantum numbers, S , L and J . It should be noted that this sheme does not 

characterize completely the energy levels, nevertheless, it is the commonly used. 27 The 4f 

electrons are not the outermost ones and they are “shielded “ from external fields by the 

two electronic orbitals 5s 2 and 5p 6 with larger radial extension. Thus 4f electrons are only 

weakly perturbed by surrounding ligands charges resulting in special spectroscopic 

properties with parity-forbidden f 

4 f 4 absorptions that have characteristic narrow-line 

emission (high purity). The emission occurs mostly in the visible and near-infrared (NIR) 

ranges and with very low molar absorption coefficients, ε (or absorption cross-sections, 

σABS). These are typically < 10 mol -1 cm -1 ( or σABS < 4 × 10 -20 cm 2 ) when compared with 

* In this notation the orbital angular momenta of the individual electrons add up to form a resultant 

orbital angular momentum L . Within the same strategy the individual spin angular momenta are 

presumed to couple to produce a resultant spin angular momentum S . Finally L and S combine 

scattering events and form the total angular momentum J . In Russel-Saunders coupling scheme, 

( 2 1)( 2L 

1 

the terms J and S define one of the ) 

S terms (or multiplets) of the configuration, 

generic represented by J , where ( 2S 

1) 

indicates the spin multiplicity. 

( 2S 

1) 

L 

S L ,..., S L S L 

J 

L 

S,..., 

L S S L 

(3.1)



the molar absorbtion coefficient for d 

3d 3 transitions values that are10 times higher, or 

for ligand-to-metal charge-transfer (LMCT) transitions that are 100 times higher. 28 

These transitions are formally parity-forbidden, the lifetimes of the excited states are 

long, which allows the use of time-resolved detection, a definitive asset for bioassays, 29-34 

11,35,36 2930313233343536 

and luminescence microscopy. 

Figure 3.2. Partial energy diagrams for the lanthanide aquo ions. The main 

luminescent levels are drawn in red, while the fundamental level is indicated in 

blue. 23 

The direct Ln 3+ photoexcitation is not very efficient as already mentioned. The 

design of lanthanide complexes, in which the ligands incorporate organic chromophores 

strongly bonded to the 4 f metal center, improves significantly the Ln 3+ luminescence 

intensity. Weissman demonstrated in 1942 that the excitation of lanthanide complexes into 

the ligand states result in metal-centered luminescence. Part of the energy absorbed by 

the organic receptor(s) is transferred onto Ln 3+ excited states and sharp emission bands 

originating from the metal ion are detected after rapid internal conversion to the emitting 

level. These chromophores typically present effective absorption cross-sections 10 4 –10 5 

85



times higher and over a much broader spectral range than the Ln 3+ corresponding ones. 

The phenomenon is termed sensitization of the metal-centered luminescence (also 

referred to as “antenna effect”) and is quite complex. 37 

When the Ln 3+ ions are introduced into a host, the ion experiences an 

inhomogeneous electrostatic field, the so-called crystal field, which is produced by the 

Ln 3+ ligands. In addition to the splitting of the energy of the f 

86 

4 orbital by spin-orbit 

coupling, individual J levels of a lanthanide ion are split further by the crystal field into a 

maximum of 2J+1 components, depending on the local symmetry of the Ln 3+ ion, making 

the lanthanide emissions extremely sensitive to their environments. Figure 3.3 displays a 

schematic representation and order of magnitude of the effects of the intra-atomic and 

ligand field interactions acting on the 

N 

4 configuration, in particular of the Eu 3+ ion. 

f 

The 4 f 4 f transitions have essentially electric-dipole (ED) and magnetic dipole 

(MD) character. ED transitions require a change of parity of the electron wavefunction 

and, then, they should be strictly forbidden within the f 

4 configuration (due to the Laporte 

selection rule). Contrarily, MD transitions between those levels are permitted. For Ln 3+ 

ions localised in crystalline sites without inversion symmetry, however, a mixing of 

opposite-parity states into the 

selection rule and then in the observation of intra- f 

N 

4 levels occur resulting in a relaxation of the Laporte 

f 

4 forced ED transitions. 

Important parameters such as radiative lifetime values, emission quantum 

efficiency and quantum yield should be determined in order to adequately 

characterize an optical material. The concepts referred in this thesis are further 

described in Carlos L.D et al. . 28 The radiative lifetime ( RAD 

f ' J 

N 

) for a given ' 

excited state is obtained as the reciprocal of the total transition radiative probability 

A T , which is the sum of all possible radiative decay rates from the state ' 

towards lower levels J . The RAD 

 

is given by: 28 

1 1 

RAD 

(3.2) 

A 

AT J 'J 

J 

where AJ ' J , the radiative transition probability, or Einstein’s spontaneous 

emission rate. 

J



Figure 3.3. Schematic representation and order of magnitude of the effect 

intra-atomic and ligand field interactions acting on the 

the Eu 3+ (Top) and Tb 3+ ions (Bottom). 38 

87 

N 

4 configuration of 

f



The emission quantum efficiency of a given excited level ' 

88 

J (also refereed by some 

authors as intrinsic quantum yield 39 ) is given by the ratio between the radiative decay rate 

and the total decay rate. This expression also includes all other processes contributing to 

the depopulation of level ' 

processes. 28 

 

 

 

J , such as nonradiative decay paths and energy-transfer 

The emission quantum efficiency ( ) is calculated by: 

EXP 

RAD 

 

(3.3) 

RAD 

A 

RAD 

A 

A 

NRAD 

where the radiative lifetime described above and the experimental lifetime 

 

 

EXP RAD NRAD are given by the reciprocal of the total decay rate. 28 

The emission quantum yield ( ) is an experimentally evaluated quantity given by 

the ratio of the number of photons emitted and the absorbed photons and can be 

described by: 

number of 

emitted 

photons 

 

(3.4) 

number of 

absorved photons 

Typically, the emission quantum yield includes the absorption efficiency, the 

intersystem crossing efficiency, the donor–Ln 3+ energy-transfer efficiency, the intra-Ln 3+ 

nonradiative decaypaths efficiency, , and the ' 

According to equation 3.4 we can denote that 

ligand excited states are transferred to the Ln 3+ excited levels. 28 

In the particular case of the Eu 3+ ions the intra- f 

J -level emission quantum efficiency . 

If all the energy absorbed by the 

4 transitions generally occur 

between the J F 

5 7 

D0 ( J 0 6 ) levels. These transitions are essentially of the induced 

5 7 

ED type, with the exception of the D0 F1 

transition that is a 100% MD transition. 

5 7 

Attending to the pure MD character of the D0 F1 

transition it is possible to determine 

intensity parameters from the emission spectrum. Since this transition does not depend on 

the local ligand field seen by Eu 3+ it can become a reference for the entire spectrum. The 

5 7 

D0 F1 

spontaneous decay rate, 01 

A , is given by 

3 

1 

A01 A'01n 

, where A'01 14. 

65s 

in



vacuum. 40 Within these parameters it is possible to express the intensity of the 6 

transitions, J 

I 0 , in terms of the area of their emission curves, J 

5 D0 

S J 

I0 J hc A0 

J N 

0 

89 

S 0 : 28 

(3.5) 

5 

N D0 

is the population of the 0 

where 

the total radiative decay rate as: 28 

A 

T 

 

6 

A hc 

S 

0J 

01 

A0 

J 

J 0 01 J 0 

0J 

6 

S0 

J 

hc 

5 

D 

7 

0 

F0 

5 D emitting level; it is possible to express 

(3.6) 

It should be noted that the radiative contribution can be calculated from the relative 

7 

0 

F0 

intensities of the 4 

5 

D transitions, as the branching ratio for the 6 

5 

7 

D F 

transitions should be neglected due to their poor relative intensity. As already mentioned 

above, the emission quantum efficiency is defined by the ratio between experimental and 

radiative lifetimes. The intensity parameters 

J 

will be given by:28 

3h 

9 1 

 

A 

4 2 3 

2 

2 0J 

(3.7) 

64 

e 2 

nn 

2 

5 7 

D U F 

with 6 

0 

J 2, 

4, 

. Values for the squared reduced matrix elements are 0.0032 and 

0.0023 for 2 

J and J 4 

, respectively. 28,41 As previously mentioned, the 

5 7 

D0 F6 

transition has very low intensity making it very difficult for the determination of 

the rank 6 parameter. 

The sensitivity of the f 

4 transitions can be used to obtain further information about 

the metal ions local environment in addition to providing information on the light-emission 

properties of the Ln 3+ -containing materials. In the case of the Eu 3+ ion this important 

feature is even more prominent, since Eu 3+ is a powerful local ion probe due to its peculiar 

spectroscopic characteristics. These characteristics include: 28 

0 

5,



A large energy difference (gap) between the 0 

90 

5 D first excited state and the 

7 -1 

high-energy F 6 level of the fundamental septet (ca. 12300cm ). Given that 

for smaller gaps the desactivation through nonradiative process (e.g., O–H 

vibrations) is more likely to happen; 

A non-degenerated first excited state that allows a simpler Stark-effect 

analysis, with the subsequent correspondence between the observed J- 

splitting degeneracy and the Eu 3+ local-site symmetry; 

The presence of the ligand-field-independent 1 

transition, as discussed above; 

The presence of a single 0 

5 7 

D0 F magnetic-dipole 

5 7 

D0 F line permitted for Cs,C1,2,3,4,6, and C2v,4v,6v 

point symmetry groups with a predominantly electric-dipole nature, explained 

by J-mixing effects. 42,43,44 The energy of this non-degenerated transition can 

be directly related to the covalency of the chemical bonds of the first 

coordination shell in Eu 3+ (the so called nephelauxetic effect); 45 

The observation of vibronic lines in a relatively large spectral region of the 

5 7 

D0 F2 

excitation transition (24400–21550 cm -1 ) allows the identification of 

vibration modes related to the Ln 3+ local environment up to ca. 3000 cm -1 ; 

The presence of ligand-to-metal charge-transfer (LMCT) in the UV-vis region 

of the excitation spectra assigned to particular ligand groups. 

Important information about the metal-ion local coordination can be extracted from 

three distinctive parameters: i) the changes in the number of Stark components of each 

intra- f 

4 manifold and the variations of their relative intensity, ii) differences observed in 

the energy of particular lines, and iii) the analysis of the excited-state decay curves. 

Examples of the information extracted from these distinctive parameters are the existence 

of more than one Eu 3+ local symmetry group, the number of coordinated water molecules, 

the magnitude of the ligand field, and the importance of the covalency of the Eu 3+ -ligands 

bonds. 28 

The large values usually found for the full width at half maximum (fwhm) of the non- 

5 7 

degenerated D0 F0 

line typically range from ca. 20 to 30 cm -1 . 46 They indicate that the 

matrix where the Eu 3+ ions are accommodated has a large distribution of similar local



sites. In the case where the Eu 3+ local symmetry group does not present an inversion 

centre, the Laporte’s rule is relaxed due to odd parity terms in the ligand field 

Hamiltonian, 38 and the emission spectrum will be dominated by the 2 

For Eu 3+ -containing materials the number of water molecules ( w 

91 

5 7 

D0 F transition. 

n ) in the first coordination 

shell may be obtained by the difference between the decay time values measured in H2O 

and D2O as: 28 

nw 

1 1 

1. 11 

0. 

31 

 

H2O 

D2O 

 

where O 

and O 

H 2 

D 2 

(3.8) 

are the decay times in milliseconds in water and in D2O, 

respectively, taken in this latter case as the purely radiative decay. As mentioned above, 

the purely radiative decay may also be calculated through the experimental emission 

spectrum and therefore RAD 

can be used in Equation 3.8 substituting O 

value obtained experimentally for the decay EXP is used to substitute O 

H 2 

; while the 

D 2 

. In conclusion 

n w can be calculated for many different Eu 3+ -containing material within different medias. 

3.3. TISSUE OPTICAL PROPERTIES 

The optical properties of tissues are characterized by the use of light within the 

ultraviolet (UV) and the near-infrared (NIR) spectral region. The interaction of photons with 

tissue is based on three processes: i) absorption of light, ii) scattering of light and iii) 

emission of fluorescence. Tissue optical properties can be characterized by using these 

parameters separately. 47 As the penetration depth of light in living tissue strongly depends 

on the wavelength used, 48 the correct wavelength range should be chosen (since the 

amount of absorption in tissue is a function of the wavelength). Normally, the extent of 

scattering in tissue decreases with increasing wavelength. For wavelengths below 600 

nm, absorption dominates scattering resulting in a small penetration depth of hundredths 

of micrometers up to a few millimetres, so that only superficial assessment of tissues is 

possible in this spectral region.



When direct visual inspection or characterization is used through microscopic 

techniques high spatial resolution images of tissue structures is obtained. Some examples 

are the examination of tissue surfaces using optical fibres incorporated into endoscopes 

or laparoscopes, as well as of ocular diseases through ophthalmoscopes and the direct 

assessment of skin diseases or tissues during surgical procedures. 43 

Light absorption in tissue originates from oxy- and deoxyhemoglobin, other 

porphyrins, melanin and structures involved in cellular metabolism (such as NADH and 

flavins) as well as from several structural tissue components (such as collagen), elastin 

and lipo-pigments. Most of these compounds exhibit characteristic fluorescence spectra 

(tissue autofluorescence) throughout the visible (VIS) spectral region up to approximately 

700 nm. 49,50 Fluorescence presented by these fluorescent markers provides additional 

information on tissue structure and pathophysiological states. 49,51 This tissue 

autofluorescence phenomenon has already been thoroughly exploited to extract spectral 

patterns which indicate diseased tissue areas, e.g. in lymph node characterization; 52 

endoscopy of the gastrointestinal tract 50,53,54 or cardiovascular diagnosis. 55 

In order to image a large tissue volume light within the spectral range (700–900 nm) 

is necessary as the tissue absorption coefficient is relatively low (resulting in penetration 

depths up to a few centimetres (figure 3.3). 56 It is also possible to identify inhomogeneities 

as they exhibit a difference in absorption or fluorescence when compared to bulk tissue. 

Due to the scattering process photons do not follow straight paths when propagating 

through tissue and mathematical models of photon transport are needed to calculate the 

optical properties of tissue. 43 This process limits the spatial resolution of the obtained 

image, where morphological and structural tissue parameters are not accessible. Tissue 

absorption is mainly determined by oxyhaemoglobin, deoxyhaemoglobin and water, which 

exhibit a well defined absorption minimum in the NIR spectral region (figure 3.3) and 

provide information that can be utilized to quantitatively calculate important physiological 

parameters, such as blood concentration (total haemoglobin) and oxygenation (ratio 

oxy/deoxy haemoglobin). 57 The absorption data, together with tissue-dependent scattering 

properties, can be fitted by mathematical models to reconstruct the most probable photon 

propagation throughout the tissue in order to generate a spatial map of tissue optical 

properties for a given illumination and detection geometry. This method has predominantly 

92



been applied in the detection of breast tumours and imaging brain functions. Photon 

migration and diffuse light imaging are comprehensively described in reference. 58 

Figure 3.3. Absorption coefficients of oxyhemoglobin and deoxyhemoglobin as 

a function of wavelength. The near-infrared window of tissue is defined as the 

spectral region between approximately 700-900 nm, where the absorption 

coefficients are at minimal levels. 59 

3.4. OPTICAL IMAGING TECHNOLOGY 

Over the past decade different types of optical imaging techniques have been 

developed for biomedical applications. These include various microscopy methods such 

as confocal microscopy, two-photon microscopy and coherent anti-Stokes Raman 

scattering (CARS) microscopy for in vitro and ex vivo applications as well as several 

methods for in vivo applications such as bioluminescence imaging, fluorescence imaging, 

diffused optical tomography and optical coherence tomography. 43 Several of these 

techniques can be used in combination, either simultaneously or sequentially, to provide 

complementary information from the same cells, tissues, organs or animals. 60,61 

Both bioluminescence and fluorescence imaging techniques have found wide 

applications for in vivo tumor optical imaging in mouse models, providing a non-invasive 

method with high resolution and a convenient real-time, high frequency visualization and 

93



measurement of tumor biomarkers. The detection of disease progression and therapeutic 

response in the same animals is also possible. This minimizes the subject-to-subject 

variability and reduces the animal number required by a traditional method. 43 Most 

importantly, these in vivo studies can bridge the gap between in vitro (cell/tissue level) and 

in vivo (the whole animal) studies, while facilitating preclinical and further translational 

studies. 61 

Bioluminescence imaging (BLI) is typically based on the ATP- and O2-dependent 

enzymatic conversion of exogenous luciferin to oxyluciferin by luciferase within living cells. 

This reaction can produce photons with a broad yellow emission spectrum with a peak 

around 560 nm, detectable with a highly sensitive charge-coupled device (CCD) camera 

at 10-12 min after intraperitoneal injection of luciferin. The bioluminescence phenomena 

can last over 60 min in mice, providing multiple images of disease progression and 

therapeutic response based on the changes in the number of cells with luciferase 

expression or transcriptional activity. BLI has allowed quantitative measurements of tumor 

loadings, treatment response, immune cell trafficking and detection of gene transfer. It is 

also possible to obtain spatio-temporal information of whole biological systems in vivo 

within a short time frame that may accelerate the development of experimental therapeutic 

strategies. 61,62,63 

On the other hand, fluorescence imaging has an entirely different methodology that 

consists of exciting certain fluorophores in a living system by using external light and 

detecting fluorescence emission with a sensitive CCD camera. 

There are several types of fluorophores: i) endogenous molecules (such as collagen 

or hemoglobin), ii) exogenous fluorescent molecules such as green fluorescent protein 

(GFP) or iii) small synthetic optical contrast agents. When compared to in vitro 

fluorescence microscope, in vivo fluorescence imaging is a complex process affected by 

many factors. 43 One of the major limitations of in vivo fluorescence imaging is light 

attenuation and scattering due to adjacent living tissues. As previously mentioned light in 

the near infrared (NIR) range (650-900 nm) can improve the light penetration. 61,62,64 It can 

also minimize the autofluorescence of some endogenous absorbers such as hemoglobin, 

water and lipids. In a full-body mouse illumination experiment, photon counts in the NIR 

range (670 nm) are about four orders of magnitude higher when compared to those in the 

green light range (530 nm) under similar conditions. Hence NIR fluorescence imaging has 

proved to be an effective solution in improving the imaging depth along with sensitivity and 

specificity. As a result NIR-fluorophores are important for successful in vivo optical 

imaging and future clinical applications. 

94



3.5. OPTICAL CONTRAST AGENTS 

As mentioned in the previous chapter, the general definition of Contrast Agents 

(CAs) is a chemical substance introduced to the anatomical or functional region being 

imaged in order to increase the differences between different tissues, or between normal 

and abnormal tissue. 

Optical contrast agents are frequently used to stain portions of biological samples in 

order to obtain a greater understanding of the molecular, cellular and physiological 

changes at hand. Two important classes of these agents are the colorimetric and 

fluorescent markers. Colorimetric contrast agents modify the light absorption resulting in a 

contrast of observed colour; while fluorescent agents also modify light agent, fluorescent 

absorption, but subsequently re-emit a portion of the absorbed light at higher wavelengths 

(lower energy). 43 This phenomenon is leveraged to achieve greater sensitivity with the use 

of fluorescent markers. Through non illumination of the sample with light at the emission 

wavelengths of the fluorophore(s) present an enhanced signal to noise ratio will be 

achieved. In recent years, the attractiveness of fluorescent microscopy in biological 

sciences has grown; however, many challenges are still to be overcome in the application 

of fluorescent markers in biological systems. For instance, organic dyes are prone to rapid 

photobleaching limiting their application to long-term bioimaging investigation. 65,66 Many 

other dyes also have relatively broad emission spectra complicating their integration into 

multicolor imaging applications. In bioimaging, additional artifacts such as high light 

scattering via tissue interfaces, autofluorescence, and absorption by hemoglobin (Hb) in a 

mid-visible wavelength range are observed. The application of NIR range (650 to 900 nm) 

contrast agents can overcome many of these issues and is preferred for bio-imaging thick 

tissues. The development of NIR imaging agents such as NIR quantum dots (QDs) and 

dyes (e.g. cyanine dyes) have recently attracted significant attention. 67,68 In the case of 

optical imaging CA, they are normally molecules or nanoparticles bearing one or more 

fluorophores. Typical optical imaging CAs are injected intravascularly or intraperitoneally 

and allowed to accumulate at the target site over several hours to days. Fluorophores can 

be endogenous molecules (such as collagen or hemoglobin), exogenous fluorescent 

molecules such as green fluorescent protein (GFP) or small synthetic optical CAs. 

Within the last decades multifunctional contrast agents have emerged substantially 

with the attempt to obtain the benefits of multiple imaging modalities with the use of a 

95



single contrast agent. 69,70 One example can be found in the fluorescent and MRI active 

probes that allow for simultaneous detection by MRI and fluorescence microscopy by 

combining the advantages of 3D anatomical resolution at cellular level from MRI with the 

high sensitivity offered by fluorescence microscopy in a single particle. 

Organic dyes 

Organic dyes, such as rhodamine and fluoresceins are the earliest and most 

classical luminescent materials employed in optical agents. Nevertheless, these dyes 

have some problematic limitations that are critical to the quality of optical imaging. Broad 

emission bands (which can pose a problem in multiplexing and sensitivity), short 

luminescence lifetimes ca. 10 -9 s, a small Stokes shift (difficulty in separating the 

excitation and emission signals), poor photochemical stability and susceptibility to 

photobleaching (photochemical destruction of the fluorophore) on the 71,72 are the basic 

drawbacks of these type of agents. Several attempts have been made in order to create 

new fluorescent dyes that are more resistant to photobleaching and insensitive to pH 

changes. The problem of short lifetimes and Stokes shift has still not been satisfactorily 

overcome. Despite these setbacks organic dyes continue to be very popular due to their 

low cost, high availability and practicality. 72 

Quantum dots 

Quantum dots (QDs) are crystalline semiconductors typically less than 10 nm in 

diameter and are considered to be another promising luminescent tag for optical imaging. 

They are reported to be several thousand times more stable against photobleaching, 

much brighter and have a spectral width reduced by up to one-third from conventional 

organic dyes. 73,74 

One of the special features of QDs is its emission wavelengths variation that is 

strongly size dependent: They can emit different colours spanning the entire visible/NIR 

region within the same material by merely changing the size of the particles. This is a 

consequence of the so called quantum confinement effect: a confinement of the hole and 

electron wave functions in the nanocrystals when the size of the particles are reduced 

96



beyond the bulk-exciton Bohr radius, and consequently, an increase in the semiconductor 

band gap. 72-78 A combination of the broad absorption bands of QDs, narrow emission 

bands and their size dependent emissions offer the possibility of using them to tag 

biomolecules in ultrasensitive biological detection and in bioapplications such as gene- 

expression studies, high-throughput screening and medical diagnostics based on optical- 

coding technology. 75 Generally QDs have longer lifetimes (20- 50 ns) than that of organic 

dyes by one order of magnitude and so QDs are highly regarded as a new generation 

76,79 75,76,77,7879 

materials capable of bringing immense benefits to medical diagnosis. 

On the other hand, there are a few limitations that QDs have to overcome in order to 

be the ideal candidates for optical tags. An example is the danger of releasing toxic 

elements into biological systems, especially in in vivo analyses. The most optically 

efficient QDs are engineered with one or more toxic element such as cadmium and 

selenium. 80 CdSe QDs have luminescence properties from near UV to NIR regions with 

size-tunable absorption, this broad absorption band provide two advantages: i) freedom to 

select any excitation wavelength below the band gap energy and ii) minimize background 

by increasing Stokes shift. Another reported limitation of QDs is their tendency to suffer 

from optical blinking emission (continued on-off emission), a situation that is detrimental to 

72,81-84 81,82,83,84 

real time imaging 

As already mentioned QDs are more resistant to photobleaching than organic dyes, 

however some studies have reported various degrees of QDs fluorescence quenching 

under irradiation in ambient environment, where some of the authors 85,86 conclusively 

pointed out photo-oxidation of the nanocrystals as the main/primary cause. In order to 

overcome photo-oxidation limitation researchers attempted to coat the QDs with a 

protective layer (shell) forming a core-shell type QD, where the inner core would be a 

semiconductor material, while the shell would be another semiconductor but with a wider 

band gap 87 . Unfortunately this attempt did not reach expectations. Within the same 

perspective emerged the coating with thiols 88,89,90 , polymers 91,92,93 or silica 94,95 serving not 

only as protective layers but as platforms for possible bioconjugations introducing 

mutimodal prespective to QD. 

In conclusion, invaluable effort is being invested by many researchers with 

uncompromised intentions to resolve the toxicity of high-quality QDs and develop non- 

toxic QDs in order to offer all the advantages of QDs to biomedical imaging and 

therapeutic interventions. 

97



Lanthanide Chelates (and time resolved fluorescence) 

Considerable interest has been allotted to the application of kinetically stable 

lanthanide complexes as probes for biological imaging and assaying. 96,97 While efforts 

have been concentrated their on the development of gadolinium agents (as MRI contrast 

agents), 35,98 luminescent trivatent lanthanide ions such as terbium (Tb), europium (Eu), 

ytterbium (Yb) and neodymium (Nd) also have much to offer. 99,100 The large Stokes shifts 

and long-lived emission commonly associated with sensitised emission from these ions 

ensures the usage of time-gated spectroscopy to separate such long-lived signals from 

the short-lived signals that arise from scattered light and biological fluorescence. 101 As a 

result considerable interest has been given the use of luminescent lanthanides in 

bioassays and more recently the possibility of using time-resolved spectroscopy and 

microscopy. 103-105 

Lanthanide complexes can be tailor-made to exhibit different luminescent lifetimes 

by the exclusion of solvent molecules from the inner hydration sphere of the lanthanide 

ion. Therefore such complexes can be used in microscopy applications to gain more 

detailed information. Lanthanide complexes can be distinguished by their characteristic 

wavelengths of emission (due to the metal centre), very sharp emission bands (typically 

with a full width at half maximum of less than 10 nm) and their luminescent lifetimes (ca. 

10 -3 71 102103104105 

s), which are controlled by their coordination environment. 

Time-resolved fluorescence (TRF), which is more or less an exclusive preserve of 

lanthanides compounds, has become by far the most effective way of eliminating 

background interferences in order to enhance sensitivity. 102-105 The introduction of a time 

delay (~ 1 or 100 μs), prior to detection of the emitted light eliminates interference from 

light scattering and auto-fluorescence enhancing greatly the signal to noise ratio 

(detection limits are conservatively in the 10 -12 – 10 -15 M range) and hence the reliability of 

detection and monitoring. 102 The principle of TRF is illustrated in figure 3.4 and involves 

sequentially an excitation of the system, cessation of the excitation allowing enough decay 

time (during which the typical fluorescence from dyes and QDs would have decayed) for 

background interferences to have disappeared and measuring the luminescence of the 

lanthanide. 

98



Figure 3.4. Principle of TRF. The excitation is removed, followed by a 

sufficient time delay, and later measurement of lanthanide fluorescence and 

the cycle can continue thereafter with new excitation. 106 

An important point in the design of an efficient lanthanide-based probe is to protect 

the lanthanide cation from solvent molecules. Particular attention should be paid in 

protecting the cation from water as the lanthanide emission is quenched by the presence 

of water molecules in the first coordination sphere of the lanthanide (non-radiative 

deactivation through O-H vibrators). NIR emitting lanthanide ions commonly have a lower 

quantum yield than those other lanthanide ions emitting in the visible range, since their 

non-radiative transitions have higher probability with regards to the lanthanide ions 

emitting in the visible range.Nevertheless, as previously mentioned, they are more 

suitable for biological applications as they can be excited at lower energy and NIR 

photons can go deeper into tissue. 

Several research groups have devoted great effort to designing organic ligands, a 

great variety of which are now described in the literature and reviewed by Bünzili. 37 In 

general, small variations in a ligand structure can lead to remarkable changes in the 

photophysical properties. Several types of lanthanide carboxylates have been reported. 107 

99



Lanthanide doped nanoparticles 

Lanthanide doped NPs have emerged as a fast-growing platform in cell imaging due 

to the generation of low background noise for their NIR emission. To date the lanthanide 

oxide phosphor system has been the widely studied. Meiser et al. reported the synthesis 

of monodisperse fluorescent LaPO4 NPs .108,109,110 These NPs, approximately 7 nm in size, 

have fluorescence that originates from their bulk properties - transitions between d and f 

electron states and their local symmetry - and is independent of their size. These NPs 

were biofunctionalized via biotin-avidin chemistry with good photostability and fluorescent 

properties. 111 

Setua et al. produced highly monodispersed Eu 3+ and Gd 3+ doped Y2O3 nanocrystals 

and presented bi-modal imaging applications of both paramagnetism that enabled 

magnetic resonance imaging and bright red-fluorescence. This aided the optical imaging 

of cancer cells, targeted specifically to their molecular receptors. 112 

Lanthanide up-converting NPs (UCNPs) are rare-earth doped ceramic-type 

materials such as oxides, oxysulfides, fluorides or oxyfluorides which convert infra-red 

light into the visible spectrum. They are usually synthesized as nanospheres and were 

introduced as probes for bioassays in the 1990s. Most of these UCNPs contain Er 3+ ions 

as two-color (green, 540 nm; red, 654 nm) emitters and Yb 3+ ions as sensitizers, but other 

Ln 3+ pairs have also been proposed (e.g., Tm 3+ /Ho 3+ /Yb 3+ /Tm 3+ ). These type of NPs 

present several advantages over classical bioprobes, including high sensitivity, 

multiplexing ability if several different Ln 3+ ions are co-doped, low sensitivity to 

photobleaching and cheap laser diode excitation, in addition to deep penetration of the 

excitation NIR light. Initially, they were used in luminescent immunoassays. Presently their 

applications have been extended to luminescence imaging of cancerous cells. 113 Novel 

imaging systems based on UCNPs have been designed 114 and improved sensitivity of 

down to single molecule detection within cells is foreseen. 115116117 

Lanthanide up-converting NPs (UCNPs) have been developed as a new generation 

of luminescent labels due to their superb optical features, long lifetimes and excellent 

photostability. Among various kinds of UCNPs, NaYF4 is the most well-known system that 

has been employed in both cellular and in vivo animal imaging. 118,119,120 

Since the growth of this field is rapidly developing it is foreseen that lanthanide 

doped NPs will find their way into even more elaborate biotechnological applications in the 

coming future due to their relatively simple nanocomposition, deep penetration depth of 

100



NIR and other advantageous physical features. Due to the potentially high toxicity of 

lanthanides, more detailed investigations will probably be required to evaluate their 

biochemical and physiological behaviours before lanthanide doped NPs are eventually 

approved for biomedical applications. 

101




1 Bornhop D.J., Contag C.H., Licha K., Murphy C.J., Advances in Contrast Agents, 

Reporters, and Detection Biomed Opitcs 2001, 6, 106 – 110. 

2 Weissleder R., A Clearer Vision for in vivo Imaging Nature Biotechnol 2001 19, 316 – 317. 

3 Martinez Maestro, L.; Martin Rodriguez, E.; Vetrone, F.; Naccache, R.; Loro Ramirez, H.; 

Jaque, D.; Capobianco, J. A.; Garcia Sole, J., Nanoparticles for highly efficient multiphoton 

fluorescence bioimaging. Optics Express 2010, 18 (23), 23544-23553. 

4 Wang, F.; Banerjee, D.; Liu, Y.; Chen, X.; Liu, X., Upconversion nanoparticles in biological 

labeling, imaging, and therapy. Analyst 2010, 135 (8), 1839-1854. 

5 Park, Y. I.; Kim, J. H.; Lee, K. T.; Jeon, K.-S.; Bin Na, H.; Yu, J. H.; Kim, H. M.; Lee, N.; 

Choi, S. H.; Baik, S.-I.; Kim, H.; Park, S. P.; Park, B.-J.; Kim, Y. W.; Lee, S. H.; Yoon, S.-Y.; 

Song, I. C.; Moon, W. K.; Suh, Y. D.; Hyeon, T., Nonblinking and Nonbleaching 

Upconverting Nanoparticles as an Optical Imaging Nanoprobe and T1 Magnetic Resonance 

Imaging Contrast Agent. Advanced Materials 2009, 21 (44), 4467-4471. 

6 Scaff W.L., Dyer D. L., Mori, K., Fluorescent Europium Chelate Stain J. Bacteriol. 1969, 98, 

246-248. 

7 Soini E., Hemmilä I., Fluoroimmunoassay: Present Status and Key Problems Clin. Chem. 

1979, 25, 353 - 261. 

8 Hemmilä, I.; Ståhlberg, T.; Mottram, P. Bioanalytical Applications of Labelling Technologies, 

2nd ed.; Wallac Oy: Turku, Finland, 1995. 

9 Bazin H., Trinquet E., Mathis G., Time-resolved Amplification of Cryptate Emission: A 

Versatile Technology to Trace Biomolecular Interactions Rev. Mol. Biotechnol. 2002, 82, 

233 - 250. 

10 Hovinen J., Guy P.M. Bioconjugation with Stable Luminescent Lanthanide(III) Chelates 

Comprising Pyridine Subunits Bioconjugate Chem. 2009, 20, 404 - 421. 

11 Connally R.E., Piper J.A. Time-gated Luminescence Microscopy Ann. N. Y. Acad. Sci. 

2008, 1130, 106- 116. 

12 Moore E.G., Samuel A.P.S., Raymond K.N. From Antenna to Assay: Lessons Learned in 

Lanthanide Luminescence. Acc. Chem. Res. 2009, 42, 542 - 552. 

13 Bünzli J.-C.G., Lanthanide Luminescent Bioprobes (LLBs) Chem. Lett. 2009, 38, 104 - 109. 

14 Song B., Vandevyver C.D.B, Deiters E., Chauvin A.-S., Hemmilä l., Bünzli J.-C.G. A 

Versatile Method for Quantification of DNA and PCR Products Based on Time-resolved 

EuIII Luminescence. Analyst. 2008 133,1749 - 1756. 

15 Hess B.A., Kedziorski A., Smentek L., Bornhop D.J. Role of the Antenna in Tissue Selective 

Probes Built of Lanthanide-Organic Chelates. J. Phys. Chem. A 2008, 112, 2397 - 2407. 

102



16 Montgomery C.P., Murray B.S., New E.J., Pal R., Parker, D. Cell-penetrating Metal 

Complex Optical Probes: Targeted and Responsive Systems Based on Lanthanide 

Luminescence Acc. Chem. Res. 2009, 42, 925 - 937. 

17 Bünzli J.-C.G., Chauvin A.-S., Vandevyver C.D.B., Song B., Comby S. Lanthanide 

Bimetallic Helicates for in Vitro Imaging and Sensing Ann. N. Y. Acad. Sci. 2008, 1130, 97- 

105. 

18 Dos Santos C.M.G., Harte A.J., Quinn, S.J., Gunnlaugsson T. Recent Developments in the 

Field of Supramolecular Lanthanide Luminescent Sensors and Self-assemblies Coord. 

Chem. Rev. 2008, 252, 2512. 

19 Fan Y., Yang P., Huang S., Jiang J., Lian H., Lin J. Luminescent and Mesoporous 

Europium-Doped Bioactive Glasses (MBG) as a Drug Carrier J. Phys. Chem. C 2009, 113, 

7826 - 7830632. 

20 Harvey E.N., A History of Luminescence. From the Earliest Times Until 1900, American 

Philosophical Society, Philadelphia, 1957. 

21 Melhuish W.H., Nomenclature, Symbols, Units and their Usage in Spectrochemical Analysis 

VI: MOLECULAR LUMINESCENCE SPECTROSCOPY Pure Appl. Chem., 1984, 56, 231 - 

245. 

22 Blasse G., Grabmaier B.C., Luminescent Materials, Springer,Verlag Berlin Heidelberg, 

1994. 

23 Bünzli J.-C.G., Piguet C. Taking Advantage of Luminescent Lanthanide Ions Chem. Soc. 

Rev. 2005, 34, 1048 – 1077. 

24 Skoog D.A., Holler F.J., Crouch S.R. Principles of Instrumental Analysis, 6th ed.; Thomson 

Brooks/Cole: Canada, pp 154, 2007. 

25 Kaltsoyannis N., Scott P., The f Elements, Oxford Chemistry Primers, ed. J. Evans, Oxford 

Science Publications, Oxford, 1999. 

26 Cotton S., Lanthanides and Actinides, McMillan Physical Science Series, McMillan 

Education, London, 1991. 

27 Hüfner S., Optical Spectra of Transparent Rare Earth, Academic Press, New York 1978. 

28 Carlos L.D., Ferreira R.A.S., Zea Bermudez V., Ribeiro S.J.L. Lanthanide-Containing Light- 

Emitting Organic–Inorganic Hybrids: A Bet on the Future Adv. Mater. 2009, 21, 509 – 534 

29 Nishioka T., Fukui K., Matsumoto K. In Handbook on the Physics and Chemistry of Rare 

Earths; Gschneidner K.A., Bünzli J.-C.G., Pecharsky V.K., Eds.; Elsevier Science BV: 

Amsterdam, Vol. 37, Chapter 234, p 171 2007. 

30 Yuan, J.; Wang, G. Lanthanide-based luminescence probes and time-resolved 

luminescence bioassays Trends Anal. Chem. 2006, 25, 490 – 500. 

31 T. Jüstel, H. Nikol, C. Ronda, New Developments in the Field of Luminescent Materials for 

Lighting and Displays Angew. Chem. Int. Ed. 1998, 37, 3084 - 3103. 

103



32 Werts M.H.V. Making sense of lanthanide luminescence Sci. Prog. 2005, 88, 101 - 131. 

33 Hasegawa Y., Wada Y., Yanagida S., Strategies for the Design of Luminescent 

Lanthanide(III) Complexes and their Photonic Applications, J. Photochem. Photobiol. C 

2004, 5, 183 – 202. 

34 Hemmilä I., Laitala V. Progress in Lanthanides as Luminescent Probes. J. Fluoresc. 2005, 

15, 529 - 542. 

35 Parker D., Dickins R. S., Puschmann H., Crossland C., Howard J.A.K. Being Excited by 

Lanthanide Coordination Complexes: Aqua Species, Chirality, Excited-State Chemistry, and 

Exchange Dynamics Chem. Rev. 2002, 102, 1977 - 2010. 

36 Mulder W.J.M., Griffioen A.W., Strijkers G.J., Cormode D.P., Nicolay K., Fayad Z.A. 

Magnietic and Fluorescent Nanoparticles for Multimodality imaging Nanomedicine 2007, 2, 

307 - 324. 

37 Bünzli J.-C.G. Lanthanide Luminescence for Biomedical Analyses and Imaging Chem. Rev. 

2010, 110, 2729 - 2755. 

38 Malta O.L., Carlos L.D. Intensities of 4f-4f Transitions in Glass Materials Quim. Nova 2003, 

26, 889 - 895. 

39 S. Comby, J.-C. G. Bünzli, in Handbook on the Physics and Chemistry of Rare Earths (Eds: 

K. A. Gschneidner, Jr, J.-C. G. Bünzli, V. K. Pecharsky), Elsevier, New York, U.S.A., pp. 

217 – 254, 2007. 

40 Werts M.H.V., Jukes R.T.F., Verhoeven J.W. The emission spectrum and the radiative 

lifetime of Eu 3+ in luminescent lanthanide complexes Phys. Chem. Chem. Phys. 2002, 4, 

1542 - 1548. 

41 Lima P.P., Ferreira R.A.S., Freire R.O., Paz F.A.A., Fu L.S., Alves S., Carlos L.D., Malta 

O.L. Spectroscopic Study of a UV-Photostable Organic–Inorganic Hybrids Incorporating an 

Eu 3+ β-Diketonate Complex ChemPhysChem 2006, 7, 735 - 746. 

42 Porcher P., Caro P. Influence of J-mixing on the Phenomenological Interpretation of the 

Europium Ion Spectroscopic Properties J. Lumin. 1980, 21, 207 - 216. 

43 Malta O.L., Azevedo W.M., Gouveia E.A., Desa G.F. On the 5 D0→ 7 F0 Transition of the Eu 3+ 

Ion in the {(C4H9)4N}3Y(NCS)6 Host J. Lumin. 1982, 26, 337 - 343. 

44 Carlos L.D., A. Videira L.L. Emission Spectra and Local Symmetry of the Eu 3+ Ion in 

Polymer Electrolytes Phys. Rev. 1994, 49, 11721 - 11728. 

45 Carlos L.D., Malta O.L., Albuquerque R.Q. A Covalent Fraction Model for Lanthanide 

Compounds Chem. Phys. Lett. 2005, 416, 238 - 242. 

46 Carlos L.D., Ferreira R.A.S., Zea Bermudez V., Molina C., Bueno L.A., Ribeiro S.J.L. White 

Light Emission of Eu 3+ -Based Hybrid Xerogels Phys. Rev. B 1999, 60, 10042 - 10053. 

47 Licha, K., Topics in Current Chemistry - Contrast Agents for Optical Imaging. Springer- 

Verlag Berlin Heidelberg 2002, 222. 

104



48 Wan S., Parrish A.J., Anderson R.R., Madden M. Transmittance of Nanionizing Radiation in 

Human Tissues Photoochem Photobiol 1981, 34, 679 – 681. 

49 Andersson-Engels S., Wilson B.C. In vivo Fluorescence in Clinical Oncology: Fundamental 

and Practical Issues Cell Pharmacol 1992, 3, 48 – 61. 

50 Wagnieres G.A., Star W.M., Wilson B.C. In Vivo Fluorescence Spectroscopy and Imaging 

for Oncological Applications Photochem Photobiol 1998, 68, 603 – 632. 

51 Andersson-Engels S.K., linteberg C., Svanberg K., Svanberg S. In vivo Fluorescence 

Imaging for Tissue Diagnostics Phys Med Biol 1997, 42, 815 – 824. 

52 Moesta K.T., Ebert B., Handke T., Nolte D., Nowak C., Haensch W.E., Pandey R.K., 

Dougherty T.J., Rinneberg H., Schlag P.M. Protoporphyrin IX occurs naturally in colorectal 

cancers and their metastases Cancer Res. 2001, 61, 991 – 999. 

53 Stepp H., Sroka R., Baumgartner R. Fluorescence Endoscopy of Gastrointestinal Diseases: 

Basic Principles, Techniques, and Clinical Experience Endoscopy 1998, 30,379 – 386. 

54 Svanberg K., Wang I., Colleen S., Idvall I., Ingvar C., Rydell R., Jocham D., Diddens H., 

Bown S., Gregory G., Montan S., Andersson-Engels S., Svanverg S. Clinical Multi-colour 

Fluorescence Imaging of Malignant Tumours: Initial Experience Acta Radiol 1998, 39, 2 – 9. 

55 Maarek J-M.I., Marcu L., Fishbein M.C., Grundfest W.S. Time-resolved Fluorescence of 

Human Aortic Wall: Use for Improved Identification of Atherosclerotic Lesions Lasers Surg. 

Med. 2000, 27, 241 – 254. 

56 Grosenick D., Wabnitz H., Rinneberg H.H., Moesta K.T., Schlag P.M. Development of a 

Time-domain Optical Mammograph and First in vivo Applications Appl. Opt. 1999, 38, 2927 

– 2943. 

57 Sevick E.M., Chance B., Leigh J., Nioka S., Maris M. Quantitation of Time- and Frequency- 

resolved Optical Spectra for the Determination of Tissue Oxygenation Anal. Biochem. 1991, 

195, 330 – 351. 

58 De Haller E.B. Time-resolved Transillumination and Optical Tomography J. Biomed. Opitcs 

1996, 1,7 – 17. 

59 http://omlc.ogi.edu/spectra/hemoglobin 

60 Moriyama E.H., Zheng G., Wilson B.C. Optical Molecular Imaging: From Single Cell to 

Patient Clin Pharmacol Ther. 2008, 84, 267 – 271. 

61 Yunpeng Y., Xiaoyuan C. Integrin Targeting for Tumor Optical Imaging Theranostics 2011, 

1, 102 – 126. 

62 Contag C.H., Ross B.D. It's Not Just About Anatomy: In vivo Bioluminescence Imaging as 

an Eyepiece into Biology J. Magn. Reson. Imaging 2002, 16, 378 – 387. 

63 Shah K., Weissleder R. Molecular Optical Imaging: Applications Leading to the 

Development of Present Day Therapeutics. NeuroRx. 2005, 2, 215 – 225. 

105



64 Tung C.H. Fluorescent Peptide Probes for in vivo Diagnostic Imaging. Biopolymers 2004, 

76, 391 – 403. 

65 Chan W.C., Maxwell D.J., Gao X., Bailey R.E, Han M., Nie S. Luminescent Quantum Dots 

for Multiplexed Biological Detection and Imaging. Curr. Opin Biotechnol 2002, 13, 40 – 46. 

66 Bruchez M., Moronne M., Gin P., Weiss S., Alivisatos A.P. Semiconductor Nanocrystals as 

Fluorescent Biological Labels Science 1998, 281, 2013 – 2016. 

67 Licha K., Riefke B., Ntziachristos V., Becker A., Chance B., Semmler W. Hydrophilic 

Cyanine Dyes as Contrast Agents for Near-infrared Tumor imaging: Synthesis, 

Photophysical Properties and Spectroscopic in vivo Characterization Photochem Photobiol 

2000, 72, 392 – 398. 

68 Pham W., Lai W.F., Weissleder R., Tung C.H. High Efficiency Synthesis of a 

Bioconjugatable Near-infrared Fluorochrome Bioconjug Chem. 2003, 14, 1048 – 1105. 

69 Huber M.M., Staubli A.B., Kustedjo K., Gray M.H.B., Shih J., Fraser S.E., Jacobs R.E., 

Meade T.J. Fluorescently Detectable Magnetic Resonance Imaging Agents Bioconjug 

Chem 1998, 9, 242 – 249. 

70 Santra S., Bagwe R.P., Dutta D., Stanley J.T., Walter G.A., Tan W., Moudgil B.M., Mericle 

R.A. Synthesis and Characterization of Fluorescent, Radio-opaque, and Paramagnetic 

Silica Nanoparticles for Multimodal Bioimaging Applications Adv. Mater. 2005, 17, 2165 – 

2169. 

71 Pandya S., Yu J., Parker D. Engineering emissive europium and terbium complexes for 

molecular imaging and sensing Dalton. Trans. 2006, 2757 - 2766. 

72 Wang F., Tan W.B., Zhang Y., Fan X., Wang M. Luminescent nanomaterials for biological 

labelling Nanotechnology 2006, 17, R1 – R13. 

73 Rhyner M.N., Smith A. M., Gao X., Mao H., Yang L., Nie S Quantum Dots and 

Multifunctional Nanoparticles: New Contrast Agents for Tumor Imaging Nanomedicine 

2006, 1, 209 - 217. 

74 Tan W., Wang K., He X., Zhao X.J., Drake T., Wang L., Bagwe, R.P. Bionanotechnology 

Based on Silica Nanoparticles Med. Res. Rev., 2004, 24, 621 – 638. 

75 Tan W. B., Zhang Y. Multifunctional Quantum-Dot-Based Magnetic Chitosan Nanobeads 

Adv. Mat. 2005, 17, 2375 – 2380. 

76 Biju V., Itoh T., Ishikawa M. Delivering Quantum Dots to Cells: Bioconjugated Quantum 

Dots for Targeted and Nonspecific Extracellular and Intracellular Imaging Chem. Soc. Rev. 

2010, 39, 3031 – 3056. 

77 Trindade T., Pickett N.L., O’Brien P. Nanocrystalline Semiconductors: Synthesis, 

Properties, and Perspectives Chem. Mater. 2001, 13, 3843 - 3858. 

106



78 Sapra S., Sarma D.D., Sanvito S., Hill N.A. Influence of Quantum Confinement on the 

Electronic and Magnetic Properties of (Ga,Mn)As Diluted Magnetic Semiconductor Nano 

Lett. 2002, 2, 605 - 608. 

79 Smith A. M., Dave S., Nie S., True, L., Gao, X. Multicolor quantum dots for molecular 

diagnostics of cancer Expert Rev. Mol. Diagn., 2006, 6 , 231 - 244. 

80 Mericle R.A., Moudgil B.M., Tan W., Cao Z., Dutta D., Bertolino C., Liesenfeld B., Santra S. 

Fluorescence Lifetime Measurements to Determine the Core-shell Nanostructure of FITC 

Doped Silica Nanoparticles: An Optical Approach to Evaluate Nanoparticle Photostability J. 

Lumin., 2006, 117, 75 – 82. 

81 Soukka T., Kuningas K., Rantanen T., Haaslahti V., Lovgren T. Photochemical 

characterization of up-converting inorganic lanthanide phosphors as potential labels. J. 

Fluoresc. 2005, 15, 513 - 528. 

82 Yuan C.T., Yu P., Ko H.C., Huang J., Tang J. Antibunching Single-Photon Emission and 

Blinking Suppression of CdSe/ZnS Quantum Dots ACS Nano 2009, 3, 3051 – 3056. 

83 Matsumoto Y., Kanemoto R., Itoh T., Nakanishi S., Ishikawa M., Biju V. Photoluminescence 

Quenching and Intensity Fluctuations of CdSe-ZnS Quantum Dots on an Ag Nanoparticle 

Film J. Phys. Chem. C 2008, 112, 1345 – 1350. 

84 Mansur H. S. Quantum Dots and Nanocomposites WIREs Nanomedicine and 

Nanobiotechnology 2010, 2, 113 – 129. 

85 Roy A., Singha A. Quantitative Analysis of Thermal Stability of CdSe/CdS Core-shell 

Nanocrystals Under Infrared Radiation J. Mater. Res. 2006, 21, 1385 - 1389. 

86 Wang, P. N., Xu L., Yang W.L., Wang C.C., Guo J., Chen J.-Y, Ma J. Photostability of thiol- 

capped CdTe quantum dots in living cells: the effect of photo-oxidation Nanotechnology 

2006, 17, 2083 - 2089. 

87 Reiss P., Protiere M., Li L. Core/Shell Semiconductor Nanocrystals Small 2009, 5, 154 – 

168. 

88 Medintz I. L., Uyeda H. T., Goldman E. R., Mattoussi H. Quantum Dot Bioconjugates for 

Imaging, Labelling and Sensing Nat. Mater. 2005, 4, 435 – 446. 

89 Medintz I. L., Clapp A. R., Mattoussi H., Goldman E. R., Fisher B., Mauro J. M. Self- 

assembled Nanoscale Biosensors Based on Quantum Dot FRET Donors Nat. Mater. 2003, 

2, 630 – 638. 

90 Rosenthal S. J., McBride J., Pennycook S. J, Feldman L.C. Synthesis, Surface Studies, 

Composition and Structural Characterization of CdSe, Core/Shell, and Biologically Active 

Nanocrystals. Surf. Sci. Rep. 2007, 62, 111 – 157. 

91 Gao X.H., Cui Y.Y., Levenson R.M., Chung L.W.K., Nie S.M. In vivo Cancer Targeting and 

Imaging with Semiconductor Quantum Dots Nat. Biotechnol. 2004, 22, 969 – 976. 

107



92 Pellegrino T., Manna L., Kudera S., Liedl T., Koktysh D., Rogach A. L., Keller S., Rädler J., 

Natile G., Parak W. J. Hydrophobic Nanocrystals Coated with an Amphiphilic Polymer Shell: 

A General Route to Water Soluble Nanocrystals Nano Lett., 2004, 4, 703 – 707. 

93 Biju V., Kanemoto R., Matsumoto Y., Ishii S., Nakanishi S., Itoh T., Baba Y., Ishikawa M. 

Photo Induced Photoluminescence Variations of CdSe Quantum Dots in Polymer Solutions 

J. Phys. Chem. C 2007, 111, 7924 – 7932. 

94 Nann T., Mulvaney P. Single Quantum Dots in Spherical Silica Particles Angew. Chem., Int. 

Ed. 2004, 43, 5393 – 5396. 

95 Selvan S. T., Tan T. T., Ying J. Y. Robust, Non-Cytotoxic, Silica-Coated CdSe Quantum 

Dots with Efficient Photoluminescence Adv. Mater. 2005, 17, 1620 – 1625. 

96 Toth E., Helm L., Merbach A. E. in Comprehensive Coordination Chemistry II, ed. M. D. 

Ward, Elsevier, Oxford, UK, pp. 841–882, 2004. 

97 Faulkner S., Matthews J.L. in Comprehensive CoordinationChemistry II, ed. M. D. Ward, 

Elsevier, Oxford, pp .913–944, UK, 2004. 

98 Caravan P., Ellison J. J, McMurry T., Lauffer R. B. Gadolinium(III) Chelates as MRI 

Contrast Agents: Structure, Dynamics, and Applications Chem. Rev. 1999, 99, 2293 – 

2352. 

99 Selvin, P.R. Principles and Biophysical Applications of Lanthanide-based Probes Annu. 

Rev. Biophys. Biomol. Struct. 2002, 31, 275 - 302. 

100 Klink S. I., Keizer H. Van Veggel F.C.J.M. Organo-d-Metal Complexes as New Class of 

Photosensitizers for Near-Infrared Lanthanide Emission, Angew. Chem., Int.Ed., 2000, 39, 

4319 – 4321. 

101 Alpha B., Lehn J.M., Mathis G. Energy Transfer Luminescence of Europium(III) and 

Terbium(III) Cryptates of Macrobicyclic Polypyridine Ligands Angew. Chem., Int. Ed. Engl., 

1987, 26, 266 - 167. 

102 Beeby A., Botchway S. W., Clarkson I. M., Faulkner S., Parker A. W., Parker D., Williams J. 

A. G. Luminescence Imaging Microscopy and Lifetime Mapping Using Kinetically Stable 

Lanthanide(III) Complexes J. Photochem. Photobiol. B: Biol. 2000, 57, 83 – 89. 

103 Charbonnière L. J., Ziessel R., Montalti M., Prodi L., Zaccheroni N., Boehme C. Wipff G. 

Luminescent Lanthanide Complexes of a Bis-bipyridine-phosphine-oxide Ligand as Tools 

for Anion Detection J. Am. Chem. Soc., 2002, 124, 7779 – 7788. 

104 Charbonnière L. J., Weibel N., Estournes C., Leuvrey C., Ziessel R. Spatial and Temporal 

Discrimination of Silica Particles Functionalised with Luminescent Lanthanide Markers 

Using Time-resolved Luminescence Microscopy New J. Chem. 2004, 28, 777 – 781. 

105 Weibel N., Charbonniere L.J., Guardigli M., Roda A., Ziessel R. Engineering of Highly 

Luminescent Lanthanide Tags Suitable for Protein Labeling and Time-Resolved 

Luminescence Imaging J. Am. Chem. Soc., 2004, 126, 4888 – 4896. 

108



106 http://www.biotekinstruments.ru/ru/resources/articles/time-resolved-fluorescent- 

compounds.html 

107 Parker D., Williams J.A.G. Responsive Luminescent Lanthanide Complexes. In Sigel A, 

Sigel H, deitors. Metal Ions in Biological Systems, Vol. 40. The Lanthanides and their 

Interactions with Biosystems. New York, NY: Marcel Dekker; 2003. 

108 Meyssamy H., Riwotzki K., Kornowski A., Naused S., Haase M., Wet-Chemical Synthesis of 

Doped Colloidal Nanomaterials: Particles and Fibers of LaPO4:Eu, LaPO4:Ce and 

LaPO4:Ce,Tb Adv. Mater. 1999, 11, 840 - 844. 

109 Riwotzki K., Meyssamy H., Kornowski A., Haase M. Liquid-Phase Synthesis of Doped 

Nanomaterials: Colloids of Luminescing LaPO4:Eu, CePO4:Tb 5 nm-Particles with Narrow 

Particle Size Distribution J. Phys. Chem. B 2000, 104, 2824 - 2828. 

110 Riwotzki K., Meyssamy H., Schnablegger H., Kornowski A., Haase M. Liquid-Phase 

Synthesis of Doped Nanomaterials: Colloids and Redispersible Powders of Strongly 

Luminescing LaPO4:Ce, Tb Nanocrystals Angew. Chem. 2001, 40, 573. 

111 Meiser F., Cortez C., Caruso F. Biofunctionalization of Fluorescent Rare-Earth-doped 

Lanthanum Phosphate Colloidal Nanoparticles Angewandte Chemie 2004, 116, 6080 – 

6083. 

112 Setua S., Menon D., Asok A., Nair S., Koyakutty M. Folate Receptor Targeted, Rare-earth 

Oxide Nanocrystals for Bi-modal Fluorescence and Magnetic Imaging of Cancer Cells 

Biomaterials 2010, 31, 714 – 729. 

113 Wang M., Mi C., Zhang Y., Liu J., Li F., Mao C., Xu S. NIR-Responsive Silica-Coated 

NaYbF4:Er/Tm/Ho Upconversion Fluorescent Nanoparticles with Tunable Emission Colors 

and Their Applications in Immunolabeling and Fluorescent Imaging of Cancer Cells J. Phys. 

Chem. C 2009, 113, 19021 - 19027. 

114 Soga, K. Development of NIR fluorescence bioimaging system through polyscale 

technology Bunseki Kagaku 2009, 58, 461 - 471. 

115 Jiang, S.; Zhang, Y.; Lim, K. M.; Sim, E. K. W.; Ye, L., NIR-to-visible upconversion 

nanoparticles for fluorescent labeling and targeted delivery of siRNA. Nanotechnology 

2009, 20 (15). 

116 Zhou, J.; Sun, Y.; Du, X.; Xiong, L.; Hu, H.; Li, F., Dual-modality in vivo imaging using rare- 

earth nanocrystals with near-infrared to near-infrared (NIR-to-NIR) upconversion 

luminescence and magnetic resonance properties. Biomaterials 2010, 31 (12), 3287-3295. 

117 Wang, M.; Mi, C.-C.; Wang, W.-X.; Liu, C.-H.; Wu, Y.-F.; Xu, Z.-R.; Mao, C.-B.; Xu, S.-K., 

Immunolabeling and NIR-Excited Fluorescent Imaging of HeLa Cells by Using 

NaYF(4):Yb,Er Upconversion Nanoparticles. Acs Nano 2009, 3 (6), 1580-1586. 

118 Li Z., Zhang Y., Jiang S. Multicolor Core/shell-structured Upconversion Fluorescent 

Nanoparticles Adv. Mat. 2008, 20, 4765 – 4769. 

109



119 Park Y. I., Kim J. H., Kim J. H. Nonblinking and Nonbleaching Upconverting Nanoparticles 

as an Optical Imaging Nanoprobe and T1 Magnetic resonance Imaging Contrast Agent Adv. 

Mat. 2009, 21, 4467 – 4471. 

120 Xiong L.-Q., Chen Z.-G., Yu M.-X., Li F.-Y., Liu C., Huang C.-H. Synthesis, 

Characterization, and in vivo Targeted Imaging of Amine-functionalized Rare-earth Up- 

converting Nanophosphors Biomaterials 2009, 30, 5592 – 5600. 

110


Lanthanide-chelate Grafted Silica Nanoparticles as Bimodal-Imaging Contrast Agents 

111 

4. 

Lanthanide-Chelate Grafted 

Silica Nanoparticles as 

Bimodal-Imaging Contrast 

Agents



4.1. Introduction 113 



4.3.1 Characterization of Nanoparticles 121 

4.3.2 Photoluminescence Properties 131 

4.3.3 Relaxivity Properties 151 

4.3.4 Cell Imaging 159 



112



113 

Chapter published as original article: 

Pinho S.L.C., Faneca H., Geraldes C.F.G.C., Delville M-H. Carlos L.D., Rocha J. 

Lanthanide-DTPA Grafted Silica Nanoparticles as Bimodal-Imaging Contrast Agents 

Publication in Biomaterials (2011), doi:10.1016/j.biomaterials.2011.09.086 

Pinho S.L.C., Faneca H., Geraldes C.F.G.C., Delville M-H. Carlos L.D., Rocha J Silica 

Nanoparticles for Bimodal MRI-Optical Imaging via Grafting of Gd 3+ and Eu 3+ /Tb 3+ 

Complexes 

4.1. INTRODUCTION 

Submitted to the European Journal of Inorganic Chemistry 

Currently, clinical diagnostics and biomedical research employ an array of 

powerful in vivo imaging techniques, including confocal and Two-Photon Microscopy, 

Magnetic Resonance Imaging (MRI), 1 X-Ray Computed Tomography (CT), 2 Positron 

Emission Tomography (PET), 3,4 Single Photon Emission Computed Tomography 

(SPECT), 5,6 and Ultrasound. 7 Each of these techniques possesses unique strengths 

and weaknesses (spatial and temporal resolution and sensitivity limits), thus providing 

complementary information. Certain fused-modality instruments, such as PET/CT, have 

already appeared in the clinic. 6 

MRI has an excellent spatial resolution but suffers from low sensitivity, often 

requiring the administration of millimolar concentrations of commercial Gd 3+ -based 

contrast agents (CAs), in order to increase the intrinsic image contrast for an efficient 

detection of pathologies. 8-12 Radioactive tracers and optical imaging probes are orders 

of magnitude more sensitive and may be detected at much lower concentrations 

(picomolar or micromolar for PET or optical agents, respectively) but the corresponding 

13,14 891011121314 

imaging modalities have low spatial resolution. 

There has been an increasing interest in the development of multimodal imaging 

agents, integrating in a single molecular entity the requirements of MRI and a second 

imaging modality. Bimodal MRI and optical imaging probes combine the spatial 

resolution and unlimited tissue penetration of MRI with the sensitivity of optical imaging. 

The efficiency of this combined imaging technique has been demonstrated in studies



with animals. Modo et al. used a gadolinium-rhodamine-dextran agent to confirm by 

fluorescence microscopy that tracking of transplanted stem cells in ischemia-damaged 

rat hippocampus was possible by MRI. 15 Recently, several examples of bimodal agents 

have been synthesised and evaluated, including Gd 3+ complexes connected to organic 

dyes, 16-18 complexes of Gd 3+ and other visible 19-24 or near infrared (NIR) 25,26 emitting 

Ln 3+ ions, or various kinds of nanoparticle (NP) based systems. 17-34 

16171819202122232425262728293031323334 

Lanthanide ions are particularly well suited for the design of bimodal MRI and 

optical agents. 35 Their unique electronic configuration affords exceptional magnetic and 

optical properties and similar chemical behaviour. Therefore, the replacement of one 

lanthanide by another, results in compounds with different physical properties but no 

major chemical differences. The advantage of using Gd 3+ complexes as MRI contrast 

agents has been largely demonstrated. 8 The Gd 3+ ion possesses seven unpaired 

electrons (highest spin density) and a symmetrical 8 S ground state, resulting in a slow 

electronic relaxation rate, and these are excellent features for reducing the longitudinal 

T ) and transverse ( 2 

( 1 

T ) proton relaxation times of tissue water, thus enhancing image 

contrast. For example, [Gd(DTPA)(H2O)] 2- has been approved for radiologic practice 

and medicine in 1988. 36 

All Ln 3+ ions, exception for La 3+ and Lu 3+ , are photoluminescent, some more 

efficient than others. Eu 3+ and Tb 3+ are the most commonly ions used because they 

emit in the visible spectrum (in the red and green regions, respectively) and have long 

luminescence lifetimes, in the millisecond range). 14,37,38 There are several advantages 

in using lanthanide complexes as luminescent probes versus organic dyes: i) 

resistance to photobleaching; ii) long-lived excited states, allowing the short-lived (ns 

range) biological background fluorescence to disperse before the lanthanide emission 

occurs; iii) absence of reabsorption; and iv) sharp emission bands (wavelengths are 

characteristic of the lanthanide). 25 Despite these positive features of lanthanides, 

reports on the use of Ln 3+ complexes in the design of combined MRI and optical probes 

are scarce. 19-23 The luminescence properties of the lanthanides ions may be enhanced 

by intramolecular energy transfer from moieties attached to the central ion, the so- 

called “antenna effect”. It is well known that the introduction of an aromatic ligand 

induces a considerable antenna effect, therefore enhancing the luminescence 

properties of the lanthanide ions. 

Our approach is to combine the potential of the lanthanide complexes with the 

properties of NPs since these have the ability to i) carry large payloads of active 

114



magnetic centres, therefore lowering the required concentrations, ii) be target-specific 

by labelling desired cells through phagocytic pathways, and iii) be grafted by molecules 

specific to cell surface markers. Many alternative designs of efficient nanosized carriers 

for MRI probes have been proposed. 39,40 Mesoporous silica nanoparticles (MSNs) have 

been shown to be very useful platforms for efficient relaxometric contrast agents 

because of their ability to carry a large payload of Gd 3+ chelates with high water 

accessibility and, thus, they have been used as multimodal probes after incorporation 

of a fluorescent dye into the silica carrier. 41-46 Core-shell hybrid nanoporous silica NPs 

containing a luminescent [Ru(bpy)3]Cl2 core (bpy = 2,2’bypyridine) and a paramagnetic 

monolayer coating of a silylated Gd(III) complex has also been studied. 30,47,48 

41,42,43,44,45,464748 

As a proof of concept, here report on the derivatization of nanoporous silica NPs 

with aminopropyltriethoxysilane (APS), followed by reaction with diethylenetriamine 

pentaacetic acid (DTPA) and complexation of Ln 3+ ions, forming the DTPA monoamide 

system SiO2@APS/DTPA:Ln (Ln = Eu 3+ , Tb 3+ and Gd 3+ ). The thermodynamic stability 

constant of the Gd 3+ -DTPA complex is quite high, with a log K value of 22.46, which is 

very similar to the values for the Eu 3+ and Tb 3+ complexes. 49 It has been previously 

shown that, although the thermodynamic stability constants for Gd-DTPA monoamides 

decrease by log K ~2.6 relative to the Gd-DTPA ones, their blood pH conditional 

constants differ from Gd-DTPA only by log K ~1.2. 50 SiO2@APS/DTPA:Ln nanoparticles 

are not toxic 51 and similar materials accumulate mostly in the liver and spleen whereas 

the lung, kidney, and heart accounted for an accumulation of less than 5%. 52 The 

luminescence and water proton nuclear relaxation properties of these NPs both, in 

aqueous suspensions, and internalized in RAW 264.7 cells (mouse macrophage cell 

line), are studied in order to evaluate their usefulness as bimodal agents for MRI and 

optical imaging. 

We also report the synthesis and grafting of nanoporous silica NPs with APS and 

a Ln 3+ complex with a ligand possessing an aromatic ring, which acts as an antenna for 

sensitizing Ln 3+ , thus improving light emission. Silica NPs are, thus, modified with 

2,2',2",2'" – [(pyridine-2,6-diyl)bis(methylenenitrilo)] tetrakis (acetic acid) (PMN), via a 

reaction with APS grafted on the NPs. Ln 3+ are complexed by PMN, forming the system 

SiO2@APS/PMN:Ln (Ln = Eu, Tb and Gd). The water proton nuclear relaxation and 

luminescence properties of such derivatized NPs in aqueous suspensions and 

internalized in RAW 264.7 cells (mouse macrophage cell line) cells are studied, in 

115



order to assess their efficacy as Ln 3+ -based bimodal agents for MRI and optical 

imaging. 

4.2. EXPERIMENTAL PROCEDURES 

Materials and purification methods 

EuCl3 (99.99%), TbCl3 (99.99%), GdCl3 (99.99%), Tetraethoxysilane (TEOS) 

(98%), 3-aminopropyltrimethoxysilane (APS) (97%), diethylenetriamine pentaacetic bis- 

anhydride (DTPAA) (99.99%), 2,6-bis(bromomethyl) pyridine (98%), di(tert-butyl) 

iminobis(acetate) (98%), dry Na2CO3 (99.5%), dry acetonitrile (MeCN) (99.8%), 

trifluoroacetic acid (CF3COOH) (99%) and diethyl ether (Et2O) (99.7%); N- 

hydroxysuccinimide (NHS) (98%), (1-Ethyl-3-[3-dimethyl aminopropyl]carbodiimide 

Hydrochloride (EDC) (99%) and Phosphate buffered saline (PBS) were purchased from 

Aldrich. Absolute ethanol (J.T. Baker) and ammonium hydroxide solution (5N) (Fluka) 

were used as received. All other reagents were of analytical grade. Water was 

deionized (resistivity larger than 18 MQ). 

Preparation of silica nanoparticles suspension 

The method used was derived from the so-called Stöber 48,53 process, widely used 

for the synthesis of silica beads with diameters from a few tens to a few hundreds of 

nanometers 54 and based on the hydrolysis/condensation of tetraethoxysilane (TEOS) 

catalyzed by ammonia in alcoholic media. Briefly, a solution of 250 mL of absolute 

ethanol and 17 mL ammonia was heated at 50 ºC and 0.035 mol TEOS was added 

allowing reflux overnight. The average particle size, determined by transmission 

electron microscopy was 67 ± 6 nm. 

Preparation of the grafted amino-nanoparticles 

To the silica NPs suspension 4.5 mmol of APS (0.8 mL) were added and stirred 

for 3 h. The suspension was then left under reflux overnight. The nanoparticles were 

116



then washed and purified by centrifugation three times with ethanol and then water to 

remove unreacted APS. 

Preparation of the DTPA-grafted amino-nanoparticles 

The amino-modified NPs suspension was centrifuged and the supernatant 

discarded, the wet NPs were slowly diluted in 20 mL of an ethanol-acetic acid solution 

50/50 v/v % and 1 g of diethylene triaminepentaacetic bis-anhydride (DTPAA) was then 

slowly added to the solution (20/1 DTPAA/-NH2) at room temperature. The system was 

heated up and left to reflux overnight. The particles were filtered off and purified three 

times by centrifugation-redispersion in acetone-water 50/50 v/v % and finally three 

times in water. 

Preparation of the Ln 3+ - DTPA chelate-grafted amino-nanoparticles 

The DTPA-grafted NPs were redispersed in 20 mL of water. At room 

temperature, 0.3 mmol of LnCl3 (0.11g for GdCl3) were slowly added to the colloidal 

suspension. This amount corresponds approximately to the quantity of molecules 

grafted onto the particles in the 20 mL of solution, assuming a coverage rate of 6 

µmol/m². After 24 h, the excess of unreacted Ln(III) was removed by centrifugation- 

redispersion three times in water. 

Preparation of the 2,2',2",2'"–[(Pyridine-2,6-diyl)bis methylene nitrilo)] 

tetrakis(acetic acid) compound (PMN) 

First the tetra(tert-butyl)2.2',2",2"'-[(Pyridine-2,6-diyl)bis(methyleneni trilo)] tetrakis 

(acetate) compound (1) was synthesized. A mixture of 2,6-bis(bromomethyl)pyridine 

(318 mg, 1.20 mmol, solid), di(tert-butyl) iminobis(acetate) (588 mg, 2.40 mmol, solid), 

dry Na2CO3 (630 mg, 6.0 mmol, solid) and dry MeCN (30 ml) was refluxed overnight 

under argon. Filtration (obtaining a yellow solution) and evaporation gave pure 

compound (yellow gel) (412 mg, 72%). 1 H-NMR (CDCl3) δ (ppm): 1.44 (s, 36 H); 3.47 

117



(s, 8 H); 4.00 (s, 4 H); 7.48 (d, J = 7.5, 2 H); 7.63 (T, J = 75, 1 H). A solution of 

compound (1) (412 mg, 0.86 mmol) in CF3COOH (10 ml) was stirred for 2 h at RT. 

After evaporation, the mixture was triturated with Et2O and filtered to give pure 

compound (dark yellow gel) (288 mg, 92%). 1 H-NMR (DMSO) δ (ppm): 3.62 (s, 8 H); 

4.24 (s,4 H); 7.75 (d, J = 7.5, 2 H); 8.23 (t, J = 7.5, 1 H). 

Preparation of the PMN-grafted amino-nanoparticles 

The amino-modified NPs suspension was centrifuged and the supernatant 

discarded, the wet NPs were slowly diluted in 10 mL of PBS. While 0.5 g of PMN (1.35 

x 10 -3 mol) were added to 15 mL of PBS, a yellow solution was formed. A solution of 

EDC was also prepared by adding 258.9 mg (1.35 x 10 -3 mol) in 4 mL of PBS. To the 

PMN solution 15.54 mg of NHS (1.35 x 10 -4 mol) was added and stirred at room 

temperature. Sequentially the EDC and NPs solutions were added, respectively. The 

system was left to react overnight. The particles were filtered off and purified three 

times by centrifugation-redispersion in water. 

Preparation of the Ln 3+ -PMN chelate-grafted amino-nano-particles 

The PMN-grafted NPs were redispersed in 20 mL of water. At room temperature, 

33.3 μmol of LnCl3 (8.78 mg for GdCl3) were slowly added to the colloidal suspension. 

This amount corresponds approximately to the quantity of molecules grafted onto the 

particles in the 10 mL of solution, assuming a coverage rate of 6 µmol/m². After 24 h, 

the excess of unreacted Ln(III) was removed by centrifugation-redispersion three times 

in water. 

Cell culture and in vivo imaging 

RAW 264.7 cells (mouse macrophage cell line) were maintained at 37 °C, under 

5% CO2, in Dulbecco's modified Eagle's medium-high glucose (DMEM-HG) (Sigma) 

supplemented with 10% (v/v) heat-inactivated fetal bovine serum (FBS) (Sigma), 

penicillin (100 U/mL) and streptomycin (100 μg/mL), and sodium bicarbonate (1.6 g/L). 

Cells were incubated with the respective NPs (2.5 x 10 15 Part/L for both types of NPs 

SAD:Ln and SAPMN:Ln ) at 37 °C, under 5% CO2, for 1 hour. After this incubation, 

cells were washed with PBS, fixed with 4% paraformaldehyde, for 15 min at room 

118



temperature, and rinsed again with PBS. Then, cells were detached from the culture 

flasks by scraping, the cell suspensions were prepared in PBS and the cell pellets were 

obtained by centrifugation at 180g during 5 min. 1 

119 

T -weighted MRI images of the cellular 

pellets were acquired on a 3.0 T Siemens TimTrio scanner, using a spin--echo 

sequence (TE = 12 ms, TR = 3000 ms, FOV= 100x 100, slice thickness = 3.00 mm, 

matrix = 128 x 256 at room temperature for SAD:Ln NPs and TE = 50 ms, TR = 3000 

ms, FOV= 100x 100, slice thickness = 3.00 mm, matrix = 128 x 256 at room 

temperature for SAPMN:Ln NPs). The optical images were obtained by submitting the 

cellular pellets to a 450 W Xe arc lamp, as the excitation source and photographs were 

taken with a Canon 550D with EF-S 18-55mm. 

Particles Characterization 

TEM was performed at room temperature on a JEOL JEM-2000 FX transmission 

electron microscope using an accelerating voltage of 200 kV. Drops of diluted 

dispersions of nanoparticles were air-dried on carbon films deposited on 200-mesh 

copper grids. The excess liquid was blotted with filter paper. Diffuse Reflectance 

Infrared Fourier-Transform (DRIFT) spectra were recorded on a Bruker IFS Equinox 

55FTIR spectrometer (signal averaging 64 scans at a resolution of 4 cm -1 in KBr pellets 

containing ca. 2 mass % of material). The zeta potential of the nanoparticles was 

measured using a Zetasizer 3000HSA setup (Malvern Instruments) equipped with a 

He- Ne laser (50 mW, 532 nm). The zeta potential measurement based on laser 

Doppler interferometry was used to measure the electrophoretic mobility of 

nanoparticles. Measurements were performed for 20 s using a standard capillary 

electrophoresis cell. The dielectric constant was set to 80.4 and the Smoluchowsky 

constant f(ka) was 1.5. The silica, europium, terbium, and gadolinium contents were 

measured by inductively coupled plasma / optical emission spectrometry ICP/OES 

(ES720, Varian) equipped with a crossflow nebulizer. Solutions for each element with a 

concentration of 1 g/L were used to prepare the standard solutions (SCP Science to 

Paris) and were used as internal standard to evaluate the instrumental drift. 

The 29 Si magic-angle spinning (MAS) nuclear magnetic resonance (NMR), 29 Si 

cross-polarization (CP) MAS NMR and 13 C CP/MAS NMR spectra were recorded on a 

Bruker Avance III 400 (9.4 T) spectrometer at 79.49 and 100.62 MHz, respectively. 29 Si 

MAS NMR spectra were recorded with 2 μs (tip angle ca. 30º) rf pulses, a recycle delay 

of 60 s and 5.0 kHz spinning rate. 13 C CP/MAS NMR spectra were recorded with 4 μs



1 H 90º pulses, 2 ms contact time, a recycle delay of 4 s and at a spinning rate of 8 kHz. 

1 H NMR spectra were recorded on a Bruker Avance-300 spectrometer at 300.13 MHz, 

using CDCl3 as the solvent. Chemical shifts are quoted in ppm from tetramethylsilane 

(TMS) and coupling constants J in Hz. Multiplicities are described with abbreviations as 

follow: s (singlet), d (doublet), t (triplet), q (quadruplet), m (multiplet). Spectra are 

described as δ (multiplicity, number of protons, assignment, coupling constant). 

1 H longitudinal and transverse relaxation times ( 1 

120 

T and 2 

T respectively) of 

aqueous suspensions of nanoparticles were measured at 20 MHz on a Bruker 

Minispec mq20 relaxometer and at 499.83 MHz (B0= 11.7 T) on a Varian Unity 500 

NMR spectrometer, at 25 and 37 ºC. T 1 values were measured using the inversion 

recovery pulse sequence, while T 2 values were measured using a Carr-Purcell- 

Meiboom-Gill (CPMG) pulse sequence. The time interval between two consecutive 

refocusing pulses (τCP) in the train of 180 o pulses applied was 1.6 ms. The values of 

* 

T 2 , the transverse relaxation time in the presence local field inhomogeneities, were 

obtained from the water spectral line widths. All the experimental values were corrected 

for the diamagnetic contributions using aqueous suspensions of their respective NPs 

carrier SiO2@APS/DTPA (SAD) and SiO2@APS/PMN (SA/PMN) under the same 

conditions. The r2 values were also measured as a function of the CP parameter in a 

CPMG pulse sequence, for aqueous suspensions of the various NPs (CP = 0.05, 0.2, 

0.4, 0.8, 1.6, 2, 3). 

The photoluminescence spectra were recorded between 14 K and room 

temperature with a modular double grating excitation spectrofluorimeter with a TRIAX 

320 emission monochromator (Fluorolog-3, Jobin Yvon-Spex) coupled to a R928 

Hamamatsu photomultiplier, using the front face acquisition mode. The excitation 

source was a 450 W Xe arc lamp. The emission spectra were corrected for detection 

and optical spectral response of the spectrofluorimeter and the excitation spectra were 

weighed for the spectral distribution of the lamp intensity using a photodiode reference 

detector. The lifetime measurements were acquired with the setup described for the 

luminescence spectra using a pulsed Xe-Hg lamp (6 μs pulse at half width and 20–30 

μs tail). The absolute emission quantum yields were measured at room temperature 

using a Quantum Yield Measurement System C9920-02 from Hamamatsu with a 150 

W Xenon lamp coupled to a monochromator for wavelength discrimination, an 

integrating sphere as sample chamber and a multi channel analyzer for signal 

detection.



4.3. RESULTS AND DISCUSSION 

4.3.1. Characterization of Nanoparticles 

Aqueous suspensions of silica NPs were synthesized by basic polymerization of 

silane monomers under Stöber conditions, 48 using an alcohol-water-ammonia medium 

and tetraethoxysilane (TEOS) as the silane precursor. TEM (Figure 4.1) reveals 

spherical, essentially monodispersed, particles with an average size of 67 ± 6 nm (100 

particles measured). 

Figure 4.1. a) TEM images of the silica NPs; b) histogram depicting the 

experimental size distribution of the NPs and the corresponding calculated 

normal cumulative distribution for the specified mean and standard 

deviation. 

100 nm 

These NPs were successively functionalized with APS, DTPA or PMN) and a 

lanthanide salt (Ln=Eu, Tb or Gd). The zeta potential titrations as a function of pH 

confirm the shift of the stability ranges of the NPs and the isoelectric points (IEPs) with 

the different modification steps (Figure 4.2). 

Population % 

After APS coating the suspension exhibited an IEP of ca.10 characteristic of the 

presence of free amino groups on the NPs surface. 55-57 Once the peptidic coupling 

(carboxyl groups from the DTPA or PMN with the free amino groups of the APS) was 

achieved, another clear shift of the IEP towards lower pH (IEP of ca. 6 for DTPA and 

ca. 5.5 for PMN) was observed. Upon coordination of the DTPA or PMN with Eu 3+ , 

Tb 3+ , Gd 3+ (ions with similar association constants), 58,59 there was no major change of 

55-59 56575859 

the IEP, which does not depend on the lanthanide chemical nature. 

30 

25 

20 

15 

10 

5 

0 

121 

[55 -60[ 

[60 -65[ 

[65 -70[ 

[70 -75[ 

Diameter (nm) 

[75 -80[ 

[80 -85[ 

[85 -90[



Zeta potential (mV) 

80 

60 

40 

20 

0 

-20 

-40 

-60 

-80 

2 4 6 8 10 12 14 

Figure 4.2. Zeta potential titrations as a function of pH for SiO2 (S, ●), 

SiO2@APS (●), SiO2@APS/DTPA (●), SiO2@APS/PMN (●), 

122 

pH 

SiO2@APS/DTPA:Ln (●) and SiO2@APS/PMN:Ln (●) 

The amount of Ln 3+ ions present in the NPs strongly depends on the amount of 

DTPA (or PMN) grafted. Therefore, the same single set of DTPA or PMN grafted NPs 

was used throughout the study, presenting a ratio of ca. 10 4 ions per NP. As an 

example, Table 4.1 and Table 4.2 summarize the ICP data obtained on each sample 

for the quantification of Gd, Eu, Tb and Si. 

Table 4.1. Elemental composition of samples SiO2@APS/DTPA:Ln 

(SAD:Ln) ascertained by ICP 

[Gd] (M) [Eu] (M) [Tb] (M) [Si] (M) 

[Gd] 

(Ions/NP) 

[Eu] 

(Ions/NP) 

SAD:Eu,Gd 1.50 × 10 -3 1.48 × 10 -3 0.820 9.59 × 10 3 9.46 × 10 3 

[Tb] 

(Ions/NP) 

SAD:Tb,Gd 2.66 × 10 -3 3.15 × 10 -3 1.840 5.73 × 10 3 6.78 × 10 3 

SAD:Eu,Tb 1.41 × 10 -3 1.24 × 10 -3 0.790 7.07 × 10 3 6.22 × 10 3



Table 4.2. Elemental composition of samples SiO2@APS/PMN:Ln 

(SAPMN:Ln) ascertained by ICP 

[Gd] (M) [Eu] (M) [Tb] (M) [Si] (M) 

123 

[Gd] 

(Ions/NP) 

[Eu] 

(Ions/NP) 

SA/PMN:Eu 2.85 × 10 -3 0.162 7.0 × 10 4 

[Tb] 

(Ions/NP) 

SA/PMN:Tb 3.37 × 10 -3 0.171 7.8 × 10 4 

SA/PMN:Gd 3.33 × 10 -3 0.160 8.2 × 10 4 

SAP/MN:Eu,Gd 1.55 × 10 -3 1.62 × 10 -3 0.173 3.6 × 10 4 3.7 × 10 4 

SAP/MN:Tb,Gd 1.71 × 10 -3 1.57 × 10 -3 0.156 4.3 × 10 4 4.0 × 10 4 

SAP/MN:Eu,Tb 1.71 × 10 -3 1.47 × 10 -3 0.157 4.3 × 10 4 3.7 × 10 4 

DRIFT spectroscopy was also used to probe the effectiveness of the chemical 

modification for the SAD:Ln NPs 48,51 (Figure 4.3). The absorption bands in the DRIFT 

spectrum of the SiO2@APS samples (Figure 4.3) are assigned to APS and bare silica, 

evidencing the efficient APS-silanization of the silica NPs. The modification of the 

shape of the band at 3500 cm -1 is due to the N-H vibrations. The 2983 cm -1 and 2908 

cm -1 bands, assigned to the asymmetric and symmetric stretching vibrations of CH2 

groups of the grafted alkyl chain, confirm the anchoring of the amino propyl groups, as 

well as the presence of the 1631 cm -1 and 1486 cm -1 (the free and hydrogen bonding 

NH2, respectively), 800 cm -1 (N-H bending mode), and 1448 cm -1 (SiCH2) bands. 

The following step was the chemical attachment of DTPA to the aminated NPs by 

the formation of an amide linkage between the primary amine group and one of the five 

carbonyl groups of DTPA. The occurrence of this chemical grafting was confirmed by 

the DRIFT spectrum with the formation of four absorption bands: one band 

characteristic of the secondary amide C=O stretch at 1685 cm -1 , two due to 

asymmetrically coupled vibration (1631 cm -1 ) and symmetrically coupled stretching 

(1396 cm -1 ) of the two C-O bonds of the carboxylate anion, and the fourth band is 

characteristic of the C-N vibration stretch at 937 cm -1 (Figure 4.3). 

Evidence for the chelation of Ln 3+ (Ln= Gd, Eu, Tb) is also forthcoming from 

DRIFT spectroscopy. 49 Band assignments of the asymmetric and symmetric stretching 

vibrations of CH2 groups of the grafted alkyl chain (2977 cm -1 , 2935 cm -1 and 2900 cm -



1 ) and the 1415 cm -1 (hydrogen bonding NH2), 804 cm -1 (N-H bending mode), and 1448 

cm -1 (SiCH2) were identified. The complexation of the Ln 3+ induced a slight 

bathochromic shift of the two CO absorption bands, from 1631 to 1579 cm -1 and 1396 

to 1415 cm -1 respectively, and the secondary amide (C=O) 1680 to 1685 cm -1 . Finally, 

the C-N vibrational stretch remained unchanged at 937 cm -1 . 

Reflectance 

1800 1700 1600 1500 1400 

3700 3400 3100 2800 2500 2200 1900 1600 1300 1000 700 400 

Wavenumber (cm -1 ) 

Figure 4.3. Diffuse Reflectance IR Fourier-Transform spectra (DRIFT) of (a) 

silica NPs (blue line), SiO2@APS (red line) and SiO2@APS/DTPA (green 

line). The inset depicts a selected region of the spectra. 

Regarding SiO2@APS/PMN:Ln, DRIFT spectroscopy provided evidence for the 

grafting of APS to the SiO2 surface 60 but no confirmation for the reaction between APS 

and PMN. Information on the latter was, however, forthcoming from solid-state NMR. 

The 13 C CP/MAS NMR spectra of the SiO2, SiO2@APS and SiO2@APS/DTPA 

NPs are shown in Figure 4.4 (see also Table 4.3). The spectrum of the SiO2 NPs 

reveals the presence of small amounts of ethanol (peaks at ca. 19 and 58 ppm). The 

spectrum of SiO2@APS displays three broad resonances, due to the Si-bonded propyl 

chains, at ca. 42 (C3), 22 (C2) and 9 ppm (C1). 61 Finally, the spectrum of 

SiO2@APS/DTPA contains the three APS peaks as well as broad resonances in the 

ranges 165-180 (carboxylate groups) and 50-65 ppm (remaining carbons) assigned to 

DTPA. Peaks from impurities are also present at ca. 25 ppm (acetic acid), 18 ppm 

(ethanol) and 30 ppm (acetone). 

124



The 29 Si CP/MAS and MAS NMR spectra of the derivatized NPs are shown in 

Figure 4.4B and C, respectively. The former exhibit resonances at ca. -92, -101 and - 

111 ppm, ascribed to Q n (4-n)(OH) local environments ( 29 Si linked to n 29 Si atoms via 

bridging O), respectively Q 2 (2OH) (such as geminal silanols), Q 3 (OH) (single silanols) 

and Q 4 (siloxane). 62 The faint and broad resonance observed at ca. -66 ppm is 

ascribed to the organosiloxane (T 3 ) atoms R’Si(OSi)3, 51 providing evidence for the 

chemical bonding of APS to the surface of the silica NPs. The presence of T 2 

environments cannot be disregarded because the peak centred at -66 ppm is 

asymmetric, and may contain an unresolved resonance at ca. -60 ppm. 63 We note that 

several different T 3 sites are possible in the aminosilane layer, as discussed, for 

example, in Albert et al., 63 (Sheme 4.1 is, thus, just illustrative). Further evidence for 

such coupling is forthcoming from both the 29 Si CP/MAS and MAS NMR spectra 

because the number of NPs surface hydroxyl groups decreases upon derivatization 

with APS: the population ratio (Q 2 +Q 3 )/Q 4 (measured from the MAS NMR spectrum) 

decreases from 0.43 to 0.37. 

Upon reaction of SiO2@APS with DTPA, the number of silanol groups rises 

again: (Q 2 +Q 3 )/Q 4 increases from 0.37 to 0.62. This is evidence for strong interactions 

between the amino group and silica surface silanols, when immobilized on silica gel via 

hydrogen-bonding interactions and formation of a five-membered cyclic intermediate. 

These results confirm previous reports, which provided evidence for a cyclic structure 

of the aminosilane layer based on models of five- or six- membered rings in which the 

nitrogen atoms interact with either the Si atom or one of the SiOH groups. 64,65 The 

existence of six-member ring models, containing either a SiO -…. NH + H(R) or a 

SiOH …. NH(R) bonding structure was also assumed using XPS results 66,67 and FTIR and 

Raman spectroscopy. 68 

The intensity of the T 3 (+T 2 ) CP/MAS peak decreases upon addition of DTPA 

(although this must be taken with caution because the CP/MAS spectra are not a priori 

quantitative). These results indicate that the reaction of SiO2@APS with DPTA has the 

side effect of also modifying somewhat the SiO2 NPs surface. 

125



DTPA DTPA DTPA 

-50 -50 -50 -60 -60 -60 -70 -70 -70 -80 -80 -80 

AA 

A 

SiO SiO SiO22 2@APS/DTPA @APS/DTPA 

@APS/DTPA 

SiO SiO SiO22 2@APS @APS @APS 

SiO SiO SiO22 2 

BB B CC 

C 

TT T33 3 

Q 2 

Q 2 

Q 2 

Q 3 

Q 3 

Q 3 

-60 -60 -60 -70 -70 -70 -90 -90 -90 -100 -100 -100 -110 -110 -110 -120 -120 -120 -130 -130 -130 

(ppm) (ppm) (ppm) 

Q 3 

Q 3 

Q 3 

200 200 200 100 100 100 50 50 50 00 

0 

Q 4 

Q 4 

Q 4 

126 

(ppm) (ppm) (ppm) 

Q 2 

Q 2 

Q 2 

* * * 

Q 3 

Q 3 

Q 3 

3' 3' 3' 

33 

3 

Q 4 

Q 4 

Q 4 

-80 -80 -80 -90 -90 -90 -100 -100 -100 -110 -110 -110 -120 -120 -120 -130 -130 -130 

d d d (ppm) (ppm) (ppm) 

 

 

 

 

2' 2' 2' * * * 1' 1' 1' 

22 

2 

* * * 

11 

1 

SiO SiO SiO22 2@APS/DTPA @APS/DTPA 

@APS/DTPA 

SiO SiO SiO22 2@APS @APS @APS 

SiO SiO SiO22 2 

Figure 4.4. 13 C CP/MAS NMR spectra of the SiO2 (blue), SiO2@APS (red), 

SiO2@APS/DTPA (green) and DTPA (black). (* Ethanol, ▲ Acetic acid and 

♦ Acetone) B) 29 Si CP/MAS NMR spectra of the SiO2 (blue), SiO2@APS 

(red) and SiO2@APS/DTPA (green) C) 29 Si MAS NMR spectra of the SiO2 

(blue), SiO2@APS (red) and SiO2@APS/DTPA (green). The inset in B 

shows an expansion of the T region, exhibiting an asymmetric peak; the 

arrow depicts the region of a possible resonance from T 2 environments.



Scheme 4.1. Representation of a SiO2-NPs functionalized with APS and 

coupled with DTPA. 

13 C 

Table 4.3. 13 C CP/MAS and 29 Si MAS NMR chemical shifts for SiO2, 

SiO2@APS and SiO2@APS/DTPA NPs and quantification of the 29 Si Q n 

resonances. 

29 Si SiO2 SiO2@APS SiO2@APS/DTPA 

Assignment δ/ppm Assignment δ/ppm δ/ppm δ/ppm 

C 1 9 Q 2 -92.7 (2.8%) -92.8 (3.0%) -91.0 (4.6%) 

C 1’ 9 Q 3 -102.1 (27.1%) -100.9 (23.9%) -100.3 (33.7%) 

C 2 23 Q 4 -111.5 (70.1%) -110.4 (73.1%) -109.6 (61.7%) 

C 2’ 22 

C 3 42 

C 3’ 42 

DTPA ~170 – 180 T 3 -65.4 -66.4 

DTPA ~58 – 62 

DTPA ~50 

DTPA ~55 

127



In the case of the SiO2@APS/PMN, the 13 C CP/MAS NMR spectra (Figure 4.5a) 

of SiO2 and SiO2@APS/PMN exhibit two main peaks at 17.3 and 57.4-57.8 ppm, 

attributed to remains of ethanol solvent. The spectrum of SiO2@APS displays three 

broad resonances, given by the APS Si-bonded propyl chains, at ca. 42 (C3), 22 (C2) 

and 9 ppm (C1). 60,61 Although fainter, these resonances are also present in the 

spectrum of SiO2@APS/PMN, thus showing that the sample contains APS (albeit less). 

The 29 Si CP/MAS and MAS NMR spectra of the derivatized NPs are shown in Figures 

4.5b and 4.5c, respectively. The former exhibit resonances at ca. - 92, - 101 and - 111 

ppm, ascribed to Q n (4-n)(OH) local environments ( 29 Si linked to n 29 Si atoms via 

bridging O), respectively Q 2 (2OH) (such as geminal silanols), Q 3 (OH) (single silanols) 

and Q 4 (siloxane). 62 

The faint resonance observed at ca. -67 ppm in the 29 Si CP/MAS and MAS NMR 

spectra of SiO2@APS and SiO2@APS/PMN is ascribed to the organosiloxane (T 3 ) 

atoms R’Si(OSi)3, 51 providing evidence for the chemical bonding of APS to the surface 

of the silica NPs. Since the peak centred at - 67 ppm is asymmetric and may contain 

an unresolved resonance at ca. - 61 ppm therefore a T 2 environments cannot be 

disregarded. 63 As already proposed, 60 both the 29 Si CP/MAS and MAS NMR spectra: 

when the number of NPs surface hydroxyl groups decreases upon derivatization with 

APS, the population ratio (Q 2 + Q 3 )/Q 4 (measured from the MAS NMR spectrum) 

decreases from 0.43 to 0.37 (Table 4.3). 

128



a 

SiO 2 @APS/PMN 

SiO 2 @APS 

SiO 2 

b 

-50 -55 -60 -65 -70 -75 -80 

* 

3 

80 70 60 50 40 30 20 10 0 

T 3 Q 2 

Q 3 

-60 -70 -90 -100 -110 -120 -130 

(ppm) 

Q 4 

(ppm) 

129 

c 

Q 2 

2 

Q 3 

Q 4 

-80 -90 -100 -110 -120 -130 

(ppm) 

* 

1 

SiO2@APS/PMN 

SiO2@APS 

Figure 4.5. (a) 13 C CP/MAS NMR spectra of the SiO2 (blue), SiO2@APS 

(red) and SiO2@APS/PMN (green). (* Ethanol) (b) 29 Si CP/MAS and (c) 29 Si 

MAS NMR spectra of SiO2 (blue), SiO2@APS (red) and SiO2@APS/PMN 

(green). The inset in (b) expansion of the T region, exhibiting an asymmetric 

peak; the arrow depicts the region of a possible resonance from T 2 

environments. 

SiO2



Table 4.3. 13 C CP/MAS and 29 Si MAS NMR chemical shifts for SiO2, 

SiO2@APS and SiO2@APS/PMN NPs and quantification of the 29 Si Q n 

resonances. 

29 Si SiO2 SiO2@APS SiO2@APS/DTPA 

Assignment δ/ppm δ/ppm δ/ppm 

Q 2 -92.7 (2.8%) -92.8 (3.0%) -92.3 (3.5%) 

Q 3 -102.1 (27.1%) -100.9 (23.9%) -100.4 (30.2%) 

Q 4 -111.5 (70.1%) -110.4 (73.1%) -109.3 (66.3%) 

T 3 -65.4 -66.4 

Upon coupling of PMN to SiO2@APS, the number of silanol groups increases 

again: (Q 2 + Q 3 )/Q 4 from 0.37 to 0.51. This suggests that some APS is removed from 

the NPs surface, in accordance with the observation that the T 3 29 Si CP/MAS peak 

decreases concomitantly. It may also indicate strong interactions between the amino 

groups and silica surface silanols. As already mentioned possible hydrogen-bonding 

interactions and formation of a five-membered cyclic intermediate may occur when the 

amino groups are immobilized on the silica gel. 

Scheme 4.2. Representation of a SiO2-NPs functionalized with APS and 

coupled with PMN. 

130



4.3.2. Photoluminescence Properties 

SiO2@APS/DTPA:Ln NPs 

The emission (steady-state and time-resolved) and excitation properties of the 

solids and samples in suspension were investigated. Figure 4.6 a displays the 300 K 

emission spectra of SiO2@APS/DTPA:Eu in the solid state excited at three different 

wavelengths. No energy shifts are observed for any transition when the wavelength is 

varied, indicating a single local environment for the Eu 3+ ions. This conclusion is also 

valid for suspensions of NPs, for which the only differences relative to the solid-state 

spectrum are the relative intensities of the intra-4f Stark components (Figure 4.7). The 

spectra comprise a series of sharp lines, assigned to the Eu 3+ 5 D0→ 7 F0–4 transitions, 

and a strong broad band between 380 and 560 nm, ascribed to the emission of the 

SiO2@APS/DTPA host. Figure 4.8 shows the emission spectra (300 K) of 

SiO2@APS/DTPA recorded with different excitation wavelengths. The spectra consist 

of two strong Gaussian-shape broad bands, at 280 - 320 nm and 320 - 600 nm, whose 

maximum shifts to the red as the excitation wavelength increases. The excitation 

spectra were monitored along with the hybrid host's emission (inset of Figure 4.8). 

They consist of two broad bands, between 240 and 300 nm and between 300 and 430 

nm whose maximum shifts to the red 

69-72 69707172 

131



Normalized Intensity (arb. units) 

exc 280 nm 

exc 360 nm 

exc 393 nm 

5 D0 7 F 0 

5 D0 7 F 1 

5 D0 7 F 2 

380 465 550 635 720 

7 F0 5 F 4 

7 F0 5 H 3 

7 F0 5 D 4 

7 F0 5 G 2-5 

7 F0 5 L 6 

250 300 350 400 450 500 550 

Wavelength (nm) 

7 F0 5 D 2 

5 D0 7 F 3 

em 420 nm 

em 614 nm 

132 

5 D0 7 F 4 

7 F0 5 D 1 

a 

b 

c 5 D0 7 F 0 

d 5 

D0 7 575 577 579 581 


F1 581 584 587 590 593 596 599 602 


e 

5 D0 7 F 2 

603 608 613 618 623 628 633 


Figure 4.6. (a) Emission spectra (300 K) of SiO2@APS/DTPA:Eu (solid state) 

excited at 280 (black), 360 (red) and 393 nm (green); (b) Excitation spectra (300 

K) of SiO2@APS/DTPA:Eu (solid state) monitored at 420 (magenta) and 614 nm 

(blue); (c) (d), and (e) show a magnification of the 5 D0 7 F0-2 transitions. 


a) 

b) 

c) 

exc 270 nm 

exc 288 nm 

exc 317 nm 

exc 317 nm 

exc 393 nm 

exc 393 nm 

5 D0 7 F 0 

5 D0 7 F 1 

5 D0 7 F 2 

380 465 550 635 720 


Figure 4.7. Emission spectra (300 K) of the SiO2@APS/DTPA:Eu in water 

solution (blue), solid state at 300K (red) excited at different wavelengths. 

5 D0 7 F 3 

5 D0 7 F 4



Intensity (arb. units) 

285 370 455 540 625 710 


133 

245 290 335 380 425 

Figure 4.8. Emission spectra (300 K) of the SiO2@APS/DTPA excited at 

270 (black), 300 (red), 330 (green), 350 (blue), 360 (cyan), 375 (magenta) 

and 400 nm (yellow). The inset shows the excitation spectra (300 K) 

monitored at 350 (black), 405 (red), 430 (green) and 450 nm (blue). 

To shed more light onto the origin of this broad band, Figure 4.9 compares the 

emission spectra of the host in SiO2, SiO2@APS, SiO2@APS/DTPA and 

SiO2@APS/DTPA:Eu. In accord with previous results, 68 the Gaussian-shape broad 

band shifts to the blue with the addition of APS, from 440 (SiO2) to 430 nm 

(SiO2@APS, SiO2@APS/DTPA and SiO2@APS/DTPA:Eu). The full-width-at-half- 

maximum (fwhm) decreases from 133.8 (SiO2) to 109.0 (SiO2@APS), 99.7 

(SiO2@APS/DTPA) and 84.8 nm (SiO2@APS/DTPA:Eu). The SiO2 and the SiO2@APS 

emission spectra are in agreement with spectra reported for analogous materials and 

are ascribed to oxygen defects in the silica skeleton. 69,70 It should be noted that no 

calcination was used, whereas previous works report luminescence properties only 

after calcination. 69,70




SiO2 ex 360 nm 

SiO2@APS ex 360 nm 

SiO2@APS@DTPA ex 360 nm 

SiO2@APS@DTPA@Eu ex 355 nm 

375 460 545 630 715 


Figure 4.9. Emission spectra (300 K) of the SiO2 (black), SiO2@APS (red), 

SiO2@APS/DTPA (green) and SiO2@APS/DTPA:Eu (blue) excited at 360 

nm. 

The SiO2@APS/DTPA:Eu emission spectra exhibit a series of sharp lines 

ascribed to the Eu 3+ 5 D0→ 7 F0–4 intra-4f 6 transitions upon 280 and 360 nm excitation 

(host excited states, Figures 4.6 and 4.10) providing clear evidence for the energy 

transfer from the host to the Eu 3+ ion. The comparison between the emission spectra of 

DTPA:Eu and SiO2@APS/DTPA:Eu, displayed in Figure 4.11 (in particular the energy 

and fwhm of the 5 D0 7 F0 line and the energy and relative intensities of the 7 F1-4 Stark 

components), indicates an effective interaction between the Eu 3+ ions and the 

SiO2@APS/DTPA host, completely different from that observed for the DTPA:Eu 

complex. 

134




250 290 330 370 410 450 490 530 570 


135 

em 616 nm Solid (300K) 

em 616 nm Solid ( 12K) 

Figure 4.10. Excitation spectra of the SiO2@APS/DTPA:Eu monitored at 

616 nm at different temperatures 300K (red) and 12K (blue). 

Normailzed Intensity (arb. units) 

5 D0 7 F 0 

5 D0 7 F 1 

5 D0 7 F 2 

5 D0 7 F 4 

570 620 670 720 

240 300 360 420 480 540 600 

5 D0 7 F 3 


DTPA_Eu exc 393.5 nm 

SiO2@APS@DTPA_Eu exc 393 nm 

DTPA_Eu em 614 nm 

SiO2@APS@DTPA_Eu em 616 nm 

Figure 4.11. Top: Emission spectra (300 K) of the DTPA:Eu (black) and 

SiO2@APS/DTPA:Eu (red) excited at 393.5 nm and 393 nm, respectively; 

Bottom: Excitation spectra (300 K) of the DTPA:Eu (black) and 

SiO2@APS/DTPA:Eu (red) monitored at 614 nm and 616 nm, respectively.



The excitation spectra (300 K) of the same SiO2@APS/DTPA:Eu system 

monitored at 420 nm (magenta) and 614 nm (blue), (Figure 4.6b), show two strong 

broad bands at 275 and 340 nm overlapping with a series of sharp lines ascribed to the 

Eu 3+ intra- 4f 6 transitions between the 7 F0 and the 5 L6, 5 D4,2,1, 5 F4, 5 H3, 5 G2-5 levels. 

Lowering the temperature from 300 to 14 K, the relative intensity of the hybrid host 

bands increases and a new band appears at 330 nm (Figure 4.10). This temperature 

dependence supports the assignment of this new excitation band to a ligand-to-metal 

charge transfer (LMCT) transition, resulting from the interaction between the host and 

the Eu 3+ ions. 73 

In order to get further insight into the Eu 3+ local coordination, the 4f 6 emission 

lines were recorded at high resolution (Figure 4.6c-e). The detection of a single 

5 D0 7 F0 line (17289.0±1.5 cm -1 ) and the J-degeneracy splitting of the 7 F1,2 levels into 

three Stark components, observed over the entire range of excitation wavelengths 

used, indicate that the Eu 3+ cations reside in a single low-symmetry site. The larger 

intensity of the electric-dipole 5 D0 7 F2 transition, relative to the intensity of the 

magnetic-dipole 5 D0 7 F1 transition, indicates the absence of an inversion centre for the 

Eu 3+ site. 

The calculated value of the 5 D0→ 7 F0 fwhm, 59 cm -1 , is much larger than the 

values (20-30 cm -1 ) 74 reported for other organic-inorganic hybrids, suggesting for the 

Eu 3+ ions a large distribution of similar local sites. The room-temperature 5 D0 emission 

decay curve, monitored within the 5 D0→ 7 F2 transition at 614 nm and excited at 393 nm, 

is well fitted by a single exponential function, yielding a 5 D0 lifetime of 0.35±0.02 ms 

(Figure 4.12). The slight deviation from a mono-exponential of the decay curve is in 

agreement with the aforementioned large distribution of the Eu 3+ ions in similar local 

sites. Therefore, a decrease in the lifetime relative to the non-grafted monomeric 

complex [Eu(DTPA)] 2- (structure already well characterized and studied 75 ) is observed 

due to changes in the local environment of the Eu 3+ ion. 

136




0 1 2 3 4 5 

137 

Time (ms) 

Figure 4.12. Emission decay curve of Eu 3+ excited states ( 5 D0) from the 

sample SiO2@APS/DTPA:Eu, which were monitored at 614 nm and excited 

393 nm (blue), at 300 K in the solid state. 

The experimental Ω2 and Ω4 intensity parameters were determined from the 

emission spectra shown in Figure 4.6 using the 5 D0→ 7 F2 and 5 D0→ 7 F4 electronic 

transitions, respectively, and expressing the emission intensity (I) in terms of the 

surface (S) under the emission curve as equation 3.5. 

The branching ratio for the 5 D0→ 7 F5,6 transitions must be neglected as they are 

not observed experimentally and their influence on the depopulation of the 5 D0 excited 

state may be ignored, and the Ω6 parameter is not determined. The 5 D0 7 F1 transition 

does not depend on the local ligand field and may be used as a reference for the whole 

spectrum. An effective refractive index of 1.5 was used leading to A01 ≈ 50 s -1 . 76 The 

radiative emission rate is given by as equation 3.6: 74,77,78,79 

The 5 D0 radiative (Ar) and non–radiative (Anr) transition probabilities were 

determined for sample SiO2@APS/DTPA:Eu and are 0.3116 ms -1 and 3.0217 ms -1 , 

respectively. The quantum efficiency (η) [ A A A 

the emission spectrum and the 5 D0 lifetime ( nr 

] was estimated based on 

r 

T 

r 

r 

nr 

1 

A A A ) as η=0.09. This small 

value is essentially due to the high Anr value. The emission absolute quantum yield ( ) 

was measured and found less than 0.01. The Judd-Ofelt intensity parameters (Ω2,4) 

were 1.10 x 10 -20 cm 2 and 1.02 x 10 -20 cm 2 , respectively.



The photoluminescence characterization of the SiO2@APS/DTPA:Tb sample was 

also carried out and gave similar results Figures 4.13, 4.14, 4.15, and 4.16). Figure 

4.13a displays the emission spectra at 300 K of the SiO2@APS/DTPA:Tb in the solid 

state excited at 285 nm (black) and 377 nm (red). The spectra consist of a large broad 

band between 380 and 520 nm, ascribed to the emission of the SiO2@APS/DTPA host, 

as for the Europium system, and of a series of straight lines assigned to the Tb 3+ 

5 D4→ 7 F6–0 intra-4f 8 transitions. No energy shifts are observed for any transition when 

the wavelength is varied, indicating that the Tb 3+ ion is in a single local environment. 

This conclusion is also valid in solution at 300 K, where the only difference detected 

relatively to the spectrum recorded in the solid state is in the relative intensities of the 

intra-4f Stark components (see Figure 4.14). 


exc 285 nm 

exc 377 nm 

380 465 550 635 720 

7 F6 5 I 8 , 5 F 4,5 , 5 H 4 

7 F6 5 H 5,6 

7 F6 5 H 7 , 5 D 1 

7 F6 5 L 7,8 

7 F6 5 G 3 

5 D4 7 F 6 

7 F6 5 L 9 , 5 G 5,4 , 5 D 2 

260 305 350 395 440 485 530 


138 

5 D4 7 F 5 


7 F6 5 L 10 , 5 G 6 , 5 D 3 

Figure 4.13. (a) Emission spectra (300 K) of the SiO2@APS/DTPA:Tb (solid 

state) excited at 285 nm (black) and 377 nm (red); (b) Excitation spectra 

(300 K) of the SiO2@APS/DTPA:Tb (solid state) monitored at 544 nm 

(blue). 

5 D4 7 F 4 

5 D4 7 F 3 

7 F6 5 D 4 

a 

b




exc 285 nm 

exc 285 nm 

exc 377 nm 

exc 377.5 nm 

5 D4 7 F 6 

5 D4 7 F 5 

380 465 550 635 720 


139 

5 D4 7 F 4 

5 D4 7 F 3 

5 D4 7 F 2-0 

635 665 695 

Figure 4.14. Emission spectra (300 K) of the SiO2@APS@DTPA@Tb in the 

liquid (blue) and solid state at 300K (red): (a) excited at 285 nm; (b) excited 

at ~377 nm. 

The excitation spectra (300K) of the SiO2@APS/DTPA:Tb monitored at 544 nm 

(blue), Figure 4.13b, show one large broad band peaking at 275 nm overlapping a 

series of straight lines ascribed to the Tb 3+ intra- 4f 8 transitions between the 7 F6 and the 

5 L10-7, 5 G6-3, 5 H7-4 5 D3-1, 5 F5,4 5 I8, levels. The broad band is ascribed to the convolution of 

the hybrid host excited states, as observed for the SiO2@APS/DTPA:Eu NPs. 

The comparison between the emission spectra of DTPA:Tb and 

SiO2@APS/DTPA:Tb, displayed in Figure 4.15 (in particular the energy and fwhm of 

the 5 D4 7 F5 line and the energy and relative intensities of the 7 F6-3 Stark components), 

clearly points out to an effective interaction between Tb 3+ ions and the 

SiO2@APS/DTPA host, different from that discerned in the DTPA:Tb complex. 

x10




5 D4 7 F 6 

5 D4 7 F 5 

5 D4 7 F 4 

475 550 625 700 

140 

5 D4 7 F 3 


5 D4 7 F 2-0 

Figure 4.15. Emission spectra (300 K) of the DTPA:Tb (black) and 

SiO2@APS/DTPA:Tb (red) monitored at 544 nm. 

The lifetime values of the Tb 3+ excited states ( 5 D4) were monitored at 546 nm and 

excited at 377 nm, at 27 ºC and solid state. The emission decay curves are well fitted 

by a single exponential function yielding τ ( 5 D4) = 1.87 ± 0.02 ms, indicating only one 

local coordination of the Tb 3+ cation (Figure 4.16). 


0 1 2 3 4 5 6 

0 1 2 3 4 5 6 

Time (ms) 

Figure 4.16. Emission decay curve of Tb 3+ excited states ( 5 D4) from the 

sample SiO2@APS/DTPA:Tb, were monitored at 546 nm and excited at 272 

nm (black) and at 377 nm (blue), at 300 K solid state.



For certain applications, it may be of interest to introduce two different lanthanide 

centres (Ln1, Ln2) optically-active in the visible range. For example, when the colours 

of the emission of Ln1 and the cell auto-fluorescence are similar, one may resort to the 

emission of Ln2. As a proof of concept, SiO2@APS/DTPA:EuTb (1:1) NPs were 

prepared, and they display the red Eu 3+ and green Tb 3+ emission (Figure 4.17a). 

Bimodal, MRI and optical imaging, nanoparticles of SiO2@APS/DTPA:EuGd (1:1) and 

SiO2@APS/DTPA:TbGd (1:1) were prepared. Their emission spectra, depicted in 

Figure 4.17c,d clearly show that the Eu 3+ and Tb 3+ emission features described above 

are not influenced by the incorporation of Gd 3+ . 


380 465 550 635 720 

250 300 350 400 450 500 550 

5 D0 7 F 0 

5 D0 7 F 1 

5 D4 7 F 6 

5 D0 7 F 2 

5 D0 7 F 3 

141 

5 D0 7 F 4 

5 D4 7 F 5 

5 

D0 7 F 

0 

5 

D0 7 F1 5 

D4 7 5 

D4 F3 7 F4 Wavelength (nm) 

5 D0 7 F 2 

380 465 550 635 720 380 465 550 635 720 

5 D4 7 F 5 

5 D0 7 F 3 

450 475 500 525 550 

Figure 4.17. Room-temperature emission spectra of 

SiO2@APS/DTPA:EuTb (1:1) in the solid state, excited at 284 (black), 317 

(blue), 330 (cyan) and 393 nm (magenta). (b) Room-temperature excitation 

of SiO2@APS@DTPA@EuTb (1:1) in the solid state, monitored at 543.5 

(green) and 697.5 nm (red). (c) Room-temperature emission spectra of 

SiO2@APS/DTPA:EuGd (1:1) in the solid state, excited at 290 (black), 

394.5 nm (red), (d) Room-temperature emission spectra of 

SiO2@APS/DTPA:TbGd (1:1) in the solid state, excited at 285 (black), 379 

nm (green). 

a 

b 

c d 

5 D4 7 F 6 

5 D4 7 F 4 

5 D4 7 F 3 

5 D0 7 F 4



SiO2@APS/PMN:Ln NPs 

Figure 4.18a displays the 300 K emission spectra of SiO2@APS/PMN: Eu in the 

solid state excited at two different wavelengths, while Figure 4.19 compares the 

emission spectra measured at three different wavelengths in solid state and water 

suspension. The spectra comprise a series of sharp lines assigned to the Eu 3+ 

5 D0→ 7 F0–4 transitions and a strong broad band centred at ca. 450 nm which, 

considering the emission spectra of SiO2, SiO2@APS and SiO2@APS/PMN (Figures 

4.20 and 4.21), is attributed to oxygen defects in the silica host (and also to the NH2 

groups of the APS layer). 70,73 This broad band shifts to the blue with the addition of 

APS, from 440.0 (SiO2) to 430.0 nm (SiO2@APS, SiO2@APS/PMN), while the full- 

width-at-half-maximum (fwhm) decreases from 133.8 (SiO2) to 109.0 (SiO2@APS), 

94.7 (SiO2@APS/PMN) and 82.6 nm (SiO2@APS/PMN:Eu). 

The excitation spectra of SiO2@APS/PMN:Eu (300 K), Figure 3b, display two 

strong broad bands, peaking at 275.0 and 355.0 nm, overlapping with a series of Eu 3+ 

intra- 4f 6 transitions between the 7 F0 and the 5 L6, 5 D4,2,1, 5 F4, 5 H3, 5 G2-5 levels. When the 

temperature is decreased from 300 to 12 K a new and very broad band in the range ca. 

300-390 nm is observed (Figure 4.22). 


380 415 450 485 520 

5 D0 7 F 0 

380 465 550 635 720 

7 F0 5 H 3 x10 

SIGNAL1 

7 F0 5 D 4 

7 F0 5 G 2-5 

5 D0 7 F 1 

7 F0 5 L 6 

7 F0 5 D 2 

5 D0 7 F 2 

250 300 350 400 450 500 550 


5 D0 7 F 3 

142 

7 F0 5 D 1 

5 D0 7 F 4 

a 

b 

c 5 D0 7 F 0 

576 578 580 582 

d 

e 

5 

D0 7 Wavelength (nm) 

F1 581 584 587 590 593 596 599 602 


5 D0 7 F 2 

603 603 608 608 613 613 618 618 623 623 628 628 633 633 638 638 

Wavelength Wavelength (nm) (nm) 

Figure 4.18. Emission spectra (300 K) of SiO2@APS/PMN:Eu (solid state) 

excited at 270.5 (black) and 393.5 nm (red); b) Excitation spectra (300 K) of 

SiO2@APS/PMN:Eu (solid state) monitored at 420 (magenta) and 614 nm 

(blue); c) d), and e) show a magnification of the 5 D0 7 F0-2 transitions.




a) 

b) 

c) 

exc 270 nm 

exc 270 nm 

exc 318 nm 

exc 317 nm 

exc 393.5 nm 

exc 393 nm 

380 465 550 635 720 


143 

5 D0 7 F 0 

5 D0 7 F 1 

5 D0 7 F 2 

Figure 4.19. Emission spectra (300 K) of the SiO2@APS/PMN:Eu in water 

solution (blue), solid state at 300K (red) excited at different wavelengths. 

Intensity Intensity Intensity (arb. (arb. (arb. units) units) units) 

Intensity Intensity Intensity (arb. (arb. (arb. units) units) units) 

aa 

285 285 285 370 370 370 455 455 455 540 540 540 625 625 625 710 710 710 

Wavelength Wavelength Wavelength (nm) (nm) (nm) 

bb 

5 D0 7 F 3 

245 245 245 290 290 290 335 335 335 380 380 380 425 425 425 

245 245 245 290 290 290 335 335 335 380 380 380 425 425 425 

285 285 285 370 370 370 455 455 455 540 540 540 625 625 625 710 710 710 

Wavelength Wavelength Wavelength (nm) (nm) (nm) 

Figure 4.20. (a) Emission spectra (300 K) of the SiO2@APS/PMN excited at 

270.0 (black), 300.0 (red), 330.0 (green), 360.0 (blue), 370.0 (cyan) and 

485.0 (magenta). The inset shows the excitation spectra (300 K) monitored 

at 340.0 (black), 425.0 (red), 440.0 (green) and 510.0 nm (blue); (b) 

5 D0 7 F 4



Emission spectra (300 K) of the SiO2@APS excited at 270.0 (black), 300.0 

(red), 330.0 (green), 360.0 (blue), 380.0 (cyan) and 400.0 (magenta). The 

inset shows the excitation spectra (300 K) monitored at 340.0 (black), 420.0 

(red), 440.0 (green) and 520.0 nm (blue). 


SiO 2 ex 360 nm 

SiO 2 @APS ex 360 nm 

SiO 2 @APS/PMN ex 360 nm 

SiO 2 @APS/PMN:Eu ex 360 nm 

375 460 545 630 715 


Figure 4.21. Emission spectra (300 K) of the SiO2 (black), SiO2@APS (red), 

SiO2@APS/PMN (green) and SiO2@APS/PMN:Eu (blue) excited at 360.0 

nm. 


250 290 330 370 410 450 490 530 570 


Figure 4.22. Excitation spectra of the SiO2@APS/PMN:Eu monitored at 614 

nm, 300 K (red) and 12 K (blue). 

144



In order to get further insight into the Eu 3+ local coordination, the 4f 6 emission 

lines were recorded at high resolution (Figure 4.18c-e). The spectra excited at 270.5 

and 393.5 nm exhibit differences in the relative intensities and fwhm of the 5 D0 7 F0-2 

transitions. The fwhm of the non-degenerated 5 D0 7 F0 line, in particular, is 43.3 ± 0.3 

cm -1 , at 270.5 nm, and 53.5 ± 0.5 cm -1 , at 393.5 nm, which evidences the presence of 

(at least) two distinct Eu 3+ local environments. A similar conclusion is reached for the 

NPs water suspensions (Figure 4.19). 

The room-temperature 5 D0 emission decay curve of solid SiO2@APS/PMN:Eu, 

monitored within the 5 D0→ 7 F2 transition at 614.0 nm and excited at 270.0 nm, is well 

fitted by a bi-exponential function, yielding lifetimes τ1 = 0.242 0.071 ms and τ2 = 

0.927 0.023 ms (Figure 4.23a), confirming the presence of two Eu 3+ local 

environments. The 5 D0 decay curve of the free [Eu(PMN)] monomeric complex, 

recorded in the solid state excited at 275.0 nm (Figure 4.24), is well fitted by a mono- 

exponential function yielding a lifetime of 0.292 0.062 ms (in solution the reported 

lifetime is ca. 0.400 ms 61-62 . 


a 

0 1 2 3 4 5 6 7 8 

b 

0 1 2 3 4 

Time (ms) 

5 6 7 

Figure 4.23. Emission decay curve of Eu 3+ excited states ( 5 D0) from the 

sample SiO2@APS/PMN:Eu, which were monitored at 614.0 nm and 

excited at a) 270.0 nm (black); b) 370.0 nm (blue), at 300 K in the solid 

state. 

145



Intensity (Counts) 

2980.95799 

1096.63316 

403.42879 

148.41316 

54.59815 

20.08554 

7.38906 

2.71828 

1 

ex = 275.0 nm; em = 614 nm 

0 1 2 3 4 5 6 

146 

Time (ms) 

Figure 4.24. 5 D0 emission decay curve of [Eu(PMN)], monitored at 614.0 

nm and excited 275.0 nm (blue) at 300 K in the solid state. 

Although the free [Eu(PMN)] complex and SiO2@APS/PMN:Eu exhibit similar 

lifetimes of, respectively, 0.29 and 0.24 ms (τ 1) there are indications that the latter is 

not attributed to the [Eu(PMN)] residue, rather to a Eu 3+ environment interacting 

strongly with the NPs surface, probably via silanol groups (which are far more 

abundant that the amino groups on the silica surface). The arguments are as follows. 

Figure 4.25 compares the emission and excitation spectra of the free [Eu(PMN)] 

complex and SiO2@APS/ PMN:Eu NPs. The 5 D0 7 F0 line of the latter can be fitted by 

two Gaussian bands, at ca. 578 and 580 nm (inset in Figure 4.25), in accord with the 

existence of two distinct Eu 3+ local environments. Whereas the low-energy-component 

(ca. 578 nm) is ascribed to a Eu 3+ site with a coordination shell similar to that of the free 

[Eu(PMN)] complex, the high-energy band (ca. 580 nm), not present in the complex, 

must correspond to a Eu 3+ coordination shell involving directly the SiO2@APS host. 

Therefore the excitation spectra monitored on both components provide information on 

the nature of the two distinct local sites. Figure 4.26 compares the excitation spectra of 

the SiO2@APS/PMN:Eu NPs, monitored at 578.0, 580.0 and 614.0 nm, and [Eu(PMN)] 

complex, monitored at 614.0 nm. While the spectra monitored at 580.0 nm and that of 

the free complex are similar (Figure 4.26), the spectrum monitored at 578.0 nm shows, 

in addition, a broad band (‘pedestal’) between 350.0 and 425.0. This band is, thus, 

assigned to the SiO2@APS host (Figure 4.18b and Figure 4.19). The 5 D0 emission 

decay curve excited at 370.0 nm (Figure 4.25b) is well fitted by a mono-exponential



function yielding a lifetime of 0.212 0.014 ms 66 providing solid evidence for the 

previous assignment of the shorter lifetime to a Eu 3+ environment interacting strongly 

with the NPs surface, that corresponds to the low-energy component of the 5 D0 7 F0 

line at ca. 578. 

576 578 580 582 


5 D0 7 F 0 

5 D0 7 F 1 

5 D0 7 F 2 

570 620 670 

PMN:Eu 614 nm 

em 

720 

7 F0 5 H 3 

7 F0 5 D 4 

7 F0 5 G 2-5 

250 300 350 400 450 500 550 

147 

PMN:Eu exc 393 nm 

SiO 2 @APS/PMN:Eu exc 393.5 nm 

5 D0 7 F 3 

5 D0 7 F 4 

SiO 2 @APS@PMN:Eu em 614 nm 

7 F0 5 L 6 


Figure 4.25. Top: Emission spectra (300 K) of [Eu(PMN)] (black) and 

SiO2@APS/PMN:Eu (red) excited at 393 nm and 393.5 nm, respectively; 

Bottom: Excitation spectra (300 K) of [Eu(PMN)] (black) and 

SiO2@APS/PMN:Eu (red) monitored at 614 nm. 

7 F0 5 L 6 

7 F0 5 D 2




7 F0 5 H 3 

7 F0 5 D 4 

7 F0 5 G 2-5 

250 300 350 400 450 500 550 

148 

7 F0 5 L 6 

7 F0 5 L 6 


Figure 4.26. Excitation spectra (300 K) of [Eu(PMN)] (black, monitored at 

614.0 nm) and SiO2@APS/PMN:Eu (monitored at 614.0, 578.0 and 580.0 

nm, red, green and blue curves, respectively). 

We assign the longer, τ 2, lifetime to Eu 3+ coordinated to the PMN ligand and 

grafted via APS to the SiO2 surface because the excitation spectra of the free 

[Eu(PMN)] complex and SiO2@APS/PMN:Eu are very similar (Figure 4.25 and Figure 

4.26) in the region of the intra-4f lines, exhibiting also the same broad band at ca. 270 

nm, assigned to the aromatic ring of PMN. The longer lifetime of SiO2@APS/PMN:Eu 

relatively to lifetime of the free [Eu(PMN)] complex may be due to differences in the 

Eu 3+ coordination spheres, in particular, brought about by the peptoidic coupling of 

PMN to APS in the former. 

The maximum emission absolute quantum yields of the free [Eu(PMN)] complex 

and SiO2@APS/PMN:Eu NPs (measured at 270.5 nm) were, respectively, 0.04 0.01 

and 0.05 0.01. Importantly, a much smaller value (0.01 0.01) has been reported for 

a similar system but that did not contain an aromatic ‘antenna’, 

SiO2@APS@DTPA:Eu .60 

SiO2@APS/ PMN:Tb NPs were also prepared an their emission spectra consist of 

(i) a broad band between 380.0 and 550.0 nm, ascribed to the emission of the 

SiO2@APS/PMN host as observed for the Eu-containing samples, and (ii) a series of 

sharp Tb 3+ 5 D4→ 7 F6–0 intra-4f 8 lines (Figure 4.27). 

7 F0 5 D 2



As mentioned before, it is of interest to introduce two different optically-active 

lanthanide ions (Ln1, Ln2) emitting in the visible range. As a proof of concept, 

SiO2@APS/PMN:EuTb (1:1) NPs were prepared, and they displayed the red, Eu 3+ , and 

green, Tb 3+ , emission (Figure 4.28). 


exc 285 nm 

exc 377 nm 

380 465 550 635 720 

7 F6 5 H 7 , 5 D 1 

5 D4 7 F 6 

7 F6 5 L 7,8 

7 F6 5 L 9 , 5 G 5,4 , 5 D 2 

260 305 350 395 440 485 530 


149 

5 D4 7 F 5 

5 D4 7 F 4 

5 D4 7 F 3 


7 F6 5 L 10 , 5 G 6 , 5 D 3 

305 330 355 380 405 430 455 480 505 530 

Figure 4.27. (a) Emission spectra (300 K) of the SiO2@APS/PMN:Tb (solid 

state) excited at 280.0 nm (black) and 377.0 nm (red); (b) Excitation spectra 

(300 K) of the SiO2@APS/PMN:Tb (solid state) monitored at 544.0 nm 

(blue). 

The emission spectra of bimodal SiO2@APS/PMN:EuGd (1:1) and 

SiO2@APS/PMN:TbGd (1:1) NPs for MRI and optical imaging (Figure 4.29) are very 

similar to the spectra of the Gd-free samples (Figure 4.18a and Figure 4.28), thus 

showing that the Eu 3+ and Tb 3+ emission is not influenced by the presence of Gd 3+ . 

7 F6 5 D 4 

a 

b




5 D4 7 F 6 

380 465 550 635 720 

250 300 350 400 450 500 550 


150 

5 D4 7 F 5 

5 

D4 7 5 

D0 F4 7 F0 5 

D0 7 F1 5 D4 7 F 3 

5 D0 7 F 2 

5 D0 7 F 3 

5 D0 7 F 4 

450 475 500 525 550 

Figure 4.28. (a) Room-temperature emission spectra of 

SiO2@APS/PMN:EuTb (1:1) in the solid state, excited at 284.0 (black), 

317.0 (blue), 330.0 (cyan) and 393.0 nm (magenta). (b) Room-temperature 

excitation of SiO2@APS/PMN:EuTb (1:1) in the solid state, monitored at 

543.5 (green) and 697.5 nm (red). 


a 

380 465 550 635 720 

b 

5 D4 7 F 6 

5 D0 7 F 0 

5 D0 7 F 1 

5 D4 7 F 5 

5 D0 7 F 2 

380 465 550 635 720 

5 D4 7 F 4 

5 D4 7 F 3 

5 D0 7 F 3 

5 D4 7 F 2-0 

Figure 4.29. Room-temperature emission spectra of a) 

SiO2@APS/PMN:EuGd (1:1), b) SiO2@APS/PMN:TbGd (1:1) NPs, excited 

at, respectively, 290 nm (black), 394.5 nm (red), and 285 (black), 379 nm 

(green). 

5 D0 7 F 4



4.3.3. Relaxivity Properties 

SiO2@APS/DTPA:Ln NPs 

The SiO2@APS/DTPA:Ln (Ln= Gd, Eu:Gd (1:1), Tb:Gd (1:1) NPs suspensions 

remained stable throughout the NMR measurements, allowing the collection of 

consistent relaxation data. Table 4.4 shows the proton relaxivity values (r1p and r2p), 

determined at two frequencies (20 MHz and 500 MHz) and two temperatures (25 ºC 

and 37 ºC) for the SiO2@APS/DTPA:Gd, SiO2@APS/DTPA:EuGd (1:1) and 

SiO2@APS/DTPA:TbGd (1:1) water suspensions. These relaxivities were calculated 

from the observed linear dependence of the ) 

151 

R i 

i 1/ T , ( i 1, 

2 relaxation rates on the 

concentration of the Gd 3+ ions present in all samples, shown in Figure 4.30 to Figure 

4.33. These values are constant over a large pH range, indicating that the 

paramagnetic NPs are stable and do not leach out Gd 3+ , Eu 3+ or Tb 3+ ions, even in 

highly basic conditions. The ) 

rip ( i 1, 

2 values measured for the 

SiO2@APS/DTPA:Gd nanoparticles with 67 nm diameter are very similar to those 

reported for the monomeric [Gd(DTPA)] 2- complex, 8 reflecting the virtually free 

rotational motion of the complex at the surface of the nanoparticles, which counteracts 

the effect of the slow global motion of the nanoparticle on the relaxivities. The 

r1 p values decrease with increasing frequency, as expected for the standard inner- 

sphere and outer-sphere dipolar mechanisms of proton relaxation. They are also 

almost constant with increasing temperature, reflecting that the T 1 relaxation process is 

limited by slow-to-intermediate water exchange, characteristic of DTPA-amide 

systems. 8,9 These r1p values are similar to those reported for nanoporous silica 

nanoparticles coated with covalently bound Gd-Si-DTPA 30 or Si-EDTA 46 derivatives, but 

smaller than when a Si-DTTA 30 derivative was used (H4DTTA = 

diethylenetriaminetetraacetic acid), mainly reflecting the different water accessibilities 

of the Gd 3+ ion in those systems. This water accessibility is much increased in 

mesoporous silica-based nanosystems covalently labelled with Gd-DTPA, Gd-DTTA or 

Gd-DOTA derivatives, 40-45 leading to r1p values 5 to 10 times larger than for the 

corresponding monomeric complexes.



For the SiO2@APS/DTPA:GdEu and SiO2@APS/DTPA:GdTb NPs, where 50% of 

the DTPA-coordinated Gd 3+ ions are replaced by Eu 3+ or Tb 3+ , the r1p values (referred 

to one mM Gd 3+ ) increase relative to the SiO2@APS/DTPA:Gd NPs (Table 4.4), 

reflecting the dipolar relaxation effect of the extra ions at the particle surface. This 

increase is larger for the Tb 3+ than for Eu 3+ ions, as the former induces stronger 1 

relaxation due to its slower electronic relaxation. The frequency and temperature 

dependence of r1p for the mixed cation nanoparticles is the same as for the Gd 3+ ones. 

The r2p values for the SiO2@APS/DTPA:Gd NPs undergo a large increase when 

the measuring frequency increases (Table 4.4). Large r2p values have also been 

observed for silica nanosystems covalently labelled with Gd 3+ complexes, particularly at 

high frequencies. 30,40-42,45-47 This indicates that the 2 

dipolar mechanism operating for 1 

152 

T -relaxation process, besides the 

T -relaxation, also has a strong outer-sphere 

contribution from field inhomogeneities created by the magnetized particles that the 

water protons experience (measured by the frequency shift at the particle surface, Δ) 

as they diffuse nearby (with a diffusion correlation time τD), and which increase with the 

square of the external magnetic field strength. 80 The presence of this contribution is 

confirmed by the increase of p 

r 2 values observed for the mixed SiO2@APS/DTPA: 

EuGd NPs (Table 4.4). This magnetic susceptibility effect is particularly strong for p 

values at 500 MHz, and can also be observed for the 20 MHz p 

r 2 

T 

r 2 values of the mixed 

SiO2@APS/DTPA: TbGd NPs. These effects of the nanoparticle-bound Tb 3+ ions are 

stronger than those observed for the Eu 3+ ions, in agreement with the larger magnetic 

moment of Tb 3+ . However, their 500 MHz r2 p values decrease, rather than increase, 

when 50% of the Gd 3+ ions are replaced by Tb 3+ (Table 4.4). This may reflect a 

breakdown of the outer-sphere relaxation model for T2 -relaxation at high magnetic field 

due to the presence of the Tb 3+ ions, when τD >> 1/ Δ. In these conditions, the static 

dephasing regime (SDR) model describes the transverse relaxation and the value of 2 

becomes dependent on the time interval between two consecutive refocusing pulses 

(τCP) in the train of 180 o pulses applied in a Carr-Purcell-Meiboom-Gill (CPMG) pulse 

sequence. 80,81 In preliminary experiments, we have observed that 2 

r of suspensions of 

these particles indeed depends on τCP (data not shown). A more complete study of the 

relaxation mechanisms of these mixed NPs is beyond the scope of the present study. 

r



Table 4.4. Calculated 1 H relaxivity values, rip (i = 1,2), determined at 20 

MHz and 500 MHz, at 25 ºC and 37 ºC for samples SiO2@APS/DTPA:Gd, 

SiO2@APS/DTPA:EuGd (1:1) and SiO2@APS/DTPA:TbGd (1:1). 

20 MHz 

r1p (s -1 mM -1 ) 

153 

r2p (s -1 mM -1 ) 

25º 37º 25º 37º 

SiO2@APS/DTPA:Gd 5.24 ± 0.04 5.66 ± 0.03 6.36 ± 0.01 6.86 ± 0.01 

SiO2@APS/DTPA:EuGd (1:1) 8.08 ± 0.03 8.39 ± 0.02 10.09 ± 0.003 10.26 ± 0.007 

SiO2@APS/DTPA:TbGd (1:1) 17.4 ± 0.1 16.6 ± 0.1 21.59 ± 0.01 20.85 ± 0.05 

500 MHz 

r1p (s -1 mM -1 ) 

r2p (s -1 mM -1 ) 

25º 37º 25º 37º 

SiO2@APS/DTPA:Gd 2.08 ± 0.04 1.93 ± 0.03 26.6 ± 0.4 34.8 ± 0.6 

SiO2@APS/DTPA:EuGd (1:1) 2.64 ± 0.08 2.50 ± 0.09 50 ± 2 55 ± 3 

SiO2@APS/DTPA:TbGd (1:1) 13.1 ± 0.6 9.5 ± 0.7 22 ± 1 22.2 ± 0.6 

R1 (s -1 ) 

a-I 

a-II 

a-III 

50 

40 

30 

20 

10 

0 

Slope=17.4 

R 

0.0 0.5 1.0 1.5 2.0 2.5 3.0 

14 

12 

10 

8 

6 

Slope=8.08 

4 

2 

R 

0 

0.0 0.5 1.0 1.5 2.0 

 

35 

30 

25 

20 

15 

Slope=5.24 

10 

5 

R 

0 

0 1 2 3 4 5 6 

 

b-I 

b-II 

b-III 

50 

40 

30 

20 

10 

0 

0.0 0.5 1.0 1.5 2.0 2.5 3.0 

14 

12 

10 

8 

6 

4 

2 

Slope=8.39 

R 

0 

0.0 0.5 1.0 1.5 2.0 

 

Concentration [Gd 3+ ] (mM) 

Slope=16.6 

R 

Slope=5.66 

R 35 

30 

25 

20 

15 

10 

5 

 

0 

0 1 2 3 4 5 6 

Figure 4.30. r1 values measured at 20 MHz; (a) 25ºC and (b) 37ºC and 

samples I- SiO2@APS/DTPA:Gd; II- SiO2@APS/DTPA:EuGd; (1:1) and 

III- SiO2@APS/DTPA:TbGd (1:1).



R2 (s -1 ) 

a-I 

a-II 

a-III 

70 

60 

50 

40 

30 

Slope=21.59 + 0.01 

20 

R 

10 

0 

0.0 0.5 1.0 1.5 2.0 2.5 3.0 

2 =0.99958 

18 

16 

14 

12 

10 

8 

Slope=10.086 + 0.003 

6 

4 

R 

2 

0 

0.0 0.5 1.0 1.5 2.0 

2 =0.99972 

35 

30 

25 

20 

15 

Slope=6.36 + 0.01 

10 

R 

5 

0 

0 1 2 3 4 5 6 

2 =0.99961 


154 

b-I 

b-II 

b-III 

Slope=20.85 + 0.05 

R 2 70 

60 

50 

40 

30 

20 

=0.99994 

10 

0 

0.0 0.5 1.0 1.5 2.0 2.5 3.0 

18 

16 

14 

12 

10 

8 

Slope=10.263 + 0.007 

6 

4 

R 

2 

0 

0.0 0.5 1.0 1.5 2.0 

2 =0.99982 

Slope=6.86 + 0.01 

R 2 35 

30 

25 

20 

15 

10 

=0.99895 

5 

0 

0 1 2 3 4 5 6 

Figure 4.31. r2 values measured at 20 MHz; (a) 25ºC and (b) 37ºC and samples I- 

SiO2@APS/DTPA:Gd; II- SiO2@APS/DTPA:EuGd; (1:1) and 

III- SiO2@APS/DTPA:TbGd (1:1). 

R1 (s -1 ) 

a-I 

a-II 

a-III 

40 

30 

20 

10 

4 

3 

2 

1 

0 

0.0 0.5 1.0 1.5 2.0 

12 

10 

8 

6 

4 

2 

Slope=13.1 

R 

0 

0.0 0.5 1.0 1.5 2.0 2.5 3.0 

Slope=2.64 

R 

Slope=2.08 

R 

0 

0 1 2 3 4 5 6 

b-I 

b-II 

b-III 

40 

30 

20 

10 


0 

0.0 0.5 1.0 1.5 2.0 2.5 3.0 

4 

3 

2 

1 

Slope=9.5 

R 

Slope=2.50 

R 

0 

0.0 0.5 1.0 1.5 2.0 

12 

10 

8 

6 

4 

2 

Slope=1.93 

R 

0 

0 1 2 3 4 5 6 



III- SiO2@APS/DTPA:TbGd (1:1).



R2 (s -1 ) 

a-I 

a-II 

a-III 

70 

60 

50 

40 

30 

Slope= 22 + 1 

20 

R 

10 

0 

0.0 0.5 1.0 1.5 2.0 2.5 3.0 

2 =0.99201 

100 

80 

60 

40 

20 

0 

0.0 0.2 0.4 0.6 0.8 1.0 1.2 1.4 1.6 1.8 2.0 

200 

150 

100 

50 

Slope= 50 + 2 

R 2 =0.99724 

Slope= 26.6 + 0.4 

R 2 =0.99355 

0 

0 1 2 3 4 5 6 


155 

b-I 

b-II 

b-III 

70 

60 

50 

40 

30 

Slope= 22.2 + 0.6 

20 

R 

10 

0 

0.0 0.5 1.0 1.5 2.0 2.5 3.0 

2 =0.99857 

100 

80 

60 

40 

20 

Slope= 55 + 3 

R 2 =0.99788 

0 

0.0 0.5 1.0 1.5 

200 

150 

100 

50 

Slope= 34.8 + 0.6 

R 2 =0.99495 

0 

0 1 2 3 4 5 6 



III- SiO2@APS/DTPA:TbGd (1:1). 

SiO2@APS/PMN:Ln NPs 

The SiO2@APS/PMN:Ln (Ln= Gd, Eu:Gd (1:1), Tb:Gd (1:1) NPs suspensions 

were studied by NMR. Table 4.5 shows the proton relaxivity values ( p 

r 2 ) 

r 1 and p 

determined at two frequencies (20 MHz and 500 MHz) and two temperatures (298 K 

and 310 K) for the SiO2@APS/PMN:Gd, SiO2@APS/PMN:EuGd (1:1) and 

SiO2@APS/PMN:TbGd (1:1) water suspensions. These relaxivities were calculated 

from the slopes of the observed linear dependence of the Ri 1/ Ti 

, ( i 1, 

2) 

relaxation 

rates on the concentration of the Gd 3+ ions present in the samples, shown in figures 

4.34 to 4.37. These values are constant over a large pH range, indicating that the 

paramagnetic NPs are stable and do not leach out the Ln 3+ ions, even in highly basic 

r i 

( i 1, 

2 

conditions. The ) values measured for the SiO2@APS/PMN:Gd nanoparticles 

with 67 nm diameter are very similar to those reported for the monomeric [Gd(DTPA)] 2- 

complex 8 and our previous studies for SiO2@APS/DTPA:Gd NPs. 60



In the cases of the mixed-metal SiO2@APS/PMN:GdEu and 

SiO2@APS/PMN:GdTb NPs, where 50% of the PMN-coordinated Gd 3+ ions are 

replaced by Eu 3+ or Tb 3+ , the r1p values (referred to one mM Gd 3+ ) increase relative to 

the SiO2@APS/PMN:Gd NPs (Table 4.5), and are in agreement with previous studies 

for SiO2@APS/DTPA:GdEu and SiO2@APS/DTPA:GdTb NPs. 60 

Table 4.5. Calculated 1 H relaxivity values, rip (i = 1,2), determined at 20 

MHz and 500 MHz, at 25 ºC and 37 ºC for samples SiO2@APS/PMN:Gd, 

SiO2@APS/PMN:EuGd (1:1) and SiO2@APS/PMN:TbGd (1:1). 

20 MHz 

r1p (s -1 mM -1 ) 

156 

r2p (s -1 mM -1 ) 

25º 37º 25º 37º 

SiO2@APS/PMN:Gd 2.70± 0.01 2.83 ± 0.01 4.32 ± 0.01 4.02 ± 0.01 

SiO2@APS/PMN:EuGd 4.14 ± 0.01 4.39 ± 0.01 7.68 ± 0.01 7.50 ± 0.01 

SiO2@APS/PMN:TbGd 8.49 ± 0.01 7.95 ± 0.02 11.98 ± 0.01 12.40 ± 0.01 

500 MHz 

r1p (s -1 mM -1 ) 

r2p (s -1 mM -1 ) 

25º 37º 25º 37º 

SiO2@APS/PMN:Gd 1.05 ±0.03 1.10 ± 0.02 80 ± 1 56.4 ± 0.8 

SiO2@APS/PMN:EuGd 2.13 ± 0.05 1.91 ± 0.06 227 ± 4 174 ± 2 

SiO2@APS/PMN:TbGd 2.26 ± 0.04 2.13 ± 0.05 89.9 ± 0.8 79.4 ± 0.9 

The p 

r 2 values for the SiO2@APS/PMN:Gd NPs increase sharply at high 

frequency (Table 4.5), as also observed for other silica nanosystems covalently 

labelled with Gd 3+ complexes. 42-44,46,5380-]82 This results from field inhomogeneities 

created by the magnetized particles (measured by the frequency shift at the particle 

surface, Δ) that the water protons experience as they diffuse nearby (with a diffusion 

correlation time τD). This outer-sphere contribution to 2 

T -relaxation increases with the 

square of the external magnetic field strength. 80 This contribution also causes an 

increase of r2 p values for the mixed SiO2@APS/PMN:GdEu NPs, in particular at 500



MHz, (Table 4.5). This effect is even stronger for the Tb 3+ -containing NPs, due to its 

larger magnetic moment. The decrease of their 500 MHz p 

157 

r 2 values when 50% of the 

Gd 3+ ions are replaced by Tb 3+ (Table 4.5), may result from a breakdown of the outer- 

sphere relaxation model for 2 

T -relaxation at high magnetic fields due to the presence 

of the Tb 3+ ions, when τD >> 1/Δ. In these conditions, the static dephasing regime 

(SDR) model describes the transverse relaxation and the value of r2 becomes 

dependent on the time interval between two consecutive refocusing pulses (τCP) in the 

train of 180 o pulses applied in a Carr-Purcell-Meiboom-Gill (CPMG) pulse sequence. 42 

In fact, preliminary experiments showed that the 2 

depend on τCP (data not shown), as expected in those conditions. 

R1 (s -1 ) 

a-I 

a-II 

a-III 

15 

10 

5 

0 

0.0 0.5 1.0 1.5 2.0 

10 

8 

6 

4 

Slope=4.137 

2 

R 

0 

0.0 0.5 1.0 1.5 2.0 

 

10 

8 

6 

4 

2 

Slope=8.493 

R 

Slope=2.699 

R 

0 

0 1 2 3 4 

b-I 

b-II 

b-III 

15 

10 

10 


5 

8 

6 

4 

2 

r of suspensions of these particles 

Slope=7.95 

R 

0 

0.0 0.5 1.0 1.5 2.0 

Slope=4.386 

R 

0 

0.0 0.5 1.0 1.5 2.0 

10 

8 

6 

4 

2 

Slope=2.830 

R 

0 

0 1 2 3 4 


samples I- SiO2@APS/PMN:Gd; II- SiO2@APS/PMN:EuGd; (1:1) and 

III- SiO2@APS/PMN:TbGd (1:1).



R2 (s -1 ) 

a-I 

a-II 

a-III 

20 

15 

10 

5 

0 

0.0 0.5 1.0 1.5 2.0 

15 

10 

5 

0 

0.0 0.5 1.0 1.5 2.0 

15 

10 

5 

Slope=11.985 + 0.006 

R 2 =0.99677 

Slope=7.679 + 0.002 

R 2 =0.99695 

Slope=4.323 + 0.002 

R 2 =0.97687 

0 

0 1 2 3 4 


158 

b-I 

b-II 

b-III 

20 

15 

10 

Slope=12.397 + 0.003 

R 2 5 

=0.98978 

0 

0.0 0.5 1.0 1.5 

15 

10 

5 

Slope=7.503 + 0.003 

R 2 =0.99661 

0 

0.0 0.5 1.0 1.5 2.0 

15 

10 

5 

Slope=4.025 + 0.002 

R 2 =0.99677 

0 

0 1 2 3 4 



III- SiO2@APS/PMN:TbGd (1:1). 

R1 (s -1 ) 

a-I 

a-II 

a-III 

5 

4 

3 

2 

1 

0 

4 

3 

2 

1 

0 

0.0 0.5 1.0 1.5 2.0 

4 

3 

2 

1 

Slope=2.26 

R 

0.0 0.5 1.0 1.5 

Slope=2.13 

R 

Slope=1.05 

R 

0 

0 1 2 3 4 

b-I 

b-II 

b-III 


4 

3 

2 

1 

4 

3 

2 

1 

Slope=2.13 

R 

0 

0.0 0.5 1.0 1.5 2.0 

Slope=1.91 

R 

0 

0.0 0.5 1.0 1.5 2.0 

4 

3 

2 

1 

Slope=1.10 

R 

0 

0 1 2 3 4 



III- SiO2@APS/PMN:TbGd (1:1).



R2 (s -1 ) 

a-I 

a-II 

a-III 

150 

100 

50 

400 

300 

200 

100 

Slope= 89.9 + 0.8 

R 2 =0.99837 

0 

0.0 0.5 1.0 1.5 2.0 

Slope= 227 + 4 

R 2 =0.95496 

0 

0.0 0.5 1.0 1.5 2.0 

350 

300 

250 

200 

150 

Slope= 80 + 1 

100 

R 

50 

0 

0 1 2 3 4 

2 =0.98537 


159 

b-I 

b-II 

b-III 

150 

100 

50 

400 

300 

200 

100 

Slope= 79.4 + 0.9 

R 2 =0.97588 

0 

0.0 0.5 1.0 1.5 2.0 

Slope= 174 + 2 

R 2 =0.97877 

0 

0.0 0.5 1.0 1.5 2.0 

Slope= 56.4 + 0.8 

R 2 350 

300 

250 

200 

150 

100 

50 

=0.96426 

0 

0 1 2 3 4 



III- SiO2@APS/PMN:TbGd (1:1). 

4.3.4. Cell Imaging 

Regarding the cellular uptake of the both NPs, the results obtained show that 

they are rapidly internalized by RAW 264.7 cells. In the case of the 

SiO2@APS/DTPA:Ln study, the T1 -weighted MRI images of cellular pellets with cells 

incubated with and without NPs, are shown in Figure 4.38a. A clear increase in the 

intensity of the pellets (positive contrast), obtained with cells incubated with 

SiO2@APS/DTPA:EuGd NPs (sample III), is observed relative to the pellets 

corresponding to unexposed cells (sample I), as opposed to the strong negative 

contrast caused by internalization of the T2 -shortening Fe2O3 NPs (sample II). The 

optical features of the NPs internalized cells were also assessed at a wavelength of 

393 nm. The results illustrated in Figure 4.38b demonstrate that the fluorescence of 

sample II is a combination of the autofluorescence of cells and the fluorescence 

exhibited by the SiO2@APS/DTPA:EuGd NPs (violet is the combination of red and



blue). These observations confirm the potential of the NPs as optical imaging contrast 

agents. 

a 

c) 

b) 

a) 

Figure 4.38. (a) 1 

III 

I 

II 

T -weighted MRI image of cellular pellets corresponding to: I- no NP 

internalization (control); II- -Fe2O3 NPs ( T 2 ) NP cell internalization and III- 

SiO2@APS/DTPA:EuGd NP cell internalization; (b) Photograph of cellular pellets, 

excited at 393 nm, corresponding to: I- no NP cell internalization (control) and II- 

SiO2@APS/DTPA:EuGd NP cell incorporation. 

In the case of the SiO2@APS/PMN:Ln study, the 1 

160 

T -weighted MRI images of 

cellular pellets with cells incubated with and without NPs, are shown in Figure 4.39a. A 

clear increase in the intensity of the pellets (positive contrast), obtained with cells 

incubated with SiO2@APS/PMN:EuGd NPs (sample III), is observed relative to the 

pellets corresponding to unexposed cells (sample I), however the sample with cells 

incubated with SiO2@APS/PMN:Gd NPs (sample II), however, displays a decrease in 

intensity of the pellets confirming the r1 values obtained above. The optical features of 

the NPs internalized cells were also assessed at a wavelength of 393 nm. The results 

illustrated in Figure 4.39c demonstrate that the fluorescence of sample II and III is a 

combination of the autofluorescence of cells (sample I) and the red fluorescence 

exhibited by the Eu 3+ ion (in Figure 4.39b),(violet is the combination of red and blue). 

These observations confirm the potential of the NPs as optical imaging contrast agents. 

b 

II 

I



a 

I 

III 

II 

Figure 4.39. (a) 1 

T -weighted MRI image of cellular pellets: I- no NP internalization 

(control); II- SiO2@APS/PMN:Gd cell internalized NPs; III- SiO2@APS/PMN:EuGd cell 

internalized NPs; (b) Photograph of SiO2@APS/PMN:Eu NPs in the solid state (top) 

and suspension (bottom), excited at 393 nm (right) and non-excited (left); (c) 

Photograph of cellular pellets, excited at 393 nm: I- no NPs cell internalization (control); 

II-SiO2@APS/PMN:Eu cell internalized NPs; III- SiO2@APS/PMN: EuGd cell 

internalized NPs. 

b 

4.4. CONCLUSIONS 

Bimodal MRI - optical probes for bio-imaging, based on SiO2 nanoparticles 

derivatized with DTPA-Ln and PMN-Ln complexes (SiO2@APS/DTPA:Gd:Ln; and 

SiO2@APS/PMN:Gd:Ln; Ln= Eu 3+ , Tb 3+ ) were developed. The incorporation of Gd 3+ 

ions (the MRI probe) in the nanosystems does not change the emission properties of 

the Eu 3+ and Tb 3+ ions, while the relaxometric features are slightly better than the 

properties of the commercially-available [Gd(DTPA)] 2- complex. 

The grafting of pyridine-based aromatic ligands (efficient Ln 3+ sensitizers) to the 

silica surface via APS, and their complexation with Gd 3+ , Eu 3+ /Tb 3+ ions affords 

relaxometry, and photoluminescent properties. 

161 

c 

III 

II 

I



Both of these bimodal probes are rapidly and efficiently uptaken by RAW 264.7 

cells (mouse macrophage cell line) and exhibit both, 1 

162 

T -weighted MRI images of 

cellular pellets increased contrast and potential optical tracking by 

fluorescence.increased contrast and potential fluorescence tracking.




1 Townsend, D. W.; Beyer, T.; Blodgett, T. M., PET/CT scanners: A hardware approach to 

image fusion. Seminars in Nuclear Medicine 2003, 33 (3), 193-204. 

2 Momose, A.; Takeda, T.; Itai, Y.; Hirano, K., Phase-contrast X-ray computed tomography 

for observing biological soft tissues. Nature Medicine 1996, 2 (4), 473-475. 

3 Shokeen, M.; Anderson, C. J., Molecular Imaging of Cancer with Copper-64 

Radiopharmaceuticals and Positron Emission Tomography (PET). Accounts of Chemical 

Research 2009, 42 (7), 832-841. 

4 Ametamey, S. M.; Honer, M.; Schubiger, P. A., Molecular imaging with PET. Chemical 

Reviews 2008, 108 (5), 1501-1516. 

5 Wadas, T. J.; Wong, E. H.; Weisman, G. R.; Anderson, C. J., Coordinating Radiometals 

of Copper, Gallium, Indium, Yttrium, and Zirconium for PET and SPECT Imaging of 

Disease. Chemical Reviews 2010, 110 (5), 2858-2902. 

6 Zhang, R.; Xiong, C.; Huang, M.; Zhou, M.; Huang, Q.; Wen, X.; Liang, D.; Li, C., 

Peptide-conjugated polymeric micellar nanoparticles for Dual SPECT and optical 

imaging of EphB4 receptors in prostate cancer xenografts. Biomaterials 2011, 32 (25), 

5872-5879. 

7 Ntziachristos, V.; Razansky, D., Molecular Imaging by Means of Multispectral 

Optoacoustic Tomography (MSOT). Chemical Reviews 2010, 110 (5), 2783-2794. 

8 Caravan, P.; Ellison, J. J.; McMurry, T. J.; Lauffer, R. B., Gadolinium(III) chelates as MRI 

contrast agents: Structure, dynamics, and applications. Chemical Reviews 1999, 99 (9), 

2293-2352. 

9 Bulte, J. W. M., The chemistry of contrast agents in medical magnetic resonance 

imaging. edited by A. E. Merbach and E. Toth. Wiley, Chichester, 2001, £135. NMR in 

Biomedicine 2004, 17 (4), 210-210. 

10 Terreno, E.; Castelli, D. D.; Viale, A.; Aime, S., Challenges for Molecular Magnetic 

Resonance Imaging. Chemical Reviews 2010, 110 (5), 3019-3042. 

11 Chen, K.-J.; Wolahan, S. M.; Wang, H.; Hsu, C.-H.; Chang, H.-W.; Durazo, A.; Hwang, 

L.-P.; Garcia, M. A.; Jiang, Z. K.; Wu, L.; Lin, Y.-Y.; Tseng, H.-R., A small MRI contrast 

agent library of gadolinium(III)-encapsulated supramolecular nanoparticles for improved 

relaxivity and sensitivity. Biomaterials 2011, 32 (8), 2160-2165. 

12 Raymond, K. N.; Pierre, V. C., Next generation, high relaxivity gadolinium MRI agents. 

Bioconjugate Chemistry 2005, 16 (1), 3-8. 

13 Jennings, L. E.; Long, N. J., 'Two is better than one'-probes for dual-modality molecular 

imaging. Chemical Communications 2009, (24), 3511-3524. 

14 Louie, A. Y., Multimodality Imaging Probes: Design and Challenges. Chemical Reviews 

2010, 110 (5), 3146-3195. 

163



15 Modo, M.; Cash, D.; Mellodew, K.; Williams, S. C. R.; Fraser, S. E.; Meade, T. J.; Price, 

J.; Hodges, H., Tracking transplanted stem cell migration using bifunctional, contrast 

agent-enhanced, magnetic resonance imaging. Neuroimage 2002, 17 (2), 803-811. 

16 Mishra, A.; Pfeuffer, J.; Mishra, R.; Engelmann, J.; Mishra, A. K.; Ugurbil, K.; Logothetis, 

N. K., A new class of Gd-based DO3A-ethylamine-derived targeted contrast agents for 

MR and optical imaging. Bioconjugate Chemistry 2006, 17 (3), 773-780. 

17 Bernhard, C.; Goze, C.; Rousselin, Y.; Denat, F., First bodipy-DOTA derivatives as 

probes for bimodal imaging. Chemical Communications 2010, 46 (43), 8267-8269. 

18 Kotkova, Z.; Kotek, J.; Jirak, D.; Jendelova, P.; Herynek, V.; Berkova, Z.; Hermann, P.; 

Lukes, I., Cyclodextrin-Based Bimodal Fluorescence/MRI Contrast Agents: An Efficient 

Approach to Cellular Imaging. Chemistry-a European Journal 2010, 16 (33), 10094- 

10102. 

19 Lowe, M. P.; Parker, D.; Reany, O.; Aime, S.; Botta, M.; Castellano, G.; Gianolio, E.; 

Pagliarin, R., pH-dependent modulation of relaxivity and luminescence in macrocyclic 

gadolinium and europium complexes based on reversible intramolecular sulfonamide 

ligation. Journal of the American Chemical Society 2001, 123 (31), 7601-7609. 

20 Crich, S. G.; Biancone, L.; Cantaluppi, V.; Esposito, D. D. G.; Russo, S.; Camussi, G.; 

Aime, S., Improved route for the visualization of stem cells labeled with a Gd-/Eu-chelate 

as dual (MRI and fluorescence) agent. Magnetic Resonance in Medicine 2004, 51 (5), 

938-944. 

21 Nasso, I.; Galaup, C.; Havas, F.; Tisnes, P.; Picard, C.; Laurent, S.; Elst, L. V.; Muller, R. 

N., Bimodal system (luminophore and paramagnetic contrastophore) derived from Ln(III) 

complexes based on a bipyridine-containing macrocyclic ligand. Inorganic Chemistry 

2005, 44 (23), 8293-8305. 

22 Nonat, A.; Gateau, C.; Fries, P. H.; Mazzanti, M., Lanthanide complexes of a picolinate 

ligand derived from 1,4,7-triazacyclononane with potential application in magnetic 

resonance imaging and time-resolved luminescence imaging. Chemistry-a European 

Journal 2006, 12 (27), 7133-7150. 

23 Picard, C.; Geum, N.; Nasso, I.; Mestre, B.; Tisnes, P.; Laurent, S.; Muller, R. N.; Vander 

Elst, L., A dual lanthanide probe suitable for optical (Tb3+ luminescence) and magnetic 

resonance imaging (Gd3+ relaxometry). Bioorganic & Medicinal Chemistry Letters 2006, 

16 (20), 5309-5312. 

24 Laurent, S.; Elst, L. V.; Wautier, M.; Galaup, C.; Muller, R. N.; Picard, C., In vitro 

characterization of the Gd complex of 2,6-pyridinediylbis(methylene nitrilo) tetraacetic 

acid (PMN-tetraacetic acid) and of its Eu analogue, suitable bimodal contrast agents for 

MRI and optical imaging. Bioorganic & Medicinal Chemistry Letters 2007, 17 (22), 6230- 

6233. 

164



25 Pellegatti, L.; Zhang, J.; Drahos, B.; Villette, S.; Suzenet, F.; Guillaumet, G.; Petoud, S.; 

Toth, E., Pyridine-based lanthanide complexes: towards bimodal agents operating as 

near infrared luminescent and MRI reporters. Chemical Communications 2008, (48), 

6591-6593. 

26 Tallec, G.; Imbert, D.; Fries, P. H.; Mazzanti, M., Highly stable and soluble bis-aqua Gd, 

Nd, Yb complexes as potential bimodal MRI/NIR imaging agents. Dalton Transactions 

2010, 39 (40), 9490-9492. 

27 Tallec, G.; Imbert, D.; Fries, P. H.; Mazzanti, M., Highly stable and soluble bis-aqua Gd, 

Nd, Yb complexes as potential bimodal MRI/NIR imaging agents. Dalton Transactions 

2010, 39 (40), 9490-9492. 

28 Mulder, W. J. M.; Koole, R.; Brandwijk, R. J.; Storm, G.; Chin, P. T. K.; Strijkers, G. J.; 

Donega, C. D.; Nicolay, K.; Griffioen, A. W., Quantum dots with a paramagnetic coating 

as a bimodal molecular imaging probe. Nano Letters 2006, 6 (1), 1-6 

29 Jin, T.; Yoshioka, Y.; Fujii, F.; Komai, Y.; Seki, J.; Seiyama, A., Gd(3+)-functionalized 

near-infrared quantum dots for in vivo dual modal (fluorescence/magnetic resonance) 

imaging. Chemical Communications 2008, (44), 5764-5766 

30 Rieter, W. J.; Kim, J. S.; Taylor, K. M. L.; An, H.; Lin, W.; Tarrant, T.; Lin, W., Hybrid 

silica nanoparticles for multimodal Imaging. Angewandte Chemie-International Edition 

2007, 46 (20), 3680-3682. 

31 Hu, K.-W.; Jhang, F.-Y.; Su, C.-H.; Yeh, C.-S., Fabrication of Gd(2)O(CO(3))(2)center 

dot H(2)O/silica/gold hybrid particles as a bifunctional agent for MR imaging and 

photothermal destruction of cancer cells. Journal of Materials Chemistry 2009, 19 (15), 

2147-2153 

32 van Tilborg, G. A. F.; Vucic, E.; Strijkers, G. J.; Cormode, D. P.; Mani, V.; Skajaa, T.; 

Reutelingsperger, C. P. M.; Fayad, Z. A.; Mulder, W. J. M.; Nicolay, K., Annexin A5- 

Functionalized Bimodal Nanoparticles for MRI and Fluorescence Imaging of 

Atherosclerotic Plaques. Bioconjugate Chemistry 2010, 21 (10), 1794-1803. 

33 Choo, E. S. G.; Tang, X.; Sheng, Y.; Shuter, B.; Xue, J., Controlled loading of 

superparamagnetic nanoparticles in fluorescent nanogels as effective T(2)-weighted 

MRI contrast agents. Journal of Materials Chemistry 2011, 21 (7), 2310-2319. 

34 Ke, J.-H.; Lin, J.-J.; Carey, J. R.; Chen, J.-S.; Chen, C.-Y.; Wang, L.-F., A specific tumor- 

targeting magnetofluorescent nanoprobe for dual-modality molecular imaging. 

Biomaterials 2010, 31 (7), 1707-1715. 

35 Bottrill, M.; Nicholas, L. K.; Long, N. J., Lanthanides in magnetic resonance imaging. 

Chemical Society Reviews 2006, 35 (6), 557-571. 

36 Runge, V. M.; Carollo, B. R.; Wolf, C. R.; Nelson, K. L.; Gelblum, D. Y., Gd DTPA: a 

review of clinical indications in central nervous system magnetic resonance imaging. 

165



Radiographics : a review publication of the Radiological Society of North America, Inc 

1989, 9 (5), 929-58. 

37 Sammes, P. G.; Yahioglu, G., Modern bioassays using metal chelates as luminescent 

probes. Natural Product Reports 1996, 13 (1), 1-28. 

38 Bulte, J. W. M.; Modo, M. M., In Nanoparticles in biomedical imaging - emerging 

technologies and applications. Springer, New York 2008. 

39 Na, H. B.; Song, I. C.; Hyeon, T., Inorganic Nanoparticles for MRI Contrast Agents. 

Advanced Materials 2009, 21 (21), 2133-2148. 

40 Taylor, K. M. L.; Kim, J. S.; Rieter, W. J.; An, H.; Lin, W.; Lin, W., Mesoporous silica 

nanospheres as highly efficient MRI contrast agents. Journal of the American Chemical 

Society 2008, 130 (7), 2154-2155. 

41 Tsai, C.-P.; Hung, Y.; Chou, Y.-H.; Huang, D.-M.; Hsiao, J.-K.; Chang, C.; Chen, Y.-C.; 

Mou, C.-Y., High-contrast paramagnetic fluorescent mesoporous silica nanorods as a 

multifunctional cell-imaging probe. Small 2008, 4 (2), 186-191. 

42 Hsiao, J.-K.; Tsai, C.-P.; Chung, T.-H.; Hung, Y.; Yao, M.; Liu, H.-M.; Mou, C.-Y.; Yang, 

C.-S.; Chen, Y.-C.; Huang, D.-M., Mesoporous silica nanoparticles as a delivery system 

of gadolinium for effective human stem cell tracking. Small 2008, 4 (9), 1445-1452. 

43 Carniato, F.; Tei, L.; Cossi, M.; Marchese, L.; Botta, M., A Chemical Strategy for the 

Relaxivity Enhancement of Gd(III) Chelates Anchored on Mesoporous Silica 

Nanoparticles. Chemistry-a European Journal 2010, 16 (35), 10727-10734. 

44 Carniato, F.; Tei, L.; Dastru, W.; Marchese, L.; Botta, M., Relaxivity modulation in Gd- 

functionalised mesoporous silicas. Chemical Communications 2009, (10), 1246-1248. 

45 Steinbacher, J. L.; Lathrop, S. A.; Cheng, K.; Hillegass, J. M.; Butnor, K.; Kauppinen, R. 

A.; Mossman, B. T.; Landry, C. C., Gd-Labeled Microparticles in MRI: In vivo Imaging of 

Microparticles After Intraperitoneal Injection. Small 2010, 6 (23), 2678-2682. 

46 Santra, S.; Bagwe, R. P.; Dutta, D.; Stanley, J. T.; Walter, G. A.; Tan, W.; Moudgil, B. 

M.; Mericle, R. A., Synthesis and characterization of fluorescent, radio-opaque, and 

paramagnetic silica nanoparticles for multimodal bioimaging applications. Advanced 

Materials 2005, 17 (18), 2165-2169. 

47 Wu, C.; Hong, J.; Guo, X.; Huang, C.; Lai, J.; Zheng, J.; Chen, J.; Mu, X.; Zhao, Y., 

Fluorescent core-shell silica nanoparticles as tunable precursors: towards encoding and 

multifunctional nano-probes. Chemical Communications 2008, (6), 750-752. 

48 Stöber, W.; Fink, A.; Bohn, E., Controlled Growth of Monodisperse Silica Spheres in 

Micron Size Range. Journal of Colloid and Interface Science 1968, 26 (1), 62-69. 

49 Rizkalla, E. N.; Choppin, G. R.; Cacheris, W., Thermodynamics NMR, and Fluorescence 

Studies for the Complexation of Trivalent Lantanides, Ca 2+ , Cu 2+ , and Zn 2+ by 

Diethylenetriaminepentaacetic Acid bis(methlamide). Inorganic Chemistry 1993, 32 (5), 

582-586. 

166



50 Sherry, A. D.; Cacheris, W. P.; Kuan, K. T., Stability-constants for Gd3+ Binding to 

model DTPA-conjugates and Proteins - Omplications for their use as Magnetic 

Resonance Contrast Agents. Magnetic Resonance in Medicine 1988, 8 (2), 180-190. 

51 Voisin, P.; Ribot, E. J.; Miraux, S.; Bouzier-Sore, A.-K.; Lahitte, J.-F.; Bouchaud, V.; 

Mornet, S.; Thiaudiere, E.; Franconi, J.-M.; Raison, L.; Labrugere, C.; Delville, M.-H., 

Use of lanthanide-grafted inorganic nanoparticles as effective contrast agents for cellular 

uptake imaging. Bioconjugate Chemistry 2007, 18 (4), 1053-1063. 

52 Kumar, R.; Roy, I.; Ohulchanskky, T. Y.; Vathy, L. A.; Bergey, E. J.; Sajjad, M.; Prasad, 

P. N., In Vivo Biodistribution and Clearance Studies Using Multimodal Organically 

Modified Silica Nanoparticles. Acs Nano 2010, 4 (2), 699-708. 

53 Lu, Y.; Yin, Y. D.; Mayers, B. T.; Xia, Y. N., Modifying the surface properties of 

superparamagnetic iron oxide nanoparticles through a sol-gel approach. Nano Letters 

2002, 2 (3), 183-186. 

54 Nozawa, K.; Gailhanou, H.; Raison, L.; Panizza, P.; Ushiki, H.; Sellier, E.; Delville, J. P.; 

Delville, M. H., Smart control of monodisperse Sotber silica particles: Effect of reactant 

addition rate on growth process. Langmuir 2005, 21 (4), 1516-1523. 

55 Dehaan, J. W.; Vandenbogaert, H. M.; Ponjee, J. J.; Vandeven, L. J. M., 

Characterization of modified silica powders by fourier transform infrared spectroscopy 

and cross-polarization magic angle spinning NMR. . Journal of Colloid and Interface 

Science 1986, 110 (2), 591-600. 

56 Sakthivel, A.; Hijazi, A. K.; Al Hmaideen, A. I.; Kuehn, F. E., Grafting of Cu(NCCH3)(6) 

B{C6H3(m-CF3)(2)}(4) (2) on the surface of aminosilane modified SBA-15. Microporous 

and Mesoporous Materials 2006, 96 (1-3), 293-300. 

57 Sakthivel, A.; Zhao, J.; Kuhn, F. E., Grafting of the eta(5)-CPMo(CO)(3) moiety on pure 

and surface modified SBA-15 molecular sieves. Microporous and Mesoporous Materials 

2005, 86 (1-3), 341-348. 

58 Harder, R.; Chaberek, S., The Interaction of Rare Earth Ions with 

Diethylenetriaminepentaacitic Acid. Journal of Inorganic & Nuclear Chemistry 1959, 11 

(3), 197-209. 

59 Lauffer, R. B., Paramagnetic Metal-complexes as Water Proton Relaxation Agents for 

NMR Imaging - Theory and Design. Chemical Reviews 1987, 87 (5), 901-927. 

60 S.L.C. Pinho, H. Faneca, C.F.G.C. Geraldes, M.H. Delville, L.D. Carlos, J.Rocha, 

Biomaterials, in press. 

61 Nunes, S. C.; Bermudez, V. D. Z.; Cybinska, J.; Ferreira, R. A. S.; Legendziewicz, J.; 

Carlos, L. D.; Silva, M. M.; Smith, M. J.; Ostrovskii, D.; Rocha, J., Structure and 

photoluminescent features of di-amide cross-linked alkylene siloxane hybrids. Journal of 

Materials Chemistry 2005, 15 (35-36), 3876-3886. 

167



62 Schulz, H.; Pratsinis, S. E.; Ruegger, H.; Zimmermann, J.; Klapdohr, S.; Salz, U., 

Surface functionalization of radiopaque Ta2O5/SiO2. Colloids and Surfaces a- 

Physicochemical and Engineering Aspects 2008, 315 (1-3), 79-88. 

63 Albert, K.; Brindle, R.; Schmid, J.; Buszewski, B.; Bayer, E., CP/MAS NMR 

Investigations of Silica-Gel Surfaces Modified with Aminopropylsilane. Chromatographia 

1994, 38 (5-6), 283-290. 

64 Plueddemann, E., Interfaces in polymer matrix composites In: Brautman LJ, Krock RH 

editors. Composite Materials. Academic Press, New York 1974, 6. 

65 Boerio, F. J.; Schoenlein, L. H.; Greivenkamp, J. E., Adsorption of Gamma- 

Aminopropyltriethoxysliane onto Bulk Iron from Aqueous Solutions. J Appl Polym Sci 

1978, 22 (1), 203-213. 

66 Anderson, H. R.; Fowkes, F. M.; Hielscher, F. H., Electron Donor-Acceptor Properties of 

Thin Polymer Films on Silicon 2. Tetrafluoroethylene Polymerized by RF Glow Discharge 

Techniques. Journal of Polymer Science Part B-Polymer Physics 1976, 14 (5), 879-895. 

67 Moses, P. R.; Wier, L. M.; Lennox, J. C.; Finklea, H. O.; Lenhard, J. R.; Murray, R. W., 

Chemically Modified Electrodes 9.X_Ray Photoelectron-Spectroscopy of Alkylamine- 

silanes Bound to Metal-oxide Electrodes. Analytical Chemistry 1978, 50 (4), 576-585. 

68 Ishida, H.; Chiang, C. H.; Koenig, J. L., The structure of aminofunctional silane coupling 

agents: 1. γ-aminopropyltriethoxysilane and its analogues. Polymer 1982, 23 (2), 251- 

257. 

69 Jakob, A. M.; Schmedake, T. A., A novel approach to monodisperse, luminescent silica 

spheres. Chemistry of Materials 2006, 18 (14), 3173-3175. 

70 Wang, L.; Estevez, M. C.; O'Donoghue, M.; Tan, W., Fluorophore-free luminescent 

organosilica nanoparticles. Langmuir 2008, 24 (5), 1635-1639. 

71 Nobre, S. S.; Lima, P. P.; Mafra, L.; Ferreira, R. A. S.; Freire, R. O.; Fu, L.; Pischel, U.; 

Bermudez, V. d. Z.; Malta, O. L.; Carlos, L. D., Energy transfer and emission quantum 

yields of organic-inorganic hybrids lacking metal activator centers. Journal of Physical 

Chemistry C 2007, 111 (8), 3275-3284. 

72 Bermudez, V. D.; Carlos, L. D.; Duarte, M. C.; Silva, M. M.; Silva, C. J. R.; Smith, M. J.; 

Assuncao, M.; Alcacer, L., A novel class of luminescent polymers obtained by the sol-gel 

approach. Journal of Alloys and Compounds 1998, 275, 21-26. 

73 Ferreira, R. A. S.; Nobre, S. S.; Granadeiro, C. M.; Nogueira, H. I. S.; Carlos, L. D.; 

Malta, O. L., A theoretical interpretation of the abnormal D-5(0)-> F-7(4) intensity based 

on the Eu 3+ local coordination in the Na-9 EuW10O36 center dot 14H(2)O 

polyoxometalate. J Lumin 2006, 121 (2), 561-567. 

74 Carlos, L. D.; Ferreira, R. A. S.; Bermudez, V. d. Z.; Ribeiro, S. J. L., Lanthanide- 

Containing Light-Emitting Organic-Inorganic Hybrids: A Bet on the Future. Advanced 

Materials 2009, 21 (5), 509-534. 

168



75 Bryden, C. C.; Reilley, C. N., Europium luminescence lifetimes and spectra for 

evaluation of 11 europium complexes as aqueous shift reagents for nuclear magnetic 

resonance spectrometry. Analytical Chemistry 1982, 54 (4), 610-615. 

76 Malta, O. L.; dos Santos, M. A. C.; Thompson, L. C.; Ito, N. K., Intensity parameters of 

4f-4f transitions in the Eu(dipivaloylmethanate)(3) 1,10-phenanthroline complex. J Lumin 

1996, 69 (2), 77-84. 

77 Carlos, L. D.; Messaddeq, Y.; Brito, H. F.; Ferreira, R. A. S.; Bermudez, V. D.; Ribeiro, 

S. J. L., Full-color phosphors from europium(III)-based organosilicates. Advanced 

Materials 2000, 12 (8), 594-598. 

78 Lima, P. P.; Ferreira, R. A. S.; Freire, R. O.; Paz, F. A. A.; Fu, L. S.; Alves, S.; Carlos, L. 

D.; Malta, O. L., Spectroscopic study of a UV-photostable organic-inorganic hybrids 

incorporating an Eu(3+) beta-diketonate complex. Chemphyschem 2006, 7 (3), 735-746. 

79 Carnall, W.; Crosswhite, H., In energy level structure and transition probabilities of the 

trivalent lanthanides in LaF3. Argonne Natl Lab. Rept 1977. 

80 Norek, M.; Peters, J. A., MRI contrast agents based on dysprosium or holmium. 

Progress in Nuclear Magnetic Resonance Spectroscopy 2011, 59 (1), 64-82. 

81 Gillis, P.; Moiny, F.; Brooks, R. A., On T-2-shortening by strongly magnetized spheres: A 

partial refocusing model. Magnetic Resonance in Medicine 2002, 47 (2), 257-263 

82 V.A. Runge, B.R. Carollo, C.R. Wolf, K.L. Nelson, RadioGraphics 1989, 9, 929-958. 

169


Core-Shell Nanoparticles for Bimodal-Imaging Contrast Agents 

170 

5. 

Core-Shell Nanoparticles for 

Bimodal-Imaging Contrast 

Agents


Core-shell Nanoparticles for Bimodal-Imaging Contrast Agents 

5.1. Introduction 172 


5.3. Results and Discussions 178 



5.3.3. Cell Imaging 199 



171



172 

Chapter published as original article: 

Pinho S.L.C., Pereira G.A., Voisin P., Kassem J., Bouchaud V., Etienne L., Peters J.A., 

Carlos L.D., Mornet S., Geraldes C.F.G.C., Rocha J, Delville M-H. 

Fine tuning of the relaxometry of γ -Fe2O3@SiO2 nanoparticles by tweaking the silica 

coating thickness. 

ACSNano 4 (9) 5339 – 5349 (2010) DOI: 10.1021/nn101129r 

Pinho S.L.C., Laurent S., Rocha J., Roch A., Delville M-H., Carlos L.D., Elst L.V., 

Muller R.N., Geraldes F.G.C. Relaxometric studies of γ -Fe2O3@SiO2 core shell 

nanoparticles: when the coating matters. 

5.1. INTRODUCTION 

Submitted J. Phys. Chem. C (2011) 

Nanoparticles (NPs) made of inorganic or organic materials exhibit many novel 

properties compared with the bulk materials. 1 Magnetic NPs have unique properties 

such as superparamagnetism, high coercivity, low Curie temperature, high magnetic 

susceptibility, etc. 2 Magnetic NPs are of great interest in a broad range of disciplines, 

from magnetic fluids to data storage, catalysis, 3 and bio-applications. 4 Examples of 

applications of NPs in the study of biology and biomedicine are magnetic 

bioseparation, 5 cell sorting, 67 detection of biological entities, 8 clinical diagnosis and 

therapy (such as MRI, magnetic resonance imaging), 9-18 MFH (magnetic fluid 

hyperthermia) 19 targeted drug delivery, 20-23 immunoassays, 24 and biomacromolecule 

purification. 25 Magnetic iron oxide NPs play an important role in these applications and 

they have been used in in vitro diagnosis for about 50 years. 26 In the last decade, 

numerous investigations have been carried out in the field of magnetic NPs, 27 especially 

on magnetite and maghemite due to their biocompatibility, FDA approval 28 and absence 

of toxicity. 29-31 

The control of the NPs size, shape, stability, and dispersibility in specific solvents 

is a technological challenge. Bio-applications, for example, require water-solubility and 

colloidal stability. However, most reported syntheses of high-quality NPs of metals, 32,33 

semiconductors, 34,35 and metal oxides 36-38 involve non-aqueous solvents and coating 

with monolayers of hydrophobic surfactants. Several strategies to tackle these



challenges have been formulated, 39 such as i) polymer coating, 40,41 ii) exchanging the 

original hydrophobic stabilizer with dendrons, 42,43 thiols or even oligomeric 

phosphines 44 and iii) silica coatings. 45-53 9,10,11,12,13,14,15,16,17,18 19 20,21,22,23242526 27 28 

, 

29 , 30 , 3132 , 33 34 , 35 36 , 37 , 38 39 40 , 41 42,4344 45,46,47,48,49,50,51,52,53 

In order to expand the scope of the iron oxide NPs in biological applications, 

biomolecules have been employed as coatings, such as amino acids, 54 vitamins, 55,56 

proteins, 57 antibodies, 58,59 polypeptides, 60 biotin, avidin 61 and saccharides. 62 However, 

silica coating remains one of the most popular and well-known techniques for NP 

surface modification, because the resulting cross-linked silica shell protects the core 

from the environment and the other way around. The silica coating also provides 

colloidal stability in biological solutions by avoiding inter-particle interactions and 

agglomeration. Furthermore, it can act as an anchor for the binding of biological 

vectors at the NPs surface. 63 Although there are several publications concerning silica 

coatings, only a few methods have been reported for the preparation of water-soluble 

silica-coated NPs with a high colloidal stability and with sizes below 20 nm. 45,50,53 

Particles with tunable size are important when considering biomedical 

applications. While small NPs exhibit reduced nonspecific interactions, minimal steric 

effects, and high clearance rates, 64 larger NPs are subjected to internalization by 

macrophages. The thickness of the silica shell has also a strong influence on the 

physical properties of the NPs, especially in terms of contrast agent efficacy for 

magnetic resonance imaging. We describe the synthesis of γ-Fe2O3@SiO2, core-shell 

NPs with tuned shell thicknesses. These particles were characterized by Transmission 

Electron Microscopy (TEM), zeta potential determinations, Diffuse Reflectance Infrared 

Fourier-Transform and Nuclear Magnetic Resonance. The longitudinal (T1) and 

transversal (T2) relaxation times of aqueous suspensions of the prepared particles were 

measured, and their cytotoxicity was investigated. We show that the shell thickness of 

γ-Fe2O3@SiO2 NPs has a significant impact on their relaxivities. This silica layer 

exhibits two regions around the core, one, which is porous to water, and a second one, 

which is not. 

173



5.2. EXPERIMENTAL PROCEDURES 

Materials and purification methods 

Iron (III) chloride hexahydrate (98%), iron (II) chloride tetrahydrate (99%), iron (III) 

nitrate nonahydrate (99%), tetraethoxysilane (TEOS) (98%), and citric acid (99.5%) 

were purchased from Aldrich. Absolute ethanol (J.T. Baker) and ammonia (Carlo Erba) 

were used as received. All other reagents were of analytical grade. All the experiments 

were performed in deionized Milli-Q water. 

Preparation of the maghemite ferrofluid suspension 

The aqueous maghemite suspension was synthesized by precipitation from iron 

chlorides. 65,66 Briefly, the Fe3O4 precipitate (black dispersion of magnetite), obtained by 

alkalinization of the FeCl2 and FeCl3 (Fe 2+ /Fe 3+ = 1/2) aqueous mixture, was 

successively oxidized with 2M HNO3 and 0.33 M Fe(NO3)3 · 9H2O solutions at 100 ºC 

in order to obtain particles with a Fe 2+ /Fe 3+ ratio lower than 0.05. With this oxidation 

process, magnetite is converted into maghemite. The brown dispersion was peptized in 

a 2 M HNO3 solution under vigorous stirring in order to create positive surface charges. 

The acidic precipitate was isolated by magnetic separation, washed with acetone and 

dispersed at pH ~ 2.5 in water with nitric acid. The iron concentration was determined 

by volumetric titration as well as by ICP measurements and the average particle size, 

as determined by transmission electron microscopy (TEM), was 10 2 nm 

Preparation of the maghemite ferrofluid core-shell suspension 

The selected method was derived from the so-called Stöber process 67 widely 

used for the synthesis of silica beads with diameters from a few tens to a few hundreds 

of nanometers. 68 It is based on the hydrolysis/condensation of tetraethoxysilane 

(TEOS) catalyzed by ammonia in alcoholic media. The surface of γ-Fe2O3 NPs was 

activated by acidic treatment: where 7.55 mL of γ-Fe2O3 colloidal suspension 

(concentration 74.4 g/L) were dispersed in 40 mL of 0.01 M citric acid. They were 

174



isolated by decantation on a magnet. The particles were dispersed in 12 mL of water 

and peptization was performed by adding 20 µL of ammonia. Then, the alkaline sol of 

citrated- γ-Fe2O3 NPs was poured in 1 L of ethanol-water-ammonia solution 75/23.5/1.5 

v/v/v %, to obtained a 0.561 g/L concentration. The appropriate amounts of TEOS 

precursors were added to the dispersion under mild stirring to reach the targeted shell 

thickness. They were set to comply with the desired thickness of the silica shell 

according to equation 5.1 and added in multiple steps. 

 

 

 

M 

4 

 

 

3 3 

r e r 

SiO2 

TEOS 

VTEOS NPart 

 

shell 

TEOS M 

SiO2 

3 

Where shell 

e is the shell thickness (the difference 

to the volume of the silica shell 2 

weight of SiO2; TEOS 

TEOS 

TEOS 

SiO2 

, SiO 

175 

3 3 

r eshell 

r 

(5.1) 

4 

then corresponding 

3 

V is the density and 2 

M the molecular 

V , , M are the volume, density and molecular weight of 

TEOS; Npart. is the number of γ-Fe2O3 NPs. The very first amount of added TEOS (763 

μL) corresponds to the smallest observable silica shell thickness (roughly 1 nm). Then, 

after 12h of the reaction, 200 mL of this solution were stocked for analysis and 

replaced by the same amount of reaction medium. For the following step, the resulting 

solution was added with the necessary amount of TEOS to increase the shell 

thickness, and left to react another 12h. 200 mL of this solution were also stocked for 

analysis and replaced by the same amount of reaction medium. This procedure was 

used to get thicker shell sizes (the number of particles in each volume being 

recalculated to estimate the right amount of TEOS). Under these conditions, no 

secondary nucleation was observed, which is in agreement with the results reported by 

Chen et al. 69 Two series of core shell NPs were synthesised and will be denoted series 

A and B. The difference between these two series was size of the NPs, in series B a 

better range of thinner coating was produced in order to better understand the 

influence of the silica shell. 

SiO



Particle characterization 

TEM was performed at room temperature on a JEOL JEM-2000 FX transmission 

electron microscope using an accelerating voltage of 200 kV. Drops of diluted 

dispersions of core-shell were air-dried on carbon films deposited on 200-mesh copper 

grids. The excess liquid was blotted with filter paper. The Diffuse Reflectance Infrared 

Fourier-Transform (DRIFT) spectra were recorded on a Bruker IFS Equinox 55FTIR 

spectrometer (signal averaging 64 scans at a resolution of 4 cm -1 in KBr pellets 

containing ca. 2 mass % of material). The zeta potential of the NPs was assessed 

using a Zetasizer 3000HSA setup (Malvern Instruments) equipped with a He- Ne laser 

(50 mW, 532 nm). The zeta potential measurement based on laser Doppler 

interferometry was used to measure the electrophoretic mobility of NPs. Measurements 

were performed for 20s using a standard capillary electrophoresis cell. The dielectric 

constant was set to 80.4 and the Smoluchowsky constant f(ka) was 1.5. The iron 

content has been measured by inductively coupled plasma / optical emission 

spectrometry ICP/OES (ES720, Varian) equipped with a crossflow nebulizer. A 1 g/L 

iron solution was used to prepare the standard solutions (SCP Science to Paris) and 

was used as internal standard to evaluate the instrumental drift. 

and 2 

Measurements of water proton longitudinal and transverse relaxation times ( 1 

T respectively) of aqueous suspensions of the NPs were carried out at 20 MHz 

on a Bruker Minispec mq20 relaxometer and at 499.83 MHz (B0= 11.7 T) on a Varian 

Unity 500 NMR spectrometer at 25 ºC. The 1 

the inversion recovery pulse sequence, while the 2 

176 

T relaxation times were measured using 

T relaxation times were measured 

using a Carr-Purcell-Meiboom-Gill (CPMG) pulse sequence and varying the time 

interval between two consecutive refocusing pulses ( CP 

applied. The values of 

) in the train of 180 o pulses 

* 

T 2 , the transverse relaxation time in the presence local field 

inhomogeneities, were obtained from the water spectral line widths. All the 

experimental values were corrected for the diamagnetic contribution using aqueous 

suspensions of hollow silica NPs, to obtain each paramagnetic contribution. These 

hollow shells where prepared by dilution of the core by addition of concentrated HCl. 

The magnetization measurements were performed on a known amount of 

ferrofluid using a vibrating sample magnetometer VSM-NUVO (MOLSPIN, Newcastle 

T



Upon Tyne, U.K.). Magnetometry of NPs allows the determination of the saturation 

magnetisation (Msat) and the radius of the superparamagnetic crystals (r) by fitting the 

1 

data with a Langevin function ( L( 

x) 

coth( x) 

where 

x 

magnetic moment at saturation, B0 as the magnetic field, b 

and T as the temperature. 

177 

satB0 x ) with sat 

k T 

b 

as the 

k as the Boltzmann constant 

The NMRD profiles were recorded at 37 °C over a magnetic field range from 0.24 

mT to 0.24 T on a field cycling Stelar relaxometer (Mede, Italy). Additional longitudinal 

(R1) and transverse (R2) relaxation rate measurements were performed at 0.47 and 

1.41 T on Minispec mq20 and Minispec mq60 relaxometers, respectively (Bruker, 

Karlsruhe, Germany). 

Toxicity tests 

Cytotoxicity of the γ-Fe2O3@SiO2 NPs was tested by counting the cells in a 

Malassez chamber and using the MTT assay to evaluate the cell viability after the NPs 

preparation process. The core-shell NPs FF, 1A, 2A, 5A and 7A had diameters ranging 

between 10 and 143 nm. Briefly, microglial cell lines were seeded at the rate of ca. 16 x 

10 3 cells/cm 2 in 35 mm diameter plates and allowed to attach for 24 h. The cells were 

then incubated for 0, 45 min, 24 h, 48 h, 72 h, 96 h, 120 h and 144 h in 1 mL of culture 

medium for control cells and supplemented with 60 µL of different NPs (0.16 mM) for 

treated cells. MTT and counting assays were performed as duplicate for each condition 

and the data were averaged. After incubation, cells were scraped from the dishes, then 

stained with trypan blue and counted with a haemocytometer. The MTT assay is a 

colorimetric assay that measures the reduction of yellow 3-(4,5-dimethythiazol-2-yl)- 

2,5-diphenyl tetrazolium bromide (MTT) by dehydrogenase mostly from mitochondria. 

The MTT enters the cells and passes into the mitochondria, where it is reduced to an 

insoluble, colored (dark purple) formazan product. After cell culturing in the presence of 

NPs, 260 μL of the MTT solution in culture medium (0.5 mg/mL) was added into each 

well. The plate was then incubated at 37 ºC in 5 % CO2 for 45 min. The medium was 

removed and 1 mL of PBS solution was added, then cells were scraped and 

centrifuged at 1000 rpm for 5 min. The supernatant was removed and 1 mL of dimethyl 

sulfoxide (DMSO) was added to the pellets to dissolve the formazan crystals and then 

it was centrifuged again at 1000 rpm for 5 min. Supernatants were taken and their



absorbance was measured with a U-2800A (UV-VIS) spectrophotometer (Hitachi, 

Japan) at 570 nm. Since reduction of MTT can only occur in metabolically active cells, 

the level of activity is an estimation of the viability of the cells as compared to untreated 

cells. The cell viability (%) was calculated according to equation 5.2: 

Cell viability % = OD570(sample) / OD570(control) × 100 (5.2) 

where OD570(sample) represents the optical density of the wells treated with various 

iron sizes, and OD570(control) represents that of the wells treated with medium culture. 

5.3. RESULTS AND DISCUSSION 

5.3.1. Characterization of Nanoparticles 

The aqueous maghemite suspension was synthesized by basic precipitation from 

iron chlorides, followed by complete oxidation of the magnetite material. For the 

coating, a polymerization of silane monomers in the presence of the NPs under Stöber 

conditions 67,70 was performed. This procedure is widely used since it provides uniform 

silica coating with a controllable thickness. Stöber’s conditions involve alcohol-water- 

ammonia as the medium and tetraethoxysilane (TEOS) as the silane monomer. A pre- 

activation of the surface of the NPs through acidic treatment was found to improve the 

silica coating, leading to a simple and highly reproducible method for producing 

monodispersed water-soluble stable colloidal NPs with silica shells whose thickness is 

tunable in the range 2-70 nm. 

To tune the silica shell thickness, the required amount of TEOS was calculated 

from the initial and the desired final particle size, 71 , 72 taking into account the number of 

γ-Fe2O3 NPs, Npart., by means of equation 5.1. 

The estimated and experimental thicknesses of the silica coatings are collected in 

Table 5.1, while Figure 5.1 displays the TEM images obtained at various stages of the 

NPs synthesis for series A and B. 

178



TEM showed that spherical core-shell (γ-Fe2O3@SiO2) NPs with different shell 

sizes were obtained; as clearly evidenced by these images, all the γ-Fe2O3 particles 

were surrounded by the silica layer. The scheme on the right of the lower row of the 

images defines the measured size or diameter (d) of the NPs, and their silica shell 

thickness (t). The average thickness of silica shells was determined from these images 

by measurements in four directions for each particle and at least 100 particles per γ- 

Fe2O3@SiO2 sample, showing that the size dispersion of the particles is very small. 

Figure 5.2 shows the relationship between the obtained shell thicknesses of the 

series A NPs and the expected ones through calculations. They are proportional to the 

amount of TEOS added during the preparation. Note the deviation from a slope of 1, 

which is significant of the errors taking place at each step as well as some aggregation 

of the maghemite particles, as can be detected by TEM. 

The ratio of Fe/Si was determined for the series B NPs by ICP (Table 5.2). When 

the diameter of the NPs increases, the concentration of Fe with respect to Si 

decreases, as expected (Figure 5.3). 

Table 5.1. Synthesis of maghemite core-shell (γ-Fe2O3@SiO2) NPs for 

series A: comparison between estimated and experimental values of shell 

thicknesses. 

Sample 

Estimated shell 

thickness (nm) a 

Experimental shell 

thickness (nm) 

Experimental 

diameter (nm) 

FF@SiO2_1A 1 ± 1 2 ± 1 14 ± 2 

FF@SiO2_2A 4 ± 1 8 ± 2 27 ± 5 

FF@SiO2_3A 10 ± 1 15 ± 4 40 ± 8 

FF@SiO2_4A 18 ± 2 20 ± 4 50 ± 7 

FF@SiO2_5A 23 ± 3 28 ± 4 66 ± 8 

FF@SiO2_6A 31 ± 3 52 ± 6 114 ± 14 

FF@SiO2_7A 56 ± 6 67 ± 5 145 ± 10 

179



a-I) a-I) a-I) a-I) a-I) a-I) a-I) 

FF FF FF FF FF FF FF (bare (bare (bare (bare (bare (bare (bare iron iron iron iron iron iron iron NPs) NPs) NPs) NPs) NPs) NPs) NPs) FF@SiO2_1A FF@SiO2_1A FF@SiO2_1A FF@SiO2_1A FF@SiO2_1A FF@SiO2_1A FF@SiO2_1A FF@SiO2_2A FF@SiO2_2A FF@SiO2_2A FF@SiO2_2A FF@SiO2_2A FF@SiO2_2A FF@SiO2_2A FF@SiO2_3A FF@SiO2_3A FF@SiO2_3A FF@SiO2_3A FF@SiO2_3A FF@SiO2_3A FF@SiO2_3A FF@SiO2_4A 

FF@SiO2_4A 

FF@SiO2_4A 

FF@SiO2_4A 

FF@SiO2_4A 

FF@SiO2_4A 

FF@SiO2_4A 

a-II) a-II) a-II) a-II) a-II) a-II) a-II) 

10 10 10 10 10 10 10 nm nm nm nm nm nm nm 

14 14 14 14 14 14 14 nm nm nm nm nm nm nm (2 (2 (2 (2 (2 (2 (2 nm) nm) nm) nm) nm) nm) nm) 27 27 27 27 27 27 27 nm nm nm nm nm nm nm (5 (5 (5 (5 (5 8.5 8.5 nm) nm) nm) nm) nm) nm nm ) ) 40 40 40 40 40 40 40 nm nm nm nm nm nm nm (8 (8 15 (8 (8 (8 15 nm) nm) nm) nm) nm) nm nm ) ) 50 50 50 50 50 50 50 nm nm nm nm nm nm nm (7 (7 20 (7 (7 (7 20 nm) nm) nm) nm) nm) nm nm ) ) 

FF@SiO2_5A FF@SiO2_5A FF@SiO2_5A FF@SiO2_5A FF@SiO2_5A FF@SiO2_5A FF@SiO2_5A FF@SiO2_6A FF@SiO2_6A FF@SiO2_6A FF@SiO2_6A FF@SiO2_6A FF@SiO2_6A FF@SiO2_6A FF@SiO2_7A 

FF@SiO2_7A 

FF@SiO2_7A 

FF@SiO2_7A 

FF@SiO2_7A 

FF@SiO2_7A 

FF@SiO2_7A 

66 66 66 66 66 66 66 nm nm nm nm nm nm nm ( ( ( ( ( ( ( nm) nm) 28nm nm) nm) nm) 28nm ) ) 114 114 114 114 114 114 114 nm nm nm nm nm nm nm (14 (14 (54 (14 (14 (14 (54 nm) nm) nm) nm) nm) nm) nm) 145 145 145 145 145 145 145 nm nm nm nm nm nm nm (10 (10 (167 (10 (10 (10 (167 nm) nm) nm) nm) nm) nm) nm) 

FF FF FF FF FF FF FF (bare (bare (bare (bare (bare (bare (bare iron iron iron iron iron iron iron NPs) NPs) NPs) NPs) NPs) NPs) NPs) FF@SiO2_1B FF@SiO2_1B FF@SiO2_1B FF@SiO2_1B FF@SiO2_1B FF@SiO2_1B FF@SiO2_1B FF@SiO2_2B FF@SiO2_2B FF@SiO2_2B FF@SiO2_2B FF@SiO2_2B FF@SiO2_2B FF@SiO2_2B FF@SiO2_3B FF@SiO2_3B FF@SiO2_3B FF@SiO2_3B FF@SiO2_3B FF@SiO2_3B FF@SiO2_3B FF@SiO2_4B 

FF@SiO2_4B 

FF@SiO2_4B 

FF@SiO2_4B 

FF@SiO2_4B 

FF@SiO2_4B 

FF@SiO2_4B 

9.6 9.6 9.6 9.6 9.6 nm nm nm nm nm (-- (-- (-- (-- (-- nm) nm) nm) nm) nm) 

50 50 50 50 50 nm nm nm nm nm 

11.2 11.2 11.2 11.2 11.2 nm nm nm nm nm (0.8 (0.8 (0.8 (0.8 (0.8 nm) nm) nm) nm) nm) 

50 50 50 50 50 nm nm nm nm nm 


FF@SiO2_5B FF@SiO2_5B FF@SiO2_5B FF@SiO2_5B FF@SiO2_5B FF@SiO2_5B FF@SiO2_5B FF@SiO2_6B FF@SiO2_6B FF@SiO2_6B FF@SiO2_6B FF@SiO2_6B FF@SiO2_6B FF@SiO2_6B FF@SiO2_7B 

FF@SiO2_7B 

FF@SiO2_7B 

FF@SiO2_7B 

FF@SiO2_7B 

FF@SiO2_7B 

FF@SiO2_7B 


100 100 100 100 100 nm nm nm nm nm 

180 

50 50 50 50 50 nm nm nm nm nm 


200 200 200 200 200 nm nm nm nm nm 


50 50 50 50 50 nm nm nm nm nm 


200 200 200 200 200 nm nm nm nm nm 


dd 

dd 

d 

100 100 100 100 100 nm nm nm nm nm 

tt tt t d d d d d ( ( ( ( ( t t t t t ) ) ) ) )



b-I) b-I) b-I) b-I) 

Population Population (%) (%) 

b-II) b-II) b-II) b-II) 

Population Population Population (%) (%) (%) 

40 40 

35 35 

30 30 

25 25 

20 20 

15 15 

10 10 

55 

00 

40 40 40 

30 30 30 

20 20 20 

10 10 10 

00 

0 

[8 [8 -10[ -10[ 

[6 [6 [6 -8[ -8[ -8[ 

[10 [10 -12[ -12[ 

[12 [12 -14[ -14[ 

[8 [8 [8 -10[ -10[ -10[ 

[14 [14 -16[ -16[ 

[10 [10 [10 -12[ -12[ -12[ 

[16 [16 -18[ -18[ 

Diameter Diameter (nm) (nm) 

Diameter Diameter Diameter (nm) (nm) (nm) 

[18 [18 -20[ -20[ 

[12 [12 [12 -14[ -14[ -14[ 

[20 [20 -22[ -22[ 

[14 [14 [14 -16[ -16[ -16[ 

[22 [22 -24[ -24[ 

181 

Population Population (%) (%) 

Population Population Population (%) (%) (%) 

30 30 

25 25 

20 20 

15 15 

10 10 

55 

00 

30 30 30 

25 25 25 

20 20 20 

15 15 15 

10 10 10 

55 

5 

00 

0 

[50 [50 -54[ -54[ 

[80 [80 [80 [80 -84[ -84[ -84[ -84[ 

[54 [54 -58[ -58[ 

[84 [84 [84 [84 -88[ -88[ -88[ -88[ 

[58 [58 -62[ -62[ 

[88 [88 [88 [88 -92[ -92[ -92[ -92[ 

[62 [62 -66[ -66[ 

[92 [92 [92 [92 -96[ -96[ -96[ -96[ 

[66 [66 -70[ -70[ 

[96 [96 [96 [96 -100[ -100[ -100[ -100[ 

[70 [70 -74[ -74[ 

Diameter Diameter (nm) (nm) 

[100 [100 [100 [100 -104[ -104[ -104[ -104[ 

Diameter Diameter Diameter (nm) (nm) (nm) 

Figure 5.1. a) TEM images showing the average size (diameter d) of 

different maghemite core-shell (γ-Fe2O3@SiO2) NPs and of their silica shell 

thickness (t): I) for series A and II) for series B; b) histograms with 

experimental size distributions and corresponding calculated normal 

cumulative distributions for the specified mean and standard deviation: I) 

samples FF@SiO2_1A (γ -Fe2O3@SiO2 14 nm) (left) and FF@SiO2_5A (γ- 

Fe2O3@SiO2 66 nm) (right) II) samples FF@SiO2_1B (γ-Fe2O3@SiO2 11.2 

nm) (left) and FF@SiO2_5B (γ-Fe2O3@SiO2 94.9 nm) (right) and 

corresponding calculated normal cumulative distributions for the specified 

mean and standard deviation 

[74 [74 -78[ -78[ 

[104 [104 [104 [104 -108[ -108[ -108[ -108[ 

[78 [78 -82[ -82[ 

[108 [108 [108 [108 -112[ -112[ -112[ -112[ 

[82 [82 -86[ -86[ 

[112 [112 [112 -116[ -116[ -116[



Measured shell thickness (nm) 

70 

60 

50 

40 

30 

20 

10 

0 

0 10 20 30 40 50 60 70 

Calculated shell thickness (nm) 

182 

' 

Slope= 1.363 0.06 

R 2 = 0.97581 

Figure 5.2. Correlation between the experimental thickness of the silica 

shell as determined by TEM and the thickness calculated with eq. 1. 

Table 5.2 Fe/Si ratio obtained by ICP and average diameter determined by 

TEM for different NPs for series B 

Sample Fe/Si Diameter (nm) 

FF 9.6 ± 1.3 

FF@SiO2_1B 1.1230 11.2 ± 1.7 

FF@SiO2_2B 0.2440 24.7 ± 3.2 

FF@SiO2_3B 0.0563 46.4 ± 8.0 

FF@SiO2_4B 0.0183 73.8 ± 8.1 

FF@SiO2_5B 0.0075 94.9 ± 7.0 

FF@SiO2_6B 0.0038 114.2 ± 12.7 

FF@SiO2_7B 0.0017 152.9 ± 13.1



Diameter nm 

180 

160 

140 

120 

100 

80 

60 

40 

20 

0 

0.0 0.2 0.4 0.6 0.8 1.0 1.2 

Fe/Si ratio 

Figure 5.3. Dependence of the core-shell NPs (series B) diameter on the 

Fe/Si ratio. 

Figure 5.4 shows the zeta potential titrations as a function of pH, and both the pH 

range of stability and the isoelectric points (IEP) of the two types of particles (2.3 for 

silica and 7.0 for γ-Fe2O3). Silica has long been used as a nonmagnetic coating 

material, in order to avoid aggregation or sedimentation of ferrofluid magnetic NPs 

because of its extraordinary stability over a wide range of polar and non-polar solvents. 

Zeta- potential (mV) 

60 

40 

20 

0 

-20 

-40 

-60 

-80 

-100 

0 2 4 6 8 10 12 14 

pH 

Figure 5.4.. Zeta potential titrations as a function of pH of γ-Fe2O3 (●) and γ- 

Fe2O3@SiO2 (●) aqueous suspensions. 

183



In particular, aqueous dispersions of silica are known to be stable over a large pH 

range (IEP at pH 2). The shift of the IEP towards lower pH values (from ~6.5-7 to ~2.5) 

upon coating (Figure 5.4) provides an additional confirmation that the coating was 

successful. The large negative zeta potential (- 80 mV) at physiological pH of the 

coated NPs suggests that the aqueous suspensions will by highly stable under in vivo 

conditions and will not flocculate at pH 7. 

DRIFT spectroscopy was also used to probe the effectiveness of the chemical 

coating of silica on the maghemite NPs (FF) (Figure 5.5). Several absorption bands in 

the DRIFT spectrum of γ-Fe2O3@SiO2 samples (Figure 5.5) are assigned to silica; and 

clearly show that this material covers the surface of the maghemite NPs. The bound Si- 

OH groups are characterized by the very broad IR absorption band in the 2800-3700 

cm -1 region whereas the so-called free Si-OH groups provide a narrow IR absorption 

band at 3630 cm -1 . The stretching band at 1635 cm -1 shows the presence of residual 

physisorbed water molecules, while the large bands centered at 1864 cm -1 , 1108 cm -1 

and 796 cm -1 are assigned to the Si-O and Si-O-Si stretching modes. 

Reflectance 

3500 3000 2500 2000 1500 1000 500 

Wavenumber (cm -1 ) 

Figure 5.4. Diffuse Reflectance IR Fourier-Transform spectra (DRIFT) of 

maghemite NPs (black), silica NPs (pink) and γ -Fe2O3@SiO2 (red) 

184



5.3.2. Relaxivity Properties 

In order to investigate the influence of the shell thickness of the silica coating on 

the MRI contrast agent (CA) efficiency of the γ-Fe2O3 NPs, the ri (i = 1,2) relaxivities 

(defined as enhancement of Ri = 1/Ti, i = 1,2, the relaxation rates per mM concentration 

of CA) of the different core-shell NPs were measured at two resonance frequencies (20 

and 500 MHz) and two temperatures (25 and 37° C). Figure 5.5 shows typical values of 

the r1 and r2 relaxivities for the aqueous suspensions of γ-Fe2O3@SiO2 NPs (series A) 

as a function of the diameter d of the NPs with a 10.0 nm diameter γ-Fe2O3 core and an 

increasing thickness of its silica layer, giving d values of 14 nm (sample 1A) to 145 nm 

(sample 7A) (Table 5.2). 

r i (s -1 mM -1 ) 

250 

200 

150 

100 

50 

0 

r 1 (s -1 mM -1 ) 

35 

30 

25 

20 

15 

10 

5 

0 

0 25 50 75 100 125 150 

0 50 100 150 200 250 

d (nm) 

185 

d (nm) 

r1 25؛C 

r2 25؛C 

r1 37؛C 

r2 37؛C 

Figure 5.5. Dependence of water relaxivities of aqueous suspensions of the 

γ-Fe2O3@SiO2 NPs on their diameter, as a result of increased silica layer 

thickness: a) inset: r1 at 20 MHz (25 ºC); b) main plot: ri (i =1, 2) at 500 MHz 

(25 ºC and 37 ºC). r2 relaxivities were measured at τCP = 1.6 ms.



Table 5.3. Parameters obtained from analysis of r2 (CP = 1.6 ms) and r2* 

values of aqueous suspensions of core-shell (γ-Fe2O3@SiO2) NPs at B0 = 

11.7 T and 25 o C. 

Sample 

Diameter 

(nm) a 

r2 

(s -1 mM -1 ) 

186 

r2* 

(s -1 mM -1 ) 

2ri (nm) 2rdif (nm) 

FF 10 ± 2 228 ± 2 230 ± 1 13 ± 1 13 ± 1 

FF@SiO2_1A 14 ± 2 100 ± 1 103± 1 29 ± 1 30 ± 1 

FF@SiO2_2A 27 ± 5 64± 2 68 ± 1 44 ± 1 46 ± 2 

FF@SiO2_3A 40 ± 8 47 ± 2 58 ± 1 52 ± 1 63 ± 3 

FF@SiO2_4A 50 ± 7 38 ± 2 57 ± 1 53 ± 1 77 ± 5 

FF@SiO2_5A 66 ± 8 23 ± 3 52 ± 1 58 ± 2 126 ± 18 

FF@SiO2_6A 114 ± 14 15 ± 2 35 ± 1 86 ± 2 192 ± 30 

FF@SiO2_7A 145 ± 10 13 ± 2 33 ± 1 90 ± 2 225 ± 33 

The r1 values obtained at 20 MHz decrease with the increase of the silica shell 

thickness. This decrease is initially quite sharp, from 32.0 s -1 mM -1 for NPs without silica 

coating (d = 10.0 nm) to 11.2 s -1 mM -1 for d = 14 nm, while the r1 values become very 

small (< 2 s -1 mM -1 ) for d > 25 nm (Figure 5.5, inset). At 500 MHz, r1 values are very 

small in all cases, even in the absence of silica shell (Figure 5.5). 

For superparamagnetic NPs, the relaxivities ri (i = 1, 2) are dominated by the 

outer-sphere relaxation mechanism, which is due to the effect of local magnetic field 

73 74 

gradients generated by the NPs on the water protons diffusing near their surface. 

Taking into account the effect of water diffusion through the non-fluctuating magnetic 

field (B0) inhomogeneities created by the time-averaged value of the magnetic moment 

() of the NPs aligned onto B0, and the effect of the fluctuation of the magnetic 

moment itself (Δμz), a theoretical model was developed, where the r1 and r2 relaxivities 

contain terms proportional to 2 , which define the Curie relaxation 75 and dominate at 

high field, and fluctuating terms proportional to Δμz 2 (Néel relaxation) that dominate at 

low field. 74,76 This model accounts quite well for the magnetic field dependence of r1 for 

ultra small particles of iron oxide (USPIO) (diameters of 10-40 nm) at high fields (B0 > 

0.02 T, corresponding to ~ 0.8 MHz Larmor frequency), where Curie relaxation



dominates, but does not account for the small r1 dispersion observed at low field (below 

1 MHz), which depends on the crystal anisotropy energy. 77 Above 1 MHz, r1 depends 

on the translational diffusion correlation time τD and decreases with increase of the 

proton Larmor frequency I, with an inflection point defined by the condition I.τD ~ 1. τD 

= rp 2 /D, where D is the relative diffusion coefficient of the paramagnetic center and the 

water molecule and rp is the radius of the particle, which determines their distance of 

closest approach. 

The decrease of the r1 values at 20 MHz with the increase of the silica shell 

thickness reflects the decrease of the outer-sphere contribution of the core to r1 due to 

the increase of the distance of closest approach of the diffusing bulk water molecules 

to the superparamagnetic core of the particle. This induces an increase of the 

translational diffusion correlation time, D. At least a large part of the silica layer is 

expected to be impermeable to water. The relative diffusion coefficient D is expected to 

be nearly constant for all NPs. Being the sum of the diffusion constants of water (DH2O) 

and of the NP (DNP), it is dominated by DH2O due to the large size of the NPs and the 

slow diffusion of water in the putative silica surface layer. The very small r1 values 

obtained at 500 MHz result from the expected field dependence of outer-sphere 

relaxation. 

The effective transverse relaxation rates (R2*) for the aqueous suspensions of the 

γ-Fe2O3@SiO2 NPs were obtained from the spectral line widths of their proton water 

resonance. Values of R2p* (the paramagnetic contribution to R2*) were calculated by 

subtraction of the diamagnetic contribution of aqueous suspensions of diamagnetic iron 

oxide-free silica NPs from each paramagnetic contribution, using the spectral line 

widths for the various samples. Finally, the corresponding relaxivities, r2*, were 

obtained (see Table 5.2). The line broadening effects reflect the dephasing of the water 

proton magnetic moments diffusing past the magnetic field gradients in the vicinity of 

the small superparamagnetic NPs, causing their T2-shortening. 

The transverse relaxation times are characterized by the correlation times D, 

() -1 , and CP. The uncoated particles have a radius of 5 nm, from which it can be 

calculated that for these particles D is 10 -8 s. From simulations reported by Gillis et 

al., 78 the transverse relaxivity may be predicted by the outer sphere theory (eq. 5.3), 

187



where Δω is the difference in the Larmor frequency at the particle surface and the 

infinity and v is the volume fraction of the particles. 

r r v 

2 

* 4 

2 2 

9 

(5.3) 

D 

Upon coating, both D and () -1 will decrease and we assume that the outer 

sphere regime remains valid. 

The r2 values were measured as a function of the time interval between two 

consecutive 180 o pulses (CP) in a CPMG pulse sequence, for aqueous suspensions of 

the various NPs of increasing diameter. Figure 5.6 shows that the transverse 

relaxivities of these NPs are virtually independent of CP for all silica shell sizes. This 

observation is not surprising, since the D values of the systems measured are all much 

smaller that the applied CP values and, consequently the refocusing pulses are fully 

effective. 

r 2 (s -1 mM -1 ) 

120 

100 

80 

60 

40 

20 

0 

0 1 2 3 4 5 

cp (ms) 

Figure 5.6. Dependence of r2 water proton relaxivities (500 MHz, 25 ºC) of 

aqueous suspensions of the γ-Fe2O3@SiO2 NPs (series A) on cp as a 

function of their diameter, as a result of increased silica layer thickness. 

188 

1A 

2A 

3A 

4A 

5A 

6A 

7A


a) 


Figure 5.5 and Table 5.3 show that the r2 relaxivity (measured at CP = 1.6 ms) 

sharply decreases when the thickness of the coating of the NPs increases. As 

discussed above for r1 effects, this results from the decrease of the outer-sphere 

contribution of the core to r2 due to the increase of the distance of closest approach of 

the diffusing bulk water molecules to the superparamagnetic core of the particle. 

Data show that r2 r2* for the smallest particles (γ -Fe2O3 NPs (core), 1A and 2A), 

but r2 < r2 * for particles with thicker coatings. It is possible that for the thicker coatings 

the silica layer is only impermeable to water up to a certain silica shell thickness. The 

diffusion of the water molecules in the permeable silica layer may be relatively slow. If 

in this layer the diffusion is so slow that the condition D >> () -1 holds, the diffusion 

correlation time is not effective when refocusing pulses are applied and, consequently, 

the phase incoherence of the water protons is fully refocused in that part of the system, 

resulting in zero contribution to r2. As far as r2 and r2 * are concerned, it will be assumed 

that the particles consist of three spheres 79 with radii rc, ri, and rdiff (Figure 5.7). 

rc 

ri 

rdiff 

Figure 5.7. a) Schematic representation of a γ-Fe2O3@SiO2 NP. Here, rc is 

the radius of the core and ri and rdiff are the radii of imaginary spheres, as 

defined in the text. b) Variation of the silica permeability to water molecules 

with the shell thickness. 

r (nm) 

140 

120 

100 

80 

60 

40 

20 

b) 

0 

0 20 40 60 60 80 80 

189 

ri 

rdiff 

exp. radius 

experimental diameter (nm)



Here, rc is the radius of the core (5 nm), ri is the radius of the sphere around the 

core, that seems to be impermeable to water, and rdiff is the radius of a sphere, in which 

any water molecule that is inside diffuses very slowly and does not contribute to r2. 

Water molecules outside the latter sphere are assumed to contribute fully to r2, 

whereas all water (including that inside the latter sphere) contributes to r2 * . 

Taking into account the distance dependence of , v and D, the following 

80 81 

scaling may be applied: 

 

rc 

 

c 

ri 

 

 

 

 

3 

i 

(5.4) 

3 

r i v 

 

c 

rc 

 

v i 

(5.5) 

2 

r i 

 

 

D i Dc 

 

(5.6) 

rc 

 

r 

* 

r 

Combination of eqs (5.4-5.6) gives: 

r c 

 

 

 

i 

* 

2, 

c 

r 

2 (5.7) 

Similarly, it can be derived that: 

 

 

r 

r r2, 

c r 

c 

diff 

 

 

 

 

2 (5.8) 

190



Using the two latter equations and the experimental values of r2 and r2 * , the 

values of ri and rdiff were calculated for the various samples (see Table 5.2 and Figure 

5.7b). These calculated rdiff values are in relatively fair agreement with the particle 

diameters obtained from the TEM measurements. The results also suggest that the 

water impermeable part of the silica coating tends to a maximum value of 40 nm, while 

the water permeable part increases with the coating thickness. 

The core-shell NPs (series B) were studied by relaxometry and magnetometry. 

As expected for particles with the same magnetic core, the size and the saturation 

magnetization obtained by magnetometry remain almost constant (Table 5.4 and 

Figure 5.8). 

Table 5.4. Magnetization values of the NPs (series B) obtained by 

magnetometry. 

Sample r (nm) a Msat (Am 2 /kg) 

FF 5.5 62.1 

FF@SiO2_1B 5.7 55.7 

FF@SiO2_2B 5.5 58.4 

FF@SiO2_3B 5.6 57.4 

FF@SiO2_4B 5.6 56.2 

FF@SiO2_5B 5.6 55.4 

FF@SiO2_6B 5.6 55.9 

FF@SiO2_7B 5.6 57.8 

191



Magnetization (Am 2 kg -1 ferrite) 

60 

40 

20 

0 

-20 

-40 

-60 

-1000 -800 -600 -400 -200 0 200 400 600 800 1000 

Field (mT) 

Figure 5.8. Magnetometry curves of the NPs:FF (magenta, ); 

FF@SiO2_1B; (olive, ); FF@SiO2_2B (wine, ); FF@SiO2_3B (blue, ); 

FF@SiO2_4B (red, ); FF@SiO2_5B (pink, ); FF@SiO2_6B (cyan,) and 

FF@SiO2_7B (green, ). 

In order to investigate the influence of the shell thickness of the silica coating on 

the MRI contrast agent (CA) efficiency of the γ-Fe2O3 NPs, the ri (i = 1,2) relaxivities 

(defined as enhancement of Ri = 1/Ti, i = 1,2, the relaxation rates per mM concentration 

of CA) of the different core-shell NPs were measured at two resonance frequencies (20 

and 60 MHz) and 37 °C (Table 5.5). Their relaxometric behavior shows a decrease of 

the relaxation rates on increasing the size of the NPs, in agreement with preliminary 

data. 82 

The 1 H NMRD profiles recorded at 37 °C confirm the decrease of r1 when the 

coating is thicker (Figure 5.9). All NMRD profiles (Figure 5.8) have been fitted using the 

phenomenological model developed by Roch et al. 77 

192



Table 5.5: Relaxivity values of the NPs aqueous solutions at 20 and 60 

MHz (37 °C) 

Sample 

FF 

FF@SiO2_1B 

FF@SiO2_2B 

FF@SiO2_3B 

FF@SiO2_4B 

FF@SiO2_5B 

FF@SiO2_6B 

FF@SiO2_7B 

r1 (20 MHz, s -1 

mM -1 ) 

r2 (20 MHz, s -1 

mM -1 ) 

193 

r1 (60 MHz, s -1 

mM -1 ) 

35.17 129.25 15.96 130.28 

21.82 114.18 7.66 117.03 

8.14 96.16 2.05 102.08 

2.38 85.89 0.72 92.91 

0.84 50.25 0.40 55.02 

0.55 39.78 0.32 43.63 

0.57 40.57 0.34 43.06 

0.55 39.18 0.34 43.06 

r2 (60 MHz, s -1 

mM -1 ) 

For each sample, the anisotropy energy is high enough to neglect the precession 

at the electronic frequency of the magnetization of the particle. This results in the 

absence of dispersion at low field in the NMRD curve. Accordingly, the fitting equation 

may be reduced to the “high anisotropy approximation”: 

 

 

 

7 

L x 2 2 L x 

F 

2 A 

R1 3 

C 

31 

L x J I , D , N 3 

L x J I 

, D 

 

x 

x 

 

(5.9) 

where R1 is the water proton longitudinal relaxation rate. The constant C is equal 

to (in cgs units) 

32 C 

2 2 p 

C 

 

 

i sp. 

405000 

 

d D 

, where i is the gyromagnetic ratio of 

proton (2.67519. 10 4 rad.G -1 .s -1 ), μsp is the magnetic moment of the NP (e.m.u) s -1 , Cp 

is the number of particles per liter, d is the diameter of the NP (cm), D is the solvent 

media self diffusion coefficient (cm 2 s -1 ) and L(x) is the Langevin function, with



satB0 x , where 0 

k T 

b 

temperature. 

B is the magnetic field, kb the Boltzmann constant and T the 

The J F (I,D,N) function is the Freed expression of the spectral density expressed 

as: 83 

J 

F 

, 

 

 

Re 

 

1 

 

1 

194 

i 

 

 

, D N 

3 2 

1 2 4 

 

i 

D 

D N i 

D 

D N 

i 

D 

D N 

 

where I is the proton Larmor angular frequency (rad.s -1 ), N 

(s -1 ) and 

as: 84 

J 

A 

2 

d 

 

4D 

D is the translational correlation time (s -1 ). 

D 

4 

9 

D 

1 2 

N 

9 

(5.10) 

is the Néel relaxation rate 

The J A (I,D) function is the Ayant expression of the spectral density expressed 

 

2 

1 2 . . 

5 D 

D 

1 

 

 

8 8 

(5.11) 

, D 

3 2 

2 

5 2 

3 

1 2 

. D . D 4. 

D . D . D 

1 

. 

D 

6 

81 

81 

648



Relaxivity (s -1 mM -1 ) 

70 

60 

50 

40 

30 

20 

10 

0 

0.01 0.1 1 10 100 1000 

proton Larmor frequency (MHz) 

Figure 5.9: 1 H NMRD profiles for the NPs aqueous solutions and their 

theoretical fittings. NPs: FF (magenta, ); FF@SiO2_1; (olive, ); 

FF@SiO2_2 (wine, ); FF@SiO2_3 (blue, ); FF@SiO2_4 (red, ); 

FF@SiO2_5 (pink, ); FF@SiO2_6 (cyan,) and FF@SiO2_7 (green, ). 

This equation provides a good fitting of the NMRD curves for NPs covered with 

the thinnest silica layers. However, when the coating becomes very thick, the 

diamagnetic contribution of the silica cannot be neglected in the fitting of the NMRD 

profile. This contribution must be added to that of the magnetic particles alone. As 

shown by Roose et al., 85 the NMRD profile of such large diamagnetic particles is 

characterized by dispersion at very low fields. 

For this reason the central region of the NMRD profiles of the particles covered 

with the thickest layers of silica (see samples FF@SiO2_4B, FF@SiO2_5B, 

FF@SiO2_6B, FF@SiO2_7B) are fitted “by eye” with the superparamagnetic relaxation 

(between 0.5 and 10MHz). A better fit could be obtained after subtraction of the 

diamagnetic contribution of the silica. The values of the parameters obtained by fitting 

the NMRD profiles are presented in Table 5.5 together with those obtained by TEM and 

magnetometry. 

195



A good linear relationship is obtained between the NPs sizes measured by 

electron microscopy and those obtained by fitting the NMRD profiles (Figure 5.9 and 

Table 5.5). Importantly, the sizes obtained by NMRD for particles coated with a thick 

layer of silica are significantly lower than those measured by transmission electron 

microscopy, indicating that a significant part of the silica coating is permeable to water. 

Indeed, the effective distance of closest approach of the water protons to the 

superparamagnetic core, as determined by equation 1, is shorter than the NPs 

thickness as given by TEM, in agreement with what was proposed before for these 

systems. 82 

Interestingly, the magnetization Msat expressed relatively to the global weight of 

the particles, and obtained from the relaxometric data decreases significantly for the 

larger particles (Table 5.5). A correlation can be drawn between the values of Msat 

obtained by both magnetometry (which gives the value of the magnetization of the 

crystal itself) and relaxometry (which gives the value of the magnetization of “the” 

particle, i.e. the crystal plus the coating). Assuming a waterproof coating, such a 

correlation is achieved by multiplying the former one by the factor in eq. 5.12: 

M 

coat 

s 

where 

3 

bp 

3 

bp 

4. 

90 

d 

bp 

M s 

(5.12) 

4. 

90 

d 2. 

2 

( d d ) 

3 

coat 

3 

bp 

coat 

M s is the average saturation magnetization per kilo of coated particle 

material, M is the saturation magnetization of the bare maghemite particle, dbp is the 

bp 

s 

diameter of the bare particle, dcoat is the diameter of the coated particle, and 4.90 and 

2.2 are, respectively, the specific mass of maghemite and silica. The Msat values of 

Table 5.5 show that, by covering the magnetic crystal and “expelling” the water further 

and further away from it creates bigger and bigger particles of the same “magnetic 

content”, so behaving like larger but less magnetized particles. However, the fact that 

the magnetization dilution of the core-shell NPs measured by NMRD is significantly 

smaller than that predicted from eq.5.12 clearly indicates again that the silica shell is 

not waterproof, but is partly water permeable. This observation is of paramount 

importance in the context of the development of contrastophores for molecular imaging. 

196



NMRD diameter 

70 

60 

50 

40 

30 

20 

10 

0 

0 20 40 60 80 100 120 140 160 180 

TEM diameter 

197 

Slope= 0.31 ± 0.04 

R 2 = 0.966 

Figure 5.19: Comparison between the NPs diameter obtained by NMRD 

fitting and TEM. 

Table 5.5: Parameters obtained for the NPs samples by different 

techniques. 

Sample number TEM 

diameter 

(nm) 

Magnetome 

tric 


(nm) 

NMRD 


(nm) 

Msat Crystal 

Magnetometry 

(Am 2 /kg of iron 

oxide) 

Measured Msat 

NMRD 

(Am 2 /kg 

of particle) 

Calculated Msat 

NMRD 

(Am 2 /kg 

of particle) a 

FF 9.6 11.0 13.6 62.1 51.1 51.1 51.1 

FF@SiO2_1B 11.2 11.4 16.2 55.7 47.9 39.0 40.4 

FF@SiO2_2B 24.7 11.0 21.4 58.4 39.0 22.2 6.2 

FF@SiO2_3B 46.4 11.2 27.6 57.4 23.9 11.9 1.00 

FF@SiO2_4B 73.8 11.2 40.0 56.2 12.3 4.3 0.25 

FF@SiO2_5B 94.9 11.2 43.4 55.4 8.2 3.4 0.12 

FF@SiO2_6B 114.2 11.2 43.0 55.9 6.4 3.5 0.07 

FF@SiO2_7B 152.9 11.2 61.6 57.8 5.1 1.2 0.03 

a,b Calculated from equation [5.12] using a NMRD diameters and b TEM diameters. 

Calculated Msat 

TEM 

(Am 2 /kg 

of particle) b



The conditions under which the equations for outer sphere relaxation are valid 

still have to be discussed. This formalism is based on a small perturbation theory. By 

simulation, Gillis and al. 78 have shown that for magnetite particles without diamagnetic 

coating, the results of the outer sphere theory are valid when τD is shorter than 10 -7 s. If 

the measurement is made at 37 °C (D = 3.10 -9 m 2 s -1 ), the diameter should, thus, be 

smaller than 30 nm. In the same work, the authors also showed that the conditions of 

application of the outer sphere theory are fixed by the condition ()r.D



coatings because the dependence on the diameter is cubic for () and of power one 

for D. 

5.3.3. Cytotoxicity 

The cytotoxicity of the γ-Fe2O3@SiO2 NPs (series A) was assessed after 

incubations with microglial cell lines for ¾ h, 24 h, 48 h, 72 h, 96 h, 120 h, and 144 h. 

For each time point, cells were incubated or not with the NPs (0.16 mM). Then, the 

cells were separated in two sets, one used as control and the other one being 

submitted to the MTT assay. This test was performed to characterize the viability of the 

cells and evaluate the residual toxicity after the internalization of the NPs. 86 The cell 

viability tests (Figure 5.20), show that with and without NPs as well as for all NPs sizes 

except for [7A], cells can survive internalization and the cell growth process is 

maintained up to 144 h. Additionally, cells internalized with both γ -Fe2O3 or γ - 

Fe2O3@SiO2 particles exhibit the same lag phase of 48 h as the control ones. 

number number of of cells cells 

1200000 1200000 

1000000 1000000 

800000 800000 

600000 600000 

400000 400000 

200000 200000 

00 

Figure 5.20. Cell viability after exposure to different core-shell NPs sizes 

incubated at different times. 

199 

Control Control cells 

Cells Cells + NPs 

3/4H 3/4H 

48H 48H 

96H 96H 

144H 144H 

3/4H 3/4H 

48H 48H 

96H 96H 

144H 144H 

3/4H 3/4H 

48H 48H 

96H 96H 

144H 144H 

3/4H 3/4H 

48H 48H 

96H 96H 

144H 144H 

3/4H 3/4H 

48H 48H 

96H 96H 

144H 144H 

FF FF FF 1A 1A 1A 2A 2A 2A 5A 5A 5A 7A 7A 7A 

[FF] [FF] [0A] [0A] [1A] [1A] [4A] [4A] [6A] [6A]



The MTT test characteristic of the mitochondrial dehydrogenase activity was 

performed after NPs cell internalization. This metabolic test is illustrated in Figure 5.21 

by the measurement of the 570 nm absorbance of incubated cells at different time 

course with different sizes of core-shell NPs. 

optical density (a.u.) 

DO / 10 6 cells 

1.2 

1 

0.8 

0.6 

0.4 

0.2 

0 

2 

1.8 

1.6 

1.4 

1.2 

1 

0.8 

0.6 

0.4 

0.2 

0 

a) 

3/4H 

48H 

3/4H 

48H 

b) 

96H 

144H 

200 

Control cells 

Cells + NPs 

3/4H 

48H 

96H 

144H 

3/4H 

48H 

96H 

144H 

3/4H 

48H 

96H 

144H 

3/4H 

48H 

96H 

[FF] [0A] [1A] [4A] [6A] 

96H 

144H 

144H 

Control cells 

Cells + NPs 

3/4H 

48H 

96H 

144H 

3/4H 

48H 

96H 

144H 

3/4H 

48H 

96H 

144H 

3/4H 

48H 

96H 

144H 

[FF] [0A] [1A] [4A] [6A] 

Figure 5.21. Variation of total activity, and b) activity per million of cells of 

mitochondrial dehydrogenase of microglial cells as given by the variation of 

the solubilized formazan optical density of the medium with incubation times 

for different sizes of NPs after incubation without (control cells) and with 

both coated and uncoated particles of different sizes (cells + NPs).



Both coated and uncoated particles induce an optical density of the cells, which 

varies with the incubation time with a maximum value for 120 h, following the same 

behaviour as that of the control cells. Like the cell growth in the preceding experiment, 

the dehydrogenase activity is not affected by NPs internalization in the cells from FF to 

5A. In these experiments, the crude analysis of the dehydrogenase activity is relevant 

for the cell viability, but is not enough to give information about modifications of the cell 

phenotype. However, the dehydrogenase activity is drastically modified for the larger 

particles (e.g. 7A). Such NPs are known to be internalized and stored in lysosome-like 

vesicles. In case of NPs 7A, the consequences of their accumulation inside such cells 

able to phagocytise particles larger than 100 nm are still unknown. Their impact on the 

local changes in the overall redox potential due to the high load of iron in the vesicles is 

still an open question. We also normalized the activity per cell as a variation of the 

optical density per million cells as shown in Figure 5.21b. This allows all along the time 

of culture, the characterization of the cell growth on the dehydrogenase expression and 

the possible contribution of the NPs uptake during the growth time. The growth of 

control cells shows a basal level decreasing until 24 h that correlates with the lag 

phase of growth. Then the expression of the dehydrogenase per cell increases in two 

major steps, the first one between 48 h and 72 h, and the second one between 96 h 

and 144 h. After this period of time, the cells are nearly confluent. When the cells are 

incubated with NPs, the activity per cell is not significantly affected except for sample 

7A after 144 h of exposure. Therefore, one can safely assume that in these conditions, 

the dehydrogenase expression does not appear to be sensitive to the vesicular load 

with these core-shell NPs. 

5.4. CONCLUSIONS 

The understanding of the relationship between the coating properties and the 

changes in relaxivity is vital for designing magnetic NP probes for MRI. This is 

important for medical applications, as a higher contrast typically leads to a higher 

sensitivity and reduces the amount of contrast agent required for imaging. Our choice 

of a silica coating was motivated by the increased stability of the resulting NP 

suspensions and the ensuing ease of conjugation of targeting molecules to the surface 

of the contrast agents for sensing and imaging. We have shown that in γ -Fe2O3@SiO2 

core-shell NPs, the coating thickness has a significant impact on their r2 and r2* 

201



relaxivities at medium and high fields and on r1 relaxivities at medium fields, as a result 

of decreased outer-sphere relaxation effects. Comparing the r2 and r2* values for the 

different sizes of particles we were able to divide the silica coating in two regions, one 

impermeable close to the γ -Fe2O3 core and one permeable to water and at the 

interface with the bulk water. We have shown that by controlling this coating we are 

able to tune the size of these two regions. The impermeable one seems to increase up 

to a maximum value of 40 nm, while the permeable region goes on increasing with the 

coating thickness. The diffusion of the water molecules in the permeable silica region is 

relatively slow resulting in zero contribution to r2. The effect of silica coating of 

increasing thickness on the r2/r1 ratio is different from that reported for nanocrystalline 

superparamagnetic iron oxide NPs (MIONs) coated with a polyethylene glycol (PEG)- 

modified, phospholipid micelle coating with increasing molecular weights which 

increase the particle diameter, where this increase causes a r2 decrease and a r1 

increase .79 

The magnetometric curves γ -Fe2O3@SiO2 have been fitted with a Langevin 

function and the NMRD profiles with the model developed by Roch et al. The latter 

provides a good fitting of the NMRD curves for the particles covered with the thinnest 

silica layers. However, with increasing coating thickness, the silica diamagnetic 

contribution cannot be neglected in the fitting of the NMRD profile and must be added 

to the contribution of the magnetic particles. A linear relationship between the NP sizes 

measured by TEM and by fitting the NMRD profiles is obtained, but the sizes obtained 

by NMRD are significantly lower than those measured by TEM. A correlation also 

exists between the values of Msat obtained by magnetometry and relaxometry, whereby 

the Msat value of the superparamagnetic core strongly decreases with the increasing 

thickness of the diamagnetic silica shell. However such a Msat dilution as measured by 

NMRD is lower than expected for water-impermeable silica layers. Both results reflect 

the fact that a significant part of the silica coating is permeable to water, in agreement 

with what was proposed before for these systems. 40 The depth of this water-permeable 

layer is uncertain, as the exchange of the more deeply penetrating water molecules 

could become too slow to influence the measured relaxivity. This observation is of 

paramount importance in the context of the development of contrastophores for 

molecular imaging. The adequate silica shell thickness may, thus, be tuned to allow for 

both, a sufficiently high response as contrast agent, and an adequate grafting of 

targeted biomolecules. 

202



Additionally, preliminary cytotoxicity studies confirmed that these contrast agents 

do not appear detrimental to microglial cells. However, as the naked NPs have the 

highest relaxivities, and the coating thickness does not play a role in their cytotoxicity, a 

preliminary conclusion is that overall optimal particles should have a minimal coating 

thickness to provide solution stability and a basis for surface conjugation without 

compromising their relaxivities. 

Therefore, our results provide clues for the design of magnetic NP based contrast 

agents and their optimization for specific applications in medical diagnosis. This is up to 

now the only technique to provide clear evidence that a silica layer used as a coating in 

a core-shell system, exhibits regions that are porous to water and regions that are not. 

The knowledge of theses systems may be extended to other systems and applications. 

203




1 LaConte, L.; Nitin, N.; Bao, G. Magnetic Nanoparticle Probes Mater. Today 2005, 8, 32- 

38. 

2 Prigogine, I.; Stuart, A. Advance in Chemical Physics; Darmann, J.L.; Fiorani, D.; Ronc 

E. Eds., , Wiley, New York, 1977. 

3 Stevens, P. D.; Fan, J.; Gardimalla, H. M. R.; Yen, M.; Gao, Y. Superparamagnetic 

Nanoparticle-Supported Catalysis of Suzuki Cross-Coupling Reactions Org. Lett. 2005, 

7, 2085-2088. 

4 Mornet, S.; Vasseur, S.; Grasset, F.; Veverka, P.; Goglio, G.; Demourgues, A.; Portier, 

J.; Pollert E.; Duguet, E. Magnetic Nanoparticle Design for Medical Applications Prog. 

Solid State Chem. 2006, 34, 237-247. 

5 Sieben, S.; Bergemann, C.; Lübe, A.; Brockmann, B.; Rescheleit, D. Comparison of 

Different Particles and Methods for Magnetic Isolation of Circulating Tumor Cells J. 

Magn. Magn. Mater. 2001, 225, 175-179. 

6 Paul, A. L.; Chandra, G. R.; Terstappen, L. W. M. M. Optimization of Ferrofluids and 

Protocols for the Enrichment of Breast Tumor Cells in Blood J. Magn. Magn. Mater. 

2001, 225, 301-307. 

7 Miller, M. M.; Prinz, G. A.; Cheng, S. F.; Bounnak, S. Detection of a Micron-Sized 

Magnetic Sphere Using a Ring-Shaped Anisotropic Magnetoresistance-Based Sensor: A 

Model for a Magnetoresistance-Based Biosensor Appl. Phys. Lett. 2002, 81, 2211-2213. 

8 Zhao, M.; Josephson, L.; Tang, Y.; Weissleder, R. Magnetic Sensors for Protease 

Assays Angew. Chem. Int. Ed. 2003, 42, 1375-1378. 

9 Okuhata, Y. Delivery of Diagnostic Agents for Magnetic Resonance Imaging Adv. Drug 

Deliv. Rev. 1999, 37, 121-137. 

10 Wunderbaldinger, P.; Josephson, L.; Weissleder R. Tat Peptide Directs Enhanced 

Clearance and Hepatic Permeability of Magnetic Nanoparticles Bioconj. Chem. 2002, 

13, 264-268. 

11 Halavaara, J.; Tervahartiala, P.; Isoniemi, H.; Höckerstedt, K. Efficacy of Sequential Use 

of Superparamagnetic Iron Oxide and Gadolinium in Liver MR Imaging Acta Radiol. 

2002, 43, 180-185. 

12 Bulte, J. W. Intracellular Endosomal Magnetic Labeling of Cells Methods Mol. Med. 

2006, 124, 419-439. 

13 Modo, M.; Bulte, J. W. Cellular MR Imaging Mol. Imaging 2005, 4, 143-164. 

204



14 Burtea, C.; Laurent, S.; Roch, A.; Vander Elst, L.; Muller R. N., C-MALISA (Cellular 

Magnetic-Linked Immunosorbent Assay), a New Application of Cellular ELISA for MRI J. 

Inorg. Biochem., 2005, 99, 1135-1144. 

15 Boutry, S.; Laurent, S.; Vander Elst, L.; Muller, R. N. Specific E-Selectin Targeting with a 

Superparamagnetic MRI Contrast Agent Contrast Med. Mol. Imaging, 2006, 1, 15-22. 

16 Babes, L., Denizot, B., Tanguy, G., Le Jeune, J. J., Jallet, P. J. Synthesis of Iron Oxide 

Nanoparticles Used as MRI Contrast Agents: A Parametric Study Colloid Interface Sci. 

1999, 212, 474-482. 

17 Sonvico, F.; Dubernet, C.; Colombo, P.; Couvreur, P. Metallic Based Nanotechnology, 

Applications in Diagnosis and Therapeutics Curr. Pharm. Des. 2005, 11, 2091-2105. 

18 Corot, C.; Robert, P.; Idee, J. M.; Port, M. Recent Advances in Iron Oxide Nanocrystal 

Technology for Medical Imaging Adv. Drug Delivery Rev. 2006, 58, 1471-1504. 

19 Jordan, A.; Scholz, R.; Wust, P.; Schirra, H.; Schiestel, T.; Schmidt H., Felix R. 

Endocytosis of Dextran and Silane-Coated Magnetite Nanoparticles and the Effect of 

Intracellular Hyperthermia on Human Mammary Carcinoma Cells in Vitro J. Magn. Magn. 

Mater. 1999, 194, 185-196. 

20 Roullin, V. G.; Deverre, J.R.; Lemaire, L.; Hindré, F.; Julienne, M. C. V.; Vienet, R.; 

Benoit, J. P. Anti-Cancer Drug Diffusion Within Living Rat Brain Tissue: An Experimental 

Study Using [3H](6)-5-Fluorouracil-Loaded PLGA Microspheres Eur. J. Pharm. 

Biopharm. 2002, 53, 293-299. 

21 Lübbe, A. S.; Alexiou, C. C.; Bergemann, C. J. Surg. Res. 2001, 95, 200-206. 

22 Jain, T. K.; Morales, M. A.; Sahoo, S. K.; Leslie-Pelecky, D. L.; Labhasetwar, V. Iron- 

Oxide Nanoparticles for Sustained Delivery of Anticancer Agents Mol. Pharm. 2005, 2, 

194-205. 

23 Chourpa, L.; Douziech-Eyrolles, L.; Ngaboni-Okassa, L.; Fouquenet, J. F.; Cohen- 

Jonathan, S.; Souce, M.; Marchais, H.; Dubois, P. Molecular Composition of Iron Oxide 

Nanoparticles, Precursors for Magnetic Drug Targeting, as Characterized by Confocal 

Raman Microspectroscopy Analyst 2005, 130, 1395-1403. 

24 Mary, M. in: Hafeli, U.; Schutt, W.; Zborowski M. (Eds.), Scientific and Clinical 

Applications of Magnetic Carriers, Plenum Press, New York, 1997. 

25 Elaissari, A.; Rodrigue, M.; Meunier, F.; Herve, C. J. Magn. Magn. Mater. 2001, 225, 

127-133. 

26 Gupta, A. K.; Gupta, M., Synthesis and surface engineering of iron oxide nanoparticles 

for biomedical applications. Biomaterials 2005, 26 (18), 3995-4021. 

27 Cornell, R. M.; Schwertmann, U. The Iron Oxides: Structures, Properties, Reactions, 

Occurences and Uses (Wiley-VCH,Weinheim, 2003). 

28 Josephson, L. Magnetic Nanoparticles for MR Imaging. in BioMEMS and Biomedical 

Nanotechnology Ferrari, M., Ed.; Springer: New York, 2006; Vol. 1, p 227. 

205



29 Sun, Y. K.; Ma, M.; Zhang, Y.; Gu, N. Synthesis and Characterization of Biocompatible 

Fe3O4 Nanoparticles Colloids Surf. 2004, 245, 15– 19. 

30 Berry, C. C.; Curtis, A. S. G. Functionalisation of Magnetic Nanoparticles for Applications 

in Biomedicine J. Phys. D. Appl. Phys. 2003, 36, R198– R206. 

31 Chan, D. C. F.; Kirpotin, D. B.; Bunn, P. A., Jr. Synthesis and Evaluation of Colloidal 

Magnetic Iron Oxides for the Site-Specific Radio Frequency-Induced Hyperthermia of 

Cancer J. Magn. Magn. Mater. 1993, 122, 374 – 378. 

32 Jana, N. R.; Peng, X. Single-Phase and Gram-Scale Routes toward Nearly 

Monodisperse Au and Other Noble Metal Nanocrystals J. Am. Chem. Soc. 2003, 125, 

14280 – 14281. 

33 Sun, S.; Murray, C. B.; Weller, D.; Folks, L.; Moser, A. A. Monodisperse FePt 

Nanoparticles and Ferromagnetic FePt Nanocrystal Superlattices Science 2000, 287, 

1989-1992. 

34 Li, J. J.; Wang, Y. A.; Guo, W.; Keay, J. C.; Mishima, T. D.; Johnson, M. B.; Peng, X. 

Large-Scale Synthesis of Nearly Monodisperse CdSe/CdS Core/Shell Nanocrystals 

Using Air-Stable Reagents via Successive Ion Layer Adsorption and Reaction J. Am. 

Chem. Soc. 2003, 125, 12567– 12575. 

35 Murray, C. B.; Norris, D. J.; Bawendi, M. G. Synthesis and Characterization of Nearly 

Monodisperse CdE (E = S, Se, Te) Semiconductor Nanocrystallites J. Am. Chem. Soc. 

1993, 115, 8706 – 8715. 

36 Sun, S.; Zeng, H. Monodisperse Mfe2O4 (M = Fe, Co, Mn) Nanoparticle J. Am. Chem. 

Soc. 2002, 124, 8204-8205. 

37 Jana, N. R.; Chen, Y.; Peng, X. Size- and Shape-Controlled Magnetic (Cr, Mn, Fe, Co, 

Ni) Oxide Nanocrystals via a Simple and General Approach Chem. Mater. 2004, 16, 

3931 – 3935. 

38 Hyeon, T.; Lee, S. S.; Park, J.; Chung, Y.; Na, H. B. Synthesis of Highly Crystalline and 

Monodisperse Maghemite Nanocrystallites without a Size-Selection Process J. Am. 

Chem. Soc. 2001, 123, 12798– 12801. 

39 Wu, W.; He, Q.; Jiang, C. Magnetic Iron Oxide Nanoparticles: Synthesis and Surface 

Functionalization Strategies Nanoscale Res. Lett. 2008, 3, 397– 415. 

40 Gao, X.; Cui, Y.; Levenson, R. M.; Chung, L. W. K.; Nie, S. In Vivo Cancer Targeting 

and Imaging with Semiconductor Quantum Dots Nat. Biotechnol. 2004, 22, 969– 976. 

41 Pellegrino, T.; Manna, L.; Kudera, S.; Liedl, T.; Koktysh, D.; Rogach, A. L.; Keller, S.; 

Radler, J.; Natile, G.; Parak, W. Hydrophobic Nanocrystals Coated with an Amphiphilic 

Polymer Shell: A General Route to Water Soluble Nanocrystals Nano Lett. 2004, 4, 703 

- 707. 

206



42 Guo, W.; Li, J. J.; Wang, Y. A.; Peng, X. Luminescent CdSe/CdS Core/Shell 

Nanocrystals in Dendron Boxes: Superior Chemical, Photochemical, and Thermal 

Stability J. Am. Chem. Soc. 2003, 125, 3901-3909. 

43 Kim, M.; Chen, Y.; Liu, Y.; Peng, X. Super-stable, High-quality Fe3O4 Dendron- 

Nanocrystals Dispersible in Both Organic and Aqueous Solutions Adv. Mater. 2005, 17, 

1429 – 1436. 

44 Kim, S. W.; Kim, S.; Tracy, J. B.; Jasanoff, A.; Bawendi, M. G. Phosphine Oxide Polymer 

for Water-Soluble Nanoparticles J. Am. Chem. Soc. 2005, 127, 4556-4557. 

45 Schroedter, A.; Weller, W. Ligand Design and Bioconjugation of Colloidal Gold 

Nanoparticles Angew. Chem., Int. Ed. 2002, 41, 3218-3221. 

46 Lee, D. C.; Mikulec, F. V.; Pelaez, J. M.; Koo, B.; Korgel, B. A. Synthesis and Magnetic 

Properties of Colloidal MnPt Nanocrystals J. Phys. Chem. B 2006, 110, 11160- 11166. 

47 Veiseh, O.; Sun, C.; Gunn, J.; Kohler, N.; Gabikian, P.; Lee, D.; Bhattarai, N.; 

Ellenbogen, R.; Sze, R.; Hallahan, A.; Olson, J. ; et al. An Optical and MRI 

Multifunctional Nanoprobe for Targeting Gliomas Nano Lett. 2005, 5, 1003 – 1008. 

48 Liu, X.; Xing, J.; Guan, Y.; Shan, G.; Liu, H. Synthesis of Amino-Silane Modified 

Superparamagnetic Silica Supports and Their Use for Protein Immobilization Colloids 

Surf. A 2004, 238, 127– 131. 

49 Yi, D. K.; Selvan, S. T.; Lee, S. S.; Papaefthymiou, G. C.; Kundaliya, D.; Ying, J. Y. 

Silica-Coated Nanocomposites of Magnetic Nanoparticles and Quantum Dots J. Am. 

Chem. Soc. 2005, 127, 4990-4991. 

50 Schroedter, A.; Weller, W.; Eritja, R.; Ford, W. E.; Wessels, J. M. Biofunctionalization of 

Silica-Coated CdTe and Gold Nanocrystals Nano Lett. 2002, 2, 1363– 1367. 

51 Bruchez, M.; Moronne, M.; Gin, P.; Weiss, S.; Alivisatos, A. P. Semiconductor 

Nanocrystals as Fluorescent Biological Labels Science 1998, 281, 2013– 2016. 

52 Gerion, D.; Pinaud, F.; Williams, S. C.; Parak, W. J.; Zanchet, D.; Weiss, S.; Alivisatos, 

A. P. Synthesis and Properties of Biocompatible Water-Soluble Silica-Coated CdSe/ZnS 

Semiconductor Quantum Dots J. Phys. Chem. B 2001, 105, 8861-8871 

53 Darbandi, M.; Thomann, R.; Nann, T. Single Quantum Dots in Silica Spheres by 

Microemulsion Synthesis Chem. Mater. 2005, 17, 5720-5725. 

54 Sousa, M. H.; Rubim, J. C.; Sobrinho, P. G.; Tourinho, F. A. Biocompatible Magnetic 

Fluid Precursors Based on Aspartic and Glutamic Acid Modified Maghemite 

Nanostructures J. Magn. Magn. Mater. 2001, 225, 67– 72. 

55 Mornet, S.; Vasseur, S.; Grasset, F.; Duguet, E. Magnetic Nanoparticles Design for 

Medical Diagnosis and Therapy J. Mater. Chem. 2004, 14, 2161– 2175. 

56 Xie, J.; Wang, C. H. Self-Assembled Biodegradable Nanoparticles Developed by Direct 

Dialysis for the Delivery of Paclitaxel Pharm. Res. 2005, 22, 2079– 2090. 

207



57 He, H.; Liu, H.; Zhou, K.; Wang, W.; Rong, P. Characteristics of Magnetic Fe3O4 

Nanoparticles Encapsulated with Human Serum Albumin J. Cent. South Univ. Technol. 

2006, 13, 6– 11. 

58 Tiefenauer, L. X.; Kuhne, G.; Andres, R. Y. Antibody-Magnetite Nanoparticles: In Vitro 

Characterization of a Potential Tumor-Specific Contrast Agent for Magnetic Resonance 

Imaging Bioconjugate Chem. 1993, 4, 347– 352. 

59 Nam, J. M.; Stoeva, S. I.; Mirkin, C. A. Bio-Bar-Code-Based DNA Detection with PCR- 

like Sensitivity J. Am. Chem. Soc. 2004, 126, 5932-5933. 

60 Lewin, M.; Carlesso, N.; Tung, C.-H.; Tang, X.-W.; Cory, D.; Scadden, D. T.; Weissleder, 

R. Tat Peptide-Derivatized Magnetic Nanoparticles Allow In Vivo Tracking and Recovery 

of Progenitor Cells Nat. Biotechnol. 2000, 18, 410 – 414. 

61 Weizmann, Y.; Patolsky, F.; Lioubashevski, O.; Willner, I. Magneto-mechanical 

Detection of Nucleic Acids and Telomerase Activity in Cancer Cells J. Am. Chem. Soc. 

2004, 126, 1073– 1080. 

62 Lartligue, L.; Oumzil, K.; Guari, Y.; Larionova, J.; Guérin, C.; Montero, J.-L.; Barragan- 

Montero, V.; Sangregorio, C.; Caneschi, A.; Innocenti, C.; Kalaivani, T.; Arosio, P.; 

Lascialfari, A. Water-Soluble Rhamnose-Coated Fe3O4 Nanoparticles Org. Lett. 2009, 

11, 2992-2995. 

63 Ashtari, P.; He, X. X.; Wang, K. M.; Gong, P., An efficient method for recovery of target 

ssDNA based on amino-modified silica-coated magnetic nanoparticles. Talanta 2005, 67 

(3), 548-554. 

64 Allemann, E.; Gurny, R.; Doelker, E., Drug-loaded Nanoparticles - Preparation Methods 

and Drug Targeting Issues. European Journal of Pharmaceutics and Biopharmaceutics 

1993, 39 (5), 173-191. 

65 Massart, R., Preparation of Aqueous Ferrofluids without using Surfactant - Behavior as a 

Function of the pH and the Counterions. Comptes Rendus Hebdomadaires Des 

Seances De L Academie Des Sciences Serie C 1980, 291 (1), 1-3. 

66 Mornet, S.; Grasset, F.; Portier, J.; Duguet, E. Maghemite@Silica Nanoparticles for 

Biological Applications European Cells and Materials, 2002, Vol. 3. Suppl. 2, 110-113. 

67 Stöber, W.; Fink, A.; Bohn, E., Controlled Growth of Monodisperse Silica Spheres in 

Micron Size Range. Journal of Colloid and Interface Science 1968, 26 (1), 62-69. 

68 Nozawa, K.; Gailhanou, H.; Raison, L.; Panizza, P.; Ushiki, H.; Sellier, E.; Delville, J. P.; 

Delville, M. H., Smart control of monodisperse Sotber silica particles: Effect of reactant 

addition rate on growth process. Langmuir 2005, 21 (4), 1516-1523. 

69 Chen, S.-L.; Dong, P.; Yang, G.-H.; Yang, J.-J. Characteristic aspects of formation of 

new particles during the growth of monosize silica seeds J. Colloid Interface Sci., 1996, 

180, 237-241. 

208



70 Lu, Y.; Yin, Y. D.; Mayers, B. T.; Xia, Y. N., Modifying the surface properties of 

superparamagnetic iron oxide nanoparticles through a sol-gel approach. Nano Letters 

2002, 2 (3), 183-186. 

71 Mornet, S.; Elissalde, C.; Hornebecq, V.; Bidault, O.; Duguet, E.; Brisson, A.; Maglione, 

M., Controlled growth of silica shell on Ba0.6Sr0.4TiO3 nanoparticles used as 

precursors of ferroelectric composites. Chemistry of Materials 2005, 17 (17), 4530-4536. 

72 Gaudon, M.; Basly, B.; Fauque, Y.; Majimel, J.; Delville, M. H., Thermochromic Phase 

Transition on CuMo(0.9)W(0.1)O(4)@SiO(2) Core-Shell Particles. Inorganic Chemistry 

2009, 48 (5), 2136-2139. 

73 Muller, R. N.; Roch, A.; Colet, J.-M.; Ouakssim, A.; Gillis, P. in: A. E. Merbach, É. Tóth 

(Eds.), The Chemistry of Contrast Agents in Medical Magnetic Resonance Imaging, 

Wiley, New York, Chapter 10, pp. 417-435 (2001). 

74 Laurent, S.; Forge, D.; Port, M.; Roch, A.; Robic, C.; Elst, L. V.; Muller, R. N., Magnetic 

iron oxide nanoparticles: Synthesis, stabilization, vectorization, physicochemical 

characterizations, and biological applications. Chemical Reviews 2008, 108 (6), 2064- 

2110. 

75 Gueron, M., Nuclear relaxation in macromolecules by paramagnetic ions: a novel 

mechanism. Journal of Magnetic Resonance 1975, 19 (1), 58-66. 

76 Roch, A.; Muller, R.N. Longitudinal relaxation of water protons in colloidal suspensions 

of superparamagnetic crystalsProc. 11th Annu. Meet. Soc. Magn. Reson. Med. 1992, 

11, 1447. 

77 Roch, A.; Muller, R. N.; Gillis, P., Theory of proton relaxation induced by 

superparamagnetic particles. Journal of Chemical Physics 1999, 110 (11), 5403-5411. 

78 Gillis, P.; Moiny, F.; Brooks, R. A., On T-2-shortening by strongly magnetized spheres: A 

partial refocusing model. Magnetic Resonance in Medicine 2002, 47 (2), 257-263. 

79 LaConte, L.E.W.; Nitin, N.; Kurkiya, O.; Caruntu, D.; O’Connor, C. J.; Hu, X.; Bao, G. 

Coating thickness of magnetic iron oxide nanoparticles affects R2 relaxivity J. Magn. 

Res. Imaging 2007, 26, 1634-1641. 

80 Norek, M.; Pereira, G. A.; Geraldes, C. F. G. C.; Denkova, A.; Zhou, W.; Peters, J. A. 

NMR Transversal relaxivity of suspensions of lanthanide oxide nanoparticles J. Phys. 

Chem. C 2007, 111, 10240-10246. 

81 Pereira, G. A.; Norek, M., Peters, J. A.; Ananias, D.; Rocha, J.; Geraldes, C. F. G. C. 

NMR transversal relaxivity of aqueous suspensions of particles of Ln(3+)-based zeolite 

type materials Dalton Trans, 2008, 2241-2247. 

82 Pinho, S. L. C.; Pereira, G. A.; Voisin, P.; Kassem, J.; Bouchaud, V.; Etienne, L.; Peters, 

J. A.; Carlos, L. D.; Mornet, S.; Geraldes, C. F. G. C.; Rocha, J.; Delville, M.-H., Fine 

Tuning of the Relaxometry of gamma-Fe(2)O(3)@SiO(2) Nanoparticles by Tweaking the 

Silica Coating Thickness. Acs Nano 2010, 4 (9), 5339-5349. 

209



83 Freed, J. H., Dynamic effects of pair correlation functions on spin relaxation by 

translational diffusion in liquids. II. Finite jumps and independent T1 processes. Journal 

of Chemical Physics 1978, 68 (9), 4034-4037. 

84 Ayant, Y.; Belorizky, E.; Alizon, J.; Gallice, J., Calculation of spectral densities for 

relaxation resulting from random translational modulation of magnetic dipolar coupling in 

liquids. Journal De Physique 1975, 36 (10), 991-1004. 

85 Roose, P.; Vancraen, J.; Finsy, R.; Eisendrath, H., Field-Cycling Proton Nuclear 

Magnetic Relaxation-Dispersion Study of Aqueous Colloidal Silica Sols. Journal of 

Magnetic Resonance, Series A 1995, 115 (1), 20-25. 

86 Voisin, P.; Ribot, E. J.; Miraux, S.; Bouzier-Sore, A.-K.; Lahitte, J.-F.; Bouchaud, V.; 

Mornet, S.; Thiaudiere, E.; Franconi, J.-M.; Raison, L.; Labrugere, C.; Delville, M.-H., 

Use of lanthanide-grafted inorganic nanoparticles as effective contrast agents for cellular 

uptake imaging. Bioconjugate Chemistry 2007, 18 (4), 1053-1063. 

210


211 

6. 

Final Conclusions and Future 

Work


Core-Shell Nanoparticles for Magnetic Resonance Imaging 

The objective of this thesis was to design and prepare bio-imaging probes, in 

particular exhibiting MRI and photoluminescence bimodality. Two different types of 

nanoparticles performing as T1 and T2 MRI contrast agents were synthesised and their 

physical and chemical properties characterised, namely texture, structure, 1 H dynamics 

and relaxometry and photoluminescence properties. The combination of the properties 

of trivalent lanthanide complexes and nanoparticles offered an excellent solution for 

bimodal imaging. The following relevant observations and conclusions may prompt 

future studies on related systems. 

T1-type contrast agents 

ͽ Bimodal MRI-photoluminescence probes for bio-imaging consisting on 

SiO2 nanoparticles derivatized with DTPA-Ln 3+ and PMN-Ln 3+ complexes 

(SiO2@APS/DTPA:Gd:Ln; and SiO2@APS/PMN:Gd:Ln; Ln 3+ = Eu 3+ , Tb 3+ ) 

were developed. These systems bear an active magnetic centre (Gd 3+ ) 

and photoluminescent ions (Eu 3+ or Tb 3+ ) on the surface of silica 

nanoparticles. 

ͽ The number of Ln 3+ (Eu 3+ , Tb 3+ and Gd 3+ ) ions on the surface the 

SiO2@APS/DTPA nanoparticles (up to ca. 10 4 ions per nanoparticle) is 

determined by the amount of DTPA grafted. Because the 

SiO2@APS/PMN nanoparticles exhibit a second type of local coordination 

for the Ln 3+ ion it is difficult to quantify the amount of PMN complex 

grafted, although it is possible to estimate the amount of Ln 3+ per 

nanoparticle (up to ca. 10 5 ions per nanoparticle). 

ͽ The tandem use of 13 C and 29 Si solid-state NMR and DRIFT spectroscopy 

was shown to be very powerful approach to study the modification of the 

surface of the nanoparticles. Upon APS derivatization the number of 

surface hydroxyl groups decreases and the 29 Si (Q 2 +Q 3 )/Q 4 population 

ratio is reduced from 0.43 to 0.37. For the SiO2@APS/DTPA system, clear 

evidence of the covalent bonding between APS and DTPA is provided by 

the secondary amide C=O stretch at 1685 cm -1 . Since the number of 

silanol groups increases (the ratio (Q 2 +Q 3 )/Q 4 increases from 0.37 to 0.62) 

reaction of SiO2@APS with DPTA has the side effect of modifying the 

surface of the SiO2 nanoparticles. Although the APS to PMN covalent 

212


bonding could not be ascertained by DRIFT or NMR evidence for it was 

forthcoming from photoluminescence spectroscopy. As observed in the 

DTPA system, the reaction between APS and PMN modified the 

nanoparticles surface, increasing the number of silanol groups, with 

(Q 2 +Q 3 )/Q 4 raising from 0.37 to 0.51. 

ͽ Complexation of DTPA and PMN to Ln 3+ ions emitting in the visible region 

afforded long-life excited states, resistance to photo-bleaching, and sharp 

emission bands. The photoluminescence properties of the systems were 

not changed by the incorporation of Gd 3+ ions. Evidence was obtained for 

the energy transfer from the DTPA/PMN ligand to Ln 3+ . The grafting of 

PMN, a pyridine-based aromatic antenna ligand enhanced this energy 

transfer. While in the SiO2@APS/DTPA:Eu system the Eu 3+ ions reside in 

a single low-symmetry site, with a large local distribution, in the 

SiO2@APS/PMN:Eu system there are two local Eu 3+ environments. We 

conjectured that in one site the Eu 3+ coordinates to the PMN chelate, 

while in a second site it interacts strongly with the NPs surface via the 

silanol groups, which are far more abundant than the amino groups on the 

silica surface. 

ͽ The incorporation of Gd 3+ ions, the MRI probe, in the nanosystems does 

not change the emission properties of the Eu 3+ and Tb 3+ ions. The 

relaxometric features of these nanoparticles are slightly better than the 

properties of the commercially-available [Gd(DTPA)] 2- complex. 

ͽ The bimodal probes are rapidly and efficiently uptaken by RAW 264.7 

cells (mouse macrophage cell line) exhibiting the 1 

213 

T -weighted MRI images 

of cellular pellets increased contrast and potential optical tracking by 

fluorescence. 

ͽ In future studies an alternative and more efficient procedure should be 

devised for grafting the PMN complex, ensuring that a single Eu 3+ local 

environment is present. 

T2-type contrast agents 

ͽ Iron oxide nanoparticles coated with silica layers of different thickness 

were prepared. The stability of the resulting nanoparticles suspensions is


Core-Shell Nanoparticles for Magnetic Resonance Imaging 

increased by this silica shell, which also allows the straightforward grafting 

of targeting molecules. 

ͽ The thickness of this silica shell changed considerably the r2 and r2* 

relaxivities at medium and high B0 fields (0.47 T and 11.7 T), and the r1 

relaxivity at medium B0 fields (0.47 T) as a result of decreased outer- 

sphere relaxation effects. 

ͽ A model was proposed to explain the impact of the silica shell on the r2 

and r2* relaxivities. This model partitions the silica coating in two regions, 

one impermeable to water and close to the γ-Fe2O3 core, the other porous 

to water and at the interface with the bulk water. 

ͽ The magnetometry curves of the γ-Fe2O3@SiO2 nanoparticles were fitted 

with a Langevin function and the NMRD profiles with the model of Roch et 

al. A good fit of the NMRD curves was obtained for particles covered with 

the thinnest silica layers. For particles with relatively thick silica layers it 

was necessary to introduce a contribution from the diamagnetic silica. 

ͽ The cell viability and mitochondrial dehydrogenase expression given by 

the microglial cells were evaluated and confirmed that these contrast 

agents are not detrimental to the microglial cells. 

ͽ The knowledge obtained with these systems may now be extended to 

other systems and applications. The results provide clues for the design of 

contrast agents based on superparamagnetic nanoparticles and their 

optimisation for specific medical diagnosis applications. 

ͽ In future work this system should be extended to bimodal, by grafting a 

photoluminescence probe. 

214

tel-00661206, version 1 - 18 Jan 2012


Résumé : 

Cette thèse décrit une stratégie de synthèse de nouvelles générations des nanoparticules (NPs) pour 

applications biomédicales, visant à une amélioration de leurs performances pour l’imagerie, le 

diagnostic thérapeutique. Ces NPs présentent plusieurs fonctionnalités leur permettant de réaliser 

des tâches multiples. Deux types de sondes bimodales ont été développés et étudiés afin d'évaluer 

leur potentiel comme agents (1) de contraste en IRM et (2) luminescents. Ces objets combinent les 

propriétés des complexes de lanthanide (Ln 3+ ) et celles des NPs de silice ou de type cœur-écorce 

Fe2O3@SiO2 pour une imagerie bimodale. Ces NPs testées sur des cellules vivantes ont permis 

d’illustrer la preuve du concept aussi bien en IRM avec une augmentation d'intensité des images et 

un impact significatif sur les relaxivities r1, r2 et r2* qu’en photoluminescence. L’étude du système 

cœur-écorce a montré que l’influence du contrôle fin de l’écorce autour du noyau d'oxyde de fer a pu 

être modélisée. 

Mots clés : 

Nanoparticules multifonctionnelles, coeur-écorce/couronne, Fe2O3, 

Lanthanides, agents de contraste IRM agents de contraste optiques 

relaxométrie, RMND, photoluminescence 

Title: Multifunctional nanoparticles for MR and fluorescence imaging. 

Abstract: 

This thesis describes a strategy of synthesis of new generations of nanoparticles (NPs) for 

biomedical applications, aiming at an improvement of their performances for the imaging, and the 

therapeutic diagnosis. These NPs present several functionalities enabling them to carry out multiple 

tasks. Two types of bimodal probes were developed and studied so as to evaluate their potential as 

contrast agents (1) in MRI and (2) and luminescence. These objects combine the properties of the 

lanthanide complexes (Ln 3+ ) and those of NPs of silica or core/shell Fe2O3@SiO2 for a bimodal 

imaging. These NPs tested on living cells were able to illustrate the proof of the concept not only in 

MRI with an increase of intensity of the images and a significant impact on the relaxivities r1, r2 and 

r2* but also in photoluminescence. The study of the core/shell system showed that the influence of 

the fine control of the shell around the iron oxide core could be modeled. 

Keywords : 

Multifunctional nanoparticles, core-shell/corona, Fe2O3, 

Lanthanides, MRI contrast agents optical contrast agents 

relaxometry, RMND, Photoluminescence

Nanoparticules multifonctionelles pour la résonance magnétique et l ...

You also want an ePaper? Increase the reach of your titles

Delete template?

Save as template?