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BIOCHIP AND LAB-ON-CHIP FOR 

GENOMIC AND PROTEOMIC 

APPLICATIONS 

Lucia Rotiroti 

Dottorato in Scienze Biotecnologiche – XXI ciclo 

Indirizzo Biotecnologie Industriali 

Università di Napoli Federico II

Dottorato in Scienze Biotecnologiche – XXI ciclo 

Indirizzo Biotecnologie Industriali 

Università di Napoli Federico II 

BIOCHIP AND LAB-ON-CHIP FOR 

GENOMIC AND PROTEOMIC 

APPLICATIONS 

Lucia Rotiroti 

Dottoranda: Rotiroti Lucia 

Relatore: Dr. De Stefano Luca 

Coordinatore: Prof. Giovanni Sannia

The most beautiful thing we can experience is the mysterious. 

It is the source of all true art and science. 

Albert Einstein 

A TE 

…che hai sempre confidato in me e nelle mie capacità… 

…che mi hai dato forza, sempre… 

…che hai sostenuto ogni mia scelta di vita… 

…illuminandone la via con straordinaria discrezione… 

…a te…grazie!

INDICE 

SUMMARY pag. i 

RIASSUNTO 

Chapter 1:Porous Silicon (PSi) 

pag. ii 

Introduction 

pag. 1 

Materials and methods 

pag. 10 

Thue-Morse: aperiodic multilayers made of PSi 

pag. 16 

Optical sensing of chemical compounds using porous silicon devices 

Chapter 2:Chemical and Biological functionalization of PSi surface 

pag. 22 

Introduction: PSi surface chemical functionalization 

pag. 29 

DNA Optical Detection Based on PSi Technology 

pag. 33 

Optical biosensors for the detection of analytes of clinical and industrial interest. pag. 40 

Optical evidence of protein structural changes in PSi devices 

pag. 55 

Oligonucleotides direct synthesis on PSi chip 

pag. 61 

PSi-based optical biosensors and biochips 

Chapter 3:Hybrid systems based on PSi and biocompatible materials 

pag. 63 

Introduction: PSi hybrid systems 

pag. 67 

PSi-polymer composite matrix for an innovative class of optical biosensors pag. 70 

A polymer modified PSi optical device for biochemical sensing 

pag. 73 

Bio/NON Bio Interfaces for a new class of proteins Microarrays 

Chapter 4: …towards the Lab-on-Chip 

pag. 77 

Introduction to Lab-on-chip technologies 

pag. 82 

Microsystems Based on PSi-Glass Anodic Bonding technologies 

pag. 87 

Laser oxidation micropatterning of a PSi based biosensor for multianalytes 

microarrays 

Appendix 

pag. 95

SUMMARY 

There is no other inorganic material that brings together such much intriguing features like 

porous silicon (PSi) for people working in the optical sensing field. Porous silicon was 

unexpectedly discovered in the late fifties when chemists attempted to electropolish silicon 

wafers with an electrolyte containing hydrofluoric acid to get better electrical contacts in 

the first integrated electronic circuit fabrications. The acid dissolution left a sponge-like 

nanocrystalline silicon structure on the wafer. Today, the electrochemical etching is a 

standard method to fabricate nanostructured porous silicon: a proper choice of the applied 

current density, the electrolyte composition, and the silicon doping allow precise control 

over the morphology and, consequently, on the physical and chemical properties of the 

porous silicon structure. Computer controlled electrochemical etching processes are 

exploited for the realization of porous silicon films of controlled thickness and porosity 

(defined as the percentage of void in the silicon volume). Nanoporous, mesoporous and 

macroporous structures can be achieved, with pore size ranging from few nanometers up 

to microns. Moreover, since the etching process is self-stopping, it is possible to fabricate 

with a single run process multilayer stacks made of single layers of different porosity. The 

dielectric properties of each PSi layer, and in particular its refractive index n, can be 

namely modulated between those of crystalline silicon (n = 3.54, porosity = 0) and air (n= 

1, porosity = 100 %); so that alternating high and low porosity layers, lot of photonic 

structures, such as Fabry-Perot interferometers, omni-directional Bragg reflectors, optical 

filters based on microcavities, and even complicated quasi-periodic sequences can be 

simply realized. 

The other key feature for a transducer material is the large area and the chemistry of its 

surface: the porous silicon exhibits a very reactive hydrogenated specific area of the order 

of 200 – 500 m 2 cm -3 , which assures an effective interaction with several adsorbates. The 

porous silicon optical sensing features are based on the changes of its photonic 

properties, such as photoluminescence or reflectance, on exposure to the gaseous or 

liquid substances. Of course, these interactions are not specific. Therefore, the porous 

silicon hydrogenated surface has to be chemically or physically modified in order to 

achieve the sensing selectivity through specific biochemical interactions. The biosensor 

reliability strongly depends on the functionalization process: how simple, homogenous and 

repeatable it can be. The substitution of the superficial Si-H bonds with Si-C ones or Si-O- 

C guarantees a much more stable surface interface from the thermodynamic point of view. 

Several functionalization strategies were investigated. Testing and demonstrating the 

porous silicon capabilities as a useful functional material in the immobilization of biological 

probes and in the optical transduction of biochemical interactions is only the first action in 

the realisation of an optical biochip based on this nanostructured material. In this case, all 

the fabrication processes should be compatible with the utilisation of biological probes and 

the feasibility of such devices must be proven. This means that the standard integrated 

circuit micro-technologies should be modified and adapted to this new field of application. 

The porous silicon low cost technology could thus provide a link between the conventional 

CMOS technology and the photonic devices in the realization of the so-called smart 

sensors and biochips. A very strong interdisciplinary approach is required to match and 

resolve all the technological problems. The purpose of this thesis was to design and 

fabricate different resonant optical structures, integrate in simple lab-on-chip devices 

based on the porous silicon nanotechnology, and it was focused on optical biosensing 

applications of social interest, such as environmental monitoring and biomedical 

diagnostics. 

i

RIASSUNTO 

Una delle più importanti scoperte scientifiche di tutti i tempi è sicuramente la decodifica del 

genoma umano portata a termine nel 2000 dal progetto internazionale Human Genome 

Organization. Lo studio del genoma consiste nel descrivere la struttura, la posizione e la 

funzione dei circa 25.000 geni che caratterizzano la specie umana. A tal fine i genetisti 

hanno dovuto sequenziare i circa due metri di DNA contenuto in ogni cellula umana e 

quindi identificare la sequenza di approssimativamente tre miliardi di coppie di basi 

azotate che ne compongono la molecola. Come in ogni grande impresa scientifica, lo 

sviluppo tecnologico ha fortemente influenzato i tempi di esecuzione e termine di questo 

lavoro: basti pensare che il primo miliardo di basi è stato decodificato in quattro anni, 

mentre il secondo miliardo in solo quattro mesi. Oggigiorno il ritmo di sequenziamento 

delle basi è dell’ordine di diverse decine di migliaia di basi al minuto, a seconda della 

tecnologia utilizzata. Lo studio dei geni permette l’identificazione non solo del loro ruolo 

ma soprattutto dei loro malfunzionamenti che sono alla base delle malattie: le prospettive 

della medicina genomica sono talmente vaste da non esser ancora state tutte catalogate e 

pienamente comprese. Il successo del progetto genoma umano è dovuto principalmente 

ad un fattore chiave: la miniaturizzazione della strumentazione di laboratorio, in particolare 

di quella dedicata alla PCR (Polymerase Chain Reaction) e delle matrici a DNA, i 

cosiddetti “microarray” o “biochip”, che permettono di esaminare il profilo d’espressione di 

un gene o di identificare la presenza di un gene specifico con elevatissima precisione e 

ripetibilità. 

I biochip sono una particolare categoria di dispositivi più complessi comunemente indicati 

come BioMEMS (Sistemi Micro-Elettro-Meccanici Biologici) che, negli ultimi anni, hanno 

trovato un numero sempre crescente di applicazioni biologiche e biomediche. Lo sviluppo 

di questi dispositivi è parallelo a quello dei Lab-on-chip (LOC) che rappresentano 

l’evoluzione tecnologica più spinta del laboratorio analitico tradizionale: i LOC sono dei 

microsistemi complessi che coniugano tutte le funzioni generalmente svolte da 

strumentazione da banco in un unico dispositivo fabbricato su di un supporto monolitico, 

tipicamente di silicio cristallino, ed assemblato con altri materiali compatibili, come il silicio 

poroso ed il vetro o i polimeri, aventi ciascuno una propria specifica funzionalità. Entrambe 

le due categorie di dispositivi, i biochip ed i LOC, beneficano degli sviluppi di quello che è 

sicuramente il campo più avanzato della tecnologia umana e cioè la microelettronica 

integrata: la produzione di computer sempre più potenti è legata alla capacità di fabbricare 

dispositivi su scala micrometrica e nanometrica con una capacità di controllo della materia 

solida prossima al limite atomico. La possibilità di miniaturizzare la componentistica di 

laboratorio (separatori, concentratori, analizzatori, sensori, e così via) fa sì che il campo di 

applicazione dei LOC abbia una evoluzione esponenziale trovando impiego dalla 

diagnostica al monitoraggio ambientale alla rivelazione di agenti pericolosi nella sicurezza 

civile e militare, fino alle applicazioni industriali in chimica sintetica (per esempio in prove 

rapidi in microreattori per la farmaceutica). 

A partire dai biochip a DNA, è stata sviluppata una serie di dispositivi che utilizzano altri 

organismi biologici quali anticorpi, enzimi, proteine, cellule. Data la capacità di queste 

biomolecole di agire come sonde altamente specifiche rispetto ai rispettivi ligandi, la 

principale applicazione dei dispositivi è come biosensori e cioè nel riconoscimento della 

presenza dei ligandi in miscele complesse quale può essere il sangue umano. Un 

biosensore, il cui schema è rappresentato in Figura 1, lavora con un elemento biologico, 

detto biorecettore, fissato su un substrato ove avviene una reazione biologica specifica tra 

il biorecettore e l’elemento ricercato. L’evento di riconoscimento biomolecolare è rilevato 

da un trasduttore che lo converte in un segnale elettrico, ottico o chimico, che può essere 

osservato e misurato. La sensibilità e la selettività dei biosensori è tale da poter applicare 

questi strumenti anche nello studio delle interazioni biomolecolari: un chiaro successo 

ii

commerciale di questi dispositivi è il Biacore ® , dispositivo ottico comunemente utilizzato 

nella caratterizzazione della cinetica molecolare e nel calcolo delle costanti di 

dissociazione e di affinità tra biomolecole. 

Internal 

Membrane 

Electrolyte 

Signal 

Transducer 

Biorecognition 

Element 

Semipermeable 

Membrane 

Figura 1 Schema generale di un biosensore. 

Sulla base del metodo di trasduzione impiegato, i biosensori possono essere raggruppati, 

in quattro gruppi fondamentali: 

1. Elettrochimici; 

2. Termici; 

3. Ottici; 

4. Acustici; 

Le prime due tecniche di trasduzione vengono spesso impiegate nella realizzazione di 

dispositivi il cui biorecettore è un enzima; mentre i trasduttori più comunemente utilizzati 

per la realizzazione di biosensori di affinità (anticorpo/antigene) sono di tipo ottico ed 

acustico. I biosensori ottici sono di gran lunga i più diffusi grazie ad alcuni vantaggi legati 

alla natura della luce: innanzitutto la possibilità di vedere a livello micrometrico e sub 

micrometrico la distribuzione delle sonde molecolari con meccanismi di contrasto che 

vanno dalla fluorescenza al contrasto di fase; la non invasività dovuta alla non interazione 

di alcune lunghezze d’onda con le biomolecole; l’elevata sensibilità e così via. Se le 

tecniche di visualizzazione (“imaging”) e spettroscopia in fluorescenza rappresentano la 

storia dei saggi biologici e lo stato dell’arte nei microarray per applicazioni genomiche, vi è 

un crescente impulso scientifico e tecnologico, testimoniato da numerosi lavori scientifici 

negli ultimi cinque anni, per la realizzazione di dispositivi ottici che permettono di misurare 

direttamente le reazioni di riconoscimento biomolecolare senza dover ricorrere a dei 

marcatori fluorescenti. Queste tecniche “label free” offrono vantaggi non secondari 

derivanti proprio dall’evitare l’operazione di marcatura con fluorofori della molecola 

interessata: benché quest’operazione non sia particolarmente complicata, non sempre è 

possibile disporre di un saggio che prevede la marcatura del ligando, inoltre legare un 

fluoroforo ad una molecola significa perturbare in maniera non banale la struttura della 

stessa cambiandone così il comportamento biologico. 

Il progetto di ricerca oggetto della presente tesi di dottorato si inserisce in questo filone di 

ricerca biotecnologica internazionale ed è stato finalizzato in particolare alla realizzazione 

di un microsistema ottico per lo studio delle interazioni biomelocolari tra singole eliche di 

DNA oppure tra proteine e ligandi, sia per applicazioni genomiche e proteomiche, ma 

anche capace di configurarsi come biosensore abile all’individuazione di analiti bersaglio 

in miscele eterogenee. La realizzazione di un dispositivo del genere è un progetto 

complesso che può essere schematizzato nelle sue varie fasi come indicato in Figura 2: si 

parte dall’individuazione di un materiale di supporto che abbia opportune caratteristiche 

ottiche di trasduzione, che possa essere opportunamente funzionalizzato con sonde 

molecolari e che possa essere integrato in un microsistema completo di microfluidica. La 

realizzazione di questo microsistema ha richiesto un approccio fortemente multidisciplinare 

nonché l’interazione con figure professionali specializzate in ambiti scientifici 

complementari ma molto distanti tra loro come la fisica dei semiconduttori, la biologia 

Analyte 

iii

molecolare, la chimica delle superfici, l’ingegneria microelettronica. Il materiale scelto 

quale supporto ed elemento trasduttore è il silicio poroso: la sua caratteristica morfologica 

è quella di una struttura spugnosa interconnessa da nanocristalli di silicio che possono 

avere disposizione ordinata o quasi caotica a seconda delle procedure di produzione 

adottate. Questa morfologia tipica gli conferisce un’elevata superficie specifica che 

assicura un’efficace interazione con agenti esterni. Inoltre varia fortemente le sue 

caratteristiche chimico-fisiche quando esposto ad agenti biochimici. 

Le problematiche tecnico-scientifiche affrontate nel corso dei tre anni del dottorato, sono 

state quindi in particolare: 

1. la fabbricazione del silicio poroso, quale elemento di supporto del biosensore e 

sistema di trasduzione del segnale ottico; 

2. lo studio di una procedura affidabile di funzionalizzazione della superficie di silicio 

poroso in grado di ancorare in modo covalente le molecole sonda al materiale 

supporto; 

3. l’individuazione di sonde molecolari particolarmente stabili e compatibili con la 

realizzazione di biosensori riutilizzabili, quali ad esempio le proteine purificate da 

organismi estremofili; 

4. la messa a punto di opportuni protocolli sperimentali per la definizione dell’attività 

delle sonde una volta agganciate al supporto trasduttore; 

5. lo sviluppo di tecniche di microlavorazione del silicio, del vetro e di altri materiali 

eventualmente impiegati per la realizzazione di un prototipo di “lab-on-chip”, 

compatibili con l’utilizzo di sonde molecolari biologiche. 

1 

2 

3 

Reflectivity (a.u.) 

1,0 

0,8 

0,6 

0,4 

0,2 

0,0 

600 800 1000 1200 1400 1600 

Wavelength (nm) 

Figura 2: schematizzazione grafica del progetto di tesi 

L’integrazione di un biosensore con le tecnologie tipiche della microelettronica e dei circuiti 

integrati infatti non è mai una evoluzione diretta del dispositivo sensibile in un 

microsistema, perché l’impiego di sonde biologiche introduce una difficoltà in più legata 

alla stabilità delle stesse in condizioni di temperatura diverse da quelle ambientali. 

Il microsistema realizzato è stato utilizzato per rivelare l’ibridazione tra eliche di DNA 

complementari e per studiare gli eventi di riconoscimento molecolare tra proteina e 

ligando. In quanto segue si riportano le metodologie ed i principali risultati ottenuti durante 

il lavoro di tesi. 

Produzione e caratterizzazione del materiale Silicio Poroso (J3, J11) 

Il Silicio Poroso (PSi) è prodotto per dissoluzione elettrochimica di silicio cristallino 

drogato, ed ha una morfologia spugnosa che gli conferisce una superficie specifica molto 

elevata, dell’ordine di alcune centinaia di metri quadri per centimetro cubo. La dissoluzione 

4 

5 

iv

del silicio coinvolge due portatori liberi di carica (elettroni e e lacune h) come riportato nello 

schema di reazione riportato in Figura 3. 

Si + 2 F + 2 h 

SiF2 + 4 HF H2SiF6 + H2 + SiF2 SiF 2 + 4F SiF 6 2 + 2e 

2H + + 2e H 2 

Figura 3: semireazioni di dissoluzione elettrochimica del silicio cristallino 

Le dimensioni dei pori della struttura sono variabili da qualche nanometro fino a più di un 

micron a seconda del tipo di substrato di silicio cristallino utilizzato, della composizione 

della soluzione chimica e della corrente erogata nella cella durante il processo di 

fabbricazione. 

La porosità P è definita come la percentuale di aria presente in uno strato di materiale (ad 

esempio uno strato di silicio poroso, con porosità del 60 %, sarà composto da silicio 

nanocristallino solo per il 40 % in massa), e può essere misurata con una tecnica di tipo 

gravimetrico a partire dal rapporto tra la massa di silicio dissolto Mdis, durante il processo 

d’anodizzazione, e la massa totale di silicio sottoposto all’attacco MSi. La determinazione 

dello spessore e della porosità di uno strato prodotto con un singolo processo può essere 

effettuata alternativamente con un metodo di caratterizzazione ottica quale l’ellissometria 

spettroscopica ad angolo variabile. Per ottenere una differenziazione tra le varie tipologie 

di PSi, può essere utilizzata una classificazione in base al diametro medio d dei pori 

(standard IUPAC): 

1. Silicio microporoso (diametro medio del poro d < 2nm); 

2. Silicio mesoporoso (diametro medio del poro d compreso nell’intervallo 2 – 50 nm); 

3. Silicio macroporoso (diametro medio del poro d > 50nm). 

Spesso la classificazione fatta utilizzando soltanto la dimensione del diametro d del poro è 

riduttiva ed insufficiente per descrivere interamente la struttura dello strato poroso 

ottenuto. La dimensione dei pori, infatti, non contiene informazioni riguardo la morfologia. 

Con questo termine s’intende la forma, l’orientazione e l’interazione tra pori. Le proprietà 

termiche, elettriche, ottiche, del silicio poroso dipendono fortemente sia dal valore della 

porosità P che dalla morfologia della matrice spugnosa. Il parametro che influenza 

maggiormente la morfologia del PSi è il drogaggio del silicio cristallino di partenza: 

campioni poco drogati danno pori piccoli e solitamente allineati lungo una direzione 

privilegiata, mentre campioni molto drogati una volta attaccati si ramificano in una matrice 

caotica di tipo spugnoso. Alcune morfologie di silicio poroso sono mostrate come esempio 

in Figura 4. 

Figura 4: Possibili morfologie di film di silicio poroso su substrati di tipo p e n. 

Per ottimizzare l’interazione del silicio poroso con le soluzioni biologiche si è scelto di 

lavorare con un materiale mesoporoso e cioè costituito da pori con dimensioni medie 

v

dell’ordine dei 10-100 nm. Questo si ottiene dalla parziale dissoluzione di silicio di tipo p + , 

con elevato numero di portatori liberi di carica dunque, in una soluzione di acido fluoridrico 

ed etanolo in rapporto 1:1. Per ottenere strati di silicio poroso con porosità e spessore 

definiti è stato necessario, una volta fissata la composizione della soluzione elettrolitica ed 

il tipo e livello di drogaggio del substrato di silicio cristallino, calibrare il sistema di 

produzione e cioè costruire delle curve sperimentali che determinano la velocità di attacco 

(“etch rate” [nm/s] ) e la porosità, ottenuta al variare della densità di corrente applicata alla 

cella. 

Una delle caratteristiche peculiari del processo di produzione del silicio poroso è di essere 

“self-stopping”: se si interrompe l’attacco elettrochimico fermando l’erogazione di corrente, 

lo strato di poroso formatosi è svuotato dai portatori liberi di carica che hanno partecipato 

alla dissoluzione del silicio cristallino, per cui riprendendo ad erogare corrente il processo 

riparte dall’interfaccia silicio poroso-silicio cristallino, ove sono presenti i portatori di carica. 

In questo modo è possibile creare, durante un unico processo di attacco, strati successivi 

e adiacenti di silicio poroso con spessori e porosità diverse l’uno dall’altro partendo 

dall’alto (prima superficie esposta) verso il basso (fondo del wafer di silicio cristallino). Il 

buon controllo del processo di attacco permette la progettazione e la realizzazione di 

strutture in PSi con specifiche caratteristiche ottiche. È possibile realizzare dispositivi a 

singolo strato, questi si comportano dal punto di vista ottico come degli interferometri 

(Fabry-Perot) il cui spettro in riflessione è riportato in Figura 5 (a) o progettare dispositivi 

multistrato a specifiche lunghezze d’onda di risonanza, quali microcavità ottiche (b), 

specchi di Bragg (c) e sequenze Thue-Morse (d). 


1.0 

0.8 

0.6 

0.4 

0.2 

0.0 

a 

Reflectivity (a.u) 

1.0 

0.8 

0.6 

0.4 

0.2 

monostrato 

600 800 1000 1200 1400 1600 


specchio di Bragg 

600 800 1000 1200 1400 1600 

b 


2.0 

1.5 

1.0 

0.5 

0.0 

microcavità ottica 

800 1000 1200 1400 


600 800 1000 1200 1400 1600 



c 

d 

Figura 5: schematizzazioni e spettri in riflessione dei dispositivi in silicio poroso ottenuti illuminando con luce 

bianca. 

Lo spettro in riflessione si ottiene facendo incidere sul dispositivo un fascio di luce 

collimato, proveniente da una sorgente ad ampio spettro (400-1600 nm), ed inviando la 

luce riflessa ad un analizzatore ottico di spettro secondo lo schema riportato in Figura 6. 


1.0 

0.8 

0.6 

0.4 

0.2 

0.0 

Thue Morse S6 

Figura 6: apparato sperimentale utilizzato per misurare gli spettri ottici di riflessione dei campioni in PSi 

vi

I sensori ottici basati sul silicio poroso sfruttano le variazioni delle proprietà dielettriche di 

questo materiale quando esso è esposto all’analita che penetra nei suoi pori. (Figura 7) 

Reflectivity (a. u.) 

1.1 

1.0 

0.9 

0.8 

0.7 

0.6 

0.5 

0.4 

cavità 

imperturbata 

Clorobenzene 

1275 1300 1325 1400 1425 1450 1475 1500 


A 

Peak shift (nm) 

12 

10 

8 

6 

4 

2 

0 

Limit of solubility in water 

Atrazine in H 2 O 

S = 1.19 (3) nm/ppm 

LOD = 0.17 ppm 

Atrazine in HA 

S = 0.51 (2) nm/ppm 

LOD = 0.2 ppm 

0 5 10 15 20 25 30 35 

C (ppm) 

B 

Figura 7: schematizzazioni del meccanismo di riconoscimento; A) spettri in riflessione di una micro cavità 

ottica sottoposta a vapori di sostanze organiche; B) curva dose-risposta del sensore sottoposto a 

concentrazioni diverse di pesticida in soluzione acquosa. 

L’elevato rapporto superficie-volume del silicio poroso e la risposta ottica dei dispositivi, 

permettono di monitorare gli effetti dei cambiamenti in indice di rifrazione dello strato di 

poroso, dovuti alla presenza di sostanze liquide o gassose.La sostituzione dell’aria 

presente nei pori con le molecole delle sostanze organiche, provoca uno spostamento 

verso lunghezze d’onda maggiori come indicato nella Figura 7A. 

Dal momento che il meccanismo di rivelazione non è selettivo e non riesce a distinguere 

analiti presenti in miscele complesse è necessario implementare la selettività del silicio 

poroso utilizzando sonde molecolari ad alta selettività che possono essere legate 

covalentemente alla sua superficie con opportune tecniche di funzionalizzazione chimica. 

Funzionalizzazione chimica della superficie del Silicio Poroso (J6, J12, J21) 

La superficie del silicio poroso appena prodotto è fortemente reattiva in quanto idrogenata, 

presenta cioè legami Si-H che sono termodinamicamente instabili: tendono infatti a reagire 

con l’ossigeno presente in atmosfera per creare legami tipo Si-O-Si molto più stabili. 

Inoltre una superficie idrogenata è altamente idrofobica, pertanto una passivazione 

chimica o fisica ha la funzione non solo di indurre specifici legami tra gli atomi di silicio ed 

altre specie chimiche in modo da essere capace di legare selettivamente composti chimici 

o molecole di origine biologica, ma anche di rendere la superficie idrofilica in modo da 

permettere una facile penetrazione del materiale biologico, spesso in soluzione acquosa. 

In Figura 8 sono riportate le misure di angolo di contatto della superficie del PSi appena 

prodotto (A), superficie altamente idrofobica (θ=131°), e dopo funzionalizzazione chimica 

(B) da cui si ottiene una superficie idrofilica (θ=42°). 

A B 

Figura 8: caratterizzazione della superficie del PSi tramite misure di angolo di contatto. A) la superficie del 

PSi appena prodotto si presenta altamente idrofobica. B) dopo il processo di passivazione la superficie del 

PSi diventa idrofilica permettendo così la penetrazione delle soluzioni biologiche. 

vii

Un caso particolarmente interessante è quello in cui i legami con l’idrogeno sono sostituiti 

con legami Si-C i quali sono molto stabili sia rispetto alle variazioni di temperatura sia 

rispetto alle interazioni con gruppi nucleofili che hanno grande potere ossidante. 

Per individuare la procedura di modifica chimica della superficie di silicio poroso in grado 

di ancorare in modo covalente le molecole sonda, sono state sperimentate quattro 

tecniche di funzionalizzazione dei dispositivi ottici: 

1. una funzionalizzazione chimica ad opera dei reattivi di Grignard; 

2. una funzionalizzazione fotochimica, indotta da radiazione ultravioletta e condotta 

in atmosfera inerte immergendo il campione in una soluzione di estere 

succinimmidico di acido undecenoico; 

3. una funzionalizzazione in situ, ottenuta introducendo il linker organico, sotto 

forma di acido carbossilico, direttamente nella soluzione di attacco 

elettrochimico; 

4. una funzionalizzazione della superficie del silicio poroso ossidata ad opera di 

alcossi silani. 

Per analizzare la composizione superficiale del PSi si utilizza la tecnica della spettroscopia 

infrarossi a trasformata di Fourier (FTIR), che permette di acquisire informazioni sia 

quantitative sia qualitative riguardo alla chimica della superficie. In Figura 9 è mostrato un 

esempio di caratterizzazione prima e dopo il processo di funzionalizzazione, si monitora la 

scomparsa dei picchi caratteristici del PSi appena prodotto (SiH a 2100cm -1 ) a favore della 

comparsa di quelli caratteristici del linker chimico utilizzato. 

% Riflettività 

140 

120 

100 

80 

60 

40 

20 

0 

Bragg appena prodotto 

Bragg dopo funzionalizzazione con UANHS 

2855-2929 cm -1 CH 2 

1723 cm -1 

O-C=O 

1641-1563 cm -1 

succinimmide 

1288 cm 

-20 

4000 3500 3000 2500 2000 1500 1000 

-1 1038 cm 

CN; C=O 

-1 

N-O 

numero d'onda (cm -1 ) 

Figura 9: spettro FTIR di un dispositivo in PSi prima (nero) e dopo (rosso) la funzionalizzazione chimica. 

Sulla base dei gruppi funzionali disponibili sulla sonda da immobilizzare (DNA singola 

eliche, proteine, enzimi), è possibile individuare una procedura di passivazione chimica 

adatta al processo di funzionalizzazione: per esempio, è possibile funzionalizzare la 

superficie del silicio poroso con un silossano -NH2 terminale in modo da consentire 

l’ancoraggio di biomolecole o crosslinker carbossi terminali. In Figura 10 si riporta uno 

schema dei passaggi di funzionalizzazione a partire dal PSi appena prodotto fino 

all’immobilizzazione delle sonde molecolari. 

2100 cm -1 SiH 

viii

Si H 

OH 

O2 5% APTES 

Si H 

SiO2 OH 

15' @900°C 20' R.T 

Si H 

OH 

O 2 Si 

O2Si O 

O 

O 

O 

Si 

O 

O 

O 

Si 

O 

C 

N 

NH 2 

2.5% GA 

30' R.T 

or PROBE-COH 

C O 

H 

PROBE-NH 2 

Figura 10: schema di funzionalizzazione chimica della superficie del PSi 

Un’alternativa ai metodi classici di funzionalizzazione è rappresentata dalla possibilità di 

infiltrare un polimero biocompatibile contenente pendagli funzionali regolarmente spaziati 

lungo la catena. In questo modo i gruppi funzionali sono reattivi e disponibili per 

l’ancoraggio delle sonde biologiche. 

Il polimero scelto per gli esperimenti in questo lavoro di tesi, è il Poli(ε-caprolattone) (PCL), 

polimero alifatico idrofobico il cui impiego è stato ampiamente studiato negli ultimi anni, in 

campi di applicazione che vanno dalla realizzazione di tessuti ingegnerizzati al rilascio 

controllato di farmaci. 

Sono state messe a punto le condizioni di deposizione per rotazione veloce (“spin 

coating”). Questa tecnica è utilizzata per depositare un film sottile ed uniforme su un 

substrato solido piano. La quantità in eccesso di una soluzione molto diluita della specie 

che si vuole depositare viene depositata sul substrato, che è successivamente messo in 

rapida rotazione, al fine di distribuire il fluido sul substrato per effetto della forza centrifuga. 

È stata quindi determinata la concentrazione ottimale del polimero per avere un film con 

uno spessore confrontabile con la dimensione dei pori, dell’ordine dei 10 nanometri. In 

Figura 11 sono riportati lo schema di deposizione per spin coating (a) ed il grafico della 

variazione dello spessore del film al variare della concentrazione di polimero (b). 

thickness (nm) 

30 

25 

20 

15 

10 

5 

0 1 2 3 4 5 

PCL concentration (mg/ml) 

a b 

Figura 11: a) Schema di deposizione di film sottili per Spin coating; b) variazione dello spessore in funzione 

della concentrazione di PCL. 

La caratterizzazione in spettroscopia ellissometrica ad angolo variabile mostra che 

l’infiltrazione di PCL porta ad una riduzione di porosità da 79% fino ad un valore di 75.2 % 

con una disomogeneità verticale nella distribuzione del polimero. Il polimero infatti penetra 

nell’intera struttura ma si accumula sul fondo dove la presenza di idrogeno è maggiore, a 

causa dell’interazione idrofobica tra polimero e superficie idrogenata del PSi. Attraverso 

misure di angolo di contatto, variando la concentrazione del polimero infiltrato, si definisce 

come varia la bagnabilità della superficie di PSi appena prodotto: all’aumentare della 

ix

concentrazione aumenta la bagnabilità della superficie passando così da una superficie 

idrofobica ad una idrofilica. 

Essendo l’idrossido di sodio in soluzione acquosa un agente altamente corrosivo per il 

silicio poroso, è stata condotta una prova di resistenza ad una soluzione di [NaOH]=0,1M; 

sono stati realizzati due campioni di PSi gemelli, in uno è stato infiltrato il PCL mentre nel 

controllo no. Il campione infiltrato ha resistito per 18 minuti alla corrosione dovuta 

all’NaOH, raggiungendo poi una stazionarietà. È stata infiltrata anche una struttura 

multistrato in PSi, ed in particolare un Bragg costituito da 15 coppie di monostrati a due 

porosità diverse. Il dispositivo ha resistito 15 minuti (Figura 12), mantenendo la classica 

risposta ottica della struttura risonante. 



1 

0 

1.5 

1.0 

0.5 

PSi/PCL 

dopo 15 min in NaOH 

600 800 1000 1200 1400 1600 

PSi 

dopo 80s in NaOH 

0.0 

600 800 1000 1200 1400 1600 


Figura 12: Confronto tra un dispositvo di silicio poroso infiltrato (A) ed uno no (B) prima (nero) e dopo 

l’immersione in una soluzione di NaOH 0.1 M (rosso). 

Sul campione è stato infine condotto un esperimento di sensing di solventi organici, per 

verificare che il trattamento alcalino non avesse alterato la struttura porosa del dispositivo. 

Il Bragg funziona ancora come sensore di composti chimici volatili, poiché sottoponendo la 

struttura ad atmosfera satura di solventi organici questi hanno provocato uno spostamento 

dello spettro di riflessione verso lunghezze d’onda maggiori, chiaro segno di una 

sostituzione dell’aria all’interno dei pori con una sostanza ad indice di rifrazione maggiore. 

Lo spettro si sposta verso il rosso in maniera proporzionale all’aumentare dell’indice di 

rifrazione delle sostanze, con una sensibilità di 585 nm/RIU (RIU: Unità di Indice di 

Rifrazione) ed un LOD (Limite di rilevabilità) di 5*10 -4 RIU. 

Sonde biologiche utilizzate 

Sono state studiate, in collaborazione con il gruppo di ricerca del Prof. P. Arcari del 

Dipartimento di Biochimica e Biotecnologie Mediche dell’Università di Napoli “Federico II”, 

le interazioni ssDNA-cDNA e DNA/PNA 

DNA: è stata utilizzata una sequenza singola di DNA (ssDNA: 

5’GGACTTGCCCGAATCTACGTGTCCA3’,Primm), questa per gli esperimenti di 

fluorescenza è stata marcata con Fluoresceina CY3.5 (con picco di assorbimento a 

581 nm e di emissione a 596 nm). 

ed in collaborazione col gruppo del dr. S. D’Auria dell’Istituto di Biochimica delle Proteine 

del Consiglio Nazionale delle Ricerche le interazioni proteina-proteina quali: GlnBP- 

Glutamina, GlnBP-Gliadina, e TMBP-Glucosio. 

Glutamine-binding protein (GlnBP): la GlnBP, purificata da E. coli è una proteina 

monomerica composta da 224 residui amminoacidici (26 kDa) responsabile del 

primo step nel trasporto attivo della L-glutamina (Gln) attraverso la membrana 

citoplasmatica. La Gln è la risorsa principale di azoto e carbonio nel mezzo di 

coltura cellulare: il suo monitoraggio è quindi fondamentale nel controllo dei 

bioprocessi. Tra tutti gli amminoacidi presenti in natura la GlnBP lega unicamente la 

Gln con una costante di dissociazione Kd di 5 x 10 -9 M. È stata utilizzata anche per 

x

verificare la sua capacità di legare con alta affinità gli amminoacidi presenti nella 

gliadina (proteina presente in quasi tutti i cereali derivata dal glutine e ritenuta 

responsabile del morbo celiaco) costituita da una sequenza polimerica di gln. 

D-trehalose/D-maltose-binding protein (TMBP): la TMBP è una componente del 

sistema rifornimento del trealosio (Tre) e del maltosio (Mal) nell’ipertermofilo T. 

litoralis. La TMBP estratta da T. litoralis è una macromolecola (48 kDa) costituita da 

due domini e contenente 12 residui di triptofano. La struttura tridimensionale della 

TMBP è comune ad un gran numero di altre proteine atti a legare gli zuccheri. È 

stato recentemente studiato che la TMBP è capace di legare anche le molecole di 

glucosio oltre a quelle dei due zuccheri, dimeri del glucosio. Poiché il sangue 

umano non contiene né Tre né Mal, si è pensato di realizzare un biochip non 

invasivo per il monitoraggio del glucosio utilizzando questa proteina estratta da 

estremofili e per questo più stabile rispetto ad altre. 

Studio delle interazioni biomolecolari con dispositivi ottici (J2, J4, J5, J8, J14, J16, 

J17, J19, J20, J23) 

Una volta messa a punto la procedura di modifica chimica della superficie del silicio 

poroso sono stati condotti esperimenti di immobilizzazione delle sonde biologiche (DNA 

singola eliche, proteine, enzimi). Inizialmente sono stati utilizzati campioni marcati con 

fluorofori, al fine di verificare, con l’ausilio di un macroscopio a fluorescenza, l’efficienza e 

la stabilità della funzionalizzazione della superficie di silicio poroso. 

L’obiettivo finale del progetto è la realizzazione di un sistema ottico di analisi delle 

interazioni biologiche che non richieda l’uso di marcatori: le interazioni specifiche sondaligando 

vengono studiate mediante riflettometria in luce bianca in strutture fotoniche con 

precisa risposta ottica, funzionalizzate con sonde non marcate. L’incremento dell’indice di 

rifrazione medio del dispositivo, dovuto all’introduzione di materiale organico nei pori del 

dispositivo, che ivi permane anche dopo lavaggi in quanto agganciato alla rispettiva 

sonda, è rilevabile sperimentalmente da uno spostamento verso lunghezze d’onda 

maggiori dello spettro in riflessione ottenuto facendo incidere sul dispositivo un fascio di 

luce collimato, proveniente da una sorgente ad ampio spettro. 

La prima parte della sperimentazione è consistita nel determinare la concentrazione 

ottimale delle sonde per ottenere un ricoprimento uniforme della superficie di silicio 

poroso. A tal fine sono state utilizzate sonde marcate con un opportuno gruppo cromoforo 

(Fluoresceina CY3.5, o Rhodamina), tramite la misura dell’intensità di fluorescenza in 

funzione della concentrazione di sonda, prima e dopo tre lavaggi nelle soluzioni buffer e 

due dialisi della durata di una notte. In questo modo, sono state stabilite le concentrazioni 

di saturazione. Passo successivo è, dopo aver legato covalentemente la sonda non 

marcata alla superficie del silicio poroso, la rivelazione dell’interazione molecolare con il 

rispettivo ligando misurando lo spostamento dello spettro di riflessione del dispositivo dopo 

il processo di interazione che avviene direttamente su chip a temperatura ambiente. 

Gli esperimenti di riconoscimento molecolare sono stati ripetuti con diverse concentrazioni 

di ligando. È stato così possibile ottenere curve dose-riposta (Figura 13b) calcolando, per 

ogni concentrazione di ligando, la differenza di cammino ottico causata dalla variazione 

dell’indice di rifrazione medio del silicio poroso dovuta all’introduzione del materiale 

biologico (Figura 13a). Per verificare che la variazione dello spettro di riflessione non fosse 

dovuta ad interazioni aspecifiche dei ligandi nella struttura porosa, è stata ripetuta la 

procedura sperimentale utilizzando dei controlli negativi, nei quali non si registrava alcuna 

variazione dello spettro di riflessione. 

xi


1.4 

1.2 

1.0 

0.8 

0.6 

0.4 

0.2 

ss-DNA 

ss-DNA+c-DNA 6μM 



0.0 

1150 1200 1250 1300 1350 


Δλ (nm) 

35 

30 

25 

20 

15 

10 

5 

LOD cDNA = 350nM 

S cDNA =0.85 (4) nm/μM 

0 

0 20 40 60 80 

c-DNA concentration (μM) 

A B 

Figura 13: A) spostamento dello spettro di riflessione dovuto all’interazione ssDNA-cDNA. B) curva Doseresposta 

in funzione della concentrazione di cDNA. 

È stato possibile determinare per ogni coppia sonda-ligando i limiti di rivelazione (LOD) del 

nostro dispositivo e la sensibilità (S) di misura dello stesso. 

Di seguito è riportata la tabella con i valori di concentrazione per le sonde ed i valori di 

LOD ed S calcolati secondo la metodologia IUPAC (IUPAC Compendium of Analytical 

Nomenclature, Definitive Rules, 3d Edition, Section 10.3.3.3.14 “Limit of Detection”,1997). 

Sonda Concentrazione Ligando Controllo 

negativo 

LOD S 

DNA 60 μM cDNA ncDNA 350nM 0.85 nm/μM 

GlnBP 1mM Gln Glucosio 9nM 3*10 -2 nm/nM 

GlnBP 1mM Gliadina Prolamina del riso 670pM 448 nm/μM 

TMBP 7.5 μM Glucosio Gln 10 μM 0.03 nm/μM 

Integrazione del silicio poroso in un lab-on-chip (J1, J7, J9, J13, J15) 

La realizzazione di sistemi multifunzionali complessi ove circuiti e componenti elettronici 

ed ottici, ed elementi di sensing coesistono e operano sinergicamente, è di importanza 

strategica in diversi settori (produzione-automazione, trasporti, medicina, biochimica, 

sicurezza e ambiente). La diffusione di tali sistemi multifunzionali, in grado di operare nei 

più svariati ambienti, anche potenzialmente pericolosi, e che svolgono molteplici funzioni 

di monitoraggio e di diagnostica, presuppone una drastica riduzione dei costi ed un 

significativo incremento della loro affidabilità. 

L’attività di ricerca relativa a quest’ultima fase del lavoro di tesi è stata volta alla 

progettazione, realizzazione e caratterizzazione di microsistemi dotati di funzionalità di 

trasduzione ottica, basati sulle tecnologie di fabbricazione del silicio e dei materiali 

tecnologicamente compatibili. Poiché la produzione di microsensori ottici, integrabili su 

chip con altri componenti optoelettronici e microfluidici, necessita della compatibilità con i 

processi ed i circuiti microelettronici, è stato affrontato il problema dell’integrazione 

dell’elemento sensibile in silicio poroso in un microsistema ottico in cui tutte le funzioni 

(alimentazione, analisi, spurgo, e così via) per il funzionamento del trasduttore sono 

miniaturizzate in un microdispositivo. 

Per la realizzazione del dispositivo sono state ottimizzate alcune delle tecniche classiche 

della microelettronica e delle microlavorazioni, quali gli attacchi anisotropi in KOH, la 

fotolitografia e gli attacchi acidi del vetro. 

È stato inoltre definito un processo di incollaggio anodico “anodic bonding” per sigillare il 

silicio con vetro trasparente compatibilmente con le caratteristiche del silicio poroso. 

Il silicio poroso infatti per la sua morfologia spugnosa è abbastanza fragile dal punto di 

vista meccanico; inoltre la sua superficie idrogenata tende ad ossidarsi facilmente alle 

temperature tipiche dell’“anodic bonding” (400-800 °C) e poiché l’ossido cresce per un 

xii

60% al di sopra della superficie occupa parte dei pori con conseguente perdita di 

superficie specifica e quindi di sensibilità. 

Una sperimentazione dedicata ha permesso di stabilire la temperatura minima (250 °C) ed 

il tempo necessario (2 min.) per ottenere un processo tale da sigillare efficacemente il 

dispositivo con il vetro. Lo schema del dispositivo realizzato, insieme con il flusso di 

processo per la fabbricazione, è mostrato in Figura 14a, mentre in Figura 14b è riportata 

una fotografia del microsistema su banco ottico. 

Figura 14: a) Schema del microsistema ottico e flusso di processo per la fabbricazione; b) fotografia del 

microsistema su banco ottico 

L’utilizzo di un microsistema di questo tipo riduce enormemente i tempi di risposta del 

dispositivo la cui durata è essenzialmente dovuta alla diffusione delle specie biochimiche 

da analizzare all’interno della camera di test dove si trova l’elemento sensibile. Risultati 

ancora migliori possono essere ottenuti alimentando il sensore con un attuatore 

pneumatico: il fluido da analizzare, trasportato da un flusso inerte, arriva sul trasduttore 

solo in determinati intervalli di tempo che possono essere resi piccoli a piacere. 

La realizzazione di una matrice di elementi trasduttori richiede la messa a punto di un 

processo fotolitografico ed uno di attacco per la definizione delle strutture micrometriche. 

La fotolitografia (come si vede in Figura 15) è quel processo che consente il trasferimento 

del disegno da realizzare sul silicio da una maschera ad un sottile strato di materiale 

sensibile alla radiazione (resist), che ricopre la superficie del substrato. 

Figura 15: Schema del processo fotolitografico 

Una delle tecniche base per la microlavorazione del silicio è quella che utilizza soluzioni 

acquose alcaline. Comunemente l’attacco umido anisotropo include soluzioni acquose di 

KOH, miscele di KOH e IPA (alcol isopropilico), oppure EDP (Etilendiammina e 

Pirocatecolo). Tali soluzioni rendono possibile l’etching del silicio, il quale viene attaccato 

con differenti velocità nelle varie direzioni cristallografiche. 

xiii

Di fondamentale importanza è la scelta della maschera fotolitografica da utilizzare durante 

l’attacco, tenendo conto anche della velocità di incisione da parte della soluzione alcalina 

sullo strato di mascheramento. Per gli attacchi in KOH si usano generalmente maschere di 

nitruro di silicio, che non sono attaccabili, diossido di silicio che sono incise con una 

velocità di 1.4 nm/min e silicio di tipo p che offrono una riduzione sul silicio lievemente 

drogato della velocità di incisione tra 10:1 e 100:1, a seconda delle concentrazioni delle 

soluzioni e dalla temperatura del processo. 

Per la realizzazione di array di microvaschette con un volume di 10 μl, si è utilizzato una 

maschera di nitruro di silicio spessa 535 nm ed una soluzione acquose di KOH al 33% in 

peso alla temperatura di 80 ° C per 160 min. 

Il patterning, tuttavia, di campioni di PSi è fortemente limitato. La realizzazione del poroso 

richiede il contatto del campione con una soluzione HF in flusso di corrente, condizioni 

molto aggressive per le maschere litografiche comunemente utilizzate. È stato, pertanto, 

messo a punto e caratterizzato un processo fotolitografico atto a produrre una maschera 

protettiva da potersi utilizzare per un attacco elettrochimico. La definizione dei parametri di 

processo finalizzati alla realizzazione della maschera, dalla fotolitografia alla deposizione 

di film sottili, sono stati di estrema importanza, essendo questi dei processi tecnologici 

estremamente sensibili all’ambiente nel quale sono effettuati (variazioni anche minime 

nelle condizioni di temperatura, pressione ed umidità e la presenza di particelle volatili non 

desiderate minano la riuscita dell’operazione), si è scelto di lavorare in clean room, un 

laboratorio ad ambiente controllato e depolverizzato. 

Partendo da lavori riportati in letteratura è stata deposta e poi fotolitografata una maschera 

di nitruro di silicio. Le maschere composte da specie al nitruro sono state scelte poiché 

offrono una buona resistenza all’attacco a fronte di un basso costo di produzione ed 

un’elevata risoluzione per quel che riguarda il design del substrato. Il campione è stato 

ricoperto da uno strato di nitruro di silicio, depositato tramite deposizione chimica in fase 

vapore assistita dal plasma (PECVD), come maschera per l’attacco elettrochimico. Lo 

stesso è stato poi processato al fine di ottenere il design desiderato. Sono stati messi a 

punto i parametri di processo per la deposizione regolando opportunamente flussi di silano 

e ammoniaca. 

Le maschere realizzate avevano 400 fori di 200μm di diametro, spaziati tra di loro in 

entrambe le direzioni di 600μm. È stato quindi condotto l’attacco elettrochimico per 

ottenere monostrati da 500 nm (spessore minimo per un’analisi all’ellissometro) con 

porosità del 50% utilizzando una J pari a 50mA/cm 2 . Una foto dell’array realizzato, 

acquisita al microscopio ottico è riportata in Figura 16. 

Figura 16: Immagine dell’array di monostrati in PSi acquisita al microscopio ottico: ingrandimento 1.25X 

L’analisi riflettometrica, come già descritto, prevede l’utilizzo di una sorgente di luce 

bianca, da far incidere perpendicolarmente sul campione, ed un analizzatore ottico di 

spettro, al quale inviare la luce riflessa dal campione stesso. Questa connessione al 

campione avviene grazie ad una fibra ottica detta ‘ad Y’ l’unica che strutturalmente ci 

xiv

consente di inviare luce e convogliare quella riflessa, apparato e spettro ottico di 

riflessione in Figura 17. La luce, in questo caso, è stata focalizzata sul campione 

utilizzando un obiettivo 20x, capace di diminuire la focale e restringere lo spot per poter 

analizzare campioni di poroso di dimensioni micrometriche. 

A B 

Figura 17: A)apparato sperimentale utilizzato per misurare gli spettri ottici di riflessione dei campioni 

dell’array; B) confronto tra lo spettro di riflessione simulato (rosso) e quello sperimentale (nero). 

Il dispositivo ottico di array in PSi così realizzato potrà essere applicato in diversi settori, 

dalla genomica, la post-genomica, la proteomica e le analisi molecolari finalizzate alla 

diagnostica ad alta automazione per la medicina, la farmacogenomica e l'industria agroalimentare. 

Le soluzioni previste porteranno ad un miglioramento significativo del 

dispositivo in termini di sensibilità, riduzioni dei costi, velocità, predisposizione 

all'automazione, affidabilità e ripetibilità, in modo da poter monitorare 

contemporaneamente diversi analiti in miscele complesse. 

xv

Chapter 1 

Porous Silicon

Introduction 

The development of biochips is a major thrust of the rapidly growing biotechnology 

industry, which encompasses a very diverse range of research efforts including 

genomics, proteomics, computational biology, and pharmaceuticals, among other 

activities. Advances in these areas are giving scientists new methods for unraveling 

the complex biochemical processes occurring inside cells, with the larger goal of 

understanding and treating human diseases. At the same time, the semiconductor 

industry has been steadily perfecting the science of microminiaturization. The 

merging of these two fields in recent years has enabled biotechnologists to begin 

packing their traditionally bulky sensing tools into smaller and smaller spaces, onto 

so-called biochips. These chips are essentially miniaturized laboratories that can 

perform hundreds or thousands of simultaneous biochemical reactions. Biochips 

enable researchers to quickly screen large numbers of biological analytes for a 

variety of purposes, from disease diagnosis to detection of bioterrorism agents. 

The development of biochips has a long history, starting with early work on the 

underlying sensor technology. One of the first portable, chemistry-based sensors was 

the glass pH electrode, invented in 1922 by Hughes [1]. Measurement of pH was 

accomplished by detecting the potential difference developed across a thin glass 

membrane selective to the permeation of hydrogen ions; this selectivity was achieved 

by exchanges between H + and SiO sites in the glass. The basic concept of using 

exchange sites to create permselective membranes was used to develop other ion 

sensors in subsequent years. For example, a K + sensor was produced by 

incorporating valinomycin into a thin membrane [2]. Over thirty years elapsed before 

the first true biosensor (i.e. a sensor utilizing biological molecules) emerged. In 1956, 

Leland Clark published a paper on an oxygen sensing electrode [3]. This device 

became the basis for a glucose sensor developed in 1962 by Clark and colleague 

Lyons which utilized glucose oxidase molecules embedded in a dialysis membrane 

[4]. The enzyme functioned in the presence of glucose to decrease the amount of 

oxygen available to the oxygen electrode, thereby relating oxygen levels to glucose 

concentration. This and similar biosensors became known as enzyme electrodes, 

and are still in use today. 

In 1953, Watson and Crick announced their discovery of the now familiar double helix 

structure of DNA molecules and set the stage for genetics research that continues to 

the present day [5]. The development of sequencing techniques in 1977 by Gilbert 

and Sanger [6] (working separately) enabled researchers to directly read the genetic 

codes that provide instructions for protein synthesis. This research showed how 

hybridization of complementary single oligonucleotide strands could be used as a 

basis for DNA sensing. Two additional developments enabled the technology used in 

modern DNA-based biosensors. First, in 1983 Kary Mullis invented the polymerase 

chain reaction (PCR) technique[5], a method for amplifying DNA concentrations. This 

discovery made possible the detection of extremely small quantities of DNA in 

samples. Second, in 1986 Hood and coworkers devised a method to label DNA 

molecules with fluorescent tags instead of radiolabels [7], thus enabling hybridization 

experiments to be observed optically. 

The rapid technological advances of the biochemistry and semiconductor fields in the 

1980s led to the large scale development of biochips in the 1990s. At this time, it 

became clear that biochips were largely a "platform" technology which consisted of 

several separate, yet integrated components. Figure 1 shows the makeup of a typical 

biochip platform. The actual sensing component (or "chip") is just one piece of a 

complete analysis system. 

1

Figure 1: schematic of a biosensor 

Transduction must be done to translate the actual sensing event (DNA binding, 

oxidation/reduction, etc.) into a format understandable by a computer (voltage, light 

intensity, mass, etc.), which then enables additional analysis and processing to 

produce a final, human-readable output. The multiple technologies needed to make a 

successful biochip — from sensing chemistry, to microarraying, to signal processing 

— require a true multidisciplinary approach, making the barrier to entry steep. One of 

the first commercial biochips was introduced by Affymetrix. Their "GeneChip" 

products contain thousands of individual DNA sensors for use in sensing defects, or 

single nucleotide polymorphisms (SNPs), in genes such as p53 (a tumor suppressor) 

and BRCA1 and BRCA2 (related to breast cancer) [8]. The chips are produced using 

microlithography techniques traditionally used to fabricate integrated circuits (see 

below). 

Biochips are a platform that require, in addition to microarray technology, transduction and signal 

processing technologies to output the results of sensing experiments 

Today, a large variety of biochip technologies are either in development or being 

commercialized. Numerous advancements continue to be made in sensing research 

that enable new platforms to be developed for new applications. Cancer diagnosis 

through DNA typing is just one market opportunity. A variety of industries currently 

desire the ability to simultaneously screen for a wide range of chemical and biological 

agents, with purposes ranging from testing public water systems for disease agents 

to screening airline cargo for explosives. Pharmaceutical companies wish to 

combinatorially screen drug candidates against target enzymes. To achieve these 

ends, DNA, RNA, proteins, and even living cells are being employed as sensing

mediators on biochips. Numerous transduction methods can be employed including 

surface plasmon resonance, fluorescence, and chemiluminescence. The particular 

sensing and transduction techniques chosen depend on factors such as price, 

sensitivity, and reusability. 

Optical transduction is more and more used since photonic devices could be small, 

lightweight and thus portable due to the integrability of all optical components. 

Furthermore, optical devices do not require electric contacts. Fluorescence is by far 

the most used optical signalling method but a wide-use sensor can not be limited by 

the labelling of the probe nor the analyte, since this step is not always possible. 

Reagentless optical biosensors are monitoring devices which can detect a target 

analyte in a heterogeneous solution without the addition of anything than the sample. 

In the fields of genomics and proteomics this is a straightforward advantage since it 

allows real-time readouts and, thus, very high throughoutputs analysis. A label free 

optical biosensor can be realized by integrating the biological probe with a signalling 

material which directly transduces the molecular recognition event into an optical 

signal. 

There are many advantages of optical sensing. Optical sensing schemes are 

sensitive. By measuring differences in wavelengths or times, optical signals can be 

readily multiplexed. Some optical techniques, such as fluorescence, have intrinsic 

amplification in which a single label can lead to a million photons. In addition, some 

optical techniques are zero or “black” background in which the only source of signal 

is due to the presence of the species being measured, thereby enabling high 

sensitivity measurements. Finally, optical signals travel in an open path—no wires or 

other transmitting conduits are necessary. This feature enables remote 

measurements to be made. 

There are a variety of optical sensing transduction mechanisms including: 

• Luminescence 

• Fluorescence—intensity, lifetime, polarization 

• Phosphorescence 

• Fluorescence resonance energy transfer 

• Absorbance 

• Scattering 

• Raman-surface enhanced, resonance 

• Surface plasmon resonance 

• Interference 

Each of these techniques may have several variations. Another rubric for classifying 

optical sensors revolves around the categories of materials, surfaces, and arrays. In 

the materials field, there is a burgeoning effort in developing new materials for 

performing optical sensing. Such materials encompass work in the fields of polymers, 

ceramics, and semiconductors as well as more conventional inorganic and organic 

materials. These materials can be used as new recognition agents, as supports for 

immobilizing sensing materials, as materials for concentrating analytes, and as 

intrinsic optical sensors with integrated functionality including binding and signal 

transduction. This latter goal of creating materials that integrate binding with optical 

signal transduction is a major area of interest and investigation. 

In the surfaces area, a host of new phenomena are being discovered and employed 

for biosensing. 

It should be noted that there is a growing trend aimed at creating nanoscale devices. 

3

In the past two decades, the biological and medical fields have seen great advances 

in the development of biosensors and biochips for characterizing and quantifying 

biomolecules. The biochemistry-based issues may include the characteristics of the 

recognition molecules in terms of specificity, binding avidity, reversibility of binding, 

stability under conditions of use and storage, functionality after immobilization, signal 

generation, biochemical signal amplification, discrimination of specific from 

nonspecific binding events, reaction time, availability and cost. 

Using bioreceptors from biological organisms or receptors that have been patterned 

after biological systems, scientists have developed new methods of biochemical 

analysis that exploit the high selectivity of the biological recognition systems [9]. 

Arrays are being employed for an ever-increasing number of analytes. Ten years 

ago, the main focus on sensors was specificity and sensitivity. The need to measure 

multiple parameters was solved by bundling several sensors together in order to 

multiplex them. Today, arrays with tens of thousands, and even hundreds of 

thousands of features are commonplace in the DNA microarray field. Most such 

microarrays use fluorescence signals as the transduction mechanism. The ability to 

measure so many analytes simultaneously has revolutionized the thinking in sensing 

in general and optical sensing in particular. It is no longer sufficient in most 

applications to measure a single analyte as such a measurement captures only part 

of the information about a complex real-world sample. A variety of interesting and 

novel methods for preparing arrays have been developed for a wide variety of 

applications. Arrays also offer the ability to make replicate measurements to minimize 

false responses and to employ less selective sensing schemes coupled with 

intelligent processing to remove the requirement for absolute specificity. 

Silicon technology is already pervasive in our everyday lives. Silicon 'chips' often help 

us shop, cook and relax. They are involved in many forms of transport and control 

most forms of communication at the workplace. 

The choice of this thesis for the support material and contemporary for the transducer 

element fallen on porous silicon. 

• FABRICATION AND CHARACTERIZATION OF POROUS SILICON (PSi) 

MATERIALS 

The current interest in porous silicon (PSi) results from the demonstration of efficient 

visible photoluminescence of this material, first reported by Prof. Canham in 1990. 

However, PSi is not a new material: it was first reported over 40 years ago by Uhlir. 

During studies of the electropolishing of silicon in aqueous hydrofluoric acid (HF), he 

observed that the surface often became black, brown or red. More detailed studies 

were performed by Turner and Archer, but these films were not recognized as being 

PSi. It was Watanabe et al. who first reported their porous nature. Porous silicon, 

then, has been investigated for applications in microelectronics, optoelectronics, 

chemical and biological sensors, and biomedical devices. 

The in vivo use of porous silicon was first promoted by Leigh Canham, who 

demonstrated its resorbability and biocompatibility in the mid 1990s. Subsequently, 

PSi or porous SiO2 (prepared from PSi by oxidation) host matrices have been 

employed to demonstrate in vitro release of the steroid dexamethasone, ibuprofen 

and many other drugs. The first report of drug delivery from PSi across a cellular 

barrier was performed with insulin, delivered across monolayers of Caco-2 cells. An 

excellent review of the potential for use of PSi in various drug delivery applications 

has recently appeared.

The basic method to fabricate PSi is electrochemical dissolution of single crystalline 

silicon wafer in a hydrofluoridic acid electrolyte solution. This is obtained by 

monitoring either the anodic current (galvanostatic) or voltage (potentiostatic). 

In general, constant current is preferable as it allows better control of the porosity and 

thickness. It also provides better reproducibility from sample to sample. 

Pore morphology and pore size can be varied by controlling the current density, the 

type and the concentration of dopant, the crystalline orientation of the wafer, and the 

electrolyte concentration in order to form macro-, meso-, and micropores. Pore sizes 

ranging from 1 nm to a few microns can be prepared. 

In the simplest setup to fabricate PSi, plates of silicon and Pt are dipped into HF 

solution and an etching current is applied between these electrodes. The porous 

layer is formed in the surface of the silicon wafer, which is used as a positive anode. 

Usually a cathode is made of platinum and the fabrication cell has to be made of HFresistant 

material, for example Teflon. Dilute HF solution or ethanolic HF are 

generally used as electrolytes. Ethanol is added to reduce formation of hydrogen 

bubbles and to improve the electrolyte penetration in the pores and, thus, uniformity 

of the PSi layer. 

The mechanism of pore formation is not generally agreed upon, but it is thought to 

involve a combination of electronic and chemical factors. The type of dopant in the 

original silicon wafer is important because it determines the availability of valence 

band holes that are the key oxidizing equivalents in the reaction shown in Figure 2. 

Figure 2: schematic of the etch cell used to prepare porous silicon. The electrochemical half-reactions 

are shown. 

In general the relationships of dopant to morphology can be segregated into four 

groups based on the type and concentration of the dopant: n-type, p-type, highly 

doped n-type, and highly doped p-type. By “highly doped”, the dopant levels at which 

the conductivity behavior of the material is more metallic than semiconducting. 

The n-type silicon substrate has to be illuminated during anodization to generate 

photo-excited holes on the surface. For this silicon type wafers, with a relatively 

moderate doping level, exclusion of valence band holes from the space charge 

region determines the pore diameter. Quantum confinement effects are thought to 

limit pore size in moderately p-doped material. For both dopant types the reaction is 

crystal face selective, with the pores propagating primarily in the direction of 

the single crystal. 

5

A simplified mechanism for the chemical reaction in reported in Figure 3 [10]. 

Figure 3: mechanism of silicon oxidation during the formation of porous silicon. 

Application of anodic current oxidizes a surface silicon atom, which is then attacked 

by fluoride. The net process is a 4 electron oxidation, but only two equivalents are 

supplied by the current source. The other two equivalents come from reduction of 

protons in the solution by surface SiF2 species. Pore formation occurs as Si atoms 

are removed in the form of SiF4, which reacts with two equivalents of F − in solution to 

form SiF6 2− . 

The porosity and the thickness of the PSi layers are among the most important 

parameters which characterize porous silicon. The porosity is defined as the fraction 

of void within the PSi layer and can be determined easily by weight measurements. 

The virgin wafer is first weighted before anodization (m1), then just after anodization 

(m2) and finally after dissolution of the whole porous layer in a molar NaOH aqueous 

solution (m3). Uniform and rapid stripping in the NaOH solution is obtained when the 

PSi layer is covered with a small amount of ethanol which improves the infiltration of 

aqueous NaOH in the pores. The porosity is given simply by the following equation: 

% 

 

1 2 

1 3 

From these measured masses, it is also possible to determine the thickness of the 

layer according to the following formula: 

1 3 

 

where d is the density of bulk silicon and S the wafer area exposed to HF during 

anodization. 

The porosity of a growing porous Si layer is proportional to the current density being 

applied, and it typically ranges between 40 and 80%. Pores form only at the 

Si/porous Si interface, and once formed, the morphology of the pores does not 

change significantly for the remainder of the etching process. 

However, the porosity of a growing layer can be altered by changing the applied 

current. The film will continue to grow with this new porosity until the current 

changes. This feature allows the construction of layered nanostructures simply by

modulating the applied current during an etch. For example, one dimensional 

photonic crystals consisting of a stack of layers with alternating refractive index can 

be prepared by periodically modulating the current during an etch. 

The ability to easily tune the pore sizes and volumes during the electrochemical etch 

is a unique property of PSi that is very useful for biological applications. Other porous 

materials generally require a more complicated design protocol to control pore size, 

and even then, the available pore sizes tend to span a limited range. With 

electrochemically prepared PSi, control over porosity and pore size is obtained by 

adjusting the current settings during the etch. Typically, larger current density 

produces larger pores. Large pores are desirable when incorporating sizable 

molecules or drugs within the pores.(Figure 4) 

Figure 4: SEM images of typical porous silicon morphologies. 

• Stain etching 

Although the electrochemical anodization is commonly used in the fabrication of PSi, 

several other fabrication methods have been introduced. One of these methods in a 

so-called stain etching. The term stain etching refers to the brownish or reddish color 

of the film of porous Si that is generated on a crystalline silicon material subjected to 

the process [11]. The stain films are produced by immersion of silicon substrate in HF 

solution without any electrical bias. This method is even simpler than the previously 

presented anodization. However, the control of the porosity, layer thickness and pore 

size of PSi is quite limited. Stain etching in fact is less reproducible than the 

electrochemical process, although recent advances have improved the reliability of 

the process substantially [12]. Furthermore, stain etching cannot be used to prepare 

stratified structures such as double layers or multilayered photonic crystals. However, 

PSi powders prepared by stain etch are now commercially available 

(http://vestaceramics.net), and a few additional vendors are poised to enter the 

market. For the biomedically inclined researcher this eliminates the need to set up a 

complicated and hazardous electrochemical etching system, and it should stimulate 

the growth of the field. 

• PSi Photosynthesis 

Another zero-current etching method is photosynthesis. In this approach, silicon 

substrate is merely illuminated with intense visible light in a HF solution, without 

electrochemical bias. This is a very interesting method since it allows the fabricating 

of PSi microstructure without a mask, which is usually used in lithographic methods. 

7

After the etching, the next step to be considered is drying of the PSi material, if a 

liquid electrolyte has been used. Drying, if the electrolyte in the pores is allowed to 

evaporate at atmospheric pressure and temperature, may induce cracking and 

shrinkage. Especially the higher, brittle porosity structures (>75%) often do not have 

mechanical strength enough to survive the electrolyte evaporation without significant 

degradation in the structure. This is because the formed liquid vapour interface can 

generate large capillary stresses. In most cases (porosity

PSi optical sensors are based on changes of photoluminescence or reflectivity when 

exposed to the target analytes which substitute the air into the PSi pores. The effect 

depends on the chemical and physical properties of each analyte, so that the sensor 

can be used to recognize the pure substances. Due to the sensing mechanism, these 

kinds of devices are not able to identify the components of a complex mixture. In 

order to enhance the sensor selectivity through specific interactions, we have to 

chemically or physically modify the PSi surface. The common approach is to create a 

covalent bond between the porous silicon surface and the biomolecules which 

specifically recognize the target analytes [29]. 

The reliability of the biosensor strongly depends on the functionalization process: 

how fast, simple, homogenous and repeatable it is. This step is also very important 

for the stability of the sensor. 

Immediately after silicon electrochemically etched, the surface is covered with 

reactive hydride species. These chemical functionalities provide a versatile starting 

point for various reactions that determine the dissolution rate in aqueous media, 

allow the attachment of biological probes. The two most important modification 

reactions are chemical oxidation and grafting of Si-C species. In this way the 

substitution of the Si-H bonds with these ones guarantees a much more stable 

surface from the thermodynamic point of view. 

The integration of sensitive elements into complex microsystems, i.e. lab on chip or 

micro-total-analysis systems, is of straightforward interest, especially for the 

miniaturisation of each component. This step is never a trivial one: the fabrication 

process often requires high temperatures and mechanical stresses which can 

damage or even destroy the bio- or chemo- sensitive transducer structures. 

The anodic bonding (AB) is a standard IC fabrication technique which is widely used 

in microfluidic due to a wide spectrum of advantages among which the hermetic 

sealing [30]. Due to the good bonding quality, glass transparency, technological 

cleanness and high passivity to most of chemicals and biological substances, AB is 

also commonly exploited in the fabrication of lab-on-chip devices. The compatibility of 

this primary integration technique with highly demanding sensing micro-components 

plays a basic role in view of the realization of lab-on-chip devices. 

Testing and demonstrating the porous silicon capabilities as a useful functional 

material in the optical transduction of biochemical interactions is only the first action 

in the realisation of an optical biochip based on this nanostructured material. In this 

case, all the fabrication processes should be compatible with the utilisation of 

biological probes and the feasibility of such devices must be proven. This means that 

the standard integrated circuit micro-technologies should be modified and adapted to 

this new field of application in order to preserve the stability of the transducer 

element and all the biological components (probes and targets). The porous silicon 

low cost technology could thus provide a link between the conventional CMOS 

technology and the photonic devices in the realization of the so-called smart sensors 

and biochips. A very strong interdisciplinary approach is required to match and 

resolve all the technological problems. 

The final objective of this thesis is to design and fabricate different resonant optical 

structures, integrate them into a simple lab-on-chip devices based on the porous 

silicon nanotechnology. Some porous silicon surface modification strategies in order 

to realize an optical biosensor were exploited; the target was the fabrication of 

sensitive label-free biosensors, which are highly requested for applications in high 

throughput drug monitoring and diseases diagnostics, unlabeled analytes require in 

fact easy sample preparation. 

9

Materials and Methods: Porous Silicon’ production calibration 

Today, the electrochemical etching is a standard method to fabricate nanostructured 

porous silicon: a proper choice of the applied current density, the electrolyte 

composition, and the silicon doping allow precise control over the morphology and, 

consequently, on the physical and chemical properties of the porous silicon structure. 

Computer controlled electrochemical etching processes are exploited for the 

realization of porous silicon films of controlled thickness and porosity (defined as the 

percentage of void in the silicon volume). 

Nanoporous, mesoporous and macroporous structures can be achieved, with pore 

size ranging from few nanometers up to microns. Moreover, since the etching 

process is self-stopping, it is possible to fabricate with a single run process multilayer 

stacks made of single layers of different porosity. The dielectric properties of each 

PSi layer, and in particular its refractive index n, can be namely modulated between 

those of crystalline silicon (n = 3.54, porosity = 0) and air (n= 1, porosity = 100 %); so 

that alternating high and low porosity layers, lot of photonic structures, such as 

Fabry-Perot interferometers, omni-directional Bragg reflectors, optical filters based on 

microcavities, and even complicated quasi-periodic sequences (Thue-Morse) can be 

simply realized, as it is shown in Figure 5. 



1,0 

0,8 

0,6 

0,4 

0,2 

2,0 

1,5 

1,0 

0,5 

0,0 

600 800 1000 1200 1400 1600 


monolayer 

optical microcavity 

800 1000 1200 1400 



600 800 1000 1200 1400 1600 

Figure 5. Experimental reflectivity spectra of different PSi optical structures. 

In this study Fabry-Perot single layer, Bragg mirrors, Thue Morse and optical 

microcavities have been used as sensors. 

The porous silicon layer was obtained by electrochemical etching in a HF-based 

solution at room temperature. The substrate used was a highly doped p + -silicon, 

oriented, 0.01 Ω cm resistivity, 400 μm thick. Before anodization the substrate 

was placed in HF solution to remove the native oxide. With this substrate we can 

obtain mesoporous layers, with a pore average dimension of about 50 nm. In figure 6 

SEM images are reported. 

1,0 

0,8 

0,6 

0,4 

0,2 

0,0 


1.0 

0.8 

0.6 

0.4 

0.2 

0.0 

Bragg 


Thue Morse S6 

600 800 1000 1200 1400 1600 

Wavelength (nm)

Figure 6. SEM images of different PSi morphologies. The substrate p+ type is the chosen for this 

work. 

The first step of this thesis work was to set the parameters to realize and control the 

porous silicon fabrication process. 

It was checked for the best HF concentration in ethanol solution (HF : EtOH 1:1 and 

3:7). The calibration curve for establish the etch rate for each current density and the 

layer porosity were performed for each HF solution. In Figure 7 the two calibrations 

curves are reported to clarify. 

P (%) 

P (%) 

64 

62 

60 

58 

56 

54 

52 

50 

48 

46 

44 

42 

50 100 150 200 250 300 350 400 450 500 550 

76 

74 

72 

70 

68 

66 

64 

62 

60 

58 

56 

54 

n + 

J (mA/cm 2 ) 

0 20 40 60 80 100 120 140 160 180 

J 

p + 

A 

B 

E. R. (nm/s) 

E.R. (nm/s) 

500 

450 

400 

350 

300 

250 

280 

260 

240 

220 

200 

180 

160 

100 200 300 400 500 

J (mA/cm 2 ) 

0 20 40 60 80 100 120 140 160 180 

Figure 7. Silicon wafer: p+ type, orientation, 8-12 mW cm resistivity. A) Etching solution: 

HF/EtOH=50/50; B) Etching solution: HF/EtOH=30/70. 

n 

J 

p 

11

The morphology of the porous silicon layers was investigated by variable angle 

spectroscopic ellipsometer (UVISEL, Horiba-Jobin-Yvon) and scanning electron 

microscope (SEM-FEG Gemini 300, Carl Zeiss). 

Both the analytical techniques have showed the presence of a top layer on the 

porous silicon structure of thickness between 20 nm and 100 nm and porosity of 20- 

40 % due to hydrogen contamination of the silicon wafer [31]. Such film prevents not 

only the pores from filling but also any biochemical interaction with the hydrogenated 

porous silicon surface after the etching process. For sensing purposes it is therefore 

mandatory to avoid the formation of the parasitic film by thermal treating the wafer at 

300 °C in nitrogen atmosphere before the electrochemical etch [32]. 

In figure 8 a cross section SEM image of the PSi top surface is reported. The thin low 

porosity layer on the top of the porous silicon device is well evident. In the thermal 

treated silicon wafer, this top layer is absent. The same result is confirmed by 

ellipsometric measurements. 

20 nm 

Figure 8: Cross section scanning electron microscopy image of a porous silicon layer. 

Due to the nanostructured nature of PSi, it is essential to improve the pore infiltration 

of biomolecular probes: to this aim, the device was optimized employing a strong 

base post etch process. By immersing the device in an aqueous ethanol solution, 

containing millimolar concentration of KOH, for 15 min. This treatment produces an 

increase of about 15-20% the porosity without affecting the optical quality of the 

device [33]. The presence of Si-H bonds on the porous silicon surface has been 

monitored by means of infrared spectroscopy with a Fourier transform spectrometer 

(FT-IR Nicolet Nexus) [33]. 

The fresh etched porous silicon device shows the characteristic peaks of Si-H bonds 

at 910 cm -1 and 2100 cm -1 . The KOH post-etch treatment used to increase the 

porosity and to improve the pore infiltration by biological solutions, unfortunately 

removes most of these bonds and oxidizes the PSi. To restore the Si-H bonds the 

porous silicon device were rinsed in a low concentration HF-based solution (5 mM) 

for 30 s. 

Figure 9 shows the FT-IR spectrum of a porous silicon microcavity after KOH and HF 

treatments. The Si-H bonds, removed by KOH, are present again and the silicon 

dioxide fingerprint (at 1050-1100 cm -1 ) disappears.


160 

140 

120 

100 

80 

60 

40 

20 

2100 cm -1 

Si-H 

after KOH treatment 

after HF treatment 

silicon oxide 

913 cm -1 

Si-H 

3500 3000 2500 2000 1500 1000 

Wavenumber (cm -1 ) 

Figure 9: FT-IR spectrum of a porous silicon microcavity after KOH and HF post-etch treatments. 

The experimental set-up (see Figure 10) to characterize the PSi devices from an 

optical point of view is very simple: a tungsten lamp (400 nm < λ < 1800 nm) 

inquired, through an optical fiber and a collimator, the sensor closed in a glass vial 

which can be fluxed by liquids or gases. The reflected beam is collected by an 

objective, coupled into a multimode fiber, and then directed in an optical spectrum 

analyser (Ando, AQ6315A). The reflectivity spectra had been measured with a 

resolution of 0.2 nm. 

Figure 10: optical set-up 

In Figure 11 the optical reflectivity (as explicative example) spectra of a porous 

silicon layer as-etched and post KOH and HF treatments are reported, in order to 

show that the optical response of the devices is not affected by the strong chemical 

treatments. 

13

Reflectivity (u.a) 

1.0 

0.8 

0.6 

0.4 

0.2 

0.0 

sample as etched 

after KOH treatment 

after refresh in HF solution 

600 800 1000 1200 1400 


Figure 11: Reflectivity optical spectra of the porous silicon layer after KOH and HF treatments. 

In this chapter, are presented some experimental results about sensing of chemical 

substances, flammable and toxic organic solvents and also some hydrocarbons, 

using several porous silicon optical structures with different pores structures. 

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15

Thue-Morse: aperiodic multilayers made of porous silicon 

This section is aimed at the optimization of the fabrication process and the 

characterization of the Thue-Morse S0-S7 sequences by using the PSi technology.

phys. stat. sol. (c) 4, No. 6, 1966–1970 (2007) / DOI 10.1002/pssc.200674342 

Optical properties of porous silicon Thue-Morse structures 

Luca De Stefano *, 1 , Ilaria Rea 1 , Lucia Rotiroti 1 , Luigi Moretti 2 , and Ivo Rendina 1 

1 

Institute for Microelectronics and Microsystems, National Council of Research, Via P. Castellino 111, 

80131 Naples, Italy 

2 

DIMET “Mediterranea” University of Reggio Calabria, Località Feo di Vito, 89060 Reggio Calabria, 

Italy 

Received 17 March 2006, revised 15 September 2006, accepted 15 November 2006 

Published online 9 May 2007 

PACS 42.70.Qs, 78.66.Db, 78.67.–n, 81.05.Rm 

Dielectric aperiodic Thue-Morse structures up to 128 layers have been realized by the porous silicon technology. 

Normal incidence reflectivity measurements have been performed to investigate the photonic 

properties of the devices. A partial photonic band gap region, centered at 1100 nm and 70 nm wide has 

been observed for the S 6 and S 7 Thue-Morse structures. The S 6 multilayer has been studied as sensor device 

on exposure to several chemical substances. 

1 Introduction 

© 2007 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim 

In recent years, dielectric photonic band gap (PBG) structures have attracted great attention for their 

property to forbid the propagation of the light at fixed wavelengths and the possibility to realize an alloptical 

integrated circuit which is the basic element of the optical computer [1, 2]. The simplest, one 

dimensional PBG structure is the Bragg mirror which is made of alternating layers of high (nH) and low 

(nL) refractive index, whose thicknesses satisfy the Bragg relation 2(nHdH+nLdL) = λBm where m is the 

order of the Bragg condition. Lot of theoretical and experimental studies on these periodic structures 

have shown that they are characterized by the presence of only one degenerate PBG in a period of the 

reciprocal space, even if it can be omnidirectional [3]. On the contrary, multiple PBGs are typical of 

aperiodic structures. The main example of 1D aperiodic multilayer is given by the Thue-Morse sequence 

which can be obtained alternating high (A) and low (B) refractive index layers following the substitution 

rules A→AB and B→BA [4]. The lower-orders of the Thue-Morse sequence are: S0=A, S1=AB, 

S2=ABBA, S3=ABBABAAB, S4=ABBABAABBAABABBA, and so on. Dielectric Thue-Morse multilayers 

have been already demonstrated using Si/SiO2 and TiO2/SiO2 as high/low refractive index materials 

[5, 6] and recently also in porous silicon (PSi) [7]. In this study we have fabricated and characterized 

the Thue-Morse S0-S7 sequences by using the PSi technology. The PSi is a very attractive material due to 

the possibility of realize in only one process multilayered structures that exploit a high quality optical 

response. PSi is obtained by the electrochemical etching of crystalline silicon in a hydrofluoridric acid 

(HF) based solution. The porosity and the thickness of a single layer are linear functions of the current 

density and the anodization time for a fixed doping level of the silicon wafer and HF concentration. The 

refractive index of the PSi film depends on its porosity, and can be calculated in the frame of several 

effective medium approximations, like the Bruggemann or Maxwell-Garnett models [8]. Moreover, the 

PSi is an ideal material for sensing applications because it is characterized by a very large specific surface 

area (200-500 m 2 cm -3 ), so that an effective interaction with several substances is assured. 

* Corresponding author: e-mail: luca.destefano@na.imm.cnr.it, Phone: +390816132375, Fax: +390816132598 

© 2007 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim

phys. stat. sol. (c) 4, No. 6 (2007) 1967 

2 Materials and methods 

A highly doped p + -silicon, oriented, 0.01 Ω cm resistivity, 400 µm thick was used as substrate to 

realize the Thue-Morse structures. Samples were fabricated in dark light at room temperature using a 

solution of 30% volumetric fraction of aqueous HF (50% wt) and 70% of Ethanol. Before anodization, 

we have removed the thin film of native oxide from the silicon wafer by rapid rinsing it in a diluted HF 

solution. High porosity layers, with an average refractive index nH ≅ 1.3 and a thickness d ≅ 135 nm, 

were obtained applying a current density of 150 mA/cm 2 for 0.88 s. Low porosity layers, with an effective 

refractive index nL ≅ 1.96 and a thickness of d ≅ 90 nm, were obtained with a current density of 5 

mA/cm 2 for 0.53 s. The thickness d of each layer was designed to satisfy the Bragg condition dn= λB/4 

where n is the average refractive index and λB=700 nm. Thicknesses and porosities have been estimated 

by variable angle spectroscopic ellipsometry measurements on the single PSi layers. 

An Y optical reflection probe (Avantes), connected to a white light source and to an optical spectrum 

analyzer (Ando, AQ6315A), has been used for the reflectivity measurements at normal incidence. The 

reflectivity spectra were measured between 600 and 1600 nm with a resolution of 0.2 nm. 

Reflectivity measurements on exposure to volatile substances have been performed in a steel test chamber 

equipped with an optical access through a quartz window and in/out channels for gas feeding. 

3 Results and discussion 

A scheme of the porous silicon Thue-Morse sequences realized is reported in Fig. 1: the number of the 

layers increases, as 2 n where n is the Thue-Morse order, while the thickness of the devices is dSn = 2dSn-1 

for n > 1. The realized samples S0-S7 have thicknesses spanning the range between 0.135 µm and 14.4 

µm. 

Fig. 1 Scheme of the porous silicon Thue-Morse sequences realized and characterized in the present work. 

In Figs. 2 and 3 the experimental (solid line) and calculated (dash line) normal incidence reflectivity 

spectra are reported for S3 (2-a), S4 (2-b), S5 (2-c), S6 (3-d), and S7 (3-e) Thue-Morse structures which 

are constituted by 8, 16, 32, 64 and 128 layers, respectively. The reflectivity spectra have been reproduced 

by a transfer matrix method [9] including the wavelength dispersion of silicon. At the increasing 

of the layers number we can observe a worsening of the agreement between the experimental and calcu- 

www.pss-c.com © 2007 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim

1968 L. De Stefano et al.: Optical properties of porous Si Thue-Morse structures 

Reflectivity 

1.0 

0.5 

0.0 

600 

1.0 

800 1000 1200 1400 1600 

0.5 

(b) 

0.0 

600 

1.0 

800 1000 1200 1400 1600 

0.5 

0.0 

600 800 1000 1200 1400 1600 


Experimental 

Simulation 

(a) 

(c) 

Fig. 2 Experimental (solid line) and calculated 

(dashed line) normal incidence reflectivity spectra for 

S 3 (a), S4 (b) and S5 (c) Thue-Morse structures. 

0.0 

600 800 1000 1200 1400 1600 

© 2007 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim www.pss-c.com 

Reflectivity 

1.0 

0.5 

0.0 

600 800 1000 1200 1400 1600 

1.0 

0.5 


Experimental 

Simulation 

lated spectra. This effect is due to small porosity and thickness dishomogeneities of the deeper layers, 

and to scattering losses, caused by interfaces roughness. The spectrum of S3 Thue-Morse is characterized 

by two band gaps separated by a transmission peak at 1000 nm. At the increasing of the order of Thue- 

Morse sequence, the photonic band gap splits and very narrow transmission peaks appear. This is a well 

known effect due to the random nature of the aperiodic Thue-Morse: the eigenstates, i.e. the stationary 

configurations of the electromagnetic field, are more localized respect to the periodic and quasi-periodic 

geometrical structures, such as the Bragg and the Fibonacci geometries, respectively [10]. 

Transmission peak FWHM (nm) 

200 

175 

150 

125 

100 

75 

50 

25 

0 

0 25 50 75 100 125 150 

Layers number 

Fig. 4 Full width at half maximum of the transmission peaks in the experimental reflectivity spectra of S3-S7 sequences. 

(d) 

(e) 

Fig. 3 Experimental (solid line) and calculated 

(dashed line) normal incidence reflectivity spectra for 

S 6 (d) and S 7 (e) Thue-Morse structures.


The Fig. 4 reports the full width at half maximum (FWHM) of the transmission peaks present in the 

reflectivity stop bands spectra of the S3-S7 Thue-Morse structures: on increasing the number of the layers 

the peak width decreases exactly as an exponential function (R 2 = 0.998). 

The presence of resonant peaks in the reflectivity spectrum suggests the use of such porous structures as 

optical transducers for gas and liquid monitoring. The S6 Thue-Morse structure has been investigated as 

sensor device on exposure to several vapors [11, 12]. The gaseous substances penetrate into the nanometric 

pores, substituting the air. The average refractive indexes of the layers change, so that the reflectivity 

spectrum of the device shifts towards higher wavelengths. In Figure 5 the red-shift of the S6 structure 

transmission sharp peak at 1000 nm is reported as a function of the refractive index of each volatile substance. 

A linear response is observed and a sensitivity of 530 (50) nm/RIU (Refractive Index Unit) can 

be estimated. 

Table 1 Central wavelength and width of the photonic and fractal theoretical stop band. 

TM 

PBG I FBG II FBG 

Structure λ (nm) Width (nm) λ (nm) Width(nm) λ (nm) Width (nm) 

S3 766 161 1307 452 0 0 

S4 880 149 1138 324 0 0 

S5 840 111 1060 210 1240 146 

S6 870 114 1084 218 1250 112 

S7 813 119 1029 230 1216 125 

Red Shift (nm) 

220 

210 

200 

190 

180 

Refractive Index Unit (RIU) Sensitivity 

Methanol 1.328 

Pentane 1.358 

Isopropanol 1.377 

Isobutanol 1.396 

1.32 1.34 1.36 

Refractive Index 

1.38 1.40 

Fig. 5 Red-shift of the S6 structure transmission peak at 1000 nm versus the refractive index of vapor substances. 

4 Conclusions 

In conclusion, in this work we have reported the fabrication and the optical characterization of aperiodic 

Thue-Morse structures up to 128 layers based on PSi technology. The multilayers are relatively easy to 

realize and exploit a high quality optical response. A splitting of photonic band gap is observed at the 

increasing of the Thue-Morse order. A partial photonic band gap region, centered at 1100 nm and 70 nm 

wide was observed for the S6 and S7 Thue-Morse structures. The PSi Thue-Morse structures can be also 

successfully used as sensor devices due to their characteristic properties. 


1970 L. De Stefano et al.: Optical properties of porous Si Thue-Morse structures 

References 

[1] E. Yablonovitch, Phys. Rev. Lett. 58, 2059 (1987). 

[2] S. John, Phys. Rev. Lett. 58, 2486 (1987). 

[3] V. Agarwal and J. A. del Rio, Appl. Phys. Lett. 82, 1512 (2003). 

[4] M. Dulea, M. Severin, and R. Riklund, Phys. Rev. B 42, 3680 (1990). 

[5] L. Dal Negro, M. Stolfi, Y. Yi, J. Michel, X. Duan, L. C. Kimerling, J. LeBlanc, and J. Haavisto, Appl. Phys. 

Lett. 84, 5186 (2004). 

[6] F. Qui, R. W. Peng, X. Q. Huang, X. F. Hu, Mu Wang, A. Hu, S. S. Jiang, and D. Feng, Europhys. Lett. 68, 

658 (2004). 

[7] M. E. Mora-Ramos, V. Agarwal, and J. A. Soto Urueta, Microelectron. J. 36, 413 (2005). 

[8] R. W. Hardeman, M. I. J. Beale, D. B. Gasson, J. M. Keen, C. Pickering, and D. J. Robbins, Surf. Sci. 152/153, 

1051 (1985). 

[9] M. A. Muriel and A. Carballar, IEEE Photon. Technol. Lett. 9, 955 (1997). 

[10] X. Jiang, Y. Zhang, S. Feng, K. Huang, Y. Yi, and J. D. Joannopoulos, Appl. Phys. Lett. 86, 201110 (2006). 

[11] P. A. Snow, E. K. Squire, P. St. J. Russell, and L. T. Canham, J. Appl. Phys. 86, 1781 (1999). 

[12] L. De Stefano, I. Rendina, L. Moretti, S. Tundo, and A. M. Rossi, Appl. Opt. 43, 167 (2004). 

© 2007 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim www.pss-c.com

Optical sensing of chemical compounds using porous silicon 

devices 

In this section are presented some very interesting results about the detection of chemical 

compounds by using porous silicon optical devices. Some porous silicon structures, were 

exposed to several substances of environmental interest: the detection in aqueous and 

humic solutions of a common pesticide and organic compounds. Very large red shifts of 

the reflectivity spectra, due to changes in the average refractive index, have been 

observed.

Introduction to the sensing mechanism 

In this section, are presented optical sensors, based on porous silicon (PSi) technology for 

liquid and vapour detection. Optical sensors for chemical contaminants can adopt a 

spectroscopic method (usually in the IR region where gas absorption spectra have specific 

signatures), or measure the changes in some physical properties (colour, refractive index, 

photoluminescence, fluorescence and so on) due to the interaction between the 

substances under investigation and the sensor itself [1, 2]. From this point of view, PSi is a 

very interesting material due to its large specific surface area (on the order of 500 m 2 /cm 3 ) 

which is a great advantage in gas sensing applications, so that this technology has been 

extensively studied in this field. When the PSi is exposed to vapours, or dip in liquid 

substances, a repeatable and completely reversible change in the reflectivity spectrum is 

observed. In fact, the substitution of air in the pores by the organic molecules causes an 

increasing of the average refractive index of the device, shifting towards longer 

wavelengths its characteristic reflectivity spectrum. 

In a first optical characterization, it was followed the approach of Sailor et al. [3] choosing 

the simple case of a single layer of porous silicon as transducer device. Figure 1 shows a 

typical reflectivity spectrum (before and after chemical infiltration) from a PSi layer under 

white light illumination. 

Reflectance (a.u.) 

1.00 

0.75 

0.50 

0.25 

1.00 

0.75 

0.50 

after infiltration 

0.25 

11000 12000 13000 14000 15000 16000 17000 


Maximum order ( Δm) 

16 

14 

12 

10 

8 

6 

4 

2 

11000 12000 13000 14000 15000 16000 17000 

PSi as etched 


n xyl d=27352 (92) nm 

n unp d=24560 (110) nm 

Figure 1. Reflectivity spectrum of PSi layer. In the inset the m-order peak are plotted as function of the 

wavenumber. 

From the optical point of view, this structure is an optical Fabry-Perot interferometer, so 

that the maxima in the reflectivity spectrum appear at wavelengths λm which satisfy the 

fallowing relationship: 

m = 2 nd / λ 

(1) 

m 

where m is an integer, d is the film thickness and n is the average refractive index of the 

layer [4, 5]. 

If it’s assumed that the refractive index is independent on the wavelength over the 

considered range, the maxima are equally spaced in the wavenumber (1/λm). When the 

maximum order m is plotted as a function of the wavenumber, each point lies on a straight 

22

line which slope is the optical path nd, as it is shown in the inset of Figure 1. When the PSi 

layer is exposed to vapors, or dip in the liquid phase of the same substance, the 

substitution of air into the pores by its molecules causes a fringes shift in wavelength, 

which corresponds to a change in the optical path nd. Since the thickness d is fixed by the 

physical dimension of the PSi matrix, the variation is clearly due to changes in the average 

refractive index. In the case of vapors and gases, the filling liquid phase is due to the 

capillary condensation phenomenon, which is regulated from the Kelvin equation: 

K sat 

BTρ ln( p / p) 

= 2γ 

cos / R 

l lg θ 

(2) 

where ρl is the density of the liquid phase, γlg is the liquid-gas surface tension coefficient at 

room temperature T, R is the radius of the pores, p/psat is the relative vapour pressure into 

the pores, and θ is the contact angle [6]. 

A quantitative model taking into account the optical path increase can be realized by 

applying the Bruggemann effective medium approximation theory. The relation between 

the volume fraction of each medium and its dielectric constant can be written as: 

⎛ ε ⎞ ⎛ ⎞ ⎛ ⎞ 

Si −εp 

εair 

−εp 

εch 

−εp 

( 1 − p) 

⎜ ⎟+ 

( p−V) 

⎜ ⎟+ 

V⎜ 

⎟ = 0 

(3) 

⎜ 2 ⎟ ⎜ 2 ⎟ ⎜ 2 ⎟ 

⎝εSi 

+ εp 

⎠ ⎝εair 

+ εp 

⎠ ⎝εch 

+ εp 

⎠ 

where p, V, εsi, εair, εch and εp are the layer porosity, the layer liquid fraction (LLF), the 

dielectric constants of silicon, air, chemical substance and porous silicon, respectively [7]. 

References 

1. P. A. Snow, E. K. Squire, P. St. J. Russel, and L. T. Canham, Vapor sensing using the optical 

properties of porous silicon Bragg mirrors, J. Appl. Phys. 86, (1999) 1781-1784. 

2. V. S. Y. Lin, K. Motesharei, K. P. S. Dancil, M. J. Sailor, and M. R. Ghadiri, A porous silicon-based 

optical interferometric biosensor, Science 278, (1997) 840-843. 

3. Gao, J.; Gao, T.; Sailor, M.J. Porous-silicon vapor sensor based on laser interferometry. Appl. Phys. 

Lett. 2000, 77, 901-903. 

4. Lin, V. S.-Y.; Motesharei, K.; Dancil, K.-P. S.; Sailor, M. J.; Ghadiri, M. R. A porous silicon-based 

optical interferometric biosensor. Science 1997, 278, 840-843. 

5. Anderson, M. A.; Tinsley-Brown, A.; Allcock, P.; Perkins, E. A.; Snow, P.; Hollings, M.; Smith, R. G.; 

Reeves, C.; Squirrell, D. J.; Nicklin, S.; Cox; T. I. Sensitivity of the optical properties of porous silicon 

layers to the refractive index of liquid in the pores. Phys. Stat. Sol. A 2003, 197, 528-533. 

6. Neimark, V.; Ravikovitch, P. I. Capillary condensation in MMS and pore structure characterization. 

Microporous Mesoporous Mater. 2001, 44-45, 697-707. 

7. Spanier, J. E.; Herman, I. P. Use of hybrid phenomenological and statistical effective-medium 

theories of dielectric functions to model the infrared reflectance of porous SiC films. Phys. Rev. B 

2000, 61, 10437-10450. 

23

phys. stat. sol. (c) 4, No. 6, 1941–1945 (2007) / DOI 10.1002/pssc.200674336 

Quantitative measurements of hydro-alcoholic binary mixtures 

by porous silicon optical microsensors 

Luca De Stefano *1 , Lucia Rotiroti 1 , Ilaria Rea 1 , Luigi Moretti 2 , and Ivo Rendina 1 

1 

Institute for Microelectronics and Microsystems, National Council of Research, Via P. Castellino 111, 

80131 Naples, Italy 

2 

DIMET – “Mediterranea” University of Reggio Calabria, Località Feo di Vito, 89060 Reggio Calabria, 

Italy 

Received 17 March 2006, revised 15 September 2006, accepted 15 November 2006 

Published online 9 May 2007 

PACS 07.07.Df, 81.05.Rm, 81.70.Jb 

In this communication, we present a simple, completely reversible and well performing optical microsensor 

for the quantitative determination in liquid hydro-alcoholic binary mixtures, based on the simple and 

low cost porous silicon (PSi) nanotechnology. In our sensor device, the transducer element is a PSi optical 

microcavity (PSMC) which is constituted by a Fabry-Pérot layer between two distributed Bragg reflectors. 

When the device is exposed to the mixture, the partial substitution of air in the pores changes the average 

refractive index of PSMC, causing a red shift of its reflectivity spectrum, characteristic of the concentration 

of each component. We have characterized, in the 0-100% composition range, three hydroalcoholic 

binary mixtures (water-Ethanol, water- Methanol and water- Isopropanol) which exhibit very 

different behaviour. 

1 Introduction 

© 2007 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim 

Hydro-alcoholic mixtures are largely used in the food industry: for example, the monitoring of the DI 

water-Ethanol compound is of extreme importance since it determines the alcoholic volume content of 

the most common beverages such as wines, beers, liquors and spirits. The standard methods, prescribed 

for the quantitative determination of the alcohol concentration, suffer some drawbacks since they require 

professional laboratories with specialized personnel, and expensive equipment, such as the HPLC, and 

sometimes also the sample pre-treatment. These procedures cannot be applied on line and are useless for 

fast routine analyses. The optical sensors offer a very attractive solution in the realization of analytical 

devices: they could exhibit high sensitivity, fast response, small dimensions, low cost and also the possibility 

of remote sensing. 

PSi based optical sensors have recently gained a well established place among the optical monitoring 

devices due to the peculiar physical properties of this material. PSi is very cheap to produce, fully integrable 

with the standard silicon technology, moreover its high surface area assures a strong and fast interaction 

with liquid or vapor substances. On these bases, several applications of PSi monolayer and 

multilayer structures have been recently considered in biological and chemical sensing [1–3, 9]. A step 

beyond the optical identification of a pure substance is the quantitative determination of the components 

concentrations in a mixture. When the PSi device is exposed to a liquid or gaseous solution, a change in 

the optical reflectivity spectrum can be observed due to the partial substitution of air in the pores by a 

liquid phase, due to the direct infiltration or to the capillary condensation, which increases the average 

* Corresponding author: e-mail: luca.destefano@na.imm.cnr.it, Phone +390816132375, Fax +390816132598 

© 2007 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim

1942 L. De Stefano et al.: Quantitative measurements of hydro-alcoholic binary mixtures 

refractive index. This phenomenon depends on both the refractive index of the liquid phase but also on 

the filling factor, i.e. how it penetrates into the spongy structure of the PSi. The dependence of the result 

on two parameters permits to resolve a binary mixture with only one optical measurement [4]. Our group 

has already published some quantitative results on gaseous and liquid binary mixtures, obtained by using 

PSi based resonant optical sensors. In particular, we have characterized the sensor response on exposure 

to vapors of several volatile compounds (Iso-propanol/Ethanol, Iso-propanol/Methanol, Xylene/Methanol), 

and after the interaction with liquid solutions of pesticides in water and humic acid [5– 

7]. We present here an all-optical PSi based chemical sensor for quantitative determination in liquid 

hydro-alcoholic binary mixtures. 

2 Material and methods 

The optical transducer we used in our experiments is a λ/2 Fabry-Pérot layer enclosed between two 

Bragg reflectors; each one fabricated by alternating seven layers with high (low) and low (high) refractive 

indices (porosities). This PSi multilayer structure is an optical microcavity (PSMC), electrochemically 

etched by a HF-based solution on a highly doped p + -silicon, oriented, 0.01 Ω cm resistivity, 

and 400 µm thick. A current density of 350 mA/cm 2 for 0.6 sec was applied to obtain the high refractive 

index layer, with a porosity of 62 %, while one of 50 mA/cm 2 was applied for 0.8 sec for the low index 

layer, with a porosity of 42 %. The Fabry-Pérot is a low refractive index layer. This structure has a characteristic 

resonance peak at 1110 nm in the middle of a 271 nm stop band. The infiltration of the aqueous 

solution into the spongy structure could be problematic since the PSi surface as etched is highly hydrophobic 

due to the presence of Si-H bonds. In order to assure an efficient interaction it is necessary to 

make this surface hydrophilic by a thermal oxidation process (at 900 °C for 45 minutes). 

White light 

source 

Collimator 

Optical fiber 

Objective 

DBR #1 

DBR #2 

Fig. 1 The experimental optical set-up: the sample was illuminated by a collimated white source; the output signal 

was directed in an optical spectrum analyzer and it was measured over the range 1000-1050 nm with a resolution of 

0.2 nm. 

We studied the resonance peak shift on exposure to the liquid mixtures as a function of the alcohols 

concentration in the de-ionized water (DI). To avoid some lens effect, we placed a drop (20 µL) of the 

solution on the PSi surface and covered it by a thin transparent glass. In Fig. 1 the experimental set-up 

used to measure the reflectivity spectrum is reported. When the liquid penetrated into the PSMC, a 

change in the cavity reflectivity spectrum was observed: the partial substitution of air by liquid in the 


OSA 

d=λ/2nl 

X7 

Low P 

High P


pores of each layer increases the average refractive index of the microcavity and, consequentially, generates 

a red-shift of the reflectivity spectrum. Measurements are stored when equilibrium conditions are 

reached, i.e. when the signal does not change any more in time. 

Fig. 2 The optical spectrum of the PSMC before and after the thermal oxidation and the infrared spectrum after 

thermal treatment. 

3 Experimental results 

The optical and infrared spectra of the device before and after the oxidation are reported in Fig. 2. The 

microcavity has a FWHM of 12 nm and the thermal oxidation causes a blue shift of the reflectivity spectrum 

of 110 nm. The optical quality of the PSi response is preserved after the thermal treatment. In the 

infrared spectrum is well evident the presence of the characteristic peaks of Si-O-Si bonds at about 1000- 

1100 cm –1 and the absence of the Si-H bonds at 2100 cm –1 . 

In a recent paper 5 , we have already used a very simple method, common in the analytical chemistry, for 

the quantitative determinations of binary mixtures: the External Standard Method. The calibration curves 

can be obtained by measuring the physical parameter under monitoring, such as the infrared absorption at 

∆λ (nm) 

Reflectivity (u.a) 

14 

12 

10 

8 

6 

4 

2 

0 

1.0 

0.5 

PSMC as etched 

PSMC after thermal oxidation 

Isopropanol-DI 

EtOH-DI 

∆λ= 110 nm 

0.0 

600 800 1000 1200 1400 1600 


0 20 40 60 80 100 

Alcohol concentration in DI (% V/V) 

0 20 40 60 80 100 

Fig. 3 The calibration curves in case of three binary mixtures (Iso-propanol/DI, Ethanol/DI and Methanol/DI). 

www.pss-c.com © 2007 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim 

Reflectance (a.u.) 

∆λ (nm) 

4.5 

4.0 

3.5 

3.0 

2.5 

2.0 

1.5 

1.0 

0.5 

0.0 

-0.5 

140 

120 

100 

80 

60 

40 

20 

0 

2400 2200 2000 1800 1600 1400 1200 1000 800 

MeOH-DI 

polinomial fit 



SiO 2

n 

1944 L. De Stefano et al.: Quantitative measurements of hydro-alcoholic binary mixtures 

a fixed wavelength or the optical intensity, while the content of one component varies between 0 and 

100%, volume by volume (V/V). Also in the case under study, the calibration curves have been obtained 

by plotting the peak shifts measured as a function of the volume percentage of one component in the 

binary mixtures. 

The results are shown in Fig. 3. The errors on the experimental points are the standard deviations on 

three to five measurements, using the same chip and the same hydro-alcoholic solution. There is a well 

evident dissimilarity respect to the linear behaviour of the optical response obtained on exposure to the 

vapours of the volatile binary compounds. Of course, even if the calibration curves are not linear, the 

unknown concentrations of the characterized mixtures can be still quantitatively determined with a single 

measurement of the peak shift: the volume concentrations of both the components can be picked out 

from the interpolation of the experimental data. Moreover, while the Isopropanol and Ethanol-DI mixtures 

shifts follow a slow parabolic behaviour, the Methanol-DI mixture exhibits a very different response. 

The main difference is that a 50 % solution of Methanol in DI causes a larger shift than both the 

pure substances. An explanation to the shape of the ∆λ curves can be found in the behaviour of the mixtures 

refractive indexes as a function of the alcohol concentration [8], shown in Fig. 4. In all the cases, 

the peak shift strictly follows the behaviour of the refractive index binary mixture. The anomalous shape 

of the Methanol-DI mixture can be ascribed to the higher dipole moment of its constituent: due to a 

strong dipolar interaction this hydro-alcoholic mixture could exhibit a denser optical phase respect to the 

two pure substances [8]. 

The time-resolved measurements have shown that sensor dynamic is of order of few seconds, also depending 

on geometrical characteristic of the experimental set-up [3]. Moreover, the phenomenon is completely 

reversible and reproducible, so that the transducer PSi element can be reusable several times. 

As a practical application of this study, we have tried to use the PSMC to determine the alcohol content 

of a red wine (Aglianico del Taburno, DOC). To this aim, we have diluted the wine, with a declared 

alcohol content of 12% V/V, in DI water and we have measured the calibration curve up to this value. 

From the graph in Fig. 3, we expected a linear behavior, due to the small range of Ethanol concentration 

in the wine. In Fig. 5, we have reported the wine calibration curve together with the first order approximation 

of the Ethanol-DI calibration curve of Fig. 3. The two curves are both linear but they have different 

slopes: in particular, the Ethanol-DI curve underestimates the wine percentage content. This is also a 

1.38 

1.37 

1.36 

1.35 

1.34 

1.33 

IPA-DI water refractive index 

EtOH-DI water refractive index 

0 20 40 60 80 100 


Fig. 4 The mixtures refractive indexes as a function of the alcohol concentration in DI. 

n 

0 20 40 60 80 100 

somewhat expected result: in fact the wine is a very complex mixture where Ethanol and water are the 

main constituents but many other volatile organic compounds are present. Nevertheless, we believe that 


1.344 

1.342 

1.340 

1.338 

1.336 

1.334 

1.332 

1.330 

1.328 

Methanol concentration in DI (% V/V)


Fig. 5 The calibration curve in case of wine-DI water solution. 

the measurement technique could be improved, perhaps by time-resolved procedures, in order to give 

more precise results. 

4 Conclusion 

We have presented a simple and well performing optical microsensor for quantitative determination in 

liquid mixtures, based on the simple and low cost porous silicon nanotechnology. The device is a microcavity 

structure that is directly realized on the silicon chip by electrochemical etching. The fabrication 

process is compatible with silicon microelectronic and micro machining technologies, so that the sensors 

could be monolithically integrated with microelectronic circuits, or be part of a micro-optical-system. In 

terms of the measurement resolution, the resonant cavity enhanced operation of the proposed sensor 

offers very good performances: the interaction between the porous silicon microcavity and the external 

mixtures detunes the optical microcavity inducing large shifts of its resonant peak wavelength. 

References 

∆λ (nm) 

5 

4 

3 

2 

1 

0 

Wine-DI calibration curve 

First order approximation of 

the Ethanol-DI calibration curve 

0 2 4 6 8 10 12 14 

Wine concentration in DI (% V/V) 

[1] P. A. Snow, E. K. Squire, P. St. J. Russel, and L. T. Canaham, J. Appl. Phys. 86, 1781 (1999). 

[2] V. S. Y. Lin, K. Motesharei, K. P. S. Dancil, M. J. Sailor, and M. R. Ghadiri, Science 278, 840 (1997). 

[3] L. De Stefano, I. Rendina, L. Moretti, A.M. Rossi, Sens. Actuators B 100, 168 (2004). 

[4] C.K. Ho, M.T. Itamura, M. Kelley, and R.G. Hughes, Review of Chemical Sensors for In-situ Monitoring of 

Volatile Contaminants, Sandia Report SAND2001-0643, Sandia National Laboratories, 2001. 

[5] L. De Stefano, I. Rendina, L. Moretti, and A.M. Rossi, phys. stat. sol. (a) 201, 1011 (2004). 

[6] L. De Stefano, L. Moretti, I. Rendina, and L. Rotiroti, Sens. Actuators B 111/112, 522 (2005). 

[7] L. Rotiroti, L. De Stefano, L. Moretti, A. Piccolo, I. Rendina, and A. M. Rossi, Biosens. Bioelectron. 20, 2136 

(2005). 

[8] Handbook of Chemistry and Physics, 68th ed. (CRC Press, Boca Raton, 1987/1988). 

[9] L. De Stefano, I. Rea, I. Rendina, L. Rotiroti, M. Rossi, and S. D'Auria phys. stat. sol. (a) 203, 886 (2006). 


Chapter 2 

Chemical and Biological 

functionalization 

of Porous Silicon surface

Introduction: Porous Silicon surface chemical functionalization 

Biosensors are nowadays technological hot topics due to the possible applications in fields 

of social interest such as medical diagnostic and health care, monitoring of environmental 

pollutants, homeland and defence security [1]. 

Reagentless optical biosensors are monitoring devices which can detect a target analyte in 

a heterogeneous solution without the addition of anything than the sample. In the fields of 

genomics and proteomics this is a straightforward advantage since it allows real-time 

readouts and, thus, very high throughoutputs analysis. A label free optical biosensor can 

be realized by integrating the biological probe with a signalling material which directly 

transduces the molecular recognition event into an optical signal. 

PSi is also very used as photonic material due to the possibility of fabricating high quality 

optical structures, either as single layers, like Fabry-Perot interferometers [2], or 

multilayers, such as Bragg [3] or rugate filters [4]. A key feature for a recognition 

transducer is a large surface area: PSi has a porous structure with a specific area up to 

200 – 500 m 2 cm -3 , so that it can be very sensitive to the presence of biochemical species 

which penetrate inside the pores. Unfortunately, the surface of the “as etched” PSi is 

highly hydrophobic due to the Si-H bonds so that aqueous solution can not infiltrate the 

sponge like matrix. A proper passivation process must be applied to stabilise the surface 

and to covalently link the bioprobe [5]. 

• Hydrosilylation to produce Si-C bonds 

Interesting results on the chemical derivatization of PSi were reported before 1998, but the 

problem was the low treatment efficiency and, thus, the remaining instability. Especially, 

the groups of Mike Sailor (Department of Chemistry and Biochemistry University of 

California, San Diego), Jean-Noel Chazalviel (Palaiseau Polytechnical Institute, France), 

and Jillian Buriak (Department of Chemistry, University of Alberta, and the National 

Institute for Nanotechnology) studied different possible treatments for the PSi surface 

modifications with Si–C bonds. 

Their work finally led to the derivatized PSi surface, in which the Si–H bonds were 

replaced with Si–C bonds using hydrosilylation of alkenes or alkynes on the PSi surface. 

The group of Buriak introduced three different approaches to obtain the chemically 

derivatized PSi surface: Lewis acid mediated hydrosilylation [6,7] white light-promoted 

hydrosilylation [8,9] and cathodic electrografting [10]. 

Carbon directly bonded to silicon yields a very stable surface species. First recognized by 

Chidsey and coworkers [11], Si–C bonded species possess greater kinetic stability relative 

to Si–O due to the low electronegativity of carbon. Silicon can readily form 5- and 6- 

coordinate intermediates, and an electronegative element such as oxygen enhances the 

tendency of silicon to be attacked by nucleophiles. Si–C bonds are usually formed on 

hydride-terminated porous Si surfaces by hydrosilylation (Figure 1). Hydrosilylation 

involves reaction of an alkene (usually terminal) or alkyne with a Si–H bond. On porous Si, 

the reaction can be thermal [12], photochemical [13], or Lewis acid catalyzed [14]. 

Figure 1: mechanism of PSi functionalization by nucleophilic substitution. 

29

Thermal hydrosilylation provides a means to place a wide variety of organic functional 

groups on a crystalline Si or porous Si surface. The main requirement of the reaction is 

that the silicon surface contains Si–H species so they can react with a terminal alkene on 

the organic fragment. Thus it is important to use freshly etched porous Si and to exclude 

oxygen and water from the reaction mixture. 

• Chemical grafting of Si-C bonds 

As an alternative to hydrosilylation, covalently attached layers can be formed on porous Si 

surfaces using Grignard (figure 2) and alkyl- or aryllithium reagents [15]. Because of their 

ease of application and dramatic improvements in stability, hydrosilylation grafting of alkyl 

halides are useful reactions for the preparation of biointerfaces. 

Figure 2: mechanism of PSi functionalization by Grignard reactive. 

It is important to note that porous Si modification reactions do not provide 100% surface 

coverage; they merely decorate the surface with the functional group. Thus infrared 

spectra show a large amount of surface Si–H groups remaining after hydrosilylation or 

grafting reaction. 

It is still an open question as to why the modification reactions impart such stability to 

the material. The stability probably derives from a combination of two factors: the Si–C 

bond is kinetically inert due to the low electro negativity of carbon and the attached organic 

species (typically a hydrocarbon chain 8 or more carbons long) is sufficiently hydrophobic 

that aqueous nucleophiles are excluded from the vicinity of their Si atom target. Reaction 

of porous Si with gas phase acetylene generates highly carbonized porous Si that is 

possibly the most stable form of Si–C modified porous Si [16]. This material is referred to 

as thermally carbonized PSi. It has been extensively investigated by Salonen at 

Department of Physics of University of Turku, (Turku, Finland) and coworkers, with many 

publications of relevance to drug loading and delivery [17]. 

• Conjugation of bio molecules to modified PSi 

Carbon grafting stabilizes porous Si against dissolution in aqueous media, but the surface 

must still avoid the non-specific binding of proteins and other biological species. 

The oxides of porous Si are easy to functionalize using conventional silanol chemistries. 

When small pores are present (as with p-type samples), trialkoxysilanes [(RO)3Si–R′] can 

be employed as surface linkers [18]. 

Whereas Si–C chemistries are robust and versatile, chemistries involving Si–O bonds 

represent an attractive alternative for two reasons. 

First, the timescale in which highly porous SiO2 is stable in aqueous media is consistent 

with many short-term biomolecular recognitions applications: typically 20 min to a few 

hours. Second, a porous SiO2 sample that contains no additional stabilizing chemistries is 

less likely to produce toxic or antigenic side effects. 

30

• Oxidation of porous silicon 

With its high surface area, porous silicon is particularly susceptible to air or water 

oxidation. The simplest way to stabilize PSi is obviously a partial oxidation. A few hours 

even at quite mild conditions, around 300°C, cause mainly the so called back-bond 

oxidation of PSi [19]. The oxygen atoms selectively attack the back-bonds of the surface Si 

atoms instead of replacing hydrogen atoms. The oxygen bridges formed between the 

surface Si atoms and the second atomic Si layer expand the local atomic structure by 

30%, causing a slight decrease in the pore diameter. In addition to the increased stability, 

the oxidation at 300°C also changes the surface from hydrophobic to hydrophilic, which is 

important in many biological applications under physiological conditions. Increased 

temperature also increases the extent of oxidation leading to completely oxidized PSi at 

around 900–1000°C [20]. Due to the structural expansion, the pore diameter and porosity 

are dependent on the extent of oxidation, and a drastic drop in the specific surface area 

has been observed in PSi oxidized around 600°C [21]. 

Once oxidized, Si-O bonds are easy to prepare on porous silicon by oxidation, and a 

variety of chemical or electrochemical oxidants can be used. There are also many other 

techniques to oxidize PSi, such as anodic oxidation, [22,23] photo-oxidation, [24-26] and 

chemical oxidation. Like the thermal oxidation, the chemical oxidation is quite simple, in 

which PSi is oxidized with inorganic or organic agents. [27-31] 

Milder chemical oxidants, such as dimethyl sulfoxide (DMSO), or piranha solutions 

(H2SO4/H2O2 7:3) can also be used for this reaction. Mild oxidants are sometimes 

preferred because they can improve the mechanical stability of highly porous Si films, 

which are typically quite fragile. The mechanical instability of PSi is directly related to the 

strain that is induced in the film as it is produced in the electrochemical etching process, 

and the volume expansion that accompanies thermal oxidation can also introduce strain. 

Electrochemical oxidation, in which a porous Si sample is anodized in the presence of a 

mineral acid such as H2SO4, yields a fairly stable oxide. Mild chemical oxidants 

presumably attack PSi preferentially at Si–Si bonds that are the most strained, and hence 

most reactive. 

Aqueous solutions of bases such as KOH can also be used to enlarge the pores after 

etching [32]. Oxidation imparts hydrophilicity to the porous structure, enabling the binding 

and adsorption of hydrophilic biomolecules within the pores. 

References 

1. Optical Biosensors, Eds. F.S. Ligler and C.A. Rowe Taitt, Elsevier, Amsterdam, The Netherlands, 

2004. 

2. Dancil, K. -P. S.; Greiner, D. P.; Sailor, M. J., A Porous Silicon Optical Biosensor: Detection of 

Reversible Binding of IgG to a Protein A-Modified Surface, J. Am. Chem. Soc. 1999, 121, 7925- 

7930. 

3. Snow, P. A.; Squire, E. K.; Russel, P. St. J.; Canham, L. T., Vapor sensing using the optical 

properties of porous silicon Bragg mirrors, J. Appl. Phys. 1999, 86, 1781-1784. 

4. Lorenzo, E.; Oton, C. J.; Capuj, N. E.; Ghulinyan, M.; Navarro-Urrios, D.; Gaburro, Z.; Pavesi, L., 

Porous silicon-based rugate filters, Appl. Opt. 2005, 44, 5415-5421. 

5. Ouyang, H.; Chrtistophersen, M.; Viard, R.; Miller, B. L.; Fauchet, P. M., Macroporous silicon 

microcavities for macromolecule detection, Adv. Funct. Mater. 2005, 15, 1851-1859. 

6. Buriak JM, Allen MJ. 1998. Lewis acid mediated functionalization of porous silicon with substituted 

alkenes and alkynes. J Am Chem Soc 120: 1339–1340. 

7. Buriak JM, Sterward MP, Geders TW, Allen MJ, Choi HC, Smith J, Raftery D, Canham LT. 1999. 

Lewis acid mediated hydrosilylation porous silicon surfaces. J AmChem Soc 121:11491–11502. 

8. Stewart MP, Buriak JM. 1998. Photopatterned hydrosilylation on porous silicon. Angew Chem Int Ed 

37:3257–3260. 

9. Stewart MP, Buriak JM. 2001. Exciton-mediated hydrosilylation on photoluminescent nanocrystalline 

silicon. J Am Chem Soc 123:7821–7830. 

31

10. Robins EG, Stewart MP, Buriak JM. 1999. Anodic and cathodic electrografting of alkynes on porous 

silicon. Chem Comm 24:2479–2480. 

11. M.R. Linford, C.E.D. Chidsey, Alkyl monolayers covalently attached to silicon surfaces, J. Am. Chem. 

Soc. 115 (1993) 12631–12632. 

12. R. Boukherroub, J.T.C. Wojtyk, D.D.M. Wayner, D.J. Lockwood, Thermal hydrosilylation of 

undecylenic acid with porous silicon, J. Electrochem. Soc. 149 (2002) 59–63. 

13. M.P. Stewart, J.M. Buriak, Photopatterned hydrosilylation on porous silicon, Angew. Chem. Int. Ed. 

Engl. 37 (1998) 3257–3260. 

14. J.M. Buriak, Organometallic chemistry on silicon and germanium surfaces, Chem. Rev. 102 (2002) 

1272–1308. 

15. J. Terry, M.R. Linford, C. Wigren, R.Y. Cao, P. Pianetta, C.E.D. Chidsey, Alkylterminated Si(111) 

surfaces: a high-resolution, core level photoelectron spectroscopy study, J. Appl. Phys. 85 (1999) 

213–221. 

16. M. Bjorkqvist, J. Salonen, E. Laine, L. Niinisto, Comparison of stabilizing treatments on porous 

silicon for sensor applications, Phys. Status Solidi A—Appl. Res. 197 (2003) 374–377. 

17. A.M. Kaukonen, L. Laitinen, J. Salonen, et al., Enhanced in vitro permeation of furosemide loaded 

into thermally carbonized mesoporous silicon (TCPSi) microparticles, Eur. J. Pharm. Biopharm. 66 

(2007) 348–356 

18. A. Janshoff, K.P.S. Dancil, C. Steinem, et al., Macroporous p-type silicon Fabry– Perot layers. 

Fabrication, characterization, and applications in biosensing, J. Am. Chem. Soc. 120 (1998) 12108– 

12116. 

19. Salonen J, Lehto VP, Laine E. 1997. Thermal oxidation of free-standing porous silicon films. Appl 

Phys Lett 70:637–639. 

20. Canham LT. 1997. Chemical composition of intentionally oxidized porous silicon. In: Canham LT, 

editor. Properties of porous silicon. EMIS Datareview series No 18. London: INSPEC. pp 158– 162. 

21. Halimaoui A. 1995. Porous silicon: Material processing, properties and applications. In: Vial JC, 

Derrien J, editors. Porous silicon science and technology. France: Springer-Verlag. pp 33–52. 

22. Herino R. 1995. Luminescence of porous silicon after electrochemical oxidation. In: Vial JC, Derrien 

J, editors. Porous silicon science and technology. France: Springer-Verlag. pp 53–66. 

23. Mulloni V, Pavesi L. 2000. Electrochemically oxidized porous silicon microcavities. Mat Sci Eng B 

69–70:59–65. 

24. Tischler MA, Collin RT, Stathis JH, Tsang JC. 1992. Luminescence degradation in porous silicon. 

Appl Phys Lett 60:639–641. 

25. Salonen J, Lehto VP, Laine E. 1999. Photo-oxidation studies of porous silicon using 

microcalorimetric method. J Appl Phys 86:5888–5893. 

26. Gole JL, Dixon DA. 1998. Evidence for oxide formation of the single and multiphoton excitation of a 

porous silicon surface or silicon ‘‘nanoparticles’’. J Appl Phys 83:5985–5991. 

27. Song JH, Sailor MJ. 1999. Chemical modification of crystalline porous silicon surfaces. Comments 

Inorg Chem 1-3:69–84. 

28. Nakajima A, Itakura T, Watanabe S, Nakayama N. 1992. Photoluminescence of porous Si, oxidized 

and then deoxidized chemically. Appl Phys Lett 61:46–48. 

29. Rao BVRM, Basu PK, Biswas JC, Lahiri SK, Ghosh S, Bose DN. 1996. Large enhancement of 

photoluminescence from porous silicon films by post-anodization treatment in boiling hydrogen 

peroxide. Solid State Comm 97:417–418. 

30. Yin F, Li XP, Zhang ZZ, Xiao XR. 1997. Investigations on the surface reactivity of luminescent 

porous silicon. Appl Surf Sci 119:310–312. 

31. Salonen J, Lehto VP, Laine E. 1997. The room temperature oxidation of porous silicon. Appl Surf Sci 

120:191–198. 

32. L. A. De Louise and B. L. Miller, Mat. Res. Soc. Symp. Proc. 782, A5.3.1 (2004). 

32

DNA Optical Detection Based on Porous Silicon Technology 

This section is aimed at the optimization of the functionalization process investigating also 

the role of thickness and porosity of the PSi chip, the time exposure to ultraviolet (UV) light 

and also the concentration of the ssDNA solution. 

In collaboration with Prof. P. Arcari of University of Naples “Federico II”, Italy

Sensors 2007, 7, 214-221 

Full Research Paper 

sensors 

ISSN 1424-8220 

© 2007 by MDPI 

www.mdpi.org/sensors 

DNA Optical Detection Based on Porous Silicon Technology: 

from Biosensors to Biochips 

Luca De Stefano 1, *, Paolo Arcari 2 , Annalisa Lamberti 2 , Carmen Sanges 2 , Lucia Rotiroti 1,3 , 

Ilaria Rea 1, 4 and Ivo Rendina 1 

1 

Institute for Microelectronics and Microsystems – Unit of Naples – National Council of Research, Via 

P. Castellino 111, 80131 Napoli, Italy (E-mail: luca.destefano@na.imm.cnr.it) 

2 

Department of Biochemistry and Medical Biotechnologies, University of Naples “Federico II”, Via S. 

Pansini 5, 80131 Napoli, Italy. 

3 

Department of Organic Chemistry and Biochemistry, University of Naples “Federico II”, Via Cinthia, 

80128 Napoli, Italy 

4 

Department of Physical Sciences, University of Naples “Federico II”, Via Cinthia, 80126 Naples, Italy 

* Author to whom correspondence should be addressed. E-mail: luca.destefano@na.imm.cnr.it 

Received: 15 February 2007 / Accepted: 26 February 2007 / Published: 28 February 2007 

Abstract: A photochemical functionalization process which passivates the porous silicon 

surface of optical biosensors has been optimized as a function of the thickness and the 

porosity of the devices. The surface modification has been characterized by contact angle 

measurements. Fluorescence measurements have been used to investigate the stability of the 

DNA single strands bound to the nanostructured material. A dose-response curve for an 

optical label-free biosensor in the 6-80 μM range has been realized. 

Keywords: Optical Biosensors, Porous Silicon, Biochip. 

1. Introduction 

Biosensors are nowadays technological hot topics due to the possible applications in social interest 

fields such as medical diagnostic and health care, monitoring of environmental pollutants, home and 

defence security [1]. Besides the signal generated by the sensing device, the biosensor is constituted by 

the molecular recognition element and the transducer material. The molecular recognition element can 

be a biological molecule, such as DNA single strand, proteins, enzymes, or a biological system, such as

Sensors 2007, 6 

membrane, cell, and tissues: in this way, the sensing mechanism takes advantage of the natural 

sensitivity and specificity of the biomolecular interactions. Optical transduction is more and more used 

since photonic devices could be small, lightweight and thus portable due to the integrability of all optical 

components. Furthermore, optical devices do not require electric contacts. Fluorescence is by far the 

most used optical signalling method but a wide-use sensor can not be limited by the labelling of the 

probe nor the analyte, since this step is not always possible [2]. Reagentless optical biosensors are 

monitoring devices which can detect a target analyte in a heterogeneous solution without the addition of 

anything than the sample. In the fields of genomics and proteomics this is a straightforward advantage 

since it allows real-time readouts and, thus, very high throughoutputs analysis. A label free optical 

biosensor can be realized by integrating the biological probe with a signalling material which directly 

transduces the molecular recognition event into an optical signal. Recently, lot of theoretical and 

experimental work, concerning the worth noting properties of nanostructured porous silicon (PSi) in 

chemical and biological sensing, has been reported, showing that, due to its morphological and physical 

properties, PSi is a very versatile sensing platform [3, 4]. PSi is an available, low cost material, 

completely compatible with VLSI and micromachining technologies, so that it could usefully be 

employed in the fabrication of micro-opto-electric-mechanical system and smart sensors. PSi is also very 

used as photonic material due to the possibility of fabricating high quality optical structures, either as 

single layers, like Fabry-Perot interferometers [5], or multilayers, such as Bragg [6] or rugate filters [7]. 

A key feature for a recognition transducer is a large surface area: PSi has a porous structure with a 

specific area up to 200 – 500 m 2 cm -3 , so that it can be very sensitive to the presence of biochemical 

species which penetrate inside the pores. Unfortunately, the surface of the “as etched” PSi is highly 

hydrophobic so that aqueous solution can not infiltrate the sponge like matrix. A proper passivation 

process must be applied to stabilise the surface and to covalently link the bioprobe [8]. The 

development of PSi based biosensor arrays critically depends on the surface functionalization process 

and how it is compatible with the microfabrication technologies. In a recently published article, we have 

exploited a photochemical functionalization process of the PSi surface to covalently bind DNA single 

strands (ssDNA) and we have also demonstrate that the device works as an all optical biosensor for 

ssDNA-cDNA interactions [9]. In the present work, we have optimised the functionalization process by 

investigating the role of thickness and porosity of the PSi chip, the time exposure to ultraviolet (UV) 

light and also the concentration of the ssDNA solution. Fluorescent measurements have been used to 

test the chip stability against the washing in aqueous solutions. 

2. Experimental Section 

In this study we used as optical transducer a porous silicon layer of fixed thickness and porosity: 

from an optical point of view, this structure acts as a Fabry-Perot interferometer. To study the influence 

of porous silicon physical parameters, we have fabricated several layers, obtained by electrochemical 

etch in a HF/EtOH (3:7) solution, at room temperature and dark light. Highly doped p + -silicon, 

oriented, 0.005 Ω cm resistivity, 400 µm thick was used. The PSi samples were characterised by 

variable angle spectroscopic ellipsometry [10]. Etching times and anodic current densities have been 

reported in Table 1. 

215


Table 1. PSi layers physical characteristics and fabrication parameters. 

Current Density 

(mA/cm 2 ) 

Porosity 

(%) 

Etch Rate 

(nm/s) 

50 60 193 

125 70 247 

150 80 137 

Thickness Etch Time 

(μm) (s) 

2.0 10.4 

4.0 20.7 

6.0 31.1 

2.0 8.1 

4.0 16.2 

6.0 24.3 

2.0 14.6 

4.0 29.2 

6.0 43.8 

The photo-activated chemical modification of PSi surface was based on the UV exposure of a 

solution of alkenes which bring some carboxylic acid groups. The PSi chip has been pre-cleaned in an 

ultrasonic acetone bath then washed in deionized water. After dried in N2 stream, it has been 

immediately covered with 10% N-hydroxysuccinimide ester (UANHS) solution in CH2Cl2. The UANHS 

was house synthesised as described in literature [11]. This treatment results in covalent attachment of 

UANHS to the porous silicon surface. The chip was then washed in dichloromethane in an ultrasonic 

bath for 10 min to remove any adsorbed alkene from the surface. The carboxyl-terminated monolayer 

covering the PSi surface works as a reactive substrate for the chemistry of the subsequent attachment of 

the DNA sequences. DNA single strands in a HEPES solution 10mM (pH=7.5) have been incubated 

overnight. 

FT-IR spectroscopy (Thermo - Nicholet NEXUS) has been used to verify the efficiency of the 

reaction. After the chemical functionalization we have also quantitatively measured the efficiency of the 

binding between the DNA and the porous silicon surface using a fluorescent DNA probe 

(5'GGACTTGCCCGAATCTACGTGTCCA3', Primm) labelled with a proper chromophore group 

(Fluorescein CY3.5, the absorption peak is at 581 nm and the emission is at 596 nm). Fluorescence 

images were recorded by a Leica Z16 APO fluorescence macroscopy system. The fluorescent chips 

have been dialyzed overnight, first in water and then in a HEPES solution, at room temperature to 

assess the binding between the bioprobe and the PSi surface. 

Contact angle measurements have been performed by using a KSV Instruments LTD CAM 200 

Optical Contact Angle Meter. 

The reflectivity measurements have been performed by a very simple experimental set-up: a tungsten 

lamp (400 nm < λ < 1800 nm) illuminates, through an optical fiber and a collimator, the sensor and the 

reflected beam is collected by an objective, coupled into a multimode fiber, and then directed in an 

optical spectrum analyser (Ando, AQ6315A). The reflectivity spectra have been measured with a 

resolution of 0.5 nm. 

216


3. Results and Discussion 

Infrared spectroscopy is a powerful tool in surfaces characterisation: it is fast, accurate and could 

also performe quantitative determinations. We have investigated all the PSi monolayers before and after 

the photochemical passivation process since we were worried about the functionalization efficiency of 

thicker samples. A thicker layer can adsorb more bioprobes than a thinner one but the UV exposure 

could also be less effective, due to light absorption by PSi at those frequencies. Our measurements 

demonstrate that up to 6 μm thick PSi monolayer, a complete passivation and functionalization of the 

surface can be obtained by exposing for a sufficiently long time the sample: in Figure 1A are shown the 

FT-IR spectra of the PSi monolayer as etched and after three different times of exposure to UV light. 

% Reflectance 

140 

120 

100 

80 

60 

40 

20 

0 

C-H 2 

6 μm 80% porosity 

after 4'20" exposure to UANHS 

after 8'40" exposure to UANHS 

after 13' exposure to UANHS 

Si-H 

O-C=O 

Succinimmide 

C-N 

C=O 

3000 2500 2000 1500 1000 


A 

Peaks Area (%) 

100 

80 

60 

40 

20 

0 

Si-C 

20 

802300 

2200 2100 2000 

900 850 800 750 700 

Figure 1. A) FT-IR spectra of the PSi monolayer as etched and after three different times of exposure 

to UV light. B) Particulars of the Si-H bond and Si-C bond peaks at 2100 and 880 cm -1 , respectively. C) 

Peaks area as function of the exposure time: the reaction yield increases monotonically with the 

exposure time. 

% Reflectance 

100 

80 

60 

40 

60 

40 

20 

Si-C 

0 2 4 6 8 10 12 14 

C 

UV exposure tim e (min) 

SiH 

SiC 

Si-H 

W avenumber (cm -1 ) 

B 

217


The characteristic peaks of the Si-H bonds (at 2100 cm -1 ) progressively disappear while the amide I 

band (at 1634 cm -1 ) and the Si-C peak (at 880 cm -1 ) become more and more evident. Figure 1B is an 

enlargement of the characteristic peaks of these resonances. The reaction yield increases quite 

monotonically with the exposure time, as it can be seen in Figure 1C. Thinner samples can be 

functionalised in faster exposure times, but the overall volume available to adsorb bioprobes is 

drastically reduced. 

A B 

Figure 2. Characterisation of the PSi surface by contact angles measurements. A) The PSi as etched 

shows a clearly hydrophobic behaviour. B) After the passivation process the PSi surface is hydrophilic 

thus allowing the penetration of biological solution. 

The surface passivation is also well evident by contact angle measurements: the hydrogenated PSi 

surface is highly hydrophobic, due to the presence of the Si-H bonds, so that a drop of deionised water 

exploits a high surface tension which prevents its diffusion in the pores. This behaviour is well depicted 

in Figure 2A: the contact angle is 137±1 degree, averaging the left and right values of three different 

drops having the same volumes, and the estimated surface tension is 179 mN/m. After the photoreaction 

with the UANHS, the Si-H bonds are replaced by the Si-C bonds and the PSi surface shows a 

hydrophilic behaviour, as shown in Figure 2B: the drop spreads on the surface and the contact angle 

drastically decreases to 43±3 degree and also the surface tension is lowered to 67 mN/m. In this 

condition, an aqueous solution can be infiltrated in the nanostructured layer. 

To test the stability of the covalent bonding between the organic linker layers, which homogeneously 

cover the PSi surface, and the biological probes we have used a fluorescent DNA single strand as an 

optical tracer. After the chemical bonding of the labelled ssDNA, the chip was observed by the 

fluorescence macroscopy system. Under the light of the 100W high-pressure mercury source, we have 

found a high and homogeneous fluorescence on the whole chip surface which still remains bright even 

after two overnight dialysis washings in a HEPES solution and in deionised water, as it can be seen in 

Figures 3 A, B, and C. We have also studied the yield of the chemical functionalization by spotting 

different concentrations of the fluorescent ssDNA and measuring the fluorescence intensities of the 

images before and after the washings. The results reported in Figure 4 confirm the qualitative findings of 

Figure 3: the fluorescent intensities decrease but remain of the same order of magnitude. From this 

218


graph we can also estimate the concentration of the DNA probe which saturates the binding sites 

available. 

The PSi optical biosensors measure the change in the average refractive index of the device: when a 

bioconjugation event takes place, the refractive index of the molecular complex changes and the 

interference pattern on output is thus modified. The label free optical monitoring of the ssDNA-cDNA 

hybridization is simply the comparison between the optical spectra of the porous silicon layer after the 

UANHS and probe immobilization on the chip surface and after its hybridization with the cDNA. Each 

step of the chip preparation increases the optical path in the reflectivity spectrum recorded, due to the 

substitution of the air into the pores by the organic and biological compounds. The interaction of the 

ssDNA with its complementary sequence has been detected as a fringes shift in the wavelengths, which 

corresponds to a change in the optical path. Since the thickness d is fixed by the physical dimension of 

the PSi matrix, the variation is clearly due to changes in the average refractive index. 

A B C 

Figure 3: A) Fluorescence of the chip surface after the binding of the labelled ssDNA; B) after the 

overnight dialysis in HEPES solution; C) after the overnight dialysis in deionised water. 

In Figure 5A the reflectivity spectra of the PSi layer for different cDNA concentration are reported, 

while in Figure 5B a dose-response curve is reported. A control measurement has been made using a 

ncDNA sequence: a very small shift (less than 2 nm) has been recorded in the reflectivity spectrum 

respect to the one obtained after the probe linking. 

Intensity (a.u.) 

225 

200 

175 

150 

125 

Chip as spotted with labelled DNA 

Chip after HEPES washing 

Chip after water washing 

75 150 225 300 375 450 525 600 

DNA Concentration (μM) 

Figure 4. Fluorescence intensities of the chip surface after the binding of the labelled ssDNA and the 

two overnight dialysis as a function of the cDNA concentration. 

219


The sensor response has been fitted by a monoexponential growth model, y=A-Be -kx , where A is the 

offset, B the amplitude, k the rate and y’=Bk the limiting sensitivity, i.e. the sensitivity in the limit of 

zero ligand concentration. In this case, we obtained for this parameter the value of 1.1 (0.1) nm/μM, 

which corresponds to a limit of detection of 90 nM for a system able to detect a wavelength shift of 0.1 

nm. 


1.4 

1.2 

1.0 

0.8 

0.6 

0.4 

0.2 

0.0 

1150 1200 1250 1300 1350 

Figure 5. A) Fringes shifts due to ssDNA-cDNA interaction. B) Dose-response curve as a function of 

the cDNA concentration. 

In conclusions, we have optimized the PSi surface functionalization process by investigating the role 

of chip thickness and porosity, on exposure to ultraviolet (UV) light for different times interval and also 

for different concentration of the ssDNA probe. Fluorescent measurements have confirmed the chip 

stability against the washing in aqueous solutions. These results will be very useful in the design and 

realisation of a PSi based label free optical biochip. 

Acknowledgements 

Authors wish to acknowledge dr. Andrea M. Rossi of National Institute of Metrological Research, 

Turin, Italy, for helpful discussions and Prof. N. Scaramuzza and dr. M. Giocondo of University of 

Calabria, Cosenza, Italy, for contact angle measurements. This work is partially supported by the FIRB 

Italian project RBLA033WJX_005. 

References and Notes 

A 

ss-DNA 





1. Optical Biosensors, Eds. F.S. Ligler and C.A. Rowe Taitt, Elsevier, Amsterdam, The 

Netherlands, 2004. 

2. Salitermkan, S. S. Fundamentals of BioMEMS and Medical Microdevices; Wiley-Interscience, 

SPIE PRESS (Bellingham, Washington USA), 2006. 

3. Lin, V. S. Y.; Motesharei, K.; Dancil, K. P. S.; Sailor, M. J.; Ghadiri, M. R., A porous siliconbased 

optical interferometric biosensor, Science 1997, 278, 840-843. 

Δλ Δλ (nm) 

35 

30 

25 

20 

15 

10 

5 

0 

B 

0 20 40 60 80 

c-DNA concentration (μM) 

220


4. De Stefano, L.; Moretti, L.; Rendina, I.; Rossi, A. M., Time-resolved sensing of chemical species 

in porous silicon optical microcavity, Sensor and Actuators B 2004, 100, 168-172. 

5. Dancil, K. -P. S.; Greiner, D. P.; Sailor, M. J., A Porous Silicon Optical Biosensor: Detection of 

Reversible Binding of IgG to a Protein A-Modified Surface, J. Am. Chem. Soc. 1999, 121, 7925- 

7930. 

6. Snow, P. A.; Squire, E. K.; Russel, P. St. J.; Canham, L. T., Vapor sensing using the optical 

properties of porous silicon Bragg mirrors, J. Appl. Phys. 1999, 86, 1781-1784. 

7. Lorenzo, E.; Oton, C. J.; Capuj, N. E.; Ghulinyan, M.; Navarro-Urrios, D.; Gaburro, Z.; Pavesi, 

L., Porous silicon-based rugate filters, Appl. Opt. 2005, 44, 5415-5421. 

8. Ouyang, H.; Chrtistophersen, M.; Viard, R.; Miller, B. L.; Fauchet, P. M., Macroporous silicon 

microcavities for macromolecule detection, Adv. Funct. Mater. 2005, 15, 1851-1859. 

9. De Stefano, L.; Rotiroti, L.; Rea, I.; Rendina, I.; Moretti, L.; Di Francia, G.; Massera, E.; 

Lamberti, A.; Arcari, P.; Sangez, C., Porous silicon-based optical biochips, Journal of Optics A: 

Pure and Applied Optics 2006, 8, S540-S544. 

10. Wongmanerod, C.; Zangooie, S.; Arwin, H., Determination of pore size distribution and surface 

area of thin porous silicon layers by spectroscopic ellipsometry, Applied Surface Science 2001, 

172, 117-125. 

11. Yin, H.B.; Brown, T.; Gref, R.; Wilkinson, J.S.; Melvin, T., Chemical modification and 

micropatterning of Si(100) with oligonucleotides, Microelectronic Engineering 2004, 73-74, 830- 

836. 

© 2007 by MDPI (http://www.mdpi.org). Reproduction is permitted for noncommercial purposes. 

221

Examples of optical biosensors for the detection of analytes of 

clinical, industrial, and home security interest. 

This section is aimed at the characterization of a PSi device functionalized with some 

proteins and enzymes. 

It was studied that Glutamine-binding protein (GlnBP) from Escherichia coli specifically 

binds Glutamine and a wheat gliadin peptide (mainly responsible of inducing celiac 

disease), and that acetylcholinesterase (AChE) activity can be inhibited with 

organophosphorous pesticides, nerve agents and nitro-aromatic compounds, very 

dangerous substances. 

In collaboration with Dr. S. D’Auria from Institute for Protein Biochemistry, National Council 

of Research, Italy

Glutamine-Binding Protein from Escherichia coli Specifically Binds a 

Wheat Gliadin Peptide Allowing the Design of a New Porous 

Silicon-Based Optical Biosensor † 

Luca De Stefano, ‡ Mauro Rossi, § Maria Staiano, | Gianfranco Mamone, § Antonoietta Parracino, | 

Lucia Rotiroti, ‡ Ivo Rendina, ‡ Mosè Rossi, | and Sabato D’Auria* ,| 

Institute for Microelectronics and Microsystems, CNR, Naples, Italy, Institute of Food Sciences, CNR, Avellino, 

Italy, and Institute of Protein Biochemistry, CNR, Naples, Italy 

Received January 20, 2006 

In this work, the binding of the recombinant glutamine-binding protein (GlnBP) from Escherichia coli 

to gliadin peptides, toxic for celiac patients, was investigated by mass spectrometry experiments and 

optical techniques. Mass spectrometry experiments demonstrated that GlnBP binds the following amino 

acid sequence: XXQPQPQQQQQQQQQQQQL, present only into the toxic prolamines. The binding of 

GlnBP to gliadin suggested us to design a new optical biosensor based on nanostructured porous 

silicon (PSi) for the detection of trace amounts of gliadin in food. The GlnBP, which acts as a molecular 

probe for the gliadin, was covalently linked to the surface of the PSi wafer by a proper passivation 

process. The GlnBP-gliadin interaction was revealed as a shift in wavelength of the fringes in the 

reflectivity spectrum of the PSi layer. The GlnBP, covalently bonded to the PSi chip, selectively 

recognized the toxic peptide. Finally, the sensor response to the protein concentration was measured 

in the range 2.0-40.0 µg/L and the sensitivity of the sensor was determined. 

Keywords: gliadin • optical biosensors • porous silicon • celiac disease 

There is a strong need for integrated and automated bioanalytical 

systems in medical, food, agricultural, environmental, 

and defense testing. Now, the main market is shared between 

blood glucose, pregnancy, antibody-based infectious diseases, 

and biological warfare agent detection. 1 The interaction between 

an analyte and a biological recognition system is 

normally detected in biosensors by the transducer element 

which converts the molecular event into a measurable effect, 

such as an electrical or optical signal. Because of its spongelike 

structure, porous silicon (PSi) is an almost ideal material as a 

transducer: its surface has a specific area of the order of 200- 

500 m 2 cm -3 , so that a very effective interaction with liquid or 

gaseous substances is assured. 2,3 PSi optical sensors are based 

on changes of photoluminescence or reflectivity when exposed 

to the target analytes which substitute the air into the PSi 

pores. 4 The effect depends on the chemical and physical 

properties of each analyte, so that the sensor can be used to 

recognize the pure substances. Because of the sensing mechanism, 

these kinds of devices are not able to identify the 

components of a complex mixture. To enhance the sensor 

selectivity through specific interactions, some researchers have 

† The authors wish to dedicate this work to the memory of Prof. Arturo 

Leone, Director of the Institute of Food Sciences, CNR, Avellino, Italy. 

* Addess correspondence to Dr. Sabato D’Auria, Institute of Protein 

Biochemistry, Consiglio Nazionale delle Ricerche, Via Pietro Castellino 111, 

80131 Naples, Italy. Fax, +39-0816132277; e-mail, s.dauria@ibp.cnr.it. 

‡ Institute for Microelectronics and Microsystems, CNR. 

§ Institute of Food Sciences, CNR. 

| Institute of Protein Biochemistry, CNR. 

proposed to chemically or physically modify the PSi surface; 

the common approach is to create a covalent bond between 

the porous silicon surface and the biomolecules which specifically 

recognize the unknown analytes. 5-7 Among different 

probes of biological nature, ligand-binding proteins are particularly 

good candidates in designing highly specific biosensors 

for small analytes; in particular, the glutamine-binding protein 

(GlnBP) from Escherichia coli is a monomeric protein composed 

of 224 amino acid residues (26 kDa) responsible for the 

first step in the active transport of L-glutamine across the 

cytoplasm membrane. The GlnBP consists of two similar 

globular domains, the large domain (residues 1-84 and 186- 

224) and the small domain (residues 90-180), linked by two 

peptides. The deep cleft formed between the two domains 

contains the ligand-binding site. The GlnBP binds L-glutamine 

with a dissociation constant Kd of 5 × 10 -9 M 8 as well as polyglutamine 

residues. In this study, we show that GlnBP is able 

to bind with high affinity the amino acid sequences present in 

the gliadin allowing the design of a new reagentless microsensor 

for the optical interferometric detection of gliadin based 

on a chemically modified porous silicon nanostructured surface 

covalently linking the GlnBP. The gliadin by the wheat gluten 

is mainly responsible of inducing celiac disease, an inflammatory 

disease of the small intestine affecting genetically susceptible 

people. 9 Presently, a strict gluten-free lifelong diet is 

mandatory for celiac patients for both intestinal mucosal 

recovery and prevention of complicating conditions such as 

lymphoma. However, dietary compliance has been shown to 

10.1021/pr0600226 CCC: $33.50 © 2006 American Chemical Society Journal of Proteome Research 2006, 5, 1241-1245 1241 

Published on Web 03/21/2006

esearch articles De Stefano et al. 

be poor in most patients mainly because of inadvertent gluten 

consumption. From this point of view, a very sensitive assay 

for detection of toxic components of prolamins in food 

represents a perspective worth pursuing. Furthermore, the 

gliadin is a mixture of many proteins. 10 This observation 

highlighted the difficulties in identifying sequences useful for 

developing an immunoassay addressed toward the toxic portions 

of gliadin. The obtained results show that, even if the 

protein is blocked on the PSi surface, the GlnBP can strongly 

bind the substrate and work as a molecular probe for the 

detection of gliadin. 

Materials and Methods 

Materials. All the chemicals used were commercial samples 

of the purest quality. 

Protein Purification. The recombinant glutamine-binding 

protein from E. coli was prepared and purified according to 

D’Auria et al. 7 The protein concentration was determined by 

the method of Bradford 11 with bovine serum albumin as 

standard on a double beam Cary 1E spectrophotometer (Varian, 

Mulgrade, Victoria, Australia). 

Mass Spectrometry Analysis. MALDI-TOF experiments were 

carried out on a PerSeptive Biosystems (Framingham, MA) 

Voyager DE-PRO instrument equipped with a N2 laser (337 nm, 

3 ns pulse width). Each spectrum was taken with the following 

procedure: 1 µL-aliquot of sample was loaded on a stainless 

steel plate together with 1 µL of matrix R-cyano-4-hydroxycinnamic 

acid (10 mg in 1 mL of aqueous 50% acetonitrile). Mass 

spectrum acquisition was performed in both positive linear and 

reflectron mode by accumulating 200 laser pulses. The accelerating 

voltage was 20 kV. External mass calibration was 

performed with mass peptide standards (SIGMA). Tandem 

mass spectrometry (MS/MS) data for the extracted peptides 

were obtained using a Q-STAR mass spectrometer (Applied 

Biosystems) equipped with a nanospray interface (Protana, 

Odense, Denmark). Dried samples were resuspended in 0.1% 

TFA, desalted using Zip-Tip C18 microcolumns (Millipore), and 

sprayed from gold-coated, ‘medium length’ borosilicate capillaries 

(Protana). The capillary voltage used was 800 V. Doubly 

charged ion isotopic cluster was selected by the quadrupole 

mass filter (MS1) and then induced to fragment by collision. 

The collision energy was 30-40 eV, depending on the size of 

the peptide. The collision-induced dissociation was processed 

using Analyst 5 software (Applied Biosystems). The deconvolution 

MS/MS was manually interpreted with the help of the 

Analyst 5 software. 

Porous Silicon Chip Preparation and GlnBP Immobilization. 

Proof of PSi in biosensing has been experimentally 

demonstrated in a variety of sensors devices, including rugate 

filters, Bragg filters, single-layer films, and protein and DNA 

microarray, due to the large internal surface area which 

depending upon depth can easily be over 1000 times that of a 

planar surface of equal diameter, thus, enabling the immobilization 

of large amounts of selective receptors. 12-14 

However, for the device to function effectively, the porous 

morphology should enable homogeneous diffusion everywhere 

in the matrix. In the case of porous silicon, the porosity, 

determined by pore size and distribution, is tunable over a 

rather wide range by properly setting the fabrication parameters, 

that is, current density, etching time, etching solution, 

and wafer type. It is therefore possible to optimize the porous 

matrix structure to achieve maximum diffusion and maintain 

high level optical features. Previous ellipsometric characteriza- 

1242 Journal of Proteome Research • Vol. 5, No. 5, 2006 

tion has shown that the surface of the porous silicon layer as 

prepared could be covered by a 100-200 nm thin parasitic film 

of very low porosity (

GlnBP Binds Wheat Gliadin Toxic Sequences research articles 

Figure 1. Photoinduced reaction used to link GlnBP to the PSi chip surface. 

Figure 2. MALDI-TOF spectrum of gliadin digest bound to GlnBP. The amino acid sequence of MH + 2351.11 was determined by nanoESI- 

MS/MS. 

37 °C, the reaction was stopped by heating for 5 min; the 

peptic-tryptic (PT) digest was freeze-dried and stored at -20 

°C. 

A solution of homogeneous 2.0 mg/mL GlnBP in 1.0 mL of 

0.1 M bicarbonate buffer, pH 9.0, was mixed with 10 µL of 

rodhamine (Molecular Probes) solution in N,N-dimethylformamide 

(DMF) (1.0 µg of rodhamine/100 µL of DMF). The 

reaction mixture was incubated for 1.0 h at 30 °C, and the 

labeled protein was separated from the unreacted probe by 

passing over a Sephadex G-25 column equilibrated in 50 mM 

phosphate buffer and 100 mM NaCl, pH 7.0. 

The immobilization of GlnBP was carried out by applying 

50 µL of a 1 mg/mL solution prepared in 100 mM PBE buffer 

at pH 6.5 to each chip. The chips were incubated at -4 °C for 

1.0 h. After the reaction, the excess of reagent was removed 

and the wafers were extensively washed at room temperature 

by PBE buffer. 

To assess the protein penetration into the pores, we spotted 

20 µL of 1.0 mM sodium bicarbonate buffer containing the dyelabeled 

protein onto the porous silicon chip. This chip was 

observed by a confocal microscopy system. 

Results and Discussion 

GlnBP is a monomeric protein that binds glutamine (Gln) 

and poly-Gln residues with high affinity. 7,8 Since gluten proteins 

are rich in Gln residues, we questioned if GlnBP was able to 

bind amino acid sequences present in gluten proteins such as 

gliadin, a protein toxic for celiac patients. To check it, GlnBP 

was covalently bound to a CNBr-activated Sepharose 4B resin 

according to the manufacturer’s instructions (Amersham Biosciences 

Europe GmbH, Cologno Monzese, Italy) and a gliadin 

PT digest was passed through the column. The column was 

washed with 3 vol of phosphate buffer saline, and the peptide- 

(s) bound to GlnBP were eluted with 0.2 M glycine/HCl, pH 

3.0. The eluted fractions were utilized for mass spectrometric 

experiments. 

MALDI-TOF analysis on the isolated gliadin digest gave a 

complex peptide profile, which was difficult to rationalize due 

to the high sequence variability among the components of the 

gliadin fraction (Figure 2). In fact, the main difficulty in gliadin 

structural analysis lies in high protein heterogeneity owed to 

numerous duplications, and subsequent divergences of the 

gliadin multigene family encoding a polymorphic set of polypeptides. 

To identify the derived PT peptides, the gliadin 

sequences in the eluted fraction were analyzed by MS/MS 

experiments. Figure 3 shows the nanoESI-MS/MS spectrum 

of the triply charged ion at m/z 784.37 (MH + ) 2351.11 Da), 

corresponding to the sequence XXQPQPQQQQQQQQQQQQL 

(X indicates an unidentified residue) corresponding to that of 

wheat R-gliadin (Swiss-Prot, accession number Q9ZP09) with 

only the last 13 amino acid residues of peptide. 

The two amino acid residues located at the peptide Nteminal, 

corresponding to the “b2” fragment, were not exactly 

identified (Figure 3). In particular, since the “b2” fragment is 

232.08 Da, the first two amino acid residues could be either 

Met-Thr or Thr-Met. For this reason, we preferred to indicate 

them as XX. 

This result prompts us to develop a new methodology for 

sensing toxic sequences for celiac patients. First, we immobilized 

GlnBP to a porous silicon wafer and tested the 

strength of the covalent bond between the labeled GlnBP and 

the porous silicon surface by washing the chip in a demi-water 

flux. In Figure 4, a confocal microscopy image of the PSi chip 

under a He-Ne laser beam irradiation is shown. We found that 

the fluorescence is very high and homogeneous on the whole 

surface. We also checked, by Fourier transformed infrared (FT- 

IR) spectroscopy, that the PSi surface covered by the protein 

shows a good stability toward aqueous oxidation: after the 

water treatment, the characteristic peak of Si-O-Si bonds did 

not appear in the FT-IR spectrum. In the presence of gliadin 

PT, GlnBP undergoes a large conformational change in its 

global structure to accommodate the ligand inside the binding 

Journal of Proteome Research • Vol. 5, No. 5, 2006 1243

esearch articles De Stefano et al. 

Figure 3. MS/MS spectra of triple ion charged 784.37 (M ) 2351.11 Da). 

Figure 4. Confocal microscopy image of the PSi chip after 

immobilization of rodhamine-labeled GlnBP. 

site. The ligand-binding event was detected as a fringe shift in 

wavelength, which corresponds to a change in the optical path 

nd. Since the thickness d is fixed by the physical dimension of 

the PSi matrix, the variation is clearly due to changes in the 

average refractive index. 

The experimental measurement to detect the binding of 

gliadin to GlnBP is a two-step procedure: first, we registered 

the optical spectrum of the porous silicon layer after the GlnBP 

absorption on the chip surface and, then, after the PT-gliadin 

solution has been spotted on it. The organic material excess 

was removed by further rinses in the buffer solution. In Figure 

5, the shifts induced in the fringes of the reflectivity spectrum 

are reported. A well-defined red-shift of 2.6 ( 0.1 nm all over 

a wide range of wavelengths, due to protein-ligand interaction, 

was registered. Control experiments were made by using two 

different solutions containing nontoxic peptides produced 

following a peptic-tryptic digestion of zein, the corresponding 

prolamines from corn, and rice prolamines, that represent safe 

1244 Journal of Proteome Research • Vol. 5, No. 5, 2006 

Figure 5. Fringes’ shift due to protein-protein interaction. 

Continuous curve, porous silicon and GlnBP; dashed curve, 

porous silicon, GlnBP, and PT-gliadin. Wavelength shift between 

the continuous curve and the dashed curve is 2.6 (0.1) nm. 

cereals for celiac patients; the fringes’ shift in the reflectivity 

spectrum, with respect to the one obtained after the GlnBP 

absorption on the chip surface, was undetectable (∆λ < 0.1 

nm) in each single measurement. 

We also measured the signal response to the protein 

concentration after the ligand interaction. Figure 6 shows the 

dose-response curve as a function of the PT-gliadin concentration. 

Since the sensor displays a linear response between 

2.0 and 8.0 µM, it is possible to calculate the sensitivity of the 

optical interferometer to the concentration of PT-gliadin by 

estimating the slope of the curve: sGli ) 421 (13) nm/µM. The 

sensor response saturates approximately at 35 µM, which 

means that about 45% of the spotted proteins have bound the 

respective peptide. 

The rationale of the proposed strategy is that gluten proteins 

(gliadins and glutenins) are characterized by a high content of 

glutamine (Gln) residues and, consequently, they represent a

GlnBP Binds Wheat Gliadin Toxic Sequences research articles 

Figure 6. Dose-response curve. Optical variations as a function 

of the PT-gliadin concentration. 

substrate for the GlnBP. As a consequence, the relevance of 

these results for the topic of gluten determination has been 

addressed. The detection of residual gluten in food is fundamental 

for celiacs patients. Among toxic prolamins, gliadin has 

been very well characterized. Gliadin is a mixture of many 

proteins; 10 biochemical analysis revealed the existence of a 

relationship among the various constituents, so gliadin fractions 

have been grouped in three classes named R, γ, and ω 

gliadins. 20 To date, several immuno-active gliadin peptides in 

the celiac population have been identified. 21,22 Some of these 

epitopes are encompassed in a 33-mer gliadin peptide found 

to be resistant to digestion by gastric and pancreatic enzymes. 22 

It is noteworthy that most of these T-cell epitopes were 

recognized following deamidation catalyzed by tissue transglutaminase 

(tTG), in which some specific glutamine residues 

were converted to glutamic acid. 23 Moreover, T-cell epitopes 

cluster in regions that are rich in proline residues. 24 Taken 

together, these observations highlighted the difficulties in 

identifying sequences useful for developing immunoassays 

addressed toward the toxic portions of gliadin. Ideally, a 

method to determine gliadin should be applicable in a wide 

range of food, irrespective of processing, and should be directly 

related with toxicity. Until now, none of the produced methods 

were considered to be fully satisfactory. 25,26 On the other hand, 

the use of reducing agents, that are present in the proposed 

solvent, can improve the extraction of prolamines but affect 

their immunochemical quantification. 27 From this point of 

view, our method can potentially work also in reducing 

conditions, thus, overcoming the problem related to prolamin 

extraction. 28 

In conclusion, in this work, we have presented useful data 

for the development of an optical protein microsensor, based 

on porous silicon nanotechnology, for an easy and rapid 

detection of gliadin. The porous silicon matrix is demonstrated 

to be very useful as transducer material in this kind of 

biosensor: its highly specific area ensures good sensitivities 

in organic molecules sensing, and at the same time, it supplies 

a fast and easy readable optical response. Even if the GlnBP is 

covalently bonded to the PSi surface, due to a proper functionalization 

process, it still selectively recognizes the target 

analyte with an overall efficiency of 45%. Genetic manipulation 

experiments for improving the protein affinity to PT-gliadin are 

in progress. 

Abbreviations: GlnBP, glutamine-binding protein; Gln, 

glutamine; PSi, porous silicon; MS/MS, tandem mass spectrometry; 

nanoESI, nanospray electrospray ionization. 

Acknowledgment. This project was realized in the 

frame of the CRdC-ATIBB and CRdc-NTP POR UE-Campania 

Mis 3.16 activities and in the frame of the CNR Commessa 

“Diagnostica avanzata ed alimentazione”. 

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Proceedings of the 17th Meeting of the Working Group on 

Prolamin Analysis and Toxicity, London, U.K., October 3-6, 2002; 

Verlag Wissenschaftliche Scripten, Zwickau. 

PR0600226 

Journal of Proteome Research • Vol. 5, No. 5, 2006 1245

INSTITUTE OF PHYSICS PUBLISHING JOURNAL OF PHYSICS: CONDENSED MATTER 

J. Phys.: Condens. Matter 18 (2006) S2019–S2028 doi:10.1088/0953-8984/18/33/S17 

Nanostructured silicon-based biosensors for the 

selective identification of analytes of social interest 


Sabato D’Auria 1,4 ,Marcella de Champdoré 1 ,Vincenzo Aurilia 1 , 

Antonietta Parracino 1 ,MariaStaiano 1 ,Annalisa Vitale 1 ,MosèRossi 1 , 

Ilaria Rea 2 ,LuciaRotiroti 2 ,Andrea M Rossi 3 ,Stefano Borini 3 , 

Ivo Rendina 2 and Luca De Stefano 2 

1 Institute of Protein Biochemistry, CNR, ViaPietro Castellino, 111 80131 Naples, Italy 

2 Institute for Microelectronics and Microsystems, CNR—Department of Naples, Via Pietro 

Castellino 111, 80131 Naples, Italy 

3 Istituto Nazionale di Ricerca Metrologica—INRIM, Via Strada delle Cacce 91, 10100 Turin, 

Italy 

E-mail: s.dauria@ibp.cnr.it 

Received 13 January 2006 

Published 4 August 2006 

Online at stacks.iop.org/JPhysCM/18/S2019 

Abstract 

Small analytes such as glucose, L-glutamine (Gln), and ammonium nitrate 

are detected by means of optical biosensors based on a very common 

nanostructured material, porous silicon (PSi). Specific recognition elements, 

such as protein receptors and enzymes, were immobilized on hydrogenated 

PSi wafers and used as probes in optical sensing systems. The binding 

events were optically transduced as wavelength shifts of the porous silicon 

reflectivity spectrum or were monitored via changes of the fluorescence 

emission. The biosensors described in this article suggest a general approach 

for the development of new sensing systems for a wide range of analytes of 

high social interest. 

(Some figures in this article are in colour only in the electronic version) 

For many decades, scientists have recognized the power of incorporating biological principles 

and molecules into the design of artificial devices. Biosensors, an amalgamation of signal 

transducers and biocomponents, play a prominent role in medicine, food, and processing 

technologies. Compactness, portability, high specificity, and sensitivity represent some reasons 

why biosensors are considered as holding promise, with high potential for replacing current 

analytical practice [1]. 

4 Address for correspondence: Institute of Protein Biochemistry, Italian National Research Council, Via Pietro 

Castellino, 111-80131 Naples, Italy. 

0953-8984/06/332019+10$30.00 © 2006 IOP Publishing Ltd Printed in the UK S2019

S2020 SD’Auriaet al 

Fluorescence detection is the dominant analytical approach in medical testing, 

biotechnology, and drug discovery. Starting in the 1980s the first chemically synthesized 

fluorescence probes for specific analytes became available. However, the development of 

probes for complex analytes present in biological specimens cannot be approached via chemical 

synthesis alone [2–4]. In fact, the number of potential ligands specifically recognized by 

different proteins is very large and ranges from small molecules to macromolecules (including 

proteins themselves). The advantages of using proteins as components of biosensors are many 

and include relatively low costs in design and synthesis, the fact that proteins are, at least in 

general, soluble in water, and finally, with the progress of molecular genetics, the possibility of 

improving/changing some of the propertiesoftheproteins by genetic manipulation [5, 6]. 

Recently, a lot of experimental work exploiting the properties of porous silicon (PSi) in 

chemical and biological sensing has been reported [7, 8]. PSi is an almost ideal material as a 

transducer due to its porous structure, with a hydrogen terminated surface, having a specific 

area of the order of 200–500 m 2 cm −3 ,soaveryeffective interaction with several adsorbates 

is ensured. Moreover, PSi is an available and low cost material, completely compatible 

with standard IC processes, and a useful component of so-called smart sensors [9]. The 

PSi is fabricated by the electrochemical etching of a silicon (Si) wafer in a hydrofluoridic 

acid solution. It is well known that the porous silicon ‘as etched’ has an Si–H terminated 

surface due to the Si dissolution process [10]. PSi optical sensors are based on changes of 

photoluminescence or reflectivity upon exposure to the target analytes [11, 12], which replace 

the air in the PSi pores. The optical changes depend on the chemical and physical properties 

of the bound analyte. As a consequence, these devices cannot be used for the identification 

of analytes in a complex mixture. In order to enhance the sensor selectivity, it is possible to 

utilize PSi wafer as substrate for the covalent immobilization of biomolecules, such as proteins, 

enzymes, and antibodies, that are able to recognize target analytes in a complex matrix [13–20]. 

In this paper, we report new findings of relevance to the design of advanced biosensors 

obtained by a multidisciplinary group of scientists, for the optical detection of analytes of social 

interest such as glucose, ammonium nitrate, and glutamine. 

2. Glucose detection 

Close control of blood glucose is essential to avoid the long term adverse consequences of 

elevated blood glucose, including neuropathies, blindness, and other consequences [21]. Noninvasive 

measurements of blood glucose have been a long-standing research target because 

they would allow the development of a variety of devices for diabetic health care, including 

continuous painless glucose monitoring, control of an insulin pump, and warning systems for 

hyperglycaemic and hypoglycaemic conditions. Hypoglycaemia is a frequent occurrence in 

diabetics, and can result in coma or death. The acute and chronic problems of diabetics and 

hypoglycaemia can be ameliorated by continuous monitoring of blood glucose. At present, 

the only reliable method for measuring blood glucose is via a finger prick and subsequent 

glucose measurement, typically using glucose oxidase. This procedure is painful and even 

the most compliant individuals, with good understanding and motivation for glucose control, 

are not willing to prick themselves more than several times per day. These medical needs 

have fostered intensive efforts to develop sensors for glucose [22, 23]. The absence of a 

suitable non-invasive glucose measurement has fostered decades of research, little of which 

has however resulted in simpler and/or improved glucose monitoring. Included in this effort is 

the development of fluorescence probes specific for glucose, typically based on boronic acid 

chemistry [24, 25]. An alternative approach to glucose sensing using fluorescence is based 

on proteins which bind glucose. Optical detection of glucose appears to have had its origin

Nanostructured biosensors S2021 

in the promising studies of Schultz and co-workers [26, 27], who developed a competitive 

glucose assay which does not require substrates and does not consume glucose. This assay used 

fluorescence resonance energy transfer (FRET) between a fluorescence donor and an acceptor, 

each covalently linked to concanavalin A (ConA) or dextran. In the absence of glucose the 

binding between ConA and dextran resulted in a high FRET efficiency. The addition of glucose 

resulted in its competitive binding to ConA, displacement of ConA from the labelled dextran, 

and a decrease in FRET efficiency. These results generated considerable enthusiasm for 

fluorescence sensing of glucose [28]. The glucose–ConA system was also studied by other labs. 

They identified several problems, namely that the system was only partially reversible upon 

addition of glucose, and that it became less reversible with time and showed aggregation [29]. 

For this reason, the use of other glucose-binding proteins as sensors was explored by several 

research groups [30–35]. If a reliable fluorescence assay for glucose could be developed, then 

the robustness of lifetime-based sensing [36]could allow development of a minimally invasive 

implantable glucose sensor, or of a sensor which uses extracted interstitial fluid. The lifetime 

sensor could measure glucose concentration through the skin [37] usingaredlaser diode or 

light emitting diode (LED) device as the light source. These devices are easily powered with 

batteries and can be engineered into a portable device. An implantable sensor can be expected 

to report on blood glucose because tissue glucose closely tracks blood glucose with a 15 min 

time lag [38]. 

Glucose oxidase (GO) (EC 1.1.3.4) from Aspergillus niger catalyses the conversion of β- 

D-glucose and oxygen to D-glucono-1,5-lactone and hydrogen peroxide. It is a flavoprotein, 

highly specific for β-D-glucose, and is widely used to estimate glucose concentration in blood 

or urine samples through the formation of coloured dyes [39]. Because glucose is consumed, 

this enzyme cannot be used as a reversible sensor. We extended the use of GO under conditions 

where no reaction occurs. In particular, in order to prevent glucose oxidation, we removed the 

FAD cofactor which isstrictly required for the reaction. 

The absorbance spectrum of apo-GO shows the characteristic shape of the coenzyme-free 

proteins, with an absorbance maximum at 278 nm due to the aromatic amino acid residues. 

The absence of absorption at wavelengths above 300 nm indicates that the FAD has been 

completely removed. The fluorescence emission spectrum of apo-GO at room temperature 

upon excitation at 298 nm displays an emission maximum at 340 nm, which is characteristic of 

partially shielded tryptophan residues. The addition of 20 mM glucose to the enzyme solution 

resulted in a quenching of about 18% of the tryptophanyl fluorescence emission (data not 

shown). This result indicates that the apo-GO is still able to bind glucose. The observed 

fluorescence quenching may be mainly ascribed to the tryptophanyl residue 426. In fact, as 

shown by x-ray analysis and molecular dynamics simulations, the glucose-binding site of GO 

is formed by Asp 584, Tyr 515, His 559 and His 516. Moreover, Phe 414, Trp 426 and Asn 514 

are in locations where they might form additional contacts with glucose [40]. 

The intrinsic fluorescence of proteins is usually not useful for clinical sensing because of 

the need for complex or bulky light sources and because of the presence of numerous proteins 

in most biological samples. ANS is known to be a polarity-sensitive fluorophore which displays 

an increased quantum yield in low polarity environments. Additionally, ANS frequently binds 

to proteins with an increase in intensity. We examined the effects of GO on the emission 

intensity of ANS. Addition of apo-GO to an ANS solution resulted in an approximate 30-fold 

increase in the ANS fluorescence intensity. Importantly, the intensity of the ANS emission was 

sensitive to glucose, decreasing by approximately 25% upon glucose addition. The ANS was 

not covalently bound to the protein. That addition of glucose results in a progressive decrease 

in the ANS fluorescence intensity suggests that the ANS is being displaced into a more polar 

environment upon glucose binding. The decreased ANS intensity occurred with a glucose-


Phase angle (deg) or modulation (%) 

Deviations 

Moduli (%) Phase 

100 

80 

60 

40 

20 

0 

2 

0 

–2 

2 

0 

0 glucose 

20 mM glucose 

40 mM glucose 

60 mM glucose 

T = 25°C 

Apo - GO - 1,8 - ANS 

–2 

1 10 100 

Frequency (MHz) 

Figure 1. Frequency domain intensity decays of 1,8-ANS-glucose oxidase at different 

concentrations of glucose. Excitation was at 335 nm while the emission was observed through 

interference filter 535/50 nm with two Corning 3–71 cut-off filters. 

binding constant near 10 mM, which is comparable to the KD of the holo-enzyme. The fact 

that the binding affinity has not changed significantly may imply that the binding is still specific 

to glucose. Intensity decays of ANS-labelled apo-GO in the presence of glucose were studied 

by frequency domain fluorometry. Addition of glucose shifts the frequency responses to higher 

frequencies, which is due to a decreased ANS lifetime (figure 1). The shorter lifetime of ANS 

apo-GO in the presence of glucose is again consistent with the suggestion that glucose displaces 

the ANS to a more polar environment. The mean lifetime decreases by over 40% upon addition 

of glucose. These results demonstrate that apo-glucose oxidase, when labelled with suitable 

fluorophores, can serve as a protein sensor for glucose. 

3. Explosivedetection 

Organophosphorous pesticides, nerve agents and nitro-aromatic compounds are very dangerous 

substances which can be used against civil or military objectives. Recently, Bourne and 

colleagues have demonstrated the mechanism of inhibition, at molecular level, of the 

acetylcholinesterase (AChE) with several inhibitor molecules [41]. The peripheral anionic site 

on AChE, located at the active centre gorge entry (figure 2), encompasses overlapping binding 

sites for allosteric activators and inhibitors; yet, the molecular mechanisms coupling this site 

to the active centre at the gorge base to modulate catalysis remain unclear. The peripheral site


A 

B 

C 

Figure 2. Structures of the DI–mAChE and PI–mAChE complexes. (A) Ribbon diagram of the 

mAChE dimer (cyan, with the four-helix bundle in magenta) bound to DI (orange bonds, blue 

nitrogen and red oxygen atoms). The carbohydrate moieties linked to residues Asn 350 and Asn 

464 in both subunits are displayed as grey bonds and coloured spheres. The side chains of the 

catalytic triad residues, Ser 203, Glu 334, and His 447, are shown as white bonds in the two dimer 

subunits. The PEG molecule bound at the centre of the four-helix bundle and the carbonate molecule 

bound to Ser 203 are shown as grey (green) and light grey (yellow) bonds, respectively. (B) Closeup 

stereo view of the DI molecule (coloured as in A) bound to the PAS, with the 2.35 ˚A resolution 

omit Fo − Fc electron density map contoured at 3.5σ (cyan) and 7.5σ (blue); the coordinates of 

this region were omitted and the protein coordinates were refined by simulated annealing before the 

phase calculation. The interacting side chains of mAChE residues His 287 and Leu 289, located 

in the loop region connecting helices α3 6,7 and α4 6,7 ,andof residues Trp 286 and Tyr 341 at the 

gorge entrance, are displayed as grey (green) bonds; those of mAChE residues Tyr 72 and Glu 

292, whose respective mutations as methionine and lysine in BgAChE abolish PI binding (see 

figure 6), are highlighted in light grey (orange). The chloride ions and solvent molecules are 

shown aspink and red spheres, respectively. The catalytic triad residues, Ser 203, Glu 334, and 

His 447, and the carbonate molecule (bottom) are shown as white and light grey (orange) bonds, 

respectively. Hydrogen bonds between mAChE and the DI molecule are shown as white dotted 

lines. (C) Close-up stereo view of the PI molecule (coloured as for DI) bound to the PAS, with the 

2.35 ˚A resolution omit Fo − Fc electron density map contoured at 3.5σ (cyan). The PI phenyl and 

diethylmethylammonio moieties, which show alternative positions in the structure, are displayed as 

red and orange bonds. The mAChE side chains interacting with the PI molecule are displayed as 

grey (green) and light grey (orange) bonds as in (B). The mAChE molecular surfaces buried at the 

DI–mAChE (B) and PI–mAChE (C) complex interfaces are displayed in transparency.


Fluorescence Intensity (au) 

300000 

250000 

200000 

150000 

100000 

50000 

0 

580 

Acetylcholine (2mM) 

Acetylcholine (2mM) + Ammonium Nitrate (20mM) 

590 600 610 620 630 


640 650 

Figure 3. Fluorescence emission spectra of AChE immobilized on Psi, in the absence and in the 

presence of ammonium nitrate. 

has also been proposed to be involved in heterologous protein associations occurring during 

synaptogenesis or upon neurodegeneration [42]. 

A crystal form of mouse AChE, combined with spectrophotometric analyses of the 

crystals, allowed us to obtain unique structures of AChE with a free peripheral site, and 

as three complexes with peripheral site inhibitors: the phenylphenanthridinium ligands, 

decidium and propidium, and the pyrogallol ligand, gallamine, at 2.20–2.35 ˚A resolution [43]. 

Comparison with structures of AChE complexes with the peptide fasciculin or with organic 

bifunctional inhibitors unveiled new structural determinants contributing to ligand interactions 

at the peripheral site, and permitted a detailed topographic delineation of this site [44]. 

These structures provide templates for designing compounds directed to the enzyme surface 

that modulate specific surface interactions controlling catalytic activity and non-catalytic 

heterologous protein associations. Since several dangerous substances inhibit AChE activity, 

first responders and personnel at risk for exposure to these compounds need sensor systems that 

operate in real time and are reliable, cost-effective, compact, portable, and sensitive [45]. 

AChE was immobilized on a porous silicon chip by applying 50 µl ofa1mgml −1 

solution prepared in 100 mM PBE buffer at pH 6.5 to each chip. The chip was incubated 

at 4 ◦ C overnight. After the reaction, the excess of reagent was removed and the wafers were 

extensively washed at room temperature with PBE buffer. The AChE activity was monitored 

by using the Amplex ® Red Acetylcholine/Acetylcholinesterase Assay Kit, by following the 

fluorescence variations at 590 nm. 

In figure 3, we report the fluorescence emission in the absence and in presence of 

ammonium nitrate. As we can see, the presence ofammonium nitrate results in a quenching of 

the fluorescence emission of about the 39%, indicating the effectiveness of the optical biosensor 

for the detection of the basic compound of common explosives. 

4. Glutamine detection 

Glutamine (Gln) is an important energy and nitrogen source used in the culturing of eukaryotic 

cells [46, 47]. Since the catabolism of glutamine leads to the build-up of ammonia which may


Figure 4. Optical microscope image of a PSi chip spotted by a labelled protein after electronic 

beam irradiation under laser light excitation. The bright area due to the protein fluorescence is 

clearly evident. The inset shows the nanometric resolution of the pattern produced. 

become toxic to the growing culture and thus inhibit growth [48, 49], monitoring glutamine 

concentrations throughout the growth cycle is an important process. Biosensors available for 

glutamine are based on immobilized glutaminase [50, 51]. Complications in the determination 

of glutamine using this type of enzymatic biosensor arise due to the high degree of chemical 

similarity to glutamate as well as other amino acids. Consequently, additional proteins and 

electrodes must be added to take into account the interfering signals. 

The Escherichia coli periplasmic space contains a diverse group of binding proteins whose 

main function is to present various molecules, such as sugars, amino acids, peptides, and 

inorganic ions, for transport into the cell [52]. Glutamine-binding protein (GlnBP) from E. 

coli is a monomeric protein composed of 224 amino acid residues (26 kDa) responsible for the 

first step in the active transport of L-glutamine across the cytoplasmic membrane [53, 54]. 

Of the naturally occurring amino acids, only glutamine is bound by GlnBP, with a Kd of 

3 × 10 −7 M[54]. The protein consists of two similar globular domains linked by two peptide 

hinges, and x-ray crystallographic data indicate that the two domains undergo large movements 

upon ligand binding. The changes in global structure of GlnBP following the Gln binding make 

it a good candidate for serving as the basis for new biosensor design. 

We detected the molecular binding between the GlnBP from Escherichia coli and Gln by 

means of an interferometric microsensor based on the utilization of unmodified porous silicon 

as solid support for the chip preparation. The sensor operates by measurement of the Fabry– 

Perot fringes in the wavelength reflection spectrum from the porous silicon layer. The ligand 

binding was detected as a fringe shift in wavelength, corresponding to a refractive index change. 

We developed a three-step procedure for the Gln detection. In particular, firstly we registered 

the optical spectrum of the porous silicon layer as etched; then we repeated the measurement 

after the GlnBP absorption on the chip surface, and, finally, a Gln solution was spotted on it. 

Figure 3 showsthe shifts induced by each step in the fringes, due to protein–ligand interaction. 

Figure 4 shows how effective the linking between the protein and the PSi surface is after 

electron beam irradiation: the bright area is emitted as a spot on an irradiated PSi area and 

photographed by an optical microscope under laser (Ar + ) excitation after several rinses in 

demi-water. The real nanometric resolution of this functionalizing method is clearly visible. 

The graft effect is due to the desorption of surface hydrogen atoms which is provoked by the


(a) 

(b) 

∆nd (nm) 

160 

140 

120 

100 

80 

60 

Protein 

Protein-ligand complex 

10 15 20 25 30 35 40 

Concentration (µg/L) 

Figure 5. (a) Confocal microscopy image of the PSi chip infiltrated with the labelled GlnBP after 

dialysis in demi-water. (The fluorescence emittedbyrhodamine-labelled GlnBP covalently bounded 

to the photochemically modified PSi surface is registered by a confocal microscope under laser beam 

(He–Ne) exposure.) (b) Dose–response curve as a function of the concentration of the protein and 

the Gln–GlnBP complex. 

electron beam. The naked surface is thus covered by silicon radicals which strongly react 

with the proteins. In figures 5(a) and (b) the features of an optical biosensor are presented. 

In particular, the fluorescence emitted by rhodamine-labelled GlnBP immobilized on the 

photochemically modified PSi surface is shown in figure 5(a). We also measured the change in 

the optical path of the PSi layer due to the protein concentration before and after the interaction 

of the ligand. Figure 5(b) shows the dose–response curve as a function of the concentration 

of GlnBP and the complex GlnBP–Gln. In this curve, each point represents the average of 

three independent measurements and the error bars represent the standard deviations. In the 

protein concentration range investigated, the sensor exhibits a linear response (R 2 = 0.99) and 

it is possible to determine the sensitivity of the optical interferometer to the organic matter by 

estimating the slopes of the two curves: sprot = 21(1) nm µg −1 landscomp = 23(1) nm µg −1 l 

for the protein and complex, respectively. As expected, the relative distance between the two 

curves is constant, the protein–complex ratio being always 1:1.


5. Conclusions 

This paper reviews examples of optical biosensors for the detection of analytes of clinical, 

industrial, and home security interest. The biosensors described, based on porous silicon 

nanotechnology, may represent a model for designing advanced optical arrays (lab-on-a-chip) 

for simultaneously determining analytes of social interest. 

Acknowledgments 

This work was supported by CRdC-ATIBB POR UE-Campania Mis 3.16 activities (SD, MR) 

and bythe CNR Commessa Diagnostica avanzata ed alimentazione (SD). 

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Optical evidence of protein structural changes in porous 

silicon devices 

The protein-ligand molecular interactions imply strong geometrical and structural 

rearrangements of the biological complex which are normally detected by high sensitivity 

optical techniques such as time-resolved fluorescence microscopy. In this section are 

reported the measurements, made by optical spectroscopic reflectometry in the visiblenear 

infrared region, about the interaction between a sugars binding protein (TMBP), 

covalently bound on the surface of a porous silicon, and the glucose, at different 

concentrations and temperatures. 

I’m grateful to Dr. S. D’Auria from Institute for Protein Biochemistry, National Council of 

Research, Italy for supplying the TMBP protein and for helpful discussion about the 

experiments.

Introduction: proteins from extremophiles 

The proteins purified from thermophilic organisms are characterized by high stability in 

harsh environment such as high temperature, high ionic strength, extreme pH values, even 

in presence of elevated concentration of detergents and chaotropic agents so that they 

can be successfully used as very effective bio-probes in sensors design[1]. 

In particular, it has been recently demonstrated that the D-trehalose/D-maltose-binding 

protein (TMBP), which is part of the sugars uptake system in the hyperthermophilic 

organism archaeon T. litoralis, is also able to bind glucose molecules. This result suggests 

the possibility of using TMBP as a stable probe within a biological recognition system for 

glucose monitoring [1]. TMBP is a monomeric 48 kDa macromolecule constituted by twodomains 

and containing twelve tryptophan residues. The presence of these fluorescent 

residues has been used to study the interaction between TMBP and glucose at different 

temperatures by means of time-resolved fluorescence spectroscopy by Herman et al. [1]: 

they found that the highest affinity between the protein and this ligand was at a 

temperature around 60 °C, with a dissociation constant (Kd) of about 40 μM. At room 

temperature TMBP still binds substrates while the activity of the TMBP based transport 

system becomes negligible [3,4], and this behaviour can be ascribed to an increased 

rigidity of TMBP structure at room temperature [5]. As other sugar-binding proteins, TMBP 

also consists of two globular lobes linked by a hinge region made of a few polypeptide 

chains. The deep cleft formed between the two lobes, which present similarly tertiary 

structure, contains the ligand-binding site [6]. 

In a recent published paper [7], we have found that, at room temperature, the reflectivity 

spectra of a PSi distributed Bragg reflector (DBR), properly functionalized with TMBP, 

undergo a red shift on exposure to glucose solutions. In particular, we observed a red shift 

of about 1.2 nm after the interaction with a 150 μM glucose solution. The estimated 

sensitivity of the monitoring method was 0.03 nm/μM. 

On the other hand, some preliminary experiments conducted at 60 °C [8] have shown that 

the reflectivity spectra of different PSi-based structures, functionalized with TMBP, 

undergo blue-shifts as a consequence of the interaction with glucose. 

Than these apparently conflicting results was explained by combining our investigation on 

two PSi devices, both TMBP functionalized: an optical microcavity (PSMC), characterized 

by optical spectroscopic reflectometry, and a thin PSi monolayer, as a model system for 

the ellipsometric characterization in order to correlate the observed blue shifts of the 

PSMC spectra to the typical conformational changes of TMBP at 60 °C. 


D-Glucose and all the other chemicals used in this experiment were from Sigma. All 

commercial samples were of the best available quality. The TMBP was purified and 

supplied by dr. S. D’Auria of the Institute for Protein Biochemistry, National Council of 

Research, Napoli, Italy. Proteins were expressed, purified and quantified by his laboratory 

as described in reference 5, and references therein. 

In this experiment were designed and fabricated two PSi-based structures: a monolayer, 

that is a porous layer with fixed porosity, and a microcavity, which is constituted by a lowporosity 

(high refractive index) layer between two Bragg reflectors, each one obtained by 

alternating low and high porosity layers for 7 times. Both the samples were produced by 

electrochemical etching of a very highly doped n 

55 

++ -silicon wafers (Siltronix Inc., USA), 

oriented, 0.001 Ω cm resistivity, 400 μm thick, in a HF-based solution. The silicon 

was etched using a 50 wt.% HF/ethanol solution with halogen lamp illumination and at 

room temperature. Before anodization the substrate was placed in HF solution to remove 

the native oxide. The monolayer, characterized by variable angle spectroscopic

ellipsometry (Horiba-Jobin-Yvon, mod. UV-VISEL) in the wavelength range between 400 

and 1700 nm, was 503±1 nm thick with a porosity of 77.5±0.3 %. The porous silicon 

microcavity (PSMC) was electrochemically etched applying a current density of 400 

mA/cm 2 for 0.3 s to obtain the low refractive index layers, with a porosity of 59±1 %, while 

one of 50 mA/cm 2 was applied for 0.5 s for the high index layers, with a porosity of 50±1 

%. This optical structure has a characteristic resonance peak at 1017 nm, centered in a 

105 nm wide stop band. 

The covalent bind of the TMBP on the porous silicon surface is based on a three steps 

functionalization process constituted by a chemical passivation of the PSi surface after 

oxidation, as it is shown in Figure 1. The PSi optical structures have been thermally 

oxidized in O2 atmosphere, at 900 °C for 15 min; then, we have treated the surfaces with a 

proper chemical linker, the aminopropyltriethoxysilane (APTES). Samples have been 

rinsed by dipping in a 5 % solution of APTES and a hydroalcoholic mixture of water and 

methanol (1:1) for 20 min at room temperature. The chips were then washed by deionized 

water, and methanol, and dried in N2 stream. The silanized devices were then baked at 

100°C for 10 min. To create a surface able to link the amino group of the proteins, we have 

immersed the chips in a 2.5% glutaraldehyde (GA) solution in 20 mM HEPES buffer (pH 

7.4) for 30 min, and then rinsed it in deionized water and finally dried in N2 stream. The 

glutaraldehyde reacts with the amino groups on the silanized surface and coats the 

internal surface of the pores. 

O 2 Si 

Si H 

Si H 

Si H 

O 2 Si 

O 

OH 

O2 5% APTES 

SiO2 OH 

15' @900°C 20' R.T 

OH 

O 

O 

O 

Si 

O 

O 

O 

Si 

O 

C 

N 

NH 2 

2.5% GA 

30' R.T 

C N TMBP 

Figure 1. Porous silicon functionalization scheme: from the as-etched material up to the organic-inorganic 

chip. 

The experimental set-up (Figure 2) used to measure the reflectivity spectra is constituted 

by a white light, as source, directed on the porous silicon chip through a Y fiber. The same 

fiber was used to guide the output signal to an optical spectrum analyzer (OSA, Ando, 

Japan). The spectrum was measured over the range 600-1400 nm with a resolution of 0.2 

nm. A hot plate, driven by a temperature controller, has been used to perform 

measurements at 60±1 °C, verified by a thermocouple placed directly on the chip. 

56

Figure 2. Experimental setup used to measure the optical reflectivity spectra of porous silicon microcavity. 

The functionalized PSi samples work as active substrates for the TMBP protein: on each 

PSi optical structures were spotted 20 μl of a 173 μM TMBP solution in 2mM phosphate 

buffer solution (pH 7.3) and incubated the system at 4 °C over night. After incubation and 

three washing steps of five minutes each, used to remove the excess of biological matter 

non covalently linked to the PSi surface, the presence of the protein in the spongy 

structure was optically verified as a red shift of about 60 nm in the reflectivity spectrum in 

the case of PSMC. 

Experimental results and discussion 

When recorded at the temperature of 25 °C, the PSMC characteristic resonance peak 

undergoes only very small red shifts, of the order of 1 nm, on exposure to glucose 

solutions with increasing concentrations, up to 200 μM, according to the low affinity of 

TMBP with glucose at room temperature [7]: at this temperature few proteins can rearrange 

their conformational structure and efficiently bind the ligand. The results of this 

biomolecular interaction change dramatically if the same measurements are performed at 

60 °C: in Figure 3 is reported the peak shift of the functionalized microcavity on exposure 

to a 200 μM glucose concentration. The optical resonance undergoes a 4.5 nm shift 

toward lower wavelengths. 


1.0 

0.8 

0.6 

0.4 

0.2 

Unperturbed 

After exposure to a 200 μM glucose solution 

960 980 1000 1020 1040 1060 1080 

Δλ 


Figure 3. PSMC resonance blue shift on exposure to a glucose solution of 200 μM concentration. 

The dose-response curve is shown in Figure 4: the curve is linear in the range of glucose 

concentrations between 0 and 200 μM, with an estimated sensitivity to solute 

57

concentration variations of 0.022 (4) nm/μM. For higher glucose concentrations, the blue 

shift did not increased any more, which means that all the active proteins have bound their 

ligands. 

Δλ (nm) 

1 

0 

-1 

-2 

-3 

-4 

-5 

0 50 100 150 200 250 300 

Glucose concentration (μM) 

Figure 4: Dose-response curve for functionalized PSMC exposed to different concentrations of glucose. 

The large blue-shift recorded corresponds to a decrease of the optical path that the light 

exploits when reflected by the porous structure. Since the optical path is defined as the 

thickness times the average refractive index, and there is not any variation of the 

multilayer thickness, the blue-shift can be attributed to a decrease in the value of the 

average refractive index. The chemical passivated surface of the PSi is highly stable 

respect to the biological chemical buffer solutions, so that there can not be any further 

oxidation nor any corrosion of the silicon matrix: the apparent increase of the porosity is 

due to the proteins volume reduction after the interaction with the glucose. Most of the keylock 

recognition mechanisms, which are at the basis of the biological interactions such as 

antibody-antigen or protein-ligand, require large rearrangement of one bio-molecule to 

host the other. The binding of glucose to TMBP is strongly related to a movement of the 

protein lobes as well as to profound conformational changes in the hinge region. The lobes 

envelope on to the binding site and cause a net reduction of protein’s volume, as 

confirmed by fluorescence measurements [1]. In order to confirm these results, a very 

sensitive optical technique was used, the variable angle spectroscopic ellipsometry 

(VASE), to quantify the porosity changes of a thin PSi layer after TMBP functionalization 

and on exposure to increasing concentration of glucose solutions. In Table 1 the porosities 

estimated by ellipsometry after each functionalization step of the monolayer, are reported. 

Porosity (%) 

As etched 77.5 ± 0.3 

After oxidation 49.1 ± 0.3 

After APTES treatment 41.7 ± 0.5 

After GA treatment 35.5 ± 0.4 

After TMBP incubation 31.9 ± 0.5 

Table 1: Porosity decrease of the PSi monolayer due to the functionalization steps. 

58

Chemical and biological functionalization steps considerably reduce the average porosity 

of the PSi layer. Ellipsometric spectra have been also acquired after each addition of 

glucose solutions with increasing concentrations. In Figure 5 the changes in porosity of the 

PSi layer, ΔP, as estimated from the ellipsometric data, are shown as function of glucose 

concentration. On exposure to ligand buffer solutions, the VASE characterization gives a 

direct measurement of the variation of each layer component. Using these experimental 

data, we have estimated, after the addition of 250 μM of glucose, an increase of more than 

the 3.5 % in the monolayer porosity. 

ΔP(%) 

4.5 

4.0 

3.5 

3.0 

2.5 

2.0 

1.5 

1.0 

0.5 

0.0 

0 50 100 150 200 250 

Glucose concentration (μM) 

Figure 5: The porosity increase of a functionalized PSi monolayer as a function of glucose concentrations. 

On exposure to a concentration of 250 μM of glucose, an enhancement of more than 3.5 % in porosity is 

obtained. 

On the other hand, it was also numerically calculate which is the effect of such a porosity 

increment on the reflectivity spectrum of the PSi microcavity. The reflectivity spectrum of a 

microcavity with the same porosities and thicknesses of the one realized was simulated. 

The organic phase of the system has been modeled by supposing a Gaussian distribution 

for both the TMBP molecules presence throughout the upper layers of the microcavity and 

also for the correspondent increment of porosity, ΔP, due to the glucose interaction [9]. In 

particular, the ΔP distribution was peaked at the top of the structure (with a maximum 

value of 3.5 %) and was characterized by a HWHM of about 3 μm, which equals the halfwidth 

of the first Bragg reflector in the microcavity structure. The reflectivity spectra have 

been simulated by standard transfer matrix method [10] and results are reported in Figure 

6. 


0.8 

0.6 

0.4 

0.2 

0.0 

950 975 1000 1025 1050 1075 1100 


(a) 

(b) 

59

Figure 6: (a) Simulated reflectivity spectrum of the PSi microcavity used in the experiment. (b) Simulated 

reflectivity spectrum of the same structure after a non-homogeneous increase of porosity. 

The result of such simulations is a blue shift of the resonant peak of 6.0±1.0 nm, which is 

consistent with the experimental one of 4.6±0.5 nm, observed after the addition of a 

glucose solution, 250 μM. 

Even if highly specific, the protein-ligand interaction is in general reversible, so that 

proteins release completely the target molecules, especially when in buffer solutions. But 

when bound on the surface of a nanostructured material, proteins could be limited in their 

activity: after an overnight dialysis in deionized water, it was found a 2 nm red-shift of the 

resonant peak, which means that the ligand has been released only by some proteins. 

Anyway, when the measurements were repeated, the same blue-shift always about 5 nm 

with a 250 μM concentration of glucose was obtained, so that can be concluded that the 

effect was reproducible, at least for five replica made in the same experimental conditions, 

but was not perfectly reversible. 

References 

1. P. Herman, I. Barvik, M. Staiano, A. Vitale, J. Vecer, M. Rossi and S. D’Auria, Biochimica et 

Biophysica Acta 2007, 1774, 540-544. 

2. J. Diez, K. Diederichs, G. Greller, R. Horlacher, W. Boos, W. Welte, J. Mol. Biol. 2001;305:905-915. 

3. K. B. Xavier, L. O. Martins, R. Peist, M. Kossmann, W. Boos, H. Santos, J. Bacteriol. 1996, 178, 

4773-4777. 

4. R. Horlacher, K. B. Xavier, H. Santos, J. Di Ruggiero, M. Kossmann, W. Boos, J. Bacteriol. 1998, 

180, 680-689. 

5. P. Herman, M. Staiano, A. Marabutti, A. Variale, A. Scire, F. Tanfani, J. Vecer, M. Rossi, S. D’Auria, 

Proteins 2006, 63, 754-767. 

6. Y. J. Sun, J. Rose, B. C. Wang, C. D. Hsiao, J. Mol. Biol. 1998, 278, 219-222. 

7. L. De Stefano, A. Vitale, I. Rea, M. Staiano, L. Rotiroti, T. Labella, I. Rendina, V. Aurilia, M. Rossi, S. 

D’Auria, Extremophiles, DOI 10.1007/s00792-006-0058-6. 

8. L. De Stefano, L. Rotiroti, E. De Tommasi, A. Vitale, M. Rossi, I. Rendina and S. D’Auria, 

Proceedings of the SPIE 2007 Volume 6585, pp. 658517 . 

9. L. De Stefano and S. D’Auria, Journal of Physics Condensed Matter 2007, 19, 395009 (7pp). 

10. M. A. Muriel and A. Carballar, IEEE P. Technol. Lett. 9, 955 (1997). 

60

Oligonucleotides direct synthesis on porous silicon chip 

In this section is report the design, fabrication and characterization of a oligonucleotide 

(ON) biochip based on the PSi nanotechnology for direct DNA solid phase synthesis. 

In collaboration with Prof. G. Piccialli of University of Naples “Federico II”, Italy

Oligonucleotides direct synthesis on porous silicon chip 

Luca De Stefano 1 , Edoardo De Tommasi 1 , Ilaria Rea 1 , Lucia Rotiroti 1 , Luca Giangrande 1 , Giorgia 

Oliviero 2 , Nicola Borbone 2 , Aldo Galeone 2 and Gennaro Piccialli 2 * 

1 

Istituto per la Microelettronica e Microsistemi - Unità di Napoli, Consiglio Nazionale delle Ricerche, via P. 

Castellino 111, I-80131 Napoli, Italy and 2 Dipartimento di Chimica delle Sostanze Naturali, Università degli 

Studi di Napoli “Federico II”, Via D. Montesano 49, I-80131 Napoli, Italy 

ABSTRACT 

A solid phase oligonucleotide (ON) synthesis on 

porous silicon (PSi) chip is presented. The prepared Si- 

OH surface were analyzed by FT-IR and the OH 

functions were quantified by reaction with 3’phosphoramidite 

nucleotide building block. Short ONs 

were synthesized on the chip surface and the coupling 

yields evaluated. 

INTRODUCTION 

The interest in DNA-based diagnostic tests has recently 

increased since they can be used in a lot of important 

applications such as gene analysis, fast detection of 

biological warfare agents, and also in forensic cases. 

Numerous DNA detection systems, both optical and 

electrochemical, based on the hybridization between a 

DNA target and its complementary probe fixed on a solid 

support have been described. 1 DNA biochips, due to 

integration and miniaturization of all components, allow 

continuous, fast and sensitive detection of DNA 

interactions. The fabrication of a DNA biochip requires the 

immobilization of the DNA probe on the support surface, 1,2 

or, as an alternative, the synthesis of the ON strand directly 

on the device surface. 3-5 

On the other hand, Porous silicon (PSi) structures are 

quite ideal support material for chemical and biological 

sensing, mainly due to their high specific area, up to 600 

m 2 cm -3 , low cost and compatibility with standard integrated 

circuit processes. PSi is fabricated by the electrochemical 

etching of a doped silicon wafer in a hydrofluoridic 

Glass 

Silicon 

Stainless steel capillary tubes 

Resin glue 

Liquid 

input/output 

channels 

Reaction chamber Porous silicon transducer 

Fig. 1. Schematic and real view of the μ-chamber for DNA 

biochip applications. 

aqueous solution. Moreover, its surface can be properly 

functionalised to covalently link the DNA probe. 

In this work, we report the design, fabrication and 

characterization of a ON biochip based on the PSi 

nanotechnology for direct DNA solid phase synthesis. 

RESULTS AND DISCUSSION 

Porous silicon formation and chip assembly. We have 

designed and fabricated a DNA biochip, just integrating 

PSi, as the transducer material, and a glass slide, which 

ensures sealing and the interconnections for fluids inlet and 

outlet. The silicon wafer used in this work is a p + type, 

crystal orientation, with a resistivity of 8-12 

mΩcm. 6 The glass is a Borofloat 33 type, 1 mm thick. The 

reaction microchamber has been realised by a two-step 

electrochemical etching. The first step is a high current 

density (800 mA/cm 2 for about 300 s) electrochemical etch 

in hydrofluoric acid and ethanol (HF/EtOH 1:1,v/v) 

solution which creates a μ-well (100 μm deep and a volume 

of 10 μL) in the bulk silicon. The second step is a 

consecutive electrochemical etch which is used to fabricate 

the PSi layer at the bottom of the μ-chamber. The process 

parameters for this last step were: current density of 400 

mA/cm 2 and etching time of 3.2 s. After the mechanical 

drilling of the flow channels, the glass slide has been 

cleaned and activated for the anodic bonding process 

following standard cleaning procedures. The silicon chip 

has also been carefully rinsed in deionized water for several 

minutes. Silicon etched wafer and glass top prefabricated 

components were anodically bonded together with mutual 

alignment: a successful bonding, in terms of mechanical 

strength and bond quality, has been obtained at a 

temperature of 200 o C, voltage of 2.5 kV, with a process 

time of 1.5 minutes. The cross section of the chip structure 

and the general-view of the real device are shown in Figure 

1. 

Functionalization of porous silicon layer and ON 

syntheses. The successive treatment of porous silicon layer 

with a solution of conc. H2SO4/H2O2 (8:2, v/v, 15 min R.T.) 

assure the formation of Si-OH groups on the solid matrix. 

The chip was, then, rinsed in deionized water for several 

minutes and dried under reduced pressure. The FT-IR 

spectrum (Figure 2) of the chip surface showed the 

appearance of the characteristic broad bands around 1050

cm -1 (Si-O-Si) and 1065, 880 cm -1 attributable to Si-OH 

bonds. To determine the surface coverage of Si-OH groups, 

we reacted the chip 1 with tetrazole activated 5’-DMT-3’phosporamidite-timidine 

nucleotide (2 Scheme 1) which 

assure the presence of a very reactive phosphorous(III) and 

T 

Si-OH + 2 

O 

Si-O P O ODMT 

i 

O 

Si-O P O P P OH 

1 

OCE 

ii 

OH 

2 

4 

5 

T T T 

Si-OH + 3 

O 

Si-O P O 

O O 

i 

S 

ODMT 

O 

Si-O P O R O 

O 

P O P P OH 

1 

OCE 

6 

OCE 7 OCE 

ii 

T 

(pi) 2N 

P O ODMT 

CEO 

O O 

(pi) 2N S 

P O ODMT 

CEO 

T 

P P 

T 

P 

T 

OH 

2 3 

8 

Scheme 1. Solid phase synthesis on PSi-OH chip 1; P: 

phosphotriester function; R: -(CH 2) 2SO 2(CH 2) 2; P: 

phosphodiester function; CE: 2-cyanoethyl; i: standard ON chain 

elongation; ii; treatment with methylamine. 

a 5’-dimethoxytrytil group (DMT). 

The reaction cycle including the standard phosphorous 

oxidation and the capping step, afforded the PSi-3’thymidine 

4 which was quantified by UV/VIS spectroscopy 

monitoring 5’-dimethoxytrytil cation released by DCA 

treatment (ε =71700, λ = 498 nm). The Si-OH 

functionalization resulted to be in the range of 35-40 

nmol/chip and no significative differences were observed 

varying the concentration of the nucleotide/activator 

solution (30-60 mg/mL) or prolonging the coupling 

reaction time (30-180-min). Alternatively, PSi chip 1 was 

reacted with the phosphoramidite 3 thus obtaining the PSi 

support 6 (30-40 nmol/chip). The linker 3 ensures a more 

long and flexible arm for the reaction of terminal alcoholic 

OH groups and allows the release of the synthesized ON 

chains by a mild basic treatment, as well. On the supports 4 

and 6 three coupling cycles were performed with 

phosphoramidite 2 thus obtaining PSi supports 5 and 7 

respectively. The DMT measurements after each cycle 

indicated a coupling efficiency of 87-90 % for support 4 

and 92-95% for support 6. These findings are in agreement 

with previous reported ON synthetic procedures on PSi 

chips which demonstrated the important role of the spacer 

linker attached on porous silicon surface. 3-5 The detachment 

of the synthesized short thymidine oligomer from support 7 

was obtained by treatment with anhydrous methylamine 

(20 min at room temperature). The combined amine 

solution and washings (methanol and then water), dried 

under reduced pressure, were analyzed by ESI-MS 

spectroscopy. The mass data confirmed the presence of 

expected 3’-phosphate-T3-5’-OH product (m/z 931.2 MH + ). 

The treatment of 5 with methylamine did not release 

significative amount of nucleotide material from the chip, 

also confirmed by DMT deprotection and measurement 

T 

T 

T 

(upper) 

Fig. 2. FT-IR profiles of the PSi-surfaces 1 (bottom) and 5 

(upper). 

performed after the amine treatment. FT-IR spectrum 

(Figure 2) of the PSi-ON chip showed the appearance of 

the characteristic band around 1100 cm -1 and 860 cm -1 

attributable, respectively, at P=O and 5’-CH2-OH. 

CONCLUSION 

In this preliminary experiments of ON synthesis we have 

tested the direct reactivity of the freshly prepared PSi OH 

functionalized chips with standard phosphoramidite 

building block and confirmed the compatibility of the PSi- 

3’-bonded thymidine with the chemical treatments required 

by standard ON solid phase synthesis. The synthesis of 

support 4 allowed as to estimate the amount of accessible 

OH groups on the silicon surface and to obtain a 

reproducible nucleotide functionalized chip. The data were 

also confirmed by the synthesis of the PSi-support 6. We 

hypothesize that, after coupling yield optimization, the 

support types 4 and 6 could be exploited both to obtain 

deprotected ON bonded to a PSi chip and to achieve a 

standard ON solid phase synthesis on PSi matrix. 

REFERENCES 

1. Sassolas, A., Leca-Bouvier, B.D., Blum, L.J. (2008) 

Chem. Rev., 108, 109-139. 

2. Di Francia, G., La Ferrara, V., Manzo, S., Chiavarini, 

S. (2005) Biosensors and Bioelectronics, 21, 661-665. 

3. Pike, A.R., Lie, L.H., Eagling, R.A., Ryder, L.C., 

Patole, S.N., Connolly, B.A., Horrocks, B.R., Houlton, 

A. (2002) Angew. Chem. Int. Ed., 41, 615-617. 

4. Lie, L.H., Patole, S.N., Pike, A.R., Ryder, L.C., 

Connolly, B.A., Ward, A.D., Tuite, E.M., Houlton, A., 

Horrocks, B.R. (2004) Faraday Discuss., 125, 235-249. 

5. Bessueille, F., Dugas, V., Vikulov, V., Cloarec, J.P., 

Souteyrand, E., Martin, J.R. (2005) Biosensors and 

Bioelectronics, 21, 908-916. 

6. De Stefano, L., Arcari, P., Lamberti, A., Sanges, C., 

Rotiroti, L., Rea, I., Rendina, I. (2007) Sensors, 7, 214- 

221. 

*Corresponding Author. E-mail: picciall@unina.it

Porous silicon-based optical biosensors and biochips 

This section is aimed at the demonstration that a PSi surface functionalized, allow the 

attachment of linkers through stable covalent Si–C bonds, and also demonstrated to be 

compatible to the high temperature anodic bonding processes used for the realization of 

silicon-glass biochips.

Abstract 

Physica E 38 (2007) 188–192 

Porous silicon-based optical biosensors and biochips 

Ivo Rendina , Ilaria Rea, Lucia Rotiroti, Luca De Stefano 

Institute for Microelectronics and Microsystems, National Council of Research, Via P. Castellino 111, 80131 Napoli, Italy 

Available online 31 December 2006 

Porous silicon multilayered microstructures have unique optical and morphological properties that can be exploited in chemical and 

biological sensing. The large specific surface of nanostructured porous silicon can be chemically modified to link different molecular 

probes (DNA strands, enzymes, proteins and so on), which recognize the target analytes, in order to enhance the selectivity and 

specificity of the sensor device. We designed fabricated and characterized several photonic porous silicon-based structures, which were 

used in sensing some specific molecular interactions. The next step is the integration of the porous silicon-based optical transducer in 

biochip devices: at this aim, we have tested an innovative anodic bonding process between porous silicon and glass, and its compatibility 

with the biological probes. 

r 2007 Elsevier B.V. All rights reserved. 

PACS: 87.80. y; 07.07.Df; 42.79.Pw 

Keywords: Optical biosensors; Porous silicon; Biochip 


Recently, lot of experimental work, concerning the 

worth noting properties of nanostructured porous silicon 

(PSi) in chemical and biological sensing has been reported, 

showing that, due to its morphological and physical 

properties, PSi is a very versatile sensing platform [1–5]. 

A key feature for a recognition transducer is a large surface 

area: PSi has a porous structure with a specific area of the 

order of 200–500 m 2 /cm 3 , so that it is very sensitive to the 

presence of biochemical species which penetrate inside the 

pores. Moreover, PSi is an available, low-cost material, 

completely compatible with VLSI and micromachining 

technologies, so that it could usefully be employed in the 

fabrication of MEMS, MOEMS and smart sensors. 

The PSi optical sensing features are based on the changes 

of its photonic properties, such as photoluminescence or 

reflectance, on exposure to the gaseous or liquid substances. 

Unfortunately, these interactions are not specific, 

so that the PSi cannot be used as a selective optical 

Corresponding author. 

E-mail address: ivo.rendina@na.imm.cnr.it (I. Rendina). 

1386-9477/$ - see front matter r 2007 Elsevier B.V. All rights reserved. 

doi:10.1016/j.physe.2006.12.050 

ARTICLE IN PRESS 

www.elsevier.com/locate/physe 

transducer material. One way to overcome this limit is to 

chemically or physically modify the PSi hydrogenated 

surface in order to enhance the sensor selectivity through 

specific biochemical interactions. The reliability of the 

biosensor strongly depends on the functionalization 

process: how simple, homogenous and repeatable it can 

be. This fabrication step is also crucial for the stability of 

the sensor: it is well known that as-etched porous silicon 

has a very reactive Si–H terminated surface due to the Si 

dissolution process [6]. The substitution of the Si–H bonds 

with Si–C ones guarantees a much more stable surface 

from the thermodynamic point of view. 

Testing and demonstrating the PSi capabilities as a 

useful functional material in the optical transduction of 

biochemical interactions is only the first action in the 

realization of an optical biochip based on this nanostructured 

material. In this case, all the fabrication processes 

should be compatible with the utilization of biological 

probes and the feasibility of such devices must be proven. 

This means that the standard integrated circuit microtechnologies 

should be modified and adapted to this new 

field of application. A very strong interdisciplinary is 

required to match and resolve all these problems.

In this work, we review our recent experimental results in 

optical biosensing using PSi-based photonic devices and 

report our newest results about the design and fabrication 

of optical biochip by exploiting the PSi nanotechnology 

[7–12]. 

2. Materials and methods 

Following the pioneering work of Mike Sailor’s group 

[13], we started using the PSi monolayers, which optically 

act as Fabry–Perot interferometers, as basic transducer in 

biosensing experiments. This is a very convenient choice 

since PSi monolayers are very easy to fabricate and to 

characterize. These devices have also a high sensitivity to 

changes in refractive index, up to 70 ppm [14]. By 

exploiting the optical features of PSi Fabry–Perot interferometers, 

we studied several bio-molecular interactions: 

DNA single strands with their complementary strands, 

Glutamine binding protein (GlnBP) from Escherichia coli 

with Glutamine; GlnBP with Gliadine [8,11,12]. The highquality 

optical response of PSi multilayers allows the 

fabrication of resonant optical devices and photonic crystals 

that are very attractive for sensing applications [15]. 

In biochemical sensors based on 1-D photonic bandgap 

(PBG) structures, usually there is both a reduced group 

velocity of light within the sample, which increases the 

interaction with the analyte and thus the sensitivity, and 

the presence of a resonant characteristic wavelength. In this 

case, sensors whose operation is based on the measurement 

of the shift of a very narrow resonance (for instance, a 

transmission or reflection peak having a full-width at halfmaximum 

of few tens of microns) are characterized by 

increased sensitivity and resolution. We have used a PSi 

optical microcavity to study the molecular binding of 

GlnBP and glutamine with a higher sensitivity with respect 

to the interferometric characterization [10]. 

Another typical PBG structure used in optical sensing is 

the distributed Bragg mirror which is obtained by 

alternating high (A) refractive index layers (low porosity) 

and low (B) refractive index layers (high porosity). The 

thicknesses of each layer satisfy the following relationship: 

nAdA+nBdB ¼ ml/2, where m is an integer and l is the 

Bragg resonant wavelength. The main optical feature of a 

Bragg mirror is the presence of a single, degenerate 

photonic band gap in a period of the reciprocal space. As 

a consequence, the reflectivity spectrum is characterized by 

a principal (m ¼ 1) stop band, centred around l, whose 

width depends on the refractive indexes contrast (nA/nB), 

while its height, depends on the number of the layers. The 

higher orders of the resonance (m ¼ 2, 3, 4,y) are 

characterized by a narrow width and a lower reflectivity 

[16]. 

We have designed, simulated and realized a low contrast, 

20-layer-Bragg mirror, with the first-order (m ¼ 1) resonant 

wavelength at 4800 nm. A highly doped p + -silicon, 

/100S oriented, 0.01 O cm resistivity, 400 mm thick was 

used as substrate to realize the structures. The Bragg low- 


I. Rendina et al. / Physica E 38 (2007) 188–192 189 

porosity layers (effective refractive index nLffi1.56, thickness 

dLffi641 nm) were produced by an etching current 

density of 125 mA/cm 2 for 2.58 s, while an etching current 

density of 150 mA/cm 2 for 3.62 s has been used for the 

high-porosity layers (n Hffi1.50, d Hffi933 nm). 

For the biosensing experiments, a strong base post-etch 

treatment was used to increase the pore size and remove the 

superficial nano-residue due to the electrochemical etch 

process, so to improve the infiltration of biomolecular 

probes into the pores [17]. Unfortunately, this process 

removes most of the native Si–H bonds, required by the 

subsequent functionalization treatment, from the porous 

silicon surface: to restore these bonds, we rinsed the porous 

silicon device in a very diluted HF-based solution for 30 s. 

The reflectivity spectra were measured over the range 

800–1600 nm with a resolution of 0.2 nm by using an Y 

optical fibre connected to a tungsten lamp, as a light 

source, and to an optical spectrum analyzer. 

We have estimated the refractive index sensitivity, i.e., 

the response of the sensor to the changes of the average 

refractive index, by measuring the peak shift of a higher 

order Bragg resonance on exposure to several organic 

volatile substances with different and increasing refractive 

indexes. Each measurement starts after that a small 

amount of volatile substance is added to reaction chamber 

and saturates with its vapours the atmosphere surrounding 

the sensor. The vapours penetrate into the nanometric 

pores, substituting the air. The average refractive index of 

the layers changes, so that the reflectivity spectrum of the 

device shifts towards higher wavelengths. 

Procedures for the immobilization of biomolecules on 

porous silicon are usually based on a chemistry that 

involves the silanization of the oxidized PSi surface [18]. A 

promising recently proposed alternative is that exploiting 

the reaction of acids molecules with the hydrogenterminated 

porous silicon surface in order to obtain a 

more stable organic layer covalently attached to the PSi 


0.6 

0.5 

0.4 

0.2 

0.1 

8 nm 

700 750 800 850 


Fig. 1. Reflectivity spectra of a porous silicon Bragg mirror.

190 

surface through Si–C bonds. We have exploited a photoactivated 

modification of PSi surface based on the UV 

exposure of an N-hydroxysuccinimide ester (UANHS) 

solution [19]. This kind of treatment is very fast since the 

time required for the complete surface passivation is of the 

order of few minutes. Moreover, after UV exposition the 

Red Shift (nm) 

200 

190 

180 

170 


Pentane 1.358 


Isobutanol 1.396 


Bragg sensitivity 

470±40 nm/RIU 

160 

1.32 1.34 1.36 

Refractive Index 

1.38 1.40 

Fig. 2. Red shift of the Bragg peak at 750 nm versus the refractive index of 

vapor substances. 

% Reflectivity 

140 

120 

100 

80 

60 

40 

20 

0 


I. Rendina et al. / Physica E 38 (2007) 188–192 

chip can be analysed by FT–IR microscopy to verify the 

state of the process. In view of a possible integration of 

the PSi-based biosensor into a lab-on-chip, we have verified 

the thermal stability of the Si–C bonds up to the 

temperature adopted in the packaging process of the chip. 

Once functionalized, the PSi optical transducer was bonded 

to a glass slide by an innovative anodic bonding process 

that takes into account the very peculiar characteristic of 

this material. Details of the process and of the microfluidic 

solutions adopted are elsewhere described [7]. 

Even if our aim is the realization of a label-free optical 

biosensor based on the PSi nanotechnology, we have used a 

fluorescent protein to control the distribution of the 

biological matter on the chip surface and to preliminary 

test the chemical stability of the covalent link between the 

protein and the PSi surface. 

3. Experimental results and discussion 

-20 

4000 3500 3000 2500 2000 1500 1000 

Fig. 1 shows the reflectivity spectrum of the PSi Bragg 

mirror in the range 700–850 nm: we have found that despite 

the high order of the resonance (m ¼ 6), the resonant peak 

at l ¼ 750 nm has a reflectivity which is the 60% of the first 

order and the width is less than 10 nm. Note that due to 

refractive index and thickness fabrication uncertainties the 

m ¼ 6 Bragg wavelength is about 750 nm instead of 

800 nm. 


Bragg before functionalization 

Bragg after UANHS treatment 

Fig. 3. FT–IR spectra of the realized porous silicon Bragg structure before and after the photoinduced functionalization process based on UV exposure, 

together with the reaction scheme.

In Fig. 2, the red shifts of the m ¼ 6 Bragg resonance 

peak at 750 nm are reported as function of the refractive 

index of each volatile substance employed. A well linear 

response is observed and a sensitivity of 470 (40) nm/RIU 

can be estimated, where RIU is the acronym for refractive 

index unit. Since the resolution of our spectrometer is 

0.05 nm, a limit of detection for the refractive index change 

of 1.06 10 4 can be estimated for this kind of optical 

device. 

The UV-induced surface passivation process results in 

covalent attachment of UANHS to the porous silicon 

surface, as clearly shown in the FT–IR spectrum reported 

in Fig. 3 together with the reaction scheme: all the 

characteristic absorption peaks due to the organic compounds 

are present in the sample after the functionalization 

process. The chip was also optically characterized: a red 

shift of about 54 nm in the reflectivity spectrum is observed. 

Reflectance % 

50 

40 

30 

20 

10 

2 micron 70% porosity with UANHS 

after treatment @250°C 2' 



Si-C 

C-N-C 

1100 1000 900 800 700 


Fig. 4. FT–IR spectra of a PSi chip after a thermal treatment at different 

times. 


I. Rendina et al. / Physica E 38 (2007) 188–192 191 

Then the device was incubated over night with the 

biological probe. The selective recognition of the biomolecular 

target is optically verified by a further red shift 

(data will be published elsewhere). 

The FT–IR spectrum of a PSi chip after functionalization 

at different temperatures is reported in Fig. 4: itis 

well evident that the link between the inorganic and 

organic phases is stable up to 250 1C. The Si–C and C–N–C 

bonds have high thermodynamic strength so that the 

passivated chip is still working after the anodic bonding 

process. 

In Fig. 5, a real view of the hybrid porous silicon-glass 

chip (A) is reported, while in the insets (B) and (C) the 

fluorescent images observed by a Leica Z16 APO 

fluorescence macroscopy system after incubation of the 

labelled bioprobes bonded to the chip, are shown. By 

illuminating the chip spotted with labelled proteins 

(fluorescent glutamine binding protein, purified by E. coli 

[8]) by a 100 W high-pressure mercury source, we found 

that the fluorescence is very high and homogeneous on the 

whole surface (see Fig. 5(B)). In Fig. 5(C) a detail of a 

scratch on the chip surface can be seen, showing that the 

fluorescent probes also penetrate into the inner pores of the 

chip. After several flushing run with demineralized water, 

injected by a micro-syringe, we found that the fluorescence 

is still bright. 


In this work, our recent experimental results in optical 

biosensing using PSi-based photonic devices are reviewed. 

Newest results about the fabrication and characterization 

of PSi sensors based on the exploitation of distributed 

Bragg reflectors operating at higher order harmonics are 

also presented. The PSi surface functionalization of such 

structures by UV stimulated processes, allowing the 

attachment of linkers through stable covalent Si–C bonds, 

is also demonstrated to be compatible to the hightemperature 

anodic bonding processes used for the 

realization of silicon-glass biochips. 

Fig. 5. Images of the basic element of a lab-on-chip based on a PSi optical transducer: (A) a real view of the porous hybrid silicon-glass chip; (B) 

fluorescent images of labelled proteins; (C) particular of a scratch on the chip surface.

192 


The authors are grateful to Prof. Mose` Rossi and Prof. 

Paolo Arcari for the helpful discussions. They also wish to 

thank Dr. Sabato D’Auria and Dr. Annalisa Lamberti for 

all the contributes given in preparing and making the 

experiments with the biological matter. 

References 

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Chapter 3 

Hybrid systems based on Porous Silicon 

and biocompatible materials

Introduction: PSi hybrid systems 

Much attention has been attracted in the field of porous silicon (PSi) for its diverse 

applications in bio- and chemical sensing [1–3]. Unique properties for theses applications 

include the significantly increased surface interaction area, simplicity and repeatability of 

fabrication, and compatibility with the well-established silicon microfabrication technology. 

One reason of this clear success is the easy fabrication of sophisticated optical 

multilayers, such as the one-dimensional photonic crystals, by a simple, but not trivial, 

electrochemical etch process, computer controlled [4, 5]. The reflectivity spectrum of the 

PSi based photonic crystals can show single or multiple resonances [6], band-gap and 

other characteristic shapes, depending on the spatial arrangement of the porous layers [7], 

which are very useful in lot of applications from biochemical sensing to medical imaging. 

But the major drawback preventing a wide diffusion of PSi devices is the instability of its 

native interface with a metastable Si–Hx termination. The metastable hydrosilicon can 

undergo spontaneous oxidation in ambient atmosphere and results in the degradation of 

surface structures. 

Moreover, it has been shown that porous silicon wafer can be dissolved under the 

physiological conditions which are very often used in biological experiments [8]. 

Therefore, surface passivation is crucial for the technological success of this material. 

Many organic molecules such as alkenes or alkynes had been covalently grafted onto the 

PSi surface by wet chemical approaches to passivate its surface [Chapter 2]. 

A very promising approach to overcome these difficulties is in fabricate hybrid systems 

based on the infiltration of polymers or biofilm resulting from the self-assembling of 

proteins in the nanometric pores of the PSi matrix. This is an alternative to the 

conventional chemical surface modification strategies. 

Recently, beside hydrocarbon monomer monolayers, polymer films, are of special interest 

because they provide a strong shield to protect the surface nanostructure. 

An example of hybrid PSi-polymer systems are the monodisperse microparticles 

containing a tunable nanostructure, these devices are desired for various high-throughput 

screening and targeted drug delivery applications [9,10]. The applicability of porous siliconbased 

particles and films has been demonstrated [11-21]. 

Moreover the field of tissue engineering entails the creation of a degradable three 

dimensional scaffold structure that has the appropriate physical, chemical, and mechanical 

properties to enable cell penetration and subsequent tissue formation for the purpose of 

disease or trauma repair at the desired site [22]. Bone engineering specifically attempts to 

replace or increase current graft approaches by using porous scaffolds that are designed 

to support the migration, proliferation, and differentiation of osteoprogenitor cells and aid in 

the organization of these cells in the proper dimensions. While these scaffolds may be 

made from a wide variety of both natural and synthetic materials, many of the existing 

porous biodegradable polymeric systems have been found to have limitations for use as 

orthopedic scaffolds [23]. 

Mesoporous Silicon, also referred to by its proprietary name BioSilicon , possesses 

several unique features highlighting its utility as a biomaterial. Previous studies by 

Bowditch et al have demonstrated its ability to be resorbed in soft tissue with a negligible 

inflammatory response in vivo [24]. When coupled with an established biopolymer such as 

polycaprolactone (PCL), it can be processed into a broad range of self-assembling 

structures useful to tissue engineering [25]. 

67

Schematic of a PSi device as etched and after polymer infiltration. 

As was pointed out in the previous chapter, biomolecules can be linked via Si–C bonds or 

Si–O. Covalent attachment provides a convenient means to link a biomolecular capture 

probe to the inner pore walls of PSi for biosensor applications. 

Hybrid materials, in which the functional material consists of an organic polymer or a 

biopolymer, forms an additional class of devices. Composites are attractive candidates for 

biological devices because they can display a combination of advantageous chemical and 

physical characteristics not exhibited by the individual constituents. Advances in polymer 

and materials chemistries have greatly expanded the design options for nanomaterial 

composites in the past few years, and synthesis of materials using nanostructured 

templates has emerged as a versatile technique to generate ordered nanostructures [26]. 

Templates consisting of micro or mesoporous membranes, zeolites, and crystalline 

colloidal arrays have been used, and many elaborate electronic, mechanical, or optical 

structures have resulted. 

In this chapter the modification of PSi structures by infiltration of chemical and biological 

polymers is describes. In particular, it was used a biocompatible multi-block copolymer 

based on PCL and a high stable and resistant biofilm resulting from the self-assembling of 

the hydrophobins (HFBs) [27-29]. 

The polymer contains pendant functional groups regularly spaced along the chain. The 

reactive groups are available for the attachment of chemical substances and their 

distribution, can be tuned according to the molecular weight of both segments. In this way, 

the physical and chemical properties of the porous silicon surface were changed and 

modulated. 

On the other hand HFBs are a family of small and cysteine-rich fungal proteins produced 

in the hyphal cell walls and that can be purified by the culture medium. The HFBs selfassemble 

into thin and amphiphilic membranes at hydrophilic-hydrophobic interfaces. In 

this way the wettability of a surface can be controlled and the hydrophobic behaviour can 

be converted in hydrophilic, and vice versa [30-32]. 

A general classification of HFBs is made on the basis of the biofilm resistance: the class I 

HFBs form high insoluble assemblies, which can be dissolved in strong acids, whereas 

class II biofilms can be dissolved in ethanol or in sodium dodecyl sulphate [27]. 

Solutions with HFB or PCL have been used at different concentration and the infiltration 

process has been monitored by variable angle spectroscopic ellipsometry (VASE) 

measurements. 

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3. S. Le´tant, M.J. Sailor, Adv. Mater. 12 (2000) 355. 

4. F. Müller, A. Birner, U. Gösele, V. Lehmann, S. Ottow, H. Föll, Journal of Porous Materials 2000, 

7, 201. 

5. H. F. Arrand, T.M. Benson, A. Loni, R. Arens-Fischer, M. Kruger, M. Thonissen, H. Luth, S. 

Kershaw, IEEE Phot. Techn. Lett. 1998, 10, 1467. 

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9. R. Duncan, Natl. Rev. Drug Discov. 126, 1470 (2003). 

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12. V. S.-Y. Lin, K. Motesharei, K. S. Dancil, M. J. Sailor, and M. R. Ghadiri, Science 278, 840 

(1997). 

13. K.-P. S. Dancil, D. P. Greiner, and M. J. Sailor, J. Am. Chem. Soc. 121, 7925 (1999). 

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15. S. Chan, P. M. Fauchet, Y. Li, L. J. Rothberg, and B. L. Miller, phys. stat. sol. (a) 182, 541 

(2000). 

16. S. Chan, S. R. Horner, B. L. Miller, and P. M. Fauchet, J. Am. Chem. Soc. 123, 797 (2001). 

17. T. A. Schmedake, F. Cunin, J. R. Link, and M. J. Sailor, Adv. Mater. 14, 1270 (2002). 

18. F. Cunin, T. A. Schmedake, J. R. Link, Y. Y. Li, J. Koh, S. N. Bhatia, and M. J. Sailor, Nature 

Mater. 1, 39 (2002). 

19. S. O. Meade, M. S. Yoon, K. H. Ahn, and M. J. Sailor, Adv. Mater. 16, 1811 (2004). 

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21. L. T. Chanham, M. P. Stewart, J. M. Buriak, C. L. Reeves, M. Anderson, E. K. Squire, P. Allcock, 

and P. A. Snow, phys. stat. sol. (a) 182, 521 (2000). 

22. Karp, J.M.; Dalton, P.D.; Shoichet, M.S., Mater. Res. Bull. 2003, 28,301. 

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T. Mater. Res. Soc. Symp. Proc.,1999, 536,149. 

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Canham, L.. Phys. Stat. Sol (a), 2005, 202,1451. 

26. K.S. Soppimath, T.M. Aminabhavi, A.R. Kulkarni, W.E. Rudzinski, Biodegradable polymeric 

nanoparticles as drug delivery devices, J. Control Release 70 (2001) 1–20. 

27. H. J. Hektor, K. Scholtmeijer, Current Opinion in Biotechnology 2005, 16, 434. 

28. H.A.B. Wosten, M. L. de Vocht, Biochimica and Biophysica Acta 2000, 1469, 79. 

29. G.R. Szilvay, A. Paananen, K. Laurikainen, E. Vuorimaa, H. Lemmetyinen, J. Peltonen, M.B. 

Linder, Biochemistry 2007, 46, 2345. 

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69

Porous silicon-polymer composite matrix for an innovative 

class of optical biosensors 

This section is aimed at the characterization of PSi surface modification, based on the 

infiltration of Poly(ε-caprolactone) (PCL) macro-monomers in the nanometric pores of its 

matrix. 

In collaboration with Prof. R. Palumbo of University of Naples “Federico II”, Italy

A polymer modified PSi optical device for biochemical sensing

Introduction 

Due to their widespread utilisation, there are all over the world tens of thousand of 

hydrocarbons contaminated sites, where the pollution is caused by leaking of storage 

tanks, by chemical wastes dumps and also by extraction and transport operations. In 

addition, there is the real problem of vapours leaks from distributed gas pipelines in 

metropolitan areas. Therefore, an accurate characterization and long-term monitoring is 

required to reduce and control health risk and public safety. Four general groups of 

chemical sensors can be categorized on the base of physics and operating mechanisms: 

chromatography and spectrometry, electrochemical sensors, mass sensors and optical 

sensors [1]. Although a number of these devices are commercially available, there is a 

strong need of low-cost, robust and on field sensors of chemical species, especially for 

long, real-time monitoring [1]. 

In particular, PSi optical sensor based on reflectivity measurements in Bragg grating 

mirrors [2] or on photoluminescence experiments in optical microcavities operating in the 

visible wavelength region [3, 4] have been already reported. 

A typical photonic band gap (PBG) structure used in optical sensing is the distributed 

Bragg mirror which is obtained by alternating high (A) refractive index layers (low-porosity) 

and low (B) refractive index layers (high-porosity). The thicknesses of each layer satisfy 

the following relationship: nAdA+nBdB=m λ/2, where m is an integer and λ is the Bragg 

resonant wavelength. 

When the PSi Bragg mirror is exposed to vapors, or dip in liquid hydrocarbons, a 

repeatable and completely reversible change in the reflectivity spectrum is observed. In 

fact, the substitution of air in the pores by the organic molecules causes an increasing of 

the average refractive index of the microcavity, shifting towards longer wavelengths its 

characteristic peaks. 

• PSi functionalization by polymer infiltration 

Covalent attachment provides a convenient means to link a biomolecular capture probe to 

the inner pore walls of PSi for biosensor applications. As was pointed out in the previous 

section, biomolecules can be linked via Si–C bonds or Si–O. 

Advances in polymer [5] and materials chemistries have greatly expanded the design 

options for nanomaterial composites in the past few years, and synthesis of materials 

using nanostructured templates has emerged as a versatile technique to generate ordered 

nanostructures. Templates consisting of micro or mesoporous membranes, zeolites, and 

crystalline colloidal arrays have been used, and many elaborate electronic, mechanical, or 

optical structures have resulted. 


Bragg multilayers were constituted by 15 couples of 58 % and 69 % porosities and 72 nm 

and 117 nm thicknesses, respectively. 

A highly doped p+-silicon, oriented, 0.01 Ω cm resistivity, 400 μm thick was used as 

substrate to realize the structures. The Bragg low-porosity layers (effective refractive index 

nL≅1.903) have been produced by an etching current density of 100 mA/cm 

73 

2 for 2.2 s, 

while an etching current density of 200 mA/cm 2 for 2.3 s has been used for the highporosity 

layers (nH ≅ 1.632). 

The electrochemical dissolution process leaves nanodimensional residues into the 

channels which partially fill the pores in the layers and prevent a homogeneous diffusion of 

biological solutions. To remove the nanostructures and improve pore infiltration of polymer, 

a KOH post-etch process elsewhere described [6] was used. It was verified that the 

alkaline treatment does not degrade the optical response of our devices. In order to leave 

an hydrophobic surface the Si-H bonds were replaced on the surface, by immersing the 

samples in a very diluted HF solution for few seconds.

The polymer solutions were prepared by solving the polymer in dichloromethane (CH2Cl2). 

A solution volume of 100 μl was directly spotted on the PSi surface, spun at 2000 rpm for 5 

seconds, and baked at 100°C for 5 minutes. 

In Figure 1 the red shift induced in the Bragg reflectivity spectrum by the presence of PCL 

is reported. 


1.0 

0.8 

0.6 

0.4 

0.2 

0.0 

800 1000 1200 


Bragg pre PCL 

post PCL 

Figure 1: PSi reflectivity spectra before and after PCL infiltration. 

• PSi/PCL hybrid system: alkaline stability 

In order to test the new properties of the PSi-polymer composite device, two twins chips, 

one polymer infiltrated and the other bare, were exposed to a 0.1 M aqueous solution of 

sodium hydroxide. The dissolution time evolution of both the samples were compared: 

initially both undergo a shift through minor wavelength, due to the removal by NaOH 

dissolution of some unprotected silicon nanocrystals, but while after 80 s the uncoated 

sample is almost completely dissolved, while the PCL coated PSi Bragg still resist for 15 

minutes. In Figure 2 are reported the optical spectra of both samples for direct 

comparison. 



1 

0 

1.5 

1.0 

0.5 

PSi/PCL 

after 15 min in NaOH 

600 800 1000 1200 1400 1600 

PSi 

after 80s in NaOH 

0.0 

600 800 1000 1200 1400 1600 


Figure 2: Comparison between a coated (A) and an uncoated (B) PSi Bragg mirrors before (black) and after 

immersion in 0.1 M NaOH solution (red). 

Even if blue-shifted the optical spectrum (A) yet shows a reflectivity stop band. 

• PSi/PCL hybrid system: chemical sensing 

After the basic etch process the hybrid organic-inorganic device is still working as chemical 

optical transducer: it was tested its ability in sensing the vapours of different volatile 

substances. In Figure 3-A, are reported the characteristic red-shift due to the capillary 

condensation of the vapours inside the nanometric pores of the PSi Bragg. The shift is 

completely reversible and the polymer coated device can be used with different 

substances (Isopropanol, ethanol and methanol) giving high reproducible results. 

74

Organic compound Refractive index 


Ethanol 1.361 


It was also calculated the sensitivity of this optical transducer to the refractive index 

changes by exposing it to substances having different refractive index. Assuming that the 

three solvents equally penetrates the nanostructured spongy multilayer, was estimated a 

sensitivity of 585 (3) nm expressed in refractive index units [7]. 


1.0 

0.8 

0.6 

0.4 

0.2 

0.0 

600 800 1000 1200 1400 1600 


A 

unperturbed Bragg 

Methanol 

Isopropanol 

Ethanol 

Δλ (nm) 

90 

80 

70 

60 


S=585 (3) nm/RIU 

1.32 1.33 1.34 1.35 1.36 1.37 1.38 

Refractive index 

B 

Figure 3: The PSi polymer modified Bragg still works as optical transducer for vapour and liquid detection. A: 

the characteristic red shift of the optical spectrum on exposure to methanol, isopropanol and ethanol. B: 

determination of the sensitivity to refractive index changes of the optical transducer. 

It was demonstrated that the passivation of some optical devices based on the porous 

silicon technology when infiltrated by polymer not only protects the nanocrystalline material 

from basic dissolution in NaOH, but also leaves unaltered the sensing ability of such 

optical transducers adding the chemical stability which can be a key feature in 

biomolecular experiments. 

It was moreover experimentally studied a feasible optical sensor for hydrocarbons 

detecting in the gaseous state. The optical experimental setup adopted and the structure 

of the sensing element do not prevent from the scaling down of each component in a 

compact sensor system with minimum effort. In fact, the white source can be replaced by 

small, wide bandwidth LEDs and the OSA bench version by an integrated fiber optic 

spectrometer, already available on the market. Moreover, the use of a resonant device 

allows “on/off” digital reading of the sensor element by a single-wavelength source. A net 

red-shift of the transmittance peak in the reflectivity spectrum has been repeatable 

observed when the sensor is exposed to the hydrocarbons species. The shift values, due 

to capillary condensation of the liquid into the pores, are characteristic of each species and 

depend on their physical and chemical properties. The sensing process is completely 

reversible. 

• PSi/PCL hybrid system:biochemical sensing 

The hybrid interface of silicon and PCL should act as an active substrate for covalently 

binding of different bioprobes, so that the quality and the characteristic of this polymer 

layer are key features for the entire device. This represents an alternative to the 

conventional chemical surface modification strategies. The reactive groups, introduced by 

infiltration of biocompatible polymer into PSi matrix, are available for the attachment of 

chemical substances and biological probes, in order to verify this feature a methanol 

75

solution of 5(6)-Carboxyfluorescein N-hydroxysuccinimide ester (FITC) 300 μM was used 

as ligand. A schematic of immobilization chemistry is reported in figure 4. 

A B C 

Figure 4: The PSi polymer modified device. A: the characteristic reactive groups on the PSi surface after spin 

coating. B: 5(6)-Carboxyfluorescein N-hydroxysuccinimide ester (FITC) molecular formula. C: Fluorescence 

image after incubation. 

The immobilization was carried in methanol solution at room temperature over night. After 

incubation, the sample was rinsed three times in methanol. By illuminating the fluorescent 

FITC samples, a bright fluorescence signal was recorded, and it was quite homogeneous 

on the whole surface. It was also qualitatively tested the strength of the binding between 

the PCL and the FITC, by washing the chip in a dialysis membrane overnight in methanol. 

Since the fluorescent intensities before and after the dialysis do not differ, can be 

concluded that the PCL-FITC system is very stable. 

References 

1. C. K. Ho, M. T. Itamura, M. Kelley and R. G. Hughes, Sandia Report SAND2001-0643, Sandia 

National Laboratories, (2001) 1-27. 

2. P. A. Snow, E. K. Squire, P. St. J. Russel, and L. T. Canham, J. Appl. Phys. 86, (1999) 1781- 

1784. 

3. V. Mulloni and L. Pavesi, Appl. Phys. Lett. 76, (2000) 2523-2525. 

4. M. Ben-Chorin, A. Kux, and I. Schechter, Appl. Phys. Lett. 64, (1994) 481-483. 

5. P. A. Snow, E. K. Squire, P. St. J. Russel, and L. T. Canham, J. Appl. Phys. 86, (1999) 1781- 

1784. 

6. L. A. De Louise and B. L. Miller (2004), Mat. Res. Soc. Symp. Proc. Vol. 782, A5.3.1- A5.3.7. 

7. L. De Stefano, I. Rendina, L. Moretti, A.M. Rossi, Materials Science and Engineering B 2003, 

100/3, 271. 

76

Bio/NON Bio Interfaces for a new class of proteins Microarrays 

This section is aimed at the characterization of PSi surface biological passivation, based 

on the infiltration by the fungal proteins called hydrophobins. The self-assembled biofilm 

changes the wettability of the PSi structures and protects the PSi structures from alcaline 

agent. Moreover these is a new active substrate as a functional one for proteins 

microarrays development. 

In collaboration with Prof.ssa P. Giardina of University of Naples “Federico II”, Italy

5 

CREATED USING THE RSC ARTICLE TEMPLATE (VER. 3.1) - SEE WWW.RSC.ORG/ELECTRONICFILES FOR DETAILS 

ARTICLE TYPE www.rsc.org/xxxxxx | XXXXXXXX 

Bioactive Modification of Silicon Surface using Self-assembled 

Hydrophobins from Pleurotus ostreatus. 

L. De Stefano* a , I. Rea a,b , E. De Tommasi a , I. Rendina a , L. Rotiroti a,c , M. Giocondo d , S. Longobardi e , A. 

Armenante e , and P. Giardina e . 

Received (in XXX, XXX) Xth XXXXXXXXX 200X, Accepted Xth XXXXXXXXX 200X 

First published on the web Xth XXXXXXXXX 200X 

DOI: 10.1039/b000000x 

A crystalline silicon surface can be made biocompatible and chemically stable by a self-assembled 

biofilm of proteins, the hydrophobins (HFBs) purified from the fungus Pleurotus ostreatus. The 

10 protein modified silicon surface shows an improvement in wettability and is suitable for proteins 

immobilization. Two different proteins were successfully immobilized on the HFBs coated chips: 

the bovine serum albumin and an enzyme, a laccase, which retains its catalytic activity even when 

bound on the chip. Variable angle spectroscopic ellipsometry (VASE), water contact angle (WCA), 

and fluorescence measurements demonstrated that the proposed approach in silicon surface 

15 bioactivation is a feasible strategy for the fabrication of a new class of hybrid biodevices. 

20 

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Introduction 

Biological molecules are more and more used in a large class 

of research and commercial applications such as biosensors, 

DNA and protein microarrays, cell culturing, immunological 

assays, and so on. The fabrication of a new generation of 

hybrid biodevices, where a biological counterpart is integrated 

in a micro or a nano-electronic platform, is strictly related to 

the bio-compatibilization treatments of the surfaces involved. 

1 

The design ad the realization of bio/non-bio interfaces with 

specific properties of chemical stability, wettability, and 

biomolecules immobilization are key features in the 

miniaturization of devices needed in the development of 

genomic and proteomic research. 2 In particular, protein 

immobilization is a hot topic in biotechnology since 

commercial solutions such as the DNA array are not still 

available. Proteins are, due to their composition, a class of 

very heterogeneous macromolecules with variable properties. 

For these reasons, it is extremely complex to find a common 

surface suitable for different proteins with a broad range in 

molecular weight and physical–chemical properties such as 

charge and hydrophobicity. A further aspect is the orientation 

of spotted compounds, that becomes of crucial relevance for 

some applications, like quantitative analysis, enzymatic 

reactions, interaction studies. Besides tethering the proteins to 

the surface by adhesion or non-oriented covalent attachment, 

recent developments for the directed immobilization of 

proteins are emerging. These efforts are addressing challenges 

such as loss of enzymatic activity due to unfavourable 

orientation of the immobilized enzyme. 3 Despite the 

development of many different surfaces in the last five years, 

notably only few systematic investigations have been 

conducted and yet, no universal surface ideal for all 

applications could be identified 4-8 . Among others, silicon is a 

very interesting platform due to its wide use in all the micro 

and nanotechnologies developed for the integrated circuits 

industry. Lot of studies about the chemical functionalization 

of silicon in order to make it compatible with the organic 

chemistry have been published in the last ten year. 9 A 

completely different approach is to use a biological substrate 

to passivate the silicon surface: recently, a nanostructured 

self-assembled biofilm of amphiphilic proteins, the 

hydrophobins, was deposited by solution deposition on 

crystalline silicon and proved to be efficient as masking 

material in the KOH wet etch of the crystalline silicon. 10 

HFBs are a group of very surface active proteins. 11 They are 

small (around 10 kDa) proteins that originate from 

filamentous fungi, where they coat the spores and aerial 

structures, and they mediate the attachment of fungal 

structures to hydrophobic surfaces and affect the cell wall 

composition. 12-13 HFBs self-assemble at the air/water interface 

and lower the surface tension of water. Furthermore, they 

shown self-assembling properties also at interfaces between 

oil and water, and between water and a hydrophobic solid. 14 

The primary structure of HFBs is characterized by a 

conserved pattern of eight cysteine residues, forming four 

intramolecular disulfide bridges. 15 HFBs are further divided 

into classes I and II based on their hydropathy patterns, 

although the amino acid sequence similarity both within and 

between the classes is small. Class I HB assemblies are 

remarkable due to their insolubility. They are insoluble even 

in hot solutions of sodium dodecyl sulfate (SDS) and can only 

be dissolved in some strong acids, such as trifluoroacetic 

acid. 11 

Some very recent applications of class II hydrophobins in 

protein immobilization have been reported, nevertheless only 

one example of enzyme immobilization on a Class I 

hydrophobin (SC3) layer formed on a glassy carbon electrode 

has been investigated. 12 Inspired by these very promising 

results, we have studied the immobilization of a fluorescent 

protein and a laccase enzyme on a Class I hydrophobin layer 

when self-assembled on silicon surface. 

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Matherials and methods

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Protein purification. The HFB and the laccase POXC were 

isolated from the culture broth of P. ostreatus (type: Florida 

ATCC no. MYA-2306). For HFBs production, mycelia were 

grown at 28 °C in static cultures in 2 L flasks containing 500 

mL of potato dextrose (24 g/l) broth with 0.5% yeast extract. 

After 10 days of fungal growth, the HFBs aggregates were 

obtained by air bubbling of the culture broth by using a 

Waring blender. Foam was then collected by centrifugation at 

4000 g. The precipitate was freeze dried, treated with 

trifluoroacetic acid (TFA 100%) for 2h, and sonicated for 30 

min. The sample was dried again in a stream of air and then 

dissolved in 60% ethanol. The precipitate formed was 

separated by centrifugation at 4000g. the purity of the HFB 

was ascertained by SDS-PAGE, by using a silver staining 

method. Mycelium growth and laccase purification were 

carried out following the procedures described by Palmieri et 

al.. 16 

Hydrophobin self-assembly on silicon chip. Silicon samples, 

single side polished, oriented (chip size: 1cm×1cm), 

after standard cleaning procedure, 18 have been washed in 

hydrofluoridic acid solution for three minutes to remove the 

native oxide thin layer (1-2 nm) due to silicon oxidation. Then 

samples have been incubated in HFBs solutions (0.1 mg/mL 

in 80% ethanol) for 1 h, dried for 10 min on the hot plate 

(80°C), and then washed by the solvent solution (80% ethanol 

in water). 

Laccase immobilization method. The laccase immobilization 

procedure on HFB modified silicon surface was carried at 4 

°C for 1 hour, followed by a over night rinsing in 50 mM 

phosphate buffer, pH 7. After 16 hours, the silicon biochip 

was washed in the same buffer at 25°C till to no laccase 

activity was found in the washing buffer. Then each chip was 

dust-protected and kept at constant humidity conditions when 

not in use. 

BSA immobilization method. Bovine serum albumin (BSA) 

was labelled by rhodamine, following the Molecular Probes © 

procedure MP06161, and solubilised in water at three 

different concentration (3, 6, 12 μM). To assess the protein 

binding on the chip surface, we have spotted on the chips 

covered by the HFBs biofilm 50 μl of water solution 

containing the labeled protein. The immobilization was 

carried in water solution at 4°C over night. After incubation, 

the samples were rinsed three times in deionized water at 

room temperature. 

Enzyme assay Laccase activity was assayed at 25°C using 2,6dimethoxyphenol 

(DMP) 10mM in McIlvaine citrate– 

phosphate buffer, pH 5. Oxidation of DMP was followed by 

the absorption increment at 477 nm (ε477=14.8×10 3 M -1 cm -1 ) 

using Beckman DU 7500 spectrophotometer (Beckman 

Instruments). Enzyme activity was expressed in International 

Units (IU). Immobilized enzyme was assayed by silicon dip in 

10 ml of McIlvaine citrate–phosphate buffer, pH 5. The 

absorption increment at 477 nm was followed withdrawing 

200 μl of reaction mixture each 30s for 10 min. The total 

immobilised activity per unit of silicon (chip) were calculated 

as U/chip = ΔA min -1 /ε × 10 4 . 

Optical Techniques. Ellipsometric characterization of the 

hydrophobin biofilm deposited on the silicon substrate has 

been performed by a variable-angle spectroscopic 

ellipsometry model (UVISEL, Horiba-Jobin-Yvon). 

Ellipsometric parameters Δ and Φ have been measured at an 

angle of incidence of 65° over the range of 360-1600 nm with 

a resolution of 5 nm. Fluorescence images were recorded by a 

Leica Z16 APO fluorescence macroscopy system. Contact 

angle measurements have been performed by using a KSV 

Instruments LTD CAM 200 Optical Contact Angle Meter: 

each contact angle has been calculated as the average between 

the values obtained from three drops having the same volume 

(about 3 μL), spotted on different points of the chip.You can 

also put lists into the text. 

2 | Journal Name, [year], [vol], 00–00 This journal is © The Royal Society of Chemistry [year] 

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The immobilization of proteins and enzymes on the silicon 

modified surface is schematized in Figure 1. 

HFB 

BSA 

POXC 

cSi cSi 

Figure 1. Schematic of immobilization of enzyme (POXC) 

and protein (rhodamine fluorescent BSA) on a crystalline 

silicon wafer biologically modified by a self-assembled 

biofilm of hydrophobin. 

The hybrid interface of silicon and HFB should act as an 

active substrate for non-covalently binding of different 

bioprobes, so that the quality and the characteristic of this 

proteins layer are key features for the entire device. The 

coating of a solid surface by chemical solution deposition is 

not a finely controlled process, even if the covering substance 

has self-assembling properties, as it is the case of HFBs. The 

result is that the HFB biofilms self-assembled on the silicon 

surface by deposition of equal concentration solutions could 

have thicknesses ranging between 10 and 30 nanometers. This 

behaviour could be ascribed to different local aggregation of 

HFB in the solution covering the silicon chip, due to the 

strong hydrophobic interactions among the proteins. In fact, 

during HFB deposition, solvent evaporation can determine 

local increases of protein concentration, thus favouring 

multimers and aggregate formation that can deposit on the 

layer. In this step there is no way to control the clustering and 

stacking of the proteins which are constituting the selfassembled 

biofilm. The association of hydrophobins in 

solution and their self assembly have both been attributed to 

the amphiphilic structure of the HFB monomer. 19-20 Before 

using the HFB biofilm as a nanostructured template for 

protein binding, we have tested its stability and some features 

by optical methods such as VASE, spectrofometry, and WCA.

5 

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25 

30 

35 

The HFB biofilm self-assembled on the hydrophobic silicon 

shows a strong adhesion to the surface: the protein layers 

always exhibit great stability to alkaline solutions even at high 

temperature. After 10 minutes washing in sodium hydroxide 

and SDS at 100 °C, a residual biofilm, which has on average a 

thickness of about 3.1±0.7 nm, is still present. The biofilm 

characterization has been performed by VASE on 12 samples 

realized in the same experimental conditions; the optical 

model for experimental data fitting has been developed in our 

previous work. 10 We believe that this is the thickness of a 

monolayer of HFBs when self-assembled on hydrophobic 

silicon: this value is also consistent with a typical molecular 

size and comparable to atomic force microscopy 

measurements. 21 According to the above described model, the 

washing step of the chip is strong enough to remove the 

aggregates deposited on the HFB monolayer that directly 

interacts with the hydrophobic silicon surface. This behaviour 

points out the stronger interactions between the silicon surface 

and the HFB monolayer with respect to those between the 

HFB aggregates and the HFB monolayer. 

In a recent article, we have already demonstrated that the 

nanometric HFB biofilm uniformly covers the silicon surface: 

10 

on exposure to the potassium hydroxide (KOH), which 

anisotropically dissolves the silicon, a silicon wafer coated by 

HFB is completely shielded by the etching agent, at least for 

several minutes. The protection against KOH can be 

prolonged at any time by increasing the thickness of the HFB 

biofilm self-assembled on the silicon surface. This is possible 

since successive depositions of proteins form a biofilm having 

thickness proportional to the number of depositions, as it is 

shown in Figure 2: we have reported the thicknesses of the 

HFB films after each one of three consecutive depositions and 

plotted them as function of the deposition cycle. 

Figure 2. Growth of a thick HFB biofilm by repeated cycles 

of deposition on a hydrophobic silicon substrate 

Moreover, the presence of the biofilms does not alter the 

optical behaviour of the silicon surface as it is shown in 

40 Figure 3, where the transmission curves of two different 

biofilms on silicon are reported. In both cases the selfassembled 

layer is transparent in a very large interval of 

wavelengths: the transmittance is still the 80 % at 290 nm. 

The improvement in the silicon wettability is also a key 

45 features in biodevices realization: the silicon surface, after 

standard cleaning process and washing in HF, is hydrophobic, 

this occurrence can constitute a severe limit in the fabrication 

of microchannels for microfluidics applications. Since HFBs 

self-assemble in stable biofilm on a variety of different 

surfaces, such as mica, Teflon, and polystyrene 6-7 due to the 

amphiphilic interaction with hydrophobic or hydrophilic 

surfaces, we can modify the silicon behaviour with respect to 

the water interaction by covering its surface by HFBs. 


50 

Figure 3. Figure caption. Transmission curves of two 

55 different samples of HFB self-assembled biofilm on silicon. 

60 

65 

The changing in silicon wettability has been monitored by 

water contact angle measurements. In Figure 4 A and B are 

reported the WCA results in case of the bare silicon surface 

and after the deposition of the HFB biofilm. The dramatic 

increase in wettability of the silicon surface is well evident: in 

the first case, the WCA results in 90°±1°, so that the surface 

can be classified as hydrophobic, while after the HFB 

deposition the WCA falls down to 25°±2° so that the surface 

is clearly hydrophilic. The HFB modified silicon surface is 

highly stable and the WCA value does not change after 

months, even if stored at room temperature without particular 

saving condition. 

70 Figure 4. Changing the wettability of c-Si: the hydrophobin 

nanolayer turns the hydrophobic surface of c-Si into a 

hydrophilic one. A: a water drop on the c-Si after washing in 

hydrofluoridic acid forms a contact angle of 90°±1°. B: a 

water drop after the hydrophobin deposition forms a contact 

75 angle of 25°±2°. 

80 

Transmittance (%) 

90 

85 

80 

75 

70 

22 nm biofilm 

64 nm biofilm 

400 600 800 1000 1200 1400 


A B 

The HFB modified silicon surface has been tested by BSA 

immobilization: solutions containing the rhodamine labelled 

protein, at different concentration, between 3 and 12 μM, have 

been spotted on the HFB film. The labelled bioprobes have 

been spotted also on some surface which have not been treated 

with the hydrophobins, as a negative control on the possible

5 

10 

15 

20 

25 

30 

35 

aspecific binding between the silicon and the probe. All the 

samples have been washed in deionized water to remove the 

excess of biological matter and observed by the fluorescence 

macroscopy system. 

A 

B 

Fluorescence intensity 

fluorescence intensity (a.u.) 

18 

16 

14 

12 

10 

8 

6 

4 

2 

0 

4,8 

4,0 

3,2 

2,4 

1,6 

0,8 

0,0 

0 5 10 15 20 

Distance (μm) 

[BSA]=300 fmol/cm 2 

0 2 4 6 8 10 12 

BSA concentration (μM) 

Figure 5. A: Fluorescence image of labelled BSA spotted on a 

silicon chip covered by HFBs. B: Negative control sample: 

BSA spotted directly on silicon. C: Intensity profile of 

fluorescencent BSA on HFB biofilm inside and outside the 

drop. D: The dose-response curve of fluorescence intensity as 

function of the BSA concentration. 

By illuminating the fluorescent BSA samples, we found that 

the fluorescence is brighter than the negative control as it can 

be seen in Figure 5A and 5B. The fluorescence signal is also 

quite homogeneous on the whole surface, as confirmed by the 

intensity profile on a treated surface reported in Figure 5C. 

We have qualitatively tested the strength of the affinity bond 

between the HFB and the BSA, by washing the chip in a 

dialysis membrane overnight in deionized water. Since the 

fluorescent intensities before and after the dialysis do not 

differ, we can conclude that the HFB-BSA system is very 

stable. We have also realized a dose-response curve of 

fluorescence intensity as function of BSA concentration, 

shown in Figure 5D, and we have estimated a saturation 

concentration equal to 3.12 μM corresponding to 300 

fmol/cm 2 . 

Protein immobilization on the HFB film has been also verified 

and analyzed using an enzyme, POXC laccase, produced by 

the same fungus P. ostreatus 16 , easily available in our 

laboratory and whose activity assay is simple and sensitive. 

The enzymatic assay on the immobilized enzyme has been 

performed dipping the chip into the buffer containing the 

substrate and following its absorbance during several minutes. 

We chose to use a pH 5 buffer and DMP as a substrate, a good 

settlement between stability and activity of the enzyme. A 

volume of 30 µl of the enzyme solution (about 700 U/ml) 

D 

C 

have been deposited on the chips (10mm x 10mm) and, after 

several washing, an activity of 0.1÷0.2 U has been determined 

on each chip (Figure 6), with an immobilization yield of 

0.5÷1%. 

4 | Journal Name, [year], [vol], 00–00 This journal is © The Royal Society of Chemistry [year] 

40 

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55 

60 

Figure 6. Enzymatic activity determination of POXC laccase 

immobilized on a chip. 

This value is comparable to that one (7%) obtained in the 

optimized conditions for laccase immobilization on 

EUPERGIT C 250L © . 22 Taking into account the specific 

activity of the free enzyme (430 Umg -1 ) and its molecular 

mass (59 kDa), 0.5µg of laccase corresponding to about 8 

pmol (5 x 10 12 molecules) have been immobilized on each 

chip. A reasonable evaluation of the surface occupied by a 

single protein molecule can be based on crystal structures of 

laccases 23-24 and homology among these proteins. This surface 

should be 28 x 10 -12 mm 2 , considering the protein as a sphere 

with radius of 3 x 10 -6 mm. On this basis the maximal number 

of laccase molecule on each chip should be 3 x 10 12 . These 

data can indicate that the number of active immobilized 

laccase molecules on each chip is of the same order of 

magnitude than the expected maximum. Laccase assays have 

been repeated on the same chip after 24 and 48 hours in the 

same conditions. Figure 7 shows that about one half of the 

activity was lost after one day, but no variation of the residual 

activity was observed after the second day. 

ΔA/min 

Ads (477 nm) 

0.32 

0.30 

0.28 

0.26 

0.24 

0.22 

0.20 

0.18 

0.16 

0.7 

0.6 

0.5 

0.4 

0.3 

0.2 

0.1 

0.0 

-0.1 

ΔA/min=0,317± 0,006 

0.0 0.5 1.0 1.5 2.0 

Time (min) 

0 10 20 30 40 50 

Time (h) 

Figure 7. Time resolved profile of laccase activity of the 

65 immobilized enzyme.

5 

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20 

25 

Conclusions 

We have demonstrated the bio-modification of the silicon 

surface by using the self-assembling features of the HFB, an 

amphiphilic protein purified by the fungus P. ostreatus. The 

nanometric biofilm of proteins, casted on the silicon surface 

by solution deposition, is very stable from the chemical point 

of view, since it is still present after strong denaturing 

washing in NaOH and SDS at 100 °C. The biofilm turns the 

hydrophobic silicon surface into a hydrophilic one, which can 

be very useful in microfluidic integrated application. 

Moreover, the biofilm can arrange both in a monolayer or in a 

multilayer stack of proteins depending on deposition 

conditions. The HFB monolayer acts as a bioactive substrate 

to bind other proteins, such as BSA and laccase. These results 

can be the starting point in the fabrication of a new generation 

of hybrid devices for proteomic applications. 

ACKNOWLEDGMENT. The authors gratefully acknowledge 

dr. N. Malagnino of ST. Microelectronics in Portici (NA), 

Italy, for water contact angle measurements, Prof. P. Arcari 

(Dept. of Biochemistry and Medical Biotechnologies, 

University of Naples “Federico II”) for fluorescent BSA, dr. 

F. Autore for the purification of POXC laccase and G. 

Armenante for technical support. 

Notes and references 

a 

Unit of Naples-Institute for Microelectronics and Microsystems, 

National Council of Research, Via P. Castellino 111, 80131, Naples, Italy 

Fax: +390816132598; Tel: +390816132375; E-mail: 

luca.destefano@na.imm.cnr.it 

b 

30 Dept. of Physical Sciences, University of Naples “Federico II”, Via 

Cintia 4, 80126, Naples, Italy. 

c 

Dept. of Organic Chemistry and Biochemistry, University of Naples 

“Federico II”, Via Cintia 4, 80126, Naples, Italy. 

d 

LYCRIL – INFM-CNR, Via P. Bucci, Cubo 33/B, 87036 Arcavacata di 

35 Rende, Cosenza, Italy. 

40 

45 

50 

55 

60 

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Chapter 4 

…towards the Lab-on-Chip

Introduction to Lab-on-chip technologies 

A heart attack is a typical situation for which a fast diagnostic is of paramount importance. 

In addition to an electrocardiogram, a blood test is often necessary, that is conventionally 

done in a laboratory. With the portable system marked by the Biosite company, this 

diagnostic can be made in only 10 min, directly with the patience [1]. The drop of blood is 

simply deposited in a reservoir in the system that then performs the entire test: minute 

quantities of blood are displaced and injected into several tiny chambers in which different 

reactions take place. 

This example clearly illustrates the principle of a Lab-on-Chip: integrating all the functions 

of a human-scale test laboratory including transferring samples, drawing off a precise 

volume of a chemical product, mixing with reagents, heating, titration, etc, on a system of a 

few square centimeters. 

A LOC includes many functions depending on the application such as pumps, valves, 

sensors, electronics, etc. Therefore, it can be considered as a complex microsystem 

including mechanical, electronic, fluid functions, etc.. It also needs a network of 

microchannels. 

Since the ‘planar’ technology was introduced into microelectronics at the end of the 1950s, 

micromanufacturing techniques have never stopped improving, continuously extending the 

limits of miniaturization. Silicon was initially chosen for its semi-conducting properties and 

the excellent quality of its oxide, and was also used to develop mechanical functions 

including resonant beams, micro-motors, etc [2-4]. 

The concept of microsystems such as (MEMS—Micro-Electro-Mechanical Systems) 

emerged during the 1980s at the University of Berkeley as the integration of the different 

functions of a complete system on a single chip [5]. The first example of a microsystem 

marketed by Analog Devices in 1993 was the ADXL50 accelerometer that consisted of a 

capacitive sensor accompanied by its associated electronics monolithically integrated onto 

a 10 mm 2 substrate [6]. MEMS technology takes advantages from microelectronic 

production tools that allow for the manufacturing of thousands of miniaturized objects in 

parallel and thus reduces drastically fabrication costs. 

Several examples of microsystems are now used in daily life: accelerometers that are 

present in modern cars (airbag trigger system) [7,8], and also in Nike shoes (integrated 

pedometer), Sony Playstation games consoles (measurement of the inclination of the 

joystick) and GPS systems [9,10], inkjet print heads [11-14], arrays of micromirrors in 

video projectors [15,16]. Many sectors are now concerned such as optical and radio 

frequency telecommunications (switches, variable capacitances and inductances) [17-19]. 

Finally, the interest in microsystems in the chemistry and biomedical fields has never 

ceased increasing, particularly since the development of the miniaturized total chemical 

analysis system (μTAS) concept by A Manz at the beginning of the 1990s and then since 

the development of LOC [20]. 

Over the last 10 years, this popularity has been a driving force for the development of new 

types of microsystems combining electrical and mechanical functions with microfluidic 

functions (microwells, microchannels, valves, pumps). It also encouraged the introduction 

of new processes and materials in microtechnologies. The glass technology widely used at 

first, combining wet etching and thermal bonding, is limited by difficulties in terms of 

integration and aspect ratio, while etching of silicon is much more developed. 

Silicon and glass technologies 

• Bulk micromachining 

Bulk micromachining refers to the situation in which patterns are defined in the substrate 

itself. The example of wet etching of glass is shown in figure 1. 

90

Figure 1: (1) An example of bulk micromachining: wet etching of glass; (2) microchannels etched by 

wet etching in glass [IMT Neuchˆatel] [21] (© 2003 IEEE). 

It clearly illustrates the generic approach to bulk machining. After the substrate has been 

cleaned, the material that will be used for protection during the etching step is deposited 

(a). A photoresist is coated using a spin coater (b). This resist is exposed to ultraviolet 

radiation through a mask with transparent and opaque areas (c). The resist is then 

developed (d) and exposed areas are eliminated (positive resist, as illustrated in figure 1) 

or remain (negative resist), depending on its polarity. The protective material is etched (e) 

and the resist is removed from the substrate (f ). The next step is the glass etching step 

(g). Finally, the mask material is removed from the substrate (h). 

• Wet etching 

The etching step is fundamental in bulk micromachining [22, 23]. Etching as presented in 

figure 2 is isotropic, i.e. the etching rate is the same in all directions. Typical examples are 

wet etching of silicon in a mixture of hydrofluoric acid, nitric acid and ethanol [24], or wet 

etching of glass in hydrofluoric acid [25, 26]. Isotropic wet etching has several 

disadvantages, including difficulty in controlling the profile due to the strong influence of 

stirring and isotropy that severely limits the possible resolution and depth, as illustrated 

in figure 2. 

Figure 2: Under-etching phenomenon during isotropic wet etching 

Anisotropic wet etching of silicon [27-30] in solutions such as potassium hydroxide (KOH) 

or tetramethyl ammonium hydroxide (TMAH) [31-33] can be used to create patterns 

defined by crystallographic planes of silicon as illustrated in figure 3. This process, 

together with etching-stop techniques (geometric, electrochemical, implantation of boron) 

[34-36] makes it easy for manufacturing very thin membranes or suspended structures. 

However, some simple shapes such as a circular hole cannot be made, the etching profile 

limits the integration density and the chemicals used are relatively aggressive. 

Figure 3: (a) Different anisotropic wet etching profiles in single-crystalline silicon; (b) examples of microwells 

made by anisotropic wet etching [GESIM] [37], © 1999, with permission from Elsevier. 

91

• Glass etching 

The technology for the microfabrication of glass is gaining importance because more and 

more glass substrates are currently being used to fabricate MEMS devices. Glass has 

many advantages as a material for MEMS applications, such as good mechanical 

properties, good optical properties, high electrical insulation, and it can be easily bonded to 

silicon substrates at temperatures lower than for fusion bonding. 

Glass is particularly useful for fabrication of microfluidic devices used for bioanalysis due 

to its high chemical resistance. Microfluidic dispensing and controlling devices, such as 

micro pumps, micro pressure/flow sensors, monolithic membrane valve/diaphragm pumps, 

have been developed with integrated glass components. DNA related microfluidic devices, 

such as micro flow cells for DNA single molecule handling, micro injectors for DNA mass 

spectrometry, and micro polymerase chain reaction (PCR) devices for DNA amplification, 

have also been presented [38]. Being transparent under a wide wavelength range ( 300– 

2000 nm for Pyrex), glass is a prime candidate for incorporation into micro bioanalytical 

devices where optical detection of the bioanalytes is used, for example, micro capillary 

electrophoresis (μCE) devices [39]. 

Extensive investigation of glass microfabrication technologies is paving the way for 

increasing utilization of glass substrates in MEMS. These technologies include wet 

chemical etching using various concentrations of hydrofluoric acid (HF) [40, 41], dry etch 

techniques like deep reactive ion etching (DRIE) using SF6 plasma [42], laser 

micromachining [43] and ultrasonic drilling [44]. Laser drilling and ultrasonic drilling are 

suitable for fabricating the small numbers of through holes in the glass. 

Wet glass etching with various concentrations of HF is the most widely used method 

because the etching rate is fast and a large quantity of glass wafers can be processed 

simultaneously. Masks for glass etching in HF are normally photoresist [40] and Cr/Au 

photoresist combinations [45]. However, the formation of pinholes through defects within 

the metal mask is a notorious problem. This becomes especially severe when deep 

etching is required. A second problem found for HF etching is rapid undercutting of the Cr 

mask, which leads to a more rapid lateral etching of the glass than vertical etching. This 

leads to a poor aspect ratio generally smaller than one for etched cavities. Although the 

application of metal masks, for deep etching with HF, has the problem of pinholes and 

large lateral undercutting as mentioned above, the approach is simple and compatible with 

standard microfabrication facilities. 

• Assembly 

An assembly step between two machined substrates is usually necessary to seal 

structures. Several types of bonding techniques have been developed for different 

applications. Wafer bonding is a cost-effective method for zero-level MEMS packaging and 

has increasingly become a key technology for materials integration in various areas of 

MEMS, microelectronics and optoelectronics. Different wafer bonding approaches are 

currently used in the MEMS industry: fusion, adhesive, eutectic, anodic, solder bonding 

[46, 47]. 

These bonding techniques are summarized in figure 4. One of the main concerns is the 

thermal stress, mainly in the case of heterogeneous assembly, i.e. assembly of materials 

with different thermal expansion coefficients. The required level of cleanness may also be 

an obstacle in a process such as thermal bonding. Adhesive bonding is a low-temperature 

process and does not require complex cleaning procedures but introduces a 

supplementary material into the structure which may be undesirable in microfluidics in 

which the control of surface properties is critical. It is often estimated that assembly is 

frequently the most expensive step in microsystem processes. 

92

Figure 4: Different types of bonding between substrates 

As fusion bonding processes require a high temperature annealing which is not always 

suitable for the devices with aluminum or copper integrated circuits; adhesive bonding is 

non-hermetic, more and more interests are focused on low temperature bonding 

(maximum process temperature below 300°), which can not only reduce process cost and 

time, but also minimize bonding-induced stress and warpage after cooling. Moreover, the 

wafers with large difference in coefficient of thermal expansion can also be bonded. The 

hybrid integration approach appears better adapted to a lab-on-chip integration. The 

challenge is then to couple the different elements, for example a sensor made using the 

silicon technology and a microfluidic network replicated in a thermoplastic, with good 

alignment and without introducing additional dead volumes, and obviously at the lowest 

cost. 

Anodic bonding is one of the most used wafer level packaging procedures. In anodic 

bonding, the substrates are heated to a typical temperature above 400° and a typical 

voltage above 600V is applied to the wafer pair to be bonded, electrostatic force and the 

migration of ions lead to an irreversible chemical bond at the boundary layer between the 

individual wafers. Conventionally, this is done in the furnace or on the hotplate, which 

require long ramp times, consume large amounts of power, and have significant 

manufacturing footprints. Currently, many efforts are focusing on improving anodic 

bonding quality [48, 49]. 

However, the bond quality is reduced with decrease in the bonding temperature, low 

bonding strength will result in the bubbles or cavities at the interface, only few literature 

reports the success in low temperature anodic bonding with high bond strength and 

bubble-free interface [49]. 

The concept of μTAS, later extended to the lab-on-a-chip concept, was introduced in 1990 

to mitigate the deficiencies of chemical sensors, particularly in terms of selectivity and life 

time [20]. Integration onto a chip can also very strongly reduce the consumption of 

chemicals that may be rare, expensive and/or polluting. Similarly, it reduces the production 

of waste. Automation of procedures illustrated by the DNA analysis station marketed by 

the Caliper Company, increases analysis rates and reduces the risk of error due to the 

human factor [50]. 

Furthermore, these products enjoy the advantages of prices for mass production 

processes. Miniaturization also provides a means of making these systems portable. 

Cost and size reductions are not the only advantages; performance gain is also important. 

For example, an immunoassay was carried out in less than 25s while more than 10 min 

are necessary under classical conditions [51]. Therefore miniaturization of volumes 

provides a means of reducing the time for the operation. 

93

Many lab-on-chip applications have been developed in genomics, under the emphasis of 

the human genome project and sequencing needs [52]. Moreover genetic analyses 

became increasingly routine operations, efforts have been made and are still being made 

on analysis of proteins. Its traditional approach includes extraction from cells, separation 

by electrophoresis on gel, digestion and analysis by a mass spectrometer. These 

operations are slow and time consuming. A system was made to perform the digestion, 

separation and analysis steps [53]. The operation that normally takes hours is done in a 

few minutes. 

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Microsystems Based on Porous Silicon-Glass Anodic Bonding 

technologies 

This section is aimed to the demonstration of the compatibility of porous silicon and anodic 

bonding technologies for the realization of sensing micro-components in lab-on-chip 

applications. In particular, it was developed and tested an innovative anodic bonding 

process: voltage, time and temperature have been properly matched in order not to affect 

the properties of the transducing material, which can be strongly modified by thermal 

treatments through oxidation.

Sensors 2006, 6, 680-687 

Full Research Paper 

sensors 

ISSN 1424-8220 

© 2006 by MDPI 

http://www.mdpi.org/sensors 

A Microsystem Based on Porous Silicon-Glass Anodic Bonding 

for Gas and Liquid Optical Sensing 

Luca De Stefano 1,* , Krzysztof Malecki 1 , Francesco G. Della Corte 2 , Luigi Moretti 2 , 

Ilaria Rea 1 , Lucia Rotiroti 1 and Ivo Rendina 1 

1 Institute for Microelectronics and Microsystems, National Council of Research, Via P. Castellino 

111, 80131 Naples, Italy (E-mail for Ilaria Rea: ilaria.rea@na.imm.cnr.it) 

2 DIMET “Mediterranea” University of Reggio Calabria, Località Feo di Vito, 89060 Reggio Calabria, 

Italy 

* Author to whom correspondence should be addressed. E-mail: luca.destefano@na.imm.cnr.it 

Received: 15 March 2006 / Accepted: 22 June 2006 / Published: 23 June 2006 

Abstract: We have recently presented an integrated silicon-glass opto-chemical sensor for 

lab-on-chip applications, based on porous silicon and anodic bonding technologies. In this 

work, we have optically characterized the sensor response on exposure to vapors of several 

organic compounds by means of reflectivity measurements. The interaction between the 

porous silicon, which acts as transducer layer, and the organic vapors fluxed into the glass 

sealed microchamber, is preserved by the fabrication process, resulting in optical path 

increase, due to the capillary condensation of the vapors into the pores. Using the 

Bruggemann theory, we have calculated the filled pores volume for each substance. The 

sensor dynamic has been described by time-resolved measurements: due to the analysis 

chamber miniaturization, the response time is only of 2 s. All these results have been 

compared with data acquired on the same PSi structure before the anodic bonding process. 

Keywords: Porous Silicon, Anodic Bonding, Microchamber, Lab-on-Chip. 

Introduction 

The physical and structural properties of porous silicon (PSi), first of all its high surface area 

matrix, have led many scientific researchers to investigate this material, using it as a transducer in 

sensing systems. In particular, high sensitivity results have been obtained by monitoring changes in 

optical properties, such as photoluminescence [1], reflectivity [2-3] and ellipsometry [4]. Both 

monolayers, which act as Fabry-Perot interferometers, and multilayer devices, like Bragg filters and

Sensors 2006, 6 681 

optical microcavities, can be easily fabricated and exploited in optical sensing experiments. The 

integration of sensitive elements into complex microsystems, i.e. lab on chip or micro-total-analysis 

systems, is of straightforward interest, especially for the miniaturisation of each component. This step 

is never a trivial one: the fabrication process often requires high temperatures and mechanical stresses 

which can damage or even destroy the bio- or chemo- sensitive transducer structures. 

The anodic bonding (AB) is a standard IC fabrication technique which is widely used in 

microfluidic due to a wide spectrum of advantages among which the hermetic sealing [5-6]. Due to the 

good bonding quality, glass transparency, technological cleanness and high passivity to most of 

chemicals and biological substances, AB is also commonly exploited in the fabrication of lab-on-chip 

devices. The compatibility of this primary integration technique with highly demanding sensing microcomponents 

plays a basic role in view of the realization of lab-on-chip devices. We have recently 

reported on the optimisation of the standard silicon-glass AB parameters, searching for low 

temperature, low voltage and short time, taking into account the electrode type and thickness of glass 

wafers [7]. 

In the present work, the above mentioned results have been exploited and combined with new 

bonding findings, to ensure AB compatibility with the features of porous silicon (PSi) transducer. We 

have studied the static and dynamic sensor response on exposure to vapours of several organic 

compounds. The results have been compared with those obtained for the same PSi layer before the 

integration process. 

Theory 

Even if we have already realised µ-chambers with porous silicon multilayer optical structures, such 

as Bragg filters and optical microcavities which are more performing in terms of sensitivity and 

resolution, in this first optical characterisation, we have followed the approach of Sailor et al. [2] 

choosing the simple case of a single layer of porous silicon as our transducer device. Figure 1 shows a 

typical reflectivity spectrum from a PSi layer under white light illumination. 

Intensity (a. u.) 

1.8 

1.6 

1.4 

1.2 

1.0 

0.8 

0.6 

0.4 

Peak Order m 

14 

12 

10 

8 

6 

4 

2 

0 

11000 12000 13000 14000 15000 16000 17000 


600 650 700 750 800 850 900 


Figure 1. Reflectivity spectrum of PSi layer. In the inset the m-order peak are plotted as function of 

the wavenumber.

Sensors 2006, 6 682 

From the optical point of view, this structure is an optical Fabry-Perot interferometer, so that the 

maxima in the reflectivity spectrum appear at wavelengths λm which satisfy: 

m = 2 nd / λ m 

(1) 

where m is an integer, d is the film thickness and n is the average refractive index of the layer [8-9]. 

If we assume that the refractive index is independent on the wavelength over the considered range, the 

maxima are equally spaced in the wavenumber (1/λm). When the maximum order m is plotted as a 

function of the wavenumber, each point lies on a straight line which slope is the optical path, as it is 

shown in the inset of Fig. 1. When the PSi layer is exposed to vapors, or dip in the liquid phase of the 

same substance, the substitution of air in the pores by its molecules causes a fringes shift in 

wavelength, which corresponds to a change in the optical path nd. Since the thickness d is fixed by the 

physical dimension of the PSi matrix, the variation is clearly due to changes in the average refractive 

index. In the case of vapors and gases, the filling liquid phase is due to the capillary condensation 

phenomenon, which is regulated from the Kelvin equation: 

K BTρ ln( psat 

/ p) 

= 2γ 

cos / R 

l lg θ 

(2) 

where ρl is the density of the liquid phase, γlg is the liquid-gas surface tension coefficient at room 

temperature T, R is the radius of the pores, p/psat is the relative vapor pressure into the pores, and θ is 

the contact angle [10]. 

A quantitative model taking into account the optical path increase can be realized by applying the 

Bruggemann effective medium approximation theory. The relation between the volume fraction of 

each medium and its dielectric constant can be written as: 

⎛ εSi 

−εp 

⎞ ⎛ εair 

−εp 

⎞ ⎛ εch 

−εp 

⎞ 

( 1 − p) 

⎜ ⎟+ 

( p−V) 

⎜ ⎟+ 

V⎜ 

⎟= 

0 

⎜ 2 ⎟ ⎜ 2 ⎟ ⎜ 2 ⎟ 

⎝εSi 

+ εp 

⎠ ⎝εair 

+ εp 

⎠ ⎝εch 

+ εp 

⎠ 

where p, V, εsi, εair, εch and εp are the layer porosity, the layer liquid fraction (LLF), the dielectric 

constants of silicon, air, chemical substance and porous silicon, respectively [11]. 

Experimental Section 

We have designed and fabricated a sensor device which can be considered the basic element of a 

lab-on-chip, just integrating PSi, as the transducer material, and a glass slide, which ensures sealing 

and the interconnections for fluids inlet and outlet. The silicon wafer used in this work is a p+ type, 

crystal orientation, with a resistivity of 8-12 mΩcm. The glass is a Borofloat 33 type, 1 mm 

thick. The reaction microchamber has been realized by a two-step electrochemical etching. The first 

step is a high current density (800 mA/cm 2 for about 300 s) electrochemical etch in hydrofluoric acid 

and ethanol (HF:EtOH/50:50) solution which creates a µ-well (100 µm deep) in the bulk silicon. The 

second step is a consecutive electrochemical etch which is used to fabricate the PSi layer at the bottom 

of the µ-chamber. The process parameters for this last step were: current density of 400 mA/cm 2 and 

etching time of 3.2 s. More details on µ-chamber fabrication can be found elsewhere [12]. The 

thickness and porosity of the PSi layer was measured by variable angle spectroscopic ellipsometry 

using a Horiba-Jobin-Yvon UVISEL ellipsometer. After the mechanical drilling of the flow channels, 

the glass slide has been cleaned and activated for the AB process following standard RCA and H2O2 

procedures. Because of the highly reactive PSi nature, the standard cleaning procedures had to be 

(3)

Sensors 2006, 6 683 

changed with a soft cleaning procedure based on trichloroethylene, acetone and ethanol. The silicon 

chip has also been carefully rinsed in deionized water for several minutes. Silicon etched wafer and 

glass top prefabricated components have been anodically bonded together with mutual alignment: a 

successful bonding, in terms of mechanical strength and bond quality, has been obtained at a 

temperature of 200 o C, voltage of 2.5 kV, with a process time of 1.5 minutes. The cross section of the 

chip structure and the general-view of the real device are shown in Figure 2. The reflectivity spectra in 

the VIS-IR wavelength region have been recorded with a very simple experimental set-up: a white 

source illuminates, through an optical fiber and a collimator, the porous silicon at nearly normal 

incidence. The reflected light is collected by an objective and coupled into a multimode fiber. The 

signal is directed in an optical spectrum analyzer (Ando, Mod. AQ-6315B) and measured with a 0.2 

nm resolution. We have also performed time-resolved measurement in order to characterize the sensor 

dynamic: using the laser beam from an IR source, we have measured, as a function of time, the signal 

of a receiving photodetector before, during and after the exposure to acetone. 


From the ellipsometric measurements we have estimated the thickness and the porosity of the 

porous silicon sensitive layer, resulting in about 9.1 µm and 79 %, respectively. 

Figure 2. Schematic and real view of the µ-chamber for lab-on-chip applications. 

Using the Bruggemann relation in Eq. (2), the refractive index is calculated as 1.34. As it can be 

noted in the reflectivity spectra reported in Figure 3, on exposure to several volatile substances in nonsaturated 

atmosphere, due to the phenomenon of capillary condensation, the average refractive index 

of the layer increases, and, as a consequence the optical thickness of the porous silicon layer also 

increases.

Sensors 2006, 6 684 

Intensity (a.u.) 

1.3 

1.2 

1.1 

1.0 

0.9 

0.8 

0.7 

0.6 

0.5 

0.4 

0.3 

840 850 860 870 880 890 900 


Unperturbed 

On exposure to: 

Xylene 

Acetone 

Ethanol 

Iso-propanol 

Figure 3. Reflectivity spectra of the unperturbed sensor and on exposure to vapours of organic 

compounds. 

In Figure 4 the maxima of order m as function of wavenumber are shown; the slopes of the lines 

give the optical thickness of the layer for each substance. In the wavelength range considered, the 

assumption of a refractive index independent on the wavelength is satisfied since the PSi refractive 

index changes less then 2 % in this interval. Using the Eq. (3), the LLF, which is the filled volume by 

the condensed liquid, has been calculated for each substance obtaining the values reported in Table 1. 

The LLF values are about 30 % lower than ones obtained in the case of the PSi transducer before of 

the AB process [13]. These differences can be ascribed to a slight oxidation of the porous silicon layer 

due to the AB temperature process. 

maximum order (∆m) 

16 

14 

12 

10 

8 

6 

4 

2 

0 

Unperturbed 

On exposure to: 

Iso-propanol 

Acetone 

Ethanol 

Xylene 

12000 13000 14000 15000 16000 17000 

wavenumber (cm -1 ) 

Figure 4. Maximum m-order as function of wavenumber (1/λm). The slopes of the straight lines are the 

optical thicknesses of the unperturbed layer and on exposure to the vapours. 

The silicon oxide can fill or obstruct the very small pores of the sponge-like structure, preventing 

the liquid to condense into them. Moreover, due to the air present during the AB process, there

Sensors 2006, 6 685 

probably exists an extra pressure into the µ-chamber because some air has been entrapped in the PSi 

matrix. As a consequence, according to the Kelvin equation (2), the capillary condensation of the 

liquid phase decreases. 

Table 1. Chemical organics substances used in sensing experiments and some relevant physicalchemical 

properties a . 

Chemical 

compounds 

n ρ 

(g/cm 3 ) 

STC 

(mN/m) 

VP 

(kPa) 

BP 

(°C) 

∆ 

(nm) 

LLF 

Iso-propanol 1.377 0.785 20.93 6.8 82.4 1000 0.22 

Ethanol 1.360 0.785 22.8 5.8 78 950 0.23 

Xylene 1.501 1.454 38.8 0.046 214 1400 0.22 

Acetone 1.359 0.791 23.46 30.8 56 998 0.23 

a 

n is the liquid refractive index; ρ is the density (@ 25 °C); STC is the surface tension 

coefficient (@ 25 °C); VP is the vapor pressure; BP is the boiling point; ∆ is the 

average optical path increase and LLF the layer liquid fraction. 

The result of time-resolved measurement is compared in Figure 5 to the data acquired on the same 

PSi layer before the AB integration process: it is well evident that, due to chamber miniaturization, the 

identification time (τid=2 s) is significantly shorter. 

Intensity (a. u.) 

1.00 

0.75 

0.50 

0.25 

0.00 

gas in 

τ id =8.7 s 

τ id =2 s 

τ rec =1.7 s 

τ rec =8 s 

gas out 

0 30 60 90 120 150 180 210 240 270 300 

Time (s) 

Figure 5. Time-resolved measurements of porous silicon layer during the monitoring of acetone in 0.4 

l test chamber (solid line) and in the integrated 14 µl µ-chamber (dashed line). 

The values of response time depends not only on the physical phenomena involved (i.e., 

equilibrium between adsorption and desorption in the PSi layer) but also on the geometry of the test

Sensors 2006, 6 686 

chamber and on the measurement procedure, i.e. static or continuous flow mode. In static condition, 

the identification time is mainly determined by the diffusion of the gas into the chamber volume: in 

fact, when vapor is in contact with the porous silicon surface, the capillary condensation takes place 

instantaneously [14-15]. For the same reason, the recovery time is longer (τrec=8 s): as soon as 

Nitrogen is introduced into the µ-chamber, the conditions for capillary condensation are not still valid 

so that the liquid phase disappears, depending on atmosphere rate exchange. As it is shown in this 

graphic, the sensor response is completely reversible. 

Conclusions 

In this work, we have optically characterized an original hybrid silicon-glass sensor system, based 

on PSi and AB technologies, which can be thought as the reaction chamber of a more complex lab-onchip. 

Reflectivity spectrum of the device has been studied on exposure to vapors of organic volatile 

solvents. An increase of the optical path, due to capillary condensation of the vapor into the pores, has 

been observed. Using the Bruggemann effective medium approximation we have calculated the liquid 

volume fraction adsorbed into the layer for each substance, obtaining values between 0.22 and 0.23. 

We have characterized the sensor dynamic by time-resolved measurement during exposure to acetone. 

The data assure fast analysis time and the complete reversibility of the sensing mechanism. The 

comparison with data obtained from the same PSi transducer before the AB, shows that the integration 

process did not degrade the sensor features. 


This work has been supported by Marie Currie Fellowship of the European Community programme 

“Technologies for micro-opto-mechanical systems” under contract number CO-IMM-NA-01-2002. 

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© 2006 by MDPI (http://www.mdpi.org). Reproduction is permitted for noncommercial purposes.

Laser oxidation micropatterning of a porous silicon based 

biosensor for multianalytes microarrays 

The aim of this section is to present the fabrication and characterization of oxidized 

micropatterned biosensors based on PSi, obtained by direct laser writing and their 

chemical functionalization for biological applications. The laser local oxidation technique is 

a good alternative to the traditional photolithographic process, in fact the standard masking 

of PSi by means of photoresist presents remarkable difficulties since the developer, being 

alkaline, causes the dissolution of the PSi.

LASER OXIDATION MICROPATTERNING OF A POROUS 

SILICON BASED BIOSENSOR FOR MULTIANALYTES 

MICROARRAYS 

L. DE STEFANO 1 , L. ROTIROTI 1,3 , I. REA 1,2 , E. DE TOMMASI 1 , M. A. NIGRO 4 , F. 

G. DELLA CORTE 4 , AND I. RENDINA 1 

1 IMM-CNR - Sezione di Napoli, Via P. Castellino 111, 80131 Napoli, Italy 

2 Dept. of Physics, “Federico II” University of Naples, Monte S. Angelo, 80126 Naples, 

Italy 

3 Dept. of Organic Chemistry and Biochemistry, “Federico II” University of Naples, 

Monte S. Angelo, 80126 Naples, Italy 

4 “Mediterranea” University of Reggio Calabria Department of Computer Science, 

Mathematics, Electronics and Transports, Località Vito di Feo, Reggio Calabria, Italy 

Abstract 

In this communication, we present the fabrication and characterization of oxidized porous silicon 

(PSi) micropatterns, obtained by direct laser writing, and chemically functionalized for biological 

applications. The laser local oxidation technique is a good alternative to the traditional 

photolithographic process since this approach allows the development of micropatterned 

biosensors. In fact, from a technological point of view, the standard masking of PSi by means of 

photoresist presents remarkable difficulties because of the low resistance to the electrochemical 

process. The micropatterned PSi oxidized surfaces can be properly functionalized by a specific 

chemical linker. The modified surface is able to link the carboxylic or aminic groups of biological 

probes. The binding between the functionalized surface and the biological matter has been tested 

by directly spotting on the PSi local oxide a fluorescent labelled antibody. The device has been 

characterized by FT-IR measurements and fluorescence macroscopy. 

Keywords 

Porous silicon, laser writing, biosensors, microarrays. 

INTRODUCTION 

Porous silicon based optical biosensors offer several advantages in label-free 

detection and identification of organic molecules due to the well known 

properties of this material [1, 2]. PSi has a spongy structure with a large 

specific surface of the order of 100–500 m 2 cm -3 [3], so that a very effective 

interaction with several adsorbates is assured. When a biological solution 

penetrates into the PSi pores, it substitutes the air, so that the average dielectric 

properties of the heterogeneous silicon-air-liquid system change inducing a 

red-shift in the reflectivity spectrum of the device. The effect depends on the 

refractive index value of the solution but also on how it penetrates into the 

pores. Due to this sensing mechanism, PSi optical devices cannot identify the 

382

383 

single components of a complex mixture. In order to enhance the sensor 

selectivity, several biological probes, which exploit very specific interactions 

only with selected biochemical molecules, can be linked to the PSi surface: 

most common examples are DNA strands, proteins, enzymes, antibodies and 

so on. In this work, the fabrication and characterization of oxidized porous 

silicon micropatterns obtained by direct laser writing [5] and their chemical 

functionalization for biological applications are reported. In particular we have 

employed a murine monoclonal antibody (IgG) UN1 previously selected for 

the specific reactivity with human thymocytes as compared to peripheral blood 

cells [4]. The antigen recognized by UN1 is a 100–120 kDa transmembrane 

glycoprotein showing biochemical features of cell membrane-associated 

mucin-like glycoproteins, a class of macromolecules that are involved in cellto-cell 

interactions and cancer progression [6]. 

EXPERIMENTAL AND RESULTS 

The device was a PSi monolayer, 1 μm thick and with a porosity of 70%, 

produced by electrochemical etch in a HF-based solution, starting from a 

highly doped p+-silicon, oriented, 0.04 Ω cm resistivity, 400 μm thick. 

The silicon was etched using a 30 wt. % HF/ethanol solution in dark and at 

room temperature. After the etching the samples were rinsed in ethanol and 

dried under a stream of N2. 

Laser diode 

M2 

Pinhole 

M1 

XY stage 

Objective X40 

PS sample 

M3 

Fig. 1: Scheme of the setup used to write the channel waveguides. 

The localized oxidation is obtained by exposing the sample to the beam of a 

laser diode (μLS Micro Laser Sistem) at 408 nm with an output power of

384 

48mW, using a setup composed of a microscope objective and a motorized xy 

micrometric stage as shown in Figure 1. The beam dimension is about 1.4/0.32 

mm. We have directly written on the sample some oxide strips, 10 μm wide, 

with a laser fluence of 38 mJ/mm 2 . 

Starting from our previous experience on chemical modification of the 

oxidised porous silicon, we have used an organosilane compound such as the 

aminopropyltriethoxysilane (APTES) and a functional crosslinker such as 

Glutaraldehyde (GA) which reacts with the amino groups on the silanized 

surface and coats the micropatterned PSi oxidized surfaces. [7]. The obtained 

modified surface works as an active substrate for the chemistry of the 

following attachment of other biological probes, such as DNA single strand, 

proteins, enzymes, and antibodies. 

The functionalization process takes place directly on the chip surface, simply 

covering it with the chemical solutions. In Fig. 2 is reported the scheme of the 

functionalisation process. The PSi device has been rinsed by immersion in a 

5% solution of APTES and an hydroalcoholic mixture of water and methanol 

(1:1), for 20 min at room temperature. After the reaction time, we have washed 

the sample with demi-water and methanol and dried in N2 stream. The 

silanized surface was then baked at 100°C for 10 min. 

The APTES modified surface is able to link the carboxylic group of a 

biological probe such as IgG or other antibodies. A further step of 

functionalization is required to bind the amino group always present in 

biological matter through the aldehydic group of opportune chemical linker. To 

this aim, we have treated the chip in a 2.5% glutaraldehyde solution for 30 

min. The glutaraldehyde reacts with the amino groups on the silanized surface 

and coats the internal surface of the pores with another thin layer of molecules. 

Infrared spectra were collected by using a Fourier transform spectrometer 

equipped with a microscope (Nicolet Continuμm XL, Thermo Scientific) and 

are reported in Fig. 3, where it is well evident that, after the functionalization 

processes, the Si-OH peaks are completely substituted by the organic linkers 

characteristics ones.

Fig. 2: functionalization steps of PSi surface by means of APTES and Glutaraldehyde and 

subsequent binding of the biological probe. 

Fig. 3: FT-IR spectra of the micropatterned PSi oxidized surfaces before and after the chemical 

surface modification treatments. 

385 

The binding between the surface and the biological matter has been tested by 

using fluorescent labelled bioprobes directly spotted on the PSi chemical 

modified structures. We have spotted on the porous silicon chip a solution 

containing the rhodamine labelled IgG antibody, 6.8 μM, and incubated 

overnight at room temperature. The device has been observed by a Leica Z16

386 

APO fluorescence macroscopy system. By 100W high-pressure mercury 

source we found that the fluorescence is very high and homogeneous on the 

region locally oxidized by the laser as shown in Fig. 4. 

Fig. 4: Post-dialysis fluorescence image of the chip after rhodamine labeled IgG binding 

CONCLUSIONS 

In this work, we have exploited the possibility to locally functionalize the PSi 

surface. 

A micropattern has been directly written on the porous silicon layer by a laser 

oxidation process. The pattern, constituted by strips 2 μm wide, has been 

functionalized by specific chemical linkers such as APTES and glutaraldehyde; 

then we have spotted on the PSi sample a fluorescent labelled antibody. The 

interaction between the modified surface and the biological matter has been 

verified by fluorescence macroscopy; a high fluorescence signal has been 

observed only in the oxidized region. 

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PUBLICATIONS: 

J1. Luca De Stefano, F. G. Della Corte, K. Malecki, Ivo Rendina, Lucia Rotiroti, Luigi Moretti, 

Andrea M. Rossi, “Integrated silicon-glass opto-chemical sensors for lab-on-chip application”, 

Sensor and Actuators B,114, 625-630, (2006). 

J2. L. De Stefano, I. Rendina, L. Rotiroti, L. Moretti, V. Scognamiglio, M. Rossi, S. D’Auria, 

“Porous silicon-based optical microsensor for the detection of L-Glutamine”, Biosensors and 

Bioelectronics 21/8, 1664-1667, 2006. 

J3. L. Moretti, I. Rea, L. Rotiroti, I. Rendina, G. Abbate, A. Marino, L. De Stefano “Photonic band 

gaps analysis of Thue-Morse multilayers made of porous silicon” Optics Express, 14 (2006) 

6264-6272. 

J4. L. De Stefano, I. Rea, I. Rendina, L. Rotiroti, M. Rossi, and S. D’Auria, “Resonant cavity 

enhanced optical microsensor for molecular interactions based on porous silicon”, Phys Stat 

Sol (A) , 203 (2006) 886-891. 

J5. S. D’Auria, S. Borino, A. Vitale, A.M. Rossi, I. Rea, I. Rendina, L. Rotiroti, M. Staiano, M. de 

Champdorè, V. Aurilia, A. Parracino, M. Rossi, L. De Stefano, “Nanostructured silicon-based 

biosensors for the selective identification of analytes of social interest”, Journal of Physics: 

Condensed Matter, 18 S2019-S2028, (2006). 

J6. L. De Stefano, L. Rotiroti, I. Rea, I. Rendina, L. Moretti, G. Di Francia, E. Massera, A. 

Lamberti, P. Arcari, C. Sangez, “Porous silicon-based optical biochips” Journal of Optics A: 

Pure and Applied Optics, S540-S544, 8 (2006). 

J7. L. De Stefano, K.Malecki, F. G. Della Corte , L. Moretti , I. Rea , L. Rotiroti, I. Rendina “A 

Microsystem Based on Porous Silicon-Glass Anodic Bonding for Gas and Liquid Optical 

Sensing” Sensors, 6 (2006) 680-687. 

J8. L. De Stefano, M. Rossi, M. Staiano, G. Mamone, A. Parracino, L. Rotiroti, I. Rendina, M. 

Rossi, and S. D’Auria, “Glutamine-binding protein from Escherichia coli specifically binds a 

wheat gliadin peptide allowing the design of a new porous silicon-based optical biosensor”, 

Journal of Proteome Research 5,5, 1241-1245, 2006 

J9. L. De Stefano, L. Rotiroti, I. Rea, F. Della Corte, D. Alfieri, L. Moretti, I. Rendina, “An 

integrated pressure-driven microsystem based on porous silicon for optical monitoring of 

gaseous and liquid substances”; Phys. Stat. Sol. (a) 204, No. 5, 1459-1463 (2007). 

J10. L. De Stefano, L. Rotiroti, I. Rea, I. Rendina, L. Moretti, "Quantitative determinations 

in hydro-alcoholic binary mixtures by porous silicon optical microsensors", Phys. Stat. Sol. (c), 

4, No. 6, 1941– 1945 (2007). 

J11. L. De Stefano, I. Rea, L. Moretti, L. Rotiroti, I. Rendina, “Optical properties of porous 

silicon Thue-Morse structures”, Phys. Stat. Sol. (c) 4, No. 6, 1966-1970 (2007). 

J12. I. Rendina, I. Rea, L. Rotiroti, L. De Stefano, “Porous Silicon Based Optical 

Biosensors and Biochips”, Physica E, 38 (2007) 188-192 invited paper. 

J13. L. De Stefano, I. Rea, L. Rotiroti, I. Rendina, L. Fragomeni, F.G. Della Corte, “An 

integrated hybrid optical device for sensing applications”, Phys. Stat. Sol. (c), 4, No. 6, 1946– 

1950 (2007) 

J14. L. De Stefano, L. Rotiroti, I. Rea, I. Rendina, M. Rossi, S. D’Auria, “Assessment of a 

porous silicon transducer for the development of optical biochips”, Rivista di Materiali 

Nanocompositi e Nanotecnologie 3, 39-41, 2006 

J15. L. De Stefano, I. Rea, L. Rotiroti, M. Iodice, I. Rendina, “Optical microsystems based 

on a nanomaterial technology”, J. Phys.: Condens. Matter 19 395008, doi:10.1088/0953- 

8984/19/39/395008, (2007). 

J16. L. De Stefano, L. Rotiroti, I. Rendina, A.M. Rossi, M. Rossi, S. D’Auria, “Biochips at 

work: porous silicon microbiosensor for proteomic diagnostic”, Phys.: Condens. Matter 19 

395007, doi:10.1088/0953-8984/19/39/395007 (2007). 

J17. L. De Stefano, P. Arcari, A. Lamberti, C. Sanchez, L. Rotiroti, I. Rea, I. Rendina, 

“DNA optical detection based on porous silicon technology: from biosensors to biochips” 

Sensors, 7 (2007) 214-221. 

J18. L. De Stefano, A. Lamberti, L. Rotiroti, M. De Stefano, “Interfacing the nanostructured 

biosilica microshells of the marine diatom Coscinodiscus wailesii with biological matter”, Acta 

Biomaterialia Volume 4, Issue 1, January 2008, Pages 126-130. 

J19. Ivo Rendina, Edoardo De Tommasi, Ilaria Rea, Lucia Rotiroti, and Luca De Stefano, 

“Porous silicon optical transducers offer versatile platforms for biosensors” SPIE NEWSROOM 

DOI: 10.1117/2.1200801.0982 invited paper

J20. L. De Stefano, A. Vitale, I. Rea, M. Staiano, L. Rotiroti, T. Labella, I. Rendina, V. 

Aurilia, M. Rossi, S. D’Auria, “Enzymes and proteins from extremophiles as hyperstable 

probes in nanotechnology: the use of D-trehalose/D-maltose-binding protein from the 

hyperthermophilic archaeon Thermococcus litoralis for sugars monitoring”, Extremophiles, 12, 

1, 69-73 (2008). 

J21. L. De Stefano, E. De Tommasi, I. Rea, L. Rotiroti, L. Giangrande, G. Oliviero, N. 

Borbone, A. Galeone and G. Piccialli, “Oligonucleotides direct synthesis on porous silicon 

chip”, Nucleic Acids Symposium Series No. 52 721–722 

J22. L. De Stefano, L. Rotiroti, M. De Stefano, A. Lamberti, S. Lettieri, A. Setaro, P. 

Maddalena, “Marine Diatoms as Optical Biosensors" has been accepted for publication in 

Biosensors and Bioelectronics, 2008. 

J23. E. De Tommasi, L. De Stefano, I. Rea, V. Di Sarno, L. Rotiroti, P. Arcari, A. Lamberti, 

C. Sanges, and I. Rendina, “Porous Silicon Based Resonant Mirrors for Biochemical Sensing”, 

Sensors 2008, Volume 8 (10), Page 6549-6556. 

PROCEEDINGS: 

P1. L. De Stefano, L. Rotiroti, I. Rea, L. Moretti, I. Rendina, “Quantitative determinations in 

liquid and gaseous binary misture by porous silicon optical microsensors”, Advances in 

Sensors and Interfaces, Proc. of IWASI 2005, Eds. D. De Venuto and B. Courtois, 178-180, 

(2005). 

P2. L. De Stefano, I. Rendina, L. Rotiroti, L. Moretti, V. Scognamiglio, M. Rossi, S. D’Auria, 

“Protein-ligand interaction detection by a porous silicon optical sensor”, Proceedings of the 

10 th Italian Conference Sensors and Microsystems 2005, World Scientific, 89. 

P3. Luca De Stefano, F. G. Della Corte, K. Malecki, Ivo Rendina, Lucia Rotiroti, Luigi Moretti, 

Andrea M. Rossi, “A microchamber for lab-on-chip applications based on silicon-glass anodic 

bonding”, Proceedings of the 10 th Italian Conference Sensors and Microsystems 2005, World 

Scientific, 450, 2005. 

P4. L. De Stefano, K. Malecki, F. Della Corte, L. Moretti, I. Rea, L. Rotiroti, I. Rendina, 

“Silicon/glass integrated optical sensor based on porous silicon for gas and liquid inspection” 

Proceedings EUROSENSORS XIX, Settembre 11-14, 2005, Barcellona, Spagna. 

P5. L. Rotiroti, L. De Stefano, I. Rea, L. Moretti, G. Di Francia, V. La Ferrara, A. Lamberti, P. 

Arcari, C. Sanges, I. Rendina; “Porous silicon based optical biochips” Book of Abstract, 

European Optical Society Topical Meeting on Optical Microsystems, Settembre 15-18, 2005, 

Capri, Italia. 

P6. L. De Stefano, I. Rea, L. Rotiroti, I. Rendina, A. Lamberti, P. Arcari, C. Sanges, A. M. 

Rossi, “Resonant cavity enhanced optical microsensor for molecular interaction based on 

porous silicon”; Book of Abstract, The 11th National Conference on Sensor and Microsystems, 

p. 139-140, Febbraio 8-10, 2006, Lecce, Italia. 

P7. L. De Stefano, L. Rotiroti, I. Rea, I. Rendina, L. Moretti, “Optical quantitative determination 

of the alcohol fraction in binary mixtures”; Book of Abstract, The 11th National Conference on 

Sensor and Microsystems, p. 141-142, Febbraio 8-10, 2006, Lecce, Italia. 

P8. L. De Stefano, I. Rea, L. Rotiroti, I. Rendina, L. Fragomeni, F. G. Della Corte, “A hybrid 

integrated optical microsystem for sensing application”; Book of Abstract, The 11th National 

Conference on Sensor and Microsystems, p. 245-246, Febbraio 8-10, 2006, Lecce, Italia. 

P9. L. De Stefano, I. Rea, L. Rotiroti, K. Malecki, L. Moretti, F. G. Della Corte, I. Rendina, 

Sensors and Microsystems, 227-231 (2006). 

P10. L. De Stefano, I. Rea, L. Rotiroti, L. Moretti, I. Rendina, “Optical properties of porous 

silicon Thue-Morse structures”; Extended Abstracts of 5th International Conference on Porous 

Semiconductuctors-Science and Technology, p. 73-74, Marzo 12-17, 2006, Sitges-Barcellona, 

Spagna. 

P11. L. De Stefano, L. Rotiroti, I. Rea, F. Della Corte, D. Alfieri, L. Moretti, I. Rendina, “An 


gaseous and liquid substances”; Extended Abstracts of 5th International Conference on 

Porous Semiconductuctors-Science and Technology, p. 126-127, Marzo 12-17, 2006, Sitges- 

Barcellona, Spagna.

P12. L. De Stefano, L. Rotiroti, I. Rea, L. Moretti, I. Rendina, “Quantitative determinations in 

hydro-alcoholic binary mixtures by porous silicon optical microsensors”; Extended Abstracts of 

5th International Conference on Porous Semiconductuctors-Science and Technology, p. 248- 

249, Marzo 12-17, 2006, Sitges-Barcellona, Spagna. 

P13. L. De Stefano, I. Rea, L. Rotiroti, L. Fragomeni, F. Della Corte, I. Rendina, ”An integrated 

hybrid optical device for sensing applications”; Extended Abstracts of 5th International 

Conference on Porous Semiconductuctors-Science and Technology, p. 250-251, Marzo 12- 

17, 2006, Sitges-Barcellona, Spagna. 

P14. L. De Stefano, P. Maddalena, L. Moretti, I. Rea, I. Rendina, “Porous biosilica from marine 

diatoms: a new brand porous material for photonic applications”; Extended Abstracts of 5th 

International Conference on Porous Semiconductuctors-Science and Technology, p. 252-253, 

Marzo 12-17, 2006, Sitges-Barcellona, Spagna. 

P15. L. De Stefano, I. Rea, L. Rotiroti, L. Moretti, I. Rendina “Resonant photonic devices 

based on porous silicon: a perfect tool for optical sensing”; Proceedings of XX 

EUROSENSORS Conference, p. 294-295, Settembre 17-20, 2006, Göteborg, Svezia. 

P16. L. De Stefano, L. Rotiroti, I. Rea, E. De Tommasi, M. A. Nigro, F. G. Della Corte, and I. 

Rendina, Laser oxidation micropatterning of a porous silicon based biosensor for multianalytes 

microarray, Proceedings of the 12 th Italian Conference Sensors and Microsystems 2007, 

World Scientific, 382-387. 

P17. L, De Stefano, L. Rotiroti, I. Rea, E. De Tommasi, A. Vitale, M. Rossi, I. Rendina and S. 

D’Auria, "Design and realization of highly stable porous silicon optical biosensor based on 

proteins from extremophiles" Proc. of SPIE (2007) Vol. 6585 658517-1 

P18. L. De Stefano, I. Rendina, I. Rea, L. Rotiroti, E. De Tommasi, and G. Barillaro, " An 

optical microsystem based on silicon-air Bragg mirror for the continuous monitoring of liquid 

substances" Proc. of SPIE (2007) Vol. 6585 65850N-1 

P19. Ivo Rendina, Edoardo De Tommasi, Ilaria Rea, Lucia Rotiroti, and Luca De Stefano, 

“Porous silicon optical transducers offer versatile platforms for biosensors” DOI: 

10.1117/2.1200801.0982 

P20. L. De Stefano, I. Rendina, I. Rea, L. Moretti, L. Rotiroti, “Aperiodic photonic bandgap 

devices based on nanostructured porous silicon”, Proceedings of SPIE, Volume 6593, 

Photonic Materials, Devices, and Applications II, Ali Serpengüzel, Gonçal Badenes, Giancarlo 

C. Righini, Editors, 65931A (Jun. 5, 2007). 

P21. I. Rendina, I. Rea, L. Rotiroti, E. De Tommasi, L. De Stefano, “Integrated optical 

biosensors and biochips based on porous silicon technology”, Proceedings of SPIE -- Volume 

6898, Silicon Photonics III, Joel A. Kubby, Graham T. Reed, Editors, 68981D (Feb. 13, 2008) 

P22. Lucia Rotiroti, Paolo Arcari, Annalisa Lamberti, Carmen Sanges, Edoardo De Tommasi, 

Ilaria Rea, Ivo Rendina, and Luca De Stefano, “Optical detection of PNA/DNA hybridization in 

resonant porous silicon-based devices” Proc. SPIE 6991, 699120 (2008). 

P23. L. Rotiroti, E. De Tommasi, I. Rendina, M. Canciello, G. Maglio, R. Palumbo, and L. De 

Stefano,” BIOACTIVE POLIMERS MODIFIED POROUS SILICON NANOSTRUCTURES”, 6th 

International Conference on Porous Semiconductuctors-Science and Technology, p. 348-349, 

Marzo 10-14, 2008, Mallorca, Spagna. 

P24. L. Rotiroti, I. Rendina, E. De Tommasi, M. Canciello, G. Maglio, R. Palumbo, and L. De 

Stefano, A Nanostructured Hybrid Device Based On Polymers Infiltrated Porous Silicon 

Layers For Biotechnological Applications, Proceedings of the Polymer Processing Society 

24th Annual Meeting ~ PPS-24 ~ June 15-19, 2008 Salerno (Italy). 

P25. L. Rotiroti, I. Rendina, E. De Tommasi, M. Canciello, G. Maglio, R. Palumbo, and L. De 

Stefano, A Hybrid Optical Biosensor Based on Polymer Infiltrated Porous Silicon Device, IEEE 

SENSORS 2008, Lecce (Italy). 

NATIONAL AND INTERNATIONAL CONGRESS 

C1. L. De Stefano, I. Rendina, L. Rotiroti, L. Moretti, V. Scognamiglio, M. Rossi, S. D’Auria, 

“Protein-ligand interaction detection by a porous silicon optical sensor”, Oral, AISEM 2005, 

Firenze, 15-17 Febbraio, 2005.

C2. Luca De Stefano, F. G. Della Corte, K. Malecki, Ivo Rendina, Lucia Rotiroti, Luigi Moretti, 

Andrea M. Rossi, “A microchamber for lab-on-chip applications based on silicon-glass 

anodic bonding”, Poster, AISEM 2005, Firenze, 15-17 Febbraio, 2005. 

C3. L. De Stefano, L. Rotiroti, I. Rea, L. Moretti, I. Rendina, “Quantitative determinations in 

liquid and gaseous binary mixtures by porous silicon optical microsensors”, Oral, IWASI 

2005, Bari, 19-20 April, 2005. 

C4. L. De Stefano, L. Rotiroti, I. Rea, L. Moretti, I. Rendina, “Microsensori ottici per applicazioni 

biochimiche basati sulla nanotecnologia del silicio poroso”, Oral, GE2005, Giardini Naxos 

(CT), 29 giugno – 2 luglio, 2005. 

C5. M. A. Ferrara, L. Sirleto, L. Rotiroti, L. Moretti, E. Santamato, I. Rendina “EMISSIONE 

RAMAN IN SILICIO POROSO A 1.5 μm”, Oral, GE2005, Giardini Naxos (CT), 29 giugno – 

2 luglio, 2005. 

C6. L. De Stefano, K. Malecki, F. G. Della Corte, L. Moretti, I. Rea, L. Rotiroti, I. Rendina, Oral, 

“Silicon/glass integrated optical sensor based on porous silicon for gas and liquid 

inspection”, Eurosensors XIX, Barcelona, Spain, 9-14 September, 2005. 

C7. L. Rotiroti, L. De Stefano, I. Rea, L. Moretti, G. Di Francia, V. La Ferrara, A. Lamberti, P. 

Arcari, C. Sanges, I. Rendina, “Porous silicon based optical biochips”; Oral, European 

Optical Society Topical Meeting on Optical Microsystems, Settembre 15-18, 2005, Capri, 

Italia. 

C8. M. A. Ferrara, L. Sirleto, G. Messina, M.G. Donato, L. Rotiroti and I. Rendina, “STRAIN 

MEASUREMENTS IN POROUS SILICON BY RAMAN SCATTERING”, AISEM 2006, 

Lecce, 8-10 Febbraio, 2006. 

C9. M. A. Ferrara, L. Sirleto, L. Moretti, L. Rotiroti, B. Jalali and I. Rendina, “RAMAN 

EMISSION IN POROUS SILICON: PROSPECT FOR AN AMPLIFIER”, SPIE Silicon 

Photonics. 

C10. L. De Stefano, I. Rea, L. Rotiroti, I. Rendina, A. Lamberti, P. Arcari, C. Sanges, A. M. 

Rossi, “Resonant cavity enhanced optical microsensor for molecular interaction based on 

porous silicon”; The 11th National Conference on Sensor and Microsystems, Febbraio 8-10, 

2006, Lecce, Italia. 

C11. L. De Stefano, L. Rotiroti, I. Rea, I. Rendina, L. Moretti, “Optical quantitative 

determination of the alcohol fraction in binary mixtures”; The 11th National Conference on 

Sensor and Microsystems, Febbraio 8-10, 2006, Lecce, Italia. 

C12. L. De Stefano, I. Rea, L. Rotiroti, I. Rendina, L. Fragomeni, F. G. Della Corte, “A hybrid 

integrated optical microsystem for sensing application”; The 11th National Conference on 

Sensor and Microsystems, Febbraio 8-10, 2006, Lecce, Italia. 

C13. I. Rea, D. Alfieri, K. Malecki, I. Rendina, L. Rotiroti, L. De Stefano, L. Moretti, F.G. Della 

Corte, “Fast optical detection of chemical substances in integrated silicon-glass chip”; The 

11th National Conference on Sensor and Microsystems, Febbraio 8-10, 2006, Lecce, Italia. 

C14. L. De Stefano, I. Rea, L. Rotiroti, L. Moretti, I. Rendina, “Optical properties of porous 

silicon Thue-Morse structures”; of 5th International Conference on Porous 

Semiconductuctors-Science and Technology, Marzo 12-17, 2006, Sitges-Barcellona, 

Spagna. 

C15. L. De Stefano, L. Rotiroti, I. Rea, F. Della Corte, D. Alfieri, L. Moretti, I. Rendina, “An 


gaseous and liquid substances”; 5th International Conference on Porous 


Spagna. 

C16. L. De Stefano, L. Rotiroti, I. Rea, L. Moretti, I. Rendina, “Quantitative determinations in 

hydro-alcoholic binary mixtures by porous silicon optical microsensors”; 5th International 

Conference on Porous Semiconductuctors-Science and Technology, Marzo 12-17, 2006, 

Sitges-Barcellona, Spagna. 

C17. L. De Stefano, I. Rea, L. Rotiroti, L. Fragomeni, F. Della Corte, I. Rendina, “An integrated 

hybrid optical device for sensing applications”; 5th International Conference on Porous 


Spagna.

C18. L. De Stefano, I. Rea, L. Rotiroti, I. Rendina, “Porous silicon biosensors and biochips”, 

Invited, EMRS 2006, May 29-June 2, Nice, France, 2006. 

C19. M. Angeloni, L. Moretti, V. Mocella, L. De Stefano, L. Rotiroti, I. Rea, S. Bernstorff, “In 

situ investigation of capillary condensation phenomena in porous silicon nanostructured by 

means of GISAX and SAXS measurements”, Poster, EMRS 2006, May 29-June 2, Nice, 

France, 2006. 

C20. L. De Stefano, I. Rea, L. Rotiroti, L. Moretti, F. G. Della Corte, I. Rendina, “Integrated 

optical microsensors based on porous silicon technology for liquid and gas detection”, 

Poster, EMRS 2006, May 29-June 2, Nice, France, 2006. 

C21. M. A. Ferrara, L. Sirleto, M. Angeloni, L. Rotiroti, B. Jalalii, I. Rendina, “Raman emission 

in porous silicon: prospects fora n amplifier”, Poster, EMRS 2006, May 29-June 2, Nice, 

France, 2006. 

C22. C. Sanges, A. Lamberti, L. Rotiroti, L. De Stefano, I. Rendina, P. Arcari, “Biochip ottici in 

Silicio Poroso”, Poster, XII GIORNATE SCIENTIFICHE - Facoltà di Medicina e Chirurgia, 

Agraria, Medicina veterinaria, Farmacia, Scienze MM.FF.NN., Scienze Biotecnologiche, 15- 

16 June, Napoli, Italy, 2006. 

C23. I. Rea, L. Rotiroti, I. Rendina. L. Moretti, L. De Stefano; “Proprietà ottiche di multistrati 

thue-morse realizzati con la tecnologia del silicio poroso” Riunione Annuale del Gruppo 

Elettronica (GE 2006), Giugno 21-23, Ischia (NA), Italia. 

C24. L. Rotiroti, I. Rea, D. Alfieri, L. Moretti I. Rendina, F. Della Corte, L. De Stefano; 

“Microsensori ottici integrati basati sulla tecnologia del silicio poroso per il riconoscimento di 

sostanze liquide e gassose” Riunione Annuale del Gruppo Elettronica (GE 2006), Giugno 

21-23, Ischia (NA), Italia. 

C25. L. De Stefano, L. Rotiroti, I. Rea, S. D’Auria, A. Vitale, M. Rossi, I. Rendina, “A biosensor 

for selective identification of analites of social interest based on porous silicon technology”, 

IMCS11, July 16-19, Brescia, Italy, 2006. 

C26. L. De Stefano, I. Rea, L. Rotiroti, L. Moretti, I. Rendina “Resonant photonic devices 

based on porous silicon: a perfect tool for optical sensing”; XX EUROSENSORS, 

Settembre 17-20, 2006, Göteborg, Svezia. 

C27. I. Rea, L. De Stefano, L. Rotiroti, M. Iodice, I. Rendina; “Optical microsystems based on a 

nanomaterial technology” Oral, Nanoscience & Nanotechnology 2006, Novembre 6-9, 

2006, Monte Porzio Catone (Rome), Italy. 

C28. I. Rea, L. De Stefano, L. Rotiroti, E. De Tommasi, A. Nigro, F.G. Della Corte, I. Rendina, 

“Laser oxidation micropatterning of a porous silicon based biosensor for multianalytes 

microarrays”, Poster, AISEM 2007, Febbraio 12-14 2007, Napoli, Italy. 

C29. L. De Stefano, L. Rotiroti, I. Rea, L. Moretti, I. Rendina, “High-sensitivity optical sensors 

based on aperiodic multilayer structures”, Poster, AISEM 2007, Febbraio 12-14 2007, 

Napoli, Italy. 

C30. L. De Stefano, L. Rotiroti, I. Rea, E. De Tommasi, A, Vitale, M. Rossi. I. Rendina and S. 

D’Auria “Design and realization of highly stable porous silicon optical biosensor based on 

proteins from extremophiles” – Oral presentation. SPIE Europe, Optics and 

Optoelectronics, 2007, Prague, Czech Republic. 

C31. L. D Stefano, I. Rendina, I. Rea, L. Rotiroti, E. De Tommasi and G. Barillaro “An optical 

microsystem based on vertical silicon-air Bragg mirror for liquid substances monitoring” – 

Oral presentation. SPIE Europe, Optics and Optoelectronics, 2007, Prague, Czech 

Republic. 

C32. L. Rotiroti, E. De Tommasi, I. Rea, A. Lamberti, C. Sanges, I. Rendina, P. Arcari and L. 

De Stefano “Optical detection of PNA-DNA hybridization in resonant porous silicon based 

devices” – Poster presentation. Europtrode 2008, Dublin, Ireland. 

C33. E. De Tommasi, I. Rea. V. Di Sarno, L. Rotiroti, I. Rendina and L. De Stefano “Porous 

Silicon Resonant Mirrors Biosensors” – Poster presentation. Porous Semiconductors 

Science & Technology, 2008, Mallorca, Spain. 

C34. E. De Tommasi, I. Rea. V. Di Sarno, L. Rotiroti, I. Rendina and L. De Stefano “Porous 

Silicon Resonant Mirrors for Optical Biosensing” – Oral presentation. 13th National 

Conference AISEM, 2008, Rome.

C35. L. Rotiroti, E. De Tommasi, I. Rendina, M. Canciello, G. Maglio, R. Palumbo, and L. De 

Stefano,” BIOACTIVE POLIMERS MODIFIED POROUS SILICON NANOSTRUCTURES”, 

6th International Conference on Porous Semiconductuctors-Science and Technology, 

Poster presentation., Marzo 10-14, 2008, Mallorca, Spagna. 

C36. L. Rotiroti, P. Arcari, A. Lamberti, C. Sanges, E. De Tommasi, I. Rea, I. Rendina and L. 

De Stefano “Optical detection of PNA-DNA hybridization in resonant porous silicon based 

devices” – Poster presentation. SPIE Photonics Europe, 2008, Strasbourg, France. 

C37. L. De Stefano, E. De Tommasi, I. Rea, L. Rotiroti, I. Rendina “Porous Silicon resonant 

devices as optical transducers for nanostructured hybrid biosensors”, Invited, New 

Frontiers in Micro and Nano Photonics, 23-26 Aprile 2008, Firenze, Italia. 

C38. L. Rotiroti, I. Rendina, E. De Tommasi, M. Canciello, G. Maglio, R. Palumbo, and L. De 

Stefano, “A nanostructured hybrid device based on polymers infiltrated porous silicon layers 

for biotechnological applications”, oral presentation, PPS-24 Meeting The Polymer 

Processing Society 24th Annual Meeting June 15-19, 2008, Salerno, ITALY. 

C39. E. De Tommasi, I. Rea, V. Di Sarno, L. Rotiroti, I. Rendina and L. De Stefano “Porous 

silicon resonant mirrors in the high sensitive monitoring of biological interactions” – Oral 

presentation. First Mediterranean Photonics Conference of the European Optical Society, 

2008, Ischia (Naples), Italy. 

C40. L. Rotiroti, I. Rendina, E. De Tommasi, M. Canciello, G. Maglio, R. Palumbo, and L. De 

Stefano, A Hybrid Optical Biosensor Based on Polymer Infiltrated Porous Silicon Device, – 

Oral presentation, IEEE SENSORS 2008, Lecce (Italy). 

PATENTS 

Patent n. RM 2005 A 0002280 " An integrated device for the optical determination of the 

alcoholic content, in particular for wines and liquors, main procedures of fabrication, and main 

procedures of measurements" 

SCHOOLS AND WORKSHOPS: 

1. 11 giugno 2008 SPETTROSCOPY SEMPLIFIED – THERMO FISCHER SCIENTIFIC, Hotel 

Selene, Pomezia 

2. 6 febbraio 2008 LA BUONA PRATICA DI PESATA - GWP®, METTLER TOLEDO, Holiday 

Inn, Napoli. 

3. 17 luglio 2007 MICROSCOPIA A SONDA NELLA SCIENZA DEI MATERIALI, seminario 

indetto da 2M STRUMENTI “Veeco”, Portici (Napoli). 

4. 26 febbraio-2 marzo 2007 ADVANCED COURSE in design, synthesis, analysis and 

evaluation of bioactive molecules, Napoli. 

5. 5-6 Luglio 2006 Ciclo di seminari sulla Luce di Sincrotrone, Napoli. 

6. 26-27 giugno 2006 corso “ENZIMI DA IPERTERMOFILI E NANOTECNOLOGIE: ASPETTI 

APPLICATIVI”, Napoli. 

7. 19-21 Giugno 2006 scuola di dottorato del Gruppo Elettronica (GE 2006), Benevento (BN), 

Italia. 

8. 15-17 maggio 2006 corso teorico/pratico “FUNZIONALIZZAZIONE DI SUPERFICI VIA 

PLASMA PER APPLICAZIONI BIOTECNOLOGICHE”, Napoli. 

9. 12 marzo 2006 Short Course “5 th International Conference on Porous Semiconductuctors- 

Science and Technology”, Sitges-Barcellona, Spagna.

Appendix

INSTITUTE OF PHYSICS PUBLISHING JOURNAL OF OPTICS A: PURE AND APPLIED OPTICS 

J. Opt. A: Pure Appl. Opt. 8 (2006) S540–S544 doi:10.1088/1464-4258/8/7/S37 

Porous silicon-based optical biochips 

Luca De Stefano 1 ,LuciaRotiroti 1 ,IlariaRea 1 ,LuigiMoretti 2 , 

Girolamo Di Francia 3 ,Ettore Massera 3 ,Annalisa Lamberti 4 , 

Paolo Arcari 4 ,CarmenSanges 4 and Ivo Rendina 1 

1 Institute for Microelectronics and Microsystems, CNR—Department of Naples, 

Via PCastellino 111, 80131 Naples, Italy 

2 DIMET, University ‘Mediterranea’ of Reggio Calabria, Località Feo di Vito, 89060 Reggio 

Calabria, Italy 

3 ENEA—Centro Ricerche Portici, 80055 Portici (NA), Italy 

4 DBBM—Università diNapoli ‘Federico II’, Via S Pansini 5, 80131 Napoli, Italy 

Received 28 December 2005, accepted for publication 17 March 2006 

Published 12 June 2006 

Online at stacks.iop.org/JOptA/8/S540 

Abstract 

In this paper, we present our work on an optical biosensor for the detection of 

the interaction between a DNA single strand and its complementary 

oligonucleotide, based on the porous silicon (PSi) microtechnology. The 

crucial point in this sensing device is how to make a stable and repeatable 

link between the DNA probe and the PSi surface. We have experimentally 

compared some functionalization processes which modify the PSi surface in 

order to covalently fix the DNA probe on it: a pure chemical passivation 

procedure, a photochemical functionalization process, and a chemical 

modification during the electrochemical etching of the PSi. We have 

quantitatively measured the efficiency of the chemical bond between the 

DNA and the porous silicon surface using Fourier transform infrared 

spectroscopy (FT-IR) and light induced photoluminescence emission. From 

the results and for its intrinsic simplicity, photochemical passivation seems to 

be the most promising method. 

The interaction between a label-free 50 µMDNAprobe with 

complementary and non-complementary oligonucleotides sequences has been 

also successfully monitored by means of optical reflectivity measurements. 

Keywords: porous silicon, optical biosensor, surface functionalization, 

DNA analysis 



In the past two decades, the biological and medical fields 

have seen great advances in the development of biosensors and 

biochips for characterizing and quantifying biomolecules. 

Abiosensor can be generally defined as a device that 

consists of a biological recognition system, often called 

a bioreceptor, and a transducer. The interaction of the 

bioreceptor with the analyte is designed to produce an effect 

measured by the transducer, which converts the information 

into a measurable effect, such as an electrical signal. 

Using bioreceptors from biological organisms or receptors 

that have been patterned after biological systems, scientists 

have developed new methods of biochemical analysis that 

exploit the high selectivity of the biological recognition 

systems [1]. 

Porous silicon (PSi) is an almost ideal material as 

a transducer due to its porous structure, with hydrogen 

terminated surface, having a specific area of the order of 200– 

500 m 2 cm −3 ,sothataveryeffective interaction with several 

adsorbates is assured. Moreover, PSi is an available and low 

cost material, completely compatible with standard integrated 

circuit processes. Therefore, it could usefully be employed in 

the so-called smart sensors [2]. Recently, much experimental 

work, exploiting the noteworthy properties of PSi in chemical 

and biological sensing, has been reported [3–6]. 

PSi optical sensing devices are based on changes of its 

physical properties, such as photoluminescence or reflectance, 

on exposure to the surrounding environment. Unfortunately, 

this interaction is not specific, so a PSi sensor cannot 

discriminate the components of a complex mixture. Some 

researchers have chemically or physically modified the Si–H 

1464-4258/06/070540+05$30.00 © 2006 IOP Publishing Ltd Printed in the UK S540

White light 

source 

Gas channels 

Sealed 

or fluxed 

reaction 

chamber 

Glass 

Silicon 

Optical fibres 

Collimator 

Objective 

Optical 

Spectrum 

Analyser 

(OSA) 

Figure 1. The experimental optical set-up. 

Porous silicon 

microcavities 

surface sites in order to enhancethe sensor selectivity through 

specific interactions. The common approach is to create a 

covalent bond between the PSi surface and the biomolecules 

which specifically recognize the target analytes [7, 8]. 

The reliability of a biosensor strongly depends on the 

functionalization process: how fast, simple, homogenous and 

repeatable it is. This step is also very important for the stability 

of the sensor: it is well known that ‘as-etched’ PSi has a Si– 

Hterminatedsurface due to the Si dissolution process which 

is very reactive [9]. The substitution of the Si–H bonds with 

Si–C ones guarantees a much more stable surface from the 

thermodynamic point of view. 

In this work, we report some PSi surface modification 

strategies in order to realize an optical biosensor: the target 

is the fabrication of sensitive label-free biosensors, which 

are highly requested for applications in high throughput 

drug monitoring and disease diagnostics; unlabelled analytes 

require in fact easier and faster analytical procedures. 


Porous silicon is a very attractive material due to the possibility 

of fabricating high quality optical structures, either as single 

layers, like Fabry–Perot interferometers, or multilayers, such 

as Bragg or rugate filters [12]. In this study we used as 

a sensor a Fabry–Perot single layer. The PSi layer was 

obtained by electrochemical etch in a HF-based solution at 

room temperature. A highly doped p + -silicon substrate, 〈100〉 

oriented, 0.01 cm resistivity, 400 µmthickwasused. Before 

anodization, the substrate was placed in HF solution to remove 

the native oxide. 

The dielectric and physical properties of the PSi layer 

were investigated by spectroscopic ellipsometry and scanning 

electron microscopy. Both the analysis methods showed the 

presence of a top layer on the PSi structure of about 50 nm 

thickness, characterized by a very low porosity. The top layer 

is due to a hydrogen contamination which passivates boron 

and changes the local resistivity. The presence of the top 

layer reduces pore infiltration. To remove the contamination 

we pre-treated the crystalline silicon substrate before the 

electrochemical etching by heating the wafer at 300 ◦ CinN2 

atmosphere for 30 min [10]. The ellipsometric measurements 

on the pre-treated PSi layer show that the top layer was 

eliminated. 


1.0 

0.8 

0.6 

0.4 

0.2 

PSi as etched 

After KOH treatment 

Ater HF treatment 


0.0 

600 800 1000 1200 


Figure 2. Reflectivity optical spectra of the porous silicon layer after 

KOH and HF treatments. 

Due to the nanostructured nature of PSi, it is necessary 

to improve the pore infiltration of biomolecular probes: with 

this aim, we optimized our device by employing a strong base 

post-etch process. We immersed the device in an aqueous 

ethanol solution, containing millimolar concentration of KOH, 

for 15 min. This treatment produces an increase of about 15– 

20% in the porosity without affecting the optical quality of the 

device [11]. 

We used a very simple experimental set-up (see figure 1) 

to characterize the PSi device from an optical point of view: 

atungsten lamp (400 nm

LDeStefanoet al 

Reflectance 

40 

35 

30 

25 

20 

15 

10 

5 

CH 2 

Si-H 

Si-C 

0 

4000 3500 3000 2500 2000 1500 1000 500 


with EtMgBr 

After post-etch treatments 

PSi H 

CH3CH2 Mg Br 

@ 85°C 

PSi CH 2CH 3 

Figure 3. FT-IR spectra of the PSi monolayer before and after the pure chemical functionalization process based on EtMgBr, together with 

the reaction scheme. 

% Reflectance 

160 

140 

120 

100 

80 

60 

40 

20 

2854.7 CH 2 

2926 CH 2 

2100 Si-H 

1563 succinimmide 

1038 N-O 

1641 succinimmide 

1723 O-C=O 

0 

-20 

1288 C-N C=O 

4000 3000 2000 1000 

Wavenumbers (cm -1 after post-etch treatments 

after UV exposition with UANHS 

) 

PS i H 

UANHS 

UV 

O 

H 

O 

PSi C 

H 

(CH2) n C O N 

O 

DNA 

PSi 

H 

C (CH2 ) n 

H 

O 

C 

H 

N DNA 

Figure 4. FT-IR spectra of the PSi monolayer before and after the photoinduced functionalization process based on UV exposure, together 

with the reaction scheme. 

2.1. Chemical functionalization by Grignard reactives 

The chemical functionalization is based on ethyl magnesium 

bromide (CH3CH2MgBr) as a nucleophilic agent which 

substitutes the Si–H bonds with the Si–C. The reaction was 

performed at 85 ◦ Cinaninert atmosphere (argon) to avoid the 

deactivation of the reactive, for 8 h. The chip was then washed 

with a 1% solution of CF3COOH in diethyl ether, and then with 

deionized water and pure diethyl ether. The modified surface 

chip was characterized by infrared spectroscopy to verify the 

efficiency of the method. The FT-IR spectrum and the reaction 

scheme are reported in figure 3. 

2.2. Photochemical functionalization 

The photo-activated chemical modification of the PSi surface 

was based on the UV exposure of a solution of alkenes which 

bring some carboxylic acid groups. The PSi chip was precleaned 

in an ultrasonic acetone bath for 10 min then washed 

in deionized water. After being dried in a N2 stream, it was 

immediately covered with 10% N-hydroxysuccinimide ester 

(UANHS) solution in CH2Cl2. The UANHS was synthesized 

in house as described in [7]. This treatment results in covalent 

attachment of UANHS to the PSi surface, clearly shown in the 

FT-IR spectrum, reported in figure 4 together with the reaction 

scheme. The chip was then washed in dichloromethane in an 

S542 

ultrasonic bath for 10 min and rinsed in acetone to remove any 

adsorbed alkene from the surface. 

2.3. Functionalization during the etch process 

We also tried to chemically modify the surface of the PSi 

by directly using a functionalizing agent during the etching 

process. We introduced some organic acids (eptinoic and 

pentenoic acid with concentrations from 0.4 M to 3 M) in an 

electrochemical cell in the presence of a dilute HF solution 

(HF:EtOH=1:2). Inthiscase a current density of 60 mA cm −2 

was applied to etch an area of 0.07 cm 2 .TheFT-IR spectrum 

is reported in figure 5. 

3. Experimental results 

We studied the infrared spectrum in each case of functionalization 

method in order to determine the best reaction conditions 

and the maximum yield in the Si–H/Si–C substitution. 

The post-etch KOH treatment, which we used to increase the 

porosity and to improve the pore infiltration, removes most of 

Si–H bonds on the fresh PSi surface and also induces its oxidation. 

To assure the formation of the Si–C bonds, we had 

first to restore the Si–H bonds by rinsing the PSi in a very dilute 

HF-based solution for 30 s. FT-IR spectroscopy confirmed 

the presence of Si–H bonds on the PSi surface after all pretreatments.

% Reflectance 

110 

100 

90 

80 

70 

60 

50 

40 

30 

20 

10 

before functionalization 

after functionalization 

2915.4CH2 

2383.7 CO2 

2348.1 CO2 

3500 3000 2500 2000 1500 1000 

Wavenumber (cm –1 ) 

2104SiH 

2064.2SiH 

1709.4COOH 

1075 SiO 

1260 CH=CH 

1032 SiO 

903 

867.4 

Figure 5. FT-IR spectra of the PSi monolayer before (solid line) and 

after (dashed line) functionalization during the etching process. 

For each method we studied, we report the infrared 

spectrum (figures 3–5). In all of them, the characteristic peaks 

of the Si–C bonds can be observed, even if with different 

intensities. In the case of the pure chemical functionalization 

method, the FT-IR spectrum (figure 3) showsanincomplete 

substitution of the Si–H bonds, since their characteristic peaks 

(at 2100 cm −1 )are still present. 

In figure 4 we can also easily distinguish all the 

characteristic peaks of the reactive (UANHS) immobilized on 

the PSi surface. The infrared analysis of the chip prepared 

by direct functionalization during the etch process, shows 

a–COOH peak, whose intensity increases on increasing the 

acid concentration, as reported in figure 5. 

Among the three procedures experimented, we believe 

that the photoinduced method is the best one for several 

reasons: the relaxed reaction conditions (atmospheric pressure 

and room temperature); the shorter reaction time; and the best 

reaction yield (from FT-IR measurements). This last result is 

somewhat expected because the reactive considered has a socalled 

‘outgoing group’ which promotes its substitution with 

the amine group of the DNA probe. 

In view of these considerations, on each chip with a 

photochemical modified surface we incubated, overnight, a 

luminescent DNA probe. After washing, we measured the 

photoluminescence emission with a microscope equipped with 

a CCD. As a reference sample, we used the signal coming from 

chips not washed after labelled DNA incubation. 

The PL measurements showed that the photochemical 

modification of the PSi surface is very efficient: we registered 

52 ± 8counts from treated chips, 5 ± 1counts from the 

unfunctionalized chips, and 63 ± 6counts from the reference 

samples. The count values are the arithmetic mean over 

three different chip samples and the errors are the standard 

deviations. A total yield of about 83% can be thus estimated. 

The optical monitoring of DNA–cDNA hybridization is a 

two-step procedure: first, we registered the optical spectrum 

of the PSi layer after the UANHS and probe immobilization 

on the chip surface and, then, after the hybridization with 

the cDNA. Each step of the chip preparation increases the 

optical path in the reflectivity spectrum recorded, due to the 

substitution of the air into the pores by organic and biological 

compounds, so that several red-shifts can be observed after the 


1.0 

0.8 

0.6 

0.4 

0.2 


PSi with DNA probe 

PSi with c-DNA 

600 800 1000 


1200 

Figure 6. Reflectivity optical spectra of the PSi device after DNA 

probe immobilization and after cDNA hybridization. 

functionalization (10 nm), and after the covalent binding of the 

DNAsingle strand (probe) with the linker on the porous silicon 

layer (31 nm). The interaction with the cDNA is also detected 

as a shift in wavelength of the optical path (33 nm). In figure 6 

the reflectivity spectra of only these last two steps are reported 

for the sake of clarity. A control measurement was made using 

an ncDNA sequence: a very small shift (less than 2 nm) was 

recorded in the reflectivity spectrum with respect to the one 

obtained after probe linking. 

4. Conclusion 

In conclusion, we have presented our experimental results 

of the study of several procedures of porous silicon surface 

functionalization that are a fundamental step in realizing an 

optical porous silicon microsensor for DNA–cDNA interaction 

detection. The pore infiltration by the DNA single-strand 

solution was improved by introducing two more steps during 

the realization process: a pre-etch cleaning of the crystal 

substrate to remove hydrogen contamination; and an alkaline 

post-etching process that increases the device porosity. Among 

several surface functionalization methods used to immobilize 

in the PSi matrix a DNA probe which selectively recognizes 

its complementary sequence, we chose a photoinduced one 

because of the simplicity of the reaction conditions and 

the well performing results. We also demonstrated that 

an optical Fabry–Perot interferometer based on a PSi layer 

can be effectively used as a transducer of the DNA–cDNA 

hybridization interaction. 


This work was realised in the frame of the CRdC-NTAP and 

in the frame of the MIUR FIRB Project 2003 Costituzione 

di un laboratorio dedicato all’analisi delle interazioni ligandorecettore 

mediante biochip di silicio. 

References 

[1] Sadana A 2002 Engineering Biosensors (San Diego, CA: 

Academic) 

S543

LDeStefanoet al 

[2] De Stefano L, Moretti L, Rendina I, Rossi A M and 

Tundo S 2004 Smart optical sensors for chemical substances 

based on porous silicon technology Appl. Opt. 43 167–72 

[3] Dancil K-P S, Greiner D P and Sailor M J 1999 A porous 

silicon optical biosensor: detection of reversible binding of 

IgG to a protein A-modified surface J. Am. Chem. Soc. 

121 7925–30 

[4] De Stefano L, Moretti L, Rossi A M, Rocchia M, Lamberti A, 

Longo O, Arcari P and Rendina I 2004 Optical sensors for 

vapors, liquids, and biological molecules based on porous 

silicon technology IEEE Trans. Nanotech. 3 49–54 

[5] Dattelbaum J D and Lakowicz J R 2001 Optical determination 

of glutamine using a genetically engineered protein Anal. 

Biochem. 291 89–95 

[6] Bradford M M 1976 A rapid and sensitive method for the 

quantitation of microgram quantities of protein utilizing the 

principle of protein-dye binding’ Anal. Biochem. 72 248–54 

S544 

[7] Yin H B, Brown T, Gref R, Wilkinson J S and Melvin T 2004 

Chemical modification and micropatterning of Si(100) with 

oligonucleotides Microelectron. Eng. 73/74 830–6 

[8] Hart B R, Letant S E, Kane S R, Hadi M Z, Shields S J and 

Reynolds J G 2003 New method for attachment of 

biomolecules to porous silicon Chem. Commun. 322–3 

[9] Canham L (ed) 1997 Properties of Porous Silicon (London: 

IEE INSPEC) 

[10] Chamard V, Dolino G and Muller F 1998 Origin of a parasitic 

surface film on p+ type porous silicon J. Appl. Phys. 

84 6659–66 

[11] DeLouise L A and Miller B L 2004 Optimization of 

mesoporous silicon microcavities for proteomic sensing 

Mater. Res. Soc. Symp. Proc. 782 A5.3.1–A5.3.7 

[12] Theiss W 1997 Optical properties of porous silicon Surf. Sci. 

Rep. 29 91–192

IOP PUBLISHING JOURNAL OF PHYSICS: CONDENSED MATTER 

J. Phys.: Condens. Matter 19 (2007) 395008 (7pp) doi:10.1088/0953-8984/19/39/395008 

Optical microsystems based on a nanomaterial 

technology 


LDeStefano 1 ,LRotiroti 1,2 ,IRea 1,3 ,MIodice 1 and I Rendina 1 

1 National Council of Research-Institute for Microelectronic and Microsystems-Department of 

Naples, Via P Castellino 111, 80131 Naples, Italy 

2 Department of Organic Chemistry and Biochemistry, ‘Federico II’ University of Naples, Via 

Cinthia, 4-80126 Naples, Italy 

3 Department of Physics, ‘Federico II’ University of Naples, Via Cinthia, 4-80126 Naples, Italy 

E-mail: luca.destefano@na.imm.cnr.it 

Received 13 February 2007 

Published 30 August 2007 

Online at stacks.iop.org/JPhysCM/19/395008 

Abstract 

In this work, we present an optical sensor for quantitative determination of the 

alcohol content in hydro-alcohol mixtures, realized by using porous silicon 

(PSi) nanotechnology. The device is an oxidized PSi micro-cavity (PSMC) 

constituted by a Fabry–Perot layer between two distributed Bragg reflectors. 

Due to the capillary condensation, a red shift of the PSMC reflectivity spectrum 

is observed on exposure to vapour mixtures. The phenomenon is completely 

reversible. Moreover, to reduce the analysis time, we have designed the 

integration of the sensor in a thermally controlled lab-on-chip, by merging PSi 

and anodic bonding technologies. Numerical calculations have been performed 

to study the thermal behaviour of the integrated device. 


The fast and intensive development of the lab-on-chip (LOC) for sensing applications is due 

to several factors that are essential for routine analytical determinations, such as very limited 

sample consumption and short analysis time, which can be obtained by measuring a transient 

signal in a flow-through detector. Due to their intrinsic features, in particular contactless 

monitoring capability, optical sensors are very attractive for many application fields. 

In recent years, many devices based on porous silicon (PSi) technology have been realized 

for several sensing applications [1, 2]. PSi, fabricated by electrochemical etching of a silicon 

(Si) wafer in HF-based solution, is an ideal material for an optical transducer, due to its spongelike 

nanostructure (specific surfaces up to 500 m 2 cm −3 )[3], that assures an effective interaction 

with several chemicals and biological molecules. On exposure to chemical substances, several 

physical parameters, such as refractive index, photoluminescence, and electrical conductivity, 

0953-8984/07/395008+07$30.00 © 2007 IOP Publishing Ltd Printed in the UK 1

J. Phys.: Condens. Matter 19 (2007) 395008 L De Stefano et al 

Figure 1. Experimental optical set-up used for the sensing experiments. 

change drastically. It is also possible to realize in a single process multilayered structures that 

exhibit a high selective optical response, which is very attractive for sensing [4]. 

Silicon–glass anodic bonding (AB) is one of the most promising techniques in the 

micromachining field; it easily allows the encapsulation and sealing on silicon chips of threedimensional 

microfluidic structures, such as channels, chambers, cavities and other complex 

gas and liquid routes. Additionally, glass transparency at optical wavelengths enables simple, 

but highly accurate, alignment of pre-patterned or structured glass and silicon wafers. 

We have recently demonstrated the compatibility of PSi and AB technologies for the 

realization of sensing microcomponents for LOC applications [5, 6]. The integration was 

not easy or straightforward: the microfabrication techniques are typically optimized for the 

microelectronic industry and are not well appropriated or compatible with the non-standard 

materials such as the ones used in biological or chemical sensing. 

We present here a simple optical sensor for the analysis of hydro-alcohol mixtures, based 

on PSi nanotechnology. Faster time analysis can be obtained by integrating the silicon–glass 

sensor for LOC applications by merging the PSi and AB technologies. 


For this study, we fabricated a λ/2 Fabry–Perot optical microcavity sandwiched between 

two Bragg reflectors of seven periods each, realized by alternating layers with high and low 

refractive indices. The porous silicon microcavity (PSMC) was obtained by electrochemical 

etch in a HF-based solution (50 wt% HF:ethanol = 1:1) in dark light and at room temperature. 

A current density of 400 mA cm −2 for 0.8 s was applied to a highly doped p + -type standard 

silicon wafer, 〈100〉 oriented, with 10 m cm resistivity, 400 μm thick, producing the highrefractive 

index layer (with a porosity of 75%), while a current density of 80 mA cm −2 was 

applied for 1.9 s in the case of the low-index layer (with a porosity of 67%). The λ/2 layer 

is a low-refractive-index one. This structure has a characteristic resonance peak at 1110 nm 

in the middle of a 280 nm wide stop band. To ensure the infiltration of aqueous solution, the 

device, which (as etched) is hydrophobic due to the presence of SiH bonds on its surface, was 

made hydrophilic by thermal oxidation at 1000 ◦ C for 30 min. The optical set-up required 

for our sensing experiments was very simple (see figure 1). The PSMC was placed in a test 

chamber equipped with a glass window. A white light source interrogated the sensor through an 

optical fibre and a collimator. The reflected beam was collected by an objective, coupled into 

2


Figure 2. Optical spectra of the PSMC as etched and after thermal oxidation. 

a multimode fibre, and then directed in an optical spectrum analyser (Ando model AQ6315A). 

The reflectivity spectra were measured over the range 600–1600 nm with a resolution of 0.2 nm. 

We studied the resonance peak shift on exposure to an ethanol/deionized water (DI) binary 

mixture at different concentrations, placing a drop of solution in the test chamber. Due to 

the capillary condensation of the solution vapours into the pores of the PSMC, a change in 

its reflectivity spectrum was observed: the partial substitution of air by liquid in each layer 

determines an increase of the average refractive index of the microcavity and, consequentially, 

a red shift of the spectrum. Measurements are stored when equilibrium conditions are reached, 

i.e. when the signal does not change any more in time. The phenomenon is completely 

reversible and reproducible. 

3. Experimental results and discussion 

In figure 2, the reflectivity spectra of the PSMC as etched and after thermal oxidation are 

reported. The oxidation process causes a blue shift in the spectrum of 110 nm due to the 

lower value of the SiO2 refractive index (nSiO2 ≈ 1.45) with respect to the Si refractive index 

(nSi ≈ 3.6). We have not observed degradation in cavity quality (Q = λ/λ ≈ 70 in both 

cases). 

A very simple method, commonly used in analytical chemistry, in quantitative 

determinations of binary mixtures is the external standard method: calibration curves can be 

obtained by measuring the parameter under monitoring, such as an optical intensity, while 

the content of one component varies between 0 and 100%. In our case, the curve has been 

obtained by plotting the peak shift measured as a function of the volume percentage of ethanol 

in the binary mixture, as shown in figure 3. The errors on the experimental points are the 

standard deviations on five measurements, using the same chip and the same hydro-alcohol 

solution. After the calibration, unknown concentrations of the same mixture can be quantitative 

determined with a single measurement of the peak shift: the volume concentrations of both 

components can be picked out from the curve obtained by interpolating the experimental 

calibration data. In order to limit the wide error bars and the time necessary for the PSMC 

to reach equilibrium with the mixture (by limiting the volume), integration in a single chip of 

the device is proposed. 

The compatibility of the porous silicon technology with the standard integrated circuit 

fabrication processes allows the design of complex microsystems in which the porous silicon 

3


Figure 3. The calibration curve in case of an ethanol/DI binary mixture. 

Figure 4. Sketch of the alcohol microsensor. 

optical device is the transduction element. In figure 4, we report a rough and somehow intuitive 

sketch of the alcohol microsensor. The PSMC can be integrated in an LOC consisting of two 

μ-chambers thermally controlled by integrated heaters, and sealed by a cover glass after the 

anodic bonding process. The mixture is injected in the left μ-chamber by means of an 

appropriate microsyringe and here the integrated heater induces the hydro-alcohol mixture 

evaporation. The vapours produced diffuse and reach the second μ-chamber, cooled by another 

integrated element, where the PSMC is and the capillary condensation happens. 

4. Numerical calculations 

Before proceeding towards the real fabrication steps of the whole device depicted in figure 4, 

it is mandatory to simulate its thermal behaviour since this is one of the key features in 

the measurement process. The precision and the accuracy in the quantitative determination 

of the mixture concentration strongly depend on the thermodynamic parameters such as the 

equilibrium time, i.e. the time required to have the same evaporation and condensation into 

the PSMC, and the temperature distribution in each layer used to realize the LOC. In order 

to simulate the thermal response of the device, a finite element model has been implemented 

with the commercial FEM code ANSYS Multiphysics 9.0 [7]. We have restricted the analysis 

to a two-dimensional (2D) model corresponding to the device middle cross section sketched 

(not to scale) in figure 4. Ignoring the minimal border effects at the device ends, a temperature 

4


Figure 5. Temperature distribution in the device when the tank μ-chamber is heated with a power 

density of 1.0Wcm −2 . 

Table 1. Physical parameters for lab-on-chip materials used in thermal simulations. 

Thermal conductivity Density Specific heat 

Material (Wm −1 K −1 ) (kg m −3 ) (Jkg −1 K −1 ) 

Silicon 156 2640 917 

Glass 1.05 2500 840 

Water 0.58 1000 4186 

Air 0.025 1.225 1005 

gradient in fact exists only in this plane. This hypothesis can be justified by the observation that, 

for the proposed geometry, the aspect ratio between its width (46 mm) and thickness (1.5 mm) 

is about 1/30. Of course, a full 3D model could be more accurate, though it is very time 

consuming during the FEM analysis. The structure has been meshed with 2242 isoparametric 

linear elements of the type plane 55, with a mesh refinement in the active regions of the device. 

In table 1 we have reported the main physical parameters of the materials used for ANSYS 

thermal simulations. 

The results of the numerical simulations depend on the boundary conditions fixed: we have 

supposed a natural convective heat exchange on the upper face of the glass slide which is in 

contact with air at 20 ◦ C. We have assumed an exchange coefficient of 27 W m −2 K −1 [8]. 

Lateral and lower surfaces have been considered adiabatic; this model could be a good 

approximation of a device mounted on a plastic case. The lower surface of the left μ-chamber, 

which is our tank for the liquid substance, is kept hot at constant temperature by a heater 

element. The right μ-chamber is cooled by a Peltier μ-cell with a maximum heat flux equal 

to 0.4 Wcm −2 (i.e. micro modules from TE Technology Inc., USA). In figure 5 we report the 

2D steady-state thermal distribution, when the tank chamber is heated with a power density of 

1.0 Wcm −2 , corresponding to a maximum temperature on the hot side of about 100 ◦ C. From 

this calculation it is clear that the PSMC region reaches a temperature of about 50 ◦ C, which 

is a little bit higher of the evaporation temperature of ethanol; in this case the alcohol cannot 

condense inside the pores of the PSMC. 

5


Figure 6. Temperature distribution versus the lateral dimension of the device for two different 

heating power densities. 

Figure 7. Temperature dynamic in the condensation chamber for two different heating power 

densities. 

We have repeated the simulation considering a lower heating power density, 0.4 Wcm −2 , 

corresponding to heating the tank μ-chamber to 35 ◦ C. In this case, the PSMC is cooled, 

thanks to the Peltier module, at about 5 ◦ C: the analysis chamber temperature decreases quite 

linearly with the heat source temperature, as expected. These results are shown schematically in 

figure 6, where the 1D lateral temperature distribution, calculated at the silicon–glass interface, 

is reported versus the lateral dimension of the device for the two different heating power 

densities. We have also calculated the PSMC average temperature as a function of time for 

the two different heating power densities. The transient solver Antype4 wasusedtoperform 

6


the analysis, using the multi load-step method provided in the code. From the plot in figure 7, 

it is well evident that, in the case of a heating power of 0.4Wcm −2 , the porous silicon–vapour 

system is at thermal equilibrium after about 50 s. In contrast, if we apply a heating power of 

1.0Wcm −2 , a much longer time is necessary. 


We have presented a simple and well performing optical sensor for ethanol/DI binary mixture 

analysis, based on PSMC. Due to capillary condensation, the reflectivity spectrum of the device 

shifts towards higher wavelengths on exposure to vapours. The shift is characteristic of the 

concentration of each component of the mixture. The sensing process is completely reversible. 

In order to limit the time needed for the PSMC to reach equilibrium with the mixture, we have 

proposed the integration of the device in a single chip that is thermally controlled. The thermal 

behaviour of the designed LOC has been studied by proper FEM numerical calculations. 

References 

[1] Dancil K-P S, Greiner D P and Sailor M J 1999 J. Am. Chem. Soc. 121 7925–30 

[2] De Stefano L, Moretti L, Rossi A M, Rocchia M, Lamberti A, Longo O, Arcari P and Rendina I 2004 IEEE Trans. 

Nanotech. 3 49–54 

[3] Herino R, Bomchil G, Barla K, Bertrand C and Ginoux J L 1987 J. Electrochem. Soc. 134 1994–2000 

[4] Theiss W 1997 Surf. Sci. Rep. 29 95–192 

[5] Malecki K and Della Corte F G 2005 Proc. Micromachining and Microfabrication Process Technology X; SPIE 

5715 180–9 

[6] De Stefano L, Malecki K, Della Corte F G, Moretti L, Rea I, Rotiroti L and Rendina I 2006 Sensors 6 680–7 

[7] Ansys 9.0 user manual see http://www.ansys.com 

[8] Sturm J C, Wilson W and Iodice M 1998 IEEE J. Sel. Top. Quantum Electron. 4 75–82 

7

Extremophiles 

DOI 10.1007/s00792-006-0058-6 

ORIGINAL PAPER 

Enzymes and proteins from extremophiles as hyperstable probes 

in nanotechnology: the use of D-trehalose/D-maltose-binding 

protein from the hyperthermophilic archaeon Thermococcus 

litoralis for sugars monitoring 

Luca De Stefano Æ Annalisa Vitale Æ Ilaria Rea Æ Maria Staiano Æ 

Lucia Rotiroti Æ Tullio Labella Æ Ivo Rendina Æ Vincenzo Aurilia Æ 

Mose’ Rossi Æ Sabato D’Auria 

Received: 30 October 2006 / Accepted: 29 November 2006 

Ó Springer 2007 

Abstract The D-trehalose/D-maltose-binding protein 

(TMBP), a monomeric protein of 48 kDa, is one 

component of the trehalose and maltose (Mal) uptake 

system. In the hyperthermophilic archaeon Thermococcus 

litoralis, this is mediated by a protein-dependent 

ATP-binding cassette system transporter. The gene 

coding for a thermostable TMBP from the archaeon T. 

litoralis has been cloned, and the recombinant protein 

has been expressed in E. coli. The recombinant TMBP 

has been purified to homogeneity and characterized. It 

exhibits the same functional and structural properties 

as the native one. In fact, it is highly thermostable and 

binds sugars, such as maltose, trehalose and glucose, 

with high affinity. In this work, we have immobilized 

TMBP on a porous silicon wafer. The immobilization 

of TMBP to the chip was monitored by reflectivity and 

Fourier Transformed Infrared spectroscopy. Furthermore, 

we have tested the optical response of the protein-Chip 

complex to glucose binding. The obtained 

Communicated by D. A. Cowan. 

The authors wish to dedicate this work to Prof. Ignacy 

Gryczynski, University of North Texas, TX, USA, for his 

outstanding contribution to the development of new sensing 

methodologies. 

L. De Stefano I. Rea L. Rotiroti I. Rendina 

Istituto di Microelettrica e Microsistemi, CNR, Napoli, Italy 

A. Vitale M. Staiano T. Labella V. Aurilia 

M. Rossi S. D’Auria (&) 

Istituto di Biochimica delle Proteine, CNR, 

Via Pietro Castellino 111, 80131 Napoli, Italy 

e-mail: s.dauria@ibp.cnr.it 

data suggest the use of this protein for the design of 

advanced optical non-consuming analyte biosensors for 

glucose detection. 

Keywords Trehalose/maltose-binding protein 

Biosensors Porous silicon Glucose Diabetes 

Archaeon 

Introduction 

The general interest in biomolecules, isolated from 

thermophilic organisms was originally due to the biotechnological 

advantages offered by the utilization of 

these highly stable molecules in industrial processes 

(Brock 1985). In fact, enzymes and proteins isolated 

from thermophilic microorganisms exhibit a high stability 

in conditions usually used to denature proteins: 

high temperature, ionic strength, extreme pH-values, 

elevated concentration of detergents and chaotropic 

agents (Jaenicke 1994). In addition to the potential 

industrial applications, it is important to highlight that 

proteins and enzymes that are stable and active over 

100°C represent good models to shed light on the 

molecular adaptation of life at high temperature 

(D’Auria et al. 1998). 

The D-trehalose/D-maltose-binding protein 

(TMBP) is one component of the trehalose (Tre) and 

maltose (Mal) uptake system, which, in the hyperthermophilic 

archaeon Thermococcus litoralis, is mediated 

by a protein-dependent ATP-binding cassette 

(ABC) system transporter (Xavier 1996). TMBP from 

T. litoralis is a monomeric 48 kDa two-domain macromolecule 

containing 12 tryptophan residues (Diez 

et al. 2001). TMBP shares common structural motifs 

123

with a number of other sugar-binding proteins. This 

class of biomolecules is composed of proteins, whose 

structure consists of two globular domains connected 

by a hinge region made of two or three short peptide 

segments. The two domains are formed by non-contiguous 

polypeptide stretches and exhibit similar tertiary 

structure. The ligand-binding site is located in the 

deep cleft between the two domains and the binding is 

accompanied by a movement of the two lobes as well 

as by conformational changes in the hinge region (Sun 

et al. 1998). 

The gene coding for the thermostable TMBP from 

the archaeon T. litoralis has been cloned, and the recombinant 

protein has been expressed in E. coli 

(Horlacher et al. 1998). The recombinant TMBP has 

been purified to homogeneity and characterized. It 

exhibits the same functional and structural properties 

as the native one (Horlacher et al. 1998). 

In a recent work (Herman et al. 2006), we have 

investigated the binding of several sugars to TMBP by 

time-resolved fluorescence spectroscopy, and molecular 

dynamics methods. We found that TMBP is also 

able to bind glucose molecules. Since human blood 

does not contain Tre and Mal, it is not outrageous to 

envisage the use of the TMBP as a probe for the design 

of a minimally invasive biochip for glucose detection. 

In this work, we have extended our investigation on 

the utilization of TMBP for sensing glucose by immobilizing 

the protein on a porous silicon chip. The 

TMBP-based chip has been characterized and tested as 

regards the response to glucose. The obtained results 

are discussed. 


D-Glucose and all the other chemicals used in the 

present study were from Sigma. All commercial samples 

were of the best available quality. 

Purification of TMBP and protein concentration 

determination 

An mount of 20 g wet weight of the E. coli cell pellet 

containing the expressed TMBP were resuspended in 

50 mM Tris–HCl, pH 7.5, 500 mM NaCl (100 ml), 

ruptured by a French pressure cell at 16,000 psi, and 

centrifuged for 15 min at 19,000 g. The supernatant 

was heated to 90°C for 20 min and centrifuged for 

15 min at 19,000 g. The supernatant was collected and 

dialyzed for 12 h against 10 mM Tris–HCl, pH 7.5 at 

4°C. The solution was loaded onto a DEAE column 

previously equilibrated in 10 mM Tris–HCl, pH 7.5. 

123 

After washing the column with 10 mM Tris–HCl, pH 

7.5, TMBP was eluted by a NaCl gradient (0.0–1.0 M 

NaCl). Centricon 30 concentrator (Amicon, MA, 

USA) was used to concentrate and to dialyze TMBP 

against 10 mM Tris–HCl, pH 7.5. The purified protein 

was passed through a Blue A affinity column (Amersham, 

NJ, USA) in order to obtain DNA-free TMBP. 

The purity of TMBP was verified by SDS-PAGE and 

absorption spectra. 

The protein concentration was determined by the 

method of Bradford (Bradford 1976) with bovine 

serum albumin as standard on a double beam Cary 

1E spectrophotometer (Varian, Mulgrade, Victoria, 

Australia). 

Chip preparation and spectroscopic 

characterization 

In this study, we used as sensor an apodized Bragg 

reflector obtained by alternating high (A) refractive 

index layers (low-porosity) and low (B) refractive index 

layers (high-porosity) whose thicknesses satisfy the 

following relationship: n AdA +nBdB =mk/2, where m 

is an integer and k is the Bragg resonant wavelength. 

The distributed Bragg reflector (DBR) was produced 

by electrochemical etch of a highly doped p + -silicon, 

Æ100æoriented, 0.04 W cm resistivity, 400 lm thick, in a 

HF-based solution. The silicon was etched using a 

30 wt.%HF/ethanol solution in dark and at room 

temperature. Before anodization the substrate was 

placed in HF solution to remove the native oxide. A 

current density of 148 mA/cm 2 for 2.31 s was applied 

to obtain the low refractive index layer (effective 

refractive index n L @ 1.505, thickness d L @ 598 nm) 

with a porosity of 72%, while one of 110 mA/cm 2 was 

applied for 2.34 s for the high index layer (nH @ 1.585, 

thickness d H @ 568 nm) with a porosity of 69%. This 

structure is characterized by a first order (m =1) 

resonance at 3,600 nm. In Fig.1, the experimental 

White Light 

Source 

OSA 

Y fiber 

fiber 

Porous Silicon Chip 

Extremophiles 

Fig. 1 Experimental setup used to measure the optical reflectivity 

spectrum of porous silicon chip

Extremophiles 

setup, used to measure the reflectivity spectrum, is 

reported. We have used as source a white light directed 

on the porous silicon chip through a Y-fiber. The same 

fiber was used to guide the output signal to the optical 

spectrum analyzer. The spectrum was measured over 

the range 800–1,600 nm with a resolution of 0.2 nm. 

The porous silicon surface after electrochemical etching 

is hydrogenated and for this reason, very reactive 

and thermodynamically instable; in order to realize a 

sensor more stable and able to link the biological 

probes it is necessary to substitute the Si–H bonds with 

the Si–C or Si–O–C ones. 

To promote a covalent link between the porous 

silicon and the TMBP, we have exploited a three-step 

functionalization process, based on the chemical passivation 

of the PSi surface after oxidation. We have 

first thermally treated the PSi DBR, in O 2 atmosphere, 

at 1,000°C for 30 min, to remove all the Si–H bonds 

and create an oxide layer on the pores’ surface to assure 

the covalent attachment with a proper chemical 

linker, the minopropyltriethoxysilane (APTES). To 

this aim, we have rinsed by immersion the DBR in a 

5% solution of APTES and a hydroalcoholic mixture 

of water and methanol (1:1), for 20 min at room temperature. 

After the reaction time, we have washed the 

chip with DI-water and methanol and dried in N 2 

stream. The silanized device was then baked at 100°C 

for 10 min. The next step consists in creating a surface 

able to link the carboxylic group of the proteins: we 

have thus immersed the DBR in a 2.5% glutaraldehyde 

solution in 20 mM HEPES buffer (pH 7.4) for 30 min, 

and then rinsed it in DI-water and finally dried in N 2 

stream. The glutaraldehyde reacts with the amino 

Fig. 2 Optical reflectivity 

spectrum of the Bragg 

reflector as-etched (continues 

curve), after oxidation 

(dashed curve), after APTES 

treatment (dot curve) and 

after glutaraldehyde 

treatment (dot-dashed curve) 

Reflectivity (u.a.) 

1.2 

1.1 

1.0 

0.9 

0.8 

0.7 

0.6 

0.5 

0.4 

0.3 

0.2 

0.1 

groups on the silanized surface and coats the internal 

surface of the pores with another thin layer of molecules. 

We have monitored all the reaction steps by Fourier 

Transformed Infrared (FT-IR) Spectroscopy and the 

consequent modification of the optical response of the 

PSi device by reflectivity spectroscopy. 

The modified surface obtained works as an active 

substrate for the chemistry of the following attachment 

of the protein: we have spotted on the PSi chip 20 ll of 

7.5 lM sodium bicarbonate buffer (pH 7.35) containing 

a rhodamine labeled TMBP and incubated the 

system at –4°C overnight. Even if our aim is the realization 

of a label-free optical biosensor based on the 

PSi nanotechnology, we have used a fluorescent protein 

to control the distribution of the biological matter 

on the chip surface and to test the chemical stability 

of the covalent link between the TMBP and the PSi 

surface. 

After this assessment phase, we have also optically 

detected the ligand-binding interaction by following 

the wavelength shift of the reflectivity spectrum. The 

experimental measurement of the TMBP–glucose 

binding is a two step procedure: first, we have registered 

the optical spectrum of the porous silicon layer 

after the TMBP immobilization on the DBR surface 

and after the Glucose solution has been spotted on it. 

Results 

In Fig. 2, we have reported the optical spectra of our 

device as etched and after each step of the chemical 

0.0 

600 800 


Bragg as etched 

after oxidation 

after APTES treatment 

after Glutaraldehyde treatment 

123

treatment in the range 600–800 nm, where the m=5 

Bragg resonance peak at about 720 nm is present. The 

oxidation process causes a blue shift of the reflectivity 

spectrum of 99.5 nm due to the lower value of the 

SiO2 refractive index (noxi @ 1.5)with respect to the 

Si refractive index (nSi @ 3.2). On the contrary, the 

silanisation steps by APTES and glutaraldehyde produce 

red shifts of the reflectivity spectrum of 28 and 

17 nm, respectively, corresponding to an increasing of 

the average refractive index of the layers due to a 

filling of the pores by the organic layers. 

Since the protein distributes uniformly on the pores 

surface, the TMBP attachment causes a new detectable 

red shift of only 9 nm in the reflectivity spectrum. 

The FT-IR spectra of the oxide PSi sample and after 

the silanization process are reported in Fig. 3: the main 

characteristic peaks of silicon dioxide (at 1,124 cm –1 ), 

of the APTES amino groups (at 3,300 and 3,352 cm –1 ) 

and of gluteraldehyde cyano group (at 1,404 cm –1 ) are 

easily recognized. 

In Fig. 4a, is shown the DBR observed by a Leica 

Z16 APO fluorescence macroscopy system after incu- 

Fig. 3 FT-IR spectra of the 

Bragg reflector after 

oxidation and APTES/ 

glutaraldehyde treatment 

Fig. 4 a Porous silicon chip 

after incubation with the 

labeled-TMBP. b Porous 

silicon chip after washings in 

demi-water 

123 

%R 

320 

300 

280 

260 

240 

220 

200 

180 

160 

140 

120 

100 

80 

60 

40 

20 

4000 

3500 

bation. By illuminating the chip spotted with the labeled 

protein by a 100 W high-pressure mercury 

source, we found that the fluorescence is very high and 

homogeneous on the whole surface. We have also 

qualitatively tested the strength of covalent bond between 

the protein and the porous silicon surface by 

washing the device in a dialysis membrane overnight in 

DI-water. Since the fluorescent intensities differ of 

only few counts, we can conclude that the PSi-TMBP 

double layer is very stable. 

We have also measured the signal response to the 

glucose concentration after the interaction with the 

protein in a range between 10 and 150 lM. The 

maximum shift of the Bragg wavelength is 1.2 nm. 

Figure 5 shows the dose–response curve to glucose 

additions. The estimated sensitivity of the TMBP- 

Chip is 0.03 nm/lM. Interestingly, the concentration 

of glucose that induces a optical response of the 

protein is very close to the amount of the sugar 

present in the human interstitial fluids. This result 

suggests the use of this protein in designing of a nonconsuming 

and minimally invasive biosensor for the 

3352.17 NH2 

3300.00 NH2 

DBR post oxidation process 

DBR post APTES functionalization 

DBR post Glutaraldehyde functionalization 

3000 

2930.43CH3 

2500 2000 

cm-1 

1124.94 SiO2 

1656.52 COH 

1500 

1404.35 CN 

Extremophiles 

1000 

500

Extremophiles 

∆λ (nm) 

1.4 

1.2 

1.0 

0.8 

0.6 

0.4 

0.2 

0.0 

70 80 90 100 110 120 130 140 150 

Glucose concentration (µM) 

Fig. 5 Dose–response curve for PSi DBR optical sensor exposed 

to several concentration of glucose 

continuous detection of the level of glucose in diabetic 

patients. 

In conclusion, we have reported an effective methodology 

for the covalent immobilization of proteins 

from extremophiles on a nanostructured material, such 

as porous silicon wafer, that can be utilized as general 

platform for the development of very stable proteinbased 

arrays for analyses of high social interest. 

Acknowledgments The authors thank Dr. Fabienne Chevance, 

University of Utah, Salt Lake City, USA, and Prof. Winfried 

Boos, University of Konstanz, Germany, for supplying a partially 

purified protein sample. This work is in the frame of the project 

‘‘Diagnostica Avanzata ed Alimentazione’’ CNR Commessa of 

the Agro-Food Department (SD). This work was also supported 

by the ASI project MoMa No. 1/014/06/0 (SD) and by a grant 

from the Ministero degli Affari Esteri, Direzione Generale per la 

Promozione e la Cooperazione Culturale (SD). 

References 

Bradford MM (1976) A rapid and sensitive method for the 

quantitation of microgram quantities of protein utilizing the 

principle of protein-dye binding. Anal Biochem 72:248–254 

Brock TD (1985) Life at high temperatures. Science 239:132–138 

D’Auria S, Moracci M, Febbraio F, Tanfani F, Nucci R, Rossi M 

(1998) Structure-function studies on b-glycosidase from 

Sulfolobus solfataricus. Molecular bases of thermostability. 

Biochemie 80:949–957 

Diez J, Diederichs K, Greller G, Horlacher R, Boos W, Welte W 

(2001) The crystal structure of a liganded trehalose/maltosebinding 

protein from the hyperthermophilic archaeon 

Thermococcus litoralis at 1.85 A. J Mol Biol 305:905–915 

Herman P, Staiano M, Marabotti A, Varriale A, Scirè A, Tanfani 

F, Vecer J, Rossi M, D’Auria S (2006) D-Trehalose/ 

D-maltose-binding protein from the hyperthermophilic archaeon 

Thermococcus litoralis: The binding of trehalose and 

maltose results in different protein conformational states. 

Proteins Struct Func Bioinform 63:754–767 

Horlacher R, Xavier KB, Santos H, DiRuggiero J, Kossmann M, 

Boos W (1998) Archaeal binding protein-dependent ABC 

transporter: molecular and biochemical analysis of the 

trehalose/maltose transport system of the hyperthermophilic 

archaeon Thermococcus litoralis. J Bacteriol 180:680–689 

Jaenicke R (1994) Studies on hyperstable proteins: crystallins 

from the eye-lens and enzymes from thermophilic bacteria. 

In: Doniach S (ed) Statistical mechanics, protein structure, 

and protein substrate interactions. Plenum, New York, pp 

49–62 

Sun YJ, Rose J, Wang BC, Hsiao CD (1998) The structure of 

glutamine-binding protein complexed with glutamine at 1.94 

angstrom resolution: comparisons with other amino acid 

binding proteins. J Mol Biol 278:219–222 

Xavier KB, Martins LO, Peist R, Kossmann M, Boos W, Santos 

H (1996) High-affinity maltose/trehalose transport system in 

the hyperthermophilic archaeon Thermococcus litoralis. 

J Bacteriol 178:4773–4777 

123

Sensors 2008, 8, 1-x manuscripts; DOI: 10.3390/sensors 

Article 

sensors 

ISSN 1424-8220 

www.mdpi.com/journal/sensors 

Porous Silicon Based Resonant Mirrors for Biochemical Sensing 

Edoardo De Tommasi 1,* , Luca De Stefano 1 , Ilaria Rea 1, 2 , Valentina Di Sarno 1, 2 , Lucia 

Rotiroti 1, 3 , Paolo Arcari 4 , Annalisa Lamberti 4 , Carmen Sanges 4 and Ivo Rendina 1 

1 

Institute for Microelectronics and Microsystems – Unit of Naples – National Council of Research, 

Via P. Castellino 111, 80131 Napoli, Italy. 

2 

Department of Physical Sciences, University of Naples “Federico II”, Via Cinthia, 80126 Naples, 

Italy 

3 

Department of Organic Chemistry and Biochemistry, University of Naples “Federico II”, Via 

Cinthia, 80128 Napoli, Italy 

4 

Department of Biochemistry and Medical Biotechnologies, University of Naples “Federico II”, Via 

S. Pansini 5, 80131 Napoli, Italy. 

* Author to whom correspondence should be addressed: edetommasi@na.imm.cnr.it 

Received: 14 Oct 2008; in revised form: 6 Oct 2008 / Accepted: 21 Oct 2008 / Published: 

Abstract: We report on our preliminary results in the realization and characterization of a 

porous silicon (PSi) resonant mirror (RM) for optical biosensing. We have numerically and 

experimentally studied the coupling between the electromagnetic field, totally reflected at 

the base of a high refractive index prism, and the optical modes of a PSi waveguide. This 

configuration is very sensitive to changes in the refractive index and/or in thickness of the 

sensor surface. Due to the high specific area of the PSi waveguide, very low DNA 

concentrations can be detected confirming that the RM could be a very sensitive and labelfree 

optical biosensor. 

Keywords: DNA Optical Biosensors, Porous Silicon, Resonant Mirrors. 


Optical label-free biosensors are ideal candidates for high throughput screening in the prevention of 

bio-threat agents or social illnesses: all the different configurations proposed in the literature, such as 

optical microcantilever [1], interferometric devices [2, 3], and also the recent photonic crystals 

geometries [4], show very high sensitivities in the recognition of specific molecules. The detection of

Sensors 2008 2 

analytes present in complex mixtures in ultra-low concentration, such as viruses in human blood, is 

possible due to two key factors: the utilization of bioprobres, properly linked on the sensor surface, 

and an optimized design to maximize the interaction between the biological matter and the 

electromagnetic field. 

Resonant Mirrors (RMs) are optical sensors which exploit the evanescent fields features to probe 

changes in refractive index and thickness taking place on their surface after exposure to gaseous or 

liquid substances. From the optical point of view, RMs are refractometric sensing devices similar to 

grating coupler sensors: they are characterized by the high sensitivity typical of waveguiding devices, 

and, on the other hand, by the simple scheme of Surface Plasmon Resonance (SPR) sensors [5]. In 

RMs, in fact, incident light gives rise to an evanescent wave at the interface between a prism and an 

optical waveguide; the coupling of light into the waveguide occurs only at those specific incident 

angles for which the propagation constant of the evanescent wave matches that of a waveguide mode. 

This matching produces a dip in the angular spectrum of the reflected light, so that each change in the 

refractive index and-or in thickness at the sensor surface produces a corresponding shift in the position 

of the dip [5, 6]. 

In this work, we report on the realization of a porous silicon (PSi) based RM prototype on chip, 

whose modal properties and characteristic angular resonances have been both numerically computed 

and experimentally measured. 

PSi is by far one of the most intriguing material in optical sensing: the refractive index is widely 

tuneable, namely between the silicon refractive index and that of the air, and its specific surface is very 

large, up to 500 m 2 cm -3 . Due to these characteristics, lot of optical structures, such as rugate filters [7] 

and microcavities [8] have been proposed in literature for biosensing. Recently, a PSi waveguide 

biosensor have been theoretically and experimentally demonstrated [9]. We have studied and reported 

in this paper the feasibility of such biosensor for DNA-DNA hybridization experiment. 

Figure 1. Experimental setup. DL: diode laser emitting s-polarized radiation at 785 nm; L: 

lens with focal length f=7.5 cm; RM: resonant mirror; RS: rotation stage; PM: power meter 

on an independent rotation stage.


2. Experimental 

A highly doped p+-silicon, oriented, 0.01 Ω·cm resistivity, 400 μm thick was used as 

substrate in the waveguide fabrication. The structure was obtained by electrochemical etching of 

crystalline silicon in a HF-based solution (50 wt. % HF:ethanol = 3:7) at room temperature. A current 

density of 10 mA/cm 2 was applied for 18.9 s to produce the core layer of thickness 2.5 μm and porosity 

65 %, while a current density of 109 mA/cm 2 was applied for 13.6 s in the case of the cladding layer 

with thickness 2.5 μm and porosity 78 %. These porosities correspond to a core and cladding refractive 

indexes of 1.749 and 1.384, respectively, calculated by the Bruggeman model [10] at a wavelength of 

785 nm. The device was then fully oxidized in pure O2 by a two step thermal treatment (400 °C for 30 

min and 900 °C for 15 min). 

In Figure 1 a wide view of our experimental set-up is shown. RM is mounted on a rotation stage 

(Melles Griot, model 07 TRS 501); the light source is a diode laser emitting s-polarized radiation at 

λ=785 nm (CrystaLaser, model RCL-080-785-5); the laser beam exiting the prism is collected by a 

photometer (Newport, model 1830-C). The waveguide has been brought in close contact with one side 

of an SF6 prism (n=1.784), still allowing the presence of a sub-micrometric air layer acting as a 

coupling zone between the prism and the waveguide. 

The covalent bond of the DNA single strand (5’-GGACTTGCCCGAATCTACGTGTCC-3’) on 

the PSi surface is based on a three steps functionalization process constituted by a chemical 

passivation of the surfaces, as it is shown in Figure 2 (all the chemicals used in the present study were 

from Sigma). We have treated the surfaces with a proper chemical linker, the 

aminopropyltriethoxysilane (APTES). Samples have been rinsed by dipping in a 5 % solution of 

APTES and a hydroalcoholic mixture of water and methanol (1:1) for 20 min at room temperature. 

After the reaction time, we have washed the chips by deionized water and methanol, and dried in N2 

stream. The silanized devices were then baked at 100 °C for 10 min. To create a surface able to link 

the amino group of the biological probes, we have thus immersed the chips in a 2.5% glutaraldehyde 

(GA) solution in 20 mM HEPES buffer (pH 7.4) for 30 min, and then rinsed it in deionized water and 

finally dried in N2 stream. The GA reacts with the amino groups on the silanized surface and coats the 

internal surface of the pores with another thin layer of molecules. All these reaction steps have been 

monitored by means of FT-IR spectroscopy to verify the presence of the characteristic peaks of the 

organic linkers. The main characteristic peaks of Si-O-Si (1040 cm -1 ) and Si-OH bonds (3740 and 

1646 cm -1 ) are present, after oxidation, on the PSi. After the silanization process, the APTES 

characteristic peaks of the ethylic (at 1626, 1529 and 1379 cm -1 ), amino (at 1060 cm -1 ) and also the – 

CH (at 2908 and 2851 cm -1 ) groups are well evident. Finally, after the GA treatment, the characteristic 

imine group (at 1627 cm-1) due to the reaction with APTES is easily recognized. 

The PSi surface was covered overnight at room temperature with a 50 μM DNA probe solution (30 

μL). For each measurement, complementary (5’-GGACACGTAGATTCGGGCAAGTCC-3’) and noncomplementary 

(5’-CACTGTACGTGCGAATTAGGTGAA-3’) DNA (10 μL) were spotted on the 

chip and incubated for two hours. Before optical measurements, all samples have been extensively 

rinsed in deionized water to remove the excess of biological matter.


Figure 2. Schematic of the functionalization process, from the oxidized PSi chip to the 

covalent attachment of the DNA single strands. 

SiO 2 

OH 

OH 

OH 

O 2Si 

O 2Si 

O 

5 % APT ES 

20' R.T 

O 2Si 

O 

O 

O 

Si 

O 

NH 2 

2.5% GA 

30' R.T 

O 

3. Results and Discussion 

O 

O 

Si 

O 

O 

Si 

O 

O 

N 

N 

C 

C 

C O 

C N 

ssDNA 

ss DN A-NH 2 

over night 

R.T 

The PSi waveguide, once oxidized in order to reduce the scattering losses [11] and to allow the 

bond with biological linkers, has been characterized by means of “m lines” technique, which 

determines the modal behavior of the waveguide at different wavelengths and estimates the refractive 

indices nc and nb for the core an the buffer layers [12]. In Figure 3, is reported a standard m-lines 

spectrum due to the excitation of four adjacent modes in transverse electric (TE) polarization of the 

electromagnetic field at 785 nm: by measuring the coupling angles position is possible to quantify the 

core and cladding thicknesses and refractive indexes. We have obtained the refractive indexes of core 

and buffer and the thickness of the core equal to nc = 1.370 ± 0.001, nb = 1.18 ± 0.01, and dc = 2.71 ± 

0.01μm respectively. It is important to notice the strong reduction of the refractive indexes after the 

thermal oxidation process, mainly due to the substitution of silicon with silicon dioxide. 

The close contact between the prism and the waveguide still allows the presence of a sub-micrometric 

air layer where the evanescent fields coming from the prism and the waveguide can overlap, thus 

transferring the light energy from the prism to the waveguide. The thickness of this air layer strongly 

affects the efficiency of the coupling of the incoming light into the waveguide. In Figure 4, some 

simulations of RM spectra corresponding to different air thicknesses are shown: the lower the 

thickness, the stronger the coupling of the evanescent wave with the waveguide. The dips present in 

the total reflection zone (i.e. for incident angle θ > 34° 

), correspond to the characteristic modes 

propagating in the waveguide. The number of these modes can be obtained from the well known 

2d 2 2 

relation M = nc 

− nb 

≈ 5, 

where d stands for the thickness of the waveguiding layer. 

λ 

i


Figure 3. Resonant peaks of the waveguide, corresponding to four guided modes in TE 

polarization at 785 nm, after functionalization with APTES and GA. 

Output power (μW) 

65 

60 

55 

50 

45 

40 

46 48 50 52 

Resonance Angle (deg) 

Figure 4. Numerical simulations of RM spectra as a function of the air gap thickness 

Normalized Output Intensity 

1.0 

0.5 

0.0 

30 35 40 

Resonance Angle (deg) 

Air thickness: 1 μm 

Air thickness: 500 nm 

Air thickness: 200 nm 

. 

In Figure 5, the results of angular measurements after incubation with probe DNA and after two 

steps of the DNA hybridization process at two different concentrations are reported: the resonance 

shifts confirm the effectiveness of the DNA binding.


Figure 5. PSi RM resonances after functionalization with probe DNA and after two DNA 

hybridization steps: the coupling angle shifts demonstrate the molecular interaction 

between the probe DNA and its complementary target at different concentration. 

Normalized Output Power 

1.0 

0.9 

0.8 

0.7 

0.6 

After incubation with 50 μM of DNA 

After incubation with 25 μM of target cDNA 

After incubation with 50 μM of target cDNA 

Δθ 

45.5 46.0 46.5 47.0 47.5 

Resonance angle (deg) 

The coupling angle shifts as a function of the concentration of the complementary DNA, as it can be 

seen in Figure 6. The high efficiency of hybridization is confirmed by the fact that, after exposure to 

50 μM of complementary DNA, the sensor response curve reaches a plateau, corresponding to a 

saturation of almost all of the active sites. The experimental points can be well fitted (R 2 =0.99, χ 2


Figure 6. Resonance shifts of the waveguide mode coupling angle as a function of the 

DNA concentration. 


Δθ (deg) 

1,8 

1,6 

1,4 

1,2 

1,0 

0,8 

0,6 

0,4 

0,2 

0,0 

0 10 20 30 40 50 

Target Concentration (μM) 

We have experimentally confirmed the ability of a PSi RM to detect a DNA-DNA hybridization 

process with very high sensitivity. The use of PSi technology in the fabrication of the waveguide, 

which is the heart of a RM sensor, allows the presence of a high specific-surface available for 

interaction with great amounts of analytes. We have demonstrated that our measuring system is 

characterized by a technical limit of about 80 nM of DNA, in good agreement with previously 

published theoretical calculations [13]. Next efforts will be aimed at the fabrication of a free-standing 

waveguide (no crystalline silicon substrate), which, in conjunction with a flow delivery system, could 

allow to reduce considerably the time of incubation of the DNA probe and targets over the 

functionalized surface. 


This work is partially supported by the FIRB Italian project RBLA033WJX_005. 

References and Notes 

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© 2008 by the authors; licensee Molecular Diversity Preservation International, Basel, Switzerland. 

This article is an open-access article distributed under the terms and conditions of the Creative 

Commons Attribution license (http://creativecommons.org/licenses/by/3.0/).

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