Polymer integrated Young interferometers for label-free...

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ISBN 978-951-42-9963-6 (Paperback)ISBN 978-951-42-9964-3 (PDF)ISSN 0355-3213 (Print)ISSN 1796-2226 (Online)

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OULU 2012

C 432

Meng Wang

POLYMER INTEGRATED YOUNG INTERFEROMETERS FOR LABEL-FREE BIOSENSING APPLICATIONS

UNIVERSITY OF OULU GRADUATE SCHOOL;UNIVERSITY OF OULU,FACULTY OF TECHNOLOGY,DEPARTMENT OF ELECTRICAL ENGINEERING;UNIVERSITY OF OULU,INFOTECH OULU

C 432

ACTA

Meng W

ang

C432etukansi.fm Friday, October 26, 2012 12:55 PM

A C T A U N I V E R S I T A T I S O U L U E N S I SC Te c h n i c a 4 3 2

MENG WANG


Academic dissertation to be presented with the assentof the Doctoral Training Committee of Technologyand Natural Sciences of the University of Oulu forpublic defence in Auditorium TS101, Linnanmaa, on 23November 2012, at 12 noon

UNIVERSITY OF OULU, OULU 2012

Copyright © 2012Acta Univ. Oul. C 432, 2012

Supervised byProfessor Risto MyllyläDocent Jussi Hiltunen

Reviewed byProfessor Seppo HonkanenProfessor Janis Spigulis

ISBN 978-951-42-9963-6 (Paperback)ISBN 978-951-42-9964-3 (PDF)

ISSN 0355-3213 (Printed)ISSN 1796-2226 (Online)

Cover DesignRaimo Ahonen

JUVENES PRINTTAMPERE 2012

Wang, Meng, Polymer integrated Young interferometers for label-free biosensingapplications. University of Oulu Graduate School; University of Oulu, Faculty of Technology, Department ofElectrical Engineering; Infotech Oulu, P.O. Box 4500, FI-90014 University of Oulu, FinlandActa Univ. Oul. C 432, 2012Oulu, Finland

Abstract

Integrated optical (IO) sensor allowing sensitive, label-free, real-time and multi-parametermonitoring of bio-molecular interactions are conventionally fabricated with inorganic dielectricsinherited from CMOS manufacturing technology. Polymers as complement materials to inorganicdielectrics are becoming to have an increasing market share for IO circuits in opticalcommunications networks owing to its good optical properties, versatile processibility and lowcost. This work aims at developing disposable low-cost biosensors based mainly on polymericmaterials, with a performance comparable to inorganic-dielectric based IO biosensors.

This thesis describes the development of polymer IO biosensors based on the Younginterferometer (YI) transducer platform for ambient noise compensation and a complete periodicintensity fringe pattern. Three different waveguide configurations were utilized, taking intoconsideration operational simplicity, fabrication simplicity and enhanced sensitivity. Among thedeveloped polymer biosensors, an unconventional interferometer structure: a vertically placeddual-slab waveguide interferometer and an inverted rib waveguide configuration were employed.To enhance the sensitivity of the waveguides, deposition of Ta2O5 high index coating wasperformed on the rib waveguide configuration. Along with the development of polymer biosensorsbased on the inverted-rib waveguide configuration, a fabrication process was also developedfeaturing UV-imprinting and spin coating. The simple two-step fabrication process demonstratedusing a polymer mold is potentially transferable to the roll-to-roll manufacture process.

Calibration of the developed sensors was performed by homogeneous refractive index (RI)sensing with glucose de-ionized water solutions. By investigating an antibody – antigen bindinginteraction involving C-reactive protein and its conjugates, this thesis confirmed the applicabilityof the developed sensors to specific molecule detection. Moreover, to establish the influence ofwater molecular absorption on measurement stability, an evaluation was carried out on thepolymeric waveguide. Finally, the thesis presented a comparison between the developed sensors,exploring their sensitivities, stabilities, limits of detection (LODs) and other aspects related tooperation and fabrication. The results indicated that the Ta2O5-coated polymer waveguide sensorhad a high sensing capability. In homogeneous RI sensing, the achieved detection limits were9×10-7 RIU (refractive index unit), i.e., three times the noise level, and 270 fg/mm2 for surfacemass density.

Keywords: immunoassay, optical biosensor, rib waveguide, slab waveguide, Ta2O5high-index coating, UV-imprinting, UV-sensitive polymer, Young interferometer

Wang, Meng, Polymeeriset integroidun optiikan Youngin interferometritmerkkiaineettomissa biomittauksissa. Oulun yliopiston tutkijakoulu; Oulun yliopisto, Teknillinen tiedekunta, Sähkötekniikan osasto;Infotech Oulu, PL 4500, 90014 Oulun yliopistoActa Univ. Oul. C 432, 2012Oulu

Tiivistelmä

Integroidulla optiikalla toteutetut anturit mahdollistavat biomolekulaarisen vuorovaikutuksentutkimisen käyttäen herkkiä moniparametrisia ja merkkiaineettomia menetelmiä. Näiden bioan-tureiden valmistukseen käytetään tavallisesti CMOS-teknologian piiristä tuttuja epäorgaanisiapuolijohteita ja eristemateriaaleja. Viime aikoina on kuitenkin polymeeristen materiaalien käyt-töä integroidussa optiikassa tutkittu merkittävästi johtuen polymeerien hyvistä optisista ominai-suuksista, monipuolisesta työstettävyydestä ja edullisista kustannuksista. Tämän työn tarkoituk-sena on kehittää edullisia, kertakäyttöisiä, pääasiallisesti polymeerisistä materiaaleista valmistet-tuja bioantureita, jotka vastaavat suorituskyvyltään epäorgaanisista materiaaleista valmistettujaintegroidun optiikan antureita.

Tässä työssä kehitetyt polymeeriset integroidun optiikan bioanturit perustuvat Youngin inter-ferometriin mahdollistaen ympäristökohinan kompensoinnin ja ne tuottavat pintavuorovaikutus-ten tutkimiseen jaksoittaisen interferenssikuvion. Työssä hyödynnettiin kolmea erilaista valoka-navarakennetta huomioiden niiden käytön helppous, valmistuksen yksinkertaisuus ja mittaus-herkkyys. Yksi kehitetyistä polymeerisistä bioantureista koostui päällekkäisistä kerrostetuistapolymeerikerroksista. Toisen tutkitun rakenteen toiminta puolestaan perustui käänteiseen harjan-nevalokanavaan. Mittausherkkyyttä parannettiin pinnoittamalla polymeerirakenne Ta2O5-pin-noitteella. Näin muodostui kerrostettu komposiittivalokanava, joka oli tässä työssä tutkittu kol-mas sensorirakenne. Itse bioanturien lisäksi kehitettiin myös valmistusprosessi, jossa hyödynnet-tiin UV-painatusta ja nestefaasipinnoitusta. Tässä työssä havaittiin lisäksi, että kehitetty yksin-kertainen valmistusmenetelmä on paitsi toimiva, myös mahdollisesti siirrettävissä rullalta rullal-le valmistus- ja tuotantoteknologiaan.

Kehitettyjen anturien kalibrointi suoritettiin homogeenisella taitekerroinmittauksella käyttä-en liuoksia, jotka valmistettiin glukoosista ja deionisoidusta vedestä. Kehitettyjen anturiensoveltuvuus spesifien molekyylien tunnistamista varten todennettiin tutkimalla vasta-aineiden jaantigeenien sitoutumisreaktioita ja vuorovaikutusta C-reaktiivisella proteiinilla ja sen konjugaa-teilla. Lisäksi työssä tutkittiin veden absorption vaikutusta mittauksen stabiilisuuteen. Tutkimuk-sessa suoritettiin vertailu kehitettyjen anturien ja niiden ominaisuuksien välillä kiinnittäen huo-miota mittausherkkyyteen, stabiilisuuteen, määritys- ja toteamisrajoihin ja muihin anturien val-mistukseen sekä käyttöön liittyviin keskeisiin piirteisiin. Tulokset osoittavat, että Ta2O5-pinnoi-tetun polymeerivalokanavan mittausherkkyys oli suurin vertailluista rakenteista. Homogeenises-sä taitekerroinmittauksessa saavutettu määritys- ja toteamisraja oli 9×10-7 taitekerroinyksikköä(RIU). Pintamassatiheysmittauksessa saavuttu tulos oli 270 fg/mm2.

Asiasanat: harjannevalokanava, immunomääritys, optinen bioanturi, planaarinenvalokanava, Ta2O5-pinnoitus, UV-painatus, UV-sensitiivinen polymeeri, Younginterferometri

7

Acknowledgements

This study is based on research work carried out at the Optoelectronics and

Measurement Techniques Laboratory of University of Oulu during the years

2007–2012. Therefore I would like to express my deepest gratitude to my

supervisor Prof. Risto Myllylä for providing me the research opportunity in his

group, and all the support and guidance throughout my undergraduate and

doctoral studies.

My sincere appreciation goes to Dr. Jussi Hiltunen from VTT Technical

Research Center of Finland. I want to thank him for leading me to his research

expertise, and for sharing his time, knowledge and ideas to guide me through the

difficulties of research. Thanks also to other colleagues from VTT, Dr. Leena

Hakalahti for her help of providing many of the facilities and resource necessary

to carry out my research work, Dr. Pentti Karioja, M.Sc. Sanna Uusitalo, M.Sc.

Christina Liedert, M.Sc. Marianne Hiltunen, and M.Sc Noriyuki Masuda for their

help and fruitful discussions.

I want to acknowledge and thank all my colleagues in the Optoelectronics

and Measurement Techniques Laboratory, especially Dr. Hanna Jobbour, Lic.Tech.

Eija Vieri, M.Sc. Karoliina Jokinen, and M.Sc. Miia Määttälä for helping me

accomplish this thesis work with their friendship, encouragement, and continuous

support.

I wish to thank co-author Prof. Jyrki Lappalainen, M.Sc. Jarkko Puustinen in

Microelectronics and Materials Physics Laboratory for AFM measurements, and

Dr. Martin Charlton and Dr. Stuart Pearce for preparing Ta2O5 coating in

University of Southampton.

I would also like to thank Prof. Seppo Honkanen (University of Eastern

Finland) and Prof. Janis Spigulis (University of Latvia) for reviewing my thesis.

I want to thank my parents for believing in me and supporting me to pursue

my degrees in Finland ten years ago. Without their encouragement, none of this

would be possible. Finally, I owe my deepest and loving gratitude to my beloved

husband Dr. Wu Chen and son Jiayuan Chen. Without their love and

understanding, I would have never finished this book.

Financial support from Infotech Oulu Graduate School, TaunoTönning

Foundation, and Tekniikan edistämissäätiö (TES) is highly appreciated.

Oulu, June 2012 Meng Wang

9

Abbreviations and symbols

1D one dimensioanl

2D two dimensional

AFM atomic force microscopy

BSA bovine serum albumin

CMOS complementary-symmetry metal–oxide–semiconductor

CRP C-reactive protein

CTE coefficient of thermal expansion

CVD chemical vapor deposition

DI de-ionized

FFT fast Fourier transform

F13 –TCS tridecafluoro-(1,1,2,2)-tetrahydrooctyl-trichlorosilane

III-V refers to the periodic table classification; material composition of

group three and five materials

IgG immunoglobulin G

IO integrated optics

KCl potassium chloride

LOD limit of detection

NaCl sodium chloride

Na2HPO4 disodium hydrogen phosphate

NaH2PO4 sodium dihydrogen phosphate

NIL nano-imprinting lithography

PBS phosphate-buffered saline

PDMS polydimethylsiloxane

PMMA poly(methyl methacrylate)

RI refractive index

RIE reactive ion etching

RIU refractive index unit

RMS root-mean-square

SEM scanning electron microscope

Si silicon

Si3N4 silicon nitride

SiO2 silicon dioxide

SiON silicon oxynitride

SNR signal-to-noise-ratio

SOI silicon on insulator

10

SPR surface plasma resonance

Ta2O5 Tantalum pentoxide

TiO2 Titanium dioxide

TE transverse electric

TIR total internal reflection

TM transverse magnetic

UV ultraviolet

YI Young interferometer

a half width of the waveguide

arb. unit arbitrary unit

°C degrees Celcius

b half length of the waveguide

D the distance from the sensor output facet to the collecting objective

lens

d the separation between the sensing and referencing waveguide at the

output facet

dn/dc the refractive index increment of the protein solution

Da Dalton, 1.66×10−27 kg

fg/mm2 femtogram per square millimeter

E electric field

E|| tangential part of the electric field

H|| tangential part of the magnetic field

I, I1, I2 Intensity

j imaginary unit

k0 wave number in vacuum

kx, ky wave number in the x- and y-direction

L interaction length

ml milliliter

nad refractive index of the adlayer

nc refractive index of the cover medium

neff effective refractive index

nf refractive index of the core

nm nanometer

nM nanomolar

ns refractive index of the substrate

P spatial period of the fringe pattern

11

pol polarization

p, q integers corresponding to the numbers of the optical power peaks in

x- and y- directions

rad radian

rpm rounds per minute

s second

t time

tad thickness of the adlayer

teff effective waveguide thickness

tf thickness of the waveguide

x, y, z Cartesian coordinates

β propagation constant

Γ coupling efficiency

δ phase change of the propagating wave in the sensing waveguide

ΔC concentration change of the analyte

ΔM change of the surface mass density

Δneff change of effective refractive index

ΔX change of one or a combination of parameters

Δφ phase difference between the referencing and sensing waveguides

Δzf, c evanescent field penetration depth into the substrate (j=s) or cover

medium (j=c)

λ wavelength

µm micrometer

ρ parameter used for theoretical sensitivity calculation

ω angular frequency

13

List of original papers

I Wang M, Uusitalo S, Määttälä M, Myllylä R & Känsäkoski M (2008) Integrated dual-slab waveguide interferometer for glucose concentration detection in the physiological range. Proc SPIE 7003: 70031N-1-10. DOI: 10.1117/12.780184.

II Wang M, Uusitalo S, Liedert C, Hiltunen J, Hakalahti L & Myllylä R (2012) Polymeric dual-slab waveguide interferometer for biochemical sensing applications. Appl Opt 51: 1886–1893.

III Wang M, Hiltunen J, Uusitalo S, Puustinen J, Lappalainen J, Karioja P & Myllylä R (2011) Fabrication of optical inverted-rib waveguides using UV-imprinting. Microeletron Eng 88(2): 175–178.

IV Wang M, Hiltunen J, Liedert C, Hakalahti L & Myllylä R (2012) An integrated Young interferometer based on UV-imprinted polymer waveguides for label-free biosensing applications. J Europ. Opt Soc Rap Public 7: 12019.

V Wang M, Hiltunen J, Liedert C, Pearce S, Charlton M, Hakalahti L, Karioja P & Myllylä R (2012) Highly sensitive biosensor based on UV-imprinted layered polymeric–inorganic composite waveguides. Opt Express 20: 20309–20317.

Author’s contribution to publications I-V:

I The author studied the slab waveguide Young interferometer with the

commercialized SiON based sensor chip provided by Farfield Scienctific

Inc. In this paper, a theoretical sensitivity analysis of three-layer waveguide

biosensors was carried out together with co-author Sanna Uusitalo.

II The author fabricated and characterized the polymer dual-slab waveguide

interferometer biosensors with glucose water solutions and demonstrated

the functionality of biosensing with non-specific adsorption of

biomolecules.

III Together with co-author Dr. Jussi Hiltunen, the author developed the

fabrication process for the inverted-rib polymer waveguide. The author

fabricated the waveguides utilizing UV-imprinting lithography.

Characterization of the waveguide was done together with Dr. Jussi

Hiltunen. Fabrication of the inverted-rib waveguide is a fundamental step

for development of sensing systems in Papers IV & V.

IV The author fabricated and characterized the polymer inverted-rib

waveguide Young interferometer biosensor. The sensor was calibrated with

bulk refractive index sensing using glucose water solutions. Biosensing

ability was demonstrated with specific antibody-antigen binding.

14

V The composite polymeric-inorganic waveguide interferometer biosensor

was developed, based on the structure of the polymer inverted-rib

waveguide prepared by the author. Ta2O5 high-index coating was fabricated

by co-author Jussi Hiltunen at Southampton University. The author

conducted both the theoretical and experimental sensitivity analysis and

performed the biosensing measurements using an immunoassay provided

by co-author Christina Liedert. The measurement results suggested a

significant enhancement of sensitivity and detection limit.

15

Contents

Abstract Tiivistelmä Acknowledgements 7 Abbreviations and symbols 9 List of original papers 13 Contents 15 1 Introduction 17

1.1 Basics of optical biosensors .................................................................... 17 1.2 Integrated evanescent field biosensors .................................................... 18 1.3 Motivation of the thesis ........................................................................... 19

2 Waveguide biosensors in theory 21 2.1 Theory behind waveguide biosensors utilizing evanescent field

probe ....................................................................................................... 21 2.2 Young interferometer .............................................................................. 23

2.2.1 Structures of Young interferometer .............................................. 23 2.3 Theoretical sensitivity analysis of polymer-based waveguide

sensors ..................................................................................................... 25 2.3.1 Determination of propagation constant by the Marcatili

method .......................................................................................... 25 2.3.2 The sensitivities to the change of bulk RI and adlayer

thickness ....................................................................................... 26 3 Sensor design, fabrication and characterization 29

3.1 Polymers in integrated optics .................................................................. 29 3.1.1 Polymeric IO biosensors ............................................................... 30 3.1.2 Ormocer - organic and inorganic hybrid polymer ........................ 32

3.2 Dual-slab Young interferometers ............................................................ 32 3.2.1 Fabrication process and waveguide characterization ................... 33

3.3 Inverted-rib waveguide Young interferometers ....................................... 35 3.3.1 Inverted-rib waveguide fabrication and characterization ............. 35 3.3.2 Rib-waveguide for Young interferometers ................................... 38

3.4 Ta2O5 high refractive index coating ......................................................... 40 4 Biosensor experiments and results 45

4.1 Measurement setup ................................................................................. 45 4.2 Effect of water absorption in polymer waveguides ................................. 46 4.3 Sensing of bulk refractive index ............................................................. 48

16

4.4 Immunoassay .......................................................................................... 51 4.4.1 Materials ....................................................................................... 51 4.4.2 Experiments .................................................................................. 52 4.4.3 Results .......................................................................................... 54

5 Discussion 57 5.1 Measurement error analysis .................................................................... 57

5.1.1 Mechanical noise .......................................................................... 57 5.1.2 Quantization errors of the image acquisition device..................... 58 5.1.3 Optical noises ............................................................................... 60 5.1.4 Drift caused by temperature variations ......................................... 61

5.2 A comparison between the three different YI configurations .................. 62 5.3 Future work ............................................................................................. 64

6 Summary 65 References 67 Original Papers 73

17

1 Introduction

During the recent decades, the benefits of biosensor analysis have begun to be

recognized in many areas of analytical science, research and development. The

current application range includes biomedical research, healthcare,

pharmaceuticals, environmental monitoring and homeland security

(Narayanaswamy 2004 & Wolfbeis 2004, Cooper 2009). In general, the function

of a biosensor can be described as follows: a biosensor utilizes a biological

recognition element to detect the presence of an analyte and creates a physical or

chemical response that is converted by a transducer to a measurable signal. A

sophisticated biosensor can provide detailed information on the binding affinity

and, in many cases, also on the binding stoichiometry, thermodynamics and

kinetics of an interaction (Cooper 2009). As illustrated in Fig. 1, a biosensor

consists of two important components: a bio-recognition element and a transducer.

The bio-recognition element is often a biomolecule used as a receptor to

selectively capture an analyte (Prasad 2003), while the physical transducer may

be based on a range of mechanisms, including electrochemical (Wang 2006),

electrical (Chen et al. 2004), radioactive (Hsu et al. 1981), acoustic (Länge et al.

2008) and optical mechanisms (Ligler & Taitt 2002). This thesis focuses on

biosensors based on optical transducers.

Physical transducer

Fig. 1. Schematic illustration of a biosensor consisting of two components: a bio-

recognition element and a physical transducer.

1.1 Basics of optical biosensors

Biosensors based on optical transducers are often favored over other types of

transducers, because they are immune to electromagnetic disturbances, capable of

performing remote sensing, and offer multiplexed detection within a single device

(Fan et al. 2008). Optical biosensors convert biochemical interactions into

information carried by light. This information i.e. intensity, wavelength, phase,

polarization, etc., can then be measured by a photodetector or a spectrometer. This

18

type of sensors is widely used in two detection protocols: fluorescent-based

detection and label-free detection. In a fluorescent-based detection scheme,

labeling of target molecules or receptors usually needs to be prepared in advance.

This process is both labor intensive and time-consuming, and the fluorescent

marker may interfere with the function of the labeled biomolecule (Ray et al.

2010). In a label-free detection scheme, on the other hand, the properties of light

are directly manipulated by biochemical interactions to produce a real-time signal

without affecting the targeted molecules. The most common method in label-free

detection involves monitoring effective refractive index (RI) changes resulting

from biomolecular interactions. Effective RI changes take place on a sensing

surface coated with immobilized receptors. By binding to the receptors, target

molecules change the thickness of the surface adlayer, thereby inducing a change

in the effective refractive index. Such changes are probed by an evanescent field

at the surface of an optic fiber or a waveguide. By using the evanescent field, a

sensitive determination of the refractive index changes on the near adjacent of the

waveguide surface can be obtained in integrated optical sensors.

1.2 Integrated evanescent field biosensors

Integrated optical (IO) sensors have drawn a great interest in the past a few

decades, because they allow sensitive, real-time, label-free, on-site measurements

of biochemical species (Lambeck 2006). Most IO sensors are based on a planar

multilayer waveguide structure, deposited on top of a silicon or glass substrate. A

multilayer waveguide structure consists of a very thin higher refractive index

layer, containing a network of optical waveguide channels, and adjacent lower

refractive index layers for light confinement. As a result of total internal reflection

(TIR) of light propagating in the layer with the higher refractive index, an

evanescent wave field penetrates the boundaries of the layer and reaches the top

of the waveguide surface with a typical penetration depth of a few hundred

nanometers, which is sufficient for the study of biochemical interactions (Ligler

& Taitt 2002). This type of planar waveguide geometry allows facile integration

with sample delivery and detection systems, as well as the functionalization and

patterning of recognition element arrays onto the surface, enabling the

simultaneous detection of multiple analytes using a single waveguide transducer

(Mukundan et al. 2009).

Ever since the introduction of the first IO biosensor based on grating couplers

(Lukosz & Tiefenthaler 1984), IO sensors of many different configurations have

19

been developed, including grating couplers (Vörös et al. 2002), interferometer

sensors (Prieto et al. 2003), resonant mirrors (George et al. 1995), surface plasma

resonance (SPR) sensors (Homola et al. 1999), ring resonators (Ksendzov & Lin

2005) and photonic crystal waveguide sensors (Skivesen et al. 2007). Typical IO

sensors are usually fabricated using inorganic dielectrics such as SOI (Densmore

et al. 2008), Si3N4 (Ymeti et al. 2002), III-V group composites (Cassidy 2007),

metal oxides e.g. Ta2O5 (Schmitt et al. 2007), with claddings of SiO2. Fabrication

of inorganic dielectric-based waveguides poses high facility requirements.

Constructing a slab waveguide with its multilayer structure involves such

processes as sputtering, ion exchange, epitaxial growth or chemical vapor

deposition (CVD), while ridge/rib waveguides require conventional

photolithography and reactive ion etching (RIE) to define the wave-guiding

channel (Hunsperger 2009). These fabrication processes are both capital-intensive

and time-consuming. Typical sensitivity of bulk RI sensing devices or biosensors

based on inorganic dielectric materials have a limit of detection (LOD) above 10-6

refractive index units (RIU), approaching 10-8 RIU for homogeneous refractive

index sensing. In surface sensing of molecule adsorption, a LOD of 13 fg/mm2 in

surface mass density has been achieved (Schmitt et al. 2007).

1.3 Motivation of the thesis

In recent years, polymers have become important materials for optical

waveguides, because of their good optical properties, versatile processability and

potential for low-cost production. Polymers can be easily processed with various

molding techniques at relatively low temperatures. As a result, they have become

suitable complementary materials to inorganic dielectrics in waveguide sensor

fabrication.

The key motivation for this work is to develop low-cost, high-performance,

disposable optical sensors based on polymeric materials for biosensing

applications in aqueous environments. To achieve that end, the following issues

were taken into consideration during the sensor development:

1. A fabrication process targeted on polymeric materials needs to be developed

based on molding technologies. The process should enable high-volume,

low-cost production of polymer IO biosensors.

2. To simplify the fabrication process and enhance usability, new sensor

structures and waveguide configurations need to be utilized and evaluated.

20

3. To achieve a performance level comparable to inorganic dielectric based

biosensors, the sensitivity of polymer-based sensors needs to be improved.

21

2 Waveguide biosensors in theory

In an optical waveguide, the higher refractive index of the core layer allows total

internal reflection of light at the core-cladding interface. A light beam coupled

into the waveguide is confined in the core and guided through to the end facet of

the waveguide. Light propagation in optical waveguides is usually described as an

electromagnetic wave having angular frequency ω propagating in the z direction

with propagation constant β. The electric field is expressed by:

( ) ( )ztjeyxEzyxE βω −= ,),,( , (1)

The propagation constant β is the z-directed component of the wave number k.

The ratio of the propagation constant β in a medium to the wave number k0 in

vacuum is called the effective refractive index neff:

0k

neff

β= . (2)

2.1 Theory behind waveguide biosensors utilizing evanescent field

probe

When light propagates in a dielectric waveguide, the boundary condition of the

electromagnetic field must be satisfied. That is, the tangential component of the

electric field E|| and magnetic field H|| should be continuous across the boundary.

In order to satisfy the boundary condition, a non-vanishing solution to Maxwell’s

equations is found, suggesting the existence of a non-travel wave at the boundary.

This means that, although light energy is totally reflected, an electromagnetic

field extends out from the interface into the lower index medium as shown in Fig.

2. This field is called evanescent field, and it decays exponentially with distance

from the surface, generally around a depth of 100 nm (Taitt et al. 2005).

22

Fig. 2. Light propagation in a three-layer planar waveguide system.

The evanescent field probes changes of the propagating condition at the near

adjacent above the waveguide. Such changes cause the propagation constant β to

vary, leading to a phase change of the propagating light. From the biosensing

point of view, the evanescent field senses two types of ambient changes: the bulk

RI change of the cover media usually caused by the concentration change ΔC of

the analyte, or the adsorption of biomolecules ∆M (change of surface mass

density) on the waveguide surface. The transduction relationship can be described

with the following chain: (ΔC or ∆M) →ΔX →Δneff →δ, where Δneff and δ are the

effective index change of the waveguide and the phase change of the propagating

wave, respectively. The effective refractive index is a function of the following

parameters (Schmitt et al. 2008)

( ).,,,,,,, polntntnnnn cadadffseffeff λ= , (3)

where λ is the wavelength of the incident light; as shown in Fig. 2 ns, nf, nad, and

nc are the refractive indices of the substrate, waveguide, adlayer and cover media,

respectively; tf and tad are the thicknesses of the waveguide and the adlayer; and

pol indicates the polarization. ΔX denotes a change in one parameter or in a

combination of parameters mentioned above. Speaking of biosensing applications,

nc and tad are the factors affected by bio-stimulations, while changes in other

parameters are usually considered as noise.

Biosensing systems based on interferometers directly measure changes in the

effective refractive index Δneff. Consisting of a reference and a sensing waveguide,

the phase change of the propagating mode in the sensing waveguide of an

interferometer is translated into a spatial shift in the generated interference pattern.

By continuous recording the interferogram, real-time monitoring of biological

events can be enabled.

23

2.2 Young interferometer

Interferometer sensors typically consist of a reference arm and a sensing arm,

offering inherent background compensation for environmental disturbances, such

as temperature fluctuations. Interferometry is considered one of the most sensitive

techniques for optical sensing (Lukosz 1991). Different from the most commonly

used Mach-Zehnder interferometer, an IO Young interferometer (YI) generates a

complete periodic intensity pattern. It allows determining the absolute sign of the

induced phase change and produces a linear correlation between the spatial

change of the interference pattern and the induced phase change (Cross et al.

1999). For these reasons, the Young interferometer was chosen as the biosensor

development platform in this thesis.

2.2.1 Structures of Young interferometer

Three different structures of YI are commonly used in IO sensors, including a

horizontally-placed double-rib/strip waveguide structure (Schmitt et al. 2007, Qi

et al. 2009), a vertically-placed double-slab waveguide structure (Cross et al.

2003, Ricard-Blum et al. 2006) and a Y-junction structure (Brandenburg et al.

2000, Ymeti et al. 2005). This thesis utilized the latter two structures in polymer

biosensor developing. The dual-slab waveguide interferometer, shown in Fig. 3a,

was first developed in 1999 by Cross et al. (Cross et al. 1999). The vertical

structure contains alternatively stacked thin layers of waveguide claddings and

cores. Unlike traditional horizontally integrated Y-junction YIs, the slab-

waveguide YI does not need to be patterned for the structure of the Y-junction.

Moreover, the vertically-integrated structure offers the benefit of high tolerance to

misalignments of in-coupling light, since horizontal alignment is not required.

The Y-junction YI shown in Fig. 3b was also used, due to the existence of such

patterning methods, as imprinting and moulding, that are dedicated to polymers

and can be applied to simplify the fabrication process.

24

a) b)

Fig. 3. Structures of Young interferometer a) vertical spaced dual-slab Young

interferometer; b) Y-junction Young’s interferometer (Paper II & V, published by

permission of (c) 2012 OSA).

In both structures, a common input light beam launched at the input facet is split

and propagates through both waveguides simultaneously. The two separated

beams diffract out when they reach the output end-facet and interfere with each

other. Preferably an equal distribution of light in both waveguides is required in

order to optimize the visibility of the interferogram. In the far-field, a fringe

pattern is generated by the overlapping of the outgoing divergent beams. The

irradiance I of the interferogram pattern can be described by the following

equation (Hariharan 2003)

)cos(2 2121 δϕ −Δ++= IIIII , (4)

where I1 and I2 are the intensities of the two interfering beams, Δφ is the phase

difference between them, and δ is the additional phase shift caused by an effective

RI change in the sensing window. This change can be quantified by the following

expression

effnL Δ⋅=

0

2

λπδ , (5)

where L indicates interaction length (i.e., the length of the sensing window) and

λ0 is the vacuum wavelength of the monochromatic light source. When a stimulus,

i.e., a change in the analyte’s RI or attachment of biomolecules on the sensing

waveguide, appears, the effective RI of the propagating mode in the sensing

waveguide varies with respect to the wave propagating in the reference

waveguide. This produces a phase change of the generated interferogram,

displayed visually as a spatial shift of the fringe pattern.

25

2.3 Theoretical sensitivity analysis of polymer-based waveguide sensors

2.3.1 Determination of propagation constant by the Marcatili method

In slab and rib waveguides, the propagation constant β can be determined by the

Marcatili Method. It is an analytical method, designed to model two-dimensional

optical waveguides. Fig. 4 shows a cross-sectional view of a buried optical

waveguide. The core has a width of 2a and a height of 2b.

Fig. 4. Schematic illustration of the waveguide model used in the Marcatili method

(Paper I, published by permission of (c) 2008 SPIE).

In this method it is assumed that the electric and magnetic fields are confined to

the core and do not exist in the four hatched regions. By solving Maxwell’s

Equations, the following relations for wave numbers are obtained for the TE and

TM modes (Kawano & Kitoh 2001):

TE & TM:

22220 yxf kknk −−=β , (6)

TE:

( ) ( )12

arctan2222

0

2220

2220 −+

−−⋅

−−

= pk

knnk

knk

knkak

x

xcf

yc

yfx

π , (7)

( ) ( ) ( )12

arctan2

1arctan

2

12222

02222

0 −+

−−+

−−= q

k

knnk

k

knnkbk

y

ysf

y

ycf

y

π , (8)

26

TM:

( )( )

( ) ( )12

arctan2222

0

2220

2

2220

2

−+

−−⋅

−−

= pk

knnk

knkn

knknak

x

xcf

xcf

xfcx

π (9)

( ) ( ) ( )12

arctan2

1arctan

2

12222

0

2

222220

2

2

−+

−−+

−−= q

k

knnk

n

n

k

knnk

n

nbk

y

ysf

c

f

y

ycf

c

fy

π (10)

where k0 = 2π/λ0 is the wave number in vacuum and kx and ky are the wave

numbers in the x and y directions. As shown in Fig. 4, a and b are the half width

and length of the waveguide, while p and q are integers corresponding to the

numbers of the optical power peaks in the x and y directions. Unlike ordinary

mode orders, they start from 1 instead of 0.

With equation (2), the effective refractive index of the guided mode can be

derived from the propagation constant. This index is required in the next section

to analyze the theoretical sensitivity of the waveguide sensor.

2.3.2 The sensitivities to the change of bulk RI and adlayer thickness

IO waveguide biosensors respond to changes in the bulk RI of the cover media

and to the surface adsorption of biomolecues. A stimulated change on the

effective RI can be described by

ad

ad

effc

c

effeff t

t

nn

n

nn Δ

∂∂

+Δ∂∂

=Δ , (11)

where the derivatives ∂neff/∂nc and ∂neff/∂tad are the sensitivity of the waveguide to

the bulk RI change ∆nc and the thickness change of the biomolecule adlayer ∆tad

correspondingly. When the wavelength of the propagating wave and the physical

parameters of the waveguide structure are known, the sensitivities ∂neff/∂nc and

∂neff/∂tad can be defined by the following equations according to (Tiefenthaler &

Lukosz 1989)

ρ

−

Δ=

∂∂

12

2

,

c

eff

eff

cf

eff

c

c

eff

n

n

t

z

n

n

n

n , (12)

( ) ( )( ) ( )

ρ

−+

−+−−−

=∂∂

1

122

22

22

2222

feffceff

adeffceff

cf

cad

effeff

efff

ad

eff

nnnn

nnnn

nn

nn

tn

nn

t

n , (13)

27

where ρ = 0 for TE modes and ρ = 1 for TM modes. nad is the refractive index of

the adlayer and teff is the effective waveguide thickness given by

=

Δ+=csj

jffeff ztt,

,, (14)

tf is the waveguide thickness and ∆zf,j is the penetration depth of the evanescent

field into the substrate (j = s) or cover medium (j = c).

( ) ( )[ ]22

0

122

220

,

11

jeff

jefffeff

jeff

jfnnk

nnnn

nnkz

−

−++

−

−=Δ−

ρρ , (15)

where ρ equals 0 for the TE mode and 1 for the TM mode. Using Ormocore for

the waveguide core, Ormocomp for the substrate and water for the cover medium,

the sensitivity of the waveguide can be studied with respect to the different

waveguide thicknesses in both TE0 and TM0 modes. Sensitivity to the RI change

of the cover medium and the adlayer thickness are plotted in Fig. 5(a) and (b),

respectively.

0.3 0.6 0.9 1.2 1.5 1.8 2.10.000

0.002

0.004

0.006

0.008

0.010

0.012

0.014

∂nef

f /∂n

c

Waveguide thickness [μm]

TE0

TM0

0.3 0.6 0.9 1.2 1.5 1.8 2.10.0

5.0x10-6

1.0x10-5

1.5x10-5

2.0x10-5

TE0∂n

eff

/∂t ad

Waveguide thickness [μm]

TM0

Fig. 5. Waveguide sensitivity to a) the change of cover medium refractive index and b)

adlayer thickness plotted against the waveguide thickness (Paper IV, published by

permission of (c) 2012 European Optical Society).

The parameters used in this simulation are λ = 633nm, nc = 1.333, nf = 1.553,

ns = 1.520, and nad = 1.45 for protein adlayers (Schmitt 2008). The refractive

indices used here correspond to the polymers used later in the fabrication process

described in Section 3. In terms of processing, it is easier to construct waveguides

with a thickness of around 1 µm, and light coupling is easier with thicker

waveguides. As a result of these considerations, the TM coupling mode was

chosen to enable higher sensitivity. According to simulations, the highest

28

sensitivities are achieved at a waveguide thickness of 580 nm with maximum

values of ∂neff/∂nc = 0.0134 and ∂neff/∂tad = 2.1×10-5 nm-1.

With longer radiation wavelength, the penetration depth would become longer.

However, in usual biosensing applications with the size of biomolecules up to a

few nanomemters, a much longer penetration depth is not necessary. Besides, in

the application point of view, visible wavelength is much easier to handle in cases

like light coupling and detection. Therefore the visible wavelength was selected

for measurements in this thesis. The near infra-red wavelength was only used for

testing purposes.

29

3 Sensor design, fabrication and characterization

This section provides a detailed description of the material, structure, fabrication

process and device characterization of the developed sensors. Owing to the

unique benefits possessed by the different Young interferometer structures, both

the vertically-spaced dual-slab and the conventional Y-junction based structure

described in Subsection 2.2.1 were utilized in the fabrication of the polymer-

based interferometric sensors.

3.1 Polymers in integrated optics

Research on the use of organic materials for integrated optics started in the early

1970s (Tien 1971). As polymers represent a new material available for optical

integrated optics, they are less studied than semiconductors and inorganic

dielectrics within the telecommunications industry (Paloczi 2005). In recent years,

the driving force of the market has greatly increased the demand for large

quantities of inexpensive IO devices, and polymeric materials certainly satisfy

this requirement.

Polymers possess many good features, including high optical transmittance,

low dispersion and a refractive index that is tunable within a limited range. Most

importantly, polymers offer versatile processability. Thus, in contrast to

semiconductor materials that require several successive thin-film growths and

photolithography with expensive equipment, optical polymers lend themselves to

flexible, large area, low-cost fabrication of waveguide devices through simple

spin-coating techniques and low-temperature processing. Moreover, they are fully

compatible with semiconductor electronics. By controlling the ratio of the

polymer diluted in the solvent and spinning speed, a coated film can achieve a

thickness in the range of 0.1 to 100 µm (Ma et al. 2002). Due to easy

processability, a variety of substrates are compatible for polymer waveguide

devices, such as glass, silicon dioxide, silicon, indium phosphide—and even

certain types of polymers (Chang-Yen et al. 2005).

The fabrication process of polymer waveguides will be summarized here

briefly. For slab waveguides, where the optical field is vertically confined, spin

coating or dip coating are usually applied to the deposition of multilayer thin

films on the chosen substrate. For rib waveguides, confinement of the optical field

in the lateral direction requires patterning of the polymer. Various techniques

30

available for patterning fall into the following categories: photoresist-based

photolithography (Usui et al. 1996), direct photolithography (Cha et al. 2004),

soft lithography (Unger et al. 2000) and moulding (Heckele & Schomburg 2004).

Photoresist-based photolithography, which utilizes masks, is directly adopted

from standard semiconductor processing technology, where photolithography is

performed on the photoresist layer coated on the polymer and the developed

pattern is transferred into the polymer by RIE. For photosensitive polymers,

direct photolithography can be used to pattern waveguides by conventional mask

photolithography or direct laser writing (Eldada et al. 1996). Photosensitive

polymers behave like negative tone photoresists, and polymerization is caused by

exposure to ultraviolet (UV) light. Being confined to the polymer fabrication

process, non-photolithographic techniques, i.e., soft lithography and moulding,

are not available for more conventional photonic materials, such as glass and

semiconductors (Eldada et al. 2000). These unique techniques for polymer

processing are simple to perform and require less complicated, low-cost

equipment. There is no clear distinction between these two techniques, since they

both use a mould to transfer patterns. However, soft lithography utilizes

polydimethylsiloxane (PDMS) moulds to transfer micropatterns. The moulds

themselves are made of hard materials, such as aluminium, nickel, silicon or even

hardened polymers (Wang et al. 2011), which allow quick prototyping. Often

used moulding techniques for thermal plastics are injection moulding and hot

embossing, while the UV nanoimprint lithography (NIL) technique is utilized for

photosensitive polymers.

UV-NIL is usually performed at low pressure and room temperature. By

pressing a transparent mould with nano-/microscale features on a UV-sensitive

thin resin layer or droplet and curing it under UV light, a replica of the mould is

formed. NIL relies on direct mechanical deformation of the resist material and can

therefore achieve resolutions beyond the limitations set by light diffraction or

beam scattering that are encountered in conventional photolithography (Guo

2007). This technique is capable of producing high-resolution nanoscale features

down to sub-10 nm with high throughput (Kuo et al. 2003).

3.1.1 Polymeric IO biosensors

Besides the easy processability mentioned above, another advantage of polymer-

based IO devices is their high potential for use as chemical and biological sensors,

31

because the organic groups in a polymeric compound can be designed and

tailored to react against a specific medium (Lifante 2003).

Several polymer-based planar waveguide sensors have been developed

recently, such as ring resonators (Chao et al. 2006, Kim et al. 2008), polarometric

interferometers (Kwon et al. 2009), Mach-Zehnder interferometers (Shew et al.

2008, Bruck et al. 2011) and grating couplers (Oh et al. 2006, Kim et al. 2010).

The sensitivity of these previously reported polymer biosensors for bulk refractive

index changes is in the order of sub-10-5 RIU. As the RI of polymer materials

ranges from 1.3 to 1.7, there is very little variation between different materials,

leading to a low RI contrast (typically < 0.3) between the core and the cladding in

a polymer waveguide. For the commonly used cladding material SiO2 (n = 1.46),

the highest sensitivity achievable in a three-layer waveguide system can be

calculated analytically, as shown in Fig. 6. . The range of core RIs is chosen from

1.5 to 2.1, since it includes typical polymer materials, as well as many commonly

used inorganic dielectrics, such as SiO2-TiO2, SiON, Si3N4 and Ta2O5. In the

same figure the optimized waveguide thickness was also plotted illustrating the

waveguide thickness at which the highest sensitivity was obtained.

Fig. 6. Sensitivity of a three-layer waveguide system, plotted as a function of the core

material refractive index. Also shown is the optimized waveguide thickness at which

the highest sensitivity is obtained.

It can be concluded from Fig. 6 that the sensitivity of a three-layer waveguide

system increases as the RI contrast between the core and cladding layers increases.

Nevertheless, relative to inorganic dielectrics, waveguides based on polymer

materials featuring a small RI difference have a reduced sensitivity. One direct

solution to increasing this sensitivity is to enlarge the RI contrast between the

32

cladding and the core. To that end, Horváth et al. utilized a nanoporous silica

substrate with a refractive index of 1.2, while other research groups exploited the

fluorinated polymer Cytop with a RI of 1.34 (Agnarsson et al. 2010, Kim et al.

2010) as cladding material. Another way to increase sensitivity is to deposit a

high RI coating above the polymer waveguide, such as TiO2 deposited on a

polymer waveguide (Kwon et al. 2009) and a Ta2O5 coating on a glass waveguide

(Qi et al. 2009).

3.1.2 Ormocer - organic and inorganic hybrid polymer

ORMOCER® is a series of inorganic-organic hybrid polymers commercialized by

micro resist technology GmbH in Germany. A combination of good optical and

dielectric properties makes it a good candidate for electro-optical applications

(Buestrich et al. 2001). It is highly transparent above the wavelength of 400 nm

and up to the near-infrared region. At 633 nm, optical loss of the series of

materials is less than 0.1 dB/cm.

The Ormocer series contains several different polymeric materials. Three of

them, Ormocore, Ormoclad and Ormocomp, were used in this work. Their

refractive indices are 1.553, 1.536 and 1.520, respectively, determined at the

wavelength of 633 nm with a Metricon prism coupler. Taking into consideration

the fabrication process, optical waveguides can be produced within one material

class, and the processing is compatible to conventional equipment used in thin-

film processing of wafer-scale devices. Moreover, Ormocer materials are easy to

handle and offer good adhesion on most substrates, such as Si wafers, inorganic

glasses and polymers. Most importantly, Ormocers are photosensitive, allowing

curing under UV exposure. The ability to be patterned by UV irradiation makes

them attractive for any kind of wafer-scale application (Houbertz et al. 2003).

Although they behave like negative photoresists, their thermal and chemical

stability is better, due to a backbone composed of inorganic Si-O-Si networks

(Streppel et al. 2003). This unique feature results in a decomposition temperature

of ≈ 270 °C, which is sufficiently high for most technical applications.

3.2 Dual-slab Young interferometers

One of the sensing structures investigated in this thesis is the dual-slab waveguide

structure. As shown in the schematic illustration in Fig. 3a, it is a five-layer

structure with alternatively stacked cores and claddings. When a light beam with

33

a diameter larger than the inter-cladding between both cores shines on the facet of

the multilayer stack, the beam split up and propagates in both cores. At the other

end of the stack, the two beams diffract out and interfere with each other,

generating a fringe pattern. The principle behind this phenomenon fits perfectly to

that of a YI. Unlike traditional horizontally-integrated Y-junction YIs, the dual-

slab Young interferometer is vertically spaced and does not need Y-junction

patterning for light splitting. Moreover, its vertically-integrated structure offers

the benefit of high tolerance to misalignment of the in-coupling light. Due to the

multi-layer structure of the slab waveguide interferometer, it is easy to fabricate

the sensor using polymer materials. Each layer can be deposited by spin-coating

and the thickness is adjustable by varying the coating speed.

3.2.1 Fabrication process and waveguide characterization

In the fabrication of the dual-slab Young interferometer-based polymer sensors

described here, Ormocomp was used as cladding layer and Ormocore as core

layer material. Single-crystal silicon 4” wafers were selected as substrates, owing

to their smooth end-facets after cleaving. Apart from mere mechanical support,

the substrate does not offer any additional functionality. Above the Si substrate, a

thick layer (>10 µm) of Ormocomp was deposited with the imprinting method by

pressing a blank glass wafer against the dispensed resin. After curing under

broadband UV exposure, the wafer was placed in a plasma etcher (Pico, Diener

Electronic GmbH) for 30 s to open the bonds on the imprinted Ormocomp layer.

Subsequent layers were deposited one after another by the spin-coating technique

with a common speed of 3000 rpm for 60 seconds. After the curing of each coated

layer, plasma etching was performed for 30 s, before coating of the next layer. Fig.

7a shows a cross-section image of the finished structure under SEM (JCM-5000

Neoscope, Jeol Ltd.). The thickness of the two waveguiding layers is ~1 µm and

that of the cladding layer between the two waveguides > 6 µm.

34

Fig. 7. Dual slab Young’s interferometer a) SEM cross-section image of the fabricated

multi-layer structure; b) the output beams at the end-facet; c) fringe pattern generated

at 150 µm distance (Paper II, published by permission of (c) 2012 OSA).

Fig. 7 b) shows the sensing and referencing beams at the output end-facet. In this

image, the light strips are wider than the actual waveguide thickness, because of

the diffraction effects in the imaging system. The slightly bright speckles above

and below the actual guiding modes come from stray light propagating through

the claddings. The interferometric fringe pattern (Fig. 7c)), taken at a distance of

150 µm from the end facet, gives a good visibility of ~0.8. A grid-like pattern,

seen inside each of the fringes, is produced by the interference with the stray light.

Based on the observed physical parameters of the multi-layer structure in Fig. 7,

theoretical sensitivities of the slab waveguide sensor can be obtained from Fig. 5.

Thus, the value of (∂neff/∂nc)TE = 0.0053 and (∂neff/∂tad )TE = 0.975×10-5 nm-1 in

TE- coupling mode, and (∂neff/∂nc)TM = 0.0072 and (∂neff/∂tad )TM = 1.16 ×10-5 nm-1

in the TM- coupling mode.

Simulation with the Fimmwave commercial software package (Photon

Design Inc.) showed that the two Ormocore waveguiding layers exhibit single

mode operation at the wavelength of 633 nm. As confirmed by the guided mode

propagation image on the left side of Fig. 8, single-mode propagation in x-

direction can be observed in both slabs. Visible as black lines in the mode profile

are the physical boundaries of the waveguides. A one-dimensional mode profile in

the x-direction was plotted beside the mode propagation xz-image, indicating that

the penetration depth above the sensing waveguide into the aqueous analyte is

less than that into the cladding. However, exceeding the 100 nm range, this

penetration depth is suitable for most of the applications involving protein

detections.

35

Fig. 8. Simulation of the guided mode through a dual slab Young’s interferometer (on

the left) and the generated fringe pattern in the far field (on the right) (Paper II,

published by permission of (c) 2012 OSA).

3.3 Inverted-rib waveguide Young interferometers

The rib waveguide Young interferometer utilizes the conventional Y-junction

splitter as the basic structure of the interferometric sensor. Although rib-

waveguide fabrication requires patterning, it can be easily implemented for

polymers using moulding techniques. With UV imprinting, patterns can be

replicated rapidly at room temperature by pressing a mould onto a UV-sensitive

resin, in this case Ormocer, and curing it under UV exposure. By using the

inverted-rib waveguide structure, the fabrication process can be further simplified.

In the fabrication of inverted waveguides, patterning is performed on the under-

cladding of the waveguide. The core layer is then simply spin-coated on the

under-cladding to fill the patterned grooves. By tuning the material’s viscosity

and the spin-coating speed, the thickness of the residual layer can be finely

controlled. In this case, RIE treatment is not required to remove the residual layer

as in a normal rib waveguide for a better confinement of the propagating mode

(Hiltunen et al. 2009).

3.3.1 Inverted-rib waveguide fabrication and characterization

Inverted-rib waveguide fabrication consists of two parts: mould fabrication and

waveguide replication, as shown in Fig. 9 a) and b), respectively. Meld fabrication

comprises a photolithographical patterning of a positive photoresist (Ultra-i 123 I-

line, Rohm and Haas), and a first imprinting step on Ormocomp to form the

36

mould for the next step. Cross-section SEM (JCM-5000 Neoscope, Jeol Ltd.)

images of the patterned positive photoresist and the imprinted mould are shown in

Fig. 10 a) and b). A patterned positive tone photoresist can only withstand one

imprinting process. To avoid repeating the lithographic steps in NIL mould

fabrication, an Ormocomp-based mould was formed with improved chemical and

mechanical stabilities compared to a mould made of a positive tone photoresist.

Fig. 9. Process flow a) mould fabrication; b) Inverted-rib waveguide fabrication (Paper

III, published by permission of (c) 2011 Elsevier).

Before the subsequent imprinting process, the Ormocomp mold was first

deposited with tridecafluoro-(1,1,2,2)-tetrahydrooctyl-trichlorosilane (F13 –TCS)

anti-adhesion material.

37

Waveguide fabrication started with patterning the waveguide lower cladding

on a 4-inch Si wafer. After dispensing Ormocomp on the Si wafer, the mold was

pressed on top of it by a laboratory-customized imprinting set-up to transfer the

pattern. Near-UV flood exposure (EFOS Lite E3000) was applied through the

transparent mould to cure the soft polymer layer. The resulting replica has a

groove that resembles the ridge in the inverted-waveguide structure. The height of

the ridge was about 1.2 µm, defined by the thickness of the spin-coated positive

tone resist in the first step. After being released from the mould, the formed

replica was rinsed with acetone and de-ionized water. Fig. 10 c) shows a cross-

section SEM image of the replica with a 2-µm wide groove. The material used for

the core layer was Ormocore. To reduce the thickness of the residual layer,

Ormocore was diluted in Ormothin with a ratio of 1:4 (v/v). The dilution was then

spin-coated onto the replica with a speed of 3000 rpm for 60 seconds and left at

room temperature for 1–2 min to allow the solvent to evaporate before UV curing.

A SEM cross-section image of the fabricated inverted-rib waveguide is shown in

Fig. 10 d).

Fig. 10. Cross-section SEM images of a) patterned positive photoresist; b) imprinted

Ormocomp mould; c) Ormocomp replica imprinted from the mould; d) spin-coated

Ormocore on the Ormocomp replica (Paper III, published by permission of (c) 2011

Elsevier).

38

The root-mean-square (RMS) surface roughness of the replica (Fig. 10 c)) and of

the inverted-waveguide structure (Fig. 10 d)) were examined using atomic force

microscopy (AFM, Veeco Dimensions 3100). An average of 6 measurements

showed that the RMS surface roughness of the replica (Ormocomp lower

cladding layer) was 0.54 ± 0.09 nm, while that of the inverted-rib waveguide was

1.61 ± 0.19 nm. Thus, the RMS surface roughness of the replica was in fairly

good agreement with the with RMS values for positive-tone photoresist (Fig. 10

a)) and the mould (Fig. 10 b)), 0.48 ± 0.02 nm and 0.8 ± 0.2 nm, respectively,

reported by Jussi et al. This demonstrates that a fabrication process with two

consecutive imprinting procedures is capable of producing a very smooth surface

on the moulded lower cladding layer in the inverted-waveguide structure. Also

thickness variations of the spin-coated core layer were inspected under SEM. In

the 4 cm range, the core layer’s thickness was about 0.9 µm ± 0.1 µm, confirming

that the spin-coated core layer is fairly uniform. Based on visual inspection, the

core layer uniformly covered the entire 4" substrate.

A detailed characterization of the fabricated inverted-rib waveguide was

presented in Paper III. With an excitation light source of 1310 nm, the waveguide

operated in single mode, achieving a transmittance of close to 60% in a 3 cm long

waveguide. This result is promising even for IO interconnection applications in

the telecommunications industry. Although scattering is expected to be higher at

632.8 nm, the wavelength used in the experiments reported in this thesis, than in

the infrared range, the transmitted light power is still sufficient for biosensing

applications.

3.3.2 Rib-waveguide for Young interferometers

The upper-cladding of the rib-waveguide Young interferometer was formed by

spin-coating a thick ~ 20 µm layer of Ormoclad above the core. A sensing

window was then opened above the waveguide using photolithography. Finally,

the fabricated waveguide sensor was hard-baked for 2 hours in an oven at 150°. A

schematic design of the Young interferometer sensor based on a Y-junction rib-

waveguide is shown in Fig. 11, with corresponding microscope images. These

microscope illustrations omit the middle section of the interferometer due to the

relatively long structure of the waveguide.

39

Fig. 11. Top view images of the rib-waveguide Young interferometer taken with a

microscope, with a schematic illustration shown above (Paper IV, published by

permission of (c) 2012 European Optical Society).

From the input to the output end-facet, the interferometer measured 20 mm in

length with a 10 mm sensing window opened above the reference waveguide. The

design is symmetrical with exactly the same dimensions for the sensing and

referencing waveguide. Apart from the sensing window, the remaining sections of

the waveguides were completely covered by Ormoclad cladding. The distance

between the sensing and referencing waveguide was 50 µm, while the dimension

of the cross-section was 2 µm (width) × 400 nm (thickness), with a 500 nm

residual slab layer over the groove, left from the spin-coating process. To

maximize sensitivity, the thickness of the waveguide was reduced in both rib and

slab composition elements. The positive photoresist was diluted in ethyl lactate

with a ratio of 2:1 (v/v) to reduce the thickness of the original mould, and the

Ormocore was diluted in ma-T 1050 with a ratio of 1:4 (v/v) to reduce the

thickness of the residual layer. With an overall thickness of 900 nm, the

waveguide has a theoretical sensitivity of (∂neff/∂nc )TE = 0.0064, (∂neff/∂tad )TE =

1.2×10-5 nm-1 in the TE coupling mode and (∂neff/∂nc )TM = 0.0087, (∂neff/∂tad )TM =

1.4×10-5 nm-1 in the TM coupling mode, according to Fig. 5.

A Fimmware simulation of the supported mode was conducted on the

fabricated rib-waveguide interferometer, both in the sensing window and in areas

with the over-cladding layer. In this simulation, the sensing window was assumed

to be filled with water. It was found that two supported TM modes exist in the

designed rib-waveguide at 632.8 nm, as shown in Fig. 12. However, only the

fundamental mode can propagate outside the sensing area, i.e., in areas with over-

cladding. As the first-order mode is filtered away outside the sensing window,

where Ormoclad cladding with a higher refractive index is applied, the

assumption of single-mode operation is considered valid.

40

Fig. 12. Simulated mode profiles at 632.8 nm of the inverted-rib waveguide a)

fundamental TM00 mode in the sensing window; b) first-order TM01 mode in the

sensing window; c) fundamental TM00 mode in the waveguide covered with Ormoclad

cladding (Paper IV, published by permission of (c) 2012 European Optical Society).

Fig. 13 a) shows light intensity at the output facet of the sensor chip. Two

localized spots can be observed, confirming proper waveguide operation. Fig. 13

b) shows the corresponding interference pattern. Imperfections in the fringe

pattern result from stray light interference.

Fig. 13. Captured images of a) output modes at the end facet of the sensing chip; b)

generated fringe pattern with 632.8 nm TM-polarized He-Ne laser input (Paper IV,

published by permission of (c) 2012 European Optical Society).

3.4 Ta2O5 high refractive index coating

The polymer sensor based on Ormocomp and Ormocore has a low RI contrast of

0.033 at 633 nm, resulting in a low sensitivity for the polymer sensor. To enhance

the sensitivity, a thin layer of Ta2O5 (n = 2.1) high index coating was reactively

sputtered using a 4N pure Ta metal target (Helios XL Pro, Leybold Optics) on the

inverted rib waveguide interferometer. Ta2O5 was deposited at near room

temperature. It causes minimal thermal stress mismatching with the underneath

a b c

a b

41

polymer layer, thus very good film quality can be achieved with no cracks. Ta2O5

has very low optical loss at both visible and infra-red wavelengths making it ideal

for planar waveguide applications (Pearce et al. 2012).

By modifying the propagation mode in the sensing window with the high

index coating, sensitivity can be enhanced. Only the section of the waveguide

inside the sensing window was in direct contact with Ta2O5, other sections were

covered with Ormoclad upper-cladding before sputtering. Fig. 14 shows a SEM

cross-section image of the Ta2O5-coated polymer waveguide inside the sensing

window. The sputtering process was performed at Southampton University

through Photosens project collaboration. A special sputtering system was used to

maintain a near-room temperature during the sputtering process to prevent cracks

forming on the polymer waveguide.

Fig. 14. Cross-section SEM image of the inverted rib waveguide inside the sensing

window (Paper V, published by permission of (c) 2012 OSA).

In the fabricated integrated YI, the distance between the sensing and referencing

waveguide was 50 µm. Part of the YI sensor is shown in the top view microscope

image in Fig. 15 a). Fimmwave simulations demonstrated that, at 633 nm, the

sensing waveguide only supports one TM propagation mode in the sensing

window, where water serves as upper-cladding, with a refractive index of 1.33.

Single-mode operation also proved the only option available for areas with

Ormoclad over-cladding. The coupling efficiency Γ2 between the two regions in

the sensing waveguide was defined as an overlap integral (Reed & Knights 2004)

according to

( )

⋅

⋅=Γ

dxdyyxEdxdyyxE

dxdyyxEyxEyx

OTaclad

OTaclad

),(),(

),(),(),(

22

2

2

52

52 . (16)

42

Eclad and ETa2O5 are the electric field distributions in the Ormoclad covered region

and in the Ta2O5 coated region, respectively. Theoretical intensity profiles

obtained by the film-mode-matching (FMM) method are shown in Fig. 15 a) at

the indicated locations in the sensing waveguide.

Fig. 15. a) microscope image showing a section of the YI sensor where the adiabatic

transition of the high-index Ta2O5 coating takes place. Displayed here are intensity

profiles of the propagating TM modes in three places: Ormoclad cladding, the front

edge of the sensing window, where the thickness of the Ta2O5 coating is about 20 nm,

and the middle section of the sensing window, where the Ta2O5 coating thickness is 80

nm. Inside the sensing window, water is considered upper-cladding material. b) The

generated interference pattern at a distance of 2.5 mm from the edge of the sensor

chip (Paper V, published by permission of (c) 2012 OSA).

By using the shadow masking technique in the Ta2O5 sputtering process, an

adiabatic region with gradually increasing Ta2O5 thickness was formed between

the Ormoclad-covered region and the Ta2O5-coated region. Without adiabatic

transition, the coupling efficiency between the two regions was calculated to be

60%. With adiabatic transition, starting from 20 nm of Ta2O5 coating at the front

edge of the sensing window, the obtained coupling efficiency was approximately

90%. Assuming that power loss occurs at the front and the back edges of the

sensing window, the overall transmission efficiency was estimated to be around

81%. As evaluated by Pearce et al., the intrinsic material loss of a sputtered Ta2O5

thin film is about 1 dB/cm. The propagation loss in a composite polymeric

waveguide with 100 nm thick Ta2O5 layer was estimated to be about 1.6 dB/cm

(Jussi et al. 2011). These attenuation values are small enough to have a good

interference fringe visibility at the output of the sensor chip facet. The generated

fringe pattern from the referencing waveguide and sensing waveguide with

43

adiabatic transitions is shown in Fig. 15 b). The clear visibility of the

interferogram gives a good resolution for analyzing fringe pattern phase shifts by

a two-dimensional (2D) fast Fourier transform (FFT). Like the fringe patterns

generated by the other two sensor structures: the polymer dual-slab and the

polymer inverted rib waveguide sensor structures, the grid-like patterns contained

in the fringe pattern was caused by the inference with the stray like. By using the

2D-FFT analyzing method, the high frequency noise does not degrade the

determination of the phase.

IO waveguide biosensors respond to changes in the bulk refractive index of

liquid samples and the surface adsorption of biomolecules. The bulk RI change of

a liquid sample is seen as a change in the refractive index of the upper-cladding nc.

The adsorption of biomolecules translates into a thickness change of the adlayer

tad, growing above the sensing waveguide. According to Equation (11), the

induced effective RI change Δneff of the sensing waveguide is a function of the

bulk refractive index change Δnc and the growth of the adlayer thickness Δtad.

Using the observed physical parameters of the inverted-rib waveguide structure

shown in Fig. 14, theoretical sensitivities of ∂neff/∂nc and ∂neff/∂tad were examined

using Fimmwave. To calculate ∂neff/∂tad, nad = 1.45 was used for the RI of the

adsorbed biomolecule layer. With increasing coating thickness, the optical field

distribution of the propagating mode is pushed up towards the waveguide surface,

resulting in increased power in the evanescent field, thus improved sensitivity. As

can be seen from Fig. 16 a) and b), the curves for homogeneous refractive-index

sensing and surface-adsorption sensing resemble each other, though there is a

slight Ta2O5 thickness difference for the maximum sensitivity. For the TE

polarization state, sensitivity increases with increasing Ta2O5 thickness up to

about 60 nm for homogenous refractive index sensing and up to about 70 nm for

adsorption sensing, after which further increases in the thickness of the high-

index layer do not improve sensitivity. Corresponding thicknesses for the TM

polarization state were 120 nm and 140 nm, respectively.

44

TE TM

0 20 40 60 80 100 120 140 160

0.00

0.04

0.08

0.12

0.16

0.20

0.24

∂nef

f/∂n c

Ta2O

5 thickness [nm]

a

0 20 40 60 80 100 120 140 160

0.0000

0.0001

0.0002

0.0003

0.0004

0.0005

∂nef

f/∂t ad

[nm

-1]

Ta2O

5 Thickness [nm]

TE TM

b

Fig. 16. Sensitivity of a) homogeneous refractive index sensing and b) surface sensing

of biomolecular adsorption as a function of the thickness of the Ta2O5 coating for both

TE- and TM- coupling modes (Paper V, published by permission of (c) 2012 OSA).

The observed enhancement in sensitivity is associated with decreased

confinement inside the inverted guiding ridge. When a much thicker high-index

coating is deposited on the polymer layers, the composite waveguide structure

becomes leaky. To optimize sensitivity and mode confinement in the sensing

region, 80 nm coating and TM polarization were chosen. According to

simulations, the obtained sensitivity of 80 nm Ta2O5 coating is 0.12 and 2.1×10-4

nm-1 for homogeneous RI and surface molecular adsorption sensing, respectively.

This sensitivity exceeded that of a polymer waveguide sensor without high-index

coating by over 40 times, while still retaining 17% power confinement inside the

inverted ridge, ensuring in proper waveguide operation.

45

4 Biosensor experiments and results

This thesis developed three polymer Young interferometer sensors based on

different waveguide configurations: dual-slab waveguide Young interferometer,

inverted-rib Young interferometer and inverted-rib Young interferometer with

Ta2O5 coating. For clarity, Fig. 17 shows schematic cross-section images of these

different structures. Several sensors were produced with each different structure,

only the best three samples were regarded in this thesis.

Fig. 17. The schematic cross-section images of the waveguide used to build a) dual-

slab waveguide Young interferometer sensor; b) polymer inverted rib waveguide

Young interferometer sensor; c) polymeric-inorganic composite waveguide Young

interferometer sensor with inverted rib waveguide structure.

Sensor performance was evaluated in terms of homogeneous refractive index

sensing and surface sensing of biomolecule adsorption. Glucose deionized water

solution was used to demonstrate the homogeneous refractive index sensing. To

demonstrate the polymer sensors’ capacity for biosensing, an immunoassay was

performed utilizing specific binding between antibodies and antigens.

4.1 Measurement setup

Shown in Fig. 18 is the setup for aqueous-based sensing measurements, using a

linearly polarized He-Ne laser (632.8 nm, 5 mW) as light source. A linear

polarizer was placed in front of the laser to excite the TM-polarized mode.

Incident light was coupled into the waveguide through a 60x objective lens by the

end-fire coupling method. At the output facet of the interferometer, a 40x

objective lens was used to collect and magnify the fringe pattern image. This

image was then captured by a CMOS camera (PixeLink) with a resolution of

1280×1024. The setup did not contain any temperature controlling units. Because

46

the experiments were carried out in an aqueous environment, a continuous-flow

syringe pump (OPAM Instruments) was used to deliver the analyte onto the

sensing window through a flow cell attached above the sensor chip. A silicone-

based gasket having an opening window with a volume of 10 µl was placed

between the sensing chip and the flow cell to prevent leakage of the liquid

samples at the interface of the chip and the cell.

Fig. 18. Biosensing measurement set-up.

Fringe pattern images were captured at a rate of 1Hz. To extract spatial changes in

the fringe pattern, representing phase changes of the interferogram, Matlab

software was used. This software implemented a 2D-FFT algorithm to calculate

the amplitude and phase spectrum of the fringe pattern image. The corresponding

phase of each image was extracted by looking up the value of the same frequency

showing the peak of the amplitude spectrum. Phase discontinuities were corrected

by adding multiples of ±2π when absolute jumps between consecutive extracted

phases were greater than or equal to the default jump tolerance of π radians. After

phase response extraction, Δneff could be derived using Equation (5).

4.2 Effect of water absorption in polymer waveguides

Due to their molecular structure, polymeric materials are permeable to water

vapour and able to absorb water molecules to a certain extent (Mark 2007). When

a polymer absorbs water molecules, its refractive index changes in response to

changes in density and molar refraction (Shioda et al. 2003). The whole process

of introducing water to the sensing window of the developed sensors is

summarized by the following description. When water is introduced to the

sensing window of a polymer waveguide sensor, an abrupt change of the phase

47

was observed in the beginning as shown in Fig. 19 for all three different

waveguide structures. This phase jump was triggered by the immediate change of

refractive index in the sensing window from air (n = 1) to water (n = 1.33). This

swift change produced large spikes at the location where a grey area is masked

over the sensorgrams. Inside these regions, phase responses could not be

explicitly defined. At the meanwhile, the polymer-based sensing waveguide

started to absorb water molecules, thereby increasing its refractive index. As seen

in Fig. 19 a) and b), this caused a gradual phase drop, which cancelled out the

initial phase rise resulting from the refractive index change in the sensing window.

0 40 80 120

-20

0

20

40

water

Ph

ase

[rad

]

Time [min]

air

0 20 40 60-4

0

4

8

12

Pha

se [

rad]

Time [min]

waterair

0 2 4 6 8 10

-3

-2

-1

0

Phase

[rad]

Time [min]

water

air

a) b) c)

Fig. 19. Sensor responses after applying water to the sensing window of a) dual-slab

Young interferometer; b) inverted-rib Young interferometer; c) inverted-rib Young

interferometer with Ta2O5 coating (Modified from Paper II & V, published by permission

of (c) 2012 OSA).

In Fig. 19 a) and b), the phase converged gradually, when the diffusion of water

molecules reached equilibrium. An explanation for the different phase changing

rates may be offered by the different surface hydrophilicity of the different

sensing surfaces. The sensing surface the inverted-rib waveguide sensor was

treated lightly with O2 plasma, before the deposition of Ormoclad upper cladding,

after which it underwent a photolithographic treatment. On the other hand, the

sensing surface of the slab waveguide sensor was untreated before applying water

to the sensing window. The different phase responses in Fig. 19 a) and b) are

caused by a different interaction length. In Fig. 19 c), the phase response

stabilized in a couple of minutes once the sensing window was completely filled

with water. This indicates that an 80 nm Ta2O5 coating effectively blocked the

penetration of water into the polymer waveguide underneath.

48

4.3 Sensing of bulk refractive index

In aqueous environments, when water molecule absorption in the polymer

waveguide reaches a saturation point, a flat baseline can be obtained for

subsequent measurements. In homogeneous sensing, the liquid sample above the

sensing window serves as cladding material. When the sample changes, a change

is observed in the homogeneous refractive index of the cladding. Experiments

focusing on homogeneous refractive index sensing assumed that no molecule

adsorption takes place on the surface of the sensing waveguide.

In this research, solutions of DI water, diluted with glucose (D-glucose,

Sigma-Aldrich) to various concentrations ranging from 0.005% ~ 1%, were used

for homogeneous bulk RI sensing. Glucose water solution was used for sensor

calibration due to certain content of glucose contained in human blood varying

between 30mg/dl (0.03%) and 500mg/dl (0.5%). Between each measurement, DI

water was used as a buffer solution to flush the sensing window and bring the

sensor response to the baseline. Each measurement on a glucose solution sample

took five minutes with a pumping rate of 0.5 ml/min. Fig. 20. shows the phase

responses of the three different sensors to low-concentration glucose solutions.

The red curves superposed on the original phase response in Fig. 20 a) and b) are

the phase responses after low-pass filtering. Complete phase responses can be

found in Papers III, IV, and V.

49

0 5 10 15 20

0.0

0.1

0.2

Pha

se [

rad]

Time [min]

0.02 %

2.9*10-5

0.04 % 0.06% 0.08 %

5.9*10-5 8.8*10-5 1.2*10-4

0 5 10 15 20

0.00

0.02

0.04

0.06

0.08

0.06% 0.08%0.04%

Pha

se [r

ad]

Time [min]

0.02%

0 5 10 15 20

0.0

0.3

0.6

0.9

Pha

se [

rad]

Time [min]

0.005%7*10-6

3*10-5

0.02%

0.04%6*10-5

7*10-5

0.05%

Fig. 20. Phase response of a) slab waveguide b) polymer inverted-rib waveguide c)

polymeric-inorganic composite waveguide interferometer sensor to the bulk refractive

index change when applying glucose de-ionized water solutions (Modified from Paper

II, IV & V, published by permission of (c) 2012 European Optical Society and OSA).

Homogeneous sensing of liquid samples with a known RI is usually performed

for the purpose of sensor calibration; a theoretically calculated sensitivity to Δnc is

compared to the actual sensitivity obtained from measurement data to evaluate

how close the fabricated sensor is to the physical model used in the theoretical

estimation. Fig. 21 plots the derived ∆neff from the phase responses calculated

using Equation (5) with respect to the actual refractive index changes obtained

from the concentration of the analyte with the empirical equation (Weast 1974):

ionConcentratnc ×=Δ 14713.0 , (17)

a

b c

50

0.0 4.0x10-4 8.0x10-4 1.2x10-3

0.0

4.0x10-5

8.0x10-5

1.2x10-4

1.6x10-4

slope = 0.0071 slope = 0.0086 slope = 0.1061

Δnef

f

Δnc

Slab Rib Rib+Ta

2O

5

1E-5 1E-4 1E-310-7

10-6

10-5

10-4 Slab Rib Rib+Ta

2O

5

Δnef

f

Δnc

Fig. 21. Effective refractive index changes ∆neff derived from the phase responses of

the three different sensors, plotted against the actual refractive index changes ∆nc of

the glucose solutions a) in the linear scale b) in the log scale.

The data points are measurement results extracted from Fig. 20, and the lines are

linear fittings of those points. The sensitivity ∂neff/∂nc of homogeneous sensing

was determined from the slope of the linear regressions in Fig. 21 a). The slope of

the Ta2O5-coated sensor is much steeper than that of the other two sensor

configurations. At lower glucose concentrations, the error bars, as shown in Fig.

21, are larger for slab and Ta2O5-coated rib waveguide sensors. This is because

the signal-to-noise-ratio (SNR) was quite low at low concentrations, making the

phase value extracted from the sensor response curves (Fig. 20) more ambiguous

than at higher concentrations. Another reason could be lower glucose

concentration accuracy at low concentrations due to errors in preparing the

glucose dilutions. The theoretical sensitivity ∂neff/∂nc of the sensors to bulk

refractive index changes and the sensitivity extracted from the sensor responses

are listed Table 1.

Table 1. Theoretical and measured sensitivity ∂neff/∂nc of all three sensor

configurations in the TM- coupling mode.

Slab Rib Ta2O5

Interaction length [cm] 1.3 0.6 0.9

∂neff/∂nc theoretical 0.0072 0.0087 0.1200

measured 0.0071 0.0086 0.1061

Sensitivities extracted from the sensor responses correspond very well with the

theoretically calculated values. This confirms that the physical models used for

b a

51

theoretical sensitivity calculations and the fabricated waveguide sensors are

closely identical.

4.4 Immunoassay

4.4.1 Materials

The performance of the sensors developed for biosensing applications were

examined by detecting specific molecular binding events between C-reactive

protein (CRP antigens) and monoclonal anti-human C-reactive protein (CRP

antibodies). A human CRP molecule (having a molecular weight of 115,135 Da)

is composed of five identical subunits (molecular weight 23,027 Da) forming an

annular configuration with cyclic pentameric symmetry (Black et al. 2004), as

shown in Fig. 22. Each subunit contains an epitope, which is specifically

recognized by the antibody. Thus, a single CRP molecule can be bound by five

different antibody molecules.

Fig. 22. Crystal structure of complexed C-reactive protein (Black et al. 2004, published

by permission of (c) 2004 ASBMA).

The median CRP concentration in healthy adult blood is around the level of 0.8

µg/ml. Following an acute tissue damage or inflammation the concentration can

reach a level over 1000-fold within 24–48 hours (Steel & Whitehead 1994). CRP

quantification in plasma or serum can also provide valuable clinical information

for the diagnosis, prognosis and monitoring of cardiovascular diseases. Research

shows that, in human blood, a CRP level between 1 µg/ml to 3 µg/ml constitutes a

52

detection limit for mildly elevated risk of a cardiac event (Pepys & Hirschfield

2003, Verma et al. 2004).

In the experiments reported here, phosphate-buffered saline (PBS) (10 mM

Na2HPO4, 1.8 mM NaH2PO4, 0.1 M NaCl, 2.7 mM KCl pH 7.4) was used as

buffer solution. C-reactive protein (Schripps Laboratories) as CRP antigens and

monoclonal anti-human C-reactive protein antibody (Medix Biochemica) as CRP

antibodies were prepared in the PBS buffer. Analyte concentrations of 2 µg/ml

(16 nM) and 0.1 µg/ml (0.8 nM) were prepared of CRP antigens, while the

concentration of CRP antibodies, functioning as receptors attached on the sensing

surface, remained constant at 0.5 mg/ml. To prevent non-specific binding on the

sensing surface, bovine serum albumin (BSA) with a concentration of 2 mg/ml

was used as blocking molecule. In negative control measurements, instead of CRP

antibodies, 0.5 mg/ml of mouse IgG (Jackson ImmunoResearch Laboratories, Inc)

was employed as receptor to examine non-specific binding. Moreover, to confirm

the specific binding of CRP antibodies and antigens, 13.2 µg/ml CRP antibodies

labeled with Alexa 546 fluorescent markers were used as secondary antibodies

with fluorescent sandwich-type immunoassay.

4.4.2 Experiments

Measurements started when the baseline of the sensorgram stabilized, reaching a

flat level in a few hours after the presence of a PBS buffer. In the case of slab

waveguide sensors, receptors were attached in advance by printing. In this

method, a Dimatix material inkjet printer was used to print CRP antibodies (0.5

mg/ml in PBS buffer) over an area of 15×2 mm2 on the sensing waveguide. For

rib waveguide sensors, the sensing surface was activated by pumping CRP

antibodies through the sensing window. When the antibodies passed through the

sensing waveguide, those flowing close by the sensing surface were non-

specifically adsorbed on the surface of the hydrophobic Ormocore waveguide

(Step 1). To increase the probability of this happening, the process was allowed to

proceed for a considerable period of time. Fig. 23 provides a schematic

illustration of the odd steps of the measurement process.

53

Fig. 23. Illustration of the measurement performed to detect the surface adsorption of

biomolecules including specific binding between CRP molecules and their antibodies,

and the negative control measurement performed with mouse IgG. (Paper V, published

by permission of (c) 2012 OSA).

Even steps, not shown in this figure, consisted of rinsing with the PBS buffer after

each injection of biomolecules to flush away loosely attached and excessive

molecules. After surface activation with CRP antibodies, BSA was injected to

block any free spots on the waveguide surface to which CRP could bind

unspecifically (Step 3). After rinsing with PBS, CRP antigens were pumped

through the sensing window to bind specifically to the CRP antibodies on the

sensing surface (Step 5). In Step 7, secondary CRP antibodies, labeled with

fluorescent markers, were injected to confirm the measurement results under a

fluorescence stereomicroscope (SteREO Discovery.V8, Carl Zeiss). This was only

done after the injection of 0.1 µg/ml CRP analyte and after the negative control

on the slab waveguide and Ta2O5-coated rib waveguide sensor due to relatively

low surface mass densities involved in these experiments.

There are two reasons to use fluorescent-labelled secondary antibodies.

Firstly, at low concentrations, signal amplification may be needed. Since each

CRP antigen contains five identical epitopes, four will be left unoccupied when an

antigen binds to a CRP antibody on the sensing surface. By applying a high

concentration of secondary antibodies, all available epitopes can be expected to

become occupied, leading to amplified surface adsorption due to the comparable

molecular weight of the CRP antigens (115 kDa) and antibodies (150 kDa).

Secondly, non-specific binding needs to be verified on both Ormocore and Ta2O5-

based sensing surfaces.

For each of the three sensors, a negative control was performed with mouse

IgG as receptor, due to its low affinity toward CRP. The purpose was to ensure

54

that when a CRP analyte was applied, the sensor response was stimulated by

specific affinity bindings between antibodies and antigens, rather than direct

attachments of CRP on the bare sensing surface. This concern was raised, because

the receptors, i.e., CRP antibodies, are attached on the waveguide surface by weak

hydrophobic interactions. As a result, they can detach from the surface and be

flushed away in the later steps of the measurement. This allows CRP antigens to

occupy the empty spaces left behind.

4.4.3 Results

Phase responses of the three different sensors for the surface sensing of

biomolecular adsorption are shown in Fig. 24.

0 20 40 60 80 100 120

0

1

2

Pha

se [

rad

]

Time [min]

CRP 2 μg/ml Negative control

6

3

45

7

0 30 60 90 120 150 180

0

2

4

6

8

10

1

Pha

se p

er u

nit i

nter

actio

n le

ngth

[rad

]

Time [min]

CRP 2 μg/ml Negative control

62 3

4 5

160 180 200 220

56,7

57,0

57,3

120 140 160

57

60

63

66

69

72

75

0 50 100 150 200 250

0

20

40

60

80

2 μg/ml 0.1 μg/ml negative

Pha

se p

er u

nit

inte

ract

ion

leng

th [r

ad]

Time [min]

1

6

2 3

47

8

5

Fig. 24. Phase responses per unit interaction length (1 cm) to surface biomolecule

adsorption and negative control of a) slab waveguide b) polymer inverted-rib

waveguide c) polymeric-inorganic composite waveguide interferometer sensor.

55

Images in the lower left corner are fluorescent microscopic images taken inside the

sensing window following a flow of secondary CRP antibodies. Numbers correspond

to the steps in the measurement process (Modified from Paper IV & V, published by

permission of (c) 2012 European Optical Society and OSA).

As shown in Fig. 24 a) for the slab waveguide sensor, the measurement started

with a BSA blocking step, since the attachment of CRP antibodies was prepared

in advance by inkjet printing. After rinsing, a clear rise of the phase occurred at

0.41 rad, following the injection of CRP antigens. In the negative control, no

obvious phase response followed the injection of CRP antigens and secondary

CRP antibodies, intended to increase signal amplification. The polymer rib

waveguide sensor, on the other hand, exhibited a large phase rise, when primary

antibodies of CRP and mouse IgG were injected. After BSA blocking, a clear

phase response was observed, following an injection of CRP antigens. After

flushing with PBS, 0.46 rad of additional phase increase remained. Similar to the

slab waveguide sensor, no signal rise was detected in the negative control. Finally,

measurements on the composite waveguide sensor began in the same manner as

those on the polymer rib waveguide sensor. However, phase responses obtained

from biomolecular adsorption were much greater. The remaining phase change

after PBS wash was about 9.1 rad and 2.3 rad for CRP analytes with a

concentration of 2 µg/ml and 0.1 µg/ml. Following the detection of 0.1 µg/ml

CRP analyte, a phase change of over 8 rad resulted from the injection of

secondary antibodies. The amplified phase response is due to multiple binding of

secondary antibodies to the primary antibody-antigen complexes. In the negative

control, only a small phase change was detected, representing minor nonspecific

binding. This nonspecific binding could only be detected by the composite

waveguide sensor, due to its higher sensitivity. Fluorescence microscopy images,

presented in Fig. 24, confirm that secondary antibodies attached well onto the

CRP antibody-activated surface, but not onto the sensing surface of the negative

control.

Taking into account the theoretical sensitivity of surface sensing, the actual

thickness growth of the adlayer can be calculated with the following equation:

ad

effeffad t

nnt

∂∂

Δ=Δ , (18)

The obtained layer thickness does not indicate exact molecular height, but an

average thickness taking account of molecular packing density. The

56

corresponding surface mass density can then be obtained with De Feijter’s

formula (De Feijter et al. 1978)

dcdn

nntM cad

ad

−Δ= , (19)

where dn/dc is the refractive index increment of the protein solution. A common

value of 0.18 cm3/g for protein adsorption was used for dn/dc (Vörös et al. 2004).

Estimated surface mass densities for the sensors are listed in Table 2.

Table 2. Phase responses per unit length (1 cm) and estimated surface mass densities

for the three waveguide sensors in the TM-coupling mode.

Slab Rib Ta2O5

2 µg/ml 2 µg/ml 2 µg/ml 0.1 µg/ml

δ [rad] 0.48 0.55 9.1 2.6

∂neff/∂tad [nm-1] 1.16e-5 1.4e-5 2.1e-4

tad [nm] 0.43 0.40 0.44 0.12

M [pg/mm2] 270 260 280 80

From Table 2, it can be seen that, for the 2 µg/ml CRP analyte, the surface mass

densities measured with the three developed waveguide sensors are in good

agreement with each other. The obtained measurement results are less than the

monolayer surface mass coverage of 1.48 ng/mm2 (Lin et al. 2006), partially due

to the random orientation of CRP antibodies. This means that only some of the

CRP antibodies have their binding sites facing out from the waveguide and are

therefore available to bind CRP antigens. To enhance the antibody-antigen

binding efficiency, surface immobilization protocols could be utilized to control

the orientation of primary antibodies attached on the sensing surface. With an

immobilization protocol, primary antibodies tend to be covalently bonded to the

sensing surface, which would greatly reduce the problem of detaching during the

measurement process.

57

5 Discussion

This section analyzes the influence of disturbance and noise on the measurement

system to determine the different waveguide sensors’ detection limit for bulk RI

sensing and sensing of biomolecule adsorption. Also the benefits and limitations

of the different sensing configurations are compared.

5.1 Measurement error analysis

Noise plays an important role in signal acquisition and processing. As the SNR

directly restricts the detection limit of the fabricated waveguide sensors, an

analysis of possible noise sources is neccessary be conducted. Noise that may

have an effect on the measurement results fall into the following categories:

mechanical, optical, environmental and electrical.

5.1.1 Mechanical noise

Two types of mechanical noise sources influence measurement results. Firstly,

there is vibration noise, which may originate from the floor, air fluctuations or

acoustic noise. It contributes to the measurement results by exciting differential

movement of the optical components (Selvin & Ha 2007). In this work,

differential movement refers to the relative movement between the waveguide

sensor and the objective lens when capturing fringe patterns (Fig. 18). Taking the

polymeric-inorganic composite waveguide sensor as an example, the period of the

generated fringe pattern is ~ 200 pixels. Using Equation (20), the actual spatial

period p can be estimated:

d

DP

λ= , (20)

where D = 2.5 mm is the distance from the sensor output facet to the collecting

objective lens, and d = 50 µm is the separation between the sensing and

referencing waveguide at the output facet. An actual spatial period of 32 µm is

found, corresponding to 5 µm for 1 radian in phase. Assuming that the generated

noise is entirely caused by vibration, the measured noise level of 0.003 radian

equals a relative movement of 15 nm between the waveguide sensor and the

objective lens.

58

The other source of mechanical noise results from the syringe pump

continuously circulating the liquid sample in the measurement system. Pressure

generated by the moving syringe causes a mechanical movement of the flow cell

attached above the sensing element, especially with a high flow rate. Pumping

pulses may also produce another kind of noise. If micro air bubbles are trapped in

the sensing window, pumping pulses shake the bubbles vibrating around the

interaction path, causing a change of the refractive index in the sensing path.

Since the composite waveguide has a higher sensitivity, micro air bubbles have a

larger influence on it. Shown in the following figure is the phase response of the

composite waveguide sensor with the pump turned on and off.

0 2 4 6 8 10

0.00

0.05

0.10

0.15

Pump off

Pha

se [

rad]

Time [min]

Pump on

Fig. 25. Phase response of the composite waveguide sensor with deionized water in

the flow circulating system with the pump turned on and off.

Mechanical noise caused by the syringe pump can be alleviated by attaching

microfluidic flow systems above the sensing surface. A microfluidic channel

driven by capillary force serves to smooth the liquid flow.

5.1.2 Quantization errors of the image acquisition device

Many different types of noise originate from the electronic image acquisition

device used in the measurement setup of this work. For example, dark current,

shot noise and electronic interference are common types of noise in digital

cameras. A detailed explanation of electronics noise can be found in textbooks by

Lee (Lee 2005). Focus of the noise analysis generated from the electronic

acquisition device is put on quantization errors, which arise when analogue

images of interference fringe patterns are digitized. Digitization may lead to

deterioration in image quality, particularly as the camera is limited both in terms

59

of spatial resolution and bit depth. Spatial resolution in the x and y direction

determines the number of data points sampled on the generated fringe patterns,

while bit depth determines the number of available grey-scale levels (Skydan et al.

2003).

To determine phase errors caused by digital quantization, bit depth and spatial

resolution were considered separately by keeping one factor constant, while

varying the other. Thus, the phase error for a known phase shift was obtained by

comparing the FFT-extracted phase shifts of the digital signal and the original

analogue signal. For simplification, a one-dimensional (1D) ideal sinusoidal

signal was used instead of a 2D image in both cases. Shown in Fig. 26 a) are the

average phase errors simulated for signals with a bit depth of 2 to 15. For all

utilized 1D sine signals, the number of pixels per period was kept at 125, similar

to experiments where fringe pattern images are set to less than 8 periods for a

length of 1024 pixels. Given a phase shift, the phase error was obtained from the

phase difference between the FFT-extracted phase shift of the analogue and

digital signals. For each bit depth in Fig. 26 a) and number of pixels in Fig. 26 b),

the average phase error was the average of 500 phase errors obtained from 500

randomly given phase shifts ranging from 10-4 to 10 rad.

0 2 4 6 8 10 12 14 16

0.0

1.0x10-3

2.0x10-3

3.0x10-3

4.0x10-3

5.0x10-3

6.0x10-3

Obtained phase error Exponential fit of

Pha

se e

rror

s [r

ad]

Bit depth

a

0 250 500 750 1000 1250

0.0

5.0x10-4

1.0x10-3

1.5x10-3

2.0x10-3

2.5x10-3

3.0x10-3

3.5x10-3

Pha

se e

rror

[ra

d]

No. of pixels in one fringe

Bit depth=6 Bit depth=8 Bit depth=10

b

Fig. 26. Phase errors plotted against a) bit depth with 125 pixels in a fringe, b) number

of pixels in one fringe with a bit depth of 6, 8 and 10.

As shown in Fig. 26 a), the phase error decays exponentially as the bit depth

increases. When the bit depth equals eight, which is the bit depth of the CMOS

camera used here, the phase error produced by digital quantization is in the range

of 10-4 rad. Fig. 26 b) presents the average phase error simulated with respect to

the number of pixels in a fringe for a bit depth of 6, 8 and 10, respectively.

60

Moreover, the phase error also decays exponentially as the number of pixels

increases. As the figure indicates, fewer pixels are required as bit depth grows to

achieve a low phase error. When the bit depth equals 8, the phase error is within

the range of 10-4 rad, which is less than the noise level of 10-3 rad present in the

measurements. Therefore, it is safe to conclude that, as far the measured noise is

concerned; quantization errors are not the main contributor.

5.1.3 Optical noises

Optical noise comes mostly from stray light in the sensor chip, generated by mode

mismatching between the coupling mode and the diameter of the waveguide

cross-section. In this research, the fabricated single-mode slab waveguide had a

thickness of ~1 µm, while the polymer rib waveguide had a cross-section

dimension of 2 µm × 0.9 µm and the composite waveguide 2 µm × 1.6 µm. Since

the focused spot diameter of the 60× microscope objective lens was

approximately 3 µm, it was larger than the dimensions of the waveguide. As

shown in Fig. 27 the resultant stray light deteriorated the visibility of the

interference pattern, rounding the top of the amplitude spectrum peak of the 2D

FFT. As a result, a degree of ambiguity surrounded the selection of peak

frequency value.

500 510 520 530pixel

arb.

uni

t ar

b. u

nit

500 510 520 530

5

10

15

x 106

pixel

arb.

uni

tar

b. u

nit

a

b

2D-FFT

2D-FFT

Fig. 27. a) The 2D-FFT is derived from a fringe pattern image with good visibility. The

center column indicated with the arrows is also plotted in one dimension. b) 2D-FFT

and its center column plot for a fringe pattern image having bad visibility.

61

On the other hand, optimizing the alignment of the optical path causes the

incident beam to be reflected back into the He-Ne laser cavity, producing

instability, such as amplitude modulation and frequency shifting (Budzyn et al.

2008). In this work, back reflection was minimized by placing an isolator in front

of the laser. However, a small portion of the back-reflected light was still able to

enter the laser cavity, introducing noise into the baseline phase response.

5.1.4 Drift caused by temperature variations

Temperature variations tend to affect measurements by causing a drift of the

sensing response in a long measurement range. This drift is a combinational effect

of changes in the bulk RI of the analyte, the RI of the used polymers and changes

in waveguide dimensions. Assume the change in the refractive index of water

caused by temperature variations in the range of 20–30 °C is linear. This

relationship can be approximated with dn/dT = -1×10-4 /°C (Lide 2000). Based on

the change rate of RI to temperature dn/dT = -1×10-4 /°C found in Poly(methyl

methacrylate) (PMMA) (Tomiki et al. 2005), the refractive index of the

waveguide core (Ormocore) and cladding (Ormoclad) was assumed to change

within a range of -1×10-6 ~ -1×10-3/°C. For Ormocore, the coefficient of thermal

expansion (CTE) is 100~130 ppm K-1. Finally, the phase change subject to

temperature raise per Celsius degree was calculated for a polymer-based

waveguide interferometer sensor having waveguide thickness of 1 µm and

interaction length of 1 cm (Fig. 28 a). According to the calculation in Fig. 28 a),

the phase change caused by 1 °C temperature variation is in a range of -0.1 ~ 0.6

rad. When the change of the polymer RI equals to -1.32×10-4/°C, the drift of

phase caused by the temperature variation can be totally cancelled out. As shown

in Table 1.Fig. 24 a) the drift do exists in the fabricated polymer sensors. It

suggests that the fabricated sensor can not completely compensate the

temperature induced drift.

62

0E-4 2E-4 4E-4 6E-4 8E-4 10E-4-0.2

0.0

0.2

0.4

0.6

Pha

se c

hang

e [r

ad]

Δn of the polymer core and cladding/°C

0 10 20 30 40 50 60

-0.06

-0.04

-0.02

0.00

0.02

0.04

Pha

se [

rad]

Time [min]

Fig. 28. a) Simulated phase response under the influence of temperature variation per

Celsius degree in a three-layer waveguide system; b) Long-term stability

measurement in the composite waveguide interferometer sensor.

The long-term stability of a composite waveguide sensor was also monitored to

characterize the behaviour of the sensor under the influence of environmental

temperature variation (Fig. 28 b)). In this measurement, the sensor was exposed to

an air flow to substantially increase the influence of air fluctuations and

temperature. At the beginning and end of the measurement, a drastic phase change

took place, owing to a temperature variation caused by the absence or presence of

an operator. In the middle of the measurement, with no large temperature

variations, the baseline remained flat. Moreover, baseline stability is improved by

a continuous flow of the liquid sample, since the sensing surface is covered with

liquid under the flow cell, making it more resistant to short-term temperature

variations caused by air turbulence.

5.2 A comparison between the three different YI configurations

Detection limits (LODs) of the three fabricated IO YI sensors were estimated

based on their sensorgram baseline noise level. The obtained results were then

compared with a commercialized dual-slab waveguide sensor made from the

inorganic dielectric SiON (Farfield Scientific Inc). For all four sensors,

measurements were performed in the same setup, in order to keep the evaluation

conditions as identical as possible. The LOD is defined by the 3σ criterion

(Schmitt et al. 2007), where the minimum detectable signal is three times larger

than the noise level. The estimated LODs for bulk refractive index sensing and

surface sensing of biomolecule adsorption are listed in Table 3.

63

Table 3. LODs for four different waveguide sensors in TM coupling mode.

Paper I

SiON

Paper III

Slab

Paper IV Rib Paper V

Ta2O5

∂neff/∂nc (measured) 0.0092 0.0071 0.0086 0.1061

∂neff/∂tad [nm-1] (thoretical) 2.1e-5 1.2e-5 1.4e-5 2.1e-4

Noise level [rad] <0.002 0.006 0.002 0.003

LOD in RIU 4e-6 2e-5 1e-5 9e-7

LOD in surface mass density [fg/mm2] 1600 7500 4500 270

Of the studied sensors, the polymer-based dual-slab waveguide YI sensor has the

lowest SNR and worst long-term stability. As shown in Fig. 24 a), a drift can be

seen after the injection of CRP antigens. In Fig. 20 a) and Fig. 24 a) a higher

noise level can be observed with the slab waveguide sensor compared to the rest

of sensors. Although sensitivity may be increased by the deposition of high-index

coating, also the noise level may be amplified at the same time. Because the

reference waveguide underneath cannot be deposited with a corresponding high-

index layer, the result in an unbalanced noise compensation. As expected, the

lowest noise level is found in the inorganic sensor, due to better thermo-

mechanical and thermal-optical stability. The composite waveguide YI sensor has

the best LOD, thanks to its high sensitivity, which exceeds that of the other

configurations by a factor of 10. The developed composite waveguide sensor can

achieve a sensitivity and LOD level comparable to some of the inorganic sensors

(Densmore et al. 2008, Ymeti et al. 2002, Sun et al. 2011, Zinoviev et al. 2011,

Tu et al. 2012).

Besides sensitivity and LOD, other aspects of the fabricated sensors can also

be compared. One of the important characteristics of IO sensors is the ability for

multiparameter analysis on a single sensor chip. For a rib waveguide-based

structure, a cascade configuration can be constructed in the sensor design, as

demonstrated by Ymeti et al. For a slab waveguide structure, however,

multiparameter analysis is more difficult, since no definite borders exist in the

vertical direction. But slab waveguide sensors have an advantage in terms of

incident light coupling, because slab configuration allows easier horizontal

aiming and a higher tolerance for misalignment. For applications in aqueous

environments, polymer-based sensors must be immersed in water before

measurements to achieve a flat baseline response. Inorganic and composite

64

waveguide sensors, on the other hand, do not require such advance preparations.

In terms of fabrication, inorganic slab waveguides need consecutive depositions

of SiON layers by a chemical vapor deposition (CVD) system, while polymer-

based waveguide sensors only require spin coating and imprinting. For the

composite waveguide sensor, only one inorganic high-index layer needs to be

deposited by sputtering at near-room temperature. Therefore, the fabrication of

polymer-based waveguide sensors, especially rib waveguide sensors, can be less

expensive and less time consuming.

Having evaluated all the aspects mentioned above, we may conclude that, of

all the developed sensors, the composite rib waveguide configuration has the most

favourable performance characteristics and a relatively simple fabrication process.

In addition, inorganic Ta2O5 coating can effectively block water penetration into

the polymer waveguide. This stabilizes the sensing response immediately after the

application of an aqueous solution, making the sensor ready for measurements.

5.3 Future work

There are several ways of improving the performance of the developed sensors.

For the detection of biomolecular binding events, the receptors can be

immobilized firmly on the sensing surface through, for example, covalent

bonding to minimize the risk of receptor detachment. To improve the system’s

SNR, integrated microfluidic channels could be used to alleviate mechanical

noise caused by the continuous syringe pump. Moreover, multiparameter sensing

could be realized by a higher level integration of rib waveguide YI sensors on a

single sensor chip. It would extend the range of references to the sensing system

and further improve noise compensation.

The next stage of the current work involves the adaption of the fabrication

process of rib waveguide YI sensors to the roll-to-roll printing technique that is

compatible for high volume production. In the future, the development of a fully

integrated lab-on-a-chip biosensor comprising the developed waveguide

interferometer together with a compact light source, a polarizer, and a detector

would be an eventual solution for a compact biosensor.

65

6 Summary

The main ambition of this thesis was to develop low-cost disposable IO

biosensors compatible for industrial mass-production. Polymers have been used

as fundamental materials in sensor development due to their good optical

properties, versatile processability and relatively low cost. This thesis used a

series of UV-curable organic and inorganic hybrid materials known as Ormocer

for sensor fabrication, largely owing to their comparatively good chemical and

thermal stability.

Three types of IO YI sensors with different waveguide configurations were

developed here: polymer dual-slab waveguide sensors, polymer inverted-rib

waveguide sensors and composite waveguide sensors based on the inverted-rib

waveguide structure with Ta2O5 coating. These three configurations were selected

to enhance the capabilities of the sensor prototype by enabling a new compact

structure, a simplified fabrication process and enhanced sensitivity.

Theoretical sensitivities were analyzed for homogenous refractive index

sensing and surface biomolecule adsorption. The results confirmed that an 80 nm

Ta2O5 layer deposited on a polymer waveguide improved its sensitivity in the

TM-coupling mode by over 40 times. Also propagation modes in each waveguide

configuration were simulated and single mode sensing was assured for all three

configurations.

Considering the configuration of each type of the sensors separately, we find

that the dual-slab waveguide IO YI sensors can be fabricated by multiple spin-

coating processes. Moreover, compared to rib waveguides, it allows easier

coupling of incident light with a greater tolerance for misalignment. For the

inverted-rib waveguide sensors, a very simply fabrication process, compatible to

roll-to-roll printing technique, was developed. Utilizing UV-imprinting, the

inverted-rib waveguide sensor can be fabricated in just two main steps:

patterning/imprinting the undercladding of the waveguide and spin-coating its

core. Being totally RIE-free, this method simplifies the fabrication process and

eliminates the risk of increased waveguide surface roughness, which could cause

additional scattering loss. The composite waveguide sensors, based on the

inverted-rib waveguide structure with an additional high RI Ta2O5 coating, enable

a simple, low-cost fabrication process with a minimum number of inorganic layer

depositions. Furthermore, the high-index inorganic layer serves not only to

improve sensitivity but, in aqueous environments, it also prevents the absorption

of water molecules in the polymer waveguide.

66

The developed sensors were evaluated in glucose DI-water for homogenous

bulk RI sensing and in an immunoassay for surface sensing of protein adsorption.

All the tested sensors fulfilled these functions in both sensing mechanisms but,

with a LOD of 9 ×10-7 RIU and 270 fg/mm2 for surface mass density, the best

results were achieved by composite waveguide sensors. A simple fabrication

procedure and high sensitivity place Ta2O5-deposited polymeric-inorganic

composite waveguide YI sensors high on the list of disposable low-cost

biosensors suitable for mass production.

67

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Original Papers

I Wang M, Uusitalo S, Määttälä M, Myllylä R & Känsäkoski M (2008) Integrated dual-slab waveguide interferometer for glucose concentration detection in the physiological range. Proc SPIE 7003: 70031N-1-10. DOI: 10.1117/12.780184.

II Wang M, Uusitalo S, Liedert C, Hiltunen J, Hakalahti L & Myllylä R (2012) Polymeric dual-slab waveguide interferometer for biochemical sensing applications. Appl Opt 51: 1886–1893.

III Wang M, Hiltunen J, Uusitalo S, Puustinen J, Lappalainen J, Karioja P & Myllylä R (2011) Fabrication of optical inverted-rib waveguides using UV-imprinting. Microeletron Eng 88(2): 175–178.

IV Wang M, Hiltunen J, Liedert C, Hakalahti L & Myllylä R (2012) An integrated Young interferometer based on UV-imprinted polymer waveguides for label-free biosensing applications. J Europ Opt Soc Rap. Public. 7: 12019.

V Wang M, Hiltunen J, Liedert C, Pearce S, Charlton M, Hakalahti L, Karioja P & Myllylä R (2012) Highly sensitive biosensor based on UV-imprinted layered polymeric–inorganic composite waveguides. Opt Express 20: 20309–20317.

Reprinted with permission from SPIE (I) OSA (II, V) Elsevier (III) and European

Optical Society (IV).

Original publications are not included in the electronic version of the dissertation.


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