Development of materials, surfaces and manufacturing ...437739/FULLTEXT02.pdf · achieve these goals, novel materials, surfaces and manufacturing methods in ... manufacturing methods

Development of materials, surfaces and

manufacturing methods for micro�uidic applications

C A R L F R E D R I K C A R L B O R G

Doctoral Thesis in Microsystem TechnologyStockholm, Sweden 2011

Development of materials, surfaces andmanufacturing methods for microfluidic

applications

CARL FREDRIK CARLBORG

Doctoral ThesisStockholm 2011

The front cover photo shows a prototype of a miniaturised point-of-care test. A car-tridge with integrated microfluidics and label-free slot waveguide ring resonator sensors.Superimposed on the image is an exploded view of the optical chip and the microfluidicdistribution layers. Top right: An electron micrograph showing a small section of asuper-lubricating microchannel. Water, supported by surface tension, is flown in the30 µm wide channel between the two rows of 2 µm thick pillars. Bottom right: Abiosensor microarray for detecting lactose intolerance, encapsulated with a microfluidicsticker. Each of the spots are 250 µm in diameter and consist of !-lactoglobulin (milkprotein from cow).

TRITA-EE 2011:058ISSN 1653-4146ISBN 978-91-7501-086-1 Microsystem Technology, KTH

Akademisk avhandling som med tillstånd av Kungl Tekniska högskolan framläggestill o!entlig granskning för avläggande av teknologie doktorsexamen den 23:e sep-tember 2011 klockan 10:00 i sal F3, Lindstedtsvägen 26, Stockholm.

Thesis for the degree of Doctor of Philosophy at the Royal Institute of Techno-logy (KTH), Stockholm, Sweden

© Carl Fredrik Carlborg, September 2011

E-mail: [email protected]

Tryck: Universitetsservice US AB, Stockholm 2011

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Abstract

This thesis presents technological advancements in microfluidics. The over-all goals of the work are to develop new miniaturized tests for point-of-carediagnostics and robust super-lubricating surfaces for friction reduction. Toachieve these goals, novel materials, surfaces and manufacturing methods inmicrofluidics have been developed.

Point-of-care diagnostic tests are portable miniaturized instruments thatdownscale and automate medical tests previously performed in the central lab-oratories of hospitals. The instruments are used in the doctor’s o"ce, in theemergency room or at home as self-tests. By bringing the analysis closer to thepatient, the likelihood of an accurate diagnosis, or a quick therapy adjustment isincreased. Already today, there are point-of-care tests available on the market,for example blood glucose tests, rapid streptococcus tests and pregnancy tests.However, for more advanced diagnostic tests, such as DNA-tests or antibodyanalysis, integration of microfluidic functions for mass transport and samplepreparation is required. The problem is that the polymer materials used inacademic development are not always suited for prototyping microfluidic com-ponents for sensitive biosensors. Despite the enormous work that has gone intothe field, very few technical solutions have been implemented commercially.

The first part of the work deals with the development of prototype point-of-care tests. The research has focused on two major areas: developing newmanufacturing methods to leverage the performance of existing materials anddeveloping a novel polymer material platform, adapted for the extreme de-mands on surfaces and materials in miniaturized laboratories. The novel man-ufacturing methods allow complex 3D channel networks and the integrationof materials with di!erent surface properties. The novel material platform isbased on a novel o!-stoichiometry formulation of thiol-enes (OSTE) and hasvery attractive material and manufacturing properties from a lab-on-chip per-spective, such as, chemically stable surfaces, low absorption of small molecules,facile and inexpensive manufacturing process and a biocompatible bondingmethod. As the OSTE-platform can mirror many of the properties of commer-cially used polymers, while at the same time having an inexpensive and facilemanufacturing method, it has potential to bridge the gap between research andcommercial production.

Friction in liquid flows is a critical limiting factor in microfluidics, wherefriction is the dominant force, but also in marine applications where frictionallosses are responsible for a large part of the total energy consumption of seavessels. Microstructured surfaces can drastically reduce the frictional losses bytrapping a layer of air bubbles on the surface that can act as an air bearing forthe liquid flow. The problem is that these trapped air bubbles collapse at theliquid pressures encountered in practical applications.

The last part of the thesis is devoted to the development of novel low fluid-friction surfaces with increased robustness but also with active control of thesurface friction. The results show that the novel surfaces can resist up to threetimes higher liquid pressure than previous designs, while keeping the samefriction reducing capacity. The novel designs represent the first step towardspractical implementation of micro-structured surfaces for friction reduction.

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Sammanfattning

Den här avhandlingen presenterar tekniska framsteg inom forskningsfältetmikrofluidik. De övergripande målen med arbetet är att utveckla nya minia-tyriserade analysinstrument för patientnära medicinsk diagnostik och stabilamikrostrukturerade ytor för friktionsreduktion i vätskeflöden över fasta ytor.För att uppnå dessa mål, presenterar avhandlingen nya material, ytor och till-verkningsmetoder för mikrofluidik.

I analysinstrument för patientnära medicinsk diagnostik förminskas ochautomatiseras medicinska tester, som tidigare gjorts på sjukhusens centralla-boratorier, till små, portabla enheter. Användningsområdet för dessa analy-sinistrument återfinns på husläkarmottagningar, i akutrum på sjukhus eller ihemmet som självtester. Genom att förflytta analysen närmare patienten, kandiagnoser och dosjusteringar av medicineringar göras snabbare och mer exak-ta. Redan idag finns enklare produkter tillgängliga på marknaden, exempelvisblodsockertester, tester för vissa bakterieinfektioner och graviditetstester. Föratt kunna utveckla nästa generations instrument och för att kunna diagnosti-sera mer avancerade sjukdomstillstånd, är integration av avancerade mikroflu-idiska komponenter, som sköter masstransport och provpreparering en förut-sättning. Trots att mycket stora forskningsinsatser har lagts ned på området,har mycket få tekniska lösningar implementerats kommersiellt. Problemet idagär att de material och metoder, som används inom akademisk forskning, inteär anpassade för ändamålet och behöver förbättras eller bytas ut för att kunnatillverka relevanta kommersiella prototyper. Den första delen av arbetet ägnasåt att utveckla nya tillverkningsmetoder för att adressera ovanstående problemsamt att utveckla av en ny polymerbaserad materialplattform, speciellt avpas-sad för de extrema krav som ställs på ytor och material i dessa miniatyrisera-de laboratoriemiljöer. De nya tillverkningsmetoderna möjliggör komplexa 3Dstrukturer och integration av material med olika ytegenskaper. Den nya mate-rialplattformen, baserad på nya icke-stökiometriska tiolen-formuleringar (OS-TE), har attraktiva material och processegenskaper för prototyptillverkning avminiatyriserade medicinska test, t.ex. kemiskt stabila ytor, låg absorption avbiologiskt material, enkel och billig tillverkningsmetod samt en biokompatibelsammanfogningsprocess. Eftersom OSTE-platformen kan spegla egenskapernapå kommersiellt använda polymerer och ha en tillverkningsprocess som är till-räckligt enkel och billig för de flesta laboratorier, har den även potential attöverbrygga gapet mellan forskning och kommersiell produktion.

Friktionsreduktion i vätskeföden över ytor är ett viktigt område inom mik-rofludik, där friktion är den dominerande kraften, men även för marina tillämp-ningar, där friktionsförluster svarar för en stor del av den totala energiåtgången,till exempel för oljetankers. Mikrostrukturerade ytor kan drastiskt minska frik-tionsförluster genom att fånga ett lager av gasbubblor på ytan som fungerarsom smörjmedel för vätsketransporten. Problemet är att dessa ytor kollapsarvid de vätsketryck som på trä!as i praktiska tillämpningar. Den andra delenav arbetet ägnas åt att utveckla nya lågfriktionsytor för att öka stabilitetenmen även för att aktivt manipulera friktionen på ytan. De nya ytorna kla-rar upp till tre gånger högre vätsketryck med bibehållen friktionsreduktion änvad som kunnat visas tidigare och representerar ett första steg mot praktiskatillämpningar.

v

Till Pappa, som hade sett fram emot den här dagen

Contents

Contents vi

List of publications ix

Objectives and Overview xiiiStructure . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . xiii

1 Introduction to lab-on-chip devices 11.1 Towards improved healthcare: point-of-care tests . . . . . . . . . . . 2

1.1.1 What is a point-of-care test? . . . . . . . . . . . . . . . . . . 21.1.2 Market and opportunities . . . . . . . . . . . . . . . . . . . . 3

1.2 Advantages of microfluidics for medical testing . . . . . . . . . . . . 41.3 Lab-on-chip development . . . . . . . . . . . . . . . . . . . . . . . . 4

1.3.1 Background on polymer technology . . . . . . . . . . . . . . . 41.3.2 Polymer materials . . . . . . . . . . . . . . . . . . . . . . . . 61.3.3 Rapid prototyping . . . . . . . . . . . . . . . . . . . . . . . . 81.3.4 Polymer microstructuring methods . . . . . . . . . . . . . . . 81.3.5 Back-end processes . . . . . . . . . . . . . . . . . . . . . . . . 10

1.4 The ideal prototyping system for labs-on-chip . . . . . . . . . . . . . 131.5 Summary and conclusions . . . . . . . . . . . . . . . . . . . . . . . . 15

2 Novel manufacturing methods for labs-on-chip 172.1 Dual surface-energy adhesive for the integration and packaging of an

optical label-free sensor . . . . . . . . . . . . . . . . . . . . . . . . . 172.1.1 Dual surface-energy adhesives . . . . . . . . . . . . . . . . . . 172.1.2 Background on optical biosensing . . . . . . . . . . . . . . . . 192.1.3 Background on mass transport . . . . . . . . . . . . . . . . . 212.1.4 Microfluidics design and manufacturing . . . . . . . . . . . . 232.1.5 Integration and bonding of the sensor cartridge . . . . . . . . 252.1.6 Biosensing results . . . . . . . . . . . . . . . . . . . . . . . . 272.1.7 Discussion and outlook . . . . . . . . . . . . . . . . . . . . . 27

2.2 High yield process of vertical interconnects in PDMS for batch man-ufacturing 3D microfluidics devices . . . . . . . . . . . . . . . . . . . 282.2.1 PDMS polymerisation process . . . . . . . . . . . . . . . . . . 30

vi

CONTENTS vii

2.2.2 Residual-free interconnects by local inhibition . . . . . . . . . 302.2.3 Direct bonding using the inhibited surface . . . . . . . . . . . 302.2.4 3D microfluidic networks for labs-on-chip . . . . . . . . . . . 302.2.5 Discussion and outlook . . . . . . . . . . . . . . . . . . . . . 31

2.3 Summary and outlook . . . . . . . . . . . . . . . . . . . . . . . . . . 33

3 OSTE: a novel material toolbox for labs-on-chip 353.1 Thiol-ene "click" chemistry . . . . . . . . . . . . . . . . . . . . . . . . 353.2 OSTE: O!-stochiometry thiol-enes . . . . . . . . . . . . . . . . . . . 36

3.2.1 Residual activity through o!-stoichiometry . . . . . . . . . . 363.2.2 Tuneable mechanical properties . . . . . . . . . . . . . . . . . 393.2.3 Direct patternable surface modification . . . . . . . . . . . . 403.2.4 Biocompatible low-temperature bonding . . . . . . . . . . . . 413.2.5 Low absorption of molecules . . . . . . . . . . . . . . . . . . . 423.2.6 Solvent resistant channels . . . . . . . . . . . . . . . . . . . . 423.2.7 A rapid and scalable manufacturing process . . . . . . . . . . 43

3.3 Facile integration of microfluidics with microarrays: the Biosticker . 433.3.1 Introduction to microarrays . . . . . . . . . . . . . . . . . . . 433.3.2 Mass-transport limitation in microarrays . . . . . . . . . . . . 443.3.3 The Biosticker: a micropatterned OSTE-sticker for microarrays 453.3.4 Preliminary results . . . . . . . . . . . . . . . . . . . . . . . . 46

3.4 Summary and outlook . . . . . . . . . . . . . . . . . . . . . . . . . . 473.5 Future work . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 48

4 Introduction to low fluid-friction surfaces 494.1 Motivation . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 494.2 Surface friction in liquid flows . . . . . . . . . . . . . . . . . . . . . . 504.3 Superhydrophobic surfaces . . . . . . . . . . . . . . . . . . . . . . . . 51

4.3.1 Mechanism of operation . . . . . . . . . . . . . . . . . . . . . 514.3.2 Stability limitations . . . . . . . . . . . . . . . . . . . . . . . 53

5 Novel robust super-lubricating surfaces 555.1 A model for friction reduction in a microchannel . . . . . . . . . . . 555.2 Fractal surfaces: temporary life support . . . . . . . . . . . . . . . . 575.3 Active switching: wet or dry . . . . . . . . . . . . . . . . . . . . . . . 575.4 Regulating the air pocket pressure to avoid collapse . . . . . . . . . 58

5.4.1 Active regulation . . . . . . . . . . . . . . . . . . . . . . . . . 585.4.2 Self-regulating air pockets . . . . . . . . . . . . . . . . . . . . 595.4.3 Performance . . . . . . . . . . . . . . . . . . . . . . . . . . . 60

5.5 Summary and outlook . . . . . . . . . . . . . . . . . . . . . . . . . . 61

6 Conclusions 63

Appendix: Tables 65

viii CONTENTS

Summary of appended papers 69

Acknowledgement 73

References 75

Paper reprints 87

List of publications

The thesis is based on the following papers in international peer reviewedjournals:

1. "A packaged optical slot-waveguide ring resonator sensor array for multiplexlabel free assays in labs-on-chip",C. F. Carlborg, K. B. Gylfasson, A. Ka#mierczak, F. Dortu, M. J. BañulsPolo, A. Maquieira Catala, G. M. Kresbach, H. Sohlström, T. Moh, L. Vivien,J. Popplewell, G. Ronan, C. A. Barrios, G. Stemme, and W. van der Wijn-gaart. Lab on a Chip, vol. 10, no. 3, pp. 281-290, February 2010

2. "On-chip temperature compensation in an integrated slot-waveguide ring res-onator refractive index sensor",K. B. Gylfasson, C. F. Carlborg, A. Ka$mierrczak, F. Dortu, H. Sohlström,L. Vivien, C. A. Barrios, W. van der Wijngaart, and G. Stemme Optics Ex-press, vol. 18, no. 4, pp. 3226-3237, February 2010

3. "Large scale integrated 3D microfluidic networks through high yield fabrica-tion of vertical vias in PDMS",C. F. Carlborg, K. T. Haraldsson, M. Cornaglia, G. Stemme, and W. vander Wijngaart IEEE Journal of Microelectromechanical Systems, vol. 19, no.5, pp. 1050-1057, October 2010

4. "Beyond PDMS: o!-stoichiometry thiol-ene (OSTE) based soft lithographyfor rapid prototyping of microfluidic devices"C. F. Carlborg, K. T. Haraldsson, K. Öberg, M. Malkoch, and W. van derWijngaart. Lab on a Chip, vol. 11, no. 18, pp, 3136-3147, July 2011

5. "Click Wafer Bonding for Microfluidic Devices"F. Saharil, C. F. Carlborg, K. T. Haraldsson, and W. van der Wijngaart.Lab on a Chip, submitted September 2011

6. "Sustained superhydrophobic friction reduction at high pressures and largeflows"C. F. Carlborg, and W. van der Wijngaart. Langmuir, vol. 27, no. 1, pp.487-493, December 2010

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x LIST OF PUBLICATIONS

The thesis is also based on the following international peer reviewedconference proceedings:

7. "Continuos flow switching by pneumatic actuation of the air lubrication layeron superhydrophobic microchannel wallsC. F. Carlborg, M. Do-Quang, G. Stemme, G. Amberg, and W. van derWijngaart.in IEEE Proceedings of the 21th Int. Conf. on Micro Electro MechanicalSystems, Tuscon, USA, January 2008, pp. 599-602

8. "Biosticker: Patterned microfluidic stickers for rapid integration with microar-rays"C. F. Carlborg, M. Cretich, K. T. Haraldsson, L. Sola, M. Bagnati, M. Chiariand W. van der Wijngaart.in Proceedings of the 11th Int. Conf. on Miniaturized Systems for Chemistryand Life Sciences (µTAS), Seattle, USA, October 2011, accepted

The contribution of Carl Fredrik Carlborg to the di!erent publications:

1. part of design, all packaging and microfluidic fabrication, major part of ex-periments and major part of writing

2. part of design, all packaging and microfluidic fabrication, part of experiments,and writing

3. major part of design, fabrication, all experiments and major part of writing

4. major part of design, all of fabrication, major part of experiments and writing

5. major part of design, part of fabrication, experiments and writing

6. major part of design, all fabrication, experiments and writing

7. major part of design, all fabrication, major part of experiments and writing

8. all design, fabrication, major part of experiments and all writing

xi

The work presented in the thesis has also been presented at the followinginternational peer reviewed conferences:

1. "Reliable batch manufacturing of miniaturized vertical vias in soft polymerreplica molding"C. F. Carlborg, K. T. Haraldsson, G. Stemme, and W. van der Wijngaart.in Proceedings of the 11th Int. Conf. on Miniaturized Systems for Chemistryand Life Sciences (µTAS), Paris, France, October 2007, pp. 257-259

2. "Microchannels with Substantial Friction Reduction at Large Pressure andLarge Flow"C. F. Carlborg, G. Stemme, and W. van der Wijngaart.in Proceedings of the 22rd Int. Conf. on Micro Electro Mechanical Systems,Sorrento, Italy, January 2009, pp. 39-42, oral presentation

3. "A packaged optical slot-waveguide ring resonator sensor array for multiplexlabel free assays in labs-on-chip",K. B. Gylfasson, C. F. Carlborg, A. Ka$mierczak, F. Dortu, M. J. BañulsPolo, A. Maquieira Catala, G. M. Kresbach, H. Sohlström, T. Moh, L. Vivien,J. Popplewell, G. Ronan, C. A. Barrios, G. Stemme, and W. van der Wijn-gaart.in Proceedings of the 13th Int. Conf. on Miniaturized Systems for Chem-istry and Life Sciences (µTAS), Jeju Island, South Korea, November 2009,pp. 2004-2006, oral presentation

4. "Large scale integrated 3D microfluidic networks through high yield fabrica-tion of vertical vias in PDMS",C. F. Carlborg, K. T. Haraldsson, M. Cornaglia, G. Stemme, and W. vander Wijngaartin IEEE Proceedings of the 23rd Int. Conf. on Micro Electro Mechanical Sys-tems, Hong Kong, P.R. China, January 2010, pp. 240-243, oral presentation

5. "Beyond PDMS: o!-stoichiometry thiol-ene (OSTE) based soft lithographyfor rapid prototyping of microfluidic devices"C. F. Carlborg, K. T. Haraldsson, K. Öberg, M. Malkoch, and W. van derWijngaart.in Proceedings of the 14th Int. Conf. on Miniaturized Systems for Chemistryand Life Sciences (µTAS), Groningen, Netherlands, October 2010, pp. 70-72,oral presentation

6. "Low temperature "Click" wafer bonding of o!-stoichiometry thiol-ene (OSTE)polymers to silicon"C. F. Carlborg, F. Saharil, K. T. Haraldsson, and W. van der Wijngaart.in Proceedings of the 15th Int. Conf. on Miniaturized Systems for Chemistryand Life Sciences (µTAS), Seattle, USA, October 2011, accepted

xii LIST OF PUBLICATIONS

Other international peer reviewed journal papers, by Carl Fredrik Carl-borg, not included in the thesis:

1. "Thermal boundary resistance between single-walled carbon nanotubes andsurrounding matrices",C. F. Carlborg, J. Shiomi, and S Maruyama.Physical Review B, vol. 78, no. 20, pp. 205406, 2008

2. "Poly(vinyl alcohol) as a temporary carrier for fabrication of fragile membranesand 3D fluidic networks"J.M. Karlsson, T. Haraldsson, C.F. Carlborg and W. van der Wijngaart.Journal of Micromechanics and Microengineering, manuscript

3. "Low-stress transfer bonding and assembly of multiple wafer-sized polymerlayers using floatation"J.M. Karlsson, T. Haraldsson, C.F. Carlborg and W. van der Wijngaart.Sensors and Actuators B, submitted August 2011

Objectives and Overview

This thesis presents technological advancements in microfluidics. The overall goalsof the work are to develop new miniaturized tests for point-of-care diagnostics androbust super-lubricating surfaces for friction reduction. To achieve these goals, novelmaterials, surfaces and manufacturing methods in microfluidics have been developed

StructureChapter 1 gives a general introduction to the development of laboratories-on-chipwith specific focus on polymer materials. Chapter 2 introduces novel manufacturingmethods for the development and integration of microfluidics with labs-on-chip, andtheir applications. This chapter also describes some of the design issues encounteredin the design of microfluidic components for medical sensors. Chapter 3 introducesa novel, highly versatile prototyping material aiming to bridge the gap betweenacademic proof of concept devices and commercial products. Chapter 4 gives anintroduction to super-lubricating surfaces, their uses and their limitations. Chapter5 presents two approaches to break the robustness limitation of current super-lubricating surfaces towards implementation in liquid flow conditions encounteredin realistic applications. The concluding chapter summarizes the work presented inthis thesis.

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Chapter 1

Introduction to lab-on-chip devices

Accidentally your leg is scratched while out jogging, there is some blood but itdoes not looks so serious so you forget about it. In the evening your leg hurts andyou notice that from the scar, dark lines are travelling up along your leg. Youget a little worried and call your local clinic. The nurse urges you to immediatelygo to the emergency room for a check-up. At the hospital they suspect bloodpoisoning and take blood samples to determine the type of the bacteria and thuswhat antibiotics they can use. Normally this test is done in the central hospitallaboratory and takes days. You do not have that much time, because by now youhave a fever and are shaking uncontrollably. The first thing the doctors do is togive you a large dose of a broad spectrum antibiotics and hope that the bacteriawill be responsive to at least one of them. Unfortunately, repeated use of broadspectrum antibiotics is one of the major causes of antibiotic resistance. Luckily,the hospital was recently equipped with the latest point-of-care system for rapidanalysis of antibiotic resistance bacteria, and with only a small amount of blood thedevice quickly determines exactly what antibiotics should be used. Twenty minutesafter you arrive, you are intravenously administered the correct antibiotics. Youare able to leave the hospital after two days of observation. In another scenario,you suspect your three years old son is allergic, but you do not know to what.He recently got red rashes over his whole body when he ate and you are worriedwhat other symptoms he will show. At the doctors o"ce, they suggest to do anallergy test by applying small droplets of di!erent allergy causing substances overthe whole back of your son and use a small needle to make it penetrate into theskin. You know that this is painful and could induce a serious allergic shock if theallergy is severe. Another clinic uses point-of-care devices for allergy testing thatneed only a droplet of blood from the index finger, to exactly determine what yourson is allergic to. These are two examples of scenarios for which labs-on-chip aredeveloped at the Microsystem Technology Laboratory at KTH. This introductiongives a brief overview of the intended usage of labs-on-chip and the most commonmaterials and manufacturing processes. Also, the reasons for the lack of success ofacademic research in the consumer market are discussed briefly.

1

2 CHAPTER 1. INTRODUCTION TO LAB-ON-CHIP DEVICES

1.1 Towards improved healthcare: point-of-care tests

1.1.1 What is a point-of-care test?

A point-of-care test (POCT) is defined as any medical test close to the patient: inthe doctor’s o"ce, by the hospital-bed, or at home. This closeness to the patientis believed to increase the likelihood of a quick and accurate diagnosis and therapyadjustment, or give the patients the convenience of avoiding hospital appointments.In contrast to centralised laboratory machines, which are usually based on large andbulky pipetting robots, a POCT should ideally be portable, easy to use and presenta first level of interpretation of the result to the user. To qualify as a successfulPOCT [1], it is usually required that the sensitivity and specificity are equivalent orbetter than centralised laboratory tests and that the cartridges are self-contained,disposable and low-cost. To realise these goals many POCT are built around atechnology called labs-on-chip; miniaturised automated laboratories that use theadvantages of microfluidics to compete with classical culture bottles, petri dishesand microtitre plates. The anatomy of a visionary lab-on-chip for point-of-care isdisplayed in Figure 1.1. It consists of a disposable part, with a loading port for theliquid sample, a sample preparation and metering unit (preconcentration, amplifi-cation, cell lysis), a microfluidic network (splitting, moving and mixing sample andreagents) and a sensor and signal transduction part with receptors for labelled orlabel-free detection. The disposable part is inserted into a reader, containing all theelectronic part (signal processing) and the user interface (display and buttons).

Figure 1.1. An example POCT as envisioned by IBM Corporation [1].

1.1. TOWARDS IMPROVED HEALTHCARE: POINT-OF-CARE TESTS 3

1.1.2 Market and opportunities

During 2009 the market for POCTs still remained only a fraction (14% or 6.9$ bil-lion) of the total in-vitro diagnostic (medical tests) market [2]. The POCT marketis dominated by a small number of established point-of-care products, most notablyglucose tests, rapid streptococcal tests and pregnancy tests. However, these POCTsare technologically relatively simple and cannot handle more complex diagnostictasks, such as nucleic acid or antibody analysis, required, in for example, antibi-otic resistance screening [1]. To realise complex diagnostic devices, microfluidiccomponents are required to handle sample transport and preparation. Miniaturisa-tion of fluidic components is currently a major focus of many academic groups anddiagnostic companies.

However, there are not only technological hurdles for the wider adaptation ofmore advanced POCTs. The complex structure of the diagnostic market as well asstringent regulation from public health agencies contribute to the slow growth expe-rienced the last years. Moreover, the reimbursement model for diagnostic devices,the acceptance by medical personnel, complex business models and large geographi-cal di!erences may contribute to the di"culties in replacing centralised testing withPOCTs [1].

Nevertheless, there are a few pioneering POCTs on the market that uses mi-crofluidics. Among these are Abbott’s i-Stat (blood gas analysis) and Biosite’sTriage System (immunoassays), both shown in Figure 1.2. As POCTs become ac-cepted as trusted alternatives to centralised medical testing, they have potentialto revolutionise the healthcare market, in particular home-tests which represents alarge untapped segment. An important user group will also be developing countrieswhere inexpensive POCTs will significantly improve the accessibility and quality ofhealthcare [3].

Figure 1.2. Two commercial POCT labs-on-chip based on microfluidics currentlyon the market. Left: Abbott’s i-Stat for electrolytes and blood gases tests. Right:Biosite Triage System for immunoassays.


1.2 Advantages of microfluidics for medical testingLabs-on-chip are built in dimensions measured in micrometers and handle liquidsmeasured in nanoliters. There are several reason why labs-on-chip are built thatsmall. In the microscale world, things are quite di!erent from how we are used toperceive them in macroscale world. With distances shrunk a million times, frommeters to micrometers, the surface-to-volume ratio increases linearly and surfacedependent forces, such as friction, dominate completely over volume related forces,such as inertia. In microfluidics, the immediate e!ect is that fluid flows are com-pletely laminar, heat transfer is very fast, the di!usion times are short, and thelikelihood of a molecule interacting with a channel surface is high. In addition tosmaller instruments and less consumption of expensive reagents, the scaling e!ectsbrings additional advantages:

• Well defined laminar flow, the sample is easy to control in the chip as the flowis completely laminar

• Faster reaction times, reactions happen faster when di!usion lengths areshorter

• High degree of parallelization, many tests can be done in parallel on the samedroplet of sample

The benefits o!ered by downscaling for medical diagnostics and chemical anal-yses are undisputed and today almost all new diagnostic kits or analytical systemsuse miniaturised components or e!ects associated with the microscale.

1.3 Lab-on-chip developmentThe development of a lab-on-chip is a multidisciplinary challenge, involving manydi!erent disciplines. This first part of the thesis deals mainly with the engineeringand integration aspects of developing prototypes for labs-on-chip, and not with theequally important biochemical and medical aspects. Successful development reliesheavily on proper choice of materials and manufacturing methods. The followingsection aims to give a short, non-exhaustive, review on the development process oflab-on-chip devices and to identify important limitations related to materials andmanufacturing methods.

1.3.1 Background on polymer technologyA short introduction to polymer science

In this section follows an introduction to some of the definitions in polymer scienceencountered later in the thesis. This is far from an exhaustive presentation ofpolymer science and the brave and curious is recommended to open Principles ofPolymerization by George Odian [4].

1.3. LAB-ON-CHIP DEVELOPMENT 5

General To start from the very begining: A polymer is a molecule composed ofseveral repeating structural units, monomers, connected by covalent bonds. Poly-merisation is the act of combining these monomers into long chains which can beeither linear or branched. Cross-linking is the formation of covalent bonds betweenpolymer chains. Curing is cross-linking of thermoset resins. The degree of poly-merisation defines to what degree the monomers have reacted. The gel point occurswhen the polymer first starts to solidify and form an infinite network, i.e. a gi-gantic "molecule". At the gel point, viscosity changes abruptly and the mobilityof the unreacted monomers decreases. If the gel point occurs at a high degree ofpolymerisation, the polymer network will have very little built in stress.

Polymer growth Polymer growth can occur by step-growth, chain-growth ormixed step-growth/chain-growth. In step-growth, the polymer network is formedby monomers reacting in a stepwise reaction between functional groups on themonomers to first form dimers than trimers, longer oligomers and finally chains.In step-growth polymerisation, the molecular weight increases at a slow rate, de-laying the gel point until very late in the polymerisation and very high degreesof polymerisation can be obtained (>95%). Step-growth polymerisation results inpolymer chains with an alternating sequence of monomers (ABABAB) and a highnetwork homogeneity. Chain-growth polymerisation involves opening up unsatu-rated monomers and adding them to the reactive end groups of the growing poly-mer chain. In chain-growth polymerisation, the individual polymer chains grow veryrapidly, creating a dense, nested, yarn-like structure that leads to an early gel-pointand eventually will hide the reactive end-groups from the unreacted monomers.This leads to a lower degree of polymerisation than for step-growth polymerisation.Moreover, during chain-growth, the polymer chains can react with themselves andform loops, resulting in a non-homogenous network. In mixed chain-growth/step-growth the two polymerisation processes compete with each other.

Radical polymerisation The polymerisation reaction can start spontaneouslyor more commonly with the help of an initiator, that creates a reactive intermediatecompound capable of successively linking monomers to polymer chains. Initiatorsare commonly triggered by heat or light irradiation. The most widely used initiatorsproduce free radicals that attack and open unsaturated bonds. The free radical poly-merisation can be divided into three steps: Initiation, when the first active centre iscreated from which a polymer chain is formed. This is usually accomplished by aninitiator that splits into two radical fragments upon actuation. Propagation, whenthe reactive end-groups of the growing chains react with new monomers in a serialfashion, where each reaction event recreates the active radical on the last addedmonomer. Termination, can be accomplished in many di!erent ways, for instancewhen there are no reactive groups left, when two radicals combine or by dispro-portionation. Chain transfer occurs when the reactive site of a growing polymerchain is transferred to another molecule. This could have the e!ect of increasing


the degree of polymerisation by transporting the reactivity of "hidden" reactive sitesout of the "yarn ball", to initiate a new chain of propagation steps.

Bulk properties The physical properties of the polymer are strongly dependenton the length and size of the polymer chains, to what degree they have branchedand cross-linked with each other, as well as on the mechanical properties of theindividual monomer molecules.

The glass transition temperature Tg, defines the temperature above which thepolymer chains vibrational energy exceeds the van der Waal forces keeping themtogether. Above the glass transition temperature the chains can slide relative toeach other, and the polymer can deform. The transition does not happen exactlyat Tg, but in an interval around this temperature. The size of this interval actuallyreveals something about the nature of the polymer network. A broad transitionindicates an in-homogenous network with chains of di!erent lengths. A narrowtransition indicates a homogeneous network, with very well-defined lengths andsizes of the polymer chains. Typically, chain-growth polymers have a broad glasstransition and step-growth polymers have a narrow glass transition, reflecting thedi!erence in their polymerisation mechanisms.

Figure 1.3. The loss tangent as a function of temperature for photopolymerised filmsof an epoxy-acrylate and a thiol-ene. The peak of the loss tangent curve indicates theglass transition temperature. The narrow glass transition of the thiol-enes comparedto most other polymers indicates exceptionally homogenous networks. Reproducedfrom Ref. [5] .

1.3.2 Polymer materialsThermoplastics

Thermoplastics are polymers that are not covalently cross-linked (Figure 1.4), andmelts at temperatures above Tm and freeze to a glassy state at temperatures belowTg. From a manufacturing point of view, the main advantage of thermoplastics


is their ability to be melted and reshaped against a mould, enabling productionof thermoplastic parts with high throughput. From a lab-on-chip perspective, theavailability of many commercial, medical grade formulations is a great advantage.The sti! mechanical properties also provide structural support and protect the sen-sor and the microfluidic network. However, many solvents, common in chemicalanalysis and separation, dissolve thermoplastics. Nevertheless, most commerciallabs-on-chip are made of thermoplastics. Common thermoplastics used in microflu-idics are poly(metylmethacrylate) (PMMA), polycarbonate (PC) and COC (CyclicOlefin Copolymer).

Thermosets

Thermosets, are polymers that are covalently cross-linked (Figure 1.4), and thus donot melt. From a manufacturing point of view, thermosets are shaped during thepolymerisation and cross-linking process. Because of the covalent bond formation,thermosets exhibits higher residual stress, shrinkage and crack-formation comparedto thermoplastics. From a lab-on-chip perspective, the main advantages of the ther-mosets are their geometrical stability and solvent resistance. Common thermosetsused in microfluidics are poly(dimethylsiloxane) PDMS (an elastomer), the hardresist SU-8 (MicroChem, USA), and the optical glue NOA81 (Norland Products,Inc, USA), that has recently been used for solvent resistant microlfuidics [6].

Figure 1.4. The di!erence in structure between thermoplastics and thermsets.

Elastomers (PDMS)

An elastomer is a rubbery and elastic polymer; it can be a thermoplastic or a ther-moset. It has few cross-links between the chains and thus a low E-modulus andhigh yield strength, compared with other materials. In microfluidics, the thermosetelastomer poly(dimethylsiloxane) PDMS is the dominating material for prototyp-ing microfluidic devices. PDMS is easy to handle in small laboratories, flexible butsturdy enough to manipulate, biocompatible, inert and easily bonded to silicon or


glass using oxygen plasma treatment. Because of the few cross-links and high fillercontent, the polymerised films experience only moderate shrinkage (1-3%) duringpolymerisation [7] and can replicate nanometer sized features. However, high vol-ume fabrication schemes for patterned PDMS layers are currently not available.From a lab-on-chip perspective, the elastic mechanical property has enabled inte-gration of pneumatically actuated valves and pumps in microfluidic chips, and directsealing to smooth surfaces. Moreover, PDMS is resistant to high temperatures, ox-idation and many chemical and biological environments. It also exhibits high gaspermeability, which is important for many living cell studies.

However, an important limitation with PDMS is the di"culty to permanentlymodify the surface, due to the high mobility of the polymer chains. Furthermore,the polymer network absorbs small molecules [8], leaches uncured monomers [9]and swells in solvents [10]. For example, it was shown that PDMS implanted intodogs absorbed 0.7 % of its own weight in small molecules, mainly lipds [11]. Sev-eral attempts have been made to improve PDMS. Notably, Rolland et al. devel-oped a fluorinated elastomer with similar mechanical properties as PDMS, but withdramatically improved solvent resistance and Kyung et al. [12] developed a pho-tocurable PDMS for faster curing. Others have tackled the absorption and swellingproblem by coating the inside of PDMS channels with di!erent polymers or inor-ganic materials to block the di!usion of small molecules [13, 14, 15].

1.3.3 Rapid prototypingThe rapid prototyping of microfluidic devices, and in particular of labs-on-chipplaces high demands on the resolution of the microstructuring process. In gen-eral high-resolution replica moulding must be used to reproduce smooth micro-and nanoscale features. In replica moulding a replication master, with the inversegeometrical features with respect to the finished device, is first produced using litho-graphic techniques. The replication master is used in a polymer replication process,such as hot embossing, injection moulding or casting. An important limiting factorin lab-on-chip prototyping is the time-consuming back-end processes. These pro-cesses are often performed in a serial fashion and comprise, for instance, drilling,surface modification, biofunctionalisation and bonding. The back-end processes areestimated to make up 80% of the total cost and time of lab-on-chip prototyping,and commercial manufacturing [16].

1.3.4 Polymer microstructuring methodsCasting

Casting involves pouring a liquid thermoset prepolymer on the replication masterand curing using heat or UV-light (Figure 1.6). Casting is an uncomplicated processwell suited for small-scale rapid prototyping, as it typically requires a very smallinvestment in equipment (UV-lamp or oven) and requires little or no process opti-misation. PDMS is casted in a process called soft lithography. However, there are


Figure 1.5. A schematics of the development process of labs-on-chip.

no high volume commercial casting processes for the replication of micron-sized fea-tures, which limits the capability to scale up production of the prototyped device.Casting is a planar process and the realisation of 3D structures requires drilling orpunching vertical interconnects between stacked channel layers.

Hot embossing

Hot embossing involves heating a thermoplastic substrate to just above Tg undervacuum (Figure 1.6) and press it against a replication master to imprint microstruc-tures. Hot embossing requires investments in expensive equipment to handle pres-sure, heat and vacuum control. The process requires optimisation of temperatureand embossing time for each new pattern and has a cycle time for a 4" wafer around4-15 minutes [16]. Hot embossing works well with small feature sizes and has ex-cellent dimensional control, but has problems with high aspect ratios and can onlyreplicate planar features [17]. Typical thermoplastic materials used are PMMA, PCand COC. The cost of a hot embossing system starts around 10k$ [16].

Microinjection moulding

Injection moulding typically involves melting thermoplastic pellets that are injectedinto a closed replication tool at a high pressure and a high temperature. As injec-tion moulding is the dominant replication techniques for plastics in general, mostcommercial labs-on-chip parts are manufactured this way. However, academic ac-cess to injection moulding equipment is limited because of very high machine costs(>75k$) as well as maintenance costs [16]. Moreover, there are many process pa-rameters to be optimised, and it usually takes considerable time to achieve a goodmicro-structured polymer. The moulding process has a high degree of automationand moulding times vary from 30 sec to 2 min [17]. Thermosets can also be injectionmolded but the process is generally more complex.


Figure 1.6. Comparison of the casting process (left) and hot embossing process(right)

1.3.5 Back-end processesPorts and interconnects

Both casting and hot embossing are planar processes producing polymer sheetswith microstructures on one side. Ports for fluidic access must be opened throughthe polymer layers to connect to the microfluidic channels, so called chip-to-worldconnections. Moreover, to realise 3D structures these layers must be stacked ontop of each other and vertical interconnects must be defined between them. Portsand interconnects can be drilled in thermoplastics and punched in PDMS, but theresolution and spacing between the holes is limited.

Injection moulding of thermoplastics can, in contrast to casting and hot emboss-ing, directly mould 3D structures.

Surface modification and control

Surface modification in labs-on-chip devices is one of the most important steps inthe manufacturing process. Common tasks for surface modification in a lab-on-


chip, exemplified in Figure 1.7, includes blocking of non-specific binding of proteins[18], spatial control of surface wetting [19, 20], and attachment of bioreceptors [21].Modification of the surface properties can be achieved in various ways: grafting ofpolymer chains on the surface [22], chemically attaching mono-layers of molecules[23] or physically adsorbing molecules on the substrate [24, 25]. Moreover, in manyapplications, surface modification must be spatially controlled to specific areas onthe chip. With the proper choice of chemistry both grafting and chemical linkingcan be patterned using UV-light and a stencil mask. Physical deposition may bepatterned using a physical mask during the deposition process or by micro-contactprinting [26, 27]. However, as the deposited material is not bonded to the surface,di!usion may lead to low resolution or adsorption into the substrate. Microfluidicdevices made of sti!er materials, such as thermoplastics or thermosets, providegood substrates for stable and permanent surface modification. On the contrary,devices prototyped in PDMS will have problems with permanent channel surfacemodifications, requiring complicated workarounds, as described earlier.

Figure 1.7. A fictive lab-on-chip with a hydrophobic valve for timing of samplereaction with reagent, anti-fouling coating to avoid loss of sample at the walls anda detection zone coated first with a linker layer and subsequently spotted with anti-bodies.

Bonding and integration

When selecting bonding method for labs-on-chip, biocompatibility is a major con-cern as bonding normally constitutes the last step in the fabrication scheme. Typ-ically, bio-reactive molecules, e.g. antibodies or antigens, are deposited prior tobonding. Functionalisation after bonding, in a closed o! device, is a complex andslow process requiring individual filling of each channel.

In Table 2.1, features of di!erent bonding techniques for polymers in microfludicsare listed.


Table1.1.

Com

parisonof

existingbonding

techniquesfor

polymer

microfluidic

devices.

Bondingm

ethodStrength

Processcom

plexityBondingtim

eBiocom

-pati-ble

Materials

Ref

Thermalbonding

Medium

LowLong

No

PMM

A,PC,CO

C[28,29,30]

SolventsH

ighLow

ShortYes

PMM

A,PC,CO

C[31,32,33,34]

Ultrasonic

weldingM

edium-high

Medium

Medium

YesPM

MA

[35]Laserwelding

Medium

-highH

ighM

ediumYes

PMM

A,CO

C[36,37]

Plasma

treatment

High

Medium

LowN

oPD

MS

[38]Clam

pingLow

Low-

YesPD

MS

Gluing

High

Medium

-high

Short-m

ediumYes

PMM

A,CO

C[39,40]

Adhesivefilm

sM

ediumLow-m

ediumShort

YesPM

MA

[41]

O!-stoichiom

etryM

edium-high

LowLow-M

ediumYes

PDM

S[42]

Surfacem

odificationM

edium-high

Medium

Medium

YesPC,PD

MS

[43,44]

1.4. THE IDEAL PROTOTYPING SYSTEM FOR LABS-ON-CHIP 13

As seen from the table, the biocompatibility requirement directly disqualifiesthermal bonding, the most common bonding technique for thermoplastics, as wellas plasma bonding, the most common technique for joining two PDMS layers. Ultra-sonic welding and laser welding are capable of focusing energy only to the interfaceand bond thermoplastics, but require substantial process optimisation, as well as ad-vanced equipment. This leaves solvent bonding, physical clamping, gluing, adhesivefilms, surface modifications and o!-stoichiometry bonding.

Out of these both clamping and o!-stoichiometry are specific to PDMS. Rub-bery materials like PDMS, can be physically clamped but the adhesion is low andhigh clamping pressures risks deforming the channels. By mixing PDMS in o!-stoichiometric ratios: one layer with excess of the base (vinyl groups) and one layerwith excess of curing agent (Si-H groups), they can be covalently bonded to eachother without using plasma treatment. [42].

In solvent bonding, the topmost surface of a thermoplastic polymer is dissolvedusing a solvent to achieve chain-entanglement across the bonding interface. Ap-proaches where the solvent is only applied to the top layer have been shown, whichmakes the bonding process biocompatible[34]. Glues [39, 40] can be used to jointhermoplastics at room temperature and do not require high temperatures [45].However, the application process is critical, and care must be taken not to acciden-tally fill channels [46]. Double-sided adhesive films [41] do not block the channels,but are sensitive to solvents and create channels with di!erent top and bottomsurfaces. A major problem is that most glues and adhesives do not adhere to therubbery surface of PDMS, which is the material of choice in academic proof ofconcept devices.

The surfaces of PDMS and thermoplastics may be modified with reactive moleculesthat can covalently link to each other under biocompatible conditions, such as poly-mer coatings [43] or organofunctional silanes [44].

1.4 The ideal prototyping system for labs-on-chipThe ideal prototyping method for labs-on-chip is fast, relies on inexpensive mate-rials, allows for 3D features (through holes) and do not require access to expen-sive/technically complicated facilities. The ideal prototyping material is a tuneablematerial that can recreate pneumatic valves and pumps but also can provide struc-tural stability and external interface, such as manifold integration and direct tubingconnections. It is chemically inert, compatible with chemical and biological sampleswithout absorbing them and allows stable and patternable surface modification forcontrol of wetting and biological functionalisation. Finally the material allows forbiocompatible bonding to a wide range of substrates.

These properties are concretized in the list below:

1. Tuneable mechanical properties. The mechanical properties of an idealprototyping material are somewhat contradicting. On the one hand, it mustmirror the sti!ness of commercial thermoplastics, to produce geometrically


stable microfludic chips with robust external chip-to-world interfaces. On theother hand, it must be soft enough to allow for the pneumatically actuatedvalves and pumps, commonly used in PDMS.

2. Chemical inertness and low interaction with the sample. To be able toanalyse low concentration samples, the ideal material must not absorb smallmolecules, such as proteins or DNA from the sample, react with the sample orleach uncured components that may interact with the sample or the sensor.

3. Solvent resistance. Critical for many chemical reactions and separationprocesses is the use of harsh solvents. The ideal material must therefore notdissolve or swell in these solvents.

4. Direct, patternable and stable surface modifications. The spatial con-trol of surface properties is instrumental for controlling liquids and immobil-lizing biological receptors on the chip. The ideal material allows for spatialcontrol of surface modifications, preferably without the need to first activatethe polymer surface by plasma or solvents.

5. Fast, scalable and utilizing inexpensive materials and processes. Theprototyping method has a fast curing/microstructuring step and uncompli-cated back-end processes, to allow a rapid development process. To be use-able in academic research, the ideal prototyping method relies on inexpensivematerials and do not require access to expensive/technically complicated fa-cilities. To allow a fast transition to commercial production, the method ispossible to scale up to medium or large-scale production.

6. Three-dimensional microfluidics. Advanced labs-on-chip must be ableto handle multiple liquids, something which often requires 3D microfluidicchannels with under- and overpasses. The ideal prototyping method there-fore allows for e"cient fabrication of multiple vertical interconnects betweenchannel layers.

7. Biocompatible bonding. Essential for labs-on-chip is an uncomplicatedand biocompatible bonding method to surfaces that are functionalized withproteins and DNA. The ideal prototyping method form a strong bond to awide number of materials under biocompatible conditions.

In academia, soft lithography in PDMS is the method predominately used forrapid prototyping of microfluidic devices. For some applications, PDMS is the per-fect material, in particular for cell studies, as it is easy to structure and permeableto oxygen. However, PDMS cannot accomplish some of the basic features of anideal lab-on-chip material, such as low absorption of molecules from the sample andstable surface modifications. Moreover, the planar fabrication method complicatesfabrication of 3D microfluidic networks. Biocompatible bonding is possible usingo!-stoichiometric mixtures, but only to other PDMS layers. Furthermore, concerns

1.5. SUMMARY AND CONCLUSIONS 15

have been raised that devices developed in PDMS will be di"cult to commercialisedue to the extensive redevelopment required to transfer them into commercial ther-moplastic devices [16, 47].

The question naturally arises why thermoplastics are not used directly in theprototyping process? After all, thermoplastic materials address many of the proper-ties of the ideal prototyping material, such as chemical inertness, direct patternablesurface modifications, biocompatible bonding (adhesive bonding, laser or ultrasonicwelding) and 3D microfluidics (injection moulding). The reasons why thermoplas-tics have not gained foothold in academic prototyping lie in its cost and complexity.Although the thermoplastic material in itself is cheap, the cost of equipment andfacilities is high. Moreover, extensive process-optmisation is required for each newdesign.

Surprisingly, there have been few, if any, attempts to develop a polymer systemspecifically suited for lab-on-chip applications that respond to the challenges listedabove.

1.5 Summary and conclusionsThe increasing availability of diagnostic point-of-care devices has potential to in-crease the quality of the healthcare system by putting the power of large centralisedlaboratories into the hands of doctors and patients. The portability of the testsand the closeness to the patient, are believed to increase the likelihood of a quickand accurate diagnosis and therapy adjustments at the hospital bed, in the doc-tor’s o"ce, or at home. Furthermore, portable point-of-care tests will improve theaccessibility and quality of healthcare in the developing world, where the access tohospital is limited.

The first generation of point-of-care tests is already available on the market:blood glucose tests, rapid streptococcal tests and pregnancy tests. However, toenable more complex tasks such as nucleic acid or antibody analysis, microfluidicsmust be integrated to handle sample preparation and transport. The development ofthese technologies is currently the focus of many academic groups. However, thereis a material and manufacturing bottleneck limiting the transfer of technologiesfrom academia to commercial products. Firstly, the most commonly used materialfor lab-on-chip proof-of-concepts PDMS, has some serious drawbacks, preventing orcomplicating the development of many important functions in labs-on-chip. Sec-ondly, prototypes in PDMS must be completely redeveloped, both from a materialand manufacturing perspective, to be transferred to into commercial thermoplasticdevice production.

The following two chapters address both of these problems. In Chapter 2, twoimportant improvements of the PDMS manufacturing process for labs-on-chip areintroduced and demonstrated. In Chapter 3, a novel prototyping material free fromthe limitations of PDMS and with potential to bridge the development gap betweenacademia and commercial products is presented.

Chapter 2

Novel manufacturing methods forlabs-on-chip

This chapter introduces and demonstrates two novel manufacturing methods forproof-of-concept POCTs, using soft lithography in PDMS. Specifically, a methodfor biocompatible bonding of PDMS to thermoplastics and a high yield fabricationmethod of 3D microfluidic devices are introduced. Both of these methods aim toremove limitations related to the PMDS material discussed in Chapter 1.

In the first section, a method for joining materials with di!erent surface energies,such as PDMS to thermoplastics, is presented. The first section also presents thepackaging and integration of an optical label-free sensor, in which the novel bondingmethod is used. Furthermore, it discusses design-issues related to the optical sensorand considerations for the mass-transport of analyte down to the sensor surface.The second section describes a high yield process for creating vertical interconnectsusing a novel method to inhibit the PDMS polymerisation at specific locations.Furthermore, the second chapter briefly examplifies a lab-on-chip application inwhich the inhibition method is used.

2.1 Dual surface-energy adhesive for the integration andpackaging of an optical label-free sensor

2.1.1 Dual surface-energy adhesivesDue to the di!erence in physicochemical properties, thermoplastic materials do notgenerally form irreversible bonds with PDMS, even after oxygen plasma treatmentor heating. However, the very di!erent material properties of thermoplastics andelastomers, o!er many opportunities for hybrid microfluidic devices using the advan-tages of both. Elastomers enable pneumatic valves, thermoplastics enable a varietyof reliable external interface options, such as manifold integration, direct tubingconnections, and gasket connections. Therefore, a general technique for integratingmaterials with di!erent surface energies in a rapid and uncomplicated fashion, is

17

18 CHAPTER 2. NOVEL MANUFACTURING METHODS FOR LABS-ON-CHIP

presented in Paper 1.

Figure 2.1. Dual surface-energy adhesive

Only a very few strategies exist to bond PDMS to thermoplastics, notably CVDprocesses or silane coatings [44, 48]. These methods typically require several lengthysteps: plasma treatment (!1 min), incubation with the silane linker (!20 min) andbonding (!10 min) [48]. Another alternative, both faster and more flexible, is theuse of patterned double-sided pressure adhesive films. However, the glue foundon most adhesives is an acrylic based glue which has no or limited adhesion tothe low-energy PDMS surface. In one attempt to overcome this problem, uncuredPDMS was spun and subsequently thermally cured on top of the adhesive film[49]. However, this technique involved several complicated steps in preparing theadhesive film and was considerd too elaborate. Instead, Paper 1 describes a rapidand uncomplicated bonding method using a patterned dual surface-energy adhesivefilm (5302A, Nitto Denko, Japan). This adhesive film, designed for the mobilephone industry to attach the key pads made out of rubber, has one side coatedwith a silicon based pressure sensitive glue, and one side with an acrylic based glue.The silicon glue bonds e"ciently with low energy surfaces like PDMS, while theacrylic glue bonds well to the PMMA . Before bonding, openings are cut out in theadhesive film for fluidic and optics ports. The bond strength of the tape to PDMSand PMMA was characterised by peeling the tape from the substrate at 180! (Table2.1).

Table 2.1. Peeling strength of Nitto Denko 5302A (N/20mm width)

Surface Acrylicside

Silicon side

PDMS 0.15 5PMMA 20 10

The advantage of the dual surface energy film is rapid bonding of PDMS withouthaving to use liquid glue or plasma treatment. It allows for a rapid and uncompli-cated bonding, under biocompatible conditions with a high yield. The drawback issolvent incompatibility.

2.1. DUAL SURFACE-ENERGY ADHESIVE FOR THE INTEGRATION ANDPACKAGING OF AN OPTICAL LABEL-FREE SENSOR 19

2.1.2 Background on optical biosensingOptical sensing is a powerful technology for label-free detection for point-of-caretests. It o!ers the sensitivity required for detecting the low concentrations of ana-lytes present in bodily fluids, is free from electrical interference, has a wide dynamicrange and multiplexing capability. One important drawback of optical sensors istheir sensitivity to temperature which, without proper temperature control, maycause unwanted drift in the output signal. Many of the analytical instruments usedin universities and pharmaceutical companies are based on optical transducers. Forexample, both the Biacore system (GE Healthcare, Uppsala), the gold standardin protein interaction measurements, and the AnaLight system (Farfield, UK), asystem for protein conformation measurements, use optical transducers. Accord-ingly, to bring these powerful analytical tools into the hands of a wider user-base,there is a strong interest in integrating optical sensors in labs-on-chip. Even thoughan abundance of optical sensor principles has been demonstrated, very few havesuccessfully been integrated in complete labs-on-chip.

Some electromagnetism

All molecules interact to a varying degree with electromagnetic fields that passthrough them. The electrons in the molecule experience a force when they areexposed to the oscillating electromagnetic fields of light. If a molecule has free elec-trons, it will be polarised by the electric field resulting in the formation of an electricdipole. The extent of polarisation is dependent on the size, shape and orientationrelative to the electric field. The quantity known as electric susceptibility, "e of amolecule quantifies its polarisability. When a polarisable molecule is placed in anoscillating electromagnetic field, such as light, the electrons within the moleculeswill start to oscillate and produce a current. This modifies the local relative permit-tivity #r of the dielectric material, which in turn modifies the local refractive indexn = "

#r = "1 + "e. It is this local change in refractive index, that results in amodified light propagation speed, that can be measured by optical sensors, not themass directly. This is possible since an optical waveguide in an aqueous mediumwill couple some of its energy into the surrounding water, this field is known as theevanescent field and extends typically some 100 nm from the surface. When biologi-cal molecules, which have a higher electrical susceptibility than water, interact withthis field they change the local dielectric constant. Optical waveguide biosensorscan thus directly probe their surroundings by a label-free detection method.

Slot-waveguide ring resonator

Great e!ort has been put into improving the resolution, or limit of detection, ofoptical biosensors through the design of resonator structures with extremely highquality factor, or Q. The quality factor is defined as Q = $0/!$, where $0 is thecentre wavelength and !$ is the spectral with of of the resonance determined at halfthe peak maximum. The higher the Q, the narrower is the resonance and a lower


limit of detection can be achieved. High Q-resonators can be achieved by couplinglight into a circular waveguide that can fit an integer number of wavelengths aroundthe perimeter of the circle. For these modes, standing waves or resonance will besupported in the ring. To excite the rings, light from a tuneable laser is coupledinto the rings through a straight waveguide passing very close to the ring. As thewavelength of the tunable laser is swept, some wavelengths will match a resonantwavelength of the ring and light of that wavelength will be coupled into the ringand disappear from the output spectra. As the analyte attaches to the receptorligand on the ring resonator surface, the local refractive index is modified, theresonance frequency of the ring changes and the dip in the output spectrum willmove. These types of sensors have been demonstrated to have high sensitivity andlimit of detection [50, 51, 52, 53].

However to increase the interaction with the sample and enhance the sensitivity,the slot-waveguide strategy was introduced [54, 55]. Essentially this strategy takesadvantage of a slot in the waveguide, small enough that a large portion of theelectromagnetic field power propagates inside the liquid filled open slot. In this waythe interaction with the sample can be increased. Moreover, as the water filled slothas a negative thermo-optic coe"cient, it can balance the positive value of bothsilicon oxide (bottom) and silicon nitride (wave guide) and contributes to renderthe waveguide temperature insensitive, as discussed in Paper 2. This solves thetemperature sensitivity mentioned earlier as one of the most important drawbacksof optical sensors, and is a huge advantage for the practical implementation ofPOCT based on optical biosensors. In Paper 1, six active and two references ringresonators were multiplexed on the chip enabling the simultaneous measurementsof several analytes.

Figure 2.2. The slot waveguide enables increased light/sample interaction and thushigher sensitivity, since up to 40% of the total optical power [54] propagates in theliquid filled slot as illustrated in the cross-section above.


Optical chip

The optical sensor, partly designed and fully fabricated by Dr. Kristinn Gylfasson atKTH, consisted of a silicon chip with patterned waveguides and an optical distribu-tion network in silicon nitiride on thermally grown oxide. The surface was coveredwith a cladding of silicon oxide, except over sensor rings, where the cladding wasremoved. An overview of the optical components is shown in Figure 2.3. Beforethe assembly, the nitride ring resonators were selectively coated with glutaralde-hyde using a process developed and applied by Dr. María José Bañuls Polo, andher collegues, at the Department of Chemistry, Universidad Politéchnica de Valen-cia (Valencia, Spain). It allows for functionalisation of only the nitride waveguideswhile leaving the silicon oxide uncoated [56].

Figure 2.3. A top view of the layout of the optical chip: light is injected at thesurface grating coupler (C) and split by a multi-mode interference splitter (B) tothe six transducer channels M1 to M6 and the two reference channels REF1 andREF2. Inset are: an optical micrograph of the splitter (B), electron micrographs ofthe grating coupler (C), and a slot waveguide ring resonator (A), with an enlargementof the coupling region between the straight waveguide and the ring. The optical chipwas partly designed and fully fabricated by Dr. Kristinn Gylfasson at KTH

2.1.3 Background on mass transportA sensor with a potentially low limit of detection and high sensitivity is no guaranteefor a good assay performance. The surface chemistry plays an essential role toe"ciently capture the analyte from the sample and the mass transport of the analyteto the sensor is also critical in governing the dynamics of the sensor response, and


ultimately the performance of the assay. Three physical processes control the masstransport to the sensor surface: di!usion, convection and surface reactions.

In the design of the microchannels, the interplay between these three e!ectsmust be carefully considered and modelled. The full problem can be numericallysolved, but an analysis with the help of dimensionless numbers, comparing therelative importance of these three e!ects, is often enough for the initial design of adevice. The transport-processes in biosensors have been reviewed in some excellentreviews, such as those by Gervais et al. [57] and Squires et al. [58]. Below followsa short introduction to three common dimensionless number used in the design ofLabs-on-chip.

Figure 2.4. Model system for mass-transport analysis [58]. Solution with targetconcentration c0 flows with a volumeric flow rate Q through a channel of height Hand width Wc over a sensor of length L and width Ws that is functionalised with !s

receptors per unit area. The binding reactions have constants kon and koff and theanalyte di!usion constant D.

Imagine a microchannel where the target analyte flows with the volumeric flowrate Q through a channel of height H, width Wc and where one wall contain asensor of width Ws, and length L. The sensor is functionalised with %s receptorsper unit area, and the solution contatins target molecules with concentration c0and di!usivity D. The analyte molecules binds to the receptors at the surface withbinding constants kon and koff .

In static conditions, when there is no flow, Q = 0, di!usion is the only transportprocess from the bulk to the sensor, and a depletion zone will be formed above thesensor with radius & =

"Dt. The collection rate at the sensor can in this case be

approximated by jD ! Dc0/&. As the depletion zone grows in the channel, thedi!usion flux becomes smaller and collection rate of analyte molecules at the sensorsurface decreases. Since di!usion in slow, the accumulation of enough analyte onthe surface may take hours or days in low concentration solutions [59].

If convective transport is added to the model, the growth of the depletion zone


is halted and gives a steady state depletion zone above the sensor. The Pécletnumber characterise relative strength of convection and di!usion and the nature ofthe mass-transport depletion zone around the sensor. The depletion zone is thincompared to the channel if Pe # 1, and the steady state di!usion zone &s abovethe sensor will be characterised by the distance at which the time for an analytemolecule to convect past the sensor is exactly equal to di!use down to the sensor. Athigher flow rates the depletion zone will be thinner and collection rate at the surfacewill increase. Contrary, if Pe $ 1, the depletion zone will extend far upstream andthe transport of analyte will be limited by di!usion, thus increasing the collectiontime scale at the sensor. At su"ciently low Péclet numbers, the di!usion zonewill exactly balance convection and all the analyte molecules in the sample can becollected by the sensor, but at the expense of decreased collection rate at the sensorsurface.

The time scale of the surface reaction compared to the di!usion time scalegive rise to another dimensionless parameter called the Damköhler number whichcaptures how quickly the sensor equilibrates. If Da # 1, the equilibration is limitedby the rate of target di!usion to the sensor, whereas the reaction itself limits thesensor kinetics if Da $ 1. Normally one strives to engineer a microfluidic systemso that equilibration is reaction-limited in order to measure the reaction kineticsrather than the di!usion kinetics.

In summary, if the chip is designed for short assay times, the mass transportshould be optimised to be reaction-limited with a high Peclet number to minimisethe size of the di!usion-limited depletion zone. This can be realised for example byusing a surface chemistry with high ligand density and use a high flow rate. If thechip must economise with a small volume of analyte, and require a high capturerate, the flow rate must be kept low enough to balance di!usion.

Table 2.2. Dimensionless numbers relating the three transport processes in Labs-on-chip.

Re Reynolds 'U0H/( inertial/viscousPe Péclet Q/DWs di!usive time/convective timeDa Damköhler kon%sH/D reaction time/di!usive time

2.1.4 Microfluidics design and manufacturingThis section deals with the design and manufacturing of the microfluidic networkfor the optical sensor chip.

Mass transport analysis

In the design of the microfuidics, the dimensionless numbers were calculated toensure the channel dimensions and flow speed would not limit the performance of


the sensor. As a model system we used the anti-BSA to BSA binding reaction [60]which was also tested during the experiments. With a flow of Q = 10 µL/min, atypical di!usion constant for small proteins of D = 10"10 m2/s and with a sensorsize of Ws = 200 nm, the Peclet number was calculated to, Pe = 8340 # 1, ensuringa thin depletion layer compared to the channel height. Using the the surface area ofa BSA molecule (56 nm2) and its molecular weight (66 kDa) [61], the BSA surfacecoverage was estimated to 1.9 ng/mm2. The Damköhler number for the system wascalculated to Da = 1 % 10"5, indicating a reaction-limited system.

However, there was a concern that the small dimension of the slot, 200 nm wideand 300 nm high, would limit the mass transport of analytes to and from the innersurfaces of the slot to such a high degree that it would impair the performanceof the sensor. A numerical investigation of the transport of analyte down to thesensor also revealed that the concentration of analyte on the inner sides of the slotcompared to the top surface of the waveguide was up to ten times lower, limitingthe advantage of the slot for fast binding kinetic measurements.

Figure 2.5. Results of the numerical investigation of mass transport. Mass transportinto the slot is di!usion limited, illustrated by the stream lines, and the concentrationof analyte (surface plot) is up to ten times lower in the slot than above the slot atsteady state. This will require longer incubation time than expected to take fulladvantage of the slot waveguide’s increased sensitivity.

Layout

The microfluidic layer, illustrated in Fig 2.6 (C) consisted of an array of six channels(M1–M6), each addressing one slot waveguide ring resonator. The channels are 200µm wide and 20 µm high over the sensor. The distance between the channels isset by the minimum spacing required to avoid channel-to-channel leakage, and theminimum channel width by the need for manual alignment of the microfluidic layerto the optics chip.

The elastomer PDMS was chosen as material for the microfluidic network partlybecause of its availability and its ease to manufacture in the laboratory, but alsofor its adhesion to silicon oxide through plasma activation of the PDMS. The mi-crofluidic layer was molded using soft lithography and through holes were punchedas liquid input and output ports as well as for laser access. The total area of the


Figure 2.6. (A) A photograph of one of the fabricated cartridges. The steel tubesglued to the hard plastic shell provide a stable fluidic interface to the microfluidicnetwork below. At the front long edge, the through hole for laser access down tothe optical chip is visible. The edge of the optical chip itself can also be seen in thecutout region of the hard plastic shell. (B) The inset shows one of the ring resonatorsin its fluidic channel. (C) The fluidic layout with individual channel to each sensor(M1 to M6).

silicon chip is 40 x 15 mm2, most of which is used as support for the microfluidicchannels and connectors.

The array format with one channel for each sensor, enable the simultaneousmeasurements of di!erent liquids, and a more reliable diagnosis. Reference channelsalso enables a thermal compensation technique that is instrumental for achievinglow detection limits in optical sensors.

2.1.5 Integration and bonding of the sensor cartridgeMicrofluidics and optical multiplexing was integrated on-chip, while fluidic pumps,tuneable laser source, read-out sensor and electronics were external in an instrumentdeveloped by Dr. Andrzej Ka$mierczak and Dr. Fabian Dortu at Multitel (Mons,Belgium). An exploded schematic of the integrated cartridge is shown in Figure 2.7and a photo of the chip, ready to be inserted into the read out system in Figure 2.6.

Reversible clamping was first evaluated as a bonding method for the microfluidiclayer to the optical chip, but it was di"cult to create an even clamping pressure overthe chip, and the 20 µm high channels were often compressed in some part of the


chip. Instead, the PDMS microfluidic layer was plasma activated to expose Si-OHgroups, which were covalently reacted with the uncoated silicon oxide surface of theoptical chip to form Si-O-Si bonds. Fig 2.6 (B) shows photo of a ring resonatorsensor in one of the microfluidic channels.

The hard plastic manifold in PMMA and the dual surface-energy adhesive film(Nitto Denko, 5302A) were micro-milled with openings for optics and fluidic ports.The adhesive film was subsequently aligned and applied to the PMMA manifold,which in turn was aligned and applied to the laminated optical and microfluidicchip.

Figure 2.7. A schematic exploded view of the sensor cartridge above the alignmentplatform, exposing the 4 permanently bonded layers of the cartridge: the opticalchip, the microfluidic layer, the adhesive film, and the hard plastic shell. Cutoutsin the hard plastic shell free the edge of the precision cut silicon optical chip foraccurate alignment against 3 pins protruding from the alignment platform of theread-out instrument. Light is coupled in from the top via a surface grating coupler,and collected at the long edge of the optical chip by imaging the output facets on a1D InGaAs photodiode array. Fluidic ports for sample injection are formed by steeltubes glued into the hard plastic shell.


2.1.6 Biosensing resultsThe performance of the cartridge was evaluated both on spotted and unspottedsensors. In Figure 2.8 (A), the sensor is functionalised only with glutaraldehydeand injected anti-BSA attaches non-specifically to the surface. In Figure 2.8 (B),the sensor has been spotted with BSA on top of the glutaraldehyde surface beforethe assembly. Anti-BSA is flushed through the channel and specifically binds to thespotted BSA molecules. The spotting of the surface was done by Dr. Gerhard Kres-bach at Zeptosens - A Division of Bayer (Schweiz) AG, (Witterswil, Switzerland).Further results, available in Paper 1, shows a surface mass detection limit of 0.9pg/mm2, which is among the best reported for integrated ring resonator sensors.

Figure 2.8. A comparison of real-time sensing results on unspotted and spottedsensors. (A) Shows the resonant wavelength shift as a function of time during injectionof an increasing concentrations of anti-BSA on a sensor with only glutaraldehyde.(B) Shows the resonant wavelength shift of a sensor spotted with BSA on top of theglutaraldehyde surface. More results are available in Paper 1.

2.1.7 Discussion and outlookThe design, fabrication, and characterisation of a packaged array of optical refractiveindex sensors, integrated with microfluidic sample handling in a compact cartridgewas demonstrated. Using the individually addressable transducers available on thechip, we separated and compensated for di!erent kinds of external disturbances,resulting in much improved noise level, compared to our previously published results.The multiplexed layout, with multiple sensor with separated flow compartment alsoallowed for temperature compensated measurements, demonstrated in Paper 2. Thetemperature compensation, combined with e"cient mass transport and the highsensitivity of the fabricated slot-waveguide ring resonators, yielded one of the bestmass detecion limis reported so far for integrated ring resonator systems.

The result is also good compared to other label-free methods, but it must bekept in mind that for a POCT ten minutes is a relatively long time and the smallestconcentration detectable during this time was 0.4 nM of anti-BSA (0.0615 µg/ml)which is higher than the concentration of many biomarkers in the blood (pM to fM)


[1]. Part of the problem is the di!usion limitation of analytes down into the 200 nmwide slot, which limits the performance of the slot waveguide for sensitive surfacemass measurements.

From a manufacturing point of view, the novel use of a dual surface-energy adhe-sive to bond PDMS to thermoplastics, enabled a rapid and biocompatible bondingmethod with tightly spaced microfluidic ports. By using the adhesive film insteadof gluing or surface modification, the time spent on back-end processes could bereduced. However, from a material point of view, PDMS itself was problematic. Ifthe PDMS adsorbed proteins, it could have led to a lower concentration of analyteover the sensor than in the initial sample. If this was the case, it would have ledto an underestimation of the sensor performance. Furthermore, PDMS absorbedthe methanol and ethanol used for the calibration of the sensor and slowly releasedthem in all channels . However, a solution for the PDMS problem has to wait untilChapter 3.

2.2 High yield process of vertical interconnects in PDMSfor batch manufacturing 3D microfluidics devices

In the quest to integrate more fluidic functions in labs-on-chip the complexity ofthe channel network inevitably increases. To freely interconnect di!erent on-chipregions requires three dimensional (3D) channel topology with under- and overpassesto allow the liquids to cross without mixing. Although, injection molding can tosome extent produce 3D microfluidic chip parts, both casting and hot embossingare planar fabrication methods. Three-dimensional microfluidic fabrication usingthese methods, must use at least two microfluidic layers with well defined vias, i.e.out of plane interconnecting fluidic channels to connect the layers (Figure 2.9).

Figure 2.9. Planar replication methods require vertical interconnects to create 3Ddevices and fluidic ports. These are often punched or drilled as a part of the back-endprocesses.

In both hot embossing and casting, interconnecting vias can only be definedafter the microstructuring process, by drilling [62, 63], laser ablation [64] or dryetching [65, 66] or punching (PMDS). This manual approach is acceptable for alow number of vias or fluidic ports spaced far apart, but for multiple tightly spacedinterconnects the via creation must be integrated in the microstructuring process.

2.2. HIGH YIELD PROCESS OF VERTICAL INTERCONNECTS IN PDMS FORBATCH MANUFACTURING 3D MICROFLUIDICS DEVICES 29

Many labs-on-chip make use of integrated pneumatic valves in PDMS[42]. Thesedevices consists of two layers of PDMS, one layer for fluidics and one layer with pneu-matic control channels. When the pneumatic channels are actuated the membrane,formed where fluidic and pneumatic channels cross, deflects and blocks the fluidicchannel. However, in many applications it is necessary to connect these two layers.The problem is to create a large number of vias for large scale integrated (LSI)microfluidic networks. In previous attempts, the strategy have been to remove theprepolymer at the location of the via holes before polymerisation. Anderson et. alintroduced the use of two-level masters with protruding posts at the via locations[67] (Figure 2.10). After pouring the liquid polymer on the master, a cover plateis pressed on top at high pressure to squeeze away the prepolymer of the top levelmold features [67, 68, 69]. After curing the plate is removed and the resulting PDMSfilm with open vias holes is aligned and bonded to a second layer. Unfortunatelyit is di"cult to achieve a high yield using this method, since it is not possible tosqueeze away all prepolymer from the protruding vias features with only clampingand a blocking, thin residual membrane often have to be removed manually or viasome dry etching process. Other techniques involve spinning [70, 71] or blowing [72]with gas to remove the prepolymer, all su!ering from uneven surface at the viaslocations due to surface tension e!ects.

Figure 2.10. Previous attempts for batch manufacturing of vertical interconnectshave focused on removing prepolymer before polymerisation.

In this work, the goal was to develop a more reliable process for the creationof residual-free interconnects for applications in labs-on-chip. Because of the dif-ficulties in completely squeezing away the prepolymer on top of the protrudingmold features, a novel strategy was used in Paper 3, to prevent the prepolymer topolymerise at the interconnecting via locations.


2.2.1 PDMS polymerisation processMost researchers working with microfluidics and labs-on-chip have at some pointmixed PDMS. However, its composition and how it polymerises is less known. Thereare two main suppliers of PDMS: Dow Corning (Sylgard 184) and Momentive (RTV615). They are both provided as two components: one base and one cross-linkingagent. The curing proceeds through hydrosilylation involving the addition reactionbetween polysiloxane in the base containing vinyl groups and the cross-linking agentcontaining Si-H functional groups [4]. The reaction relies on a platinum catalyst.

2.2.2 Residual-free interconnects by local inhibitionBy removing the platinum from the prepolymer exactly at the via locations, poly-merisation will be inhibited at those specific locations. In the literature, manysubstances have been reported to inhibit the polymerisation of PDMS, but one ofthe most e"cient are the tertiary amines. The tertiary amines create a chelatingcomplex with the platinum atom and prevents it from catalysing the polymerisation[73]. When a glass plate covalently coated with a layer of aminosilanes is clamped ontop of the PDMS prepolymer, the amines capture the platinum close to the surface.Because the di!usion of new platinum catalyst is limited in the squeeze film on topof the protruding vias features platinum will be completely depleted only at thoselocations. The result is polymerisation everywhere except at the via locations. Theinhibiting plate can be removed after polymerisation and with it follows the liqudprepolymer at the vias location, leaving membrane-free vias in the PDMS film.

2.2.3 Direct bonding using the inhibited surfaceA bonus, resulting from the above technique, is an alternative bonding method. Atthe surface of the inhibiting glass plate, enough new platinum can be resuppliedfrom the bulk to polymerise the prepolymer, but to a lesser degree of cross-linking.When demolded, this surface is a little sticky and tape-like. It has good adhesionto most surfaces and if new platinum is resupplied, for example by stamping, it canbond covalently to another PDMS layer.

2.2.4 3D microfluidic networks for labs-on-chipThe inhibition technology has so far been successfully implemented in lab-on-chipprojects at KTH by Mikael Karlsson. In a project to detect antibiotic resistancebacteria from whole blood, the technology is used to create vias and ultra-thinmembranes [74]. The project integrates on-chip PCR with optical detection ofbacterial DNA. Figure 2.13 shows a dual layer microfluidic device for on-chip PCR,fabricated in PDMS using the inhibition technology. Another application is 3Ddistribution network for cell sorting during sample preparation [75].

2.2. HIGH YIELD PROCESS OF VERTICAL INTERCONNECTS IN PDMS FORBATCH MANUFACTURING 3D MICROFLUIDICS DEVICES 31

Figure 2.11. PDMS is catalysed by Pt atoms and polymerises at elevated temper-atures. When a chelating agent (AEAPS) is coated on glass plate it captures the Ptatoms and prevent PDMS from polymerising.

2.2.5 Discussion and outlook

Using localised inhibition of the PDMS polymerisation we have demonstrated anuncomplicated batch process with a high yield to create membrane-free verticalvias in PDMS using soft lithography. The method enables complex 3D microfluidicdevices with small and tightly spaced interconnects (Figure 2.12) by eliminating theneed for manual punching of holes in PDMS layers or cleaning blocked vias fromresidual membranes and has already been successfully implemented in a numberof applications. While succesful in many applications, the procedure is limited tosurface inhibition and requires two-level molds.


Figure 2.12. This 3D basketweave structure was created using the localised inhibi-tion technique described in this section.

Figure 2.13. An example application of the via inhibition technology. A dual layermicrofluidic chip for on-chip PCR with vertical interconnects, fabricated by MikaelKarlsson (Microsystem Technology, KTH)

2.3. SUMMARY AND OUTLOOK 33

2.3 Summary and outlookThis chapter introduced two novel improvements of the manufacturing of PDMSmicrofluidic devices and exemplified them through applications.

The dual surface-energy adhesive film enabled the integration of PDMS microflu-idic layers with thermoplastic cartridge shells and allowed for rapid and leak-tightassembly with multiple, tightly spaced fluidic ports without leakage. The localisedinhibition technology of PDMS enabled complex 3D microfluidic devices to enablehandling of multiple liquids on a single chip.

However, many material issues still remain with PDMS as a lab-on-chip materialand it is time to set the stage for a successor.

Chapter 3

OSTE: a novel material toolbox forlabs-on-chip

This chapter introduces a novel prototyping polymer system aiming to bridge thegap between academic proof-of-concepts and commercial products. The novel poly-mer toolbox was developed with consideration to the requirements of an ideal pro-totyping system discussed in Chapter 1 (Section 1.4). The main objective was todevelop a versatile toolbox for academic development of lab-on-chip proof-of-conceptdevices, free from many of the adverse properties of PDMS, while at the same timebeing compatible with commercial prototyping, i.e. to mirror the properties ofcommercial thermoplastics.

In the first section of this chapter, a background of the versatile thiol-ene poly-mer system is given. Next, the novel o!-stoichiometry thiol-ene (OSTE) polymeris presented along with its powerful features for rapid prototyping of microfluidicsfor labs-on-chip (Paper 4 and Paper 5 ). Finally, an example application of theOSTE-polymers for microfluidic integration of microarrays is presented.

3.1 Thiol-ene "click" chemistryThe field of thiol-ene chemistry has during the last decades obtained a renewedinterest, due to the unique chemistry, an the versatility of thiol-ene chemistry as apolymer forming reaction in many applications.

Thiol-enes are formed by the radical initiated reaction between a multifunctionalthiol and an "ene" monomer. The "ene" stands for alkene, which is a hydrocarbonwith at least one carbon-to-carbon double bond. The unique feature of the thiol-ene systems is that they polymersise mainly through step-growth instead of tochain-growth, which is more common in radical polymerisation reactions. This isespecially important during network formation of thermoset polymers where thestep-growth meachnism will lead to high control of the polymerisation process, ahomogenous structure, a late gel-point, minimal amount of unreacted spieces, andsignificantly lower stresses than in other thermosets.

35

36 CHAPTER 3. OSTE: A NOVEL MATERIAL TOOLBOX FOR LABS-ON-CHIP

A typical thiol-ene polymer is made up of two types of monomers, one with athiol-functional group xR1-(SH)m, and the other with an "ene" functional group,yR2-(CH=CH2)n, where x and y are the number of monomers of each type andm and n are the number of functional group on each. As with all previouslydemonstrated thiol-ene systems, one normally strives to maximize the mechani-cal strength, and the ultimate goal is to have an exact equal amount of functionalgroups xm = yn, with few or no reactive groups remaining. The resulting polymerhas a non-reactive surface that requires surface treatment, e.g plasma exposure, tochange the surface properties.

Bowman and Hoyle [76] have shown that the kinetics can be varied to a largeextent by appropriate choice of "ene" monomers. One particular interesting groupof "enes" is the allyls. They polymerise uniquely via step-growth (no homopolymeri-sation via chain-growth), and have an almost complete (99.9%) degree of polymeri-sation, and minimal homopolymerisation (Figure 3.1 [76]. The reaction betweenallyls and thiols also belongs to a family of especially well-behaved lock-and-key re-actions often called "click" reactions. The term click chemistry, coined by Sharpless[77] is a class of e"cient and very selective chemical reactions that are used to joinmolecules together in a rapid manner with high yield, high purity and little or noby-products. In essence, two specific monomers are joined as easily as "clicking"together two matched buckle pieces. This is particularly important during surfacemodifications, when only one specific surface reaction is desired without unexpectedside-reactions.

To picture the network structure produced by a stoichiometric mix of multifunc-tional thiols and allyls, it could be useful to compare it with the crystal structuresof solids. For example, the ideal polymer network produced by a tri-functional thioland di-functional allyl, would have similarities with haematite (Fe2O3) shown inFigure 3.2. Each thiol monomer has three covalent bonds and each allyl monomerhas two covalent bonds.

Recently, using commercial thiol-ene glues (e.g. NOA 81, Norland Products)microfluidic devices have been fabricated with multiple layers, each of which is par-tially cured prior to lamination creating a "sticky" interface that, when re-exposedto UV-light, complete the polymerisation. The partially cured layers are useful forbonding but the the method allows little control of the thickness of the inhibitedlayer. Furthermore, it provides no discrimination between thiol or "ene" groups onthe surface and the sticky surface has limited shelf life.

3.2 OSTE: O!-stochiometry thiol-enes3.2.1 Residual activity through o!-stoichiometryIn Paper 4, o!-stoichiometric formulations of thiol-enes are introduced for the firsttime in microfluidics. These materials have an excess of one of the functional groups,xm &= ym, to achieve a polymer with remaining unreacted functional groups bothin the bulk and on the surface (Figure 3.3). The novel use of intentional large

3.2. OSTE: OFF-STOCHIOMETRY THIOL-ENES 37

Figure 3.1. Mechanism of an ideal radical thiol-ene coupling (e.g. thiol-allyls andthiol-vinyls).

Figure 3.2. Because of the pure step-growth mechanism in the thiol-ene reaction, thestructure of haematite (Fe2O3) can be qualitatively compared with to the networksproduced by tri-functional thiol monomers and di-functional allyls.


o!-stoichiometry mixing ratios in thiol-enes opens up the possibility to fine-tunethe mechanical properties and the glass transition temperature as well as producingrobust and rapid surface modification and bonding processes. Instead of a fullycross-linked network, some of the multifunctional monomers in the OSTE-polymersare not fully reacted and are only partially attached to the network, similar toFigure 3.2 but with some links free and detached from the network. This hasimmediate e!ects on the mechanical properties of the polymer. Furthermore, theunreacted groups are available to use as anchors for further click surface modificationor bonding after complete polymerisation, thus circumventing the need for breakingup the polymer surface before grafting. The unreacted groups remain stable inambient atmosphere after months of shelf-time.

Figure 3.3. Excess reactive groups using o!-stoichiometric ratios

The pure step-growth polymerisation of the thiol-allyl systems enables a highdegree of control of the number of unreacted groups available after polymerisa-tion. By choosing the appropriate multifunctional monomers, surprisingly largeo!-stoichiometric ratios, up to 100% excess of either thiol or allyl functionality canbe achieved without losing mechanical stability and without excessive leakage ofnon-crosslinked material.

Under the assumption of total mobility of all the monomers and equal reactivityof all the functional groups, the minimum number of monomers that has no bondsto the network can be approximated by multiplying the probability that each of thefunctional groups on the monomer in excess will not find a corresponding partner.For instance, in a polymer composed of tetrathiol and triallyl, with 50% excessof thiol groups (xm/yn = 1.5), the probability is (1/3)4 = 1.3% that a tetrathiolmonomer will not attach to the network and be leachable. Lower functionality ofthe monomer in excess increases the leaching. Figure 3.4 shows the theoreticalamount of non-crosslinked monomer for the two dual monomer systems used inPaper 4 and Paper 5 together with the experimentally determined leakage for thesame systems. The leaching from both systems is slightly higher than the minimumtheoretical value which is due to the fact that not only monomers but also oligomersand initiators that are not attached to the network can be leached. This increasesthe amount of extractable material.


Figure 3.4. Theoretical leaching of non-crosslinked components from the OSTE-polymers as a function of o!-stoichiometry ratio. The line represents excess ofmonomers with n=4 functional groups and the dotted line with n=3 functional groups.The experimental results for OSTE Allyl (30) (n=3) and OSTE Thiol (90) (n=4) isalso plotted.

3.2.2 Tuneable mechanical properties

The advantage of the OSTE-polymers over standard thiol-ene polymers is that themechanical properties can be fine tuned to the exact demands of the applicationby adjusting the o!-stoichiometric ratio, without changing the type of monomers.In o!-stoichiometry, the monomers have fewer cross-links to the network, whichin turn a!ects the E-modulus and the glass transition temperature. Mechanicalproperties have been demonstrated ranging from harder than PMMA to soft likePDMS, using o!-stoichiometry of thiol groups, and glass transition temperaturesTg, ranging from below 30 !C to 84 !C (Figure 3.5 left).

Particularly useful is the ability to use heat to tune the sti!ness of the polymers,which is explained further in Paper 8. By choosing the glass transition temperatureslightly above room temperature, the mechanical properties of the OSTE-polymerscan be tuned in a narrow temperature interval so that the polymer rapidly trans-forms from a glassy material (E-modulus in the GPa range) at room-temperature,to a rubbery material (E-modulus in the MPa range) at a temperature only a few


tens of degrees higher. For example, the OSTE-polymer with 90% excess describedin Paper 4 remains medium sti! (200 MPa) up until 30 !C, when it suddenly startsto soften and is ten times softer at 50 !C (! T = 20 !C). This is also the case forOSTE Thiol (70), which is developed with biocompatible bonding in mind (Figure3.5 right) and which softens at Tg = 37 !C.

Figure 3.5. Left: E-modulus and glass transition temperature of the OSTE Thiol asthe o!-stoichiometric mixing ratio is varied. Right and bottom photos: Temperaturetuning of the OSTE Thiol (70). Heating to Tg allows for perfect sealing duringbonding.

3.2.3 Direct patternable surface modificationThe OSTE-polymers provide anchors on their surfaces that can be used to chemi-cally link or graft functional groups or polymer chains using UV-initiated thiol-eneclick chemistry. With the OSTE-polymers, the surface coverage can be controlledby the ratio of o!-stoichiometry. At 5% o!-stoichiometry, macroscopic surfaceproperties such as contact angle, can be modified when grafting hydrophilic PEGmonomers. At higher o!-stoichiometry ratios the PEG grafting results in an evenlarger modification of the contact angle up to a threshold value, when steric e!ectslimits the surface coverage. The direct modification is an advantage compared tothermoplastics and PDMS, where the polymer surface first must be broken up toexpose reactive groups, for example by plasma treatment and the density of activesites, as well as their surface homogeneity, is less controlled. An additional advan-


tage is that the grafting of functional groups can be UV-patterned using a stencilmask, creating areas with di!erent surface properties on a single chip, for examplehydrophobic and hydrophilic areas, as shown in Figure 3.6.

Figure 3.6. Principle of direct UV-grafting on the OSTE-polymers

3.2.4 Biocompatible low-temperature bondingThe OSTE-polymers can form covalent bonds directly with other materials, usingthe anchors of unreacted groups available on the surface. In contrast to solventbonding and tape bonding, no leachable compounds that reduce bond integrity re-main in the bond area. In one type of bond, the OSTE-polymer is directly covalentlybonded to another OSTE-polymer with the opposite type of anchors exposed. Forexample, an OSTE-polymer with an excess of thiol groups bonds covalently withan OSTE-polymer with an excess of allyls groups after unfiltered UV-exposure. Alower wavelength (< 250 nm) [78] is required for the reaction to occur since noinitiator is available at the interface. Another type of bonding situation occurswhen the OSTE-polymer is bonded to a surface with a coating capable of reactingwith thiols or allyls. This can for example be a surface of isocyanate (a commoncoating on sensors) or vinyl silane (Paper 5 ), a gold surface [79] (a common sensorsurface) or activated esters (NHS) (Paper 8 ). The particular advantage with theOSTE- polymers compared to thermoplastics in bonding, is the temperature tuning


as described above. When heated above Tg, the OSTE-polymers conform perfectlyto micro-irregularities on the substrate and can react with a very high yield andform strong bonds (Figure 3.7).

Figure 3.7. Principle of bonding the OSTE-polymers at their glass transition tem-perature.

3.2.5 Low absorption of moleculesThe dense network of the OSTE-polymers results in a significantly lower absorptionof small molecules than in PDMS. The low absorption enables handling of analytesin very small concentrations, as very little of the sample is lost into the channelmaterial. In Figure 3.8, the absorption of Rhodamine B in OSTE and PDMS wascompared after 24 h of exposure. Only small amounts are adsorbed at the wall of themicrochannel in the OSTE-polymer wheras a considerable amount of RhodamineB has been absorbed by the PDMS.

Figure 3.8. Comparison of absorption of small molecules (Rhodamine B) betweenOSTE and PDMS after 24 hours of exposure. No di!usion is observed in the OSTEsample. In the PDMS sample a large portion of the Rhodamine B has di!used intothe walls.

3.2.6 Solvent resistant channelsSolvent resistance is an important feature in particular for applications in chemicalanalysis and separation. The OSTE polymers based on the monomers used in Pa-per 4, 5 and 8 are compatible with common solvents such as toluene, iso-propanol,

3.3. FACILE INTEGRATION OF MICROFLUIDICS WITH MICROARRAYS: THEBIOSTICKER 43

ethanol, methanol and glycerol. As seen in table 3.1, an OSTE-polymer microchan-nel bonded to silicon showed good compatibility with alcohols and toluene, butsolvents with large dipole moments, i.e. DMSO and acetone was not well tolerated.To increase the compatibility with these solvents, the monomers must be replaced.

Table 3.1. Solvent compatibility of the OSTE-polymer using the monomer compo-sition from Paper 5.

Solvent After 24 hoursIsopropanol No visible e!ectMethanol No visible e!ectAcetone Bulk material failureToluene No visible e!ectGlycerol No visible e!ectDMSO Bulk material failureDI water No visible e!ect

3.2.7 A rapid and scalable manufacturing processThe OSTE-polymers can be casted on the same type of masters as PDMS, andpolymerised in seconds using light from a standard table top UV-lamp (365 nm).During demolding, the master can be heated to the glass transition temperatureto facilitate the release of the polymer. Very sti! OSTE polymers, with an E-modulus close to 2 GPa, was successfully demolded from SU-8 masters using thistechnique. After demolding, the chip can be cut using the same trick to heat it tothe glass transition temperature during processing or diced at room temperatureusing an automated dicing machine. Moreover, due to the low built-in stresses,the OSTE material lends itself well to machining and milling producing smoothedges without cracks. The low shrinkage of the OSTE-polymers also allows forwafer-level integration of silicon substrates (Paper 5 ), also shown in Figure 3.9,which significantly decreases the back-end process time and enables integrationwith CMOS electronics [80].

3.3 Facile integration of microfluidics with microarrays:the Biosticker

3.3.1 Introduction to microarraysMicroarrays have become powerful tools in biochemical analysis, primarily becauseof their capability for highly multiplexed analysis.

A microarray consists of a solid surface with a dense matrix of receptor ligands,probes, deposited in certain volumes by a liquid dispensing robot, producing "spots"


Figure 3.9. The simultaneous manufacturing of multiple bonded chips reduces theback-end process time.

on the surface. Briefly, a sample containing the analyte, target, is loaded on thearray for specific capture to a matching probe. Next, a secondary a"nity reagent,usually tagged with a fluorescent dye, is added to visualise which spots were reactingwith the analyte.

Microarrays are used in a number of growing research areas within the life sci-ences related to DNA and protein analysis. Everyday users of microarrays are foundin universities, medical companies, hospitals and central laboratories (e.g clinicaldiagnostics or cancer diagnostics).

Microarrays are divided into two groups: DNA-arrays and protein-arrays.

DNA microarrays

In DNA-arrays, DNA molecules are deposited on the surface and used for specificcapture, hybridisation, of complementary strands in the sample. DNA-arrays areprimarily used for monitoring gene expressions, but other applications such as single-nucleotide polymorphism (SNP) genotyping, where anomalies in the DNA strandsare detected, have also been demonstrated [81]. The use of DNA-arrays has becomeubiquitous in biochemical research and DNA-arrays are the largest segment in themicroarray market.

Protein microarrays

A protein microarray comprises many di!erent types of probes (typically antibodiesor antigens) that are deposited on the surface. Each probe captures its targetprotein from a sample, typically a serum or cell lysate, and the captured proteinsare subsequently detected and quantified using a"nity reagents [82]. Protein arraysare used primarily in proteomics, the study of the function and structure of proteins,and recently also in diagnostics [83, 84]

3.3.2 Mass-transport limitation in microarraysThe performance of microarrays depends on the conditions of the assay, for instancethe bu!er and temperature, and on the substrate surface on which the probes are

3.3. FACILE INTEGRATION OF MICROFLUIDICS WITH MICROARRAYS: THEBIOSTICKER 45

immobilised [23]. Another key factor for the performance of microarrays is the masstransport of analyte down to the spots. Normal incubation times for microarrays,which relies solely on di!usion for sample transport, is several hours up to days [85].Because the di!usion coe"cient, D, for nucleic acids in aqueous solutions is on theorder of 10"11 m2s"1 [86], the typical distance traversed soley by di!usion, L =

"Dt,

in 24 h is 1 mm. Considering that the horizontal length scale of a microarray is onthe order of a few centimeters, hemispherical depletion volumes will form aroundeach probe and the e"ciency of the binding will be very low [59]. These e!ects aresimilar or worse in the case for protein microarrays since proteins vary in shape andsize, resulting in di!erent di!usion speeds for di!erent analytes. Thus, there is astrong demand for technologies enabling more e"cient mass transport and reliablemixing to avoid sample depletion over the spots and break the di!usion limitation.

Microfluidic integration

The integration of microfluidics with microarrays provides better mass transport byadding controlled convection over the microarray to avoid depletion of the sampleand circumvent the di!usion limitation. Microfluidics also o!ers the advantage ofmulti-sample capabilities on one single chip, which may be of importance in geneticmutation analysis or clinical diagnosis, where direct comparison between di!erentsamples on the same chip would be preferable because the quality of slides withprobe arrays varies from batch to batch [87]. Cross talk between di!erent spots isalso eliminated by separating the sample in compartments [88].

Microfluidics have previously been integrated with microarrays, and assay timeshave been significantly reduced. In some cases a more than thirty-fold reductionin assay time has been achieved [59, 89]. However, previous integration methodsutilized either a clamped or plasma bonded PDMS channel layer on a spotted sub-strate of bare silicon/glass/PDMS [90, 91], double sided adhesive film [92, 93] or aplastic foil covering a thermoplastic substrate containing spots in custom-machinedchannels [94, 95]. All these methods are limited in practice, either because of poormaterial properties (adsorption of small molecules, channel deformation or tapedissolving monomers), complicated assembly processes, or because of limitation tocustom-made plastic substrates.

3.3.3 The Biosticker: a micropatterned OSTE-sticker for microarraysThe Biosticker, developed in Paper 8, is a microfluidic add-on for microarrays, thatcan be bonded to almost any substrate available on the market. It was developedbased on the o!-stoichiometry thiol-ene (OSTE) polymer platform presented earlier.By making use of the narrow glass transition temperature, the active surface, andthe excellent micromolding capability of the OSTE polymer, integration, handlingand bonding are greatly simplified.

The probes are typically immobilised to the microarray surface via an epoxy-silane, amino-silane, lysine or polyacrylamide linker. These linkers can also be


Figure 3.10. A Biosticker flow cell attached to a protein microarray. The spotsare visible though the polymer. To guarantee an even flow profile over the spots,branched inlet and outlet channels are used

used for attaching the microfluidic flow cell. The OSTE-polymer has free thiolgroups on its surface that can either directly covalently react with many standardmicroarray surfaces, or be designed to react via secondary functionalization (e.g.epoxyallyl or amineallyl monomers) with virtually any microarray surface. Heatingthe Biosticker to 37 !C softens it enough for easy application to the microarraysurface, where it seals conformally and reacts with the micoarray surface. Thebonded fluidic layer protects the microarray during handling and the assay. TheBiosticker can be removed after the assay is completed by heating above its glasstransition temperature Tg=37 !C, when the Biosticker softens and the interface tothe surface is easily broken. This allows for read-out of the results in standardfluorescent scanners.

3.3.4 Preliminary resultsThe feasibility of the Biosticker concept was tested on a high performance three-dimensional microarray surface developed by Prof. Marcella Chiari and her group atIstituto di Chimica del Riconoscimento Molecolare (ICRM), C.N.R, in Milano, Italy.This particular surface consists of a copoly(DMA-NAS-MAPS) polymer receptorlinker layer, recently demonstrated to improve sensitivity and limit-of-detection[96]. To guarantee an even flow, with homogenous concentration of the analyte overall the spots on the microarray surface, a branched microfluidic layout was designedfor the Biosticker (Figure 3.10).

With the help of Dr. Marina Cretich and Dr. Laura Sola, also at ICRM, twobioassays were tested with the Biosticker, a fluorescent protein experiment withspotted !-lactoglobulin detecting 1 ng/ml of anti !-lactoglobulin antibody, and aDNA hybridization test using spotted 23 mer 5’- amine modified oligonucleotides


and target complemetary oligonucleotide (1 µM) (Figure 3.11). For a first compari-son, similar incubation times as for static experiments were used. The result showeduncomplicated application, no leakage and excellent signal and spot homogeneity,demonstrating the potential for using the Biostickers to optimise microarray as-says while avoiding the need for special tools, complicated clamping, lamination orsub-optimal materials.

Figure 3.11. The results of the scanned microarrays. The results from the "-lactoglobulin protein assay and the DNA hybridization are very promising and showhomogenous spots and excellent intensity.

At the time of the writing of this thesis, the assay protocol and the channelgeometry are optimised to achieve a reduce assay time, while keeping the sameuncomplicated handling, high sensitivity and specificity as already demonstrated.

3.4 Summary and outlook

In this chapter, a novel polymer platform, based on o!-stoichiometry thiol-enes(OSTE) was introduced and characterised. A comparison with the wish-list inthe end of Chapter 1, shows that the OSTE-polymer addresses most of the listedcriteria for an ideal prototyping system for labs-on-chip. In Table 3.2 below, and inand Table 1 in the Appendix, the OSTE-polymer is compared with other materialscurrently available for microfluidic components prototyping.

The OSTE-polymer platform is in many respects a better alternative thanPDMS for rapid prototyping of labs-on-chip devices. The OSTE-polymers do notsu!er from the adverse properties of PMDS, but are processed in an inexpensiveand uncomplicated process, which is very similar to soft lithography. OSTEs signif-icantly simplify the back-end processes by using built in anchors for surface func-tionalisation and biocompatible bonding. By also being able to mirror many of theproperties of commercial thermoplastics the OSTE-polymers potentially minimisethe redevelopment required to transfer a research proof-of-concept into a commercialthermoplastic device.


Table 3.2. Comparison of how common materials for lab-on-chip, as well as thenovel OSTE polymer, prototyping compare to important material and processingproperties for rapid prototyping of labs-on-chip

OST

EPD

MS

Ther

mop

last

ics

NO

A81

Gla

ss

Mat

eria

l 1) Tuneable mechanical proeprties (sti! to elastomer) x - - - -2) Chemical inertness, low interaction with sample x - x x x3) Solvent resistance x - - x x4) Direct, patternable and stable surface modifications x - - - -

Proc

ess 6) Three-dimensional microfluidics - x x - -

7) Fast, scalable and inexpensive x - x - -8) Biocompatible bonding x x x x x

3.5 Future workThe development of the OSTE-polymer is hopefully the first step to an even morepowerful and versatile polymer platform for labs-on-chip prototyping and produc-tion.

There are still problems that need to be addressed. Most significant is maybe theleakage of non-crosslinked monomers in the materials with high o!-stoichiometricratio. Moreover, high temperature applications such as PCR requires a higher Tg

than currently achievable in OSTE. However, both of these problems could be solvedusing ternary systems where a third monomer could react with the excess of thethiol [97], using a two step curing process [76]. First the thiol and allyl monomersare polymerised as previously, leaving a polymer with excess of thiol groups that canbe used for surface functionalisation. Secondly, after lamination of the patternedfluidic layers, the thiol excess is reacted with the third monomer, using an appro-priate initiator. This would minimise the leaching but still allow surface function-alisation before initiation of the second polymerisation. Although 3D microfluidicdevices with interconnecting via was not shown in the OSTE-polymer in this work,it has been demonstrated in preliminary laboratory experiments and with similarphotocurable stoichiometric thiol-ene polymers [98]. For a truly seamless transi-tion from a research proof-of-concept to a commercial prototype, the possibilities ofadapting the OSTE-polymer for commercial mass production must also be inves-tigated, for instance its compatibility with injection moulding. All of the above iscurrently under development at KTH Microsystem Technology.

Chapter 4

Introduction to low fluid-frictionsurfaces

In this work, surface e!ects in liquid flows over solid surfaces is studied. It isshown, that by microstructuring the solid surface, microscopic surface e!ects canbe manipulated and enhanced to have an impact on the macroscale flow patterns.The concept of super lubricating surfaces is introduced in this chapter, while mywork on improving the stability of these surfaces for flow conditions encountered inrealistic applications is presented in Chapter 5.

4.1 MotivationDrag, or fluid resistance, in liquid flows is the critical limiting factor in many micro-and macroscale applications, such as: lubrication, energy conversion, marine propul-sion, flow switching, chemical separation and mixing on lab-on-chips. In a macro-scopic flow, the drag has two components: the turbulent and the laminar losses.The laminar component is due to surface friction and is directly a!ected by thesurface properties. In microchannels, where the flow is completely laminar, thelaminar drag is the only source of energy loss. Furthermore, since the surface tovolume ratio is large, surface friction becomes a major limitation for throughputin micro and nanofluidic devices. Also on the macroscale, there exist a thin layerat the solid-liquid interface, called the laminar sub-boundary layer, where the flowis laminar, as illustrated in Figure 4.1. The drag from this layer could be as highas 60-70 % of the total drag for large sea vessels [99]. The goal of this work wasto develop e!ective means to reduce and manipulate the laminar drag using micro-machined surfaces. The experiments, performed in microchannels, could also beapplied to the laminar sub-boundary layer of macroscale flows.

Research on super-lubricating surfaces, also known as superhydrophobic surfaces,is based on two approaches: the approach using self-assembled surface coatings andthe approach using micro-structured surfaces. This work uses the latter approach.Self-assembled coatings have many advantages, such as application to a wide number

49

50 CHAPTER 4. INTRODUCTION TO LOW FLUID-FRICTION SURFACES

Figure 4.1. Left: illustration of a macroscopic flow, with a thin laminar sub-boundary layer. Right: a microchannel with laminar flow.

of materials and uncomplicated manufacturing. Micro-structuring, however, allowsfor greater control and, as we will see, more friction reduction. Recently, micro- andnanostructured surfaces with close to zero surface friction were fabricated [100].Nevertheless, there are still technological hurdles remaining, before these super-lubricating surfaces can be used in practical and commercial applications.

First, the robustness dilemma must be solved. The lubricating state of thesurfaces is fragile and collapses already at low liquid pressures. This makes thesurfaces di"cult to use, in particular since a collapsed superhydrophobic must bedried out before it can be reused, which is the second problem. Moroever, a collapsedsuperhydropohbic surface will quickly be fouled and loose its hydrophobicity. Paper6 and Paper 7 of this thesis introduce two potential solutions to these limitations;robustness, to counter collapse at liquid pressure levels encountered in practicalapplications, and active control that can actively restore and collapsed surfaces toa super-lubricating state.

4.2 Surface friction in liquid flowsIn fluid mechanics, zero-slip is the universally accepted boundary condition, at asolid-liquid boundary. It postulates that the fluid is stationary at the solid surface,because of high friction at the interface. In reality, this condition has proven untrue,but for most practical applications the surface velocity is so small that it can safelybe ignored. For the situations in which the zero slip boundary condition does nothold, the slip boundary condition model was introduced. It relates the fluid velocityat the surfaceÚ u0, the slip velocity, to the magnitude of the shear rate experiencedby the fluid at the surface:

u0 = $

!!!!)u

)y

!!!! , (4.1)

where $ is called the slip length. Its geometrical interpretation is illustrated inFigure 4.2.

4.3. SUPERHYDROPHOBIC SURFACES 51

Figure 4.2. Illustration of the slip length # at a solid-liquid interface.

The slip length is a practical measure to compare the friction reducing capacity ofdi!erent surfaces. However, due to di!erences in measurement setups, the reportedslip length on similar surfaces may vary. Smooth hydrophobic surface, for example,have a reported slip lengths of zero to a few nanometers [101, 102]. Is that enoughto significantly a!ect the total fluid resistance in a microchannel? By using a simplemodel of a pressure-driven flow between two infinite parallel plates separated by adistance H, where one surface supports slip, a relation between the volume flowrate per unit depth q, and the slip length $ can be derived:

q = H3

4µ

"' dp

dx

# $13 + $

$ + H

%. (4.2)

From the above equation, it can be seen that for a given pressure gradient dp/dx,and fluid viscosity µ, the volume flow rate q can be significantly enhanced only ifthe slip length $ on the same order of magnitude as the height of the channel H.

However, on a smooth hydrophobic surface, such as Teflon, the slip length isonly a few nanometers. This seems discouraging, since typical microchannels inlab-on-chip applications are at least a few micrometers in height, and the laminarsub-boundary layer in turbulent flows can be up to several milimeters [103]. Howcan surfaces be made more hydrophobic than Teflon?

4.3 Superhydrophobic surfaces4.3.1 Mechanism of operationSuperhydrophobic surfaces were originally inspired by the water-repelling featuresof plants, such as the lotus leaf. As shown in Figure 4.3, lotus leaves have a roughhydrophobic surface that is able to trap air between itself and a liquid, to create alubricating air cushion, very much like an inverted hoover craft.

As shown in Figure 4.3, the liquid on a superhydrophobic surface can be in twostates. It can fully wet and penetrate the roughness (Wenzel state, no lubrication)or it can rest on top of the roughness, suspended by surface tension (Cassie or fakir

52 CHAPTER 4. INTRODUCTION TO LOW FLUID-FRICTION SURFACES

Figure 4.3. A superhydrophobic surface. Left: Water droplets on a lotus leaf cannot wet the rough hydrophobic surface and experience extremely low friction. Theinset shows a close up of the surface of the lotus leaf. Right: Schematic figure of asuperhydrophobic surface. The water is suspended on air between the solid fractionsand the water can accelerate, on top of the air pockets, to a finite slip velocity beforeit reaches the next solid area where it is slowed down to zero.

state, air lubrication). In the latter state, the liquid will experience zero frictionon top of the air-pockets created in the rough surface and accelerate to a finite slipvelocity. However, at the solid-liquid contact points, the liquid will be slowed downby friction. On average, this gives rise to an e!ective slip velocity and an e!ec-tive slip length over the entire surface. The specific geometry of superhydrophobicsurfaces can vary: posts, ribs, spikes or unstructured. In this work, we use ribsoriented perpendicular to the flow, as they are robust in the sense that if one air-pocket collapses, it will only fill a small section of the surface. Moreover, they areuncomplicated to fabricate with high precision using standard UV-lithography. Ina parallel plate flow, the e!ective slip length on such a surface was predicted byLauga and Stone [104]:

$ = L

2*ln

" 1cos(Fc*/2)

#, (4.3)

where the air fraction (shear free fraction) Fc is expressed as Fc = a/L, i.e. thegas-liquid interface length between two ribs a divided by the pitch of the ribs L.From this equation, which has only two factors to play with, we see that the airfraction Fc must be maximised to reduce drag, but beyond a certain point (Fc ( 1),the slip can only be made larger by increasing the pitch L. This is why the sliplength on nano-structured surfaces will be limited (Fc high but L low).

Luckily, it is practically possible to engineer superhydrophobic surfaces with sliplengths reaching several micrometers [105, 106, 107, 108, 109, 110]. Although themeasurement techniques and measured results can vary for similar surfaces, the slipon a microstructured surface can reach hundreds of micrometers under optimisedconditions [100]. The next section describes in more detail how much the frictioncan be lowered in microchannels using these surfaces — and to what cost.

4.3. SUPERHYDROPHOBIC SURFACES 53

4.3.2 Stability limitationsThere is a fundamental limitation to all superhydrophobic surfaces. At some point,when the air pockets are made larger and larger, the surface tension will not bestrong enough to support the liquid meniscus. The liquid will then penetrate intothe superhydrophobic surface, and ruin the lubricating properties. This limitationis described by the Young-Laplace equation, which relates the maximum sustainablestatic pressure drop over the air-liquid interface, !Pmax, to the surface tension andgeometry of the air pocket. In the case of parallel ridges, the maximum interfacialpressure is

!Pmax = PL ' PG = 2+ cos(%)a

, (4.4)

where PL is the liquid pressure, PG is the air pocket pressure, + is the surfacetension, % is the contact angle of the smooth hydrophobic material and a is the airpocket cavity size.

Because the air pockets are isolated from the surroundings and their air pressureis fixed to ambient conditions during priming, this introduces a dilemma: large airpocket cavities a are needed to generate large slip, in order to have a large e!ect onthe frictional losses of the flow, but large air pockets collapse easily. For instance,microfluidic channels are typically around 10 µm in height and if the air pocketssize is 5 µm, the slip will be enough to a!ect the flow (4.2). However, it will onlysupport a liquid pressure up to 15 kPa (4.4), which is less than encountered in mostpractical applications. This reduces the advantage of using a superhydrophobicsurface, since the flow rate is also limited by the pressure.

Chapter 5

Novel robust super-lubricating surfaces

With the understanding of the mechanism and limitations of superhydropohobicsurfaces from Chapter 4, this chapter introduces two ways to circumvent the Young-Laplace limitation and allow sustained lubrication at high pressures and large flows.The main theme is to pneumatically connect the air pockets and actively or passivelycounter the increased interfacial pressure drop at high liquid pressures.

In the first section, a model is presented for the practically achievable friction re-duction in a superhydrophobic channel limited by the Young-Laplace collapse pres-sure. In the second and third sections, previous attempts to increase the robustnessand to manipulate the lubrication of superhydrophobic surfaces are discussed. Thefourth section presents my work on a simple active external regulation of the airpocket pressure (Paper 7 ) to allow switching between superhydorphobic states, anda mechanism for passive adaptation of the air pocket pressure to the liquid pres-sure by a moving piston e!ect (Paper 6 ). Both were first-time demonstrations ofsuperhydrophobic surfaces stable beyond the theoretical Young-Laplace pressure.

5.1 A model for friction reduction in a microchannelTo explore the robustness dilemma, an expression for the total friction , in a su-perhydrophobic microfluidic channel, relative to that of a smooth channel, limitedby the Young-Laplace collapse pressure was derived in Paper 6 :

, )$12ln(1/ cos(Fc*/2))

*WF Fc+ 1

%"1. (5.1)

This expression is valid for parallel plate flows using a superhydrophobic surfacewith ribs perpendicular to the flow. It contains only two parameters, the air fractionof the surface Fc and a novel dimensionless number WF , also introduced in Paper 6,dubbed the channel’s energy carrying capacity;

WF = PLDh

+ cos(%) . (5.2)

55

56 CHAPTER 5. NOVEL ROBUST SUPER-LUBRICATING SURFACES

Here, Dh is the generalised hydraulic diameter of an arbitrary cross-section, de-fined as Dh = 4A/P , where A is the cross-sectional area and P the wetted perimeter.Whereas the slip length can only describe the friction reducing properties of the sur-face, WF can be thought of as a figure of merit for the complete superhydrophobicflow system and makes it possible to compare the flow capacity of di!erent flowsystems that have di!erent geometries and surface energies. The higher WF , thehigher is the capacity for large flows and highly pressurised liquids (large Dh andPL). In a sense WF , is the energy carrying capacity of the flow.

Figure 5.1. Graph of the relative friction factor $ as a function of the dimensionlessfigure of merit WF for some values of the shear free fraction Fc. Increased performanceof the channel, that is, higher friction reduction and higher flows, is achieved bymoving toward the bottom right of the figure. For each fixed value of Fc, there is aminimum obtainable surface friction factor $ uniquely expressed by the dimensionlessparameter WF . The area to the right of each line represents the forbidden values,where friction reduction is not allowed because of collapse of the superhydrophobicsurface.

The minimum channel friction ,, for a few di!erent values of air fractions Fc,is plotted as lines in Figure 5.1 using (5.1). This figure visualises the limitationsof friction reduction due to air pocket collapse. The area to the left of each curverepresents possible combinations of channel geometry, pressure and surface energythat are stable with regard to surface tension. Points to the right side of eachFc curve represent configurations exceeding the Young-Laplace pressure limitationwhere the flow system will be unstable and loose its superhydrophobic properties.

5.2. FRACTAL SURFACES: TEMPORARY LIFE SUPPORT 57

The black area in the bottom right corner, limited by Fc = 1 (only air, no solid)represents the ultimate goal — extreme lubrication with highly pressured liquidsand large flows. The plot gives a good idea of what is practically possible whendesigning a superhydrophobic flow system.

5.2 Fractal surfaces: temporary life supportPreviously, the way to avoid the complete loss of super-lubricating properties abovethe Young-Laplace collapse pressure was to create multiple layers of air pockets, i.e.a surface with a roughness on multiple length scales. This is a well tested strategyalso used by many water repelling leaves. When examining a lotus leaf (Figure 4.2)in a microscope, several levels of roughness can be observed. The largest roughnesswill trap the biggest air pockets, but when it collapses air will still remain in thesmaller air pockets, which are more robust to higher pressures. This can go onseveral levels as the pressure increases. In this way the lubrication properties arenever completely lost, only reduced step by step. This ideas was used by Onda etal. [111] when they presented a fractal superhydrophobic surface made in wax.

Recently however, Lee et al. [100] demonstrated that fractal surfaces, contraryto popular belief, actually can reduce the lubcrication in cases when the air fractionFc of the largest roughness scale already is large. Adding a second roughness scalewill smoothen the sharp tips of the solid fractions and allow the liquid interfaceto recede slightly into the cavity of the largest roughness, which will decrease thelubrication. Thus, fractal surfaces can limit the e!ect of a collapse, but cannotprevent it and in extreme cases also reduce the lubrication.

5.3 Active switching: wet or dryOnce a superhydrophobic surface is wetted, the liquid must somehow be expelledfrom the liquid filled pockets before it can regain its super-lubricating properties.Although there are many ways to change the wettability of a surface, such as thermo-sensitive polymers [112], pH switchable surfaces [113], roughness switching of asurface [114] or optical switching [115, 116], these techniques can only transform adry super-lubricating state into a wetted, non-lubricating state. Not vice versa.

The other direction, from wetted state to a dry state, has also been separatelydemonstrated [117, 118, 106]. Here, larger amounts of external energy must beprovided to move the liquid out of the pockets and dry them. For instance, a shortburst of electrical current through wetted droplets on a superhydrophobic surfacewill evaporate the water in the cavities [117] and push up the droplets. Acousticshas also been used to shoot wetted droplets o! lotus leaves [118] and very recentlyalso dewetting using electrolytically generated gas bubbles [106] was shoen.

There have been no demonstrations combining these two e!ects to actively ma-nipulate the states of a superhydrophobic surface between wet and dry in a contin-uous flow.


5.4 Regulating the air pocket pressure to avoid collapse

Figure 5.2. Experimental setup for active regulation of the air pockets. Two parallelchannels on top of a superhydorphobic surface with perpendicular gratings. The airpockets formed in the grooves are connected to the air channel through the grating.In a first experiment the same driving pressure was connected to both channels, thuscreating a zero pressure drop over the air liquid interface. In a second experiment,the air channel pressure was manipulated enabling switching from dry to wet statesin the liquid channel.

5.4.1 Active regulationThe first solution for circumventing the Young-Laplace collapse limitation, pre-sented in Paper 7, pneumatically connects the air pockets in a superhydrophobicsurface to the same pressure source as the the liquid driving pressure. Two flowchannels were placed perpendicular to a ribbed superhydrophobic surface, in sucha way that the channels could communicate with each other, through the groovesin the surface (Figure 5.2). Water was flowed in one channel and air was flowed inthe other. As both air and liquid flows were connected to the same upstream pres-sure source, the pressure drop over the air-liquid interface, PL ' PG, in the liquidchannel could be kept constant even as the liquid pressure was increased above thetheoretical Young-Laplace collapse pressure for the specific surface geometry. Thisactive regulation of the air pocket pressure, prevents collapse of the air pockets.The principle is presented in Figure 5.2.

5.4. REGULATING THE AIR POCKET PRESSURE TO AVOID COLLAPSE 59

In theory there is no limit to the liquid pressure such a system can handle, butat elevated upstream pressures the air-pockets may burst if the pressure in the gasfilled channel exceeds the pressure in the liquid channel. The bursted air pocketsintroduces a small bubble in the liquid flow that disrupts the measurements.

In another configuration, the air channel was connected to a computer controlledpressure source, while the upstream liquid pressure was connected to a separatepressure source. As the air pressure in the air control channel was manipulated, theair-liquid interface could be retracted into the deep grooves (wet state) or restoredto full lubrication (dry state). The active switching of the surface friction enabledmanipulation of the liquid flow rate through the channel while keeping a constantupstream liquid pressure.

5.4.2 Self-regulating air pockets

Figure 5.3. The princple behind the self-regulating superhydrophobic design. Theair pockets are pneumatically connected to the liquid through feedback channels.When liquid enters the feedback channels the liquid pressure compresses the air be-hind the grating and lowers the pressure drop over the interface.

The second solution, presented in Paper 6, pneumatically connects the air pock-ets in the superhydrophobic surface to the bulk liquid pressure through feedbackchannels, as illustrated in Figure 5.3. When the liquid pressure increases, some wa-ter will enter feedback channels, which, like a piston, will compress the air in the airpockets, increasing the air pressure and reducing the interfacial pressure drop overthe air-liquid interface. In this way, liquid pressure exceeding the Young-Laplacecollapse pressure for the pattern can be tolerated without loss of super-lubricationand without the need for external pressure control.


Figure 5.4. The performance gain of active regulation and self-regulation of theair pocket pressure compared to previous superhydrophobic surfaces with isolated airpockets. The active regulation can in theory balance much higher pressures but athigh pressures air bubbles are easily introduced into the liquid flow channel. Theself-regulating design is ultimately limited by air di!usion from the air pockets intothe liquid.

5.4.3 PerformanceActive regulation

The active configuration was tested to up to three times higher liquid pressure thanthe collapse pressure of an identical superhydrophobic surface, with isolated airpockets. Figure 5.4 illustrates to performance gained by pneumatically connectingthe air pockets to an extrenal pressure source. The novel surfaces, with regulatedair pockets, are compared to reference superhydrophobic surfaces with isolated airpockets, in terms of maximum achievable energy carrying capacity WF , at fixedrelative channel friction ,. The novel designs with regulated air pockets, can moveinto the theoretically forbidden area, with higher pressures and larger flows, thanpreviously possible.

In Figure 5.5, the feasibility of the active switching configuration is demonstratedby actively switching the superhydrophobic state of the surface. In this example,


the mass flow manipulation capacity was only 5%, due to the the non-optimisedgeometry and low air fraction.

Figure 5.5. Switching of the flow rate in a superhydrophobic channel by manipulat-ing the air pocket pressure to either retract or push back the air-liquid interface intothe deep pockets. At air channel pressures under the Young-Laplace pressure, thewater is sucked all the way into the air filled channel, through the connecting grooves,and introduces liquid droplets. These droplets can not be removed and terminate thefunction of the device.

Self-regulating

The self-regulating design was likewise abel to withstand roughly three times theYoung-Laplace collapse pressure for an identical superhydrophobic surface with-out the feed back mechanism, as illustrated in Figure 5.4. However, the feedbackmechanism is ultimately limited by the gas exchange over the air-liquid interfaceand at elevated air pressures gas starts dissolving into the liquid and slowly emp-ties the feedback channels, as illustrated by Figure 5.6, in which the life-time ofsuperhydrophobicity is plotted as a function of relative over-pressure, above theYoung-Laplace collapse pressure.

5.5 Summary and outlookTwo approaches were demonstrated to increase the robustness of superhydrophobicsurfaces and, for the first time, allow sustained friction reduction at liquid pressuresexceeding the Young-Laplace collapse pressure. The first principle pneumaticallyconnected the air pockets to an external pressure source, which not only allows highliquid pressure with sustained lubricating properties, but also, for the first time, ac-tive switching between lubricating and non-lubricating states in a liquid flow. Thesecond principle regulated the pressure of the air pockets by pneumatically connect-ing them to the liquid flow pressure. This self-regulating configuration required noexternal actuation. However, due to the di!usion of air into the liquid at elevated


Figure 5.6. Life time of the super-lubricating property at pressures exceeding theYoung-Laplace collapse pressure. At prolonged times exceeding the Young-Laplacepressure, air di!usion over the interface empties the air pockets.

pressures, super-lubrication at high liquid pressures and large flows could only besustained for a limited time.

The work represent the first steps towards constructing superhydrophobic sur-faces with high lubrication, that can also be applied at liquid pressure levels en-countered in real life. Later, the self-regulating principle has been further developedand refined in the work by Lee et al. [106], where air is electrolytically generated inthe air pockets to stabilise the air pressure for longer time periods at high pressure.

Chapter 6

Conclusions

This thesis presented four novel concepts in microfluidics related to the developmentof new materials, surfaces and manufacturing methods.

The following two novel manufacturing improvements of PDMS prototyping forlab-on-chip devices were introduced and demonstrated:

1. The first adhesive packaging method based on dual-surface adhesivefilms to bond substrates with di!erent surface energies.The novel concept of dual surface-energy adhesive films to bond substrateswith di!erent surface energies enabled bonding of a microfluidic layer inPDMS, with a thermoplastic cartridge shell. The method was demonstrated inthe microfluidic integration and packaging of a label-free optical ring resonatorsensor. Biosensing results showed a mass detection limit of 0.9 pg/mm2, oneof the best reported values for integrated ring resontor sensors.

2. The first microfluidic manufacturing method for 3D devices basedon localised inhibition of the polymerisation reaction of PDMS.The novel concept of localised inhibition of the PDMS polymerisation reac-tion enabled high yield fabrication of small, tightly spaced vertical intercon-nects using soft lithography with PDMS. This technology enabled the facileand rapid fabrication of three-dimensional microfluidic channel systems sub-sequently implemented in several labs-on-chip.

The following new material for prototyping lab-on-chips devices was introduced anddemonstrated:

3. The first polymer material platform specifically developed to meetthe complete needs of lab-on-chip prototyping..The novel concept of o!-stoichiometry thiol-enes (OSTE) enabled the devel-opment of a highly versatile prototyping platform, aiming to bridge the gap

63

64 CHAPTER 6. CONCLUSIONS

between academic proof of concept devices and commercial products. Theprototype platform features attractive features for lab-on-chip rapid protoyp-ing:

a) Tuneable mechanical properties (PDMS-like to thermoplastics-like)b) Chemical inertness and low interaction with the sample (minimal absorp-

tion of small molecules)c) Solvent resistance (compatible with common solvents)d) Direct, patternable and stable surface modifications (UV-patternable us-

ing "click" chemistry)e) Low-temperature, biocompatible bonding (bonding at 37 !C without the

need of plasma or glue)f) Rapid, inexpensive and scalable (compatible with soft lithography)

The novel polymer platform was demonstrated in wafer bonding for microflu-idic devices and biocompatible integration with protein and DNA coated mi-croarrays.

The following novel surface design for friction reduction in liquid flows was intro-duced:

4. The first superhydrophobic surface that can withstand liquid pres-sures exceeding the Young-Laplace limitation.The novel designs of superhydrophobic surfaces pneumatically connected thetrapped air pockets either to an external pressure source to demonstrate ac-tive manipulation of the surface friction, or to the bulk liquid pressure througha feedback channel to demonstrate self-regulation of the air pocket pressure.The results showed the novel surfaces resist up to three times higher liquidpressures than previous designs, while maintaining the same friction reduc-ing capacity. The novel designs represented the first step towards practicalimplementations of micro-structured surfaces for friction reduction.

Appendix: Tables

65

66 APPENDIX: TABLES

Table1.

Com

parisonofsom

em

echanicalandchem

icalpropertiesofcomm

onpoly-

mer

materials

usedin

microfluidic

prototyping.

Polymer

AcronymT

gT

mYoung’sm

odulus(M

Pa)

Solventresis-tance

Acid/baseresitance

Water

contactangle

Optical

trans-m

issiv-ity

UV

Previouswork:

Polymethylm

ethacrylate 1PM

MA

100-122250-260

1800-3100G

oodG

ood72

ExcellentG

oodPolycarbonate 1

PC145-148

260-2702000-2400

Good

Good

82Excellent

PoorCyclic

olefin(co)polym

er 1CO

C70-155

190-3202000-2400

ExcellentG

ood82

ExcellentExcellent

Polydimethylsiloxane 2

PDM

S-135

-400-900

Poor(swells)

Excellent90-110

ExcellentExcellent

Norland

Adhesive81(thiol-ene

based) 2N

OA

8172

-1300

ExcellentExcellent

70-80Excellent

Excellent

Currentwork:O

STE(o!-stoichiom

etrythiol-enes) 2

OSTE

30-90-

20-1800Excellent

Excellent70-100

ExcellentExcellent

1http://www.m

atbase.com/

2Author’sD

MA

experiments

67

Tabl

e2.

Com

paris

onof

the

mic

rore

plic

atio

npr

oces

sesf

orm

icro

lfuid

icpr

otot

ypin

g.A

dapt

edfo

rm[1

6]

Proc

ess

Proc

essa

ndeq

uipm

ent

setu

p

Inve

st.

cost

Tool

ing

re-

quire

men

tsCy

cletim

esG

eom

.fle

xibi

l-ity

Prod

uct

auto

ma-

tion

Cast

ing

Sim

ple

(hou

rs)

Non

eVe

rylo

w(<

2k$

)Lo

ng(m

in-

hour

s)

Hig

hLo

w

Hot

em-

boss

ing

Med

ium

(hou

rs-

days

)

Med

ium

(>10

k$)

Low

(2-1

5k$

)M

ediu

m(m

in)

Med

ium

(2D

)M

ediu

m

Inje

ctio

nm

oldi

ngD

i"cu

lt(d

ays)

Hig

h(>

75k$

)H

igh

(20-

150

k$)

Hig

h(s

ec-

min

)

Hig

h(3

D)

Hig

h

Summary of appended papers

Paper 1 : A packaged optical slot-waveguide ring resonator sensor array for mul-tiplex label-free assays in Labs-on-chipWe present the design, fabrication, and characterisation of an array of opti-cal slot-waveguide ring resonator sensors, integrated with microfluidics in acompact cartridge, for multiplexed real time label-fee biosensing. We obtaina volume refractive index detection limit of 5 % 10"6 refractive index units(RIU) and a surface mass density detection limit of 0.9 pg/mm2.A novel bonding method based on dual surface-energy adhesive films allowedfor fast and leak-tight assembly of cartridges with multiple tightly spaced flu-idic interconnects, decreasing the time spent on back-end processes.

Paper 2 : On-chip temperature compensation in an integrated slot-waveguide ringresonator refractive index sensor arrayWe study the temperature dependence of an integrated slot-waveguide re-fractive index sensor array packaged in a microfluidic cartridge. The slot-waveguide ring resonator sensors show a low temperature dependence of '16.6pm/K, while at the same time a large refractive index sensitivity of 240 nmper refractive index unit. Furthermore, by using one channel as on-chip tem-perature reference, a di!erential temperature sensitivity of only 0.3 pm/K isobtained.We demonstrate refractive index measurments during temperature drift andshow a detection limit of 8.8 % 10"6 refractive index units in a 7 K tempera-ture window, without external temperature control.

Paper 3 : Large scale integrated 3D microfluidic networks through high yield fab-rication of vertical vias in PDMSIn this article we introduce, experimentally demonstrate and characterise anovel, uncomplicated single-step method for creating membrane free verticalvias in PDMS. It enables batch manufacturing of large scale integrated 3D mi-crofluidic networks or densely perforated membranes. The method, which has

69

70 SUMMARY OF APPENDED PAPERS

a 100% yield, is based on inhibiting the polymerisation of commercial PDMSonly at the via locations, circumventing the problem of residual membranesblocking the vias due to inadequate or uneven clamping.

Paper 4 : Beyond PDMS: o!-stoichiometry thiol-ene (OSTE) based soft lithogra-phy for rapid prototyping of microfluidic devicesWe introduce a novel polymer platform based on o!-stoichiometry thiol-enes(OSTEs), aiming to bridge the gap between research prototyping and commer-cial production of microfluidic devices. We demonstrate important featuresfor a prototyping system, such as one-step surface modifications; tuneablemechanical properties; direct leakage free sealing through direct UV-bonding;rapid prototyping; uncomplicated processing and the ability to mirror the me-chanical and chemical properties of both PDMS as well as commercial gradethermoplastics.

Paper 5 : Low temperature "click" wafer bonding of o!-stoichiometry thiol-ene(OSTE) polymers to siliconWe introduce a novel wafer bonding concept designed for permanent attach-ment of micromolded polymer structures to functionalized silicon substrates.The method, designed for simultaneous fabrication of many identical Labs-on-chip devices, utilizes a chemically reactive polymer microfluidic structurewhich rapidly bonds to a functionalized substrate wafer via "click" chemistryreactions. The microfluidic structure consists of an o!-stoichiometry thiol-ene (OSTE) polymer with a very high density of surface bound thiol groupsand the substrate is a silicon wafer that has been functionalized with com-mon bio-linker molecules. The method is biocompatible and is well suited forwafer-level microfluidic packaging of pre-functionalized surfaces. In this arti-cle, we demonstrate void free and low temperature (<37 !C) bonding in thefabrication of a complete batch of microfluidic devices consisting of a microflu-idic OSTE polymer layers bonded to a silane functionalized silicon wafer. Thediced devices showed a burst pressure exceeding 4 bars, are compatible withmost organic solvents, are easily surface modified and have excellent solventbarrier properties.

Paper 6 : Sustained superhydrophobic friction reduction at high pressures and largeflowsWe first introduce a figure of merit to describe and compare the total dragreduction of di!erent superhydrophobic flow systems. We then show it im-

71

possible to achieve high friction reduction at high liquid pressure other thanin thin channels only a few micrometers in height due to the Laplace pressurelimitation. Secondly, we introduce a novel self-regulating design for frictionreduction in laminar flow systems with large channels and high pressure. Thedesign is based on pneumatically connecting the air pockets to the liquid pres-sure through feedback channels that can regulate the air pocket pressure. Wedemonstrated that this self-regulating design can give up to three times betterperformance than standard superhydrophobic channels.

Paper 7 : Continuos flow switching by pneumatic actuation of the air lubricationlayer on superhydrophobic microchannel wallsWe introduce a method for robust, active control of friction reduction in mi-crochannels, enabling new flow control applications and overcoming previouslimitations with regard to sustainable liquid pressure. The air pockets trappedat a superhydrophobic micrograting during liquid priming are coupled to anactively controlled pressure source, allowing the pressure di!erence over theair/liquid interface to be dynamically adjusted. This allows for manipulat-ing the friction reduction properties of the surface, enabling active controlof liquid mass flow through the channel. It also permits for sustainable airlubrication at high liquid pressures, without loss of superhydrophobic proper-ties. With the non-optimized grating used in the experiment, a di!erence inliquid mass flow of 5% is obtained by alternatively collapsing and recreatingthe air pockets using the coupled pressure source. The method also allows forsustainable liquid pressure 3 times higher than the Young-Laplace pressure ofa passive device.

Paper 8 : Continuos flow switching by pneumatic actuation of the air lubricationlayer on superhydrophobic microchannel wallsWe present a one-step, reversible, and biocompatible bonding method ofa sti! patterned microfluidic "Biosticker", based on o!-stoichiometry thiol-ene (OSTE) polymers , to state-of-the-art spotted microarray surfaces. Themethod improves and simplifies the batch back-end processing of microarrays.We illustrate its ease of use in two applications: a high sensitivity flow-throughprotein assay; and a DNA-hybridization test. Read-out was performed in astandard high-volume array scanner, and showed excellent spot homogeneityand intensity. The Biosticker is aimed to be a plug-in for existing microarrayplatforms to enable faster protein assays and DNA hybridizations throughmass transport optimization.

Acknowledgement

My doctoral studies have taken me through many di!erent areas of research, andthe present work would have been impossible without the guidance and assistancefrom many people.

First of all I would like to thank my main supervisor Wouter van der Wijngaart,for accepting me as a student in his team, for giving me the time to learn and givingme the freedom to make mistakes and the guidance not to repeat them. I wouldalso like to thank Tommy Haraldsson for introducing me to the world of polymerscience. His guidance and support have made the last part of my time as a PhD-student a truly ex(c)iting journey. Many thanks also to Göran Stemme for havingcreated a truly motivating and unpretentious research environment in his group.

Financial support for this work was provided partly through grants from theSwedish Research Council and the European Commission via the FP6-IST-SABIOproject.

During my time as PhD student at the Microsystem Technology Laboratory, Ihave been lucky to work together with many extraordinary individuals. First I wouldlike to thank people who have directly helped out with the experiments: KristinnGylfason, for our collaboration in the bio-sensing project, for being a great sourceof knowledge and inspiration, and for becoming a good friend; Mikael Karlsson,for sharing long days of experimenting in the laboratory at MST and for alwaysbeing so positive; Farizah Saharil, for good team work with the OSTE experiments;Matteo Cornaglia for all the PDMS you inhibited for me; Thomas Moh, for helpingme with the microfluidic assmeblies; Liu Yitong, for the excellent work on improvingthe OSTE-polymers; Kim Öberg for help with the thiol-enes and for always havingtime; Andrzej Ka$mierzak and Fabian Dortu at Multitel, for long days and nightsof experiments in Mons; Marcella Chiari, Marina Cretich and Laura Sola at ICRMfor the microarray experiments; Minh Do-Quang at KTH Mechanics for help withfinite element simulations; Junichiro Shiomi at the Tokyo University for introducingme to molecular dynamics, and academic research in general.

Special thanks also to those who have contributed with valuable discussions andhelp: Mikael Sterner, my o"ce-mate, for patiently sharing valuable tips and trickson many diverse topics; Niklas Sandström, for discussions on exciting applicationsof our research, Aman Russom, Sergey Zelenin and Sahar Ardabili at KTH CellPhysics for all the discussions on cells, DNA and proteins and for the help withthe microscopes, Gerry Ronan at Farfield and Gerhard Kreshbach at Zeptosens for

73

74 ACKNOWLEDGEMENT

sharing their expertise in system integration and biosensing; Michael Malkoch andMauro Claudino at KTH Fiber and Polymer Technology for valuable discussions onpolymer science and for help with the machines.

All my friends, colleagues and former colleagues at MST: Adit, Andreas, Björn,Erika, Farizah, Frank, Gaspard, Göran, Hans, Henrik, Hithesh, Joachim, FredrikF, Fritzi, Kjell, Martin, Niclas, Niklas, Nutapong, Mikael A, Mikael K, Mikael S,Sta!an, Stefan, Thomas, Tommy, Umer, Wouter, Liu Yitong and Zargam. Thankyou for a supporting atmosphere and the good times.

I also very much appreciate the help from Tommy, Kristinn, Ebba, Hans, Wouterand my mom for proofreading di!erent parts of this thesis.

I am forever in debt to my parents for supporting me all the way, in particularto my dad who had very much looked forward to this day but was not given thetime to experience it.

Last but not the least, I would like to thank my beloved wife Selina, for lovingme and being all that I could ever wish for. This would never have been possiblewithout your understanding and encouragements.

Carl Fredrik Carlborg, Stockholm, August 29th, 2011

References

[1] L. Gervais, N. de Rooij, and E. Delamarche, “Microfluidic Chips for Point-of-Care Immunodiagnostics,” Adv. Mater., vol. 23, no. 24, pp. H151–H176,2011.

[2] F. Breussin, “Microfluidic technologies for point of care testing,” in Yole Work-shop lab-on-chip for diagnostics, Yolé Développement, Mar. 2010.

[3] J. Gantelius, Novel diagnostic microarray assay formats towards comprehen-sive on-site analysis Novel diagnostic microarray assay formats towards com-prehensive on-site analysis. PhD thesis, Royal Institute of Technology (KTH),2009.

[4] G. Odian, Principles of Polymerization. Hoboken, New Jersey: John Wiley& Sons, Inc., 2004.

[5] J. A. Carioscia, H. Lu, J. W. Stanbury, and C. N. Bowman, “Thiol-eneoligomers as dental restorative materials,” Dental Materials, vol. 21, pp. 1137–1143, Dec. 2005.

[6] D. Bartolo, G. Degre, P. Nghe, and V. Studer, “Microfluidic stickers,” LabChip, vol. 8, no. 2, pp. 274–279, 2008.

[7] S. Lee and S. Lee, “Shrinkage ratio of PDMS and its alignment method forthe wafer level process,” Microsystem Technologies, vol. 14, pp. 205–208, Feb.2008.

[8] M. W. Toepke and D. J. Beebe, “PDMS absorption of small molecules andconsequences in microfluidic applications,” Lab Chip, vol. 6, no. 12, pp. 1484–1486, 2006.

[9] S. Yunus, de Looringhe, C. Poleunis, and A. Delcorte, “Di!usion of oligomersfrom polydimethylsiloxane stamps in microcontact printing: Surface analysisand possible application,” Surf. Interface Anal., vol. 39, no. 12-13, pp. 922–925, 2007.

[10] J. N. Lee, C. Park, and G. M. Whitesides, “Solvent Compatibility ofPoly(dimethylsiloxane)-Based Microfluidic Devices,” Analytical Chemistry,vol. 75, pp. 6544–6554, Dec. 2003.

75

76 REFERENCES

[11] J. W. Swanson and J. E. Lebeau, “The e!ect of implantation on the physicalproperties of silicone rubber,” J. Biomed. Mater. Res., vol. 8, no. 6, pp. 357–367, 1974.

[12] K. M. Choi and J. A. Rogers, “A Photocurable Poly(dimethylsiloxane) Chem-istry Designed for Soft Lithographic Molding and Printing in the NanometerRegime,” Journal of the American Chemical Society, vol. 125, pp. 4060–4061,Apr. 2003.

[13] B.-Y. Kim, L.-Y. Hong, Y.-M. Chung, D.-P. Kim, and C.-S. Lee, “Solvent-Resistant PDMS Microfluidic Devices with Hybrid Inorganic/Organic Poly-mer Coatings,” Advanced Functional Materials, vol. 19, no. 23, pp. 3796–3803,2009.

[14] A. R. Abate, D. Lee, T. Do, C. Holtze, and D. A. Weitz, “Glass coating forPDMS microfluidic channels by sol-gel methods,” Lab Chip, vol. 8, pp. 516–518, Apr. 2008.

[15] J. Lee, M. J. Kim, and H. H. Lee, “Surface Modification ofPoly(dimethylsiloxane) for Retarding Swelling in Organic Solvents,” Lang-muir, vol. 22, pp. 2090–2095, Feb. 2006.

[16] H. Becker and C. Gärtner, “Polymer microfabrication technologies for mi-crofluidic systems,” Analytical and Bioanalytical Chemistry, vol. 390, pp. 89–111, Jan. 2008.

[17] L. J. Lee, M. J. Madou, K. W. Koelling, S. Daunert, S. Lai, C. G. Koh,Y.-J. Juang, Y. Lu, and L. Yu, “Design and fabrication of cd-like microflu-idic platforms for diagnostics: Polymer-based microfabrication,” BiomedicalMicrodevices, vol. 3, no. 4, pp. 339–354, 2001.

[18] N. Le, R. Gubala, V andGandhiraman, S. Daniels, and D. Williams,“Evaluation of di!erent nonspecific binding blocking agents deposited in-side poly(methyl methacrylate) microfluidic flow-cells,” Langmuir, vol. 27,pp. 9043–51, June 2011.

[19] R. Ledesma-Aguilar, R. Nistal, A. Hernandez-Machado, and A. Pagonabar-raga, “Controlled drop emission by wetting properties in driven liquid fila-ments,” Nature Materials, vol. 10, pp. 367–371, 2011.

[20] T. Rohr, D. Ogletree, F. Svec, and J. Fréchet, “Surface functionalization ofthermoplastic polymers for the fabrication of microfluidic devices by photoini-tiated grafting,” Advanced Functional Materials, vol. 13, pp. 264–270, April2003.

REFERENCES 77

[21] D. Witters, N. Vergauwe, S. Vermeir, F. Ceyssens, S. Liekens, R. Puers, andJ. Lammertyn, “Biofunctionalization of electrowetting-on-dielectric digital mi-crofluidic chips for miniaturized cell-based applications,” Lab Chip, vol. 11,pp. 2790–2794, 2011.

[22] D. Wu, B. Zhao, Z. Dai, J. Qin, and B. Lin, “Grafting epoxy-modified hy-drophilic polymers onto poly(dimethylsiloxane) microfluidic chip to resist non-specific protein adsorption,” Lab Chip, vol. 6, pp. 942–947, 2006.

[23] D. W. Grainger, C. H. Greef, P. Gong, and M. J. Lochhead, “Current microar-ray surface chemistries,” Methods in Molecular Biology, vol. 381, pp. 35–57,2007.

[24] C. F. Carlborg and W. van der Wijngaart, “Sustained Superhydrophobic Fric-tion Reduction at High Liquid Pressures and Large Flows,” Langmuir, vol. 27,pp. 487–493, Jan. 2011.

[25] E. Eteshola and D. Leckband, “Development and characterization of an elisaassay in pdms microfluidic channels,” Sensors and Actuators B: Chemical,vol. 72, no. 2, pp. 129–133, 2001.

[26] S. A. Ruiz and C. S. Chen, “Microcontact printing: A tool to pattern,” SoftMatter, vol. 3, pp. 168–177, 2007.

[27] M. Geissler, A. Bernard, A. Bietsch, H. Schmid, B. Michel, and E. Dela-marche, “Microcontact-printing chemical patterns with flat stamps,” Journalof American Chemical Society, vol. 112, no. 26, pp. 6303–6304, 2000.

[28] L. P. Hromada, B. J. Nablo, J. J. Kasianowicz, M. A. Gaitan, and D. L.DeVoe, “Single molecule measurements within individual membrane-boundion channels using a polymer-based bilayer lipidmembrane chip,” Lab Chip,vol. 8, pp. 602–608, 2008.

[29] Z. Chen, L. Zhang, and G. Chen, “A spring-driven press device for hot em-bossing and thermal bonding of pmma microfluidic chips,” Electrophoresis,vol. 31, pp. 2512–2519, 2010.

[30] B.-Y. Pemga, C.-W. Wub, Y.-K. Shenc, and Y. Lind, “Microfluidic chip fab-rication using hot embossing and thermal bonding of cop,” Polymers for Ad-vanced Technologies, vol. 21, no. 7, pp. 457–466, 2009.

[31] R. KW, L. H, and K. DR, “Plastic microchip liquid chromatography-matrix-assisted laser desorption/ionization mass spectrometry using mono-lithic columns,” Journal of Chromatography A, vol. 1111, no. 1, pp. 40–47,2006.

78 REFERENCES

[32] H. Klank, J. P. Kutter, and O. Geschke, “Co2-laser micromachining and back-end processing for rapid production of pmma-based microfluidic systems,” LabChip, vol. 2, pp. 242–246, 2002.

[33] L. Brown, T. Koerner, J. H. Horton, and R. D. Oleschuk, “Fabrication andcharacterization of poly(methylmethacrylate) microfluidic devices bonded us-ing surface modifications and solvents,” Lab Chip, vol. 6, no. 66-73, 2006.

[34] Y. Hsu and T. Chen, “Applying taguchi methods for solvent-assisted pmmabonding technique for static and dynamic micro-tas devices,” Biomedical Mi-crodevices, vol. 9, no. 4, pp. 513–22, 2007.

[35] R. Truckenmuller, R. Ahrens, Y. Cheng, G. Fischer, and V. Saile, “An ultra-sonic welding based process for building up a new class of inert fluidic mi-crosensors and -actuators from polymers,” Sensors and Actuators A, vol. 132,no. 1, pp. 385–392, 2006.

[36] A. Boglea, A. Olowinsky, and A. Gillner, “Fibre laser welding for packaging ofdisposable polymeric microfluidic-biochips,” Applied Surface Science, vol. 254,pp. 1174–1178, Dec. 2007.

[37] T. Ussing, L. V. Petersen, C. B. Nielsen, B. Helbo, and L. Højslet, “Microlaser welding of polymer microstructures using low power laser diodes,” TheInternational Journal of Advanced Manufacturing Technology, vol. 22, no. 1-2,pp. 198–205, 2007.

[38] J. C. McDonald, D. C. Du!y, J. R. Anderson, D. T. Chiu, H. Wu, O. J. A.Schueller, and G. M. Whitesides, “Fabrication of microfluidic systems inpoly(dimethylsiloxane),” ELECTROPHORESIS, vol. 21, pp. 27–40, Jan.2000.

[39] J. Kentsch, S. Breisch, and M. Stelzle, “Low temperature adhesion bonding forbiomems,” Journal of Micromechanics and Microengineering, vol. 16, no. 4,pp. 802–807, 2006.

[40] F. Dang, S. Shinohara, O. Tabata, Y. Yamaoka, M. Kurokawa, Y. Shinohara,M. Ishikawa, and Y. Baba, “Replica multichannel polymer chips with a net-work of sacrificial channels sealed by adhesive printing method,” Lab Chip,vol. 5, pp. 472–478, 2005.

[41] D. F. Pulak Nath, Y. A. Kunde, A. Zeytun, B. Branch, and G. Goddard,“Rapid prototyping of robust and versatile microfluidic components using ad-hesive transfer tapes,” Lab Chip, vol. 10, pp. 2286–2291, 2010.

[42] M. A. Unger, H.-P. Chou, T. Thorsen, A. Scherer, and S. R. Quake, “Mono-lithic Microfabricated Valves and Pumps by Multilayer Soft Lithography,”Science, vol. 288, pp. 113–116, Apr. 2000.

REFERENCES 79

[43] H.-Y. Chen, A. A. McClelland, Z. Chen, and J. Lahann, “Solventless adhe-sive bonding using reactive polymer coatings,” Analytical Chemistry, vol. 80,no. 11, pp. 4119–4124, 2008.

[44] K. S. Lee and R. J. Ram, “Plastic-PDMS bonding for high pressure hydrolyt-ically stable active microfluidics,” Lab Chip, vol. 9, no. 11, pp. 1618–1624,2009.

[45] A. Gerlach, H. Lambach, and D. Seidel, “Propagation of adhesives in jointsduring capillary adhesive bonding of microcomponents,” Microsystem Tech-nologies, vol. 6, pp. 19–22, Nov. 1999.

[46] C.-W. Tsaoa and D. L. DeVoe, Lab on a Chip Technology: Volume 1: Fabri-cation and Microfluidics, vol. 1. Caister Academic Press, 2009.

[47] D. Mark, S. Haeberle, G. Roth, F. Stetten, and R. Zengerle, “Microfluidic lab-on-a-chip platforms: requirements, characteristics and applications,” Chem.Soc. Rev., vol. 39, no. 3, pp. 1153–1182, 2010.

[48] V. Sunkara, D.-K. Park, H. Hwang, R. Chantiwas, S. A. Soper, andY.-K. Cho, “Simple room temperature bonding of thermoplastics andpoly(dimethylsiloxane),” Lab Chip, vol. 11, no. 5, pp. 962–965, 2011.

[49] J. Kim, R. Surapaneni, and B. K. Gale, “Rapid prototyping of microfluidicsystems using a PDMS/polymer tape composite,” Lab Chip, vol. 9, no. 9,pp. 1290–1293, 2009.

[50] H. Sohlström and M. Öberg, “Refractive index measurement using inte-grated ring resonators,” in The 8th European Conference on Integrated Optics,pp. 322–325, Apr. 1997.

[51] E. Krioukov, Klunder, A. Driessen, J. Greve, and C. Otto, “Sensor based onan integrated optical microcavity,” Opt. Lett., vol. 27, pp. 512–514, Apr. 2002.

[52] C. Y. Chao and L. J. Guo, “Biochemical sensors based on polymer microringswith sharp asymmetrical resonance,” Applied Physics Letters, vol. 83, no. 8,pp. 1527–1529, 2003.

[53] A. L. Washburn, L. C. Gunn, and R. C. Bailey, “Label-Free Quantitationof a Cancer Biomarker in Complex Media Using Silicon Photonic MicroringResonators,” Analytical Chemistry, vol. 81, pp. 9499–9506, Nov. 2009.

[54] V. R. Almeida, Q. Xu, C. A. Barrios, and M. Lipson, “Guiding and confininglight in void nanostructure,” Optics Letters, vol. 29, no. 11, pp. 1209–1211,2004.

[55] Q. Xu, V. R. Almeida, R. R. Panepucci, and M. Lipson, “Experimentaldemonstration of guiding and confining light in nanometer-size low-refractive-index material,” Optics Letters, vol. 29, no. 14, pp. 1626–1628, 2004.

80 REFERENCES

[56] C. A. Barrios, M. J. Bañuls, V. González-Pedro, K. B. Gylfason, B. Sánchez,A. Griol, A. Maquieira, H. Sohlström, M. Holgado, and R. Casquel, “Label-free optical biosensing with slot-waveguides,” Opt. Lett., vol. 33, pp. 708–710,Apr. 2008.

[57] T. Gervais and K. F. Jensen, “Mass transport and surface reactions in mi-crofluidic systems,” Chemical Engineering Science, vol. 61, pp. 1102–1121,Feb. 2006.

[58] T. M. Squires, R. J. Messinger, and S. R. Manalis, “Making it stick: convec-tion, reaction and di!usion in surface-based biosensors,” Nature Biotechnol-ogy, vol. 26, pp. 417–426, Apr. 2008.

[59] K. Pappaert, J. Vanderhoeven, P. Vanhummelen, B. Dutta, D. Clicq,G. Baron, and G. Desmet, “Enhancement of DNA micro-array analysis us-ing a shear-driven micro-channel flow system,” Journal of Chromatography A,vol. 1014, pp. 1–9, Oct. 2003.

[60] B. I. Inc, Application notes #107: Binding Kinetics Analysis with SPR: In-teraction between Bovine Serum Albumin (BSA) and Anti-BSA, Mar. 2010.

[61] A. K. Wright and M. R. Thompson, “Hydrodynamic structure of bovine serumalbumin determined by transient electric birefringence.,” Biophysical journal,vol. 15, pp. 137–141, Feb. 1975.

[62] L. Chen, F. Azizi, and C. H. Mastrangelo, “Generation of dynamic chemicalsignals with microfluidic C-DACs,” Lab Chip, vol. 7, no. 7, pp. 850–855, 2007.

[63] M. Rhee and M. A. Burns, “Microfluidic assembly blocks,” Lab Chip, vol. 8,no. 8, pp. 1365–1373, 2008.

[64] C. Neils, Z. Tyree, B. Finlayson, and A. Folch, “Combinatorial mixing ofmicrofluidic streams,” Lab Chip, vol. 4, no. 4, pp. 342–350, 2004.

[65] K. Atsuta, H. Noji, and S. Takeuchi, “Micro patterning of active proteins withperforated PDMS sheets (PDMS sieve),” Lab Chip, vol. 4, no. 4, pp. 333–336,2004.

[66] J. Cha, J. Kim, S.-K. Ryu, J. Park, Y. Jeong, S. Park, S. Park, H. C. Kim, andK. Chun, “A highly e"cient 3D micromixer using soft PDMS bonding,” Jour-nal of Micromechanics and Microengineering, vol. 16, pp. 1778–1782, Sept.2006.

[67] J. R. Anderson, D. T. Chiu, R. J. Jackman, O. Cherniavskaya, J. C. Mc-Donald, H. Wu, S. H. Whitesides, and G. M. Whitesides, “Fabrication ofTopologically Complex Three-Dimensional Microfluidic Systems in PDMS byRapid Prototyping,” Analytical Chemistry, vol. 72, pp. 3158–3164, July 2000.

REFERENCES 81

[68] B. H. Jo, L. M. Van Lerberghe, K. M. Motsegood, and D. J. Beebe, “Three-dimensional micro-channel fabrication in polydimethylsiloxane(PDMS) elas-tomer,” Microelectromechanical Systems, Journal of, vol. 9, pp. 76–81, Mar.2000.

[69] C.-H. Hsu, C. Chen, and A. Folch, “"Microcanals" for micropipette access tosingle cells in microfluidic environments,” Lab Chip, vol. 4, no. 5, pp. 420–424,2004.

[70] E. P. Kartalov, C. Walker, C. R. Taylor, W. F. Anderson, and A. Scherer,“Microfluidic vias enable nested bioarrays and autoregulatory devices in New-tonian fluids,” Proceedings of the National Academy of Sciences, vol. 103,pp. 12280–12284, Aug. 2006.

[71] R. J. Jackman, D. C. Du!y, O. Cherniavskaya, and G. M. Whitesides, “Us-ing Elastomeric Membranes as Dry Resists and for Dry Lift-O!,” Langmuir,vol. 15, pp. 2973–2984, Apr. 1999.

[72] J. H. Kang, E. Um, and J.-K. Park, “Fabrication of a poly(dimethylsiloxane)membrane with well-defined through-holes for three-dimensional microfluidicnetworks,” Journal of Micromechanics and Microengineering, vol. 19, no. 4,pp. 045027+, 2009.

[73] N. Sabourault, G. Mignani, A. Wagner, and C. Mioskowski, “Platinum Ox-ide (PtO2): A Potent Hydrosilylation Catalyst,” Organic Letters, vol. 4,pp. 2117–2119, June 2002.

[74] M. Karlsson, T. Haraldsson, N. Sandström, G. Stemme, A. Russom, andW. van der Wijngaart, “On-chip liquid degassing with low water loss,” inProceedings Micro Total Analysis Systems (muTAS), pp. 1790–1792, 2010.

[75] J. Hansson, J. M. Karlsson, T. Haraldsson, W. van der Wijngaart, and A. Rus-som, “Inertial particle focusing on parallel microfluidic channels for high-throughput filtration,” in The 16th International Conference on Solid-StateSensors, Actuators and Microsystems, pp. 1777–1780, IEEE, 2011.

[76] C. E. Hoyle and C. N. Bowman, “Thiol-Ene Click Chemistry,” AngewandteChemie International Edition, vol. 49, no. 9, pp. 1540–1573, 2010.

[77] H. C. Kolb, M. G. Finn, and K. B. Sharpless, “Click Chemistry: DiverseChemical Function from a Few Good Reactions,” Angewandte Chemie Inter-national Edition, vol. 40, no. 11, pp. 2004–2021, 2001.

[78] N. B. Cramer, J. P. Scott, and C. N. Bowman, “Photopolymerizations of thiol-ene Polymers without Photoinitiators,” Macromolecules, vol. 35, pp. 5361–5365, July 2002.

82 REFERENCES

[79] N. Sandström, R. Z. Shafagh, C. F. Carlbor, T. Haraldsson, G. Stemme,and W. van der Wijngaart, “One step integration of gold coated sensors withOSTE polymer cartridges by low temperature dry bonding,” in The 16thInternational Conference on Solid State Sensors, Actuators and Microsystems,pp. 2778–2781, IEEE, 2011.

[80] P. Man, D. Jones, and C. Mastrangelo, “Microfluidic plastic capillaries onsilicon substrates: a new inexpensive technology for bioanalysis chips,” inMicro Electro Mechanical Systems, 1997. MEMS ’97, Proceedings, IEEE.,Tenth Annual International Workshop on, pp. 311 – 316, 1997.

[81] C. A. Afshari, “Perspective: Microarray Technology, Seeing More ThanSpots,” Endocrinology, vol. 143, pp. 1983–1989, June 2002.

[82] G. MacBeath, “Protein microarrays and proteomics,” Nature Genetics,vol. 32, pp. 526–532, Dec. 2002.

[83] S. F. Kingsmore, “Multiplexed protein measurement: technologies and ap-plications of protein and antibody arrays,” Nature Reviews Drug Discovery,vol. 5, pp. 310–321, Mar. 2006.

[84] M. Hartmann, J. Roeraade, D. Stoll, M. Templin, and T. Joos, “Proteinmicroarrays for diagnostic assays,” Analytical and Bioanalytical Chemistry,vol. 393, pp. 1407–1416, Mar. 2009.

[85] M. Schena, Protein Microarrays. Jones and Bartlett Publishers, Inc, July2004.

[86] G. L. Lukacs, P. Haggie, O. Seksek, D. Lechardeur, N. Freedman, and A. S.Verkman, “Size-dependent DNA Mobility in Cytoplasm and Nucleus,” Jour-nal of Biological Chemistry, vol. 275, pp. 1625–1629, Jan. 2000.

[87] S. Russell, L. Meadows, and R. Russel, Microarray Technology in Practice.Elsevier, 2008.

[88] C. Situma, M. Hashimoto, and S. Soper, “Merging microfluidics withmicroarray-based bioassays,” Biomolecular Engineering, vol. 23, pp. 213–231,Oct. 2006.

[89] H. H. Lee, J. Smoot, Z. McMurray, D. A. Stahl, and P. Yager, “Recirculatingflow accelerates DNA microarray hybridization in a microfluidic device,” LabChip, vol. 6, no. 9, pp. 1163–1170, 2006.

[90] D. Erickson, X. Liu, U. Krull, and D. Li, “Electrokinetically Controlled DNAHybridization Microfluidic Chip Enabling Rapid Target Analysis,” AnalyticalChemistry, vol. 76, pp. 7269–7277, Dec. 2004.

REFERENCES 83

[91] J. Liu, B. A. Williams, R. M. Gwirtz, B. J. Wold, and S. Quake, “EnhancedSignals and Fast Nucleic Acid Hybridization By Microfluidic Chaotic Mixing,”Angewandte Chemie International Edition, vol. 45, no. 22, pp. 3618–3623,2006.

[92] N. B. Adey, M. Lei, M. T. Howard, J. D. Jensen, D. A. Mayo, D. L. Butel,S. C. Co"n, T. C. Moyer, D. E. Slade, M. K. Spute, A. M. Hancock, G. T.Eisenho!er, B. K. Dalley, and M. R. McNeely, “Gains in Sensitivity witha Device that Mixes Microarray Hybridization Solution in a 25-Î!m-ThickChamber,” Analytical Chemistry, vol. 74, pp. 6413–6417, Dec. 2002.

[93] Y. Liu, “DNA probe attachment on plastic surfaces and microfluidic hybridiza-tion array channel devices with sample oscillation,” Analytical Biochemistry,vol. 317, pp. 76–84, June 2003.

[94] M. Noerholm, H. Bruus, M. H. Jakobsen, P. Telleman, and N. B. Rams-ing, “Polymer microfluidic chip for online monitoring of microarray hybridiza-tions,” Lab Chip, vol. 4, no. 1, pp. 28–37, 2004.

[95] C. Situma, Y. Wang, M. Hupert, F. Barany, R. McCarley, and S. Soper,“Fabrication of DNA microarrays onto poly(methyl methacrylate) with ultra-violet patterning and microfluidics for the detection of low-abundant pointmutations,” Analytical Biochemistry, vol. 340, pp. 123–135, May 2005.

[96] M. Cretich, G. di Carlo, R. Longhi, C. Gotti, N. Spinella, S. Co!a, C. Galati,L. Renna, and M. Chiari, “High sensitivity protein assays on microarray siliconslides,” Analytical Chemistry, vol. 81, no. 13, pp. 5197–5203, 2009.

[97] J. A. Carioscia, J. W. Stansbury, and C. N. Bowman, “Evaluation and controlof thiol–ene/thiol–epoxy hybrid networks,” Polymer, vol. 48, pp. 1526–1532,Mar. 2007.

[98] J. F. Ashley, N. B. Cramer, R. H. Davis, and C. N. Bowman, “Soft-lithographyfabrication of microfluidic features using thiol-ene formulations,” Lab Chip,vol. 11, no. 16, pp. 2772–2778, 2011.

[99] K. Fukuda, J. Tokunaga, T. Nobunaga, T. Nakatani, T. Iwasaki, and Y. Kuni-take, “Frictional drag reduction with air lubricant over a super-water-repellentsurface,” Journal of Marine Science and Technology, vol. 5, pp. 123–130, Dec.2000.

[100] C. Lee and C.-J. Kim, “Maximizing the Giant Liquid Slip on Superhydropho-bic Microstructures by Nanostructuring Their Sidewalls,” Langmuir, vol. 25,pp. 12812–12818, Nov. 2009.

[101] C. C. Bizonne, B. Cross, A. Steinberger, and E. Charlaix, “Boundary Slipon Smooth Hydrophobic Surfaces: Intrinsic E!ects and Possible Artifacts,”Physical Review Letters, vol. 94, pp. 056102+, Feb. 2005.

84 REFERENCES

[102] C. H. Choi, Johan, and K. S. Breuer, “Apparent slip flows in hydrophilic andhydrophobic microchannels,” Physics of Fluids, vol. 15, no. 10, pp. 2897–2902,2003.

[103] T. C. Island, W. D. Urban, and M. G. Mungal, “Mixing enhancement incompressible shear layers via sub-boundary layer disturbances,” Physics ofFluids, vol. 10, no. 4, pp. 1008–1020, 1998.

[104] E. Lauga and H. A. Stone, “E!ective slip in pressure-driven Stokes flow,”Journal of Fluid Mechanics, vol. 489, no. -1, pp. 55–77, 2003.

[105] Y. C. Jung and B. Bhushan, “Biomimetic structures for fluid drag reduction inlaminar and turbulent flows,” Journal of Physics: Condensed Matter, vol. 22,pp. 035104+, Jan. 2010.

[106] C. Lee and C. J. Kim, “Underwater Restoration and Retention of Gaseson Superhydrophobic Surfaces for Drag Reduction,” Physical Review Letters,vol. 106, pp. 014502+, Jan. 2011.

[107] C. Lee and C.-J. Kim, “Influence of Surface Hierarchy of SuperhydrophobicSurfaces on Liquid Slip,” Langmuir, vol. 27, pp. 4243–4248, Apr. 2011.

[108] R. J. Daniello, N. E. Waterhouse, and J. P. Rothstein, “Drag reduction inturbulent flows over superhydrophobic surfaces,” Physics of Fluids, vol. 21,no. 8, pp. 085103+, 2009.

[109] B. Woolford, D. Maynes, and B. Webb, “Liquid flow through microchan-nels with grooved walls under wetting and superhydrophobic conditions,” Mi-crofluidics and Nanofluidics, vol. 7, pp. 121–135, July 2009.

[110] P. Joseph, C. C. Bizonne, J. M. Benoit, C. Ybert, C. Journet, P. Tabeling,and L. Bocquet, “Slippage of Water Past Superhydrophobic Carbon NanotubeForests in Microchannels,” Physical Review Letters, vol. 97, pp. 156104+, Oct.2006.

[111] T. Onda, S. Shibuichi, N. Satoh, and K. Tsujii, “Super-Water-Repellent Frac-tal Surfaces,” Langmuir, vol. 12, pp. 2125–2127, Jan. 1996.

[112] T. Sun, G. Wang, L. Feng, B. Liu, Y. Ma, L. Jiang, and D. Zhu, “Re-versible Switching between Superhydrophilicity and Superhydrophobicity,”Angewandte Chemie International Edition, vol. 43, no. 3, pp. 357–360, 2004.

[113] F. Xia, L. Feng, S. Wang, T. Sun, W. Song, W. Jiang, and L. Jiang, “Dual-Responsive Surfaces That Switch between Superhydrophilicity and Superhy-drophobicity,” Adv. Mater., vol. 18, no. 4, pp. 432–436, 2006.

[114] J. Lee, B. He, and N. A. Patankar, “A roughness-based wettability switchingmembrane device for hydrophobic surfaces,” Journal of Micromechanics andMicroengineering, vol. 15, pp. 591+, Mar. 2005.

REFERENCES 85

[115] A. Athanassiou, M. I. Lygeraki, D. Pisignano, K. Lakiotaki, M. Varda,E. Mele, C. Fotakis, R. Cingolani, and S. H. Anastasiadis, “Photocontrolledvariations in the wetting capability of photochromic polymers enhanced bysurface nanostructuring.,” Langmuir : the ACS journal of surfaces and col-loids, vol. 22, pp. 2329–2333, Feb. 2006.

[116] H. S. Lim, J. T. Han, D. Kwak, M. Jin, and K. Cho, “Photoreversibly Switch-able Superhydrophobic Surface with Erasable and Rewritable Pattern,” Jour-nal of the American Chemical Society, vol. 128, pp. 14458–14459, Nov. 2006.

[117] T. N. Krupenkin, J. A. Taylor, E. N. Wang, P. Kolodner, M. Hodes, and T. R.Salamon, “Reversible Wetting Dewetting Transitions on Electrically TunableSuperhydrophobic Nanostructured Surfaces,” Langmuir, vol. 23, pp. 9128–9133, Aug. 2007.

[118] J. B. Boreyko and C. H. Chen, “Restoring Superhydrophobicity of LotusLeaves with Vibration-Induced Dewetting,” Physical Review Letters, vol. 103,pp. 174502+, Oct. 2009.

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