Development and Application of Microfluidic Valves

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Title:
Development and Application of Microfluidic Valves
Physical Description:
1 online resource (122 p.)
Language:
english
Creator:
Gu, Pan
Publisher:
University of Florida
Place of Publication:
Gainesville, Fla.
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Thesis/Dissertation Information

Degree:
Doctorate ( Ph.D.)
Degree Grantor:
University of Florida
Degree Disciplines:
Mechanical Engineering, Mechanical and Aerospace Engineering
Committee Chair:
Fan, Zhonghui H
Committee Co-Chair:
Nishida, Toshikazu
Committee Members:
Segal, Corin
Roy, Subrata
Denslow, Nancy D

Subjects

Subjects / Keywords:
biomarkers -- immunoassay -- microfluidics -- microvalves -- pdms -- polyurethane -- tbi -- thermoplastic
Mechanical and Aerospace Engineering -- Dissertations, Academic -- UF
Genre:
Mechanical Engineering thesis, Ph.D.
bibliography   ( marcgt )
theses   ( marcgt )
government publication (state, provincial, terriorial, dependent)   ( marcgt )
born-digital   ( sobekcm )
Electronic Thesis or Dissertation

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Abstract:
Microfluidics has attracted the attention from a large number of researchers and is a promising technology in biotechnology, biomedical engineering to biochemistry. Comparing with microfluidic devices made from silicon and glass, a mass-produced, plastic microfluidic device is of low cost and can be disposable. One of the critical components in a LOC (lab on a chip) system is microvalve, which tend to be problematic when integrated with a device. Since commercialization calls for the reliability, it is important to develop a plastic microfluidic system with reliable microvalves. This thesis presents and demonstrates the integration of two types of microvalves in plastic devices, as well as their downstream applications. The primary problem comes from the bonding reliability caused by the low plastic surface energy. A strong bond between the elastomer layer and channel layer is a precondition to create an elastomer based microvalves. As a solution, we developed a strong, reliable and universal bonding method to bond the thermal plastic such as polymethyl methacrylate (PMMA) and cyclic olefin copolymer (COC), with elastomer material such as polydimethylsiloxane (PDMS) and polyurethane (PU). The bond between can sustain a pressure more than 100 psi without any delamination. PDMS and PU were used as the elastomer material in two types of valves: pneumatic valve and thermally actuated valve. A concept of using the valves integrated platform for simultaneous immunoassay detection of multiple analytes is presented. A platform including 13 thermally actuated valves was proposed to detection up to six analytes. A printed circuit board (PCB) board was created as a core component of power control system to handling the flow direction by valve operating combination. A surface modification method was developed to make COC adaptable to the bio-molecule immobilization. This concept was demonstrated by the detection of two types of biomarks, C-creative protein (CRP) and transferrin, at a concentration of 100 micrograms/mL.
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In the series University of Florida Digital Collections.
General Note:
Includes vita.
Bibliography:
Includes bibliographical references.
Source of Description:
Description based on online resource; title from PDF title page.
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This bibliographic record is available under the Creative Commons CC0 public domain dedication. The University of Florida Libraries, as creator of this bibliographic record, has waived all rights to it worldwide under copyright law, including all related and neighboring rights, to the extent allowed by law.
Statement of Responsibility:
by Pan Gu.
Thesis:
Thesis (Ph.D.)--University of Florida, 2012.
Local:
Adviser: Fan, Zhonghui H.
Local:
Co-adviser: Nishida, Toshikazu.
Electronic Access:
RESTRICTED TO UF STUDENTS, STAFF, FACULTY, AND ON-CAMPUS USE UNTIL 2014-05-31

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Applicable rights reserved.
Classification:
lcc - LD1780 2012
System ID:
UFE0044103:00001


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1 DEVELOPMENT AND APPLICATION OF MICROFLUIDIC VALVES By PAN GU A DISSERTATION PRESENTED TO THE GRADUATE SCHOOL OF THE UNIVERSITY OF FLORIDA IN PARTIAL FULFILLMENT OF THE REQUIREMENTS FOR THE DEGREE OF DOCTOR OF PHILOSOPHY UN IVERSITY OF FLORIDA 2012

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2 2012 Pan Gu

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3 To my grandfather

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4 ACKNOWLEDGMENTS First and foremost I would like to thank my supervisor, Dr. Z.Hugh Fan for his inspiration, guidance, and support during my Ph.D. study. Dr. Fan agreed to take me on as a graduate student, instilled in me a scientific problem solving approach, and provided a comfortable research environment for me I would also like to thank Dr. Toshikazu Nishida from the Department of Electrical and Computer Engineering for his regul ar feedback and evaluation which help ed me to get the right direction. I thank Dr. Corin Segal, Dr.Nancy Denslow and Dr. Subrate Roy for their time to serve on my supervisory committee and provide valuable comments and advice. I would like to thank all my friend s and colleague s at Dr. Z.Hugh. Fan s Research Group for their support and encouragement. I would especially like to thank my collaborator Shancy Augustine who design ed the PCB board for the project and help ed me to make the micro heaters. Finally, I would like to thank my family for their love and support. Without their encouragement, support and comfort I w ould n ot have be en able to finish this dissertation.

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5 TABLE OF CONTENTS page ACKNOWLEDGMENTS ................................ ................................ ................................ .. 4 LIST OF TABLES ................................ ................................ ................................ ............ 8 LIST OF FIGURES ................................ ................................ ................................ .......... 9 LIST OF ABBREVIATIONS ................................ ................................ ........................... 12 ABSTRACT ................................ ................................ ................................ ................... 13 CHAPTER 1 INTRODUCTION ................................ ................................ ................................ .... 15 Microfluidics ................................ ................................ ................................ ............ 15 Microfluidic D evice F abrication ................................ ................................ ............... 16 Silicon and Glass ................................ ................................ .............................. 16 Plastics ................................ ................................ ................................ ............. 17 Microvalves ................................ ................................ ................................ ............. 18 V alve a normally open microvalve. ................................ .................. 19 M icrovalve on non PDMS M icrofluidic D evices ................................ .. 19 Overview and Organization of This Thesis ................................ ............................. 21 2 FABRICATION AND TESTING OF MICRO PLASTIC MICROFLU IDIC VALVES USING PDMS AS AN ELASTOMER ................................ ................................ ...... 25 Valves in Thermoplastic Microchip ................................ ................................ .......... 25 Bonding Thermoplastic with PDMS ................................ ................................ ......... 25 Experiment ................................ ................................ ................................ ....... 26 Reagents and Material ................................ ................................ ............... 26 Bonding m ethod ................................ ................................ ......................... 26 Bonding s trength m easurement ................................ ................................ 27 Discussion ................................ ................................ ................................ ........ 28 Bonding c ondition o ptimization ................................ ................................ .. 28 PMMA versus COC ................................ ................................ .................... 29 Bonding s trength s urvey ................................ ................................ ............ 3 0 Integration and Operati on of COC/PDMS/COC Microvalve ................................ .... 31 Experiment ................................ ................................ ................................ ....... 31 Device integration ................................ ................................ ...................... 31 Device operation ................................ ................................ ........................ 32 Discussion ................................ ................................ ................................ ........ 33 Application Protein Separation ................................ ................................ ............. 34 2 D P rotein S ep a ration ................................ ................................ ..................... 34 The N ecessity of V alves ................................ ................................ ................... 35

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6 Experiment and Result ................................ ................................ ..................... 36 Bonding Mechanism ................................ ................................ ............................... 37 Experiment ................................ ................................ ................................ ....... 37 Discussion ................................ ................................ ................................ ........ 38 Surface t reatment and c haracterization ................................ ..................... 38 Reaction m echanism and b onding ................................ ............................. 39 Summary ................................ ................................ ................................ ................ 41 3 DESIGN AND OPTIMIZATION OF THERMALLY ACTUATED MICROFLUIDIC VALVE ................................ ................................ ................................ .................... 52 Overview ................................ ................................ ................................ ................. 52 Bonding B etween PU and COC ................................ ................................ .............. 53 Experiment ................................ ................................ ................................ ....... 53 Reagents and m aterials ................................ ................................ ............. 53 PU prepar ation ................................ ................................ ........................... 54 Bonding method between COC and PU ................................ .................... 55 Discussion ................................ ................................ ................................ ........ 55 Bond ing c ondition o ptimization ................................ ................................ .. 55 PU versus PDMS ................................ ................................ ....................... 57 Pneumatic Valves Based on PU film ................................ ................................ ....... 58 Experiment ................................ ................................ ................................ ....... 58 Discussion ................................ ................................ ................................ ........ 59 Thermally Actuated Microvalve based on PU film ................................ ................... 60 Experiment ................................ ................................ ................................ ....... 60 Device f abrication ................................ ................................ ...................... 60 Operation ................................ ................................ ................................ ... 61 Discussion ................................ ................................ ................................ ........ 62 Study of PU P roperty ................................ ................................ .............................. 64 Experiment ................................ ................................ ................................ ....... 64 Design ................................ ................................ ................................ ........ 64 Test method ................................ ................................ ............................... 65 Conclusion ................................ ................................ ................................ ........ 65 Modulus ................................ ................................ ................................ ..... 65 Creep ................................ ................................ ................................ ......... 66 Creep recovery ................................ ................................ .......................... 66 Summary ................................ ................................ ................................ ................ 67 4 MULTIPLEXED IMMUNOASSAY IN A VALVE ARRAY SYSTEM .......................... 80 Overview ................................ ................................ ................................ ................. 80 Immunoassay ................................ ................................ ................................ ... 80 Surface M odification ................................ ................................ ......................... 81 Challenges ................................ ................................ ................................ ....... 83 Materials for microfluidics ................................ ................................ ........... 83 Surface modification ................................ ................................ .................. 84 Introduction of samples/reagents ................................ ............................... 84

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7 Immobilization of antibody ................................ ................................ .......... 84 Detection ................................ ................................ ................................ .... 84 Microfluidic design ................................ ................................ ..................... 84 Device Design ................................ ................................ ................................ ......... 84 Assay Design ................................ ................................ ................................ ... 84 Device L ayouts ................................ ................................ ................................ 85 Flow layer ................................ ................................ ................................ ... 85 Heater layer ................................ ................................ ............................... 86 Photobiotin Immobilization ................................ ................................ ...................... 86 Experimental Methods ................................ ................................ ...................... 86 Photobiotin coating on COC surface ................................ .......................... 86 Immunoassay ................................ ................................ ............................. 87 Discussion ................................ ................................ ................................ ........ 87 Solvent e vaporated versus in solution on photobiotin coating .................... 87 Photobiotin coating on different substrates ................................ ................ 88 Immunoassay ................................ ................................ ............................. 88 Immunoassay in a 13 valve controlled microchip ................................ ................... 89 Experimental M ethod s ................................ ................................ ...................... 89 Device assembly ................................ ................................ ........................ 89 Control system ................................ ................................ ........................... 89 Immunoassay process ................................ ................................ ............... 90 Valve operation ................................ ................................ .......................... 90 Discussion ................................ ................................ ................................ ........ 91 Summary ................................ ................................ ................................ ................ 92 5 CONCLUSION AND FUTURE DIRECTIONS ................................ ....................... 106 Conslusion ................................ ................................ ................................ ............ 106 Future Drections ................................ ................................ ................................ ... 108 Valve Fabrication ................................ ................................ ............................ 108 Material ................................ ................................ ................................ .... 108 Valve design ................................ ................................ ............................ 109 Surface M odification ................................ ................................ ....................... 109 LIST OF REFERENCES ................................ ................................ ............................. 112 BIOGRAPHICAL SKETCH ................................ ................................ .......................... 122

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8 LIST OF TABLES Table page 2 1 The bonding strength in devices fabricated by different methods ....................... 42 4 1 List of traumatic brain in jury (TBI) biomarkers. ................................ ................... 93

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9 LIST OF FIGURES Figure page 1 1 Structure difference of anisotropic and isotropic etching. ................................ 23 1 2 Three fabrication methods commonly used for plastic MEMS. ........................... 23 1 3 ................................ ....... 24 1 4 Diagram of two channels: control channel and flow channel. ............................. 24 2 1 T he process flow of bonding a thermoplastic substrate with a PDMS layer (a d), followe d by additional steps for valve fabrication (e g). ................................ 43 2 2 The illustration of the set up of tensile strength test. ................................ .......... 44 2 3 The ef fects of the annealing temperature on the bonding strength. .................. 45 2 4 Two arrays of pneumatic valves actuated under applied pressure. .................... 46 2 5 The pressure required to close valves as a function of the width of control channels. ................................ ................................ ................................ ......... 47 2 6 The procedure of 2 D protein separation. ................................ ........................... 47 2 7 Layout of a microfluidic device designed for 2 D protein separation.. ................. 48 2 8 A device designed for 2 D protein separation. ................................ .................... 48 2 9 IEF separation of two proteins, R phycoerythrin (RPE) and green fluorescent protein (GFP) in the presence of PDMS based valve arrays. ............................. 49 2 10 ATR FTIR spectra of COC in the native form and those after various surface treatments.. ................................ ................................ ................................ ......... 49 2 11 Mechanism of the reaction between PDMS and COC. ................................ .... 50 2 12 (a) XPS spectra of native COC and surface modified COC. (b) Exploded view of the narrow band of C1s region in the XPS spectrum of the surface modified COC. ................................ ................................ ................................ .... 51 3 1 3 dimension structure of a thermally actuated microvalve.. ................................ 69 3 2 The relationship between thickness of PU membrane and spin coating speed.. ................................ ................................ ................................ ................ 70 3 3 P rocess to bond PU and COC film. ................................ ................................ .... 71

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10 3 4 Peeling force between a PU film and COC film as a function of the annealing temperature. ................................ ................................ ................................ ....... 72 3 5 The operation of PU based pneumatic valve arrays.. ................................ ......... 73 3 6 The dye test of a thermally actuated microvalve. ................................ ................ 73 3 7 The cha nges of current and temperature as a function of time. ....................... 74 3 8 The valve actuation time as a function of the input power. ................................ 75 3 9 The strain value of PU film changed with the closing of valve.. .......................... 76 3 10 The experiment set up of the creep test.. ................................ ........................... 77 3 11 Sto rage modulus of PU changes with temperature. ................................ ........... 77 3 1 2 Strain time diagram of PU film under the constant stress of 0.21MPa. .............. 78 4 1 Schematic illustration of the immunoassay principle ................................ .......... 94 4 2 Schematic illustration of an assay procedure ................................ ..................... 94 4 3 The structure of photobiotin used to modify COC surfaces. ............................... 95 4 4 Experimental process of antigen detection. ................................ ........................ 95 4 5 The layout of two channel lay ers designed for immunoassay test ...................... 96 4 6 The layout of two heater layer and their corresponding channel layers. ............. 96 4 7 Photobiot in coating protocol ................................ ................................ .............. 96 4 8 Performance comparison of photobiotin coating using different methods. ......... 97 4 9 Comparison of photobio tin performances on different substrates. ...................... 98 4 10 Immunoassay process on biotinylated COC surface. ................................ ......... 99 4 1 1 Calibration curve of immunoassay on COC surface. ................................ ........ 100 4 1 2 The three dimension model of the valve controlled immunoassay microchip. .. 102 4 13 A assembl ed device made by COC. ................................ ................................ 102 4 1 4 The set up of control system for the multiple valve operation. .......................... 103 4 1 5 The PCB designed for mul tiple valve microchip. ................................ .............. 103 4 16 Valve operation protocol in a 3 TBI biomarker test. ................................ .......... 104

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11 4 1 7 The result of immunoassay detection of 3 biomarkers in mu l tiple valve s controlled microchip. ................................ ................................ ......................... 105 5 1 The 3 dimension design of a 4 6 valves array for 24 analytes detection. ........ 111

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12 LIST OF ABBREVIA TIONS COC c yclic o lefin c opolymer CRP C reactive protein GAPDH Glyceraldehyde 3 phosphate dehydrogenase LOC lab on chip PCB printed circuit board PDMS p olydimethylsiloxane PMMA p oly (methyl methacrylate) PU polyurethane TMSPMA 3 (trimethoxysilyl) propyl me thacrylate

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13 Abstract of Dissertation Presented to the Graduate School of the University of in Partial Fulfillment of the Requirements for the Degree of Doctor of Philosophy DEVELOPMENT AND APPLICATION OF MICROFLUIDIC VALVES By Pan Gu May 2012 Ch air: Hugh Fan Cochair: Toshi Nishida Major: Mechanical Engineering Microfluidics has attracted the attention from a large number of researchers and is a promising technology in biotechnology, engineering and medicine. Among microfluidic devices, those mad e from plastics are advantageous compared to those from silicon and glass because plastic devices can be mass produced, are low cost, and may be disposable. A microfluidic device consists of many critical components, and one of them is the microvalve, whi ch tends to be problematic when integrated with the device. Since commercialization calls for reliability, it is important to develop a plastic microfluidic system with reliable microvalves. This study presents and demonstrates the integration of two type s of microvalves in plastic devices, as well as their applications. A concept of using a valve integrated platform for simultaneous immunoassay detection of multiple analytes is presented. A surface modification method is developed to make COC adaptable t o the biomolecule immobilization. This concept is then demonstrated by the detection of the biomarkers associated with traumatic brain injuries. The primary challenge of integrating valves into a microfluidic device comes from the bonding reliability cau sed by the low plastic surface energy. A strong bond between different plastic materials is a precondition to create elastomer based microvalves. To

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14 address the challenge, we developed a reliable and universal bonding method to assemble thermoplastic mate rials, including poly(methyl methacrylate) and cyclic olefin copolymer (COC), with an elastomer such as polydimethylsiloxane and polyurethane.

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15 CHAPTER 1 INTRODUCTION In this chapter, microfluidics and the device fabrication process will be discussed f ollowed by an introduction of the basic principle and technology of microvalves in microfluidic devices. The overview and organization of this thesis will also be presented Microfluidics Miniaturization and MEMS ( Micro Electro Mechanical Systems) gave bir th to the discipline of microfluidics in the late 1990s. 1 3 The concept of miniaturization was proposed by Richard Feynman as early as 1959. However, in the early days miniaturization was focused on scaling down dimensions to achieve small mechanical structures (MEMS) and high performance integrated circuits 4 5 In the 1990s the concept of miniaturization was expanded due to the development of technology and now it i s possible to miniaturize all kinds of system, e.g. biological fluids, thermal, down to micrometer size. 6 10 MEMS are the miniaturized systems or devices that integrate electrical and mechanical components through microfabrication technology In the past 15 years, the domain of MEMS diversified and MEMS devices were applied for biological, chemical and biomedical fields. 11 13 Those applications employ fluid flows on a scale that was not previously studied before leading to the appearance of a new discipline -microfluidics. M icrofluidics is the science of controlli ng, manipulat ing and testing fluid flows at the micro scale. While originally a subset of MEMS in the early day s today microfluidics is a unique field in itself and is mostly related to lab on a chip ( LOC) or micro total analysis system applications

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16 M icrofluidic D evice F abrication As an extension of the MEMS field, it is not surprising that microfluidic device s in the early age was made from silicon 14 15 and glass 16 17 and the photolithography process is followed for fabricat ion, since it has been well established and widely applied in the MEMS industry for years and is considered as a mature system. O f late, plastics and polymers are employed in microfluidic devices 18 21 and their fabrication process i s different from silicon and glass related technology. Silicon and Glass Silicon and glass associated microfabrication process es are normally carried out in a clea n environment known as a clean room. A microfabrication process based on silicon and glass normally includes the following steps: design, photolithography and etching 22 24 P attern ing is first designed using software such as Autocad. D uring the photolithography process, a mask with the designed pattern i s generated on a quart z plate. Chrome i s deposited using electron sputt er ing. S ometimes, other methods are empl oyed for mask preparation such as a high quality print transparency 25 26 if the resolution requireme nt is greater than 10 mi crometer s. The pattern on the photomask i s then transferred to a silicon or glass wafer covered with a photosensitive layer, using the process of exposure and development. By this step, the pattern is formed in the photoresist layer on the silicon or glass substrate. E tching process will then carve the pattern into the substrate. 27 30 T here are two types of etching: isotropic and anisotropic. 31 34 They are catalogued with the selectivity of etching direction during the etching process. T he isotropic etching is unselective, and it will be carried out i n all directions with the same rate. In contrast,

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17 anisotropic etching is known as selective since it prefers to be carried out along one crystallographic plane, and the etching rate will be faster on this plane rather than others. T he shapes of cavities fo rmed by different etching methods are different. A s shown in F igure 1 1, the isotropic etching could be used to generate a spher ical cavity or a cavity with a round angle on the sidewall surface while anisotropic etching leads to clear facets on the side w all. W e need to note that anisotropic etching could be possible only in a material which contains more than one crystallographic plane Thus, only isotropic etching could be realized on a glass y material due to its amorphous structure. Plastic s P lastic dev ice s play a more and more important role in the domain of microfluidics in recent years because of their low cost, high manufacturing output, fast replication, optical clarity and flexible deformable structures. P lastic devices are disposable due to their cheap price. C ompar ed with the complicated fabrication process on silicon and glass, a plastic microstructure is easier to generate and mass produce using the existing technology employed in the modern manufacturing industry. A s shown in F igure 1 2, the re are three current methods of plastic replication (1) Casting, a mixture of catalyst and monomers is poured on the top of a mold and peeled off after curing under room temperature. P olydimethylsiloxane (PDMS ) devices are commonly prepared using this me thod. (2) Molding. T his method is also known as hot embossing. A thermoplastic is heated above its glass transition temperature (T g ) and pushed into a mold under the required pressure. A fter cooling and separation a structure is obtained. (3) Injection Molding. A heated thermoplastic in liquid form i s injected into a mold cavity and fill s all of the space in the cavity N ote that all the structures obtained by these three methods are the negative of the mold shape.

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18 Molds are generally fabricated using the photolithography technology involved in the silicon and glass related patterning process. F or the cas t ing technique it is possible to use polymer molds. F or the method of molding and microinjection, a metallic mold, often generated by electrodeposit ion is usually necessary since it needs to withstand the high temperature and pressure used during the fabrication process. Microvalves W ith the recent success of the human genome project and the potential of bioMEMS, microfluidic systems are increasingl y discussed for its possibility to obtain a commercial success in biomedical applications. 35 However, this commercialization process is often not very smooth. O ne main reason i s the lack of reliable components such as pumps, valves and mixers which are necessary to integrate a successful Lab On a Chip ( LOC ) platform. T his thesis will focus on microvalves which will address the problem M icrova l ves can be roughly classified into passive and active ones based on the actuation principle as discussed in the literature 36 A n active microvalve i s driven by an external source M ost active microvalves contain a membrane layer which can be deformed under a stimulus in the form of magnetic 35 electrostatic 37 electrokinetic 38 39 piezoelectric 40 bimetallic 41 thermopneumatic 42 48 modular and pneumatic 49 50 etc. Besides, some active microvalves can be actuated by the utilization of smart materials such as phase change hydrogel or sol gel 51 52 An passive microvalve is driven by the inherent principle of the valve system, for example, by the geometries of the microchannel and capillarity based on the material s surface property 53 54

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19 In addition microvalve s may also be divided into normally open and normally closed valves depend ing on its initial status. I n the next section, a normally open microvalve will be introduced. Quake s V alve a normally open microvalve 49 Quake s group proposed the concept of a PDMS valve in 2000. I n the last ten year s, this valve has become more and more popular and widely used in microfluidic applications. T h e structure of Quake s valve is shown as Figure 1 3. I t mostly consists of three layers: the flow layer, control layer and thin film layer. The f low layer and the control layer contain microchannels that were obtained by the molding technique as described previously. T hree layers were integrated together in the sequence of control layer ( top), thin film ( middle), and flow layer ( bottom). B onding between PDMS and PDMS can be easily realized with plasma treatment on both bonded surfaces. Quake s valve was ope rated as shown in Figure 1 4. T he upper chann e l was the microchannel in the control layer and the lower channel in the flow layer was the microchannel filled with testing fluids. The valve was normally open and fluids in the flow layer could flow freely in the channel without any obstruction. T o close the valve a pressure is applied to the control layer using external equipment. T he flexible thin film elastomer separating the flow layer and the control layer is deformed under the externally applied press ure and, if the pressure is high enough, the microchannel is blocked in the flow layer. F luids will re flow again inside the channel when the valve is re opened by turning off the pressure. Quake s M icrovalve on non PDMS M icrofluidic D evices A fter research group reported elastic membrane based PDMS valves in 2000, several high throughput applications were developed based on their research

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20 using thousands of integrated microvalves B ut most of those studies used pure PDMS material, due to the follow ing reasons Firstly l ow Young s modulus of PDMS materials. S ince the membrane will be deformed to close the microchannel during the valve closing process, the flexibility of the membrane is a critical property which determines the minimum pressure requir ed. The Y oung s modulus is a reflection of the flexibility of the material A low Young s modulus material is preferred in Quake s microvalve concept. T he Young s modulus of PDMS is around 750kPa, more than one magnitude smaller than other elastomer materi als such as rubber Secondarily, convenient bonding method and strong bonding effect between PDMS and PDMS. C onsidering a high pressure applied during valve operation, Quake s valve s require a strong bond between two layers. A normal microfluidic system in cluding microvalves will be operated under a pressure between 0 to 25psi; while PDMS/PDMS bond with surface modification of plasma treatment can sustain a pressure as high as 50 psi, which is high enough for most microfluidic applications. The elastomer v alve has been adopted by a number of researchers A fter solving the bonding problem between PDMS and glass, with a glass device to form glass/PDMS/glass structures for DN A analysis and other applications. 55 57 However, the construction of a plastic microfluidic system including microvalves is not fully implemented mainly due to th e challenge of development a reliable bonding method between the plastic and the elastomer material. T his thesis aims to provide a bonding solution between the plastic and elastomer material, and to demonstrate integrated microvalve system for a multiplexe d immunoassay array

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21 Overview and Organization of This Thesis The objective of this work is to fabricate a pneumatic and a thermopneumatic microvalve in a plastic microfluidic device, and to achieve simultaneous detection of multiple analytes in a microfl uidic device through successful valve operation. A pneumatic microvalve in a miniaturized device is composed of a flow layer, elastomer layer and control layer, driven with compressed air A thermoelastic microvalve is composed of a flow layer, elastomer layer, valve layer and heater layer, actuated by expansion of a fluid in a cavity layer that sheeted by a resistor on the heater layer when an external power is applied. An immunoassay was carried out in a microchip integrated with multiple thermally actu ated valves. T he reliability of the valve function under long time operation is also dem o nstrated. The thesis is organized as follows. In Chapter 2, bonding between PDMS and different kinds of thermoplastic was studied and discussed As a proof of the bo nding effectiveness, a structure of two pneumatic microvalve arrays was evaluated on a 2 D protein separation device. In Chapter 3, bonding between polyurethane (PU) and thermoplastic was evaluated as the basic technique for a thermopneumatic microvalve i n a plastic microfluidic device. A thermopneumatic microvalve was fabricated to prove the overall concept and the operation, while the power input was controlled using a labview program. In Chapter 4, a 13 microvalve microfluidic device was designed, fabr icated and characterized for a multiplexed immunoassay array. A surface modification method was developed to enable the immunoassay detection. The detection of C reactive protein ( CRP ) and transferrin was demonstrated in this platform indicat ing a potenti al of large scale multi analyte detection.

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22 In Chapter 5, the conclusion of the dissertation and the future directions are presented

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23 F igure 1 1. Structure difference of anisotropic and isotropic etching. (a) Comparison between anisotropic struc ture and isotropic structure. (b) Anisotropic wet etching lead to faceted forms. (c) Isotropic wet etching lead to round form. Figure 1 2. Three fabrication methods commonly used for plastic MEMS. (a) (b) (c)

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24 Figure 1 3. Basic structu re of valve developed by Quake s group. Figure 1 4. Diagram of two channels: control channel and flow channel. W hen the pressure applied into the control channel, the membrane existing between the control channel and flow channel is deformed and the fl ow channel is closed.

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25 CHAPTER 2 FABRICATION AND TEST ING OF MICRO PLASTIC MICROFLUIDIC VALVES USING PDMS AS AN ELASTOMER Valves in Thermoplastic Microchip Thermoplastics have been increasingly employed for microfluidics applications. 58 Compared with glass and silicon, thermoplastics offer several advantages including manufacturab ility, low cost, and biocompatibility. 18 58 60 Among several thermoplastics, cyclic olefin copolymers (COC) and polymethylmethacrylate (PMMA) are frequently exploited for making microfluidic devices. These devices were often fabricated by bonding a cover sheet with a substrate containing microchannels and other microfeatures using various bonding methods including thermal fusion, 18 60 63 solvent bonding, 64 70 surface treatment, 71 76 and adhesives. 77 78 Each of these methods has advantages and disadvantages as reviewed in the literature. 58 59 M icrovalves in a microfluidic system are used to regulate flows, contain fluids, and isolate one region from the other 36 79 In C hapter I describe fabrication of the elastomer valve in a COC device with a focus on the optimization of the PDMS/COC bonding. The valves were designed for developing two dimensional protein separation device s, which were reported using two layer COC devices with gel based pseudo valves. 80 To realize the goal of integrating ela stomer based true valves in the device, the key challenge is to achieve strong bonding between COC and PDMS so that the device will not delaminate when a pressure is built up after valves are closed. Bonding Thermoplastic with PDMS The method we develo ped to bond COC with PDMS is based on surface modification using 3 ( trimethoxysilyl ) propyl methacrylate (TMSPMA). TMSPMA has been used as a coating on glass slides 81 and for grafting on fiber/polypropylene A part of this chapter has been published in (1) P. Gu, K. Liu, H. Chen, T. Nishida and Z. H. Fan, Analytical Chemistry 2010, 83 446 452. (2) K. Liu, P. Gu, K. Hamaker and Z. H. Fan, Journal of Colloid and Interface Science 2012, 365 289 295.

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26 composites, 82 but never be exploited for facilitating thermoplastic bonding. Prior to the chemical treatment, surfaces to be bonded were activated using corona discharge. Thermal annealing was then performed to facilitate the formation of chemical bonds between functional groups on the surfaces and enhance the bonding strength. We found that the method is also applicable to PMMA. The bonding strength between COC and PDMS was represented by the peeling force, which could be quantitatively measured. Experiment Reagents and m aterial Cyclic olefin copolymer (COC) films (Zeonor 1420R, 188 m thick) and COC resins (Zeonor 1020R) were purchased from Zeon Chemicals (Louisville, KY) while a 100 m thick COC (Topas 8007) was from PLITEK (Des Plaines, IL). Poly(methyl methacrylate) (PMMA) films (250 and 100 m thi ck) were obtained from Evonik Industries (Essen, German). Polydimethylsiloxane (PDMS, Sylgard 184) was bought from Dow Corning (Midland, MI). A solution of 98% 3 ( trimethoxysilyl ) propyl methacrylate (TMSPMA) was purchased from Acros organics (Fair Lawn, NJ ) while solvents and common chemicals were from Fisher Scientific (Atlanta, GA). A corona discharge generator (BD 10A) was purchased from Electro Technic Products Inc. (Chicago, IL). Bonding m ethod The process flow of bonding a thermoplastic substrate wit h a PDMS layer is shown in Figure 2 1. A mixture of PDMS pre polymer was prepared according to the instructions of the manufacturer. After eliminating air bubbles via vacuum, the mixture was spin coated onto a clean 100 m thick Topas film (or 188 m th ick Zeonor 1420R

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27 when high annealing temperature was used) using a spinner (Laurell Technologies) at a speed of 4000 RPM for 30 seconds (Figure 2 1a). The Topas thin film served as a sacrificial layer (discarded later), facilitating the formation of a thi n layer of PDMS via spin coating. PDMS on the sacrificial layer was cured in an oven at 60 C for 4 hours. The resulting PDMS thickness was 15 m, measured by a profilomer (Dektak IIa, Veeco Instruments). A plastic sheet (188 m thick Zeonor 1420R) cont aining microchannels was first activated by corona discharge treatment, followed by immersing it in a solution of 6% TMSPMA for 20 minutes (Figure 2 1b). After the sheet was taken out, it was blow dried using nitrogen. The sheet was then placed in contac t with the PDMS/sacrificial layer (prepared in Figure 1a) after both surfaces were activated by corona discharge (Figure 2 1c). The assembly was placed in an oven at 90 C overnight for annealing, which facilitated the formation of chemical bonds between activated functional groups on the surfaces and promoted bonding. At the end of annealing, the sacrificial layer was removed and irreversible bonding of thermoplastics/PDMS was achieved (Figure 2 1d). Bonding s trength m easurement The bonding strength be tween thermoplastics/PDMS was measured in order to study the effects of different materials and bonding conditions. The standard peel test established by the International Organization for Standardization (ISO) was employed to determine the bonding streng th. 83 Samples were prepared by following the ISO 180 peel test protocol and i llustrated in Figure 2 2a. The size of the COC film is 250 mm 25.4 mm and the bonding area of PDMS is 25.4 mm x 25.4 mm. The COC film was bonded with PDMS using the procedure described above. The COC/PDMS assembly was then bonded to a plastic plate tha t was fixed to the bottom grip of a two column,

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28 floor standing tensile machine made by MTS Systems Corp. (Eden Prairie, MN). The COC film was bent at 180 and the other end of it was attached to the top grip, which was programmed to be pulled up at a cros shead speed of 5 mm/minute (Figure 2 2b). The pulling force was measured by a load cell that was mounted at the top grip. The load cell had a maximum force of 100 N with a detection sensitivity of 0.02 N. Figure 2 2c shows an example of a plot that recor ded the force as a function of the moving distance of the load cell. Large fluctuations near the edges of the device (the initial and ending points) were not used for calculation per the recommendation by ISO. 83 The stable peeling forces in the middle were employed to calculate the average peeling force as indicated in Figure 2c. E ach bonding condition was studied three times to enhance the precision. D iscussion Bonding c ondition o ptimization The bonding strength, represented by the peeling force, is an objective and quantitative indicator. Thus it can be used to determine the ne cessity of each step in a bonding protocol as well as the optimum condition of each step required. The bonding protocol described in the Experimental Section for COC/PDMS was obtained after a large number of experiments. An example of the bonding conditi on optimization is the study of the annealing temperature. After corona discharge activation and TMSPMA treatment, COC was placed in contact with PDMS, followed by annealing at a certain temperature. During annealing, chemical bonds could be formed betwe en the activated functional groups on the surfaces of the thermoplastic and PDMS layers. It also promoted bonding as suggested by our experimental results.

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29 Figure 2 3 shows the effects of the annealing temperature on the bonding strength. The results i ndicate that the peeling force required to delaminate COC/PDMS linearly increased with the annealing temperature from 50 C to 90C. The bonding strength was negligible when annealing took placed at a lower temperature. We also found that the corona disc harge exposure is necessary before TMSPMA treatment. Our results indicated that the corona discharge exposure increased the bonding strength by about 20% compared to those without corona discharge treatment. ctive, we also carried out the blister test. 84 This test was carried out by observing dye spreading when delamination took place and the two layers formed bulges at the edges of a dye filled channel. The maximum pressure a device can sustain without blistering is an important pa rameter, especially for a microfluidic device containing valves since bonding must withstand the holding pressure when valves are closed. We did not observe any blistering in valve containing channels up to 100 psi (~689 KPa) when the optimized bonding pro tocol was used. The device might sustain an even higher pressure but our setup was designed to have a maximum pressure of 100 psi as discussed in the Experimental Section. Considering the fact that most microfluidic devices are operated under 50 psi, 85 we conclude that the chemical assisted COC/PDMS bonding is well suited for microfluidic applications. PMMA versus COC We also determined if TMSPMA assisted bond ing is applicable to another thermoplastic material, PMMA, which is used extensively in microfluidics. The bonding strength measurement suggested that (1) TMSPMA assisted bonding is applicable to

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30 PMMA and (2) there are some differences in the required ste ps and the optimum operation parameters between COC and PMMA. One difference is shown in Figure 2 3, in which the peeling forces are plotted as a function of the annealing temperature. Although PMMA/PDMS showed a similar trend that the peeling force line arly increased with the annealing temperature, the temperature required was much lower than COC/PDMS. Significant bonding between PMMA/PDMS was obtained even after overnight annealing at room temperature (25 C) whereas COC/PDMS gained the similar degree of bonding when annealing took place at 55 C. The PMMA/PDMS bonding strength reached a plateau when the temperature was higher than 70C. The difference in the required annealing temperature is likely related to their plasticity, since PMMA has a lower glass transition temperature than COC. It may also be related to the difference in their chemical structures and surface property. The second difference is that PMMA did not require corona discharge before chemical treatment. We did not observe any sign ificant improvement in bonding strength if it was performed for PMMA/PDMS bonding. As mentioned above, it was required for COC when it was bonded with PDMS. This difference likely resulted from the existence of acrylate functional groups in PMMA. Howev er, corona discharge treatment was required for both COC and PMMA after chemical treatment since it enhanced the bonding strength. Note that these are experimental observations, and further theoretical reasoning will be explored in the future. Bonding s tre ngth s urvey Since the peeling force measured here is an excellent indicator of the bonding strength between two layers in a microfluidic device, we measured and compared a

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31 number of bonding methods that are related to PDMS and often used in the microfluidi cs community. Table 2 1 lists the bonding materials and methods we bonding area. Soft bonding between PDMS/glass refers to the situations where plasma treatment was not used so that PDMS can be peeled away from a glass substrate for reuse. In hard bonding between PDMS/glass, plasma treatment was carried out on the surface of PDMS and permanent bonding between the two were generally realized. 85 We used the corona discharge to treat PDMS surface as previously reported. 86 87 Our results suggested that long term annealing resulted in stronger bonding between PDMS/glass. Annealing for 3 days at room temperature produced bonding strength almost twice higher than that after overnight annealing. Usin g PDMS prepolymer as glue, PDMS/PDMS generated strong bonding as expected. For comparison, we also included in the table the bonding strength between PMMA/PDMS and COC/PDMS at the optimum bonding conditions as discussed above. Integration and Operation of COC/PDMS/COC Microvalve Experiment D evice integration To fabricate a microvalve, a control layer (a 188 m thick COC sheet with microchannels) and PDMS were bonded as discussed above. Channels in the control layer were fabricated by laser ablation (JPS A IX 260 ArF excimer laser). The width of the channels was either 90 m or 150 m. Alternatively, channels in the control layer were fabricated by a computer numerically controlled (CNC) milling machine (LPKE ProtoMat S100) and the designed channel widt h was 250 m or 300 m.

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32 The fluid layer in the microvalve was fabricated using compression molding as reported previously. 61 Briefly, a pattern of microchannels was created in a glass plate via photolithography. Electroplating on the glass plate generated a nickel mold, which was employed to produce plastic parts from COC resins (Zeonor 1020) using a hydraulic press (Carver, Wabash, IN). The surface of the fluid layer was activated by corona discharge, followed by TMSPMA treatment as discussed above (Figure 2 1e). T he fluid layer was then placed in contact with the PDMS/control layer assembly after both surfaces were treated with corona discharge (Figure 2 1f). The channels in the control layer were aligned to the desired valve locations in the fluid layer. Afterwa rds, the 3 layer assembly was annealed in an oven as described above. When a pressure was supplied in the control layer, PDMS deflected and blocked the channel in the fluid layer (Figure 2 1g). D evice operation As illustrated in Figure 1g, a microvalve wa s actuated by a pressure supplied in a channel in the control layer. A pressure supply system was connected to the inlet of the control channel via an Upchurch Nanoport (Oak Harbour, WA). The pressure system was built in house, similar to what was report ed previously. 88 It consisted of a pressure controller (Cole Parmer 68502 10), a nitrogen cylinder, and a Dynamco dash valve (Model D1B1202, Commerce, GA) that was controlled by a DC power supply (Model E 3630A, Agilent Technologies, San Jose, CA). The pressure at the nitrogen cylinder was set at 100 psi, and the actual pressure to the device was fine tuned by the dash valve for values ranging from 0 to 100 psi. The DC power supply for each dash valve was controlled by a computer.

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33 To show the operation of a microvalve in the device, a solution of a food dye was introduced into channels in the fluid layer. Operation of the valve was indicated by the disappearance of the dye in the channel when the valve w as closing. An inverted microscope (IX51, Olympus) equipped with a Hamamatsu CCD camera (model C4742 80 12AG) was used for recording the dye signals. Discussion As mentioned above, COC/PDMS bonding was developed for fabricating a reliable valve array for protein separation. Figure 2 4a shows a picture of a COC substrate with the layout we used for two dimensional protein separation. 80 The device consists of one AB channel for the first dimension (IEF) and a number of CD channels for the second dimension (PAGE). Since different separation media are employed to achieve different separation mechanisms in each dimension, a valv e array is required at the intersections to prevent two separation media from contaminating each other. The valve array is represented by the dashed lines on each side of the IEF channel in Figure 2 4a. To demonstrate the operation of two valve arrays, we created two 300 m wide channels in the control layer. These two channels were pneumatically controlled using the pressure supply setup described in the Experimental Section. These control channels were not aligned with the IEF channel in this experimen t for simplicity. All channels in the fluid layer were filled with a food dye solution. Figure 2 4b shows an exploded view picture of five channels when valves were open (no pressure in control channels). Rough surface in the control channels was due t o the fabrication method used in this work; they would disappear if we fabricated them using the same method for fabricating the fluid channels. We then supplied a pressure of 30 psi into the control channel on the right side; the valves on the right side were closed as indicated in Figure

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34 2 4c. When the pressure was also applied to the control channel on the left, both valve arrays were then closed (Figure 2 4d). When the pressure in the control channels was relieved, valve arrays were opened again and dye flowed back into the channels (Figure 2 4e). Cyclic operations of valve closing opening were demonstrated. Furthermore, we found that the same device properly closed and opened 6 months later, without any indication that the bonding failed. The dept h of microchannels in our device is 45 m, and the width is 110 m. As a result, the aspect ratio (depth/width) is 0.41. This number is much higher than what is typically used in elastomer had an aspect ratio of 0.1. 3 Deeper channels with higher aspect ratios are harder to close than those with smaller aspect ratios, thus requiring a h igher pneumatic pressure. In addition, the pressure required to close valves in a fluid channel is also related to the width of the control channels. 3 A wider control channel possesses more actuation area, thus it requires a lower pressure to close the valves. We studied four different control channels widths and determined the minimum pressure required to block the same fluid channel. The r esult in Figure 2 5 shows a power relationship between the control channel width and the pressure required to close the valves. This result offers another guideline in the design of microvalves. Application Protein Separation 2 D Protein Separation Proteo mics is a branch of biotechnology which stud the structure, function and interaction of the proteins produced by the genes of a particular cell, tissue, or organism 89 Accompanying with the development of proteomics for modern drug

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35 diagnostics 90 92 two dimensional (2D) protein separation for an efficient implement ing technology is more and more studied 80 93 A 2 D protein separation consists of a first dimensional separating using integrating isoelectric focusing (IEF) and a second di mensional separation using polyacrylamide gel electrophoresis (PAGE) Proteins were separated by their various isoelectric points (pI) in the first dimension, and then sorted by their different molecular weights (MW) in the second dimension. Fig ure 2 6 sh ows an example of this process. A 2 D protein separation device was prepared as shown in Fig ure 2 6 (a). AB direction is the first dimension and CD direction is the second dimension. A monomer solution was filled into the whole device (Fig ure 2 6 (b)) and po lymerized to form a gel excluding channel AB ( Fig ure 2 6 (c)). The unpolymerized monomer in channel AB was then replaced with IEF medium, forming a pH gradient (Fig ure 2 6 (d)). Fig ure 2 6 (e) shows the IEF process. A voltage was applied on channel AB and fo rmed an electric field between point A and point B. Proteins were focused under the influence of electric field based on values of their pI. In this example, two focused protein bands were formed and these proteins were then transferred into the second dim ension. In Fig ure 2 6 (f), the arrow shows the flow directions towords the second dimension. The dash line indicates the position of detecting camera. The N ecessity of V alves A fter the second dimension separation, thousands of detected protein s pots can be compiled into a 2 D gel image called 2D map. However, there is a challenge in the separation media in the first and second dimensions I n the 2 D protein separation process, the protein mixture was driven through the IEF channel and the PAGE c hannel by applying an electric potential between two ends

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36 of channels. However, due to the geometry of the device, electric field varies around ever y intersection of cross channel. The potential existing in the vertical channel bring problems to IEF separa tion. A simulation result show n that fluid sample can get into PAGE channels under the effect of electric field when running IEF process. T hese unfocused and unseparated proteins are still the uncertainty for second dimension separation. An array of micro valves is added to the device designed for the two dimensional protein separation to address this problem. Figure 2 7 shows the device layout (a) and the pneumatically actuated valves next to the IEF channel s (b) In addition, valve arrays are applied on b oth sides of first dimension channel to avoid any contamination during the IEF process. ( Fig ure 2 7 (c)). Experiment and Result The image s of the fabricated protein separation microfluidic device are shown in Fig ure 2 8 The fluid layer consists of one IE F channel and 29 PAGE channels as reported previously. 80 PDMS based elastomer valves were fabricated using the bonding pro cedure described above. The control layer was made in PDMS since it contained two control channels that must be aligned with the IEF channel in the fluid layer. The device was rinsed with 1% KOH and DI water and the channel surfaces was coated with 2.3% hydroxypropyl cellulous solution. When the valve arrays were closed, IEF separation medium was then introduced into the IEF channel. The separation medium was comprised of 2% carrier ampholytes, 8% glycerol, 2.3% hydroxypropyl cellulous (MW 80,000), and two proteins, GFP (60 ng/L) and RPE (40 ng/L). 80 10 mM acetic acid and 10 mM ethanolamine served as anolytes and cathol ytes,

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37 respectively. An electric field of 150 V/cm was applied across the IEF channel for 1 minute. Detection was carried out by an inverted epifluorescence microscope (Olympus IX51) equipped with a charge coupled device (CCD) camera (Hamamatsu C4742 80 1 2AG) using a 10x objective, similar to what we reported previously. 94 After two valve arrays were placed on both sides of the IEF channel as indicated in Figure 2 9 we demonstr ated separation of two model proteins, R phycoerythrin (RPE) and green fluorescence protein (GFP), both of which are naturally fluorescent. The pH gradient was established using carrier ampholytes with a pH gradient of 3 10. The separated proteins were de tected using a CCD camera installed on a microscope. The separation pattern obtained is the same with what we observed when identical proteins were separated in a single channel 21 or when a pseudo valve array was used. 93 This result suggests that elastomer based valves can be exploited for two dimensional protein separation. Bonding M echanism Experiment A ttenuated 95 total reflect ion Fourier transform infrared (ATR FTIR ) spectroscopy was us ed to investigate the mechanisms of the physical and chemical treatment on the COC surfac e All spectra were collected over a range of 650 4000 cm 1 in a Magna 760 spectrometer (Thermo Scientific, USA) equipped with a ZnSe ATR unit (Gemini, Spectra Tech). For each spectrum, 64 scans were obtained at 45 angle of incidence with a resolution o f 4 cm 1 The spectrum from a blank ATR cell was used as a reference. The c hemical composition on the surface was analyzed using an X ray photoelectron spectrometer (PHI5100, Perkin Elmer) with a m agnesium X ray source A part of this chapter has been published in (1) K. Liu, P. Gu, K. Hamaker and Z. H. Fan, Journal of Colloid and Interface Science 2012, 365 289 295.

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38 (1253.6 eV). The Mg anode was oper ated at 20 mA and 15 kV (300 W). The p ass energy and take off angle were 89.5 eV and 45, respectively. Discussion Surface t reatment and c haracterization Attenuated total reflection Fourier transform infrared (ATR FTIR) spectroscopy was employed to st udy the COC surface property after physical and chemical treatment. We obtained the spectra of COC surfaces in the native form and after each st ep of surface treatment. Fig ure 2 10( a ) shows ATR FTIR spectrum of the native COC while that of the same surfa ce after activation using corona discharges is in Fig ure 2 10( b ) Note that ATR FITR spectra should be obtained immediately after the surface treatment, so that the surface property restoration over time 96 could be avoided. Comparison of the spectra in Fig ure 2 10( a ) and 2 10( b ) shows an increase in the signals of the C H stretches at the wavenumbers of 2915 cm 1, 2855 cm 1 and 1455 cm 1. This increase indicates enhanced surface energy. However, we did not observe a peak around 3400 cm 1 associated with the pr esence OH groups. Several studies on surface treatment using plasma or UVO suggested the creation of OH groups, though these efforts were on PDMS or PMMA. 73 76 After the corona discharge treatment, COC was exposed to TMSPMA After removing unbound reagents, the COC sheet was subjected to ATR FTIR and the spectrum is shown in Fig ure 2 10( c ) The sharp peaks of C O and Si O bonds at 1165 cm 1 and 1085 cm 1 suggest the presence of silicon containing organic molecules on the COC surface. The stretching bands at 1720 cm 1 and 1640 cm 1 are the characteristics of C=C and C=O bonds in the methacrylate portion of TMSPMA. The chemical structure of TMSPMA is s hown in Fig ure 2 11( a ) For comparison, the chemical

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39 structu re of APTES is shown in Fig ure 2 11( b ) APTES was used as a silane agent in the several surface modification studies 73 97 99 and the major difference between TMSPMA and APTES is the double bonds in methacrylate. After the chemical treatment, COC surface was exposed to the second dose of corona discharges, followed by ann ealing at 80 C overnight. The resulting ATR FTIR spectrum of the COC surface is shown in Fig ure 2 10( d ) The Si O Si bridge s between neighboring monomers and some C O C in their pristine ester proportions can be taken into account for the broad band ar ound the wavenumber of 1130 cm 1 A wide hump near 3400 cm 1 indicates the transformation of some of Si OCH3 to Si OH during this surface modification step. The decrease in the peak areas at 1720 cm 1 and 1640 cm 1 indicates the reduction of C=C and C=O double bonds due to reactions to the COC surface and cross linking among TMSPMA molecules. R eaction m echanism and b onding The ATR FTIR spectra discussed above led us to propose the reaction mechanism illustrated in Fig ure 2 11 While COC is a saturated hydrocarbon polymer as shown in Fig ure 2 11( c ) the C=C bonds in the methacrylate moiety of TMSPMA is less sustainable under the electrical arc induced by the coronal discharge, which has an output ranging from 10 kV to 50 kV according to the manufacturer. Therefore, radicals are likely induced in the T MSPMA coating as shown in Fig ure 2 11( d ) These radicals would initiate grafting polymerization to the COC substrate, in a way similar to those reported in photo initiated grafting in a plastic surface. 100 101 The structure after the grafting polymerization process is shown in Figure 2 11( e ) To confirm the c ovalent bond between TMSPMA and COC, we performed the following experiments. A treated COC surface was extensively rinsed using ethanol to

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40 remove any non reacted TMSPMA molecules. The sample is then dried, followed by XPS analysis. The obtained XPS spec trum is shown in Figure 2 12( a ) and we observed a strong peak at 118 eV for Si (2p). In contrast, there is no Si peak in the XPS spectrum of a native COC surface. This result clearly indicates that TMSPMA molecules have been successful grafted onto the COC surface. In addition, there is a significantly stronger peak at 537.5 eV for O (1s) in the treated COC than in the pristine COC spectrum. The increase is also attributed to the covalent bonding of TMSPMA molecules to the COC surface. We also observ ed a large decrease in the C (1s) signal, which indicates partial conversion of Si OCH3 to Si OH as indic ated in Fi g ure 2 11( d ) and 2 11( e ) The exploded view of this carbon peak is shown in Figure 2 12( b ) and it is resolved into three subtle peaks, unve iling the distinct chemical environments associated with carbon atoms in the scanned area. The bonding mechanism between the treated COC with PDMS is shown in Fig ure 2 11( f ) while the surface of the activated PDMS is illustrated in Fig ure 2 11( g ) It is w ell studied that silanol groups can be formed on the surface of PDMS via high energy activation methods such as plasma, UV ozone treatment, or Tesla coil (Fig ure 2 11( g ) ). 86 102 103 When the activated PDMS surface was placed in contact with the TMSPMA treated COC, dehydration reactions took place between hydroxyl groups on two surfaces and Si O Si bond were formed (Fig ure 2 11( f ) ). An annealing process at a high temperature for a period of time helped complete the dehydration reaction and promote bonding between COC and PDMS.

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41 Summary A reliable method has been described in this chapter to bond thermoplastics with PDMS. The optimized protocol resulted in a strong bond; the bonding strength is higher than those using several bonding methods associated with PDMS. With microfluidic valves in place, the CO C/PDMS bond could sustain at least a pressure of 100 psi. After the integration of t wo valve s in a two dimensional device protein separation was enabled in one dimension without the possibility to interfere with the other dimension. A large number of re ports exist on a variety of bonding methods and materials. However, it is difficult to compare them and tell which one would produce the strongest bond. The bonding strength measurement using a method established by ISO offered an objective and quantita tive indicator for comparing different methods as we demonstrated in this work. More concerted efforts are needed in this front to generate a consensus and guidelines. Universally accepted bonding methods are extremely important for commercialization of m icrofluidic devices since the fabrication reliability is essential.

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42 Table 2 1. The bonding strength in devices fabricated by different methods Bonding Materials and Methods Peeling Forces Soft bonding between PDMS and glass 0.02 N Hard bonding between PDMS and glass, overnight annealing (1.9 0.3) N Hard bonding between PDMS and glass, 3 day annealing (5.7 0.5) N Hard bonding between PDMS and PDMS (10.3 0.6) N Chemical assisted bonding between PMMA/PDMS (24 2) N Chemical assisted bonding between COC/PDMS (29 3) N

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43 Figure 2 1.The process flow of bonding a thermoplastic substrate with a PDMS layer (a d), followed by additional steps for valve fabrication (e g). (a) PDMS prepolymer was spin coated on a sacrificial plastic layer, follo wed by curing. (b) The surfaces of a plastic sheet were activated, followed by treatment with TMSPMA The plastic sheet was represented by a control layer containing a microchannel in the direction into the paper. (c) After activation, PDMS and the plast ic sheet were in contact, followed by annealing. (d) Irreversible bonding of thermoplastics/PDMS was formed and the sacrificial layer was removed. (e) The surface of a fluid layer containing microchannels was activated and treated with TMSPMA. (f) The f luid layer was bonded with PDMS/control layer as in (c). (g) The valve was closed when a pressure was supplied to the channel in the control layer and PDMS deflected to block the channel in the fluid layer. The drawing is not to scale.

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44 Figure 2 2 Th e illustration of the set up of tensile strength test. (a) Three dimensional view of a sample prepared for bonding strength measurement. A 15 m thick PDMS was bonded with a 188 m thick COC film using the procedure described in Figure 1. The assembly wa s then bonded to a 4 mm thick plastic plate, which was then fixed to a peel test machine. (b) Schematic showing the plastic plate was fixed on the bottom grip of the peel test machine and kept stationary. The COC film was attached to the top grip that wa s programmed to move up at a speed of 5 mm/minute. The force was measured by a load cell attached to the top grip. (c) A representative plot showing the force as a function of the traveling distance of the load cell. The range of the data used for calcu lating the peeling force is indicated.

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45 Figure 2 3. The effects of the annealing temperature on the bonding strength. The bonding strength is indicated by the peeling force required to delaminate thermoplastics/PDMS. Thermoplastics include COC (closed c ircles) and PMMA (open circles). Each line is the linear regression of the data points within the line. Each value is the average of three repeat experiments. Error bars indicate one standard deviation.

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46 Figure 2 4. Two arrays of pneumatic valves act uated under applied pressure. (a) Picture of a COC substrate that functioned as a fluid layer. The device was designed for two dimensional protein separation. Channel AB is for the first dimensional (IEF) while channels CD are for the second dimension (P AGE). Valve arrays are indicated by the dashed lines on the both sides of the IEF in the atmosphere pressure and five fluid channels filled with a dye solution. (c) Valves o n the right side were closed when a pressure was supplied through the right control channel. (d) Both valve arrays were closed when a pressure was applied to both control channels. (e) Both valve arrays were opened again and dye flowed back to channels when the pressure in control channels was relieved.

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47 Figure 2 5. The pressure required to close valves as a function of the width of control channels. The fluid channel dimension was the same in all experiments. The line is the best fit power relatio nship among experimental data. Figure 2 6 The procedure of 2 D protein separation

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48 Figure 2 7 L ayout of a microfluidic device designed for 2 D protein separation. T wo arrays of valves were added on both sides of IEF ch annel. Figure 2 8 .A device designed for 2 D protein separation. The channel between AB is the IEF channel Channels between CD are PAGE channel. T he depth of channel is 40 m. The width of IEF channel (AB) is 120 m and the width of PAGE channel (CD) is 100 m.

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49 Figure 2 9 IEF separation of two proteins, R phycoerythrin (RPE) and green fluorescent protein (GFP) in the presence of PDMS based valve arrays. The anode was on the left and the cathode was on the right. The scaling bar is 360 m Figure 2 10 ATR FTIR spectra of COC in the native form and those after various surface treatments. (a) Native COC; (b) COC surface activated by corona discharges; (c) COC surface activated and then treated with TMSPMA; and (d) COC surface activated, TMSPMA tre ated, and exposed to the second corona discharges, followed by annealing.

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50 Figure 2 11 Mechanism of the reaction between PDMS and COC. (a) Chemical structure of 3 ( trimethoxysilyl ) propyl methacrylate (TMSPMA). (b) Chemical structure of (3 aminopropyl)t riethoxysilane (APTES). (c) Chemical structure of Zeonor polymer according to the manufacturer. R 1 to R 4 are functional groups while n & m are the number of monomer units in the polymer. (d) Generation of TMSPMA radicals at C=C bond after corona discharg es while some Si OCH 3 were transformed to Si TMSPMA to COC through the formation of covalent bond between the radical and C H groups on the surface via grafting polymerization. (f) Generation of OH groups on PDMS surfaces b y activation. (g) Bonding of COC and PDMS through the formation of covalent bond via the dehydration reaction between OH groups on both COC and PDMS surfaces.

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51 Figure 2 12 (a) XPS spectra of native COC and surface modified COC. (b) Exploded view o f the narrow band of C1s region in the XPS spectrum of the surface modified COC.

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52 CHAPTER 3 DESIGN AND OPTIMIZAT ION OF THERMALLY ACT UATED MICROFLUIDIC V ALVE Overview Chapter 2 describes the fabrication of a pneumatic microvalve in a thermoplastic micro fluidic device. I t consisted of three layers: a COC flow layer on the top, an elastomer PDMS film in the middle and a COC control layer on the bottom. T he valve was opened by introducing a pressure into the channel in the control layer and causing deflect ion of the PDMS film into channel on flow layer. T he core of this valve was the middle elastomer layer. PDMS was chosen as the elastomer layer due to its low Young s modulus (around 75000Kp) and the modified surface properties for bonding. However, a pneu matically d riven microvalve requires a bulk of external equipments to supply pressures. A thermally actuated microvalve controlled by a printed circuit board (PCB) has been developed by our group to address this problem 79 T he structure of this newly microvalve is illustrated in Figure 3 1. A thermal sensitive fluid in the valve g enerates forces to deform an elastomer film into microchannels when the fluid expands due to increasing temperature. H eat is supplied to the fluid from a microheater at the bottom of valve. T he actuation of the microvalve is controlled with a PCB that is c onnected and operated with a computer. In the earlier work, a double sided tape was used as an elastomer layer and glue bonded to a device. T hat microvalve functioned well but it could be still improved in two ways: (1) B onding. T he glue used for bonding i n the previous effort has a possibility to block microchannels. (2) Short a ctuation time. D ouble sided tape film was not a good elastomer material. I ts high Young s modulus makes it hard to be deflected and a longer actuation time is needed. T o address the se two issues a new material needs to be formed to replace the double

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53 sided tape. PDMS was a potential choice as demonstrated in C hapter 2. However, PDMS has disadvantages such as porous structure, chemical incompatibility 104 Thus, an alternative material is required and it needs to be flexible enough to be deformed and bonded well with the flow layer and control layer. I n this chapter, we report our effort in usi ng polyurethane ( PU) as the elastomer layer in the microvalves. PU was classified as an elastomer material and its Young s modulus was between rubber and plastic depends on specific product 105 T he properties of PU vary with its compositions. COC was used to fabricate the other parts of device (control layer and flow layer). T o realize the goal of elastomer based microvalve in a pure polymer microfluidic device, the key challenge is to achieve a st rong bonding between COC and PU enabling this bonding sustains a pressure more than 50 psi. A nother concern when using PU as the elastomer material is its Young s modulus. A material with highYoung s modulus means a higher pressure required to close microchannels under the same channel geometry 106 T o make a co mparison between PU and PDMS both PU and PDMS based pneumatical microvalves were fabricated and the pressures to close valves were measured. Further a PU based thermal actuated microvalve was integrated and its properties were measured. Bonding between PU and COC Experiment Reagents and m aterials Cyclic olefin copolymer (COC) films (Zeonor 1420R 188 m thick ) and Zeonor 1600R COC resins were purchased from Zeon Chemicals (Louisville, KY). 100 m thick COC films (Topas 8007) were obtain from PLITEK (Des Plaines, IL). A solu tion of 98% 3 ( trimethoxysilyl ) propyl methacrylate (TMSPMA) was purchased from Acros

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54 organics (Fair Lawn, NJ) while solvents and common chemicals were from Fisher Scientific (Atlanta,GA). Temperature sensitive fluorinert liquid FC 40 was from 3M (MN, US). Gold etchant were from Transene Company ( MA, US) PU (Bayhydrol 110) was obtain from Bayer Material Science (PA US ). A 6 %( volume ratio) TMSPMA solution was prepared by diluting TMSPMS solution into ethanol solvent Purified water was obtained using Bar nstead Nanopure Water System (Model: D11911, Dubuque, Iowa) A UVO cleaner ( Model number : 342 ) was purchased from Jolight Company ( CA, US) PU preparation A PU film was prepared according to the instructions of the manufacturer. Th is PU is a commercial pr oduct, in the form of a liquid mixture contain in g PU resins, water and n methyl 2 pyrrolidone To obtain a PU thin film, Bayhydrol 110 was slowly poured onto a clean 100 m thick Topas 8007 film (or 188 m thick Zeonor 1420R when high annealing temperature was used) and then spin coated with a spinner (Laurell Technologies) at a constant speed for 30 seconds. The 8007 Topas thin film was used as a sacrificial layer and would be discarded later. PU film on the sacrificial layer was dried at room temperature overnight. T he resulting thickness of PU film varied with the spin coating speed. The relationship between the PU film thickness and spin coating speed is shown in Figure 3 2; the thickness was measured by a digital micrometer (Mitutoyo Company). A s the elastomer in a valve, a thinner film is preferred since it leads to a lower pressure to close the valve I n this work, all the PU film is spin coated at 4500 rpm for 30 seconds and its thickness is 8.9 m0.7 m.

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55 B onding method between COC and PU The proc ess of bonding between thermoplastics and PU is illustrated in Figure 3 3. A plastic film ( 188 m thick Zeonor 1420R film) or 4mm thick sheet fabricated by Zeonor 1 60 0R was washed with DI water using a sonication machine ( model:3510, Branson U ltrasonic C orporation, CT, US ) firstly. After dried the plastic sheet was activated with UVO cleaner for 60 seconds ( F igure 3 3( a ) ), and then immersed into 6% TMSPMA solution for 20 minutes. This chemical treated plastic sheet was blow off dried with nitrogen and a ctivated again using UVO cleaner for 60 seconds. A PU/ sacrificial layer was required to be treated with UVO cleaner for 60 seconds at the same time. T hese two activated layers were thermally laminated (GBC Catena) and then placed in an oven at 80 C overni ght, which f acilitate d the formation of chemical bond s between two surfaces and increase d the bonding strength A fter annealing, the sacrificial layer could be removed and a permanent bonding was formed between the COC layer and PU layer. Discussion Bondi ng c o ndition o ptimization T he functions of a PU layer in both pneumatically driven microvalve and thermally actuated microvalve are similar. Valves were closed by the deformation of the PU film under a n appropriate pressure, thus the bonding between PU a nd valve/control layer must be stable and strong enough to afford this pressure. T he bonding strength, represented by the peeling force when delaminating COC/PU bonding samples, was used for comparison of different bonding protocol s as well as for obtainin g the optimum conditions for each step. The bonding protocol described in the e xperimental section for COC/PU was determined based on a large number of experiments. As an example,

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56 annealing temperature, which was recognized as an important parameter for th e bonding strength was studied for the optimum condition. D uring annealing, chemical bond was formed between activated group on COC and PU surfaces. Figure 3 4 shows the influence of annealing temperature on the bonding strength. B onding strength was a q uantitative indicator when comparing the bonding effects between two samples. A s shown in Figure 3 4 the bonding strength was negligible when the annealing temperature was lower than 40 whereas it increased quickly when temperature was higher than 40 W hen annealing temperature is higher than 50 the bonding strength is more than 55N and the accurat e data was unavailable to obtain due to the limit with COC film strength itself. T he s trip that connects the sample with the moving part of the tensile machine couldn't afford a 60N peeling force and was tended to break around180 bended area. A blister t est was carried out for a direct observation and better comprehension of the bonding strength. A 2mm space diameter through hole was drilled in a 50 mm 50 mm sized 4 mm thick COC sheet ( Zeonor 1600 R), and then bonded with a PU layer as discussed above. A N anoport which facilitated the introduction of a pressure was installed on the through hole in the COC sheet surface. This assembly was immersed into water and air bubble would be observed when delamination happens under a high pressure. For a sample p repared with the optimum protocol we did not observe any blistering when a pressure was up to 100 psi (~689 KPa) the maximum pressure which was able to be supplied by the pressure system discussed in the Experimental section. T his device may sustain a pr essure higher than 100 psi. Considering the fact that most microfluidic devices are operated under 2 0 psi and most microvalve function was

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57 realized less than 50 psi we conclude that the chemical assisted COC/ PU bonding is well suited for microfluidic ap plications. P U versus PDMS I n C hapter 2, we discussed the method to bond thermoplastic including COC with PDMS. For both COC/PDMS bonding and COC/PU bonding, the chemical and protocol used was similar. There are four treatment steps in the optimum proces s before COC film was in contact with PU ( or PDMS) film : (1) COC film activation; (2) TMSPMA treatment; (3) activate the treated COC film (4) activate PU(or PDMS) film. Although the protocol seems the same when bonding COC with PU and PDMS, further study s hows that differences exist between two materials. For COC/PDMS bonding, the function of step 1 enhanced the final bonding strength by 20%. I n other works, COC film could still be bonded with PDMS without step 1. W hile step 1 is necessary in COC/PU bonding COC was not bonded with PU at all without this step. PDMS was also used as an elasmtomer layer in a thermally actuated microvalve. T he result shows that PDMS was not a good choice on this microvalve. V alve functioned well in the first few days, but fai led after a period of storage. After one week, an air bubble formed in heater cavity. T he size of the air bubble increased over time and it will occupy the whole cavity after one month. T his phenomenon can be explained by the porous structure of PDMS. T he molecule s of the thermal sensitive fluid could permeate the PDMS film at room temperature and cause the fluid loss out of the cavity. A coating on PDMS may solve this problem, but it makes the device fabrication complicated.

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58 Pneumatic V alve s Based on P U film Experiment A pneumatic microvalve includes a control layer, a fluid layer and a flexible PU layer. T he PU layer was prepared as described above. T he fluid layer and control layer were fabricated using compression molding as reported previously. Brie fly, a photomask was designed with AUTOCAD. The pattern on photomask was then reproduced on a glass master by photolithography process Electroplating on the glass plate generated a nickel mold, which was employed to produce plastic parts from COC resins ( Zeonor 1 600 ) or COC film( Zeonor 1 420) using a hydraulic press (Carver, Wabash, IN). To assemble a pneumatically drive microvalve the control layer was bonded with PU/sacrificial layer by following the method discussed above. The fluid layer was then activated using UVO cleaner, followed by TMSPMA treatment for 20 minutes. T he fluid layer was placed in the contact with PU /control layer by a laminator after both surfaces were treated with UVO cleaner. Afterwards, the three layer structure device was an nealed in oven overnight. As illustrated in Figure 3 5 ( a) the microvalve was operated by applying a pressure into channels in the control layer. The flexible PU film deflected under pressure and block ed the channel in the flow layer. A pressure supply sy stem was connected with the inlet of the channel in the control layer via Upchurch Nanoport (Oak Harbour, WA). The pressure system was similar with what reported in C hapter 2. I t consisted of a nitrogen cylinder, a pressure controller (Cole Parmer 68502 1 0) and a Dynamco dash valve (Model D1B1202, Commerce, GA) The Dynamco dash valve was connected with an external DC power supply (Model E3630A, Agilent Technologies, San Jose, CA) which

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59 could be controlled by a computer. T he pressure from the nitrogen cyl inder was set as 100 psi. The pressure controller was adjusted between 0 100psi. In order to observe the operation of the microvalve, a food dye solution was introduced into the channels in the flow layer. T he closing of microvalves was confirmed by the disappearance of the dye in the valve area when a pressure was applied into the control channels. An inverted microscope (IX51, Olympus) equipped with a Hamamatsu CCD camera (model C4742 80 12AG) was used for capturing the images Discussion T wo 300 m wid e channels were created in the control layer as valve arrays. These two channels were pneumatically controlled using the pressure supply setup described in the Experimental Section. A blue food dye solution was introduced into the channel s in the flow laye r to confirm the closing and opening of the microval v es. T he whole device contains 39 parallel channels An exploded view of five channels is shown in Figure 3 5 (b). A t the beginning valves were open. Then a pressure of 35 psi pressure was supplied to th e control channel on the right side, the closing of microvalve was observed in Figure 3 5( c ) A fter pressure was also applied on the control channel on the left, two microvalve arrays were closed as shown in Figure3 5( d ) W hen pressure is released, the dye solution refilled the valve area of the microchannels ( Figure3 5( e ) ) T his operation cycle was repeated a number of times and the reliability of the PU microvavle were determined To put the valve operation a PDMS elastomer based valve. I t consisted of three layers and the control layer and flow layer were same as that used in the PU valve. T he control channel is 300 m wide while the

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60 channels in the flow layer are 110 m wide and 45 m deep. T he PDMS layer was prepared by spin coating at 4000rpm for 30 seconds and then cured at 60 for 4 hours. T he thickness of PDMS layer was 15 m, measured by a profilomer (Dektak IIa, Veeco Instruments) T he pressure to close this PDMS valve is 30 psi. Comparing to the PD MS valve, the pressure to close PU valve is 35 psi when the dimension of control channel and flow channel are the same. Thermally Actuated Microvalve based on PU film Experiment D evice f abrication T he structure of a thermally actuated valve is similar to what was reported previously 79 As shown in Figure 3 1 it includes four lay ers: channel layer, elastomer layer, valve layer and heater layer. Zeonor 1 600 resin was used to obtain a channel layer which contains microfluidic channels. The procedure to fabricate a plastic channel layer is the same as for fabricating a flow layer th at was described above. A 9 m thick PU film was used as an elastomer layer, which was prepared following the process discussed above. Both the valve layer and heater layer were made with 188 m thick COC 1420R film. Holes of 2 mm diameter were drilled in the valve layer where a thermal actuated valve is required. A1000 thick serpentine Au resistor was patterned on the heater layer by photolithography and etching. A 20 m diameter through hole was drilled on heater layer for filling with the cavity with the thermally sensitive fluid. To assemble a thermally actuated microvalve, a flow layer with microchannels was permanent ly bonded with PU/ sacrificial layer that were assembed by the method described above. A fter peeling off the sacrificial layer, the flo w layer/PU assembly was then bonded with a valve layer using the same approach. The cavities in the valve layer

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61 need to be aligned with the microchannels in the flow layer where a valve is desired. T he bonding between the valve layer and heater layer was a chieved by solvent bonding using microstamping approach which was reported previously 107 A temperature sensitive liquid, FC40 was introduced into th e cavities in the valve layer through filling holes in the heater layer. T he assembly was immersed in a Fluorinert solution in a beaker and then placed into a vacuum desiccator where the solution was driven into cavities under 20 psi vacuum pressure. A s reported before, vacuum filling often leaves a tiny air bubble in the cavity and the average size of the air bubble is about 6% of the cavity volume. T o remove this air bubble, the beaker with the assembly was placed into a water bath and kept for 10 minu tes. T he temperature of this water bath was set as 70 T he volume of the thermal sensitive liquid FC40 in the cavity increased when heated due to expel ling air bubbles out of cavity. Afterwards, the beaker was taken out of water bath and cooled to room t emperature while FC40 will fill in all spaces of the cavity. A thermally actuated microvalve was achieved after sealing the filling hole with epoxy glue, the microvalve fabrication and assembly was completed. Operation T he operation of the thermally actua ted valve was same as what was reported before. T he contact pads of microheaters were connected to a 5Vpower supply (Agilent E3644A CA, US ). A solution of sodium chloride or a food dye was driven into microchannels by hydrostatic head. T emperature was mea sured using with a termperature sensor (a 76 m diameter type K thermocouple from Omega), which was placed on the bottom of the heater. A constant voltage (8V) was supplied to the microchannel using two platinum electrodes which were inserted into the in let and outlet.

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62 Electric current and temperature were measured through Labview and a DAQ card NI PCI 6229( National Instruments, TX, US), which were controlled with a computer. Discussion A s reported before, there is an air bubble in the heater cavity when introducing the thermal sensitive fluid in a vacuum chamber. T he size of the air bubble was around 6% of the cavity volume. I t decreased the heating efficient during the valve actuation. T his air bubble was removed by a water bath heating process as des cribed in experiment part. Picture in Figure 3 6 show that there were no air bubbles trapped in the valve. V alve actuation was similar to what was reported previously. A conductive solution ( NaCl) was filled into a channel with a flow rate of 333 nL/min. V alve operation was confirmed by measuring the electric ionic conduction current of the solution in the channel. W hen power was supplied to the microheater, the volume of the thermal sensitive fluids increased. T he expansions of the fluid in the heater cav ity deflected the PU film and block the flow channel. T he minima in the conduction current through the channel indicate the closure of microvalve. C omparing with the PET valve actuation, PU elastomer layer microvalves has advantages of shorter response tim e and re open time. T he actuation time was defined as the time to close valve which was calculated from the time when heater power was switched on to the when the electric ionic channel current dereased to a value less than 0.1% of the original value. The actuation time was associated with the power supplied Y modulus of the elastomer film, depth of channels and property of thermal sensitive fluid, etc. In the comparison between PU and PET thermal valves, Young s modulus of the elastomer film plays the most important role. Figure 3 7 shows the valve actuation time as a function of the input heater power

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63 for a PU valve. The minimum power required to close the PU valve is 40 m w and the valve was closed in 20 seconds, whereas a PET valve required 42 m w power and the valve was closed around 500 seconds. W hen power was more than 50 m w a PU valve was closed in 2.5 second to 8.5 seconds, whereas a PET valve required 70 seconds to 80 seconds. When the heater power was turned off the valve was reopened and fluid flowed back into the channel again. The valve reopen time was calculated from time when heater power was switched off to the time when the electric ionic channel current increased to a value more than 99% of the original value. F or a PET film based v alve a couple of minutes are required to reopen microchannels. F or a PU film based valve, the reopen time was less than 0.5 seconds. T here were two reasons that cause this difference. O ne was the lower young s modulus and smaller thickness of PU film than the PET film. A nother was that the PET film used in the previous report contained glue on both sides, which may make the PET film stick to the microchannel and require an extra effort to separate them. The result show that the valve actuation time was s hortened to 4% to 10% of the previous PET based microvalve and the valves reopen time was shortened to less than 1%. The valve reopen time was mainly associated with the flexibility of the elastomer film while the valve actuation time needs to consider bot h flexibility and the heating efficiency. A s to PU valve, we observed that the valve reopen time was smaller than valve actuation time. PU valve could be reopened in less than 0.5 seconds while PU valve actuation time is 2.5 to 8.5 seconds when the power s upply is more than 50m w (Fig ure 3 8 ). The extra time for valve closing was used for the heating and expansion of

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64 the thermal sensitive fluid. If the valve actuation time need to be further decreased, heating efficiency should be improved. Study of PU prop erty When subjected to continuous stresses, the strain of a viscoelastic material is not a fixed value but will change overtime This phenomenon is called creep. Normally, creep rate increased with increasing temperature. In our case, considering that th e valve working temperature was as high as 50 C and each cycle of valve operation lasted 20 to 30 minutes under the incubation requirement, it is necessary to study the creep property of elastomer film PU during this process. This study also provided an i nstruction of precisely valve power control in future application. Experiment Design When valve was closed, a strain occurred on PU surface due to the tensile strength from the increasing volume of thermal expanded fluid FC 40, which located inside the ca vity below PU film. Figure 3 9 shows the theoretical calculation of strain changes on PU film during the valve actuation. When valve was open, the strain was 0 since no film deformation happened. If the valve began to close, a pressure was applied on the P U film at the direction shown in red arrow (Fig ure 3 9 (a)). The PU film was deformed and strain began to increase. When the valve reached the minimum closed (Fig ure 3 9 (b)), the strain is 14.49%. When the valve was totally closed (Fig ure 3 9 (c)), the stra in is 29.12%. A preferred valve control should keep the strain change between 14.49% and 29.12%. If the strain was lower than 14.49%, valve was not closed. If the strain was larger than 29.12%, the FC 40 expanded into the channel outside of valve area and brought destruction to the microchip.

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65 Thus, the valve needs to perform a precise actuation to keep the change of strain value between 14.49% and 29.12%. However, the uncertainties of creep phenomenon, which lead to a strain shift under a certain stress, b ring troubles to reach this purpose. Two sets of experiments were designed to investigate it. One was employed to find the strain changes under a constant stress as a function of time. The other one was employed to find the stress required to keep a const ant strain as a function of time. Test method A dynamic shear rheometer was employed to perform the creep test as shown in Figure 3 1 0 (a). The PU sample was cut into the size of 1 cm 4 cm as shown in figure 3 1 0 (b) and its thickness was measured with ca liper before each test. This PU stripe was placed into the center of chamber. One end of the PU stripe was fixed on the bottom of machine and the other end was hold by a sophisticated clamp for the shear stress loading in experiment. This clamp feature d a non friction force loading which ensure the measurement accuracy. To simulate valve working environment, all the tests proceeded under a constant temperature of 50C. Conclusion Modulus The storage modulus was measured with a dynamic modulus test performed by a dynamic shear rheometer shown in Figure 3 10. I n a dynamic modulus test, the stress was applied on the specimen following equation: Stress= Asin ( t) W here is the frequence, t is time, A is constant. I n this experiment, is set at 1Hz. D ifferent from the elastic modulus, the storage modulus is obtained in a dynamic modulus test and represent the change of elastic portion during this process. Therefore,

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66 the storage modulus represents the trend of a elastic modulus and is helpful to evaulate the mat erial properties. I n a valve application, the modulus of elastomer is a key property that effects the valve performance. S ince the enviroment temperature of a working thermal pneumatic valve is from 25 C to 70 C, it is necessary to study the elastomer mo dulus chages during temperature domain. As the result shown in Figure 3 11, the modulus of PU deceased with temperature increasing. At 25 C the storage modulus of PU is ~ 1.8 10 7 Pa. A t 50 C, the valve closing temperature, the storage modulus is ~6.2 10 6 Pa. Comparing with PDMS, its modulus decreased with temperature increasing 108 It indicates that PU is a better material in a thermally actuated valves and PU based valve is easier to be deformed under higher temperature. Creep To investigate the creep property of PU, we performed a stress controlled test for 30 minutes at a temperature of 25 C and 50C. A pressure of 0.21Mpa was applied to the sample as a she ar stress. Th e value of this pressure was used in a pneumatic PU valve as described in C hapter 3 performing an instantaneous valve closing action when observed under microscope The result was shown in Figure 3 1 2 When the temperature is 50C a t the fir st five minutes, the strain represented an exponential increment with time increasing and increased 16.3% in this stage. After five minutes, the increasing rate of strain decreased and closed to a constant. Strain represented a linear increment with increa sing time and had 4.7% increment in total during the following twenty five minutes. Creep recovery To investigate changes of stress to keep a certain strain, a strain controlled test was performance at a temperature of 25 C and 50C. This test was divided into two

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67 parts. In this first 20 minutes, the shear strain was kept at 10%, which is equal to a 20% 30% tensile strain in our valve application if the same value of stress was applied. The minimum stress required to keep the 10% strain was measured and re corded. In the second 20 minutes, the applied stress was released and the change of strain was measured at the creep recovery process The result of this test was shown in Figure 3 1 3 When the temperature is 50C i n the first 20 minute part, the stress e xponentially decreased from ~46000Pa to ~27000Pa at the beginning 5 minutes, and then shows a linear decrement at the next 15 minutes. The second 20 minutes part showed the creep recovery process without the help of extra stress. The strain recovered 48% i n the first 2 minutes, and then the recovery rate turned to be slow. After 20 minutes, the strain recovered 82% in total. Summary We have developed a reliable method to bond COC with PU. A n optimized protocol resulted in a strong bonding that can sustain at least a pressure of 100 psi. This strong bonding makes it possible to use a PU film in a microfluidic device. To demonstrate this possibility, a pneumatically driven microvalve and a thermally actuated microvalve were fabricated and tested. W e also expe rientially determined the pressures required to close PDMS based and PU based microvalve. W e also demonstrate d that PDMS was not suitable in a thermally actuated valve due to its porous structure. PU based thermal ly actuated valve w as fabricated and compar ed with a PET based thermally actuated valve that was reported before. R esults show that the valve actuated time in PU based valve was shorted to 4% 10 % of that in PET based valve and the valve reopen time was shorted to less than 1 %. Since valve need to k eep closing during the incubation time, the long term creep properties of elastomer was investigated to

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68 ensure the valve function as well as achieving a precise and accurate valve control

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69 Figure 3 1. 3 dimension s tructure of a thermally actuated micro valve. From top to bottom are the channel layer, elastomer film, valve layer and heater layer.

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70 Figure 3 2. The relationship between thickness of PU membrane and spin coating speed. T he PU was spin coated for 30 second for ea ch testing sample.

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71 Figure 3 3. Process to bond PU and COC film.

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72 Figure 3 4 Peeling force between a PU film and COC film as a function of the annealing temperature.

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73 Figure 3 5 The operatio n of PU based pneumatic valve arrays (a)structure of a pneumatically driven microvalve. I t consists of a 188 m thick flow layer integrated with a 4 mm thick control layer and an elastomer layer (b) Picture of two control channels in the atmosphere p ressure and five fluid channels filled with a dye solution. (c) Valves on the right side were closed when a pressure was supplied through the right control channel. (d) Both valve arrays were closed when a pressure was applied to both control channels. (e) Both valve arrays were opened again and dye flowed back to the channels when the pressure in control channels was relieved Figure 3 6 The dye test of a thermally actuated microvalve. Valve is open ( left). Valve is closed (right). (a )

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74 Figure 3 7 The changes of current and temperature as a function of time. T he power supplied for three cycles were: 41 mW, 44 mW and 47 mW.

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75 Figure 3 8 The valve actuation time as a function of the input power.

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76 Figure 3 9 The strain value of PU film changed with the closing of valve. ( a) valve is open;(b) minimum close of valve.(c) totally close of valve. The red arrow shows the direction of applied shear stress aimed to deform the sample.

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77 Figure 3 1 0 : The e xperiment set up of the creep test. (a) A dynamic shear rheometer with opened chamber; (b) A model of testing sample. Figure 3 11. Storage modulus of PU changes with temperature.

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78 Figure 3 1 2 St rain time diagram of PU film under the constant stress of 0.21MPa. Testing temperature is 50 C

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79 Figure 3 13: Strain and stress changes as a function of time. (a) 25 C. (b) 50 C (a) (b)

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80 CHAPTER 4 MULTIPLEXED IMMUNOASSAY IN A VALV E ARRAY SYSTEM Overview Immunoassay Immunoassay is a technology that aims to detect the presence and concentration of a specific antigen or antibody from a sample 109 This detection method is widely used in the field of medical diagnostics, biological research and pharmaceutical research in order to quantify the specific protein or small molecules in various applications. T he realizat ion of an immunoassay is based on the specific bonding of an antibody with its homologous antigen. Based on the assay format t he test of immunoassay can be classified into two types: homogenous and heterogeneous 110 In h omogenous immunoassay s 111 113 antige n and antibody react inside a solution, and may then separate by physical or chemical methods such as electrophoresis 114 118 and fluorescence polarization 119 122 for further detection. In heterogeneous immunoassay s reaction of antigen and antibody happens on a modified surface where the antibody is often immobilized, and the unbounded antigen can be eas il y removed by a rinse with a washing buffer. Figure 4 1 shows the principle of immunoassay. A specific antibody (orange) was immobilized on a substrate (gray). I f the sample contains the homologous antigen, the antibo dy would interact with it and the antigen is then fixed on the substrate. A labeled secondary antibody is used to detect the existence of the antigen. The common approach to perform a heterogeneous immunoassay is a microtiter plate with multiple sample we lls 123 126 An optimized procedure of microtiter plate and related technology enables a test with high detection accuracy. However the long incubation tim e and the consumption of a large volume of samples lead to high cost and

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81 low throughput in one testing cycle. In order to circumvent this problem researchers attempted to integrate immunoassay in a microfluidic system 127 This method becomes more and more attractive and appears to be a possible solution to the traditional immunoassay technology with great improvement of miniaturization in the last decade 128 131 T he chip based immunoassay system has advantages compar ed with the traditional plate based method 132 : (1) reduction in the consumption of reagent and sample due to the micro scale channel dimension in a microchip; (2) a decrease in the analysis time since the high surface to volume ratio reduces the antigen/antibody diffusion time and speed s up the reaction process and (3) portability due to the smaller size of a microfluidics chip. In particular the coupling of immunoassay with accessories such as valves and pumps integrated in a microfluidics system would enable the accurate and precise fluid control in an experiment and enhance the det ection performance and reproducibility, leading to a point to care operation. Fig ure 4 2 illustrates a typical chip based immunoassay process 133 This disposa ble microfluidic device could detect one antigen in each test. I f the sample contains multiple target antigens, multiple devices and test s are required. To simplify the multiple antigen detection procedure, we integrate a microfuidic device with valves In one design, up to 6 antigens can be detected in one test. Surface M odification In the heterogeneous immunoassay, an antibody is immobili zed on the surface where the assay solution flows through. The surface immobilization method of the antibody is so important because it dramatically influence s the assay performance and detection sensitivity In a microfluidic system 134 antibody immobilization was realized

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82 either on the substrate of a microchannel 135 or on an embedded microbead 136 139 Although microbeds have some advantage s, it is not a good choice in our valve integrated microfluidic system due to the following r easons : (1) microbeads have a risk of surface adsorption which may prevent the valve from closing if this adsorption occurs on the valve area ; and (2) valve operation requires a narrow channel depth, while a microbeads embeded device prefers a high depth channel which can alleviate the increasing flow resistance Also, the possibility of channel clogging by the embedded beads increases T hus, we focus our work on the immobilization of antibodies on to the substrate surface inside a microchannel. A ntibod ie s can be immobilized inside a channel through physical 140 142 or chemical modification 143 145 In the physical modification proteins are immobilized via a non specific adsorption on the surface. The physical adsorption is simple to be achieved inside a channel, but it may significantly reduce the binding activi ty of the adsorbed protein 135 because of the possible denaturation and different adsorption orientation. O n the other hand, chemical modification in which the substr ate surface is covalently bonded with a compatible reagent normally represents a uniform antibody immobilization as well as homogenous binding between antigen and antibody 146 Methods of chemical modification are diversified depending on the substrate material such as silicon, glass, polymer s gold, etc. Here, we only discuss the surface modification on polymer since this work is focused on the design, fabrication, and characterization of a plastic microchip due to the low cost and the potential of mass production.

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83 Among polymer materials, PDMS is a popular substrate for microfluid ic immunoassay because of its properties such as bio compatible surface, low cost, simple molding and accurate duplication. P olymethymethacrylate (PMMA) is another popular choice in microfluidic device because of its easily modified surface property. Recen tly, COC receives more and more attention and thus its surface modification method has been studied by many people as well 147 149 Dr. MacCraith 149 and his group presented a strategy to immobilize antibody on a C OC (Zeonor 1060R) substrate by coating polyelectrolytes after plasma and chemical treatment on the surface. Dr. Klapperich 148 and his group r eported a chemiluminescence detection of the C reactive protein (CRP), an inflammation and cardiac biomarker in a s erum sample In this chapter, we report a COC surface modification method by covalently bond ing with a nitrogen generated photo activated biotin (Figure 4 3). C hallenge s W e present the fabrication and characterization of a multiple valve integrated microc hip designed for multiplexed immunoassay detection Through the combination of multipl e microvalves and optimized operation process, this microchip can detect up to 6 antigens in one sample. A printed circuited board (PCB ), coupled with a labview program was made for accurate control of each valve through a computer. Dr. Heineman 132 had summarized six fundamental that must be addressed by all micro devices aiming for an immunoassay application The following is our sol ution to these challenge s: M aterial s for microfluidic s PU was used as the elastomer because of its role in valve function The thermoplastic COC was chosen as the channel layer due to its properties such as

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84 optical transparence and chemical resistance. Th e bonding method between PU and COC, as described in Chapter 3, is also an important reason to choose COC. Surface modification The COC surface was covalently bonded with the photo activated biotin under UV exposure The photobiotin enable d the binding w ith streptavidin followed with the immobilization of antibod ies. I ntroduc tion of sampl es /reagent s A syringe pump was used to introduce sample s and buffer s into channels. A controller, which allowed the operation of up to 4 pumps at one time was connected with the syringe pump to ensure the assay reagents were injected at the setting speed. I mmobiliz ation of antibody After the modification of photobiotin on the COC surface s treptavidin was then fixed on the modified surface through the strong bond betwee n streptavidin and biotin. It then enabled the immobilization of a biotinylated antibody. D etection The fluorescence labeled antibody was used for detection. M icrofluidic design Thirteen thermally actuated valves were integrated in the microchip and the y were controlled by PCB through a computer. The operation process will be discussed in detail below. Device Design Assay Design The concept of the microfluidics enabled, multiplexed immunoassay array based on thermally actuated microvalve s will be demon strated in a 23 array as illustrated in

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85 Figure 4 4 in which 6 antigens can be simultaneously detected. The antibodies ( Ab ) will be introduced from the t wo vertical microchannels A t the channel intersections the antibody solution will not flow into horiz ontal channels when appropriate microfluidic valves are closed. I n the left vertical channel, three capture antibodies (Ab 1,Ab 2, and Ab 3) will be introduced. T he channel surface will be immobilized with these three antibodies. T hese antibodies are speci fic to antigen 1, 2 and 3. Similarly, three different antibodies (Ab 4, 5, and 6) specific to antigen 4, 5, and 6 will be introduced in the right vertical channel. After incubation and washing the channels, a sample solution is pumped into all of the channels. T he six analytes of interest will be captured in the corresponding intersections if they exist in the sample. A fter washing these channels, two secondary antibodies specific to antigen 1, and 4 are introduced in horizontal channel 1 when the appr opriate valves are closed T he secondary antibodies (detection antibodies) are modified with a fluorescent dye to facilitate detection. Similarly, two different antibodies (2nd Ab 2, and 5) specific to antigen 2, and 4 are introduced in horizontal chann el 2. Lastly, two other antibodies (2 nd Ab 3, and 6) are introduced in horizontal channel 3. A fter washing, a fluorescence signal at each intersection indicates specifically the presence of the corresponding antigen in the sample and which can be observed under microscope Device L ayouts Flow layer Fig ure 4 5 shows the layout of a channel layer. I t consists of two patterns T he channel in the layer is designed to have a dimension of 100 m width by 35 m depth A fter optimization, the inlet and outlet in each pattern is reduced from 10 in Figure 4 4 to

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86 6 I n the left pattern, the distance between each detection point which is the intersection of vertical and horizontal channel s is 9 mm ( center to center), matching the 96 well microplate standard. This st andard was created by the Society for Biomolecular Screening and accepted by the American National Standards Institute T he pattern on the right side was similar to the one on the left except for curved channels. Heater layer T wo heater layers were desig ned for their corresponding channel layouts (Fig ure 4 6). E ach heater layer consists of 13 heaters ( or valves ) that allow flow control in the channel layer. Each heater is 3.05 mm long and 50 m wide. The resistance of the heater depends on the thickness o f deposited gold film on the heater layer surface. W hen the gold layer is 1000 the resistance is 225 15 his heater layer will be connected to a programmed PCB and operated using a Labview program. Photobiotin Immobilization Experiment al Methods Photobiotin coating on COC surface Photobiotin was diluted with dimethyl sulfoxide ( DMSO ) for 20 mg/m L and then diluted with DI water to obtain 1 m g /m L The c oating substrate was rinsed with ethanol for 5 minutes, and then DI water for 20 minutes before use. For each photobiotin coating, 1 L of 1 mg/m L photobiotin was dispensed on the substrate surface, and then dried overnight in a dark room to avoid light. To form a covalent bonding between the photobiotin and substrate the photobiotin coated surface was placed under a portable UV light at the wavelength of 365 nm and exposed for 2.5 hours. After exp osure, each sample was rinsed with DI water for 30 minutes to remove the unbounded photobiotin. Figure 4 9 shows the photobiotin coating protocol.

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87 Immunoassay The b iotinylated surface was immersed in 10 m L of 1% bovine serum albumin ( BSA) for 30 minutes to prevent non specific adsorption and then washed with PBS containing 0.5% Tween 20. 1 mg /m L streptavidin was coated on the biotinylated surface and incubated 30 minutes at room temperature to prepare the surface for the followed application. In the immun oassay, the biotiny l ated antibody was immobilized on streptavidin activated surface through the conjunction of biotin streptavidin biotin The antigen to be detected then reacted with the fixed antibody on the surface through the binding of antigen/antibod y. A fluorescence labeled secondary antibody was used to detect the presence of the antigen under microscope. PBS solution with 0.5% Tween was used as the washing buffer to rinse the sample between each step Discussion Solvent e vaporated versus in solutio n on photobiotin coating The form of photobiotin after the immobilization on to the target surface is an interesting topic and has been discussed by many people. In Kuhr s paper 150 they emphasized that the photobiotin need s to be totally dried before UV exposure to get a better immobilization performance; while in Seong s paper 151 they exposed the glass surface immediately after photobiotin coating and got a uniform pattern as presented in their paper. The diversity of coating methods ma y come from the different choices of photobiotin and substrate material. To optimize the coating method in our application, we characterized the immobilization performance of photobiotin in solution and in dry The dried sample was prepared as described in the method section. The in solution sample was prepared by dropping 10 m L of 1 m g /m L photobiotin on the COC surface and then covering with a clean glass slide. 1 mg/m L streptavidin and 100 g/m L

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88 fluorescen tly labeled biotin were used to characterize the c oating performance The result was shown in Figure 4 7. After the coating steps of streptavidin and fluorescence biotin, the signal from the dried sample is more than 7 times higher than that prepared in the solution It indicated that a better immobiliza tion was obtaned with the coating of dried photobiotin. Photobiotin coating on different substrates The structure of photobiotin was shown in F igure 4 3. The aryl azide group contained in this structure is photoactive and generates nitrene when it is expo sed under a UV light ranging from 350 370nm. The generated nitrene react s with N H, O H C H or C=C bond contained in the substrate res ulted in a connection between the biotin group and the substrate This property enables it to be a universal tool to mod ify the hydrocarbon polymer such as COC, polystyrene, etc, as indicated in F igure 4 8 indicated. This result also indicated that the TMSPMA treatment did not decrease the efficiency of photobiotin immobilization because of the presence of C H, C=C bond in TMSPMA. Immunoassay The immunoassay process was shown in F igure 4 10 T he surface was modified with photobiotin and then activated with streptavidin and biotinylated antibodies C RP or transferrin, ranging from 2 g/m L to100 g/m L was tested and the ca libration curves were plotted in F igure 4 1 1 FITC CRP and FITC transferrin was employed for signal detection under fluorescence microscope.

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89 Immunoassay in a 13 valve controlled microchip Experiment al M ethods Device assembly The structure of microchip was shown in Figure 4 1 2 Thirteen holes were drill on COC valve layer, and they are matched with the location of each valve. PU was spin coated on a COC sacrificial layer at the speed of 4500 rpm and dried in an oven at 70 C overnight. Six through holes were drilled on the PU sacrificial layer, corresponded to the 6 observation spots in channel layer. The PU layer was bonded with valve layer using TMSPMA as described in chapter 3. T he bonded valve PU sacrificial layer was rinsed with DI water for 30 minutes to remove the unreacted TMSPMA. Following the process described in section 4.2.1 the photobiotin was coated on the valve layer, limited to the through holes area. A 3 mm thick COC channel layer, made with a mold through hot embossing method, was bonded with the other side of PU surface after peeling off the sacrificial layer. The fabrication and assembly procedure of the heater layer is the same as the thermally actuated valve as described in C hapter 3. A assembled device is shown in Figure 4 13. Control sys tem The valve control system consisted of a power supply, printed circuit board ( PCB), and computer. Figure 4 1 4 shows the connection among the connections of this system. PCB is the core component in this system. B oth the laptop and microchip were connec ted to the PCB. Instructions from the computer were delivered to the microchip through PCB to control the open ing and clos ing of each valve T he power required to drive valves was supplied through PCB as well. A power supply was used to enable the

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90 valve fu nction and another supply was used to power PCB. Figure 4 1 5 shows the picture of the PCB. Immunoassay process The assay reagents were introduced into channels with a syringe pump through the nanopore connector, which was installed on each inlet to keep the whole fluidic system sealed and prevent it from the leakage under the applied pressure. A controlle r which can connect up to four syringe pumps, was used to driven and control the flow of assay reagents at a certain speed. 1% BSA and 1mg/m L streptavi din were flooded through channels before test to block the non specific adsorption and activate surface. B iotinylated antibodies, antigens and detection antibodies were introduced inside channels at the speed of 200 nL/min with the necessary valve operatio n described in the experiment design section. T he incubation time of each step is 30 minutes. A PBS containing 0.5% Tweet 20 was used as the washing buffer to rinse the channel after each step. Valve operation The open and close of each valve on each step is shown in Figure 4 16.Before experiment, a mixture of 1% BSA and 1mg/m L streptavidin were flooded through the channel surface for 30 minutes to enable the immobilization of antibodies in next step A fter rins ing with PBS, valves were clo sed as shown in Figure 4 16 (a). T he mixture of 7 0 g/m L biotinylated anti CRP 70 g/m L biotinylated anti GPADH and 1% BSA was introduced into the left vertical microchannel. The mixture of 70 g/m L transferrin and 1% BSA was introduced into the right vertical microchanne l. A micropump was used to load samples and the flow rate was set at 200nL/min. After 20 minutes, air was filled inside the channel to empty channels and then valves were reopened. A rinse buffer

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91 wa s introduced inside channel to remove the unreacted antib odies. T hen valves were closed as shown in Figure 4 16 (b) and a mixture of 40 g/m L CRP, 50 g/m L transferrin, 30 g/m L GAPDH and 1% BSA was introduced inside channel at a flow rate of 200 nL/ min After 20 minutes, valves were reopened. C hannels were wash ed with rinsing buffer. FITC antibodies were then pumped into channel for detection. To loading detection FITC antibodies, v alves were closed as shown in Figure 4 16 (c). A mixture of 30 g/m L FITC anti CRP and 1%BSA was filled into the top horizontal chan nel. A mixture of 50 g/m L FITC anti transferrin and 1%BSA was introduced into the middle channel and a mixture of 30 g/m L FITC anti GAPDH and 1%BSA was driven into the bottom channel. Discussion Traumatic brain injury (TBI) biomarkers were used to demon strate the device function Table 4 1 is a list of TBI biomarkers 152 CRP, transferrin and GAPDH were used for detection in our experiment. CRP is a biomarker in response to inflammation Transferrin is related to the brain iron rate. GAPDH is a biomarker associated with metabolic function 152 The abundance of GAPDH decreased after TBI and abundances of CRP and transferrin increased after TBI. Three biomarkers were simultaneous detected in the immunoassay microchip. 70 g/m L biotinylated anti CRP, 7 0 g/m L biotinlylated anti transferrin and 70 g/m L biotinylated an ti GPADH was pattern ed into the first dimension of channels as shown in Figure 4 1 6 with corresponded valves closed. A mixture contain ing 4 0 g/m L CRP 5 0 g/m L transferrin and 30 g/m L GAPDH was flowed into the appropriate channel and reacted with the sur face bonded antibod ies 4 0 g/m L FITC antiCRP, 50 g/m L FITC antiTransferrin and 30 g/m L FITC anti GAPDH were pumped into the channel to detect

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92 the presence of antigens. A representative fluorescence result of this three biomarker detection was shown in Figure 4 1 7 1% BSA was used as a negative control in this experiment. This experiment was designed to identify the existence of biomarkers. Summary In this chapter, we designed created and characterized a prototype of an immunoassay microchip that enabl ed a simultaneous detection of up to 6 analytes in one sample. T his disposable microchip was made with a biocompatible, optically transparent thermoplastic to allow a low cost, mass production manufacture. A surface modification method was developed on COC using photo activated biotin for antibody immobilization through the interaction between streptavidin and biotin. To enable a simultaneous detection of multiple analytes, 13 thermal ly actuated microvalves were integrated in the microchip and they were con trolled by a PCB through computer. CRP and Transferrin was employed to demonstrate this platform and the result shows the potential of the simultaneous detection of up to 6 analytes. A large scale design may be developed for detection of 24 or 48 analytes or more

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93 Table 4 1 List of traumatic brain injury (TBI) biomarkers. Post TBI Protein MW(kDa) Decreased abundance Increased Abundance Cofillin 1 18.5 Profillin 1 14.9 Enolase 1 protein 47.1 Hexokinase 1 102.4 Aldolase A 39.3 G APDH 36 Hsc70 ps1 70.4 MAP2A/B Carbonic anhydrase Transferrin Fetuin Haptoglobin Actin C reactive Protein Brain creatine Kinase 200 29.6 75.8 41.5 38.5 42 25.5 42.6

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94 Figure 4 1 Schematic illustration of the immunoassay principle Figure 4 2. Schematic illustration of an assay procedure

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95 Figure 4 3. The structure of photobiotin used to modify COC surface s Figure 4 4. Experiment al process of antigen detection.

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96 Figure 4 5. The layout of two channel layers designed for immunoassay test Figure 4 6. Th e layout of two heater layer and their corresponding channel layers. (a) (b)

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97 Figure 4 7 Photobiotin coating protocol Figure 4 8 Performance comparison of photobiotin coating using different methods.

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98 Figure 4 9 Comparison of photobiotin performance s on different substrates. After the coating of photobiotin, 1 mg/m L streptavidin and 100 g/m L FITC biotin was flooded over surface to characterize the photobiotin immobilization performance.

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99 Figure 4 10 Immunoassay process on biotinylated COC surface.

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100 Figure 4 1 1 Calibration curve of immunoassay on COC surface. (a) CRP (b) transferrin. (c) GAPDH ( b ) (a)

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101 ( c )

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102 Figure 4 1 2 The three dimension mo del of the valve controlled immunoassay microchip. (a) Four layers of microchip (top to bottom) : channel layer, PU elastomer layer, photobiotin coated valve layer, heater layer. (b) The top view of an assembled device. Figure 4 1 3 A assembled device mad e by COC.

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103 Figure 4 1 4 The set up of control system for the multiple valve operation. Figure 4 1 5 The PCB designed for multiple valve microchip.

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104 Figure 4 16. Valve operation protocol in a 3 TBI biomarker test. Transferrin, CRP and GAPDH were identified in this device. T he red bar indicates the position where valve was closed. (a) Introducing the mixture of biotinylated antibodies. (b) introducing the mixture of protein. (c) intoducting the mixture of FITC antibodies for detection.

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105 Figure 4 1 7 The result of immunoassay detection of 3 biomarkers in mu l tiple valve s controlled microchip.

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106 CHAPTER 5 CONCLUSION AND FUTURE DIRECTIONS Conclusion A microfluidic syste m normally consists of channels, valves, pumps and mix er s. The microvalve is one of the critical components since it is required for containing fluid, flow regulation and region isolat ion This dissertation focused on the design, fabricat ion and characteri zation of an elastomer based valve in a low cost, disposable plastic microfluidic device. PDMS and PU were tested as the elastomer material, and pneumatic valves and thermally actuated valve s were made and characterized For their applications thirteen PC B controlled thermally actuated microvalves were integrated with an immunoassay microfluidic device to demonstrate the platform allowing simultaneous detection of multiple analy tes. Several key conclusions are summarized as follows. Bonding Bonding is t he critical technology in an elastomer based valve fabrication since valve closing in a microchannel requires a relative high pressure, ranging from 30 psi to 80 psi. We developed a bonding method to enable a strong bond between COC and elastomer materials such as PDMS and PU. This hydrolysis resistan t bonding method could sustain a pressure more than 100 psi. PDMS versus PU When Quake 3 et al. d eveloped the elastomer based valve, PDMS was used as the elastomer layer due to the low Young s modulus of PDMS and a strong bond between PDMS and the flow layer. Although PDMS is a good material in the academic setting, it is not readily or economically formed in high throughput production 153 I n particular,

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107 the porous structure of PDMS result in the evaporation and leakage of FC 40 the essential thermal sensitive fl uid used to deform the elastomer and enable the channel closing in our the rmally actuated valve. Compar ed with PDMS, PU is a manufacturable material which was widely produced in industry and is used in our everyday life. Although its elastic modulus is not as low as PDMS the PU based pneumatic valve could be driven under the sa me pressure as PDM S based one by decreasing the thickness of the elastomer layer. Furthermore, PU was successfully employed in a thermally actuated valve due to its chemical resistance as well as the material structure where PDMS failed. Pneumatic versus Thermal An elastomer based valve could be pneumatically driven or thermally actuated. The pneumatic valve has a simple structure, short response time, but requires bulk y accessories for the pressure supply T he thermally actuated valve has a longer respons e time due to the heat diffusion time constant, but it is portable since it is riven by electric power supply such as a battery Applications In this work, t hermally actuated valve s were employed in a microfluidic device to carry out simultaneous detection of multiple immunoassays. T hirteen valves were integrated to handle the flow s required for process. A lithographically patterned heater layer was designed to connect the valves with a PCB enabling the control of individual valves through a computer The detection of 100 g/m L CRP and 100 g/m L t ransferrin in this platform demonstrate d the reliability of valve function and the feasibility of multiple analyte detection. This proof of concept experiment indicated the potential of large scale simultaneous de tection in a disposable, low cost mass produ ction chip.

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108 Future Drections I n C hapter 4, we described a design to carry out simultaneous immunoassay detection of multiple analytes in a platform integrated with microvalves. W e have demonstrated this idea an d characterized a prototype that can test up to six analytes in one samples. However, a microtiter plate can employ 96 or even 384 samples in one plate, depending on the number of well s the plate contains. Therefore, as a replacement of microtiter plate, t he microfluidic platform should be further developed to detect more analytes at one time as shown in Figure 5 1. To achieve this goal dozens or even hundreds of microvalves need to be integrated in one microchip which will bring some new challenge s such a s system stability, power efficiency and detection sensitivity. Valve Fabrication As pointed in section 5.1, both the pneumatic valve and thermally actuated valve we have developed are not perfect yet in some aspects. A s valves us ed in large scale immuno assay detection, they should meet two requirements: (1) S imple. Comparing to a sophisticated structure, a simple structure may result in a higher yield and more stable system. T his is important for integrating dozens of valve s in one platform. (2) Efficien cy E lastomer based valve s require electric power. Considering the number of valves required and the small size of microchip designed a valve with low power consumption valve is necessary To achieve these two requirements, valves could be improved in the following aspects : M aterial We have compared PU and PDMS as the elastomer of valves and conclu ded that PU is better than PDMS when aiming for a low cost, mass produc ed disposable microvalve. Despite the successful application of PU in the valve assembly, we did not

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109 have a chance to do a detailed study of PU proper ties O n the other hand, as one of the most widely used material in our daily life, PU has great diversification in material properties such as elastic modulus, polymerization degree and chemical resistance. Since the elastomer properties dramatically influence the valve performance material investigation could be carried out to obtain PU with optimized properties such as low elastic modulus and high biocompatibility V alve design The thermally actuated valve results in a low fabrication yield, while the simple pneumatic valve requires a complicated control system. O n the other hand, a thermally actuated valve has low energy efficiency since the electric energy needs to be transferred into the th ermal energy and both the energy transfer and heat diffusion lead to an energy loss. Thus, the next generation valves can be improved in two aspects: simple structure and high energy efficiency A pin pressed valve, which consists of two layers and blocks channel by pressing a pin may address the issue Surface modification T he surface modification method plays a cr itical role in immunoassay performance and the detection sensitivity. I n C hapter 4, we described a method that uses photobiotin to modify the COC surface, enabling bio molecule immobilization With this modification method, CRP and transferrin were detected. H owever, this method can be optimized in two aspects for large scale immunoassay detections in the future: (1) patterned surface coating. W e performed the photobiotin coating on a flat surface such as valve layer as shown in Figure 4 15. A uniform coating inside channel layer is preferred in the future since it will simplify the device assembling process and enhance the device performance; ( 2) Sensitivity I n C hapter 4, the detection of CRP, transferrin and

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110 GAPDH were carried out at a concentration of 50 g/m L while the CRP concentration in serum 154 is much lower than this. The long term goal of this platform is to achieve d etection of biomarker s in serum s so that this platform may become a competitive tool in biomedical application s

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111 Figure 5 1. The 3 dimension design of a 4 6 valves array for 24 analytes detection. (a) Assembled device. (b ) Assembled device controlled with a PCB board.

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112 LIST OF REFERENCES 1. S. K. Sia and G. M. Whitesides, Electrophoresis 2003, 24 3563 3576. 2. D. J. Beebe, G. A. Mensing and G. M. Walker, Annual Review of Biomedical Engineering 200 2, 4 261 286. 3. M. A. Unger, H. P. Chou, T. Thorsen, A. Scherer and S. R. Quake, Science 2000, 288 113 116. 4. S. Kaneda and T. Fujii, Adv Biochem Eng Biotechnol 2010, 119 179 194. 5. J. Wang, Talanta 2002, 56 223 231. 6. A. Khademhosseini, R. Langer, J. Borenstein and J. P. Vacanti, Proc. Natl. Acad. Sci. U. S. A. 2006, 103 2480 2487. 7. M. C. Wu, Proceedings of the IEEE 1997, 85 1833 1856. 8. E. W. H. Jager, E. Smela and O. Inganas, Science 2000, 290 1540 1545. 9. D. Falconnet, G. Csu cs, H. M. Grandin and M. Textor, Biomaterials 2006, 27 3044 3063. 10. S. C. Jakeway, A. J. de Mello and E. L. Russell, Fresenius Journal of Analytical Chemistry 2000, 366 525 539. 11. N. Pamme, Lab on a Chip 2006, 6 24 38. 12. N. Lion, T. C. Rohne r, L. Dayon, I. L. Arnaud, E. Damoc, N. Youhnovski, Z. Y. Wu, C. Roussel, J. Josserand, H. Jensen, J. S. Rossier, M. Przybylski and H. H. Girault, Electrophoresis 2003, 24 3533 3562. 13. N. A. Lacher, K. E. Garrison, R. S. Martin and S. M. Lunte, Electr ophoresis 2001, 22 2526 2536. 14. K. E. Petersen, Proceedings of the IEEE 1982, 70 420 457. 15. F. Schneider, T. Fellner, J. Wilde and U. Wallrabe, J Micromech Microeng 2008, 18 16. A. Abbaspour Tamijani, L. Dussopt and G. M. Rebeiz, Ieee Trans actions on Microwave Theory and Techniques 2003, 51 1878 1885. 17. J. B. Yoon, B. K. Kim, C. H. Han, E. S. Yoon and C. K. Kim, Ieee Electron Device Letters 1999, 20 487 489.

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122 BIOGRAPHICAL SK ETCH Pan Gu was born and brought up in Anqing, China. She entered Shanghai Jiao Tong University in 200 0 and earned h er Bachelor of Engineering degree in the Department of Material Science and Engineering in 200 4 She continued a graduate study in the same university and obtain ed her Master of Engineering degree in the Research Institute of Micro/Nano Science and Technology in 200 7 Sh e came to University of Florida in August 2007 in pursuit of a Ph.D. degree in Department of Mechanical and Aerospace Engine ering.