ABSTRACT
MELVIN, ELIZABETH MICHELLE. AC Electric Field Separation, Collection and Detection of Cells and Synthetic Particles. (Under the direction of Dr. Orlin D. Velev).
The goal of this dissertation is to develop new tools for the manipulation and characterization of specific cell types in AC electric fields. Cells are polarizable and resistive, therefore, it is possible to manipulate and detect cells using AC electric fields. Novel bio-microfluidic devices using two AC field techniques, AC electrokinetics and impedance spectroscopy, were developed to collect, trap and detect cells and particles in electrolyte suspensions. First, a new device was constructed to collect cells and particles in a predetermined area on the chip, using AC fields to drive fluid flows with AC electrokinetic phenomena over asymmetric planar electrodes. When an electric field is applied to asymmetric electrodes, the ionic charge mobility above the electrodes creates small eddies of fluid flow, drawing the bulk of the fluid in a manner analogous to a conveyor belt. The gold-on-glass planar asymmetric electrode pattern uses AC electrosmotic (ACEO) pumping to induce equal, quadrilateral flow directed towards a stagnant region in the center of the electrode chip. A number of design parameters affecting particle collection efficiency were investigated including: electrode and gap width; chamber height; applied potential and frequency; number of repeating electrode pairs and electrode geometry. By optimizing these parameters, the device was proven capable of rapidly collecting particles in the stagnant region within six minutes. The robustness of the design was evaluated by varying electrolyte concentrations, particle types and particle sizes. During the evaluation with electrolyte suspensions of particles, the device was prone to electrochemical reactions at the electrode surface and cell suspensions tended to adhere to the glass substrate, reducing the overall particle collection efficiency. Several coatings (silanes, hydrogels and polyelectrolytes) over two electrode compositions (gold and platinum) were investigated to reduce or eliminate this effect. To confirm that the electrochemical reaction had been reduced for each coating type, cyclic voltammetry studies were performed. Platinum exhibited the most promising performance for device robustness. Preliminary analysis suggests polyelectrolytes to be the most effective for reducing electrochemical effects on the electrodes; however the silanes were superior in
the reduction of biofouling. These devices are amenable to integration with a variety of cell detection lab-on-a-chip methods, and specifically, are compatible with optical evanescent waveguide detection technique. The second part of the dissertation focuses on impedance-based cell detection, and the characterization of interdigitated gold electrode devices as well as methods to detect specific cell types by agglutination. Impedance is the resistance and capacitance in an electric circuit while impedance spectroscopy is the measure of impedance over varying applied frequencies at a fixed current and voltage. The effects of agglutination techniques and suspension characteristics on the impedance spectroscopy of yeast/latex particle, yeast/gold nanoparticle and yeast/magnetic particle suspensions were investigated. All of the agglutinated systems demonstrated a distinct, repeatable impedance change versus the non- agglutinated systems, showing impedance spectroscopy across interdigitated electrodes is sufficient to detect distinct agglutinated suspensions. Improving upon existing impediometric detection devices, biospecific ligand-functionalized magnetic particles were used to draw the agglutinated suspensions to the electrodes. To improve the device further, it was inverted so that cells bound exclusively to specifically functionalized magnetic particles. The bound particle and cell conglomerates were manipulable using magnetophoresis and brought to the electrode surface. The bio-chip impedance spectroscopy technique could be extended to determine the feasibility of using impedance spectroscopy methods in a system with fluid flow, to systems with multiple cell types and be of potential use in new novel biomaterials, biosensors, and microdevices.
AC Electric Field Separation, Collection and Detection of Cells and Synthetic Particles
by Elizabeth M. Melvin
A dissertation submitted to the Graduate Faculty of North Carolina State University in partial fulfillment of the requirements for the degree of Doctor of Philosophy
Chemical Engineering
Raleigh, North Carolina
2012
APPROVED BY:
______Dr. Orlin D. Velev Dr. Peter Fedkiw Committee Chair
______Dr. Sonia Grego Dr. Glenn Walker
______Dr. Michael Dickey
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DEDICATION
This dissertation is dedicated to:
My mom and dad For teaching me the value of education
And to: Adam, Abigail and my future child. You are my life and my life is now complete. iii
BIOGRAPHY
Elizabeth M. (Lynch) Melvin was born on June 17, 1978 in Springfield, IL. The daughter of an engineer and an elementary school teacher, Liz moved a bit as a child between Illinois and Ohio, eventually ending up in Lexington, OH. After graduating high school in 1996, she attended Ohio University for a quarter, majoring in biochemistry determined to pursue a career in medical research. After careful consideration of a close friend’s advice, she decided to switch majors to chemical engineering. While in ChE at Ohio State, she was involved in AIChE, science outreach, and undergraduate research with Dr. Robert Brodkey developing fluid dynamics computer models for impingement mixing. She also worked part-time for the OSU Athletic Ticket Office. She co-oped with Dow Corning (a world leader in silicon products) for four quarters in Midland, MI. One of her most favorite jobs while at Ohio State was being a TA for Mass Balance, Fluid Dynamics and Heat Transfer classes. In doing so, she discovered she had a passion for helping others understand chemical engineering principles. In August 2002, she graduated from Ohio State and started working for Johns Manville (JM) in their Fast Start program in Hazleton, PA. Johns Manville is the #2 producer of building insulation in the US, with world-wide operations. The Fast Start program sent her traveling all over the country for training and to help start up several poly- isocyanurate foam insulation plants. She started as a plant engineer and was promoted to a Product Development Engineer role in appliance and aerospace insulation in early 2004 in Denver, CO. There, she discovered a love for hiking and mountains and even found her way to the top of Mt. Princeton (14,167 ft). She enrolled in the Six Sigma program and graduated with a DMAIC Black Belt in 2005. Her career was going well, but she realized that she still had a passion for teaching and mentoring. A PhD would position her to be promotable to teaching and mentoring roles in both industry and academia. Impressed by the closeness of the NCSU Chemical and Biomolecular Engineering Department, she quit her job (after another promotion to Research Engineer), and moved from Denver, CO to Raleigh, NC. During her 5 and a half years at NCSU, she has had the opportunity to TA for mass transfer, Thermo I, Thermo II and unit operations lab. She has also been a tutor in the IMSD iv
program, held a GSA office, held a science day assembly for a local elementary school, judged poster competitions, and assisted with the 82nd ACS Colloids and Surface Science conference. While performing her graduate work, she also bought and refinanced her first house, gotten engaged / married (to the same person), adopted a beautiful mutt named Summer, joined a soccer team, joined a kickball league, gotten a master’s degree and watched her husband write and defend his thesis. Last, but certainly not least, she became a mother for the first time on May 21, 2011. After graduation, Liz will be starting her PhD career as a New Product Development division Process Engineer at Hospira in Rocky Mount, NC. v
ACKNOWLEDGMENTS
I would like to thank the Velev group for their continued support and teaching. It’s nice to know that there is always someone around to ask for help. Burak Ucar, Stephanie Lam, Elena Blanco and Hyung-Jun Koo in one way or another have all been instrumental in keeping the lab running as well as helping me whenever it was needed. They all, most certainly, made being the lab safety manager much less difficult. Thank you to the Graduate School for choosing me to be a part of their flagship Dissertation Writing Grant program. The funding was instrumental in helping me to write my dissertation and to help take care of my daughter for the first 7 months of her life. Being able to be with Abigail while finishing my PhD will be an experience I will hold dear for the rest of my life. I’d like to give an extra special thanks to Brandon Moore for his dedication and hard work with gathering data for the cell pre-concentrator project. Brandon is so enthusiastic and bright and a complete joy to work with in the lab. Brandon is now in the master’s BTEC program and is doing excellent work. I know that his career will go far! Thank you, Brandon, for being a fantastic help! The Fedkiw lab, Dr. Fedkiw and Rameswar Yadav have been incredibly helpful in teaching me about and allowing me to use the IM6 Impedance Measurement System in their lab. Yadav has been exceptionally helpful and patient with scheduling time on the equipment and answering my many questions. I’d also like to thank Andrew Loebl and his wife, Nora Schuette for their friendship and for helping me navigate the Fedkiw lab. A very special thank you to Dr. Fedkiw for being on my dissertation committee. Thank you for carefully editing my prelim as it was instrumental to the completion of my dissertation. I will also never forget that it was your class (ChE 711) that officially started my PhD journey. Dr. Sonia Grego has been instrumental in making sure I have the electrode chips I need to perform experiments. Thank you Sonia and RTI for providing our group with microelectrode chips and for helping to fund this PhD work. I also want to thank Dr. Michael Dickey and Dr. Glenn Walker for graciously agreeing to be on my PhD committees. Both have had very busy schedules and have always vi
been quick to respond to all inquiries, even while on vacation and being a new dads. Thank you, also, Dr. Dickey for making my last TA-ship fun and fulfilling. You are a great professor! Thank you to Dr. Chase Beisel, who within two weeks of my defense, offered to be a proxy for a committee member who was unable to attend my defense. A very special thank you to the ChBE office staff, but especially to Sandra Bailey, Saudra Doby, Shelia Hayes and Diane Harper. You guys made me feel at home every day and helped me to navigate the NCSU system. You also kept me well fed, especially during my pregnancy. You are all truly special and will always hold a dear place in my heart. The department is blessed to have you and I feel fortunate to have gotten to know you. I feel very fortunate to have Dr. Lisa Bullard as a mentor and friend. She has helped me to navigate my graduate career and giving me perspective into how to balance work and family as well as finding creative ways to pursue my life goals. Dr. Bullard, I appreciate the times you’ve welcomed me to your office, taken time to have lunch with me and for having enough faith in me to let me help you with the TA training workshops. I will never forget those Saturday morning sessions! Dr. Brodkey was my undergraduate research advisor, and I’m convinced, the reason I got into grad school. He is one of the most brilliant and life-loving people I know. His frank stories of his grad school experience and his constant reassurance that, “You will make it”, is always with me. Who knew that saving someone’s Ohio State Season Football tickets would turn into a grad school career? One of the things I was most nervous about in coming to grad school was choosing an advisor. All of my friends who had been in grad school unanimously agreed that your advisor selection could make or break your grad school career. I have never regretted my decision in choosing Dr. Velev as my number one choice in advisors. The group members and I are constantly in awe of the enthusiastic dedication, leadership and scholarship Dr. Velev has for the group. Thank you, Dr. Velev, for making the time for this dissertation defense to happen despite your incredibly busy schedule. Adam, Abby and I all appreciate how welcome you and Anka have made us feel over the years. vii
I would like to thank my parents for not killing me when I said I was going to quit my job at JM and come back to grad school. My parents have always shown me the importance of education and sticking with it. My mom returned to grad school when I was 16 and received her master’s in education by attending evening classes in 4 years. My Papaw, may he rest in peace, also got his PhD in Mechanical Engineering later in life when he was 48, showing me that it’s never too late to go back. I admire their dedication to education. Thank you, mom and dad for patiently listening to me complain about my devices that didn’t work. I’d also like to thank Lindsey (Jerrim) Stamm, who, despite the fact that she has enough of her own life-events looming, always has time to be my best friend. I am grateful for most of my experiences in graduate school and meeting Lindsey is in the top five. She and I met because we were in the same lab, but also because we share a deep affection for our pets. Lindsey has helped me through some of the most difficult times in my life and shown me what a true, gentile southern woman is like. Lindsey, thank you for being there for me, befriending me even in times when I was not the most friendly, and for teaching me how to make brownies that didn’t come from a box mix. I would like to thank my dog, Summer, for being ecstatic about the simple things in life. Whether it’s dinnertime or time for a walk, my dog reminds me that life isn’t as complex as we make it out to be. She also reminds me that fear of the unknown can be crippling if you don’t go and discover the root of it. Thank you for being my writing buddy, Summer! Thank you for lapping up the tears and dancing in celebration with me. I would like to thank my husband, Adam, for showing me love and patience through this process. He has had the unique opportunity to have been here before, so he has been coaching me along this process, trying to keep me focused, and mentally sound enough to keep going. He is such an inspiration to me as I see all he has to go through daily, especially over the last few years. He has taken on so much responsibility at home, making sure that Abby and I have clean floors, clothes and bathrooms. He steps in where ever he is needed and has made sure that I have everything I need without being asked. Adam (aka Capt. Awesome), without you, I would probably wearing a strait jacket in Raleigh’s Center for the Mentally Ill. You have faith in me when I don’t have faith in myself. Thank you for staying by my side, especially because I know it wasn’t always easy to do. You are my strength and viii
the best husband any woman could ask for. These have been the most challenging 5.5 years of my life and I can’t believe it is finally over. If I take nothing else away from this experience, being with you and having our daughter has definitely made this whole experience worth it. We made it. I can’t wait to start the next chapter in our lives together. Lastly, I’d like to thank my daughter, Abigail. Abby, although you were only present for the last few months of my PhD work, you were instrumental to helping me finish. Although others, even me, might have said you were a hindrance, don’t listen. You reignited the ability to get things done in a certain amount of time and kept me focused. When I had “free time” I wanted to make sure you were able to get my full attention. You aren’t going to remember this time, but I always will. Writing my dissertation while you were so young gave me a unique opportunity to be with you that a lot of working new mothers don’t get. My favorite memory will always be that you wouldn’nap on your own. So, for the first few months of your life, you napped on me while I did literature searches and read papers. You and your daddy are my life and I am so excited to see where our life takes us together now that the PhD graduate school chapter of my life is finished. I love you both so very much.
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TABLE OF CONTENTS
LIST OF TABLES ...... xi LIST OF FIGURES ...... xii CHAPTER 1: Introduction to Microfluidics, Electric Biosensors and AC Electrokinetics ...... 1 1.1. Introduction ...... 2 1.2. History of Lab-on-A-Chip Devices ...... 2 1.3. Electrical and Surface Properties of Cells ...... 5 1.4. AC Electrokinetics ...... 7 1.4.1. Electrokinetic Fluid Flow: Electrothermal and AC Electroosmosis .....8 1.4.2. Electrokinetic Particle Motion: Dielectrophoresis (DEP) ...... 11 1.5. Microfluidic Methods for Particle and Cell Manipulation and Separation...... 12 1.5.1. Manipulation ...... 12 1.5.2. Separations on a Chip ...... 15 1.6. Electrochemical Analytical Methods ...... 19 1.6.1. Impedance and Electrochemical Impedance Spectroscopy ...... 19 1.6.2. Cyclic Voltammetry ...... 22 1.7. Microfluidic Methods for Particle and Cell Detection ...... 25 1.7.1. Optical Detection ...... 25 1.7.2. Electrical Detection ...... 26 1.8. Layout of Dissertation...... 29 1.9. References ...... 31 CHAPTER 2: On-Chip Collection of Particles and Cells by AC Electroosmotic Pumping and Dielectrophoresis Using Asymmetric Microelectrodes ...... 46 2.1. Introduction ...... 47 2.2. Materials and Methods ...... 51 2.2.1. Sample Preparation ...... 51 2.2.2. Patterned Electrode Design ...... 51 2.2.3. Electrode Chip Fabrication ...... 52 2.2.4. Experimental Apparatus and Data Acquisition ...... 52 2.3. Results ...... 53 2.3.1. Bulk Fluid Flow Generation… ...... 55 2.3.2. Particle Collection Tuning Parameters ...... 57 2.3.3. Effects of Medium Composition and Particle Type ...... 72 2.4. Conclusions ...... 73 2.5. Acknowledgements ...... 74 2.6. References ...... 74 CHAPTER 3: Optimizing Coating and Composition of Metal Electrodes for AC Field Manipulation and Collection of Cells and Particles in Electrolytic Suspensions...... 78 3.1. Introduction ...... 79 3.2. Materials and Methods ...... 82 3.2.1. Particle and Cell Preparation ...... 82 3.2.2. Electrode Chip Fabrication ...... 82 x
3.2.3. Electrode Coating ...... 83 3.2.4. Experimental Apparatus and Data Acquisition ...... 86 3.3. Results and Discussion ...... 87 3.3.1. Effect of Type of Electrode Metal ...... 88 3.3.2. Hydrogel-Based Coatings ...... 92 3.3.3. Silane and LbL Polyelectrolyte Coatings ...... 95 3.3.4. On-Chip Collection of Large Particles and Cells...... 102 3.4. Conclusions ...... 104 3.5. Acknowledgements ...... 106 3.6. References ...... 106 CHAPTER 4: Particle-Cell Agglutination Based Impediometric Biosensor ...... 115 4.1. Introduction ...... 116 4.2. Materials and Methods ...... 119 4.2.1. Sample Preparation / Agglutination Techniques ...... 119 4.2.2. Yeast Viability Study ...... 120 4.2.3. Interdigitated Electrode Chip Fabrication ...... 121 4.2.4. Interdigitated Electrode Chip Types ...... 121 4.2.5. Impedance Measurements ...... 123 4.3. Results… ...... 125 4.3.1. Sample Preparation Observations ...... 125 4.3.2. Biosensor Design Considerations: Impedance Analysis ...... 131 4.3.3. Agglutination by Electrostatic Binding ...... 132 4.3.4. Impedance Analysis of the Process of Functionalized Particle Agglutination...... 135 4.3.5. Magnetic Field Manipulation of Functionalized Particles ...... 137 4.3.6. Equivalent Circuit Modeling of Agglutination Methods ...... 140 4.3.7. Interdigitated Electrode Chip Redesign...... 149 4.4. Conclusions ...... 151 4.5. References ...... 153 CHAPTER 5: Summary ...... 159 xi
LIST OF TABLES
Table 2.1 Electrode Design Configuration and Dimensions ...... 56
Table 2.2 Relative Particle Density for Electrode Repeat Designs at Various Chip Positions ...... 66
Table 3.1 Optimal Collection and Maximum Applied Potential and Frequency Settings for Platinum Electrodes...... 99
Table 3.2 Equivalent Salt Concentrations of Common Biological Buffers to Raleigh City Tap Water ...... 103
Table 4.1 Electrostatic Agglutination Samples of Yeast Cells and Particles as Pictured in Figure 4.8 ...... 127
Table 4.2 Bio-specific Binding Cell Viability for Varying Cell / Particle / Biological Salt Solutions Residence Times ...... 131
Table 4.3 Settling Times for Cells and Particle Types ...... 132
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LIST OF FIGURES
Figure 1.1 Integrated microfluidic devices combine microfluidic manipulation, separation and detection techniques. (a) Lab on a chip systems have been designed to bring whole cells into a detection chamber and perform all cell treatment steps: separation, incubation, purification lysis and gene identification.[21] (b) Lab on a chip assays treat the analyte with the proper antibodies and have more accurate result readings than traditional microarrays.[23] (c) Gel electrophoresis uses bulky equipment and gels that are friable, difficult to load and difficult to read. Burns et al. developed a device to perform all gel electrophoresis steps from sample preparation through analysis on one chip.[24]...... 4
Figure 1.2 Cells can act like a (a) traditional capacitor which is comprised of an insulator bounded by two conductive plates. (b) Cell membranes are comprised of amphiphillic molecules that form a thermodynamically favorable structure called a vesicle. Proteins are contained within the lipid bilayer and typically have a net negative charge. The charged surfaces surround the uncharged lipid bilayer like a capacitor ...... 6
Figure 1.3 AC electroosmosis occurs in asymmetric electric fields as a result of the charging of the electrical double layer. Fluid flow is tangential to the direction of the electric field.[49] ...... 9
Figure 1.4 Model of AC electroosmotic flow for varying distances from electrode (x) and frequencies.[49] For every electrolyte system, an optimal fluid velocity exists at a particular distance from the electrode, x, and frequency ...... 10
Figure 1.5 (a) AC electroosmotic flow is induced by asymmetric field gradients (red) that are driven by local eddies that form just above the electrodes and substrate. Bulk fluid flow is driven by the local eddies similar to that of a conveyor belt. Electric field asymmetry can be introduced by using (b) asymmetric planar electrodes and (c) asymmetric step electrodes. It is hypothesized that the step electrodes will induce higher bulk flow rates than that of the planar electrodes as the step electrode vortices do not impede forward bulk flow.[37] ...... 11
Figure 1.6 (a) When the particle is more polarizable than the media the particle will experience positive DEP and be drawn to the highest electric field gradient. (b) When the particle is less polarizable than the media, the particle experience negative DEP and is repelled from the electric field ...... 12 xiii
Figure 1.7 In microfluidic flows, x-direction and y-direction forces need to be overcome in order to direct cell manipulation in the positive x-direction. Negative x- direction forces include viscous drag or friction. The y-direction forces include buoyancy (positive) and gravity (negative) ...... 13
Figure 1.8 Examples of microfluidic cell and particle manipulation devices. (a) An electrically-driven fluid flows induced by asymmetric AC fields at the photoresist and silicon interface where particles and cells were drawn to a stagnant area in the center of a corral.[36] (b) Optical tweezers will generally collect one cell or particle at a time, but a system of multiple beams can trap and direct movement of a set of microparticles. Such a system was designed by Chiou et al.[79] (c) Using a zipper-like buss bar system, particles and cells on both the nano and microscale were collected using positive and negative DEP.[34]...... 15
Figure 1.9 Examples of microfluidic cell and particle separation devices. (a) A multi- branched microfluidic device separated particles of different sizes and densities. Two fluids of different densities, one of which is carrying the particle sample, are introduced to the chamber, meeting at a pinch point. Once the particle sample reaches the expanded, branched section at the end of the pinch point, the particles are drawn to different branches based on buoyancy.[87] (b) Specifically functionalized microwells used to trap HeLa and pancreatic cancer cells. Once trapped, the cells reproduced. Each well can also be actuated and removed individually for further analysis.[100] (c) Magnetic fields can be used to separate cells and particles with differing magnetic affinities by varying the magnetic field across a microfluidic channel.[102] (d) Cells and charged particles are polarizable and able to be separated using castellated electrodes by dielectrophoresis.[16] ...... 18
Figure 1.10 (a) Simple direct current (DC) circuit. (b) Simple alternating current (AC) circuit ...... 19
Figure 1.11 Phasor diagram for impedance. The real and imaginary impedance are related by the phase angle, ...... 20
Figure 1.12 Representative impedance spectra plots (a) Bode plot (b) Nyquist plot ...... 21
Figure 1.13 Experimental apparatus for studies involving thin-film metal electrodes and thin film coatings. (a) Top view. The Ag/AgCl reference electrode and the counter electrode (not pictured) are inserted at the top of the device into the xiv
electrochemical cell. The brass working electrode connection contacts the sample at the bottom of the electrochemical cell. (b) Side view. The electrochemical cell is the hollow tube at the center of the glass reservoir. The reference electrode and counter electrode do not contact the working electrode ...... 22
Figure 1.14 A typical cyclic voltammagram. The forward scan represents the reduction of the redox probe. At Ec, the oxidized form of the redox probe is depleted and the reaction shifts to oxidation. As the reduced form of the redox probe is depleted, the current becomes more negative until Ea is reached and the reduced form of the probe is depleted. Emax is typically chosen such that the potential at which Ec and Ea occur are within the bounds of the potential sweep...... 24
Figure 1.15 Examples of detection methods in microfluidic chambers. (a) Strip assays are the most common microfluidic detection systems, but are not as sensitive to dilute analytes. Rastogi et al. developed a droplet-based bioassay system immobilized using dielectrophoresis that can detect dilute samples of ricin.[145] Bringing cells and particles between electrodes increases the capacitance close to the electrodes and is detectable by measuring impedance. This has been done using (b) DEP and specifically functionalized particles and (c) a combination of DEP and specifically functionalized surfaces.[4, 31] (d) and (e) Optical waveguides use changes in the refractive index of the surface to detect cells and particles. A wavelength or incidence angle shift specific to particular particles and cell types indicates the presence of a desired analyte. These devices are most effective if the particle or cell sample is concentrated over the grating.[146-147]...... 29
Figure 1.16 Capabilities of AC electric fields cells at low voltage / high frequency and high voltage / low frequency ...... 29
Figure 2.1 (a) Top view of fully assembled planar electrode collection device. The square-wave AC current is applied to the magnified 5-repeat segment electrode design while the others are grounded. (b) Micrograph of 5-repeat segment electrode design. The collection area is 1 mm for each electrode repeat patterns (1, 3, 5, 10 repeats). (c) Enlargement of electrode configuration with electrode dimensions in microns. The repeat segments, starting at the outermost edge, include a thin electrode followed by a 5 m gap, and a larger 30 m electrode. Each segment is separated by a 15 m gap. (d) Enlarged side view of planar electrodes on glass substrate and induced ACEO flow. The xv
ACEO flow over the asymmetric electrodes is in the form a large eddy induced just above the edge of the large electrode adjacent to the small gap (adapted from Ramos et al.[29] ...... 50 Figure 2.2 Micrographs of typical particle collection results for 300 Hz, 800 mV at (a) 0 s, (b) 10 s, (c) 180 s, and (d) 360 s. A line of grouped particles, or particle front, is formed within the first few seconds of applying the AC voltage. The particle front traverses from the edge of the electrodes towards the center of the chip over the course of the trial, forming a rectangular deposit of densely packed particles in the middle of the collection area. The collection patterns are similar to those attained by Bhatt et al. [26], but thanks to the different principle are attained in 10% of the time ...... 55
Figure 2.3 Effect of chamber height on particle collection. (a) – (c) depict the top view in real time of particle collection device. The width of the collection area is 1mm. (a) At a relatively low height of 250 m, small, rapid eddies form directly over the electrodes and erratic eddies exist along the edge of the chip. Little, if any, particle collection was observed at the center of the collection area. (b) The optimal chamber height is approximately at a 1:1 ratio to the width of the collection area on the chip (750 m). This thickness leads to strong vortices and a well-defined stagnant are the center of the chip. (c) Higher chamber height allows for larger vortices to form over the electrodes, however, these vortices are strong enough to draw the particles away from the electrode chip surface. (d)-(f) show the side view of the COMSOL model output for geometry and dimensions analogous to corresponding particle collection experiments and present the corresponding fluid dynamics simulations ...... 58
Figure 2.4 Particle front velocity at 300 Hz as particles suspended in DI water travel from the electrode edge to the center of the collection area. The electrode edge is at 0, and the center of the chip is at 500 m. For lower potentials (250 – 300 mV), the bulk fluid velocity is insufficient to drag the particles to the center of the collection area within the 15 min trials. As potential increases, the bulk fluid velocity increases, drawing the particles to the center ...... 62
Figure 2.5 Pixel density analysis of ACEO particle collection in DI water at 300 Hz, (a) 300 mV, (b) 400 – 600 mV and (c) 700 – 800 mV after 280 s of field application in terms of percent deviation from the initial average pixel color value. The solid points are experimental data while the dashed line is its trend. (a) At 300 mV, very little particle movement and a small, slow particle front was observed, resulting in little change in the particle density throughout the collection area. (b) From 400 mV to 600 mV, a faster moving, more distinct, particle front was observed, but it was not rapid enough to reach the xvi
collection area center. (c) At 700 mV to 800 mV the efficient collection of particles at the center was confirmed by a positive %DFW ...... 63
Figure 2.6 Particle front velocity analysis in the low-frequency regime at 800 mV. DI water suspensions require low potentials and frequencies to induce ACEO bulk fluid flow. Very little change in the particle front velocity is observed for small frequency changes between 300 Hz and 1 kHz. As the frequency was raised above 2.5 kHz, DEP was observed, reducing the amount of particles in the collection area and reducing the apparent particle front velocity ...... 65
Figure 2.7 Pixel density analysis of observed ACEO phenomena in DI water at 300 Hz, 800 mV for 1, 3, 5, and 10 electrode repeats. The bulk fluid flow velocity driven by on electrode repeat will collect cells just beyond the large electrode, but is not sufficient to draw particles to the collection area center. Three repeats will induce flows that draw the particles only 250 m towards the collection area. Although 10 electrode repeats does act much like the 5 repeat device at first, after 100 s, the particles appear to be lifted from the corral center surface ...... 67
Figure 2.8 COMSOL finite element analysis models and micrographs taken at 280 s for increasing number of electrode pattern repeats: one (a,b), three (c,d), and ten (e,f). The bulk fluid flow eddies formed as a result of ACEO increase in size and velocity as the number of repeats increases. None of the repeat designs resulted in densely packed collection of particles as in the five repeat design (Figure 2.3b,e and Table 2.2) ...... 68
Figure 2.9 Micrographs and pixel analysis for circle geometry electrode segment repeat trials at 300 Hz, 800 mV. The %DFW analysis indicates similar results to the rectangular electrode geometry in that an optimum x-direction fluid velocity is needed to form a densely packed bed of particles in the center of the collection area. The 5-repeat design appeared to induce the optimum fluid velocity to draw 1 m particles to the collection center. Despite the lower %DFW range than in the rectangular case, (a) integrating the area under the %DFW curves for both the rectangular and circular geometries indicates the circular geometry is capable of drawing a larger amount of particles into the collection area. This is further seen by comparing the micrographs of the (b) 5-repeat rectangular design and (c) 5-repeat circular design results at 280 s ...... 70
Figure 2.10 Particle front velocity for latex particles suspended in DI water and 0.001 mM PBS at 500 Hz, 1 V ...... 72
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Figure 3.1 (a) Top view of planar electrode collection chip. The collection area is 1 mm for each electrode repeat patterns (1, 3, 5, 10 repeats). The square-wave AC current is applied to one repeat segment while the others are grounded. (b-f) Side view enlargement of electrode chip with coating schematic for each type of thin film coating. Each coating includes a thiolated crosslinker to strengthen the bond of the polymer to the metal electrode. (b) The AquaSil silane coating is crosslinked with MPTES. (c) The LbL polyelectrolyte coating is crosslinked using MUA and comprised of 5 layers of alternating PSS and PDAC. (d) To ensure that the agarose is spread to a thin film during the spin coating process, the chips were first coated with MPTES and AquaSil. (e-f) Polyelectrolyte agarose coatings were doped with PSS and PDAC respectively before being layered over the silane. Not featured: polyacrylamide coatings ...... 85
Figure 3.2 Micrographs of uncoated gold and platinum electrodes before and after electrode damage. (a) Undamaged gold electrodes. (b) Gold electrodes damaged at 3V, in 0.001 TE electrolyte, evident by the light areas along the length of the electrode. Some of the damage on the small electrode is enough to render the device useless. (c) Undamaged platinum electrodes. (d) Platinum electrodes damaged in 10 mM PBS at 800 mV, evident by the dark purple and brown areas of the large electrode. For a reference to scale, large electrode is 30 m wide ...... 89
Figure 3.3 Electrokinetic phenomena diagram for gold electrodes. Experiments for uncoated gold in 1 TE and 1 mM PBS as well as silane coated gold in 1 mM PBS were not performed as electrodes were prone to rapid damage ...... 91
Figure 3.4 Electrokinetic phenomena diagram for platinum electrodes. Experiments for uncoated platinum in 1 TE and 1 mM were not performed as electrode damage was eminent ...... 92
Figure 3.5 Cyclic voltammetry curves following ferric cyanide reaction on bare gold (black) as well as silane (red) and LbL polyelectrolyte (green) coated gold. The peak at 0.3 V and 0.01 V for the gold curve exemplifies the oxidation and reduction reaction respectively. The lack of peak for the silane and LbL polyelectrolyte coated slides suggests the faradaic reaction has been sufficiently suppressed ...... 97
Figure 3.6 Particle collection for LbL polyelectrolyte coated platinum electrodes at (a) 0 s, (b) 10 s, (c) 140 s, and (d) 280 s at 800 mV, 300 Hz. The particle collection xviii
at the centrally located stagnant point is typically preceded by a distinct particle front that forms at the edges of the innermost electrode and traverses towards the center over time ...... 98
Figure 3.7 Particle density analysis at 800 mV, 300 Hz for uncoated (black), silane coated (red) and LbL polyelectrolyte coated (green) platinum electrode chip in DI water after 280 s of particle collection ...... 100
Figure 3.8 Time dependence analysis of particle density at 4 V, 10 kHz for uncoated, silane coated and LbL polyelectrolyte coated platinum electrode chip in 1 mM PBS ...... 101
Figure 3.9 (a) Collecting yeast cells with the electrode device was attempted with uncoated gold electrodes at 5V, 200 Hz in DI water. The cells tended to stick to the substrate. Within 80 s of applying potential, bubbles formed over the electrodes. (b) Gold lift-off occurred on the smaller electrodes, especially at the southwest corner. (c-f) Micrographs of particle collection trials of large particles and yeast cells in tap water with platinum polyelectrolyte coated electrodes. The 5 m latex particle trials were performed by toggling between 2 V and 4 V at 50 kHz. Particle collection is apparent after (d) 15 mins over (c) the initial micrograph. (e-f) Yeast cells were collected at 6 V 50 kHz after 15 mins. No electrode damage was apparent after either particle or cell collection trial with the coated electrodes ...... 104
Figure 4.1 Matrix depicting agglutination via pH. At a pH above the isoelectric point, the surface charge of cells is negative and negatively charged particles will be repelled. At a pH below the isoelectric point (pH = 4 for yeast cells), the surface charge of the cells will be positive and attract negatively charged particles ...... 118
Figure 4.2 Yeast viability study and growth judging examples. Judging criteria was as follows: 0 – no observable growth, 1 – growth on less than 50 % of the line, 2 – growth on 50%-75% of the line, 3 – full line with a few spots (greater than 75% of the line full) and 4 – full line of growth ...... 121
Figure 4.3 Gold on glass interdigitated electrode chip used in agglutination impedance spectroscopy measurements. The dual interdigitated electrode sets consisted of 25 electrodes on each side separated by 10 m gaps ...... 122 xix
Figure 4.4 Gold on glass interdigitated electrode chip used in agglutination impedance spectroscopy measurements. The dual interdigitated electrode sets consisted of 25 electrodes on each side separated by 10 m gaps ...... 123
Figure 4.5 Photograph of electrode chip with soldered connections used for impedance spectroscopy measurements ...... 124
Figure 4.6 Side view of chamber for settling impedance experiments. The suspensions were loaded in the chamber from the top of the chamber and allowed to settle ...... 125
Figure 4.7 Magnetic particles can be manipulated with magnetic fields. The electrode sample chamber was inverted and a magnetic field induced just over the electrodes. (a) Unfunctionalized particles only were manipulated by the magnetic field and drawn between the electrodes. (b) For Con A magnetic particles, networks were formed with the yeast cells. Both magnetic particles and yeast cells were brought between the electrodes when the magnetic field was applied...... 125
Figure 4.8 Sample vials for agglutination via pH change. (a) Dispersed sample vials of (1) yeast + latex at high pH, (2) yeast and gold at high pH, (3) yeast + latex at low pH, (4) yeast and gold at low pH. (b) Settled sample vials of (1) yeast + latex at high pH, (2) yeast and gold at high pH, (3) yeast + latex at low pH, (4) yeast and gold at low pH. Yeast only samples not pictured ...... 127
Figure 4.9 Microcetrifuge tubes of (a) dispersed cells and particles and (b) settled cells and particles for (1) yeast only, (2) yeast and unfunctionalized gold nanoparticles and (3) yeast cells and Con A functionalized gold nanoparticles ...... 128
Figure 4.10 (a) Yeast cells and Con A functionalized magnetic particles after 5 minutes, (b) yeast cells and Con A functionalized magnetic particles after 30 mins, (c) yeast cells and unfunctionalized magnetic particles after 5 minutes ...... 130
Figure 4.11 NPIC versus time for low pH agglutination trials at 60 kHz. The yeast and latex sample NPIC changes little after 70 minutes. The yeast / latex sample xx
(red) is distinguishable from the bare yeast (black) and gold nanoparticle / yeast (green) samples as they settled faster than the other agglomerates ...... 133
Figure 4.12 NPIC vs. frequency for pH agglutination technique (a) real impedance, (b) imaginary impedance at 70 minutes. The low pH samples are distinguishable from the high pH samples with the agglutinated yeast and latex particles exhibiting the greatest NPIC ...... 134
Figure 4.13 NPIC vs. frequency for bio-specific binding with gold nanoparticles (a) real impedance, (b) imaginary impedance at 30 mins. Yeast only (black) and yeast / unfunctionalized (red) particle suspensions exhibit similar impedance spectra, indicating little difference between the two suspensions. The electrochemical impedance spectra of the yeast / functionalized particle is distinct from the other two suspensions, especially above 10 kHz. The slope of the imaginary impedance curve is also less steep with respect to frequency and can also to be used to distinguish the agglutinated suspension from the non-aggluinated suspensions ...... 136
Figure 4.14 Inverted interdigitated electrode chip with unfunctionalized magnetic particles and yeast cells. When a magnetic field was applied, the magnetic particles formed chains. A few cells were observed to adhere to the glass substrate ...... 138
Figure 4.15 NPIC vs. log frequency for biospecific binding with magnetic particles (a) real impedance, (b) imaginary impedance at 15 mins. Yeast only (black) and yeast / unfunctionalized (red) particle suspensions exhibit similar impedance spectra, indicating little difference between the two suspensions. The electrochemical impedance spectra of the yeast / functionalized particle is distinct from the other two suspensions, especially above 100 kHz...... 139
Figure 4.16 (a) Equivalent circuit and (b) model fit for pH agglutination study at 70 mins. This model is similar to those used in other surface bound impedeometric biosensors[16-17] and is relatively simple involving two constant phase elements (CPE), solution resistance (Rs) and electrode capacitance (Celectrode) ...... 142
Figure 4.17 (a) Equivalent circuit and (b) model fit for pH agglutination study. This model is similar to those used in other surface bound impedeometric xxi
biosensors[16-17] and involves two constant phase elements (CPE), solution resistance (Rs) and electrode capacitance (Celectrode) ...... 143
Figure 4.18 (a) Equivalent circuit and model fit for (b) Con A functionalized gold nanoparticle and (c) Con A functionalized magnetic particle agglutination studies at 30 and 15 mins respectively. This model includes a resistance and capacitance component near the electrode surface (R3,C3) and two sets of components between the electrode and bulk solution (R1, R2, C1, C2) ...... 145
Figure 4.19 (a) Bulk solution resistance of the ConA functionalized gold experiments with PBS alone, yeast alone, yeast and unfuctionalized particles, and yeast / ConA functionalized particles. Due to settling and the networks formation, the Con A functionalized particles and yeast cell networks exhibit just slightly higher resistances and contribute to the solution resistance. (b) The capacitance of the components from the electrode surface to just above the electric double layer. The capacitive component closest to the electrode surface C1 is predominant from 100 kHz to 500 kHz. Components at and just above the electric double layer, C2 and C3 are evident in the NPIC from 10 kHz to 100 kHz and 1 kHz to 10 kHz respectively ...... 147
Figure 4.20 (a) Bulk solution resistance of the ConA functionalized gold experiments with PBS alone, yeast alone, yeast and unfuctionalized particles and yeast / ConA functionalized particles. Due to settling and the networks formed, the Con A functionalized particles and yeast cell networks exhibit just slightly higher resistances and contribute to the solution resistance. (b) The capacitance of the components from the electrode surface to just above the electric double layer. The capacitive component closest to the electrode surface C1 is predominant from 100 kHz to 500 kHz. Components at and just above the electric double layer, C2 and C3 are evident in the NPIC from 10 kHz to 100 kHz and 1 kHz to 10 kHz respectively ...... 148
Figure 4.21 Imaginary impedance spectra from 1 kHz to 1 MHz for yeast only (black), unfunctionalized magnetic particles / yeast (red) and Con A functionalized magnetic particles (green) suspended in DI water at 15 minute time point. The distinctive difference between the functionalized and unfunctionalized particle samples between 10 kHz and 100 kHz indicates cells are present on the electrode array surface ...... 150
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1
CHAPTER 1
Introduction to Microfluidics, Electric Biosensors and AC Electrokinetics
2
1.1 Introduction In a time when cancer, pandemics, food contamination, environmental contamination, and biological terrorism are getting more public attention, there is a growing demand for rapid and accurate detection of certain cell types and viruses. In 2003, biotection was reported to be a rapidly growing market at $563 million with a compound annual growth rate of 4.3% with four major sectors: food, military, environmental and medical.[1] The analytes vary in size and shape from DNA to specific cell types for each application, however they each face the same challenges when trying to detect biological species: the tests available to date are slow, tedious, and inaccurate. Most of these lab bench macroscale techniques require large instrumentation and produce large amounts of biological waste. Developing new microscale devices for cell sample preparation, separation, collection and detection is becoming increasingly important. Microscale methods, replacing those performed on the macroscale, could lead to lower material costs and power consumption. In most cases, media, nutrients and cell lines can cost hundreds of dollars per unit or can only be collected in small volumes. Lower power consumption makes possible hand-held, battery-powered devices. Further, studying cells in lab-on-a-chip devices allows researchers to take advantage of the microscale properties of cells. For example, biological particles have many of the characteristic properties of their synthetic particles and have been found to be easily manipulated by electrically induced forces.[2-17] By taking advantage of these and other microscale properties, biological systems can be integrated with man-made systems to build new devices, develop new detection techniques, and make advancements in medical research. Although the task of taking these research-driven micro biodetection devices to market is challenging[18], the foundation has certainly been laid over the last 20 years to develop micro biodetection devices to make a deep impact in the biodetection market.
1.2 History of Lab-on-A-Chip Devices Advances in chemical sensing and electrons ultimately lead to the technology used in most modern biosensors. Chemical sensors were being developed as early as the late 1940’s, but it wasn’t until 1953 when LC Clark developed a sensor to continually measure blood 3
oxygen levels during cardiovascular surgery that the biosensor race began.[19] Less than a decade later, doctors were using similar electrode systems to measure other gases and glucose levels for diabetics. The advent of the microchip and fabrication methods used to make them gave rise to fabrication techniques like photolithography, chemical etching and micropatterning. By the late 90’s, miniature biosensor development was in full swing. Non- electrically driven biosensors were becoming commonplace in households with the great improvements in the strip assays used for home pregnancy tests. These tests have become so popular that they are now used in environmental, food pathogen and military applications. As biosensor innovation continues into the new millennium, so does the progress in micro total analysis systems (TAS), where entire lab functions are combined on one device. Micro total analysis system (TAS) devices were introduced as early as 1989 for monitoring calcium ions.[20] Since then, the analytes and microfluidic functions have become increasingly complex. New advances in microfluidic device fabrication make possible the miniaturization of the entirety of lab functions on one chip. Cell, DNA and protein research typically requires many tedious steps such as cell lysis, separation of cell contents, and DNA or protein immobilization. Lab on a chip devices have been designed to perform each of these steps within minutes as opposed to the hours or days required by comparable macroscale lab analysis steps.[21-22] Lien et al. recently introduced a lab on a chip device that performs all functions necessary to detect cancerous cells from isolation to lysis and gene identification within 10 minutes[21] (Figure 1.1a). Lab on a chip integrated devices have also been designed to replace tedious microarray assays that are difficult to read[23] (Figure 1.1b). Much biological research uses gel electrophoresis which is expensive, difficult to load properly and difficult to analyze. Burns et al. developed an all-in-one DNA sample gel electrophoresis system[24] (Figure 1.1c). Such systems have also been used to help identify and determine effective treatments for highly drug-resistant cancers.[25-26] In using microfluidic devices for cell study, the miniaturization makes it possible to combine many of the preparation, manipulation and analysis steps in one tiny location. Microscale electric manipulation, separation and detection methods can be used to take advantage of a cell or particle surface charge or resistivity to enhance performance. 4
Figure 1.1. Integrated microfluidic devices combine microfluidic manipulation, separation and detection techniques. (a) Lab on a chip systems have been designed to bring whole cells into a detection chamber and perform all cell treatment steps: separation, incubation, purification lysis and gene identification.[21] (b) Lab on a chip assays treat the analyte with the proper antibodies and have more accurate result readings than traditional microarrays.[23] (c) Gel electrophoresis uses bulky equipment and gels that are friable, difficult to load and difficult to read. Burns et al. developed a device to perform all gel electrophoresis steps from sample preparation through analysis on one chip.[24] 5
1.3 Electrical and Surface Properties of Cells Any charged particle has manipulatable and measurable electric properties. The ability for a substance to conduct or polarize in electric fields has also been widely used in biological cell and particle system characterizations.[3,9,11-12,14-15,27-32] Although direct current (DC) can be used, many electric properties are most apparent in alternating current (AC) electric fields. AC electric fields have been used to manipulate ions in solution to induce microfluidic flows (AC electroosmosis and electrothermal) or to manipulate polarizable particles (dielectrophoresis).[4-7,10,13-14,17,33-38] The parameters used in AC circuit measurements are discussed in Section 1.6. The electrical properties of live cells are commonly discussed in the context of neural activity. The conductivity of neurons is apparent in their function, but all cells have distinct electrical properties based on their membrane structures. The distinct difference between the two main cell types, prokaryotic and eukaryotic is that the less evolved prokaryotes have an outer plasma membrane comprised of polysaccharides and that they lack organelles. Some bacteria are classified as either gram-negative or gram-positive in which gram-negative bacteria are surrounded by an additional lipopolysaccharide layer. All cells, whether eukaryotic or prokaryotic, are surrounded by a lipid bilayer membrane. Lipids are amphiphillic molecules with a polar head and non-polar tail (Figure 1.2a). At certain concentrations, lipids will form a thermodynamically favorable vesicle structure due to the interaction of the polar head and non-polar tail of the molecule[39] (Figure 1.2b). Because the vesicle is formed with strong, hydrophobic and electrostatic interactions, the bilayer is stable. Proteins crucial to cell function reside in the bilayer of both cell types as well as the polysaccharide layer of prokaryotes, including receptors that target specific molecules and channels that allow substances to traverse the membranes (Figure 1.2d). The presence of integral and surface proteins gives rise to diverse dielectric, capacitive, resistive and electrostatic properties.[40]
6
Figure 1.2. Cells can act like a (a) traditional capacitor which is comprised of an insulator bounded by two conductive plates. (b) Cell membranes are comprised of amphiphillic molecules that form a thermodynamically favorable structure called a vesicle. Proteins are contained within the lipid bilayer and typically have a net negative charge. The charged surfaces surround the uncharged lipid bilayer like a capacitor.
Proteins move freely within the lipid bilayer and, in the presence of an AC electric field are partially responsible for the polarization of the cell surface. Further, the integrated proteins have large dipole moments and are highly polarizable.[41-43] Similar to their polarizable synthetic particle counterparts, cells have the propensity to be more polarizable than the media in which they are contained and can be manipulated using DEP (See Section 1.4.2).[4-7, 10, 13-14, 17, 33-38] Cells can also act as capacitors and resistors.[40, 44-45] In solid state electronics, resistors are electric circuit components that resist electron flow while capacitors are comprised of an insulating layer between two conductive plates that is able to hold charge when subjected to a current (Figure 1.2c). Whether a cell is resistive, capacitive or a combination of the two depends on the frequency of the applied potential and the suspension media.[40] Increased salt concentration in the media equilibrates electrolyte diffusion across the membrane. At low frequencies, the membrane insulates the cell interior from applied 7
potential and acts as a resistive sphere.[40] At higher frequencies, the electric field can penetrate the lipid bilayer, diminishing the cell membranes resistive properties and allowing charge to be stored in the membrane and to transfer into the cytoplasm.[40] However, due to the small concentration differences of ions and charged molecules, the membrane capacitance is only apparent within micrometers of the membrane.[46] For most suspended cells, the accepted value for membrane capacitance is approximately 1 F/cm[47], but can vary by ± 0.5 – 2 F/cm when in suspension or nutritive media.[39] For example, the membrane capacitance of yeast cells was measured experimentally to be 0.66 F/cm2 .[45] The surface charge of a cell can also be altered by changing the net charge of the suspension media, typically by increasing or decreasing the pH. Commonly, the terminus of the proteins contained within the lipid bilayer is negative; therefore the surface charge on a cell is negative. Most proteins are amphoteric and have negative net charge at a neutral pH. When a protein is subjected to a lower pH, the net charge on the molecule will decrease until the isoelectric point, the point at which a molecule has zero net charge, is reached. Beyond this point, as the pH continues to decrease, the net charge of the protein is positive.
1.4 AC Electrokinetics Electrokinetics is the manipulation of charged species by electrical fields, resulting in fluid or particle motion.[48] Varying electrokinetic phenomena can be seen for both direct current (DC) and alternating current (AC) based on the directionality of electron flow. For example DC electroosmosis and AC electroosmosis (ACEO) involve the motion of ions in solution that drag solution molecules, resulting in fluid flow. The difference between DC and AC electroosmosis is in the mechanisms of charge build-up due to the unidirectional and alternating ion flows. Due to the alternating direction of the current, AC electrokinetic phenomena are frequency dependent and each exist in somewhat unique frequency domains. They can be categorized in two major types: those that result in fluid flow and those that result in particle motion.
8
1.4.1. Electrokinetic Fluid Flow: Electrothermal and AC Electroosmosis Ionic solutions are commonly measured using conductivity, the ability to conduct electrical current; and permittivity, the flux of an electric field. When in contact with an electrode with a potential applied, the conductivity and permittivity of the solution will change locally as the ions move. Applied potentials can also induce local non-uniform heating. Electrothermal flow results from the conductivity and permittivity gradients produced from this non-uniform heating. This typically occurs at relatively high frequencies above 10 kHz in high electrolyte concentrations or at low frequencies in lower electrolyte concentrations.[49] AC electroosmosis involves the electrically induced movement of ions in solution leading to fluid flow. When a charged surface is in contact with an aqueous electrolyte solution, the charged surface will attract the oppositely charged ions from the electrolyte.[39] The adsorption of oppositely charged ions ultimately leads to the formation of the electric double layer comprised of the Stern layer and diffuse layer. The Stern layer is one hydrated ion thick, immobile and comprised of counterions to the charged electrode. The highly concentrated mobile counterionic layer just above the Stern layer is the diffuse layer. When potential is applied to the electrodes, charge builds in the electric double layer. When a particular potential threshold is reached, the net negative and positive charges on the electrodes induce an electric field from the negatively charged electrode to the positively charged electrode. The tangential component of the electric field exerts a force on the ions in solution, inducing fluid movement occurring tangential to the electric field. AC electroosmosis only takes place in a non-uniform electric field. The velocity of the fluid is dependent on the potential drop across the double layer, the frequency of the electric field and the charge of the double layer[49] (Figure 1.3).
9
Figure 1.3. AC electroosmosis occurs in asymmetric electric fields as a result of the charging of the electrical double layer. Fluid flow is tangential to the direction of the electric field.[49]
AC electroosmosis typically occurs between frequencies of 10 Hz and 10 kHz. Fluid flow could occur at higher frequencies, only due to electrothermal effects. The fluid velocity in AC electroosmosis is frequency-dependent as shown in Equation 7. For a non-dimensional frequency,
1 m 2 (Eq. 8) where is the reciprocal of the Debye length, m is the permittivity of the medium, is the conductivity of the medium, and is the angular frequency. The diffuse layer potential, d, for an applied voltage, Vo is:
V o (Eq. 9) d 21 i the average AC electroosmotic fluid velocity is:
1 22 u mo 2 (Eq. 10) 8 x 12
10
where x is the relative distance from the electrode, o is the surface potential and is the fluid viscosity. AC electroosmotic flow has both an upper limit and optimal velocity that is frequency and spatially dependent (Figure 1.4).
Figure 1.4. Model of AC electroosmotic flow for varying distances from electrode (x) and frequencies.[49] For every electrolyte system, an optimal fluid velocity exists at a particular distance from the electrode, x, and frequency.
AC electroosmosis can be used to promote bulk fluid flow in microfluidic devices, as long as asymmetry in the electric field gradient exists. The electric field asymmetry can be induced by physical roughness of a substrate surface or by the electrode arrangement.[50] In the case of electrode arrangement, the asymmetric field creates a vortex near the edge of the electrode adjacent to the electrode and substrate surface. Such eddies drive bulk fluid flow similar to a conveyor belt (Figure 1.5a). The most common electrode configurations for microfluidic pumping are based on planar asymmetric electrodes (Figure 1.5b) and by step asymmetric electrodes (Figure 1.5c). Ajdari modeled systems of planar electrodes for their relative fluid flow velocities. At low voltages, the bulk flowrate is related to the square of 11
the current potential.[50-51] Bazant et al. later modeled step electrode configurations for microfluidic pumping.[52] The advantage of using step electrodes is that the vortices created by the asymmetric electric field will be induced between the walls of the step electrodes. The vortices still contact the bulk fluid and induce flow in one direction only, reducing backflow. The step configuration is reported to possibly increase bulk flowrates up to 17- fold.[52]
Figure 1.5. (a) AC electroosmotic flow is induced by asymmetric field gradients (red) that are driven by local eddies that form just above the electrodes and substrate. Bulk fluid flow is driven by the local eddies similar to that of a conveyor belt. Electric field asymmetry can be introduced by using (b) asymmetric planar electrodes and (c) asymmetric step electrodes. It is hypothesized that the step electrodes will induce higher bulk flow rates than that of the planar electrodes as the step electrode vortices do not impede forward bulk flow.[37]
1.4.2. Electrokinetic Particle Motion: Dielectrophoresis (DEP) All particles are able to polarize, when inserted inside an electric field. The forces on the particle directing its motion via an induced dipole are called dielectrophoretic (DEP) 12
forces. When subjected to a non-uniform AC electric field, a dielectric particle can be drawn towards or away from the region of electric field with the greatest intensity. If a particle is more polarizable than the media into which it is suspended, it is drawn towards the region of the highest electric field (positive DEP) (Figure 1.6a). Conversely, if a particle is less polarizable than the media, the particle is repelled from the highest region of the electric field (negative DEP) (Figure 1.6b). The dielectrophoretic force is dependent on the intensity of the electric field (|E|), the radius of the particle (r) and the effective permittivity of the [49] particle in the media (p, m) (Eq. 7).
pm 2 F r3 Re E (Eq. 7) DEP m 2 pm
Figure 1.6. (a) When the particle is more polarizable than the media the particle will experience positive DEP and be drawn to the highest electric field gradient. (b) When the particle is less polarizable than the media, the particle experience negative DEP and is repelled from the electric field.
1.5 Microfluidic Methods for Particle and Cell Manipulation and Separation 1.5.1. Manipulation One of the first challenges in designing biosensors is delivering the dispersed cell and particles to the detection area of the device. All cells are prone to sticking to microfluidic walls and substrate surfaces. Few cell types are suspension-based. Adherent cell lines require treatment as well as certain media additives for the cells to remain in suspension. Further, some common materials used to construct or coat microfluidic devices are toxic to biological cells.[53-54] Thus, microfluidic devices for biological applications need to be designed with 13
biocompatible materials that reduce biofouling and direct cells to their desired destination on the biosensor chip while keeping cells in suspension. Directing cell motion by fluid flow is one of the most common modes for cell transport. Natural examples of such flows can be found in natural systems like the human cardiovascular system of capillaries as well as the vascular systems of leaves. In microfluidic chambers cells and particles are usually carried by fluid flows in the laminar regime[2] and therefore must overcome gravitational forces, buoyancy forces, friction and viscous drag forces[39,49] (Figure 1.7). Generating microfluidic flows, requires a gradient or pressure differential across the channel, generally instigated with a pump. Just as with macroscale pumps, microfluidic pumping systems have used positive displacement and dynamic systems[55] and commonly include piezoelectric[56-60], electric[37-38, 61-67], and magnetic[68-74] actuation.
Figure 1.7. In microfluidic flows, x-direction and y-direction forces need to be overcome in order to direct cell manipulation in the positive x-direction. Negative x-direction forces include viscous drag or friction. The y- direction forces include buoyancy (positive) and gravity (negative).
Once a cell or particle sample is inserted into a biosensing device, dilute cell samples must be concentrated to increase the effectiveness of the detection, especially in surface bound methods. Just as particles can be carried by unidirectional fluid flows, they can be focused using specific flow patterns. In these cases, electrically driven flow or AC electroosmosis is generally most desirable as the electrodes that drive low velocity flows are relatively easy to pattern and, therefore, produce predictable and tunable flow fields. Electrically driven fluid flows manipulated with precisely placed and designed electrodes can 14
be used to collect cells and particles in predictable locations on substrates[36, 75-76] and have been especially effective in increasing cell counting efficiency[76] (Figure 1.8a). Some cell focusing techniques take advantage of cells and particle’s unique surface, size or shape properties. Both optical and electrical techniques have been used successfully to precisely guide cell placement. Because the directed assembly takes advantage of inherent properties of cells and particles, these methods eliminate the need for optical tags used to help identify cells that may alter a cell’s function. Directed lasers, optical tweezers, use the refractive index differences to trap dielectric microparticles[77-81] (Figure 1.8b), but have a very small collection area and may not be effective for dilute or highly concentrated cell samples.[79]. Optical focusing techniques are also effective in trapping one cell or particle at a time. Electrical techniques are much more effective at trapping multiple particles at once. Dielectrophoresis could trap cells and particles based on the dielectric properties of cells and particles with respect to the suspension media. Just as with AC electroosmotic focused flow, electrodes can be arranged to direct cell and particle trapping. The electric field generated between the electrodes will drive cells and particles towards or away from the substrate surface. DEP can be used to trap particles from the nanoscale and microscale. Hoettges et al. developed a dual circular electrode system to trap micro- and nanoparticles in circular electrode patterns by DEP [34-35](Figure 1.8c). Krishnan et al. successfully trapped DNA- coated nanoparticles using hydrogel coated platinum electrodes.[82-84]. Trapping desired analytes in raw samples using electric fields as it can be done with little to no sample pre- treatment. Balasubramanian et al. developed a simple gold plated electrode device capable of concentrating various bacterial cells from water samples.[85].
15
Figure 1.8. Examples of microfluidic cell and particle manipulation devices. (a) An electrically-driven fluid flows induced by asymmetric AC fields at the photoresist and silicon interface where particles and cells were drawn to a stagnant area in the center of a corral.[36] (b) Optical tweezers will generally collect one cell or particle at a time, but a system of multiple beams can trap and direct movement of a set of microparticles. Such a system was designed by Chiou et al.[79] (c) Using a zipper-like buss bar system, particles and cells on both the nano and microscale were collected using positive and negative DEP.[34]
1.5.2. Separations on a Chip Raw cell samples taken from humans, animals, food or the environment typically require some degree of pretreatment before detecting or studying a cell type. Blood samples contain many different cell types while food and environmental samples may contain particulates that could be mistaken for a desired analyte. Isolating certain cell types is essential to increasing the accuracy of many microscale detection methods. One of the most commonly used methods to sort and count cells is flow cytommetry. Cells are treated with 16
phospholuminescent tags. Once these tags trigger a detector, a valve is actuated that changes the flow direction and the desired cell type is trapped in a chamber separate from other cells. These methods are only about 30% efficient[86] in separating desired cell types. Further, in some cases, the optical tags used alter the cell’s function. Just as with manipulation methods, separation methods utilize physical, chemical, magnetic and electrical properties to divide cells and particles into desired groupings, ideally without pre-treatment or optical tags. Other mechanical and acoustical separation techniques involve using size or density differences, typically by varying flow fields with varying channel widths[87-89] or microfilters.[90] Takagi et al. developed a multi-channel microfluidic device to redistribute samples of varying particle diameters and densities[87] (Figure 1.9a). The raw sample of particles is introduced into one channel while a fluid of a higher density is introduced in an adjacent channel. When the two flows meet in a constricted portion of the channel, the particles are all aligned at the top of the channel. Branched microfluidic channels at the end of the constricted segment collect particles of different sizes based on buoyancy.[87] Hydrodynamic flow filters use membranes and blocked flow chambers of different sizes to sift particles based on diameter.[91-94] Particulates and cells can also be separated based on the elasticity of their surfaces. For animal cells, the lipid bilayer is relatively elastic. When sound waves are applied to suspension of cells or particles with varying surface elasticities, they will separate based on how the sound wave is adsorbed or reflected from the membrane or particle surface. This acoustical separation method has proven effective in isolating cells from solid particles as well as lipids from eyrthrocytes.[95-97] Many biological particle separation methods use biospecific or chemical differences of ligands on cell surfaces to differentiate them. In many cases, a substrate is coated with a biologically active agent that is specific to the cell type being targeted.[4, 30, 98-99] These methods have proven especially useful for chromatography-like applications.[98-99] More recently, bioseparation methods using specifically functionalized microwells have been used to trap HeLa and pancreatic cancer cells (Figure 1.9b). The microwells allowed for the cells to reproduce, but what makes this technique unique is each microwell was detachable from the substrate for further analysis.[100] 17
Both magnetic and electrical separations involve manipulation by an induced field. Magnetic functionalized particles are commonly used to separate proteins and ligands from cell lysates in chromatography columns. In microfluidic devices, magnetic fields have been used to separate particles of different size[101-102] (Figure 1.9c). Red blood cells have also been isolated from whole blood samples with magnetophoresis, the manipulation of magnetic particulates with magnetic fields, by the paramagnetic properties of the heme groups.[101] Particles of different shapes, diameters, and materials were easily manipulated between electrodes excited by AC currents via dielectrophoresis.[33, 103] Cells have varying sizes, surface properties and shapes that allow each species to have unique electrical properties. In taking advantage of these electric properties, both interdigitated and castellated electrodes have been developed to separate cell types suspended in small chambers by dielectrophoresis[8, 16, 104-106] (Figure 1.9d). Two DEP methods in particular have been proven effective in separating homogeneous cell samples: positive vs. negative DEP and DEP field flow fractionation (DEP-FFF). Separating cells using positive and negative DEP or, migration DEP, typically occurs in stagnant chambers and is especially effective at separating cells of different sizes and shapes. Blood cells have been successfully separated from whole blood samples and live bacterial cells from dead bacterial cells[8, 17] as well as cancerous cells from healthy red blood cells.[107-108] More sophisticated designs couple DEP with viscous drag to separate cells into different flow streams.[5-7, 10, 103, 109-110] An electric field inducing DEP forces is applied perpendicular to a fluid flow and particles are separated based on their dielectric properties and diffusion.[108] This technique has been used to successfully separate cancer cells from T-cells, and T-lymphocytes from B- lymphocytes.[5-7, 10, 109]
18
Figure 1.9. Examples of microfluidic cell and particle separation devices. (a) A multi-branched microfluidic device separated particles of different sizes and densities. Two fluids of different densities, one of which is carrying the particle sample, are introduced to the chamber, meeting at a pinch point. Once the particle sample reaches the expanded, branched section at the end of the pinch point, the particles are drawn to different branches based on buoyancy.[87] (b) Specifically functionalized microwells used to trap HeLa and pancreatic cancer cells. Once trapped, the cells reproduced. Each well can also be actuated and removed individually for further analysis.[100] (c) Magnetic fields can be used to separate cells and particles with differing magnetic affinities by varying the magnetic field across a microfluidic channel.[102] (d) Cells and charged particles are polarizable and able to be separated using castellated electrodes by dielectrophoresis.[16]
19
1.6 Electrochemical Analytical Methods
1.6.1. Impedance and Electrochemical Impedance Spectroscopy In DC circuits electron flow is unidirectional and the current, I, is characterized only by the potential, V, and resistance, R (Ohm’s Law Eq. 1) (Figure 1.10a).
(a) (b) V
Z(w) I
Figure 1.10. (a) Simple direct current (DC) circuit. (b) Simple alternating current (AC) circuit.
V (Eq. 1) I R
The oscillating AC current depends on the elements of a circuit that resist electron flow or hold electric charge. Most generally, this is expressed by the impedance, Z, the complex ratio of the voltage to the current in an AC circuit, similar to that of the resistance of the DC circuit. For a simple AC circuit (Figure 1.10b) in which the resistor and capacitor are connected in series, the current is similar to that of Ohm’s Law where impedance takes the place of the resistance term (Eq. 2). V I Z (Eq. 2)
For impedance measurements, alternating currents are typically sinusoidal. The impedance is therefore, dependent on the angular frequency while the resistance and capacitance are related by a phase angle, (Figure 1.11). To distinguish the phasors, the resistance is considered to be the real part of the impedance and the capacitance to be the imaginary part (Eq 3, 4 and 5). 20
Im
Z
Re
Figure 1.11. Phasor diagram for impedance. The real and imaginary impedance are related by the phase angle, .
(Eq. 3) Z ZRe jZ Im
where
ZRRe (Eq. 4) 1 ZIm (Eq. 5) C
The real and imaginary phasors are related by the tangent of the phase angle (Eq. 6). Impedance measurements are typically presented as Bode plots, the log Z() versus the phase angle and Nyquist plots, the ZRe versus ZIm, for increasing angular frequencies (Figure 1.12a,b). Z tan Im Z (Eq. 6) Re
21
(a) 7.00 0
-10 6.00 -20 5.00 -30
4.00 -40
3.00 -50
Log |Z| (Ohms) |Z| Log -60 2.00 (deg) Angle Phase -70 1.00 -80
0.00 -90
0.00 0.76 1.35 1.81 2.21 2.61 3.00 3.40 3.80 4.20 4.60 5.00 5.40 5.80 Log Frequency (Hz) Impedance Phase Angle (b) 7.00
6.00
5.00
4.00
(Ohms)
Im 3.00 -Z
2.00
1.00
0.00 0.00 1.00 2.00 3.00 4.00 5.00 6.00 ZRe (Ohms) Figure 1.12. Representative impedance spectra plots (a) Bode plot (b) Nyquist plot
For electrochemical impedance spectroscopy measurements performed in aqueous environments, the charges that form on the surface of the electrodes must be taken into account. Even when no current is flowing through the electrode a layer of ions and water molecules will form on the surface. This layer will be capacitive because of the distribution of charges over the electrode surface and is responsible for reducing the efficiencies of electrochemical reactions. Impedance measurements can be broken down into two types: faradaic and non-faradaic. Faradaic impedance refers to the resistance to charge transfer based on a redox electrochemical reaction. Non-faradaic impedance is the impedance due to 22
the capacitive layer on the electrode, the resistance of the solution and the capacitance of a solution’s components. For this study, the interaction from the faradaic impedance is reduced by limiting the electrolyte concentration and performing experiments at low potentials.[111]
1.6.2. Cyclic Voltammetry Cyclic voltammetry is an electrochemical analytical method that is used to study the environment at the metal – solution interface. Cyclic voltammetry studies not only reveal information on reactive species in solution, but respond to very small changes to the surface of metals.[112-128] Traditionally, these experiments are performed in large glass vessels, but as the technique has been refined to study surface interactions, the apparatus has been miniaturized to limit side reactions (Figure 1.13).[111, 122, 129-132] The design of the apparatuses used for these studies varies, but ideally have the same basic components: a working electrode, redox probe, reference electrode and counter electrode.
Figure 1.13. Experimental apparatus for studies involving thin-film metal electrodes and thin film coatings. (a) Top view. The Ag/AgCl reference electrode and the counter electrode (not pictured) are inserted at the top of the device into the electrochemical cell. The brass working electrode connection contacts the sample at the bottom of the electrochemical cell. (b) Side view. The electrochemical cell is the hollow tube at the center of the glass reservoir. The reference electrode and counter electrode do not contact the working electrode.
Cyclic voltammetry is widely used because analysis of the data is relatively simple.[133] These studies are typically performed in electrochemical cells that contain an aqueous solution of an electrolyte and a redox probe. The counter electrodes are made of 23
inert noble metals that do not participate in the electrochemical phenomena being studied at the working electrode.[111, 134] The potential is applied directly to the working electrode and the resulting current is measured with respect to the reference electrode. Most commonly, reference electrodes are saturated concentrations of an electrolytes paired with its elemental metal (eg. Ag/AgCl) of a known potential ranging from 0 V to 0.4 V.[111] The saturation of the reference electrode is important in that it must remain ideally non-polarizable so that only the reaction at the working electrode is apparent in the resulting voltammagram. The voltammagram follows the redox reaction of the redox probe, an electrochemically active species that freely accepts or donates electrons at predictable potentials depending on the polarization of the electrodes. Potential is applied to the working electrode and is swept positive to negative (forward scan) from an initial potential, Eini, and reversed (reverse scan) once a particular potential maxima, Emax, is reached. If the reference electrode and working electrode are both ideally nonpolarizable, then the voltammagram will represent the electrochemical reaction that occurs in the cell only. However, in surface reaction studies, the working electrode must be polarizable such that the voltammagram represents what is occurring on the surface of that electrode.[111] Figure 1.14 is an example of a typical voltammagram that follows the reduction and oxidation reaction of the redox probe where O is the oxidant, R is the reductanct, e- is the electron and n is the sociometric number to balance the electrons transferred (Eq. 7). (Eq. 7) During the forward, negative scan, the voltammagram represents the cathodic current at the working electrode or reduction of the oxidized form of the redox probe (O). As the potential decreases, a peak emerges (ic, Ec), indicating that the concentration of the oxidized form of the redox probe is becoming depleted. Once Emax has been reached; the reverse, positive scan demonstrates the electrochemical reaction from the reduced form of the redox probe (R) to the oxidized form. As the reaction progresses, the concentration of the reduced form of the redox probe continues to diminish until the anodic current peaks (ia, Ea), when the reduced form concentration is lowest. The anodic current then reduces until the scan returns to Eini. If the process is reversible, it will follow the Nernst Equation (Eq. 8): 24
[ ] (Eq. 8) [ ] where E is the potential of the cell which is dependent on the standard potential for the redox species reaction, Eo’, temperature, T, and the concentration of the redox probe species, [R] and [O].
Figure 1.14. A typical cyclic voltammagram. The forward scan represents the reduction of the redox probe. At Ec, the oxidized form of the redox probe is depleted and the reaction shifts to oxidation. As the reduced form of the redox probe is depleted, the current becomes more negative until Ea is reached and the reduced form of the max probe is depleted. E is typically chosen such that the potential at which Ec and Ea occur are within the bounds of the potential sweep.
Cyclic voltammetry experiments can be performed in a number of variations to collect information about the surface properties of metals, coatings and electrochemical reactions. If experiments are performed at varying scan rates, one can determine important information regarding a reaction’s kinetics.[133, 135] By carefully choosing an electrode metal for the working electrode, researchers have used cyclic voltammetry to study biological reactions and molecules.[123] Small changes in voltammagrams have also been used to determine small changes at the solution metal interface.[124-128] Because cyclic voltammetry outputs are so sensitive to surface changes, some groups have reported methods to determine the orientation of self-assembled mono-layers[115, 136-138] as well as the adhesion[12] and functionality[114] of molecules on the working electrode surface For this study, cyclic 25
voltammetry was used to determine the effectiveness of coatings to reduce electrochemical reactions on an electrode surface.
1.7 Microfluidic Methods for Particle and Cell Detection Microfluidics combined with detection has been in use for over 40 years with the advent of the strip or linear flow assay.[139] By drawing a sample by capillary forces through a strip of cotton or other non-woven fibrous materials through an area pretreated with specific antibodies, an anaylte is detected once it is immobilized on an area with concentrated tagged antigens. A line of bright blue or red indicates a positive result when compared to a control line. These strips are common and commercially available in home pregnancy and drug testing kits, but have also made major strides in detection of diseases such as STD’s[140- 141] and diphtheria[142-144] in third world countries. Though relatively inexpensive to fabricate, distribute and use, strip assay kits are prone to fouling in humid conditions and are not ultra- sensitive to dilute analytes. More recently, researchers have examined more sensitive optical and electrical detection methods in microfluidic devices for cell, particle and biological molecule detection. Droplet assays have been developed for the rapid and sensitive detection of ricin using similar antigen-antibody recognition, with smaller sample sizes than those used for strip assays[145] (Figure 1.15a).
1.7.1 Optical Detection Detection by light reflectance, refractance and adsorption is widely used in chemical and surface analysis and has recently been extended to cell study. Optical waveguides have been shown to be effective devices for detecting the presence and concentration of certain cell types.[146-148] This method uses a laser beam that is focused on an optical grating which directs the laser path into a detector located on the side of the waveguide substrate. When substances such as particles or cells are positioned over the grating, the change of refractive index at the surface of the grating is detected as an incidence angle shift or wavelength shift of the laser beam[146] (Figure 1.15d and e). Optical waveguide detection is most effective when the cells are focused and concentrated directly over the optical grating.
26
1.7.2. Electrical Detection Virtually all electric detection methods are derived from the electrochemical analysis methods mentioned in Section 1.6. Impedance and cyclic voltammetry are most commonly used as they are able to detect minute differences within tens of microns from the electrode or substrate surface.[27-28, 111] Cyclic voltammetry (CV) is used to measure the mobility of ions between a metal electrode and solution. The CV response is unique for each electrochemically active species in solution and for metal types. It is sensitive enough not only to detect cells, but also to determine small changes of the cell’s extracellular environment.[149-151] CV has also been used to determine minute changes in that of surface coatings, immobilized cells[152-153] or changes to surface-bound proteins.[154] Impedance has also proven effective as a surface-bound detection method for cells, particles and biological molecules. As a result, many techniques have been studied to bring a cell type of interest to sensing electrodes in impedance cell detection devices. Antibodies specific to certain cell type have been used to functionalize the surfaces between electrodes (Figure 1.15c). The desired cell is brought between the electrodes while other cells in the sample are washed from the detection surface. The immobilized cells are detected by a distinct impedance difference.[27, 29-31] In addition, studies have shown that it is possible to detect differences in cell secretions by impedance of immobilized cells on electrodes.[155] Though effective, the functionalized surfaces used to immobilize the cell of interest are subject to bio-fouling, have short shelf-lives, and are difficult to reuse. Functionalized particles have also been used to enhance impediometric detection of cells in suspension. Cells are brought between the electrodes by DEP and gold nanoparticles functionalized with the antibody of the cell of interest are introduced to the sample chamber. The particles bind specifically to the cells of interest, producing a distinctive electrical response.[4, 11, 156-157] Velev et al. used silver to further enhance the change in electrical response[4] (Figure 1.15b). As surface bound techniques are sensitive to biofouling, methods have been developed to detect freely suspended cells in suspension. Gupta et al. demonstrated that Immunoglobulin G (IgG) functionalized latex particles exhibited a distinctive impedance spectra when agglutinated with the IgG antigen.[14] However, detection time for this system 27
was as high as 70 minutes as the agglutinated particles needed to settle on the electrodes for detection. Later studies used functionalized magnetic particles to bring networks of bacterial cells and particles between electrodes.[15, 32, 158]
28
Figure 1.15. Examples of detection methods in microfluidic chambers. (a) Strip assays are the most common microfluidic detection systems, but are not as sensitive to dilute analytes. Rastogi et al. developed a droplet- based bioassay system immobilized using dielectrophoresis that can detect dilute samples of ricin.[145] Bringing cells and particles between electrodes increases the capacitance close to the electrodes and is detectable by measuring impedance. This has been done using (b) DEP and specifically functionalized particles and (c) a combination of DEP and specifically functionalized surfaces.[4, 31] (d) and (e) Optical waveguides use changes in the refractive index of the surface to detect cells and particles. A wavelength or incidence angle shift specific to particular particles and cell types indicates the presence of a desired analyte. These devices are most effective if the particle or cell sample is concentrated over the grating.[146-147] 29
1.8 Layout of Dissertation Many microfluidic techniques exist for the manipulation, collection, separation and detection of cells and particles, yet the focus of this dissertation is to develop and improve methods involving AC electric fields. AC electrical field methods for manipulation and detection show tremendous promise for the collection and detection of cells. These devices can be easily integrated with other microfluidic manipulation and separation methods, can usually be applied with little additional sample preparation and are typically simple to operate so long as adequate potentials and frequencies can be attained and maintained. At low voltages (on the order of mV), high frequencies (1 kHz to 1 MHz), and low electrolyte concentration, the impedance spectra can be easily measured without background from the solution or from interactions with the cell membrane. At high voltages, low frequencies and high electrolyte concentration, cell or fluid motion would be induced, making it possible to manipulate, collect or separate cells. In this study, two complementary phenomena are examined: developing an impediometric based biosensor and the collection and manipulation of cells by DEP and AC electroosmosis (Figure 1.16).
Figure 1.16. Capabilities of AC electric fields cells at low voltage / high frequency and high voltage / low frequency. In Chapter 2, we introduce a device that optimally uses both DEP and ACEO to direct cell and particle collection to a stagnant region on the chip. This device has the advantage 30
that the optically clear substrate is left unaltered making the technique amenable for coupling with evanescent waveguide sensing techniques. The method uses asymmetric AC field gradients to induce cell or fluid movement near the electrodes. Each electrode type is designed such that a stagnation point exists in the center of the electrode configuration. Cells are manipulated by taking advantage of cell and suspension media electrical properties. Several electrode configurations, media types, cell types and chamber conditions were investigated for cell concentration. While investigating the optimal particle collection parameters described in Chapter 2, we discovered that the functionality of metal electrodes in high electrolyte conditions and with cells was limited. When used with cells suspensions, the substrate and electrodes were subject to biofouling. Cells were fixed to the substrate surface and difficult to collect densely. Further, the metal electrodes tended to succumb to electrochemical reactions at the metal-solution interface in solutions with high electrolyte concentrations. In Chapter 3, we introduce coatings and coating techniques that reduce the propensity for electrochemical reactions and biofouling at the metal-solution interface. These coatings were used to create a barrier between the electrode and electrolyte solution to suppress lift-off of the metal electrodes while promoting the AC electrokinetic effects that drive particle and cell collection. The coated electrode composites were also used to successfully trap and collect cells and particles in tap water. In Chapter 4, we explore an impedance-based approach for the detection of whole cells in suspension. It has been shown that agglutinated systems of IgG functionalized latex particles in the presence of the IgG antigen exhibit a distinguishable difference in their impedance spectra to that of the non-agglutinated system.[14] Two different agglutination techniques were employed to determine the relative impedance difference amongst the systems: agglutination by altering the pH and agglutination by bio-specific recognition. For agglutination via pH, systems of yeast / latex particles and yeast / gold nanoparticles were tested at high pH (~7.8) and low pH (~3.1). A pH of 3.1 is below the isoelectric point of yeast cells. The surface charge of the yeast would be positive and the cell membrane will attract negatively charged particles. Hence, at a pH of 7.8, there would be no agglutination. For agglutination by bio-specific binding gold and magnetic particles were used 31
functionalized with a lectin, Concanavalin A (Con A). The impedance spectra from 1 Hz to 1 MHz of functionalized particles and yeast systems were compared to these of yeast and non-functionalized particles. The techniques introduced in this dissertation can be combined in new generations of AC electric field driven lab-on-a-chip devices that would be beneficial for the advancement of handheld point-of-care and environmental diagnostics.
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CHAPTER 2
On-Chip Collection of Particles and Cells by AC Electroosmotic Pumping and Dielectrophoresis Using Asymmetric Microelectrodes*
* Based on: Melvin, Moore, Gilchrist, Grego and Velev, Biomicrofluidics, 2011, 5, 47
2.1. Introduction Lab-on-a-chip or micro total analysis systems (TAS) offer great potential and numerous advantages over conventional benchtop techniques for biological analysis including reduced sample size, lowered possibility for contamination, decreased waste, and reduced size of the analytical equipment. Recently research has been focused not only on streamlining the steps of sample preparation and analysis, but on combining sample preparation and analysis in a single device. In doing so, TAS devices are, in many cases, more efficient and can achieve detection more rapidly than their macro-scale laboratory counterparts. A common complication with the use of TAS devices for biological detection in samples of less than 100 L is the need to pre-concentrate the analyte. The analyte in the samples, especially particulates of biological origin, may be very dilute, introducing the need to collect enough particles and direct them to a detection area. The typical microfluidic devices require cells or other analytes to be physically bound or brought close to the surface of a particular transducer to increase the accuracy of detection. Surface-bound microfluidic devices have been successfully coupled with chemical,[1-5] and optical[6-13] transducers to rapidly detect cells and particles in microliter suspensions. However, in order to achieve rapid detection, the time for establishing diffusion equilibrium of the analyte to the transducer should be kept to a minimum, requiring small fluidic chamber heights and, consequently, small sample volumes. The analyte should also bind to the transducer surface rapidly and with a high affinity. A large portion of methods for particulate collection is based on electric fields. Microfabricated electrodes have been extensively used to manipulate analytes using AC fields and can be easily integrated with microfludic devices. When an AC electric field is applied to a suspension of particles in an electrolyte solution, the particles and the medium are subjected to AC electric field-induced forces such as dielectrophoresis (DEP) and AC electroosmosis. DEP is the electrically induced movement of a particle based on its relative polarizability compared to the suspension media. When subjected to a non-uniform electric field, a particle becomes polarized and is drawn towards or away from the region penetrated 48
by the electric field. Particles that are more polarizable than the suspension media are drawn toward the region of the highest electric field (positive DEP). Conversely, if a particle is less polarizable than the media, it is repelled from the region of the strongest electric field (negative DEP). The polarized particles also commonly form chains in the direction of the electric field. The dielectrophoretic force is dependent on the intensity of the electric field (|E|), the radius of the particle (r) and the effective permittivity of the particle in the media [14] (p, m) (Eq. 1)
pm 2 F r3 Re E (1) DEP m 2 pm The AC fields also induce fluid flows in some electrode configurations by AC electroosmosis (ACEO). The motion is initiated by the additional ionic charges attracted by the field to the counterionic double layer.[14] When a non-uniform electric field is applied, fluid movement occurs tangential to the electric field (Et). The velocity of the fluid (ux) is dependent on the potential drop across the double layer (, the frequency of the electric field ( and the charge of the double layer (
1 1Et () ux Re (2) 2 The fluid velocity in AC electroosmosis is potential and frequency dependent. AC electroosmosis typically occurs between frequencies of 10 Hz and 10 kHz; fluid flow occurring at higher frequencies is typically due to electrothermal effects.[14] Asymmetric electrode patterns have been shown to produce flows that can be used for microfluidic pumping.[15-24] Electrodes can be arranged on a substrate surface in such a way as to induce asymmetric electric field gradients at the electrodes, thereby generating a local and bulk liquid flow. The asymmetric field creates a vortex near the edge of the electrode, driving a bulk flow analogous to a conveyor belt. At applied potentials from 0.1 - 10 V, planar asymmetric electrodes have been used to produce flows that are capable of transporting particles.[17-19,21-22,25] 49
In this chapter we introduce a TAS device using ACEO flow induced by co-planar asymmetric electrodes arranged in a “corral” configuration, exploited not only for pumping the liquid but for collecting particles in a target location on the device. Previously in our group, an ACEO corral device was used to collect particles and cells in the center of an energized electrode surrounded by a dielectric layer of an insulative photoresist.[26] The asymmetrical AC field induced small AC electroosmotic flows at each edge of the dielectric photoresist that drove a bulk fluid flow towards the center of the square corral. Densely packed particles were collected via a particle front originating at the photoresist edge and migrating over time inward towards the center. Particle collection in this device required over one hour, which may be too long for some practical applications. The present new configuration uses a quadrilateral arrangement of repeating asymmetrical electrodes surrounding a 1×1 mm collection region (Figure 2.1). All electrode segments surrounding the corral are simultaneously energized with AC fields, driving bulk fluid flow from each direction and collecting the particles within a centrally-located stagnant region. A key advantage of this configuration is that it uses coplanar electrodes surrounding the collection area as opposed to our earlier patterned sandwich configuration. The new configuration is compatible with evanescent wave sensing devices.[6-8,12,27] We report the characterization of this particle/cell collection device with regards to chamber height, applied potentials, frequency, electrode geometry, electrode and gap width, repetition of asymmetric electrode patterns and electrolyte concentration. We also introduce techniques combining ACEO with DEP to improve the performance of the device for collecting particles and cells from suspension.
50
Figure 2.1. (a) Top view of fully assembled planar electrode collection device. The square-wave AC current is applied to the magnified 5-repeat segment electrode design while the others are grounded. (b) Micrograph of 5- repeat segment electrode design. The collection area is 1 mm for each electrode repeat pattern (1, 3, 5, 10 repeats). (c) Enlargement of electrode configuration with electrode dimensions in microns. The repeat segments, starting at the outermost edge, include a thin electrode followed by a 5 m gap, and a larger 30 m electrode. Each segment is separated by a 15 m gap. (d) Enlarged side view of planar electrodes on glass substrate and induced ACEO flow. The ACEO flow over the asymmetric electrodes is in the form a large eddy induced just above the edge of the large electrode adjacent to the small gap (adapted from Ramos et al.[29]).
51
2.2. Materials and Methods 2.2.1. Sample Preparation Fluorescent 1 m sulfate latex particles (Molecular Probes, Eugene, OR) were washed three times by agitation and centrifugation with deionized (DI) water from a Millipore RiOs purification system to remove any impurities and surfactants. The particles were diluted to 0.1 wt% in DI water, 0.001 and 0.01 mM PBS, 10mM PBS or tap water. The conductivities of the DI water, PBS dilutions and tap water were: 0.25 S/cm (DI water), 8.0 S/cm (0.001 mM PBS), 36.0 S/cm (0.01 mM PBS), 17.0 mS/cm (10 mM PBS), and 235 S/cm (tap water). Yeast cells (Fleischmann’s Ankenny, IA) samples were hydrated in DI water and were also washed by agitation and centrifugation to remove any residual proteins and sugars dried with the yeast cells during the manufacturing process. Yeast cell suspensions were diluted to 0.1 wt% to compare with the latex particle dispersions.
2.2.2. Patterned Electrode Design The devices were designed to collect particles onto the sensing region of an evanescent optical waveguide sensor. The concentrator consists of a quadrilateral arrangement of repeating asymmetrical electrodes surrounding a collection region. A collection area as large as 1 mm × 1 mm was selected because this is the approximate size of the measurement area of a grating-based sensor using a collimated laser beam as previously reported by our group.[6] By using electrode gaps ranging from 5 to a few tenths of microns, the concentrator was designed to operate with low applied potentials, of the order of volts. A single metal layer process was chosen to simplify fabrication. The smaller electrode of the asymmetric pair was placed in the outermost position because ACEO fluid flow is directed from the smaller electrode towards the larger electrode.[15-17,29] Electrode corral configurations consisting of a different number (e.g 3×, 5×, 10×) of concentric repeats of the asymmetric electrode pairs were designed. In order to make the electrical connection, small openings were patterned to allow routing of metal traces of the opposite electrode. In the quadrilateral electrode arrangement the openings are in the corners while in circular electrodes they are in diametrically opposed positions. The overall length of the opening is 52
larger than functional electrode gaps to avoid undesired AC field effects, yet it is a small percentage of the overall electrode length so it does not impact performance.
2.2.3. Electrode Chip Fabrication Gold-on-glass planar electrode chips were designed based on literature on ACEO and DEP particle collection.[15-24,26,27]. The microelectrode devices were fabricated by RTI International using a single mask layer on 4”, 700 um thick borofloat glass wafers using a standard photolithographic lift-off process. Negative photoresist coated wafers (NFR016D2, 3.5 m thick, (JSR Micro, Inc. Sunnydale, CA) were UV exposed (Suss MicroTec MA8, Garching, Germany). Following a post-exposure bake, wafers were developed in PD523AD (JSR Micro, Inc. Sunnyvale, CA). Chromium (100 Ǻ) and gold (1000 Ǻ) were deposited by E-beam evaporation, followed by lift-off in N-Methyl-Pyrrolidinone (NMP) (JT Baker / Mallinckrod Chemicals, Phillipsburg, NJ). The wafers were cut to device size of 26 x 17 mm with a dicing saw. Each die consisted of 3-4 electrode corrals configurations with independently addressable contact pads and a common ground (Figure 2.1a).
2.2.4. Experimental Apparatus and Data Acquisition Electric wire leads were soldered to the electrode terminal pads of the devices. Fluidic chambers with depths of 250, 750 and 1600 m were constructed with a 250 m thick Grace Biolab Hybriwell chamber and a spacer of the designed thickness (Grace BioLabs Bend, OR) (Figure 2.1a). After loading approximately 80-250 L of sample and evacuating bubbles, 3M adhesive seals were placed over the loading ports to prevent leakage and evaporation. For each test, only one patterned electrode collection set was energized while the other patterns were grounded. A 33120A Agilent function generator (Agilent Inc., Santa Clara, CA) was connected to the chip through a 4.7 F capacitor (Radio Shack, Fort Worth, TX) to filter DC current component. The potential was monitored using a GDM-8034 GW digital multimeter (Instek Inc., Chico, CA). The square wave AC voltage was set to 0.05 - 6 V at frequencies ranging from 10 Hz to 1 MHz. 53
The particle concentration processes were observed with an Olympus BX-61 microscope (Tokyo, Japan) at 4×, 10× and 50× magnifications. Time lapse images were collected every 10 s for a total of 15 min per trial. Relative particle velocities and changes in concentration density were analyzed from the digital images by using Adobe Photoshop software. Relative particle velocities were determined by measuring particle displacement frame to frame and dividing this displacement by the time elapsed. For the particle concentration density, the color micrographs were adjusted to 8-bit grayscale image. The average grayscale value of a 140 square-pixel area (5 pixels × 28 pixels) was evaluated at 25 m increments starting from the electrode edge and ending at the collection area center. The analyzed region included particles and bare glass exclusively, eliminating electrodes, which appear black. The latex particles appear gray in the micrographs and, with increased particle density, the pixels darken and the grayscale values of the pixels decreased (255 is white). The gray value percent Deviation From White (%DFW) was calculated by averaging the grayscale pixel color number for each point along the width of the collection area and normalizing it against the grayscale value from same area in the image at time t=0.
2.3. Results The goal of this work is to collect the dispersed particles and concentrate them in a dense deposit within a central 1 mm2 collection region. This could allow effective coupling of such a particle collection device with an evanescent optical wave guide or other type of detection or analysis device. Further, any practical handheld device necessitates rapid collection (under 15 min) at low applied potentials and frequencies. For simplicity, we used a singular mask photolithography process to create the quadrilateral concentric electrode pair patterns without overlaps. The 1×1 mm dimensions for the collection area were chosen to correspond to the minimum size of a typical optical waveguide grating.[6] Each electrode pair is positioned such that the small electrode is the outermost and the large electrode is the innermost feature. When an AC field is induced over the asymmetric pairs, eddies form over the electrodes.[15,17,29-30] These eddies draw the fluid flow directionally from the small electrode towards the large electrode (Figure 2.1d). 54
The characterization of the planar electrode particle collection system was performed by systematically varying the system parameters. Overall, the highest concentration of particles in the collection area was observed when a denser line of particles, or particle front, was formed within the first few seconds of applying the AC current (Figure 2.2). The particle front traversed towards the center of the chip, similarly to the on-chip collection device by Bhatt et al.,[26] but at much faster rate. Not that the velocities reported here are not the bulk fluid flow velocities as reported in previous ACEO pumping literature[17-24], but are the rate of motion of the particle front. A major goal of this study was the determination of design and operation parameters affecting particle collection efficiency. Rapid detection is required in most lab-on-a-chip devices, thus we established a target collection time of < 15 min. Our strategy for collecting particles was to establish an electrode configuration and operation that induces a bulk, unidirectional ACEO flow toward the center of the corral, then to tune other device design parameters to increase particle collection efficiency. The AC potential was varied from 50 mV to 6 V, while the frequency was varied from 10 Hz to 1 MHz. We discuss the parameters that effect ACEO bulk fluid flow generation in DI water (electrode and gap width) and 1 m latex particle collection rate (chamber height, applied potential / frequency, electrode repeat segments, electrode geometry, and toggling between optimal DEP and ACEO settings). Lastly, we discuss the effect of changing analyte composition and its relation to the design parameters. 55
FIigure 2.2. Micrographs of typical particle collection results for 300 Hz, 800 mV at (a) 0 s, (b) 10 s, (c) 180 s, and (d) 360 s. A line of grouped particles, or particle front, is formed within the first few seconds of applying the AC voltage. The particle front traverses from the edge of the electrodes towards the center of the chip over the course of the trial, forming a rectangular deposit of densely packed particles in the middle of the collection area. The collection patterns are similar to those attained by Bhatt et al. [26], but thanks to the different principle are attained in 10% of the time.
2.3.1. Bulk Fluid Flow Generation Electrode and Gap Width. Electrode asymmetry, imposed by varying electrode width and spacing or introducing roughness to the surface, is what drives ACEO flow in one direction just above the electrode surface.[15,28-29] The spacing between electrodes as well as the electrode width is crucial to drive bulk, unidirectional ACEO pumping and to induce positive and negative DEP. The size of these eddies above the electrodes is dependent on the AC field gradient and is therefore affected by the size of the electrode and the distance between them.[29-30] Previous ACEO pumping devices using asymmetric electrodes have identified 300 Hz, 800 mV as favorable parameters for inducing bulk ACEO flow.[17-18] Thus, the ACEO fluid flows and their efficiency for particle collection for three electrode pattern designs were first evaluated at 300 Hz, 800 mV (Table 2.1).
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Table 2.1: Electrode Design Configurations and Dimensions
(in m) Small Small Large Large Configuration Electrode Gap Electrode Gap 1 5 5 30 15 2 5 25 30 75 3 5 5 50 15
For each design, some degree of fluid flow was apparent. When an AC was applied to configuration 1, a particle front quickly formed at the edge of the innermost large electrode and proceeded towards the center of the array for approximately 6 min. The particles collected at the center in a well packed structure as desired (Figure 2.2d). For configuration 2, both the large and small gap sizes were increased 5-fold. Rapid particle movement between the electrodes was observed, indicating intensive fluid flow above them. However, no particle front formed at the electrode edge and no particle collection ensued. Similar results were obtained by increasing the width of the large electrode by 66% in configuration 3. Observations by reflectance and fluorescence microscopy revealed the particles were being drawn towards and trapped over the large electrodes for both designs 2 and 3. Typically, particle trapping and chaining in the gap between the electrodes is evidence of dielectrophoresis,[31-33] however, at this electrode length scale and for 1 m latex in DI water, 300 Hz, 800 mV was not sufficient to induce DEP.[14] The reason for the collection on the electrodes was that the eddies formed over the electrodes for configuration 2 and 3 were not large enough to drive a bulk fluid flow, so the particles were trapped above the electrodes only. Similar results have been reported using devices with large circular electrode pads over which cells and particles were trapped.[7-11,25] The results are consistent with ACEO pumping models predicting that the optimal pumping widths to be 1.5×, 6.55× and 4.74× for the small electrode, large electrode and large gap respectively,[34] so configuration 1, with feature dimensions of 1×, 6×, and 3×, is most effective in generating strong unidirectional fluid velocity.[17,19,23,31,34-35]
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2.3.2. Particle Collection Tuning Parameters Once unidirectional ACEO flow is induced by the asymmetric electrodes, the flow pattern must be maintained and directed to a stagnant region over the central collection area. The ACEO flow must also be strong enough to drag the suspended cells or particles from the bulk to the stagnant region. The optimal conditions for the process were established by studying the effect of chamber height, the AC field potential and frequency, number of electrode repeats, electrode geometry and electrode pattern inversion on particle collection to determine the optimal settings for ACEO flow and DEP particle collection. These conditions were then used to demonstrate a combined ACEO / DEP collection technique.
Chamber Height. It has been reported in the ACEO pumping literature that chamber height has no significant effect on bulk fluid flow or particle motion.[34] While this might be true for delivering a particle from one point to another in a rectangular microfluidic channel, through this study, it was discovered that chamber height is critical when collecting particles in a central stagnant area. We investigated three chamber heights, 250 m, 750 m and 1600 m (Figure 2.3) using the 5 electrode repeat configuration. For the 250 m chamber height, at each investigated AC frequency and potential, eddies were observed to form along the periphery of the collection area at the leading edge of the innermost large electrode (Figure 2.3a). These eddies appeared to have no predictable shape or behavior and to be most turbulent at the corners of the collection area. The particle motion increased as the potential was increased as predicted by previous ACEO asymmetrical electrode pumping studies.[17- 22,23] However, no combination of low potential (250 mV – 1 V) or frequency (10 Hz – 5 kHz) produced a unidirectional bulk fluid flow that would achieve the desired densely packed formation of particles in the center of the chip.
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Figure 2.3. Effect of chamber height on particle collection. (a) – (c) depict the top view in real time of particle collection device. The width of the collection area is 1mm. (a) At a relatively low height of 250 m, small, rapid eddies form directly over the electrodes and erratic eddies exist along the edge of the chip. Little, if any, particle collection was observed at the center of the collection area. (b) The optimal chamber height is approximately at a 1:1 ratio to the width of the collection area on the chip (750 m). This thickness leads to strong vortices and a well-defined stagnant area at the center of the chip. (c) Higher chamber height allows for larger vortices to form over the electrodes, however, these vortices are strong enough to draw the particles away from the electrode chip surface. (d)-(f) show the side view of the COMSOL model output for geometry and dimensions analogous to corresponding particle collection experiments and present the corresponding fluid dynamics simulations.
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For the 1600 m chamber height, a particle front initially formed at the electrode edge within seconds of applying the electric field. The distinct front traversed inward for approximately 1 min but the particles appeared to be drawn away from the center and upwards into the chamber. The particles collected at the center were not densely packed (Figure 2.3b). The 750 m chamber appeared to perform best at all combinations of potentials and frequencies. A particle front clearly formed within a few seconds and traveled towards the center of the collection area (Figure 2.3c). This result was unexpected as previous ACEO pumping literature reported optimum chamber heights for maximum bulk fluid flow to be approximately half of the length of one electrode repeat segment;[15,30,34] therefore we expected the 250 m thick chamber to give us the optimal results for particle collection and bulk fluid flow. We interpreted these results by simulations using a 2-D fluid dynamics model developed by COMSOL finite element analysis software package. The model examined a side-view cross section of the electrode chip and chamber, bisected across the width of two sides of the equilateral electrode segments. The total electrode segment width and chamber height were scaled with respect to the center collection area and to the dimensions tested experimentally. No-slip boundary conditions were applied at the top of the chamber, the collection area in between the electrode segments, and beyond the outer edge of the electrodes on the glass surface. A linear velocity scaled to the linear bulk fluid flow velocity observed experimentally was programmed along the surface of the electrode segments. The relative velocities are depicted by the color schematic in Figure 2.3d-f with red being the region of highest velocity. White arrows represent velocity gradient direction. For all chamber heights, bulk circular flow patterns formed originating at the electrode segments and flowing towards the center of the chip. The no-slip boundary conditions at the center of the chip, chamber top, and at the outer edge of the electrode area, give rise to the overall circular motion and the formation of updrafts and downdrafts. Figures 2.3d, 2.3e, and 2.3f represent the results for 250 m, 750 m and 1600 m chamber heights respectively. In Figure 2.3d, at the 1:0.25 collection area to chamber height scale, the bulk circular fluid flow develops just over the electrode segments. Although there 60
is a distinct stagnant area at the middle, the circular fluid flow does not reach far enough to push particles into the center of the collection area. The particles are most likely drawn into the bulk circular flow and settle along the edges of the small bulk eddies. For chamber heights at a 1:2 scale, such as those featured in Figure 2.3f, the linear bulk fluid flow is less restricted by the no-slip boundary conditions of the chamber top forming a bulk circular flow pattern with larger diameter. There is a distinct, concentrated stagnant area in the center of the corral and a tangential component of the velocity beyond the electrode segments, which may be responsible for the observed partial collection of particles. Just above the stagnant region in the center of the collection area, however, there is a strong velocity gradient in the y-direction where the bulk circular flow patterns meet. This updraft is likely responsible for the observed partial sweeping of particles away from the collection area. The 1:1 chamber model shown in Figure 2.3e also has a distinct stagnant area at the center of the collection area as in the 1:0.25 and 1:2 models. The bulk circular flow patterns originating from the electrodes do not meet in the center to form a distinct updraft and a strong tangential velocity is present beyond the edge of the electrode segments. The lack of updraft and the strong linear velocity beyond the electrode segment edge are likely responsible for the initial pulse of particles being moved to the center. Without a central updraft, the particles are pushed laterally over the collection area, where they settle and assemble in densely packed formation. This is in agreement with the experimental observation that the 750 m chamber height offers optimal particle collection. This height was used through the remainder of the study.
Potential and Frequency. The maximum ACEO pumping velocity requires an optimal potential and frequency.[17-22,24] First, we determined the effect of the potential on our ACEO flow velocity and particle collection efficiency, by varying the applied potential from 50 mV to 800 mV while holding the frequency constant at 300 Hz (the lowest frequency at which ACEO flow was consistently noted). We discuss the effect of increasing frequency later in this section. Particle motion observations were initiated at an applied potential of 50 mV and increased at 50 mV increments until moving particles were observed just above the 61
electrodes. Time lapse images of the process were analyzed to calculate particle front velocity and particle collection efficiency. The resulting particle front velocity is graphed against the position on the corral in Figure 2.4. The particle front velocity is maximal at the electrode edge and decreases as the particle front traverses from the electrode edge towards the center of the collection area. Particle movement was detected at potentials as low as 250 mV, however, the slow front velocity during those trials did not afford sufficient time for the particles to reach the center within 15 min. As the potential increased, the particle front velocity increased as expected. At potentials above 400 mV, the particle front reached the center within the 15 min trial, with the particles reaching the center the fastest (within 6 min) at 800 mV.
8
250mV 7 300mV
) 400mV 500mV 6
m/s 600mV
( 700mV 5 800mV
4
3
2
Particle Front Velocity 1
0 0 100 200 300 400 500 Chip Position (m)
Figure 2.4. Particle front velocity at 300 Hz as particles suspended in DI water travel from the electrode edge to the center of the collection area. The electrode edge is at 0, and the center of the chip is at 500 m. For lower potentials (250 – 300 mV), the bulk fluid velocity is insufficient to drag the particles to the center of the collection area within the 15 min trials. As potential increases, the bulk fluid velocity increases, drawing the particles to the center.
The pixel density along the width of the collection area was calculated for applied potentials of 250 mV – 800 mV at 280 s, the determined optimum collection time from the particle velocity analysis. Figure 2.5 presents a plot of the normalized Deviation From White (%DFW, see Methods) with respect to collected front position on the chip at 280 s. A 62
negative percentage indicates fewer particles than initially present (e.g. they become swept away by the flow), while a positive percentage indicates particles have accumulated. At 300 mV, just above the collection potential threshold, particle movement and some degree of collection is apparent with a negative %DFW (Figure 2.5). The range of %DFW, however, is between -30 and 30%, indicating very slow particle movement. For 400 mV to 600 mV, a distinct particle front could be observed, and resulted in a much higher %DFW range of - 100% to 75%. The particles are drawn from the electrode edge toward the center but at 280 s they were maximally concentrated at 350 m from the edge, short of the center of the collection region. At 700 mV, we observed more densely packed particles and thereby obtained a positive %DFW value in the middle of the collection area. The particles at the perimeter of the 1 mm2 corral at t=0 reached the stagnant region within 220 s. At longer times, more particles outside the bounds of the electrode corral are drawn into the center. The 60% DFW peak at 500 m results from the initial collection of particles while the second peak at 250 m represents the particles drawn from outside the corral after the electrodes are energized. The outermost particles get collected in the corral center in 280 s. At 800 mV the %DFW exhibited the largest observed variance, from -75% to 125% and a maximum % DFW was measured at the collection area center. Potentials above 1 V did not appear to collect particles as densely as 800 mV. The higher flow rate and resulting collection pattern resembles that of the 1600 m chamber height (Figure 2.3) as an upward directed flow drives particles away from the collection area. Therefore, 800 mV was selected as the base potential for subsequent parameter optimization. An increase in the applied AC potential results in an increase in the magnitude of the electric field, thereby increasing ACEO flow rate.[17-22,24] However, its effect on the particle collection process is much more complex, given the interplay of fluid flow, viscous drag, velocity gradients and particle sedimentation. Thus, it is possible to tune the particle collection front velocity, the density of the collection and the collection time by altering the applied potential; while the most efficient process does not necessarily occur at the fastest ACEO flow rate. 63
Figure 2.5. Pixel density analysis of ACEO particle collection in DI water at 300 Hz, (a) 300 mV, (b) 400 – 600 mV and (c) 700 – 800 mV after 280 s of field application in terms of percent deviation from the initial average pixel color value. The solid points are experimental data while the dashed line is its trend. (a) At 300 mV, very little particle movement and a small, slow particle front was observed, resulting in little change in the particle density throughout the collection area. (b) From 400 mV to 600 mV, a faster moving, more distinct, particle front was observed, but it was not rapid enough to reach the collection area center. (c) At 700 mV to 800 mV the efficient collection of particles at the center was confirmed by a positive %DFW. 64
The particle movement induced by frequencies of 300 Hz, 1 kHz and 2.5 kHz was measured at potentials ranging from 250 mV to 800 mV. At frequencies approaching zero the capacitance of the electric double layer impedes the ionic flows necessary to induce ACEO flow. As the frequency increases, the electron flow through the circuit happens rapidly enough to overcome the electric double layer capacitance and induce flow.[14] For our device, ACEO flow appeared to begin above 100 Hz and dissipated after 2.5 kHz, similarly to the [17] data of Brown et al. with latex suspensions in NaNO3 solutions. The particle front velocity for each frequency increased as expected when the potential was increased. Little difference in the particle front velocity was detectable between 300 and 1 kHz. However, when the frequency was increased above 1 kHz, fewer particles were collected in the center of the corral than at lower frequencies (Figure 2.6). Inspection of the electrode gaps under higher magnification revealed that the particles were being drawn between the smaller gaps in gold electrodes and forming chains by DEP. Particle chaining occurred more rapidly at frequencies above 2.5 kHz. DEP became the predominant effect at frequencies above 10 kHz. DEP trapping and chaining between the electrode pairs impeded ACEO particle collection in the center. However, we were able to combine ACEO and DEP in a synergistic manner as described in the section Toggling between DEP and ACEO Regimes below.
65
300 Hz 7 1000 Hz 2500 Hz
) 6
m/s ( 5
4
3
2
1 ParticleFront Velocity 0
200 300 400 500 600 700 800 Potential (mV)
Figure 2.6. Particle front velocity analysis in the low-frequency regime at 800 mV. DI water suspensions require low potentials and frequencies to induce ACEO bulk fluid flow. Very little change in the particle front velocity is observed for small frequency changes between 300 Hz and 1 kHz. As the frequency was raised above 2.5 kHz, DEP was observed, reducing the amount of particles in the collection area and reducing the apparent particle front velocity.
Number of Electrode Repeats. The particle collection performance of the electrode chip design was further characterized by changing the number of asymmetric electrode pairs surrounding the collection corral. According to the model proposed by Ramos et al.,[29] a large vortex forms over the large electrode at the edge adjacent to the small gap drawing the fluid towards and across the electrode. The two vortices formed over the small electrode are rapid, but are in opposite directions and have little influence over the bulk fluid flow velocity. For chamber heights greater than the width of the electrode repeat segments, the fluid will flow in the direction from the small electrode to the large one.[29,34] Therefore, each electrode pair generates one large vortex that drives the overall fluid flow. As the number of electrode repeats increases, the number of driving vortices and the linear velocity above the electrodes also increases. We observed the particle collection processes at 300 Hz, 800 mV for electrode chips containing 1, 3, 5, and 10 electrode repeats over 15 mins. As expected, devices with 1 and 3 electrode repeats exhibited slower particle collection than chips with 5 electrode pairs. Neither chip design induced a particle front capable of drawing particles to 66
the center of the collection area. One electrode repeat is sufficient to draw the particles just beyond the large electrode as indicated by the %DFW, which becomes slightly positive at the electrode edge, but remains unchanged towards the center of the collection area. In the 3 electrode pair repeat device, a faint particle front formed initially and slowly traversed approximately 250 m towards the center of the collection area (Figure 2.7d). The 10 electrode repeat device behaves like that of the 5 repeat device; upon field application, a dark, distinct particle front was formed immediately upon application of the AC field and was drawn towards the center of the chip. After approximately 100 s, the center of the chip began to lighten, indicating that particles were being drawn away from the collection area. At 280 s, a -60 %DFW at the chip center indicated fewer particles are in the center of the collection area than initially present. Thus, the 10 repeat does not effectively collect 1 um particles under these conditions, though it might be useful for larger particles or higher electrolyte concentrations where a greater velocity is required.
Table 2.2: Relative Particle Density for Electrode Repeat Designs at Various Chip Positions
Chip Position (m) Number of Repeats 0 100 200 300 400 500 1 ++ - - - - - 3 - - + 0 - - 5 -- -- - 0 ++ ++ 10 + + 0 + -- --
(++ =greatest increase of particle density, + = slight increase in particle density, 0 = no particle density change, - = slight decrease in particle density, -- = greatest decrease in particle density) 67
Figure 2.7. Pixel density analysis of observed ACEO phenomena in DI water at 300 Hz, 800 mV for 1, 3, 5, and 10 electrode repeats. The bulk fluid flow velocity driven by on electrode repeat will collect cells just beyond the large electrode, but is not sufficient to draw particles to the collection area center. Three repeats will induce flows that draw the particles only 250 m towards the collection area. Although 10 electrode repeats does act much like the 5 repeat device at first, after 100 s, the particles appear to be lifted from the corral center surface.
A fluid dynamics model of the electrode repeat microfluidic chambers was developed in COMSOL to interpret these results. No-slip boundary conditions in the x-direction were implemented at the chamber top and to either side of the electrode sets. A linear velocity was programmed just above the electrodes for a length proportional to that of the total width of the electrode pair segments (Figure 2.7a,c,e). For the 1-repeat system, a small eddy develops just over the electrodes, but the flow in the x-direction is dissipated just beyond the electrode edge. The eddies over the electrodes are more developed in the 3- and 5-repeat designs. In the simulation for the 10-repeat design, the x-direction fluid velocity is much higher than that of the 5-repeat design, but a strong y-direction flow in the middle of the chip where the eddies meet is likely responsible for drawing particles into the bulk fluid flow eddy.
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Figure 2.8. COMSOL finite element analysis models and micrographs taken at 280 s for increasing number of electrode pattern repeats: one (a,b), three (c,d), and ten (e,f). The bulk fluid flow eddies formed as a result of ACEO increase in size and velocity as the number of repeats increases. None of the repeat designs resulted in densely packed collection of particles as in the five repeat design (Figure 2.3b,e and Table 2.2).
The simulations are in good correlation to the experimental data. Similarly to the dependence on applied potential, the optimal particle collection occurs for a linear velocity over the electrode segments creating circular bulk eddies that do not meet above the collection center. Overall, the number of electrode repeats and the applied potential both play a major role in inducing linear flow in the devices. Although the characterization was only performed for 1 m particles, the linear velocity can be adjusted to optimize collection of larger and smaller particles and cells.
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Electrode Geometry. In this cycle of experiments we compared the efficiencies of the chip design with rectangular corral with chips having circular electrode pattern. The diameter of the circular collection area was made large enough to fit the 1×1 mm square detection area in the center (d = 1.414 mm). Circular electrode pairs with 1, 3, 5, and 10 repeats were energized by an AC voltage of 800 mV, 300Hz. The ACEO flow on each of the circular repeat pairs acted similarly to their rectangular counterparts. In the 1-repeat circular pattern device, small eddy lines formed along the electrode edge similar to those observed on the rectangular design, however no localized areas of particle collection developed near the electrode edge and the % DFW varied very little. With 3-repeat segments, a distinct particle front was formed at the electrode edge that propagated towards the center of the chip over 15 mins. This particle front did not reach the center, exhibited nearly identical behavior to that of the rectangular design and did not disappear until approximately 600 s. The 10-repeat design immediately produced a distinct particle front, visibly drawing particles towards the center. However, after 120 s, the particles appeared to lift off the surface of the collection area, as seen with the rectangular electrodes (Figure 2.5). The % DFW at 280 s indicates a small band of particle deposits beginning 150 m from the electrodes. The 5-repeat design induced the optimal x-direction fluid velocity to draw the most particles to the center of the collection area (Figure 2.8c). The %DFW range is significantly smaller than that of the rectangular design, however, the area under the 5-repeat circle curve is 1.5 times larger than that of the 5-repeat rectangular curve (Figure 2.8a). The collection patterns result from the flow pattern and dynamics established by the electrode “pumps”. In the rectangular design, the quadrilateral unidirectional flows impinge along the diagonals of the collection area, resulting in a discernable diamond-like shape in the particle front and collection patterns. The flow generated in a radial pattern over the circular electrode chip uniformly converges at the center of the collection area. It may appear that the circular electrode pattern is the better candidate for optimal particle collection, but each design has specific application-dependent advantages. The rectangular design is optimal for collecting particles in a concise, small area of the collection center while the circular design is better for collecting a larger number of particles in a uniform deposit over the entire corral surface. 70
Figure 2.9. Micrographs and pixel analysis for circle geometry electrode segment repeat trials at 300 Hz, 800 mV. The %DFW analysis indicates similar results to the rectangular electrode geometry in that an optimum x- direction fluid velocity is needed to form a densely packed bed of particles in the center of the collection area. The 5-repeat design appeared to induce the optimum fluid velocity to draw 1 m particles to the collection center. Despite the lower %DFW range than in the rectangular case, (a) integrating the area under the %DFW curves for both the rectangular and circular geometries indicates the circular geometry is capable of drawing a larger amount of particles into the collection area. This is further seen by comparing the micrographs of the (b) 5-repeat rectangular design and (c) 5-repeat circular design results at 280 s.
Electrode Pattern Inversion. Some particle and cell collection studies have had some success with collecting cells with devices that induced a fluid flow in the opposite x-direction of the collection area. A COMSOL fluid dynamics model of a scaled 5-electrode repeat segment with the x-direction fluid velocity induced away from the center shows that a large bulk fluid flow forms much like that of the model with the fluid flow induced towards the collection center. A stagnant region exists at the center of the collection area. Any particles that are in or near the vortices formed over the electrodes could be drawn to and settle in the stagnant region between the electrodes. 71
When the 300Hz, 800 mV electric field was applied to the inverted, 5-repeat segment electrodes, the particles were visibly drawn away from the center of the collection area. At approximately 120 seconds, the center of the collection area started to lighten. Over the course of the 15 min trial, the lightened area of the center shaped to be similar to that of the collected particles in the baseline electrode configuration. As expected, the particles formed a halo around the outside of the electrode repeats while on the inside of the collection area the particles collected about the base of the bulk fluid eddies formed above the electrode repeats.
Synergistically Combining DEP and ACEO Regimes. After observing DEP effects in DI water suspensions of latex particles at frequencies above 1 kHz and optimal ACEO particle collection at 800 mV and 300 Hz, we evaluated the joint use of both DEP and ACEO effects to improve the ACEO-driven particle collection observed at 800 mV & 300 Hz. The goal was to use DEP for rapid trapping of the particles in the gaps and then push the collected particles in the corral by ACEO. The potential was fixed at 800 mV while the frequency was toggled between 300 and 5,000 Hz at varying intervals of time. For each trial, the particles were collected by DEP at 5 kHz for a period of time ranging from 5 - 40 s, followed by ACEO driven collection in the center of the chip for an additional 5 - 40 s. The cycle was repeated for a total of 3.5 min. Toggling intervals of 5 s were neither adequate to collect particles by DEP nor long enough to bring a significant amount of particles to the center of the chip. Increasing the DEP collection time to 30 s appeared to be sufficient to draw particles between the electrodes in the small gaps. For the ACEO interval, it was observed that 10 s fluid motion was sufficient to induce particle motion to the center of the corral. Alternating 30 s DEP intervals and 10 s ACEO intervals resulted in drastically decreased collection times and an increase in the amount of particles collected as compared to ACEO alone.
2.3.3. Effects of Medium Composition and Particle Type Electrolyte Concentration. Many point-of-care and cell detection devices operate with media of high electrolyte concentrations. In AC electrohydrodynamics, ions in electrolyte solutions reduce double layer thickness and therefore suppress ACEO flow.[14] We investigated the ACEO particle collection in dilute PBS (phosphate buffered saline) for 72
applied potentials of 800 mV – 2 V and frequencies ranging from 100 Hz to 1 MHz. For concentrations above 0.01 mM PBS, no particle movement was detected at all frequencies and potentials below 1.2 V, the thermodynamic threshold for electrochemical reactions.[36] Above 1.2 V and below frequencies of 100 kHz, bubbles formed on the electrode surface in evidence of undesired faradic reactions at the gold electrodes. Particle motion was observed below 1.2 V for electrolyte suspensions concentrations below 0.01 mM PBS, however, the resulting particle front velocity decreased sharply and particles did not reach the center of the collection area (Figure 2.9). As with the DI water suspensions, increasing frequency only slightly increased the particle front velocity. To determine if the particle collection device could be used for unmodified environmental water analysis, latex particles were suspended in tap water (conductivity 235 mS/m). The collection pattern observed at 900 mV and 1 kHz was similar to that observed in 0.001 mM PBS. The decrease in pumping velocity in high ionic strength media and undesired faradic reactions have been widely reported in ACEO pumping devices.[17,19-20,22] Our device can be used as designed presently to collect cells and particles in low electrolyte solutions (conductivities below 36.0 S/cm), but work is in progress to develop electrode surface modifications that could suppress or eliminate the faradic reactions that occur in higher conductivity solutions.
Figure 2.10. Particle front velocity for latex particles suspended in DI water and 0.001 mM PBS at 500 Hz, 1 V. 73
Particle Size and Type. We verified the ability of the device to collect larger latex particles (5 m) and yeast cells in DI water suspensions. For the 5-repeat electrode device, the applied 800 mV, 300 Hz AC current did induce an ACEO flow, but the velocity was not great enough to carry the particles to the collection area. Increasing the potential to 2 V and the number of repeats from 5 to 10 induced a flow strong enough to carry the larger particles to the collection area. Likewise, the yeast cells having a diameter of approximately 5.0 m also required 2 V applied potentials and 10 electrode repeats to induce flows great enough to collect the cells in the center of the device. The yeast cells also were prone to adhering to the glass substrate surface outside the electrodes, in between the electrodes and in the collection area. Overall, the device can reliably collect particles within broad ranges of sizes and types.
2.4. Conclusions We report the principle of operation and in-depth characterization of a co-planar asymmetric electrode device that induces a simultaneous, quadrilateral unidirectional flow, which is capable of rapid, dense particle collection. The asymmetrical planar electrode design collects the particles suspended in low-electrolyte water over a 1×1 mm collection area at relatively low applied potentials and frequencies. This configuration provides effective operation by using electrodes surrounding the collection area without obscuring the optical sensor that may be paired with the devices. This feature makes possible the design of collection devices that may be compatible with evanescent wave sensing, grating couplers, ring resonators, and interferometers This device is flexible in that design parameters affecting collection time, full flow development, and the x-direction velocity over the electrodes can be tuned so as to be applied to varying suspensions of both particles and cells. The chamber height was observed to affect the flow pattern over the electrodes and into the collection area, and a 1:1 ratio of chamber height to collection area were found to be optimal. Both the applied potential and the number of electrode repeat segments can be increased and decreased to tune collection parameters for varying suspensions of particles and cells. The geometry of the electrode system does not alter the induction of ACEO flow, but can be altered to achieve more uniform deposition of the particles in the center of the device. These tunable parameters and 74
the photolithographic fabrication of the chip enable it to be easily coupled with surface- bound optical detection methods. The development of such devices needs to overcome one additional stage. High electrolyte suspensions continue to be problematic. As most biological suspension use salt buffer solutions or are saline in nature it is essential to develop collection devices operating in such media. It has been suggested that modifying the electrodes may reduce or eliminate the propensity for the electrochemical reactions to occur[37-38]. For now, we have shown that surface-pattern directed ACEO flow is useful for collecting analytes in tap water suspensions, leading us to believe the device can be used for environmental water contamination applications.
2.5. Acknowledgements This research was supported by RTI International, Research Triangle Park, NC and a NIRT grant from the US National Science Foundation (CBET-0609087).
2.6 References
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[13] Felten, M., Staroske, W., Jaeger, M. S., Schwille, P., & Duschl, C. (2008). “Accumulation and filtering of nanoparticles in microchannels using electrohydrodynamically induced vortical flows,” Electrophoresis. 29(14), 2987- 2996.
[14] Morgan, H., Green, Nicholas G. (2003). AC Electrokinetics: colloids and nanoparticles (1st ed.). Williston, VT USA: Research Studies Press LTD.
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[16] Gonzalez, A., Ramos, A., Green, N. G., Castellanos, A., & Morgan, H. (2000). “Fluid flow induced by nonuniform ac electric fields in electrolytes on microelectrodes. II. A linear double-layer analysis,” Physical Review E. 61(4), 4019-4028.
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[17] Brown, A. B. D., Smith, C. G., & Rennie, A. R. (2001). “Pumping of water with ac electric fields applied to asymmetric pairs of microelectrodes,” Physical Review E. 63(1). 016305-1 - 016305-8.
[18] Studer, V., Pepin, A., Chen, Y., & Ajdari, A. (2002). “Fabrication of microfluidic devices for AC electrokinetic fluid pumping,” Microelectronic Engineering. 61-2, 915-920.
[19] Studer, V., Pepin, A., Chen, Y., & Ajdari, A. (2004). “An integrated AC electrokinetic pump in a microfluidic loop for fast and tunable flow control,” Analyst. 129(10), 944-949.
[20] Urbanski, J. P., Levitan, J. A., Burch, D. N., Thorsen, T., & Bazant, M. Z. (2007). “The effect of step height on the performance of three-dimensional ac electro-osmotic microfluidic pumps,” Journal of Colloid and Interface Science. 309(2), 332-341.
[21] Hilber, W., Weiss, B., Mikolasek, M., Holly, R., Hingerl, K., & Jakoby, B. (2007). “Particle manipulation using 3D ac electro-osmotic micropumps,” Paper presented at the 18th European Workshop on Micromechanics (MME 07), Guimaraes, PORTUGAL.
[22] Bligh, M., Stanley, K. G., Hubbard, T., & Kujath, M. (2008). “Two-phase interdigitated microelectrode arrays for electrokinetic transport of microparticles,” Journal of Micromechanics and Microengineering. 18(5). 055007-1 - 055007-9.
[23] Mpholo, M., Smith, C. G., & Brown, A. B. D. (2003). “Low voltage plug flow pumping using anisotropic electrode arrays,” Sensors and Actuators B-Chemical. 92(3), 262-268.
[24] Lian, M., & Wu, J. (2009). “Ultrafast micropumping by biased alternating current electrokinetics,” Applied Physics Letters. 94(6). 064101-1 - 064101-3.
[25] Wong, P. K., Wang, T. H., Deval, J. H., & Ho, C. M. (2004). “Electrokinetics in micro devices for biotechnology applications,” IEEE-ASME Transactions on Mechatronics. 9(2), 366-376.
[26] Bhatt, K. H., Grego, S., & Velev, O. D. (2005). “An AC electrokinetic technique for collection and concentration of particles and cells on patterned electrodes,” Langmuir. 21(14), 6603-6612.
[27] Grego, S., Huffman, A., Lueck, M., Stoner, B. R., & Lannon, J. (2010). “Nanoimprint lithography fabrication of waveguide-integrated optical gratings with inexpensive stamps.” Microelectronic Engineering. 87(10), 1846-1851.
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[28] Bazant, M. Z., & Ben, Y. X. (2006). “Theoretical prediction of fast 3D AC electro- osmotic pumps,” Lab on a Chip. 6(11), 1455-1461.
[29] Ramos, A., Gonzalez, A., Castellanos, A., Green, N. G., & Morgan, H. (2003). “Pumping of liquids with ac voltages applied to asymmetric pairs of microelectrodes,” Physical Review E. 67(5), 056302-1 - 056302-11.
[30] Garcia-Sanchez, P., Ramos, A., Green, N. G., & Morgan, H. (2006). “Experiments on AC electrokinetic pumping of liquids using arrays of microelectrodes,” IEEE Transactions on Dielectrics and Electrical Insulation. 13(3), 670-677.
[31] Velev, O. D., Gupta, S., Alargova, R. G., & Kilpatrick, P. K. (2006). “Dielectrophoretic assembly of bionanocomposite materials from live cells and functionalized particles.” Abstracts of Papers of the American Chemical Society. 231, 505-COLL.
[32] Gupta, S., Alargova, R. G., Kilpatrick, P. K., & Velev, O. D. (2008). “On-chip electric field driven assembly of biocomposites from live cells and functionalized particles,” Soft Matter. 4(4), 726-730.
[33] Gupta, S., Alargova, R. G., Kilpatrick, P. K., & Velev, O. D. (2010). “On-Chip Dielectrophoretic Coassembly of Live Cells and Particles into Responsive Biomaterials,” Langmuir. 26(5), 3441-3452.
[34] Olesen, L. H., Bruus, H., & Ajdari, A. (2006). “ac electrokinetic micropumps: The effect of geometrical confinement, Faradaic current injection, and nonlinear surface capacitance,” Physical Review E. 73(5), 056313-1 - 056313-16.
[35] Debesset, S., Hayden, C. J., Dalton, C., Eijkel, J. C. T., & Manz, A. (2004). “An AC electroosmotic micropump for circular chromatographic applications,” Lab on a Chip. 4(4), 396-400.
[36] Bard, A. J., Faulkner, Larry R. (2001). Electrochemical Methods: Fundamentals and Applications. Hoboken, NJ USA: John Wiley & Sons, Inc.
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CHAPTER 3
Optimizing Coating and Composition of Metal Electrodes for AC Field Manipulation and Collection of Cells and Particles in Electrolytic Suspensions*
* Based on: Melvin, Moore, Grego and Velev, in preparation
79
3.1. Introduction Lab-on-a-chip microfluidic devices offer advantages for cell detection such as reduced detection time, sample size, sample contamination and waste. Many microfluidic manipulation and separation techniques exist, but electrically driven devices are a focus of much cutting-edge research. The miniaturization of electronic devices and improvements in microelectrode fabrication make these devices simple to create and use. These devices are appealing also because the particles, cells and fluids can be manipulated in their native state or with little pretreatment. Electric field methods are also highly selective and take advantage of unique properties of cells and particles. Asymmetric electrode devices that produce flows suitable for microfluidic pumping as well as cell sorting, collection and detection have recently been developed.[1-28] These alternating current (AC) electrokinetic devices use AC electroosmosis (ACEO) and dielectrophoresis (DEP) or a combination of the two to manipulate and trap cells and particles. DEP is the electrically induced movement of a particle based on its relative polarizability compared to the suspension media. When subjected to a non-uniform electric field, particles that are more polarizable than the suspension media are drawn toward the region of the highest electric field (positive DEP). Conversely, if a particle is less polarizable than the media, it is repelled from the region of the highest electric field (negative DEP). The polarized particles also commonly form chains in the direction of the electric field. The dielectrophoretic force is dependent on the intensity of the electric field, the radius of the particle and the effective permittivity of the particle in the media.[29] The AC fields also induce fluid flows in some electrode configurations by AC electroosmosis (ACEO). The motion is initiated by the ionic charges attracted by the field to the counterionic double layer.[29] When a non-uniform electric field is applied, fluid movement occurs tangential to the electric field. The velocity of the fluid is dependent on the potential drop across the double layer, the frequency of the electric field and the charge of the double layer. The fluid velocity in AC electroosmosis is potential- and frequency- dependent. AC electroosmosis typically occurs between frequencies of 10 Hz and 10 kHz; fluid flow occurring at higher frequencies is typically due to electrothermal effects.[29] 80
Electrodes can be arranged on a substrate surface in such a way as to induce asymmetric electric field gradients at the electrodes, thereby generating a local and bulk liquid flow. At applied potentials from 0.1 - 10 V, planar asymmetric electrodes have been used to produce flows that are capable of transporting particles.[2,9-13] Although the electric double layer is needed to evoke ACEO pumping and DEP, it also is responsible for promoting electrochemical reactions at the electrode surface with catastrophic effects on the electrodes. Two types of electrochemical processes occur at the metal-solution interface, non-faradaic and faradaic processes. A non-faradaic process refers to the adsorption or desorption of ions from the solution onto the metal surface and, therefore, is corollary to the formation of the double layer. The formation of the double layer impedes faradaic reactions in that it is capacitive and holds a certain amount of charge. In order for a faradaic reaction to occur, the electric double layer must reach a particular potential threshold before ions pass the metal-solution interface. The time that it takes to reach this threshold is called the charging time, which is dependent on the applied potential. In experiments, it would be on the order of microseconds. Faradaic reaction is the transfer of ions across the metal-solution interface in accordance with Faraday’s law. Faradaic reactions are limited by the availability of reactants, mass transfer of the ions from the bulk to the interface and the adsorption / desorption of ions at the interface.[30-31] Faradaic reactions have thermodynamic potential thresholds that are specific to each reactant chemical species. Interrupting any of the faradaic reaction precursors would increase these thresholds and prevent the faradaic reaction from occurring within a range of certain applied potentials.[30-31] Pumping highly electrolytic solutions by electric means poses challenges in that electrochemical reactions are prone to occur at the metal-ionic solution interface at relatively low potentials. These potentials tend to be lower than the ones that are necessary to induce ACEO and to manipulate particles in the micron length scale.[2,9-12,14-15,32-33] Increasing electrolyte concentrations also reduces the propensity for DEP effects when the ions in solution saturate the particle surface to such a degree as to reduce polarization of the particle, thus supressing one electrokinetic phenomena useful for trapping particles.[29] It is well known that using metal electrodes in high electrolyte concentration suspensions results in electrode degradation due to electrochemical reactions at the electrode surface.[12,33] 81
Although much work has been done to pump fluids in low electrolyte concentration solutions, few devices have been developed to successfully pump fluids or manipulate cells or particles by electric means in high electrolyte solutions. Suppressing or eliminating electrochemical reactions has been of interest particularly for corrosion applications. Fisher investigated and defined four main modes of electrochemical reaction and corrosion inhibition including: interface inhibition, electrolyte layer inhibition, membrane inhibition and passivation.[34] Electrolyte layer and passivation inhibition refer to the addition of certain chemicals to the electrolyte solution that adjacent to the electrode surface, inhibiting the formation of the electric double layer. A chemical gradient barrier is used in some applications to inhibit electrochemically active species from reaching the metal-solution interface. However, the chemicals typically used in such methods, like chromates[35], may not be biocompatible and therefore not useful for cell detection and manipulation applications. Interface and membrane inhibition involve using self-assembled monolayers (SAMs) or thin layers of molecules as a physical barrier between the electrode metal and the electrolyte solution. [34] As long as these barriers form relatively strong bonds with the metal, organic monolayers and thin films have been shown to be effective electrochemical reaction inhibitors.[35-37] Noble metals are typically chosen for microelectrode manufacturing. Although relatively inert, gold and platinum have been involved in electrochemical reactions in electrolyte solutions even at neutral pH.[12,33,38] Using both interface inhibition and membrane inhibition, layered polyelectrolytes[39], alkanethiols[37,40-45], thiolated benzene[46], polyanaline[47-49], silica[50] and silanes[37,40,44] have been successful over gold[39,43-46,48-50] and platinum[47-48] at reducing electrochemical reactions at the metal-solution interface. It has been suggested that using SAM or polymer coatings over electrode surfaces could reduce electrochemical reactions without impeding electrokinetic phenomena[33,51-52], but little work has been performed to date. The goal of this study was to develop coatings and coating techniques to reduce noble metal electrode damage at the potentials required to drive ACEO flow and DEP in high electrolyte conditions. We also sought to reduce non-specific binding of the cells and particles to the substrate surface as seen in previous work.[33] Using the device platform 82
introduced in Chapter 2, we investigated methods to prevent the electrochemical reaction that occurs in high electrolyte conditions at the metal-solution interface. In this paper we report coating techniques and potential sweep techniques that maximize particle collection while minimizing the propensity of electrode degradation using gold and platinum microelectrodes and coatings of several types such as: hydrogels, silanes and layer-by-layer (LbL) polyelectrolytes. We also demonstrate the devices’ ability to collect cells and particles up to 5 m in diameter.
3.2. Materials and Methods 3.2.1. Particle and Cell Preparation Fluorescent 1 m and 5 m sulfate latex particles (Molecular Probes, Eugene, OR) were washed three times by agitation and centrifugation with deionized (DI) water from a Millipore RiOs purification system to remove any impurities and surfactants. The 1 m particles were diluted to 0.1 wt% in each of the following media: DI water, TE buffer diluted 0.001, 0.01, 0.1 and 1.0, 1mM PBS and 10 mM PBS. The 5 m particles were diluted to ~0.05 wt% in DI water and tap water. Yeast cells (Fleischmann’s Ankenny, IA) samples were hydrated in DI water and were also washed by agitation and centrifugation to remove any residual proteins and sugars dried with the yeast cells during the manufacturing process. Yeast cell suspensions were diluted to 0.05 wt% to compare with the 5 m latex particle dispersions. The conductivities of the DI water, TE buffer and PBS dilutions, and tap water were as follows: 0.25 S/cm (DI water), 12.2 S/cm (0.001 TE), 19.6 S/cm (0.01 TE), 117.0 S/cm (0.1 TE), 953.7 S/cm (1 TE), 2.16 mS/cm (1mM PBS), 17.4 mS/cm (10mM PBS) and 189.9 S/cm (tap water).
3.2.2. Electrode Chip Fabrication Gold-on-glass and platinum-on-glass planar electrode chips were designed based on literature on ACEO and DEP particle collection.[1-2,7-12,14-15,32,53] The photolithography techniques to fabricate the electrode chips are discussed elsewhere.[33] Each die consisted of 83
3-4 electrode corrals configurations with independently addressable contact pads and a common ground. Wires were soldered to each contact pad to facilitate connection to circuit.
3.2.3. Electrode Coating Electrode Chip Preparation. Before coating, each electrode chip was washed in Nochromix, rinsed with DI water and allowed to dry in a 70°C for an hour. The coatings were applied to both platinum and gold electrode chips illustrated in the schematic in Figure 3.1.
Polyacrylamide Coating. Polyacrylamide coated electrode chips were prepared by Research Triangle Institue in Durham, NC using a 15 mm 15 mm 10 m micromold fabricated by photolithography techniques. The acrylamide solution was prepared as reported elsewhere[54] and was pipetted into the mold and pressed against the electrode surface with 1 N of force. The coating was polymerized by UV radiation for 15 s, washed with DI water and dried with nitrogen.
Silanes. To strengthen the bond between the metal electrodes to a layer of silane covering the whole chip surface, a thiolated silane was used.[55-60] A 0.02 M solution of 3- mercaptopropyltrimethoxysilane (MPTMS) (Sigma Aldrich. St. Louis, MO) in ethanol was stirred continuously by magnetic stir bar in a vessel. The upwards facing electrode chips were submerged in the solution for two hours. The chips were rinsed with ethanol and allowed to dry overnight in a covered Petri dish. A 1:100 dilution of AquaSil (Pierce Biotechnology, Rockford, IL) in of DI water was agitated in a well-ventilated hood. The pH was brought to 5.0 using HCl (Sigma Aldrich. St. Louis, MO). An MPTMS coated electrode chip was immersed face-up in the solution for 10- 15s, rinsed with methanol and cured in a 100°C oven for 1-2 hours.
Agarose and Agarose-Doped with Polyelectrolytes. Agarose was dissolved in DI water at 110°C to make a 10 wt% gel. Immediately before application, the agarose was melted in a microwave, layered on silane coated electrode chips[58-59,61-62] by spin coating and allowed to 84
cure in a covered Petri dish overnight. The thickness was checked using a Veeco D-150 Profilometer (Plainview, NY) and found to be approximately 5-9 m. For the polyelectrolyte-doped agarose chips, the agarose was prepared as previously mentioned, but poly(4-styrenesulfonic acid) ammonium salt (PSS) or poly(diallyl dimethyl ammonium) chloride (PDAC) were added at a concentration of 10 wt %. Agarose, PSS and PDAC were all purchased from Sigma Aldrich (St. Louis, MO)
Layer-by-Layer Polyelectrolytes. Washed and dried electrode chips were immersed in a 1 mM 11-mercatoundecanoic acid (Sigma Aldrich. St. Louis, MO) and ethanol solution for 2 hours. The coated chips were rinsed with ethanol and allowed to dry overnight. PSS and PDAC were diluted in separate Petri dishes to 3 mg/mL. The pH of the PSS solution was adjusted to a pH 2 with 1M HCl and PDAC to a pH of 7 with 1M NaOH. Five layers of PSS and PDAC were fabricated on the chip using a layer-by-layer (LbL) technique by alternately immersing the electrode chip face up in each polyelectrolyte solution for 20 mins. In between each experiment, the chip was rinsed with DI water with sonication and dried with nitrogen. After achieving a five layer LbL polyelectrolyte coating, the chips were allowed to cure in covered Petri dish overnight. For the Aquasil terminated coating, a LbL polyelectrolyte coating was made as stated above, with the exception of the final layer. Polyacrylic acid (PAA) (Sigma Aldrich. St. Louis, MO) was used as a negative electrolyte so as to bind to the vinyl groups on the Aquasil silane. PAA was diluted to 3 mg/mL in DI water and adjusted to pH 7 using NaOH. After being allowed to dry overnight, the electrode chip was treated as mentioned above with Aquasil, rinsed with methanol and dried in a 100°C oven 85
Figure 3.1. (a) Top view of planar electrode collection chip. The collection area is 1 mm for each electrode repeat patterns (1, 3, 5, 10 repeats). The square-wave AC current is applied to one repeat segment while the others are grounded. (b-f) Side view enlargement of electrode chip with coating schematic for each type of thin film coating. Each coating includes a thiolated crosslinker to strengthen the bond of the polymer to the metal electrode. (b) The AquaSil silane coating is crosslinked with MPTES. (c) The LbL polyelectrolyte coating is crosslinked using MUA and comprised of 5 layers of alternating PSS and PDAC. (d) To ensure that the agarose is spread to a thin film during the spin coating process, the chips were first coated with MPTES and AquaSil. (e- f) Polyelectrolyte agarose coatings were doped with PSS and PDAC respectively before being layered over the silane. Not featured: polyacrylamide coatings. 86
3.2.4. Experimental Apparatus and Data Acquisition Electric wire leads were soldered to the electrode terminal pads of the devices. Fluidic chambers with depths of 750 and 1600 m were constructed with a 250 m thick Grace Biolab Hybriwell chamber and a spacer of the designed thickness (Grace BioLabs Bend, OR). After loading approximately 120-200 L of sample and evacuating bubbles, the loading ports were covered with 3M adhesive seals were placed over to prevent leakage and evaporation. For each test, only one patterned electrode collection set was energized while the other patterns were grounded. A 33120A Agilent Function Generator (Agilent Inc., Santa Clara, CA) was connected to the chip through a 1-4.7 F capacitor (Radio Shack, Fort Worth, TX) to filter DC current. The potential was monitored using a GDM-8034 GW Digital Multimeter (Instek Inc., Chico, CA). The square wave AC voltage was set to 0.01 - 6 V at frequencies ranging from 10 Hz to 100 kHz. The particle distribution was observed with an Olympus BX-61 microscope (Tokyo, Japan) at 4×, 10× and 50× magnification. Time lapse images were collected every 10 seconds for a total of 15 minutes per experiment. Relative particle velocities and changes in concentration density were analyzed from the digital images by using Adobe Photoshop software. The color micrographs were adjusted to 8-bit grayscale image. The average grayscale value of a 140 square-pixel area (5 pixels × 28 pixels) was evaluated at 25 m increments starting from the electrode edge and ending at the collection area center. The analyzed region included particles and bare glass exclusively, eliminating electrodes which appear black. The latex particles appeared gray in the micrographs and, with increased particle density, the pixels darken and the grayscale values of the pixels decreased (255 is white). The gray value percent Deviation From White (%DFW) was calculated by averaging the grayscale pixel color number for each point along the width of the collection area and normalizing it against the grayscale value from same area in the image at time t=0. Bare and coated chips were analyzed using cyclic voltammetry with a VersaSTAT 4 potentiostat (Princeton Applied Research. Oak Ridge, TN) to determine the electrochemical activity at the electrode surface. Glass microscope slides were coated with a 1000 Å of gold 87
using a metal evaporator (Cooke Vacuum Products. South Norwalk, CO) and diced into 1 cm 1 cm squares. The gold slides were used as a working electrode, a platinum wire (Alfa Aesar, Ward Hill, MA) was used as the counter electrode and a micro Ag/AgCl electrode (Microlectrodes, Inc. Bedford, NH) was used as the reference electrode. Measurements were performed in a solution of 1 mM PBS and 5 mM K4Fe(CN)6 (Sigma Aldrich. St. Louis, MO) with a potential sweep from -0.8 V to 1 V and a scan rate of 0.1 V/s.
3.3. Results and Discussion It has been previously reported that using metal microelectrodes in electrolyte suspensions to manipulate particles and cells is problematic.[12,33,53-52] Electrochemical reactions are prone to occur in high electrolyte suspensions at potentials above 1.2 V.[30] Means of electrochemical reaction suppression at the electrode surface were investigated here by coating the electrodes with polymers and self-assembled monolayers (SAMs). These layers would ideally not interfere with electric double layer formation nor would they prevent AC electrokinetic phenomena. Additionally, the coatings were also used to suppress non- specific binding of the cells to the electrode and substrate surfaces. While varying electrolyte concentrations for each coating type and electrode metal, the robustness of the device’s ability to promote AC electrokinetic phenomena for particle collection and electrochemical reaction suppression was determined. In investigating the electrokinetic phenomena during potential and frequency sweeps, the ideal frequency for ACEO and DEP effects occured below 100 kHz. Above 100 kHz for electrolyte suspensions of less than 19.6 S/cm, particles tended to coalesce and stick to the substrate surface in the small electrode gaps due to electrothermal effects.[29] Therefore, the study was narrowed to frequencies between 100 Hz and 100 kHz. Electrolyte-containing suspensions with conductivities below 12 S/cm did not cause electrode degradation. Additionally, ACEO was rapid enough for particle collection and DEP occurred at potentials below 1.2 V. However, as the electrolyte concentration is increased above a conductivity of 12 S/cm, bare electrodes succumb to degradation. 88
Several cues that electrode damage was eminent during the frequency / potential became apparent over the course of the study. To preserve the limited number of chips available, as these cues arose, the corresponding trials were cut short to reduce or eliminate electrode damage. During optical microscopy observations, electrode damage was evident by the gray areas in Figures 3.2b and 3.2d. For each potential setting, the frequency was set to its highest value (100 kHz) and slowly dropped in decade increments. At potentials below 3V, as the frequency approached the point where electrode damage was observed, the ACEO flow would change direction. Above 3V, electrode damage was typically preceded by the particles vibrating just above the electrode surface. Vibration was also noted at higher frequencies and is likely a result of the transition between DEP and AC electrothermal effects.[29] The last and usually most drastic cue for electrode damage was bubble formation, indicating that a gas-releasing electrochemical reaction at the surface of the electrode is occurring. Bubble formation was also indication of electrode metal lift off, destroying the circuit and rendering the chip useless. While remaining cognizant of these cues, the stability of platinum and gold electrode metals was investigated. The applied potential and frequency limits for each metal type and coating type in high electrolyte conditions were determined while carefully watching for indication of electrode degradation.
3.3.1. Effect of Type of Electrode Metal Past investigations were done solely using gold electrodes[33], but other metals have shown promise in ACEO pumping such as platinum[12] and titanium.[10] For the present study, gold and platinum microelectrodes were chosen for their high conductivity and relatively inert behavior in high electrolyte solutions.[9,63] The electrode metal had no effect on the AC electrokinetic phenomena in deionized water. As seen in Figure 3.3 and 3.4, in DI water from 800 mV to 5 V both Au and Pt exhibited ACEO at frequencies below 3 kHz, a region of DEP and ACEO from 3kHz to 30 kHz and DEP at low potentials (800 mV – 1V) above 30 kHz. As electrolyte was added, the Au and Pt electrodes varied in the degree to which they degraded. Platinum exhibited electrode damage by a brown or purple discoloration, usually occurring on the large electrodes at the junction of the small electrodes in the bottom-right corner or at the ground lead (Figure 3.2a). Gold electrode damage was 89
best identified by lift-off of the thin gold film from the substrate, leaving behind the silhouette of the chromium under-layer. The damage typically occurs on the small electrodes or at the ends of the large electrodes (Figure 3.2c).
Figure 3.2. Micrographs of uncoated gold and platinum electrodes before and after electrode damage. (a) Undamaged gold electrodes. (b) Gold electrodes damaged at 3V, in 0.001 TE electrolyte, evident by the light areas along the length of the electrode. Some of the damage on the small electrode is enough to render the device useless. (c) Undamaged platinum electrodes. (d) Platinum electrodes damaged in 10 mM PBS at 800 mV, evident by the dark purple and brown areas of the large electrode. For a reference to scale, large electrode is 30 m wide.
Neither of the uncoated metal types were able to withstand the potentials and frequencies necessary to collect 1 m particles in electrolyte concentrations over 0.1 TE. From an electrode damage standpoint however, the platinum electrodes appeared to be more robust than the gold electrodes. Uncoated platinum electrodes were able to function in electrolyte concentrations up to 1.0 TE (Figure 3.4), whereas uncoated gold electrodes were only useful in electrolyte concentrations up to 0.1 TE before bubbles appeared at the surface (Figure 3.3). The gold electrodes also exhibited reverse ACEO behavior in 0.001 TE at 3 V below 1 kHz indicating that an electrochemical reaction was eminent. Platinum electrodes did not show any indication of reverse direction ACEO. Bubbling did not occur on the Pt electrodes until 5 V 500 Hz, but formed at potentials as low as 4 V on Au electrodes. 90
Faradaic reactions initiation is based on the formation of electric double layer and the charge build-up on the electrode surface. When the metal ions and solution ions are in contact with each other, a faradaic reaction will occur at a particular potential dependent on the species involved in the reaction. The rate of the faradaic reaction is based on the capacitance and thickness of the electric double layer, as well as the propensity of the ions to leave their respective layers (solution vs. metal or substrate). As Pt and Au are relatively inert noble metals and are commonly used in electrode-based applications, but have been shown to undergo faradaic reactions in high-electrolyte environments. There are two widely accepted theories as to how this might occur; the classical electrochemistry chemisorption model and incipient hydrous oxide adatom mediator (IHOAM) model. In the chemisorption model, it is assumed that the faradaic reaction occurs so long as the proper reactants from the solution are sufficiently adsorbed on the metal surface.[30] Gold has been commonly chosen for electrodes in AC electrokinetic pumps as a poor chemisorber that would not form an oxide layer.[2,9,11,15,17,32] Platinum was also successfully used in AC electrokinetic pumps at lower potentials and at lower electrolyte concentrations[12]. However, all studies performed reported some electrode damage at higher potentials and higher electrolyte concentrations.[2,9,33,63] Using the chemisorber model as a basis, platinum has been found to be less selective to electrochemically active ions than gold in molecular biosensor applications.[64] Platinum, therefore, has a tendency to form its own protective coating by adsorbing a variety of ions at the metal-solution interface.[65-67] Gold, conversely, has a strong affinity and selectivity to chloride ions which may explain the accelerated faradaic reactions in these experiments.[65] The IHOAM model suggests hydroxide intermediates are formed as a result of available metal ions at the metal-solution interface. Imperfections of the polycrystalline structure of the metal allow for localized reactions between the free metal ions and OH- available in the aqueous solution. The metal-OH- intermediate acts as a catalyst for faradaic reactions at the metal-solution interface, allowing for the reactions to occur at potentials lower than expected. This model has historically been used to explain anomalies in unexpected faradaic reactions between metals and solutions in acidic and basic conditions[38,68-71], but has been recently used in neutral-pH studies.[23] Burke et al. demonstrated gold the formation of an AuOH 91
intermediate that facilitate faradaic reactions at the electrode surface in neutral pH solutions[23], similar to that in slightly basic and acidic conditions. Platinum showed similar activity in acids and bases in forming an intermediate, but did so at higher potentials than gold.[70] Although, little work has been done on platinum directly in neutral solutions, if the trend observed in acidic and basic solutions can be extended to neutral solutions, the IHOAM model could explain the superior performance of platinum in high electrolyte conditions. Regardless of which theory is used to explain the electrochemical reaction phenomena, it is apparent that physically impeding electrolyte ions from reaching the metal-solution interface with coatings should reduce the likelihood of electrode damage.
Figure 3.3. Electrokinetic phenomena diagram for gold electrodes. Experiments for uncoated gold in 1 TE and 1 mM PBS as well as silane coated gold in 1 mM PBS were not performed as electrodes were prone to rapid damage. 92
Figure 3.4. Electrokinetic phenomena diagram for platinum electrodes. Experiments for uncoated platinum in 1 TE and 1 mM were not performed as electrode damage was eminent.
3.3.2. Hydrogel-Based Coatings Hydrogel coatings should allow for free ion movement necessary for ACEO and DEP to occur yet provide a barrier to suppress the electrochemical reaction between the metal electrodes and the electrolyte suspension. Polyacrylamide gels are commonly used in gel electrophoresis for the separation of proteins and have been shown to enhance DEP effects in high electrolyte conditions.[54,72-74] The polyacrylamide coatings for this study were formed by micromolding over gold electrodes[54] and were relatively thick, 2 – 10 m as compared to the other coatings. Although others reported success with polyacrylamide coated platinum electrodes[72-74], we did not observe any ACEO movement or DEP collection at any of the potentials or frequencies. In DI water above 1.2V, electrode damage was immediate and appeared to be accelerated as compared to uncoated gold electrodes. 93
Agarose is commonly used as a hydrogel scaffold for biological cell growth and is used to immobilize cells and was chosen for this study for its biocompatibility with a wide range of cell types. As no agarose thin film coating procedure readily exists, a design of experiment (DOE) was performed to determine the effect of surface silanization, preheating the chip, agarose volume and spin coater speed on coating thickness. ANOVA analysis of the data showed that pretreating the chip with a silane monolayer, using 750 L of agarose, and a spin speed of 3000 rpm yielded most consistently the thinnest coatings of 4-8 m. Preheating the chip showed little improvement. Although agarose coated electrodes exhibited ACEO flow, it generally occurred at higher potentials than those required for uncoated electrodes. For all trials, no movement or local ACEO eddies were observed at potentials below 2V. In DI H2O, for both Pt and Au agarose-coated electrodes, little fluid movement was observed at any frequency until 2 V. Above 2V, ACEO flow was in the reverse direction, no particle collection was noted, while electrode damage was observed in DI water trials. AC electrokinetic phenomena were evident at slightly higher electrolyte concentrations as well. At 0.001TE, both agarose- coated metal electrode types exhibited ACEO flow optimal for particle collection at 3 V, but electrode damage occurred at 3 V. The electrode damage was so devastating to the agarose- coated gold electrodes at 0.001 TE that no trials were attempted for higher electrolyte concentrations however; the platinum coated electrodes were robust enough to withstand potential sweeps up to 1.0 TE. For 0.01 and 0.1 TE, particle collection ACEO was observed at 800 mV and 2V respectively, but electrode damage was noted at potentials as low as 3V for both concentrations. At 1.0 TE, no movement was apparent until 3V, but damage to the electrodes and local ACEO were observed at 50 kHz. Overall, agarose coated electrodes showed little improvement over uncoated electrodes and were exceptionally difficult to clean. Particle sticking to the coating was problematic at potentials as low as 800 mV in at electrolyte trials below 0.01 TE. The particles adhered to the substrate so strongly that the chip was stripped of the agarose coating, washed in Nochromix and completely recoated between trials. 94
To determine if residual electrolytes were being leached from the agarose coating into the DI water, we tested the conductivity of DI water used to swell an agarose gel. An agarose gel was swelled in DI water at the same ratio as present in the fully assembled particle pre-concentrator. The conductivity of the solution was found to be 14.16 S cm, on the same order of magnitude as 0.001 TE, not nearly high enough to explain the electrode damage at low potentials. The added silane layer used to facilitate the agarose thin film also appeared to do little to protect either electrodes metal from electrochemical attack.
PDAC doped agarose coatings appeared to be promising in DI H2O. Only localized ACEO was observed below 3V, but particles were collected at 3V. The chip was able to withstand potentials up to 5V without any damage. As the electrolyte concentration was increased, the performance of the chip degraded and was only tested to 0.01 TE. Collection was noted at 3V in the 1 – 3 kHz range, but the electrode damage for the polyelectrolyte doped agarose left entire sections of large electrodes damaged by electrochemical reactions. ACEO occurs as a result of the formation and mobility of the electric double layer. The potential drop across the electric double layer is responsible for the surface charge while the potential at the edge of the diffuse layer is responsible for the tangential portion of the electric field. The charging time, , the time for the ions to form the electric double layer, is proportional to frequency (f) ( ~ 1/f ) therefore, at lower frequencies the charging time increases. Increasing the thickness of the double layer or impeding the formation of the Stern layer results in increased charging time and impedes double layer formation thereby dampening of ACEO effects, especially in the the 300 Hz to 10 kHz range where ACEO typically occurs. The device engineered by Krishnan et al. relied on DEP for particle collection alone.[54,72-74] DEP forces will be induced so long as the electric field generated is sufficiently strong and should not be diminished if the electric double layer formation is suppressed. The polyacrylamide coating was likely too thick to allow for proper electrode polarization and double layer formation for ACEO to occur at low potentials. The agarose coating was much thinner than that of the polyacrylamide coatings and exhibited ACEO phenomena, but still thick enough to require much higher potentials than bare electrodes to induce ACEO flow with a velocity sufficient to collect particles. 95
The increased propensity of electrochemical reactions at the surface of the electrode was most likely due to the porosity of the hydrogel coatings. Both polyacrylamide and agarose promote biological cell growth and storage as the porous gels store vital nutrients and moisture. The porosity of the hydrogels is likely too large for electrochemical reaction membrane inhibition applications as the pore size is large enough to allow solvated electrochemically active ions access to the metal-solution interface. The hydrogels might also act as a medium to hold these ions to the metal-solution interface, increasing the likelihood of electrochemical reactions. The agarose gel lowered the pH of DI water from 7.0 to 6.0. The slightly acidic environment might also contribute to the electrochemical reaction at the metal-solution interface.[68,69-70] A CV analysis was also performed on the agarose- based hydrogel coatings, but since thiolated silanes were used to crosslink the agarose to the gold working electrode, the results resembled those of silane coated electrode (Figure 3.5).
3.3.3. Silane and LbL Polyelectrolyte Coatings Silane and LbL polyelectrolyte coatings were formed over gold and platinum electrodes by dip coating. During the dip coating process, the surfaces and immersion fluids are kept clean and free of debris to form uniform barriers over the electrodes. Thiolated crosslinkers, mercaptopropyltriethoxysilane and mercaptoundecanoic acid, were used to strengthen the bond between the coatings and the metal surface and prevent lift-off.[55-56] The thickness of both coatings was on the order of tens of nanometers. Silanes are commonly used to prevent biofouling and have been used to protect electrodes in bio-detection applications.[75] Thiolated self-assembled monolayers have also been shown to be good insulators on gold electrodes.[43,45] Because of their ionic conductance, polyelectrolytes have been used to enhance the performance of ionic and micro-electronic devices in electronic- based microfluidic applications.[76-81] Because the coatings were thin and, in the polyelectrolyte’s case, possibly conductive, they should act as a barrier to prevent the adsorption of electrochemically active species from adsorbing on the metal surface without dampening the formation of the electric double layer and subsequent AC electrokinetic phenomena.[39] 96
The coated electrodes performed better than their uncoated counterparts when they were subjected to potential and frequency sweeps from 300 mV to 5 V, 100 Hz to 100 kHz (Figure 3.3 and 3.4). Whereas uncoated electrodes of both metal types were virtually unusable in the higher electrolyte concentrations of 1.0 TE and 1mM PBS, both silanes and polyelectrolytes were effective in allowing AC electrochemical phenomena to occur without electrode damage. At 5V, bubbling was not observed even at frequencies above 1 kHz for both coating types. The LbL polyelectrolyte coated chips were able to achieve and maintain higher applied potentials at lower frequencies than those of the silane coated chips. LbL polyelectrolyte coatings exhibited signs of possible electrode damage below 1 kHz at 5 V whereas the silane coatings consistently exhibited bubble formation on the electrodes at 4 V, 1 kHz in 1.0 TE. If operated beyond these limits, the electrodes were damaged and could not be salvaged for later experiments. To confirm these findings a cyclic voltammetry (CV) analysis was performed on both coating types and compared to bare gold. CV is commonly used to determine characteristics of electrochemical reactions, to develop anti-corrosion coatings[45] and to study self- assembled monolayers.[44,47,55-56,82-83] In tracing the redox probe at the surface of the working electrode on bare gold, two peaks are apparent on the forward and reverse scan indicating that the oxidized and reduced form of ferric cyanide concentrations were sufficiently depleted at 0.3 V and 0.01V respectively (Figure 3.5). This indicates an electrochemical reaction at the gold-solution interface is possible under the right conditions. However, for the silane and LbL polyelectrolyte coated electrodes, no peaks are evident, demonstrating the electrochemical reaction is sufficiently suppressed at the metal-solution interface.
97
Figure 3.5. Cyclic voltammetry curves following ferric cyanide reaction on bare gold (black) as well as silane (red) and LbL polyelectrolyte (green) coated gold. The peak at 0.3 V and 0.01 V for the gold curve exemplifies the oxidation and reduction reaction respectively. The lack of peak for the silane and LbL polyelectrolyte coated slides suggests the faradaic reaction has been sufficiently suppressed.
Given that the silanes and LbL polyelectrolyte coatings suppress electrochemical reactions at the metal surface, it was essential to determine whether or not the coatings affected AC electrokinetic phenomena and if coated chips were able to collect particles as effectively as bare chips. The most efficient particle collection began with a particle front that formed at the innermost electrode upon application of the electric field (Figure 3.6 a and b). The front traversed towards the center of the device forming densely packed collection of particles (Figure 3.6 c and d). Both the silane and LbL polyelectrolyte coatings exhibited ACEO and DEP as seen in Figures 3.3 and 3.4, however platinum coated electrodes were generally able to collect particles in higher electrolyte concentrations than that of coated gold electrodes. The results reported for the remainder of this section will be for coated platinum electrodes. 98
Figure 3.6. Particle collection for LbL polyelectrolyte coated platinum electrodes at (a) 0 s, (b) 10 s, (c) 140 s, and (d) 280 s at 800 mV, 300 Hz. The particle collection at the centrally located stagnant point is typically preceded by a distinct particle front that forms at the edges of the innermost electrode and traverses towards the center over time.
In DI water, as expected, both uncoated and coated electrodes exhibited AC electrokinetic phenomena and were able to collect particles in the geometric center of the collection area. Silane-coated electrodes generally required higher potentials and lower frequencies to achieve the same collection as observed in LbL polyelectrolyte trials (Table 3.1). For example, silane-coated electrodes required an increased frequency of 1 kHz, greater than the 300 Hz required by LbL polyelectrolyte coated electrodes to collect particles in the center of the electrode (Figure 3.4). Conversely, LbL polyelectrolyte coated electrodes appeared to enhance DEP and ACEO phenomena. Figure 3.7 shows the % Deviation from White (DFW) over the length of the electrode from the innermost edge of the electrodes (chip position equal to 0 m) to the geometric center of the collection area (chip position equal to 500 m). The bare electrodes exhibit significant particle collection with a % DFW 99
of 35% but fails to drive the particles to the center of the collection area as seen in the LbL polyelectrolyte curve in which the maximum DFW is approximately 10% greater. The maximum % DFW for the LbL polyelectrolyte coated chip occurs at 500 m, proving that more particles are drawn to the center of the collection area than for the uncoated chip.
Table 3.1: Optimal Collection and Maximum Applied Potential and Frequency Settings for Platinum Electrodes Optimal Collection Electrode Damage Coating Electrolyte Potential (V) Frequency (kHz) Potential (V) Frequency (kHz) Di Water 0.8 1 - - 0.001 TE 1 1 - -
0.01 TE 1 2 - -
0.1 TE 2 10 4 0.5 Silane 1 TE - - 4 1 1mM PBS - - 4 30
Di Water 0.8 0.3 - -
0.001 TE 0.8 0.3 - - 0.01 TE 1 1 - - 0.1 TE 1.2 5 - - 1 TE 4 40 5 < 100
LbL Polyelectrolyte 1mM PBS 4 10 4 10
100
Figure 3.7. Particle density analysis at 800 mV, 300 Hz for uncoated (black), silane coated (red) and LbL polyelectrolyte coated (green) platinum electrode chip in DI water after 280 s of particle collection.
This trend continued as the electrolyte concentration was increased. Since high applied potential and low applied frequency tend to cause electrochemical reactions at the surface of the electrode in high electrolyte concentrations, operating the chip at higher frequencies is most desirable when potentials above 1.2 V are required. In Table 3.1, the LbL polyelectrolyte coating consistently required higher frequencies than that of the silane coating to optimally collect particles. For electrolyte concentrations above 953 S/cm, silane-coated electrodes exhibited electrochemical degradation at potentials lower than ones needed to achieve collection. ACEO flow was observed for both silane and LbL polyelectrolyte coated electrodes in electrolyte concentrations up to 1 mM PBS. However, the ACEO flow was only adequate to collect particles on the LbL polyelectrolyte coated chips (Figure 3.4). Silanes coat the surface of the electrodes as a self-assembled monolayer (SAM) with the silane end attracted to the substrate (or the Si end of the thiolated silane that coats the metal electrodes). The monolayer coating inhibits electrically active species from reaching the metal electrode surface, but may also resist the potential drop necessary to generate ACEO flow. Higher potentials are required to drive particle collection using silane coated 101
electrodes. The polyelectrolyte has alternating layers of positively charged and negatively charged polymers. Although the polymer itself may limit the ionic motion within the coating, both ionic charges are available in the layer just above the electrodes as they would be in free solution. This would allow for positive charges to be attracted to negative electrodes and negative charges to be attracted to positive electrodes, just as they would be in solution. A LbL polyelectrolyte coating is more capacitive as it is made of multiple layers of polymers with alternating charges. The alternating charge layers allow ionic mobility more freely than the SAM silane layers and may, therefore contribute to the capacitive Stern layer as opposed to resist it. Studies have also shown LbL polyelectrolytes to be more durable than alkanthiols as alkanethiols alone are more prone to reductive and oxidative desorption.[39,56,84-87]
Figure 3.8. Time dependence analysis of particle density at 4 V, 10 kHz for uncoated, silane coated and LbL polyelectrolyte coated platinum electrode chip in 1 mM PBS.
Beyond enhancing particle collection and reducing electrochemical reaction at the metal-solution interface, the coatings also reduce adhesion of particles and cells. At frequencies above 10 kHz in low electrolyte concentration, particles collected by DEP in the small electrode gaps tended to stick to the substrate surface and uncoated chips fouled easily. The silane interfered with the electrostatic binding between the substrate and particle, thus 102
making the particles much easier to re-suspend. LbL polyelectrolyte coated chips exhibited a reverse flow phenomena that occurred at higher frequencies (Figure 3.4). The flow would change directions in the 10-100 kHz decade then resume center-directed flow below 10 kHz. ACEO flow is dependent on the ionic charge formation in the electric double layer. In the layered polyelectrolyte, the ions on the polymer are significantly less mobile than those in free suspension. From 10 to 100 kHz, the ions just above the electrode surface are diffusion limited and may not have sufficient time to reach the outermost edge of the electric double layer, altering the charge and therefore would be unable to maintain unidirectional ACEO flow. Understanding the regimes in which the reverse phenomena occurs is crucial to optimal particle collection and could possibly be used to clear the collection area between test samples.
3.3.4. On-Chip Collection of Large Particles and Cells Using the particle collection parameters optimized previously[33] and the new LbL polyelectrolyte coated platinum electrode chips, we further optimized collection parameters for 5 m latex particles. These parameters were used to collect yeast cells from suspension. The optimization parameters were first examined using DI water and later extended to tap water applications. Many environmental water pathogen-detection applications require sensors able to detect microbes in water with some degree of salinity. Most municipal water treatment facilities in add salts such as NaCl or CaCO3 to soften water or to remove particulate from drinking water. The tap water was collected from local Raleigh City Water systems and tested immediately for conductivity, which was equivalent to 0.1 TE or 0.1 mM PBS (Table 3.2).
103
Table 3.2: Equivalent Salt Concentrations of Common Biological Buffers to Raleigh City Tap Water
0.1 TE Buffer 0.1 mM PBS Buffer Component Concentration (mM) Component Concentration (mM) Tris (THAM)* 1 NaCl 1.37 EDTA** 0.1 KCl 0.027
Na2HPO4 · H20 0.1
KH2PO4 0.02 * tris(hydroxymethyl)aminomethane **ethylenediaminetetraacetic acid
Chamber heights were altered initially to determine the best flow field induction for collection. In DI water, 1 mm chamber heights to induce a bulk fluid flow profile sufficient for dense collection in the center of the chip while 750 m was required for tap water. Once favorable bulk fluid flows were established, optimal ACEO velocity was determined by altering parameters such as potential, frequency and number of electrode repeats. As reported earlier with uncoated electrodes, yeast cells had a tendency to stick to the substrate surface and required potentials beyond the electrochemical reaction threshold for uncoated electrodes resulting in gold lift-off of the electrodes (Figure 3.9 a and b). With the new LbL polyelectrolyte coated electrodes, higher ACEO velocities were necessary to carry yeast cells to the collection area than those used to transport the 5 m particles. In DI water, while the latex particles required 5 electrode repeats and 1 V, 5 kHz for collection, yeast cell collection used 10 repeats and 2V, 5 kHz. Similarly, in tap water at 50 kHz, 2 – 4 V were necessary for latex particle collection whereas 6 V was required for yeast cells (Figure 3.9 c-f). Both 5 m latex particles and yeast cells are collected at the center of the collection area within 15 mins. The LbL coating also appeared to reduce fouling in comparison to the previous uncoated chips.[33] To examine the antifouling properties of silanes, a silane terminus was introduced to the LbL polyelectrolyte coating. The first four layers were coated onto the electrode chip as previously mentioned, but the final layer was completed using polyacrylic acid (PAA). The OH- group on the polyelectrolyte attracted the vinyl groups in the Aquasil, thereby making the coating more hydrophobic. Hydrophobic surfaces attract non-ionic 104
surfacants such as Tween 20 commonly used to reduce non-specific binding; therefore the combination of both the vinyl silane and Tween 20 can be used to reduce bio-fouling.[88] Preliminary data showed the coating reduced particle sticking between the electrodes. The combined polyelectrolyte and silane coating could possibly be used to improve biosensors for suspended cells. However, the silane-LbL coating also dampened the AC electrokinetic effects, requiring higher potentials and lower frequencies to drive particle collection just as was seen in the silane-only trials.
Figure 3.9. (a) Collecting yeast cells with the electrode device was attempted with uncoated gold electrodes at 5V, 200 Hz in DI water. The cells tended to stick to the substrate. Within 80 s of applying potential, bubbles formed over the electrodes. (b) Gold lift-off occurred on the smaller electrodes, especially at the southwest corner. (c-f) Micrographs of particle collection trials of large particles and yeast cells in tap water with platinum polyelectrolyte coated electrodes. The 5 m latex particle trials were performed by toggling between 2 V and 4 V at 50 kHz. Particle collection is apparent after (d) 15 mins over (c) the initial micrograph. (e-f) Yeast cells were collected at 6 V 50 kHz after 15 mins. No electrode damage was apparent after either particle or cell collection trial with the coated electrodes.
3.4. Conclusions Most biological lab-on-a-chip applications including medical diagnostics, food-borne pathogens and environmental water testing require biosensors to be operational in high electrolyte conditions. Further, electrically driven microdevices can produce the flows 105
necessary to carry cells and particles to surface-bound detection areas, but typically require metal electrodes to do so. Generally, operation of metal electrode ACEO driven flow devices has been limited to low electrolyte and low potential regimes to prevent electrochemical reactions, thereby limiting their use for biological application. The coating techniques proposed here use membrane and interference inhibition to prevent electrochemically active ions from reaching the metal-solution interface. Ideally, the coating should prevent electrochemical reactions without dampening the AC electrokinetic phenomena used to drive microfluidic flows and particle collection. Cells and particles of varying sizes in electrolyte suspensions with conductivities up to 2 mS/cm has been shown using gold and platinum coated electrode particle collection device. Platinum appears to be more robust than gold in withstanding electrochemical reactions at high potentials in high electrolyte conditions especially, when used in conjunction with an effective electrochemical reaction inhibitor. As membrane inhibitors, hydrogels are too thick and porous to allow for charging of the electric double layer and to act as a barrier to electrochemically active species. Even the thinnest agarose coating was able to withstand potentials higher than 3V at 0.01 TE as compared to the silane and LbL polyelectrolyte coated electrodes. The thinnest coatings of silanes and LbL polyelectrolytes showed promise for electrochemical reaction inhibition in high electrolyte conditions. Using CV analysis, both coatings impeded the oxidation and reduction of ferric cyanide, proving that silane and LbL are effective interference and membrane inhibitors respectively. Most remarkably, the silane and LbL polyelectrolyte coated electrodes exhibited AC electrokinetic phenomena at relatively low potentials, proving their efficacy in ACEO and DEP applications. Both layers were thin enough to allow for sufficient electrode polarization and to drive ACEO flows for particle collection. The ions integrated in the LbL polyelectrolyte made it that much more effective than the silane coating, allowing for adequate formation of the electric double layer while being able to withstand potentials up to 6V in some conditions. Although LbL coatings are successful at promoting particle collection while reducing electrochemical reactions, more work could be done to investigate using other hydrophobic electroactive coatings such as polypyrrole or polyanaline.[47-48] 106
Using the design and operational characteristics reported in previous work[33], the LbL polyelectrolyte coated device’s ability to collect large latex particles and cells was established. The LbL polyelectroyte coating allowed for higher potentials to be applied to the electrodes to achieve the flow rate necessary to collect these cells and particles, even in tap water. The hydrophobicity of the PDAC in conjunction with a non-ionic surfactant also reduced biofouling of the substrate and electrode surfaces. Knowing the size, shape, membrane elasticity of the desired species to collect will allow the user to tune the device parameters and applied potential and frequency, making the coated electrode device applicable to a wide variety of cells and particles.
3.5 Acknowledgements
This research was supported by RTI International, Research Triangle Park, NC and a NIRT grant from the US National Science Foundation (CBET-0609087).
3.6 References
[1] Bhatt, K. H., Grego, S., & Velev, O. D. (2005). “An AC electrokinetic technique for collection and concentration of particles and cells on patterned electrodes”. Langmuir, 21(14), 6603-6612.
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CHAPTER 4
Particle-Cell Agglutination Based Impediometric Biosensor
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4.1 Introduction Electrochemical analytical methods offer sensitive and rapid means of detecting chemical species on or near metal electrodes or the metal-solution interface. Initially these methods were developed to better understand catalysis, electrochemical reactions and reaction kinetics of inorganic and organic chemicals. More recently, such techniques, most notably electrochemical impedance spectroscopy, have been extended to biological applications and used to better understand cell signaling pathways[1], DNA sequencing[2-4] and enzymatic reactions.[5-7] Biospecific functionalized surfaces are relatively simple to fabricate and have facilitated microscale research of DNA[2-4], whole cell[1,8-12] and tissue[13] activity for surface-bound electrochemical impedance. Electrochemical impedance spectroscopy is a measure of the combined resistance and capacitance of an electrochemical cell detected by a small amplitude signal over a range of frequencies. The impedance itself is the ratio of the potential phasor to the electric current phasor in an AC circuit and for a simple RC circuit can be described by the following: (Eq. 4.1) Z ZRe jZ Im
where
ZRRe (Eq. 4.2) 1 ZIm (Eq. 4.3) C
where ZRe is the real portion of the impedance, ZIm is the imaginary component of the impedance, R is the resistance, is the angular frequency and C is the capacitance. The real and imaginary components of the impedance are related to the phase angle shift of the applied amplitude signal. As the components of the electrochemical cell are changed, be it from a chemical reaction, enzymatic activity, alteration of the electrode surface, or agglutination of suspended particles, the resulting electrochemical impedance spectra is altered. The electrochemical impedance and phase angle can be modeled with an equivalent circuit in which each component represents a particular element in the electrochemical cell. The impediometric spectra and the equivalent circuit offer two means by which to detect cells and particles in suspension as long as the diameter of the analytes is in the same order of
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magnitude of the conductive working microelectrodes.[14] Therefore, a practical impediometric biosensor must have a rapid response rate; distinguishable impedance differences amongst control, negative, and positive samples; a simple analytical detection algorithm; high repeatability; and reproducibility. Impedance spectroscopy sweeps are typically conducted at low voltages making possible the detection of biological analytes using hand held devices. Whole cell study with electric biosensors is especially important for the development of rapid, microscale microbial detection for medical, environmental and food contamination sensors. While surface bound techniques have proven quite effective for cellular studies, the devices used are prone to biofouling and subject to error in electric outputs caused by settling of suspension components. These devices are nearly impossible to reuse. Developing label- free methods to detect freely suspended cells and biomolecules is desired to reduce biofouling and any effects the label might have on the function of the cell. Many of the impedance detection techniques use bio-specific functionalized particles to target a cell or particle of interest.[15-17]. These conglomerates are detected while in suspension[18], allowed to settle onto metal electrodes[19-22] or are manipulated to the electrode surface by electric[8,23] or magnetic[15-17] fields. Although all these techniques have been used successfully for whole cell detection, little is known about how particular agglutination methods affect impediomentric outputs. Two agglutination strategies in particular involve electrostatic binding and bio-specific binding. Agglutination by electrostatic binding involves altering the pH of a suspension to change the surface charge of one suspended species to promote electrostatic attraction between positive and negative particles. The pH at which the surface charge changes from positive to negative, the isoelectric point, varies for different charged particle types. The isoelectric point of yeast cells is at a pH of 4[24] while the surface charge of sulfate particles remains negative from pH 2 to pH 11.[25] When the pH is brought below the isoelectric point of yeast, the negatively charged latex and gold particles are attracted to the yeast cells (Figure 4.1). This method has been used to coat both gold nanoparticles and sulfate latex particles with proteins[26-28] but could be extended to coat extracellular proteins on cells with freely suspended particles.
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Latex Gold Yeast Only Particle Nanoparticle Yeast +Gold Yeast Yeast +Latex Cell
High pH (6.6)
Negative Surface Charge
Low pH (3.1)
Positive Surface Charge
Figure 4.1. Matrix depicting agglutination via pH. At a pH above the isoelectric point, the surface charge of cells is negative and negatively charged particles will be repelled. At a pH below the isoelectric point (pH = 4 for yeast cells), the surface charge of the cells will be positive and attract negatively charged particles.
Biospecific binding involves biologically derived attractions used to target specific ligands, proteins, antigens or cell types. Many biospecific binding methods use “lock and key” mechanisms in which two molecules are attracted to each other based on the shape of the molecule and the positions of the functional groups. As the name suggests, molecular attractions such as these are unique, reducing the likelihood that other, non-target molecules will be bound. Therefore, the greatest advantage to bio-specific binding is that particles can be designed to specifically target nearly any biological molecule. Antibody-antigen attractions are widely used as the active ends of the molecules that can be cleaved and used to functionalize substrate and particle surfaces.[8-9,29-30] Likewise, protein-specific binding of molecules such as streptavidin-biotin and actin is also used to functionalize surfaces for cell, protein, DNA and enzyme targeting for immunoassays.[31] Conductive particle tags have also been effective in enhancing electric signals for whole cell detection.[8,20-22] Bio-specific functionalized magnetic particles have been used in conjunction with magnetic fields in microfluidic systems for cell manipulation.[15-17] In those studies, cell and particle networks have been brought to the electrodes by magnetophoresis and the response time has been
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drastically improved. However the electrodes have been located at the bottom of the sample chamber and did not eliminate effects from settling of the suspension components. Any species that is within a distance on the same order of magnitude of the electrode width (~10 m), will alter the electrochemical impedance spectra.[14] If undesired cells or particles are allowed to settle in this region, they will be detectable in the impedance spectra and reduce the effectiveness of the device. In this chapter, several methods to improve upon existing surface bound and suspension-based impediometric biosenors are introduced by studying electrostatic and biospecific binding techniques. The devices described here use either settling or magnetic field manipulation to draw the cells and particles within the required proximity for detection by impedance. [14] As dead and live cells exhibit varying electrical properties[32], the effects of agglutination technique on cell viability and impedance spectra were extensively examined. Much work has been done to date using conductive particles to enhance electric detection outputs, however, little is known about the effect of using resistive particles. The effects of settling, particle material when cells are electrostatically bound to negatively charged particles on the impedance spectra and equivalent circuit model components was investigated. Using biospecifically functionalized particles with Con A lectin, we determined how the mode of transporting conglomerates of cells and particles to the electrode surface affects the sensitivity of impedance outputs. Finally, we introduce various simple solutions to increase the sensitivity of detecting a desired analyte in suspension with other particulates.
4.2 Materials and Methods
4.2.1 Sample Preparation / Agglutination Techniques The particle and cell suspensions were prepared separately for the electrostatic binding and bio-specific binding experiments. In the electrostatic binding agglutination trials, a 0.1 mM phosphate buffer solution (PBS) (Sigma Aldrich. St. Louis, MO) was prepared by dissolving the 20 mg tablet in 200 mL of DI water and diluting the resulting solution 100 times. To lower the pH to 3.1, 1 M acetic acid (Sigma Aldrich St. Louis, MO)
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was added to the 0.1 mM PBS solution (7.3 vol. %). An equal amount of Millipore DI water was added to 0.1 mM PBS to keep the electrolyte concentration equal between the low and high pH trials. Both high pH and low pH solutions were used as media to disperse Saccharomyces cerevisea, baker’s yeast (Fleischmann’s – ACH., Memphis, TN), 1m diameter surfactant-free white sulfate latex particles (Interfacial Dynamics Corporation. Eugene, OR) and 15 nm gold nanoparticles. The gold nanoparticles were synthesized using the procedure developed by Slot et al.[33] For all samples, the yeast cell content was 0.1 wt% with a 1:100 cell : particle ratio for yeast/latex particle samples and 1:1000 yeast/gold nanoparticle samples. For bio-specific binding promoted agglutination, each solution was prepared with 0.1 wt% yeast cells in 0.1 mM PBS solution with 0.01 wt % bovine serum albumin (BSA) and Tween 20 (PBS-BSA-T20) to prevent non-specific binding and adherence to the electrode chamber surfaces. Anhydrous calcium chloride (Sigma Aldrich, St. Louis, MO) was added to the PBS solution at a concentration of 15 mM.[34] Concanavalin A (Con A) functionalized 15 nm gold (Bangs Laboratory. Fishers, IN) and 1 m ferromagnetic particles (Bangs Laboratory, Fishers, IN) were used to promote the lectin binding. Suspensions of yeast cells alone, yeast cells and unfunctionalized particles and Con A functionalized particles were each prepared and subjected to a frequency sweep and impedance measurements. The cell / particle ratios were fixed at 1:1000 and 1:100 for gold nanoparticle and magnetic particle suspensions respectively. All samples were stored for 1 to 24 hours at 5ºC.
4.2.2 Yeast Viability Study The particle and yeast cell suspensions were made by either adding the yeast cells to a vial filled with the PBS-BSA-T20 mixture or by adding the PBS-BSA-T20 solution to a vial filled with yeast cells. The cells were then plated in three lines on maltose enriched agar plates at intervals of 5 minutes, 10 minutes, 30 minutes, 1 hour, 24 hours, 48 hours and 120 hours. After 24 hours, each plate was observed for yeast growth and judged using the rating scheme in Figure 4.2: 0 – no observable growth, 1 – growth on less than 50 % of the line, 2 – growth on 50%-75% of the line, 3 – full line with a few spots (greater than 75% of the line full) and 4 – full line of growth.
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Figure 4.2. Yeast viability study and growth judging examples. Judging criteria was as follows: 0 – no observable growth, 1 – growth on less than 50 % of the line, 2 – growth on 50%-75% of the line, 3 – full line with a few spots (greater than 75% of the line full) and 4 – full line of growth.
4.2.3 Interdigitated Electrode Chip Fabrication The gold interdigitated electrode chips were fabricated on 4-inch diameter, 1.0 mm thick borofloat silicon dioxide wafers (Mark Optics, Inc. Santa Ana, CA) at both the NCSU Nanofabrication Facility and at Research Triangle Institute using photolithography techniques. Negative photoresist, NFR016D2 (JSR Micro, Inc. Sunnydale, CA), was applied to the borofloat wafers via spincoating to 3.5 mm thickness and exposed to UV light through a photomask. The wafers were developed in Microposit® MF-319 (Shipley, Marborough, MA) Chromium (100 Ǻ) and Gold (1000 Ǻ) were applied to the wafer by E-beam evaporation. The wafers were treated with N-Methyl-Pyrrolidinone (NMP) (JT Baker / Mallinckrod Chemicals, Phillipsburg, NJ) to lift-off excess gold and chromium, washed, and the features of the wafers separated with a dicing saw.
4.2.4 Interdigitated Electrode Chip Types The first electrode chip design (Figure 4.3) consisted of two identical sets of interdigitated electrodes that were both energized during the frequency sweep and impedance measurements. Each set of interdigitated electrodes had 13 pairs of 15 m wide and 1 mm long electrode with a gap of 10 m. To contain the samples over the electrodes a 2 mm 5
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mm 150 m adhesive perfusion chamber (Grace Bio-Labs. Bend, OR) was placed over the electrode array. The redesigned electrode chip was approximately 2 cm 3 cm in size to facilitate chamber inversion. The single set of interdigitated electrodes was comprised of 30 to 50 pairs of 15 m wide electrodes separated by a 5 m gap (Figure 4.4).
25 Electrodes – each side Gap Width – 10 m
1mm Glass 13 mm
Gold
Figure 4.3. Gold on glass interdigitated electrode chip used in agglutination impedance spectroscopy measurements. The dual interdigitated electrode sets consisted of 25 electrodes on each side separated by 10 m gaps.
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Figure 4.4. Redesigned gold-on-glass interdigitated electrode chip used in impedance spectroscopy measurements. The interdigitated electrodes consisted of 30 to 50 electrode pairs. The electrodes were 15 m wide separated by 5 m gaps.
4.2.5. Impedance Measurements Electrochemical impedance spectroscopy (EIS) measurements were taken using an IM6e Impedance Measurement Unit with Thales software (Zahner Eletrik, Kronach, Germany) and a VersaStat potentiostat (Princeton Applied Research, Oak Ridge, TN). The potential was held constant at 10 mV and the AC current frequency range was varied from 1 Hz to 1 MHz. Bode and Nyquist plots were recorded and analyzed to determine relative impedance differences between the background solutions and test solutions. Suspended cells and particles were allowed to settle over the electrodes in the gold nanoparticle and latex trials for both the bio-specific and pH change agglutination methods. The chip was connected to the IM6 impedance measurement system via wires soldered to the electrode pads (Figure 4.5). The chip was placed electrode-side-up and the sample chamber was affixed over the electrodes. Before each particle and yeast cell sample was tested, DC current, DC potential, impedance and phase angle measurements were taken at 1 kHz after filling the chamber with deionized water to ensure the chip was functioning properly and that all residue from the previous trial had been removed. Approximately 20 L of sample was
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pipetted into the chamber three times to ensure the concentration of cells and particles housed in the sample chamber were representative of whole cell/particle suspension (Figure 4.6). A micropump was used to suction any residual suspension left at the chamber inlets. Once the sample was loaded in the chamber, another reading was taken at 1kHz. When a favorable impedance reading was confirmed, the chamber was sealed with 3M adhesive dot seals (Grace Bio-Labs, Bend, OR) and the first frequency sweep was started. Impedance spectroscopy measurements were taken every 10 minutes for a total of 90 minutes. The electrode chips used for the magnetic particle bio-specific binding trials were prepared in the same manner as those used in the gold nanoparticle and latex particle experiments. However, after confirming chip function and surface purity, the chamber was sealed and the entire chip was inverted such that the chamber side of the chip was facing the lab bench. A 1-inch rectangular ceramic magnet (Radio Shack, Raleigh, NC) or a 0.5 cm diameter rare earth metal magnet (Radio Shack, Raleigh, NC) was contacted with the borofloat glass substrate such that the interdigitated electrodes were below the magnet (Figure 4.7a and b). Impedance spectroscopy readings were taken every 5 minutes for a total of 20 minutes.
Figure 4.5. Photograph of electrode chip with soldered connections used for impedance spectroscopy measurements.
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Gold NP Yeast Cells Latex
Glass slide
Gold microelectrodes Figure 4.6. Side view of chamber for settling impedance experiments. The suspensions were loaded in the chamber from the top of the chamber and allowed to settle.
(a) (b)
Magnet
Figure 4.7. Magnetic particles can be manipulated with magnetic fields. The electrode sample chamber was inverted and a magnetic field induced just over the electrodes. (a) Unfunctionalized particles only were manipulated by the magnetic field and drawn between the electrodes. (b) For Con A magnetic particles, networks were formed with the yeast cells. Both magnetic particles and yeast cells were brought between the electrodes when the magnetic field was applied.
4.3 Results 4.3.1. Sample Preparation Observations.
Yeast cells were chosen for this study as model system of the behavior of various fungal and bacterial cells.[24,35] The cell size is large (~5 m) and easily observed with an optical microscope. Further, yeast cells can be conveniently maintained viable as they can be kept in a non-growing state for periods longer than a couple of weeks.[24,35] Fresh yeast cell suspensions were made for each trial conducted to ensure viable cells were in suspension. The yeast and particle suspension concentrations were prepared such that the yeast would form a monolayer over the electrodes and substrate in the sample chamber to ensure the yeast
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cells would be within 25 m of the electrodes upon settling, the requisite proximity for detection by impedance with the electrodes used in this study.[14] After the yeast cells had settled and before loading the yeast and particle suspensions in the sample chamber for impedance spectroscopy analysis, each of the microcentrifuge tubes was observed visually for evidence of agglutination (Table 4.1). Figure 4.8a displays each sample as fully dispersed while Figure 4.8b was taken after the samples had been allowed time to settle, approximately 15 minutes. After 15 minutes, sample 1 and 2 appeared to have latex and gold particles dispersed in the supernatant. However, the samples at pH 3.1, below the isoelectric point of yeast cells, have virtually no particles dispersed in the supernatant and the particles appeared to be dispersed amongst the settled yeast cells. This is particularly evident in the gold nanoparticle and yeast suspensions. When dispersed, gold nanoparticles with 15 nm diameters will appear red as they adsorb the lower wavelength light (green, blue and violet). As these gold nanoparticles aggregate or coat larger diameter cells, the color will darken to purple as the red wavelength light is absorbed while the lower wavelengths are reflected. In both the dispersed and settled systems, sample 2 appeared red while sample 4 appeared purple, indicating the gold nanoparticles were electrostatically bound to the cell surface (Figure 4.8).
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Table 4.1: Electrostatic Agglutination Samples of Yeast Cells and Particles as Pictured in Figure 4.8 Sample # pH Particle Type 1 7.8 Latex 2 7.8 Au Nanoparticle 3 3.1 Latex 4 3.1 Au Nanoparticle
(a)
(b)
Figure 4.8. Sample vials for agglutination via pH change. (a) Dispersed sample vials of (1) yeast + latex at high pH, (2) yeast and gold at high pH, (3) yeast + latex at low pH, (4) yeast and gold at low pH. (b) Settled sample vials of (1) yeast + latex at high pH, (2) yeast and gold at high pH, (3) yeast + latex at low pH, (4) yeast and gold at low pH. Yeast only samples not pictured.
The gold nanoparticle and magnetic microparticles for the biospecific binding agglutination technique were functionalized with concanavalin A (Con A), a ligand that binds specifically to the surface polysaccharides of yeast cells. The gold nanoparticles were used to easily compare results of the biospecific binding study with those of the pH
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agglutinated samples. Magnetic particles were used as they are easy to manipulate from outside the suspension chamber and are observable by optical microscopy. Similar to electrostatic agglutination trials, the unfunctionalized particles remained in suspension when the yeast cells were settled. Figure 4.9 depicts suspensions of (a) yeast cells alone, (b) yeast and unfunctionalized gold nanoparticles and (c) yeast and Con A functionalized gold nanopartilces. As in the pH agglutinated gold nanoparticle and yeast sample, the Con A functionalized particles appeared to be distributed amongst the yeast cells after settling indicating that the functionalized particles are bound to the cells.
(a)
1 2 3
(b)
1 2 3
Figure 4.9. Microcetrifuge tubes of (a) dispersed cells and particles and (b) settled cells and particles for (1) yeast only, (2) yeast and unfunctionalized gold nanoparticles and (3) yeast cells and Con A functionalized gold nanoparticles. The unfunctionalized and Con A magnetic particle / yeast cell suspensions were observed by optical microscopy at 5 minutes and 30 minutes after mixing to determine if the particles were bound to the cells. The microscope was focused on the top of the sample
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chamber at 10x magnification. A magnet was placed on the top of the chamber and particles and / or cells were drawn into the region of highest magnetic field. The functionalized particles appeared to bind to the cells and form networks of cells and particles within 5 minutes and form larger networks after 30 minutes. As expected, the unfunctionalized particles formed chains by magnetophoresis in the direction of the magnetic field without yeast cells attached (Figure 4.10c). The unfunctionalized particles and Con A functionalized particle and yeast cell networks were manipulable with the magnets and were brought to the top of the 150 m observation chamber within 5 minutes (Figure 4.10a and b). A brief viability study was conducted to determine if the biospecific binding or any of the cell preparation steps impeded the yeast cell’s viability. Functionalized particles bound to the yeast cell surface seemed to have little effect on the viability or cell growth of the yeast cells over 5 days as seen in Table 4.2. The unfunctionalized particles also appeared to have little effect on cell growth as compared to yeast alone. The cells, however need to be hydrated by the PBS-BSA-T20 solution for at least ten minutes without agitation. Agitating the cells before they become fully hydrated yielded little to no growth of yeast cell colonies. The manner in which the cells were hydrated, be it pouring the PBS-BSA-T20 solution over the cells or adding the cells to a vial of PBS-BSA-T20, had little distinguishable effect on cell growth.
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(a)
(b)
(c)
Figure 4.10. (a) Yeast cells and Con A functionalized magnetic particles after 5 minutes, (b) yeast cells and Con A functionalized magnetic particles after 30 mins, (c) yeast cells and unfunctionalized magnetic particles after 5 minutes.
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Table 4.2. Bio-specific Binding Cell Viability for Varying Cell / Particle / Biological Salt Solution Residence Times. Three lines of cells were plated and judged per criteria in Figure 4.2. Yeast + Yeast + Con A Unfunctionalized Functionalized Time Yeast Only Magnetic Particles Magnetic Particles 5 min 1.67 1.67 2.00 10 min 2.67 3.67 3.33 30 min 3.00 3.33 4.00 1 hr 2.67 3.33 3.67 24 hr 3.33 3.67 4.00 48 hr 3.67 3.67 3.67 120 hr 2.67 2.33 2.67
4.3.2. Biosensor Design Considerations: Impedance Analysis. All yeast and particle suspensions were subjected to an impedance measurement frequency sweep from 1 Hz to 1 MHz over gold interdigitated electrodes. An impedance sweep was performed every 10 minutes for a total of 90 minutes. The real and imaginary impedances for each sample’s data sets were normalized against the background solution to eliminate electrolyte effects and then used to calculate a normalized percent impedance change (NPIC) (Eqn. 4.4)
Z t,, Z t NPIC x PBS_ solution 100% Zt , (Eq. 4.4) PBS_ solution
where Zx is the real or imaginary impedance of a sample at a particular frequency () and time point (t) and ZPBS_solution is the real or imaginary impedance of the background PBS solution at the same frequency () and time point (t). The electrical drift from the capacitive double layer is apparent at frequencies below 1 kHz for most gold interdigitated electrode systems, therefore, frequencies of less than 1 kHz were excluded from the analysis. The NPIC for both the real and imaginary impedances were characterized for each agglutination technique and particle type to determine their unique impedance spectra and to evaluate which technique showed the most promise for a rapid, reproducible and easily analyzable solution for biosensing.
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4.3.3. Agglutination by Electrostatic Binding The particle to cell ratio was chosen to ensure coverage over half of the yeast surface once agglutinated. The particles would be present in the path of the electric current along with the yeast cells, resulting in an impedance difference between the agglutinated and non- agglutinated systems. The settling time (tsed) of bare particles and cells in suspension was calculated from the balance of the gravitational force of the particle with Stokes Law: tsed=h/u where h is the chamber height. The particle velocity, u, was calculated using Eq. 4.5 where p is the density of the particle, m is the density of the media, g is gravity, r is the radius of the particle, is the viscosity of the fluid. The calculated settling times (tsed) for each particle type are listed in Table 4.3.
2 2 휌푝 − 휌푚 푔푟 (Eq. 4.5) 푢 = 9휂
Table 4.3. Settling Times for Cells and Particle Types
3 Particle Type Radius (m) Density (g/cm ) tsed (min) Baker’s Yeast Cell 4.0 1.11 6.5 Sulfate Latex Particle 1.0 1.05 109 Gold Nanoparticle 1.5 x 10-2 19.33 26,450
The data in Table 4.3 indicates that the yeast cells should sediment faster than both 1 m latex particles and gold nanoparticles; therefore, for non-agglutinated systems, yeast cells alone should be within 25 m of the electrodes. In non-agglutinated cases, the particles should also sediment over the yeast cells, if they settle at all. Conglomerates of cells and particles should also settle faster than their unagglutinated counterparts. Settling played the greatest role in bringing the cells within the electric field in the pH agglutination experiments, exemplified by the drastic differences in NPIC (Figure 4.11). Electrokinetic forces such as dielectrophoresis (DEP) would not have occurred as the entire frequency sweep lasted for 134.9 seconds using the Zahner IM6e potentiostat. Therefore, current was only applied for approximately 1 minute at frequencies above 1 kHz, the characteristic range for DEP induction. Due to their larger conglomerate sizes, the latex and
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yeast cell samples settled fastest in the low pH samples. As a result, after 70 minutes, the NPIC changed less rapidly than that of the yeast cells alone or the yeast cell / gold nanoparticle solutions.
Yeast Only 750 Yeast + Latex Particles Yeast + Gold Nanoparticles
700
650
600
550
500
450 Normalized ChangeImpedance %
0 20 40 60 80 100 Time (mins)
Figure 4.11. NPIC versus time for low pH agglutination trials at 60 kHz. The yeast and latex sample NPIC changes little after 70 minutes. The yeast / latex sample (red) is distinguishable from the bare yeast (black) and gold nanoparticle / yeast (green) samples as they settled faster than the other agglomerates.
Figure 4.12 a and b present the data for the NPIC versus frequency for real impedance and imaginary impedance respectively at 70 minutes. Any significant difference in either impedance spectra was not evident until 60 min of the measurement. There did not appear to be much of a distinguishable difference in the real impedance amongst the varying particle systems for the low (solid shape) and high pH trials (open shapes). However, the high and low pH data sets are each clustered together with the low pH data set registering a lower resistance than that of the high pH data set. The real impedance is a function of the resistance in the system. Between 1 kHz and 100 kHz it is likely representative of the electrochemical activity on the gold electrode surface.[36] The addition of the weak acid increases the propensity of faradaic reactions at the gold-solution interface and is likely responsible for the real impedance shift.[37-39]
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(a)
Low pH Yeast Only 1000 Low pH Yeast + Latex Low pH Yeast + Gold High pH Yeast Only 800 High pH Yeast + Latex High pH Yeast + Gold
600
400
200
0
NormalizedImpedance% Change
-200 3 4 5 6 Log Frequency (Log (Hz))
(b)
Low pH Yeast Only 700 Low pH Yeast + Latex Low pH Yeast + Gold High pH Yeast Only 600 High pH Yeast + Latex High pH Yeast + Gold 500
400
300
200
100 NormalizedImpedance% Change
0 3 4 5 6 Log Frequency (Log (Hz))
Figure 4.12. NPIC vs. frequency for pH agglutination technique (a) real impedance, (b) imaginary impedance at 70 minutes. The low pH samples are distinguishable from the high pH samples with the agglutinated yeast and latex particles exhibiting the greatest NPIC. The most distinct NPIC differences within each pH grouping can be seen in the imaginary impedance spectra, representative of the capacitance of the electric double layer and any capacitive surface modifications.[36] By evidence of Table 4.3, yeast cells settled
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on the electrodes within 70 min. In the agglutinated suspensions of yeast cells and particles, the cells/particle aggregates would be settled on the electrodes within 70 min. For instance, the yeast only sample exhibits a greater NPIC than that of the other particle-cell suspensions in the dispersed samples at high pH (Figure 4.12b). The yeast cells, residing on the gold electrode surface will contribute to the capacitance component of the imaginary impedance, resulting in a larger increase in NPIC. In the low pH experiments, the NPIC of the latex/yeast suspension was significantly higher than the other two suspensions. Just as in the yeast-only sample at high pH, the larger conglomerates settled over the electrodes more rapidly with a larger surface area, thus resulting in the largest NPIC difference. The gold/yeast cell conglomerates exhibited similar NPIC to the yeast only system at low pH. This is mostly likely because the gold nanoparticles / yeast cell conglomerates did not have a diameter much greater than the bare yeast cell are, therefore, not distinguishable from yeast cells alone.
4.3.4. Impedance Analysis of the Process of Functionalized Particle Agglutination Gold Nanoparticles. The Con A functionalized gold nanoparticle trials were conducted similarly to these of the pH agglutination experiments with settling bringing the analyte to the gold electrode surface. However, distinguishable impedance differences were observed amongst the agglutinated and non-agglutinated systems for the functionalized gold nanoparticle/yeast cell conglomerates as early as 30 min, approximately half the time of the pH agglutination trials. At frequencies above 1 kHz, the real impedance of each cell/particle suspension is within +/- 10% of the PBS solution (Figure 4.13a), indicating that the resistance from the real impedance spectra can be attributed to the electrolyte solution. The imaginary impedance, again, demonstrates the greatest impedance differences amongst the suspensions. Above 10 kHz, the imaginary impedance of the agglutinated gold and yeast was 5 – 25% lower than the other two suspensions (Figure 4.13b). The formation of gold and yeast networks by bio-specific binding likely generates a more capacitive layer over the gold than the yeast cells alone in suspension (Katz). Since capacitance is inversely related to the imaginary impedance (Eqn 4.3), an increase of capacitance brings forth a decrease in the imaginary impedance as seen in other bio-specific binding applications.[15-17]
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Yeast Only Yeast + Unfunctionalized Gold (a) 40 Yeast + ConA Gold
20
0
NormalizedImpedance% Change
-20 3 4 5 6 Log Frequency (Log (Hz))
(b) Yeast Only Yeast + Unfunctionalized Gold 40 Yeast + Con A Gold
30
20
10
0
-10
-20
NormalizedImpedance% Change
-30 3 4 5 6 Log Frequency Figure 4.13. NPIC vs. frequency for bio-specific binding with gold nanoparticles (a) real impedance, (b) imaginary impedance at 30 mins. Yeast only (black) and yeast / unfunctionalized (red) particle suspensions exhibit similar impedance spectra, indicating little difference between the two suspensions. The electrochemical impedance spectra of the yeast / functionalized particle is distinct from the other two suspensions, especially above 10 kHz. The slope of the imaginary impedance curve is also less steep with respect to frequency and can also to be used to distinguish the agglutinated suspension from the non-aggluinated suspensions. The impedance spectra for the gold-yeast conglomerates is unique. Above 10 kHz, the imaginary NPIC crossed zero at different frequencies: yeast only at 16 kHz, yeast and unfunctionalized gold at 30 kHz and yeast with Con A gold at 50 kHz. Above 60 kHz, the
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imaginary NPIC of the agglutinated network of cells and Con A gold nanoparticles was approximately 15% lower than the non-agglutinated system. Additionally, the slope of the imaginary NPIC was less steep in the agglutinated suspension than that of the non- agglutinated one. Simple analytical programs can be written using both the NPIC slope and when the NPIC crosses zero to determine when a desired analyte is in suspension.
4.3.5. Magnetic Field Manipulation of Functionalized Particles Electrodes at Chamber Top. Using bio-specific binding with gold nanoparticles for cell agglutination improved the response time of the detection system, however, 30 min is still a longer detection time than desired for a practical biosensor. Since settling continued to affect the impedance spectra of the functionalized gold nanoparticles, the chip was inverted such that the electrodes were located at the top of the sample chamber instead of the bottom. Magnetophoresis of Con A functionalized particles bound to yeast cells was used to draw the cell-particle conglomerates to the electrode surface, reducing the sensing time to approximately 5 minutes. The impedance sweep was performed every 5 minutes for a total of 20 minutes. Distinct impedance spectra differences were observed amongst the suspensions after the first 5 minute time point. The impedance differences became more distinct over time and showed the greatest distinction after 15 minutes. Although BSA was added to prevent non-specific binding, a few cells did adhere to the glass surface of the electrodes (Figure 4.14). The real impedance of the unfunctionalized magnetic particles and yeast only trials did not vary with frequency, as the Con A-functionalized gold experiments, but the shapes of the curves were similar (Figure 4.15a). The similarities in curve shapes for the unfunctionalized particles and yeast cells indicate that the solution resistance is affected by the dispersion of the particles and cells. The 5-10% variance between the yeast only and unfunctionalized particle / yeast samples is likely due to the presence of the magnetic particles. The Con A functionalized magnetic particle trial did exhibit frequency dependence for the real impedance unlike the other two suspensions (Figure 4.15a). This indicates that other components in the suspension lent their electric properties to the equivalent circuit making the model more complex. The 3D networks of particles and cells
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that were formed and the resistance of the cells likely contributed to the resistance of the solution.[15-17]
Figure 4.14. Inverted interdigitated electrode chip with unfunctionalized magnetic particles and yeast cells. When a magnetic field was applied, the magnetic particles formed chains. A few cells were observed to adhere to the glass substrate. The magnetic particle/cell system showed the greatest promise for a rapid, repeatable, and reusable detection device. The imaginary impedance spectra for the unfunctionalized magnetic particle/yeast suspension and the yeast only suspension did not differ from the PBS solution as evident by the 0% NPIC (Figure 4.15b) due to the relatively low capacitance of magnetic particles. The imaginary impedance of the Con A magnetic particle trial differs from the unfunctionalized magnetic particle and yeast only trial by 10% in the 1 kHz – 10 kHz frequency decade, but as much as 150% above 500 kHz. The data analysis algothritm drawn from these results are more simple than those in the Con A gold nanoparticle trials using an if/else loop program based on the distinct NPIC difference above 100 kHz. For example, if the NPIC is greater than 5%, the cell type of interest would be assumed to exist in the test sample.
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(a) Yeast Only Yeast + Unfunctionalized Magnetic 70 Yeast + Con A Magnetic 60 50
40
30
20
10
0
-10
NormalizedImpedance% Change -20
-30 3 4 5 6 Log Frequency (Log (Hz))
(b) Yeast Only Yeast + Unfunctionalized Magnetic Yeast + Con A Magnetic 160
140
120
100 80
60 40
20
NormalizedImpedance% Change 0
-20 3 4 5 6 Log Frequency (Log (Hz))
Figure 4.15. NPIC vs. log frequency for biospecific binding with magnetic particles (a) real impedance, (b) imaginary impedance at 15 mins. Yeast only (black) and yeast / unfunctionalized (red) particle suspensions exhibit similar impedance spectra, indicating little difference between the two suspensions. The electrochemical impedance spectra of the yeast / functionalized particle is distinct from the other two suspensions, especially above 100 kHz.
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4.3.6. Equivalent Circuit Modeling of Agglutination Methods
Equivalent circuit models were developed using ZSimpWin software to further analyze the patterns noted in the NPIC impedance spectra. Such models have been used in other impedance biosensor works and have been proven effective to determine changes just above the electrodes for surface bound techniques.[9,16-17,36] Varshney et al. suggested a model including two constant phase elements (CPE) and solution resistance parallel to the capacitance just above the electrode. This model was used for the pH agglutination study at 70 mins (Figure 4.16a) and showed good fit for frequencies above 1 kHz (Figure 4.16b). More complex circuits involving series of resistors and capacitors were attempted; however, the noise in the data below 1 kHz resulted in significantly higher error in the curve fits. The error for the model in Figure 4.16a, although lower than most equivalent circuit fits was still quite large for one of the CPE parameters at greater than 100%; however the fit for the resistive and capacitive elements were less than 10%. If the phase angle approaches -90°, the circuit is capacitive dominant while a phase angle of 0° is resistance dominant. In a purely capacitive circuit, the voltage sine wave is -90° out of phase with the current sine wave while in a purely resistive current, the voltage and current waves are proportional and in phase.[40] In the pH agglutination study, the phase angle does not reach either extreme, thus, the resistance and capacitance play a key role together in the impedance spectra. At the extremes of this data set, it appears the pH agglutinated suspension impedance spectra is more dominated by the capacitive elements below 10 kHz and is dominated by the resistive elements above 10 kHz (Figure 4.16b). Each branch of the circuit can be defined by the parameters in the following equations:
1 푍 = 푅2 + (Eq. 4.6) 1 푆 휋푓푄 2
1 (Eq. 4.7) 푍2 = 2 2휋푓퐶푒푙푒푐푡푟표푑푒
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where Rs is the resistance of the solution, Q is the constant phase element (CPE), f is the frequency and Celectrode is the capacitance at the electrode surface. The total impedance of the system is,
1 1 1 = + (Eq. 4.8) 푍푇표푡푎푙 푍1 푍2
As expected, the solution resistance and electrode capacitance for the PBS solution was less than that of the suspension samples as the cells and particles in suspension will contribute to the bulk resistance and capacitance near the electrode surface (Figure 4.17 a and b). Even more remarkable is the result that the solution resistance varies little for the yeast only and yeast / gold nanoparticle suspension between the high pH and low pH solutions. This confirms the particle diameter is not increased to the degree of that of the latex- surrounded particles as was gleaned from the NPIC difference in Figure 4.12a. In both the solution resistance and electrode capacitance, the yeast / latex suspension shows the greatest difference at 100 ohms and 3 x 10-9 F respectively. Although their structure makes cells a natural capacitor, certain cell types, like yeast, have been shown to be resistive in electrolyte solutions.[32] The capacitance drops for all cell and particle agglutinated suspensions suggesting more cells are present on the electrodes.
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Figure 4.16. (a) Equivalent circuit and (b) model fit for pH agglutination study at 70 mins. This model is similar to those used in other surface bound impedeometric biosensors[16-17] and is relatively simple involving two constant phase elements (CPE), solution resistance (Rs) and electrode capacitance (Celectrode).
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Figure 4.17. (a) Equivalent circuit and (b) model fit for pH agglutination study. This model is similar to those used in other surface bound impedeometric biosensors[16-17] and involves two constant phase elements (CPE), solution resistance (Rs) and electrode capacitance (Celectrode).
The model developed for the bio-specific agglutination suspensions takes into account the more complex cell-particle networks and addition of cells and particles at varying distances from the electrodes through the chamber. The solution resistance or bulk resistance (Rs) is the furthest distance from the electrodes, residing at the outer edge of the electrode components (C3, R3), electrical double layer components (C2, R2), and the components just beyond the electric double layer (C1, R1) (Figure 4.18a). Data from the 30
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min gold Con A trial and from the 15 min magnet Con A trial were fit to this model to better understand the results from the NPIC analysis. The PBS alone curve resulted in the largest error at 30% for each experiment. However, the model proved to provide a good fit for the suspension samples. The error for the model fit for both the Con A functionalized gold and magnetic particle suspensions ranged from 2% to 15% and 3% to 17% respectively. Both the Con A functionalized experiments appear to be capacitance dominated below 100 kHz and resistance dominated above 100 kHz (Figure 4.18b and c). Equations similar to those used for the pH agglutination model can be written to describe the behavior of the Con A functionalized particle trial model with the total impedance being a summation of each series circuit branch impedance:
1 1 1 1 1 = + + + (Eqn. 4.9) 푍푇표푡푎푙 푍푆 푍1 푍2 푍3 where
푍1 = 푅푆 (Eqn. 4.10) and
1 2 푍푛 = 푅푛 + 2 (Eqn. 4.11) 2휋푓퐶푛
The variables Zn, Rn, Cn are the impedance, resistance and capacitance of each series branch with n = 1 - 3.
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Figure 4.18. (a) Equivalent circuit and model fit for (b) Con A functionalized gold nanoparticle and (c) Con A functionalized magnetic particle agglutination studies at 30 and 15 mins respectively. This model includes a resistance and capacitance component near the electrode surface (R1,C1) and two sets of components between the electrode and bulk solution (R2, R3, C2, C3).
Just as in the NPIC analysis, settling appeared to have a major impact on the parallel circuit capacitances as well as the solution resistance. In the gold Con A study, the solution resistance appeared to be much greater for the Con A functionalized gold particle suspension than that of the PBS, yeast only and unfunctionalized particle / yeast suspensions (Figure 4.19a). The yeast only and unfunctionalized particle suspensions should still have a number of cells and particles in suspension after 30 mins, so these will contribute to the solution resistance. The sizes of the conglomerates formed by the Con A functionalized particles and
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yeast can vary widely to range in size from 2-cell networks up to networks of 10+ cells. Therefore, the rate at which they settle is also relatively unpredictable. Smaller conglomerates may still be suspended after 30 minutes as well and will also contribute to the solution resistance. Since the diameter of the conglomerate is greater than the nanoparticles and yeast cells alone, the agglutinated particle-cell suspension will contribute to increasing the solution resistance. This is consistent with the observations made in the real NPIC in Figure 4.13a in that the shapes of the two non-agglutinated suspensions are similar to each other while the agglutinated suspension curve is unique. The capacitance analysis offers some insight as to the behaviors seen in the imaginary NPIC as well. The capacitance just above the electric double layer (C3) does not vary among the suspensions (Figure 4.19b). The greatest capacitance differences amongst the suspensions can be seen in the electric double layer capacitance (C2) and the capacitance at the electrodes (C1) and most likely corresponds to the phenomena seen in the NPIC between 10 kHz and 500 kHz. High frequencies are required for the current to penetrate the cell membrane[16-17,41], therefore it is unlikely that the cells could be detected by impedance spectroscopy below 1 kHz. The equivalent circuit model parameter outputs for the magnetic particle experiments showed similar trends to the equivalent circuit model for the functionalized gold NPIC. When the magnetic field is applied for 15 min, the solution resistance shows no statistical difference amongst the samples (Figure 4.20a). Due to the size of the chamber (250 m) and the inversion of the chamber, very few of the cells, particles or conglomerates should either settle on the chamber cover or be drawn to the electrodes side of the chamber by the magnetic field. The capacitance components also follow similar trends to the gold Con A experiments in that C1, C2 and C3 exemplify the behavior seen in the NPIC in Figure 4.15b from 100 kHz to 500 kHz, 10 kHz to 100 kHz and 1 kHz to 1 kHz respectively (Figure 4.20b). The greatest capacitance difference between the agglutinated and non-agglutinated samples is 1.5 x 10-8 F between the two magnetic particle suspensions at the electrode surface
C1.
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Figure 4.19. (a) Bulk solution resistance of the ConA functionalized gold experiments with PBS alone, yeast alone, yeast and unfuctionalized particles, and yeast / ConA functionalized particles. Due to settling and the networks formation, the Con A functionalized particles and yeast cell networks exhibit just slightly higher resistances and contribute to the solution resistance. (b) The capacitance of the components from the electrode surface to just above the electric double layer. The capacitive component closest to the electrode surface C1 is predominant from 100 kHz to 500 kHz. Components at and just above the electric double layer, C2 and C3 are evident in the NPIC from 10 kHz to 100 kHz and 1 kHz to 10 kHz respectively.
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Figure 4.20. (a) Bulk solution resistance of the ConA functionalized gold experiments with PBS alone, yeast alone, yeast and unfuctionalized particles and yeast / ConA functionalized particles. Due to settling and the networks formed, the Con A functionalized particles and yeast cell networks exhibit just slightly higher resistances and contribute to the solution resistance. (b) The capacitance of the components from the electrode surface to just above the electric double layer. The capacitive component closest to the electrode surface C1 is predominant from 100 kHz to 500 kHz. Components at and just above the electric double layer, C2 and C3 are evident in the NPIC from 10 kHz to 100 kHz and 1 kHz to 10 kHz respectively. Although it was assumed that the inverted chamber would show more distinct differences in the capacitance among the samples, cells were observed adhering to the substrate surface between the electrodes for the unfunctionalized magnetic particle experiments and may have affected the impedance spectra (Figure 4.15). Further, the particle
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and cell suspensions used for these experiments were dense enough to form a monolayer of cells and particles over the electrodes if they were to settle evenly across the electrodes. The smaller conglomerates were easily manipulated with magnets, yet some of the larger particle / yeast cell conglomerates were difficult to maneuver around other dispersion media in the dense suspension and may not have been present at the electrodes during detection, thereby reducing the sensitivity to a capacitance change.
4.3.7. Interdigitated Electrode Chip Redesign The interdigitated electrodes were redesigned to facilitate coupling with microfluidic chambers and chamber inversion. Using more electrode pairs, with smaller gaps was expected to increase the sensitivity of the impedance measurement, such that cell / particle conglomerates closer to the electrodes could be detected. This increased sensitivity was also thought to result in greater detectable impedance differences, thereby resulting in detection algorithms robust enough to require little data processing. Impedance measurements for these experiments were also performed using the VersaStat 6 potentiostat system. As magnetic particle conglomerates exhibited the fastest response time and greatest impedance differences, the magnetic particles were solely studied with the redesigned chip. An initial examination of the yeast only, unfunctionalized particles / yeast and Con A functionalized particles / yeast suspended in DI water was performed to determine the impedance functionality of the chip without interference from electrolytes in the media. Just as in the agglutination study, once the chamber was filled with sample, the chamber was inverted and impedance spectra data was collected every 5 minutes for a total of 20 minutes both with and without an applied magnetic field. As expected, the impedance spectra of the magnetic particle suspensions with the new chip redesign was sensitive enough to detect cells on the electrode surface with little data processing as expressed by the absolute imaginary impedance (Figure 4.21). Between 5 kHz and 100 kHz, the phase angle of each suspension’s spectra dropped below 45°, but never reached zero. This frequency range was resistance dominated. The imaginary impedance decreased when cells were brought to the electrode surface, but increased when magnetic particles alone were brought to the electrode surface. The yeast cells are conductive in DI water and caused a drop in the absolute impedance in the
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resistance-dominated frequency range. The magnetic particles alone are comprised of flakes of magnetite imbedded in a latex polymer particle. By themselves, the magnetic particles are resistive and increase the impedance in the resistance-dominated frequency range when brought to the electrode surface. In terms of a detection algorithm, each of the samples displays unique imaginary impedance spectra with the non-agglutinated suspension giving rise to an increase of 50 ohms and the agglutinated suspension a decrease of 100 ohms for the agglutinated system between 10 kHz and 100 kHz. This is consistent with magnetic particle- based impedance biosensors reported elsewhere.[15-17]
Figure 4.21. Imaginary impedance spectra from 1 kHz to 1 MHz for yeast only (black), unfunctionalized magnetic particles / yeast (red) and Con A functionalized magnetic particles (green) suspended in DI water at 15 minute time point. The distinctive difference between the functionalized and unfunctionalized particle samples between 10 kHz and 100 kHz indicates cells are present on the electrode array surface.
In order to develop a protocol and algorithm for determining cell concentrations, these experiments were repeated with the suspensions in a PBS buffer with Tween 20 and BSA at varying particle and cell concentrations. Diluting the cell and particle suspensions facilitated conglomerate manipulation and a greater number of large Con A particle and cell conglomerates were observed between the electrodes when the magnetic field was applied. However, for all cell / particle dilutions with each particle type, the solution resistance completely overshadowed any impedance signal from any capacitive component. From 100 kHz to 1 MHz, the phase angle for each suspension dropped below -45° and approached -20°
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with the circuit being more resistance dominated in this frequency range. The phase angle for each trial remained between -80° and -60° from 1 to 50 kHz, indicating the circuit was dominated by a capacitive component. However, analysis of each spectra produced no distinguishable differences in NPIC, slope of the impedance with respect to frequency or phase angle shift in any frequency decade. Each suspension spectra shape matched that of the PBS background spectra at all frequencies as well. Although equivalent circuit modeling was attempted with the data, none of the models developed for the agglutination study fit the data for the redesigned chip when electrolyte was present. The electrolyte in the media likely increased the faradaic impedance response as the total surface area of the gold electrodes was increased in the redesigned chip.[36] Similar issues were reported by Yang et al when trying to detect E. coli using magnetic particle manipulation to enhance impedance spectroscopy outputs.[15] To reduce the faradaic impedance response and any cell adhesion, the electrodes were coated with a thiolated silane and Aquasil. It was speculated the silane would enhance the capacitive and resistive components of the electric double layer, however, the impedance spectra only indicated the presence of the silane instead of the particles or cells brought to the electrode surface.
4.4. Conclusions The impediometric agglutination experiments show promise for forming the core of rapid, reproducible and reusable biosensors. No particles or antibodies are bound to the substrate surface, which allows the electrode chip to be reused for multiple sample tests. The most emphasized impedance differences are seen in the analysis of imaginary component of the impedance. Although all agglutination methods showed some degree of imaginary impedance difference, the magnetic particle bio-specific binding combined with magnetic field manipulation exhibited the most rapid and distinct response with up to a 150% impedance difference over non-agglutinated systems. The magnetic particle method also has the largest potential for an easily programmable algorithm for use in a hand-held device. Inverting the chamber appears to improve the selectivity of the magnetic particle technique,
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however, more work needs to be performed in order to determine how cell and particle concentrations affects the overall impedance performance of the device. The redesigned chip would be effective for detecting analytes in applications requiring low electrolyte concentrations or deionized water as the suspension media. Faradaic impedance will continue to be problematic so long as the gold electrode surface area is high as compared to the surface area of the analyte. Although silanes completely dampened any signal from the electric double layer, layer-by-layer polyelectrolyte coatings or other conductive coatings might reduce the faradaic impedance response but allow for cells and particles to record a distinguishable impedance spectra difference. Yet another way to eliminate faradaic impedance effects of the electrolyte is to use a known redox probe. As the surface area is covered with cells and particles, the signal from the redox probe will be measurably diminished. The redox probe would add components to the equivalent circuit model and may help in determining and modifying to the electrode surface.[36] The impedance spectra and equivalent circuit modeling results could also be confirmed with cyclic voltammetry (CV) analysis. Cyclic voltammetry is commonly used to determine alterations to a working electrode surface, especially for corrosion applications.[37-39,42-47] This method is sensitive enough to determine the orientation of molecules in self-assembled monolayers (SAM) on metal electrode surfaces[48-60] and should be able to determine surface area coverage of cells and particles on the electrodes. Most devices, including the one introduced here use stagnant chambers for impedance spectra analysis. To make these devices more amenable to being used in hand-held applications and reusable, integrating microfluidic chambers will become necessary. Detecting cells carried by microfluidic flows will allow the detection chamber to be more easily cleared between samples when the magnetic field is removed. Further, the desired analyte will form conglomerates and be immobilized by the magnetic field while other species in suspension will be washed from the microfluidic chamber. Such improvements will vastly increase the number of applications for the magnetic particle impediometric biosensor to include samples with multiple cell types or other targets of interest such as blood, environmental water samples or raw food samples.
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CHAPTER 5
Summary
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Summary Two novel AC electric field driven cell manipulation and detection devices have been introduced in this dissertation work. The first device uses AC electrokinetic phenomena typically applied in microfluidic pumping to collect particles and cells. Dielectrophoresis (DEP) and AC electroosmosis (ACEO) trap cells and induce bulk fluid flows that carry a particular analyte to a predetermined stagnant collection area. This simple device was constructed by repeating sets of narrow and wide gold-on-glass microelectrodes separated by gaps of varying widths, using asymmetry of electrode geometry to promote unidirectional flow to each set. These electrodes were designed in a rectangular or circular shape to direct ACEO flow to the center stagnant region of the chip where the cells and particles were collected. To successfully collect cells and particles, a number of performance parameters were optimized: electrode width, gap width, chamber height, potential, frequency, number of electrode repeats, and electrode geometry. The robustness of the device’s capability of collecting cells and particles of varying sizes was also determined while increasing the electrolyte concentration of the suspensions. While the device is able to collect cells and particles of varying sizes in DI water, increasing the electrolyte concentration resulted in electrochemical reactions to deteriorate the metal electrodes. Particles and cells were observed adhering to the electrode and substrate surface further diminishing the device’s collection functionality. Observation of electrochemical reactions at the metal-solution interface lead to a study of electrode materials and coatings to improve the performance and robustness of the device in high-electrolyte concentration conditions. Gold and platinum were examined as they are inert noble metals which are less likely to succumb to electrochemical reactions. The platinum electrodes performed better than the gold and were able to withstand potentials and frequencies much higher than gold in electrolyte concentrations with conductivities equivalent to 953 S/cm. Overall, the silane and polyelectrolyte coatings reduced the propensity of electrochemical reactions while hydrogel-based coatings did not. Although hydrogel coatings were promising in theory, they appeared to enhance electrochemical reactions, even in DI water. Hydrogels are pourous and allow easy transport of 161
electrochemically active species to and from the electrode surface. Any changes in pH caused by the hydrogel also facilitated ion exchange at the solution-metal interface. Both the polyelectrolyte coating and the silane coating acted favorably as a barrier to the electrochemical reactions, but the polyelectrolyte did not inhibit AC electrokinetic phenomena necessary to drive particle collection. The LbL polyelectrolyte coating technique was used on chips to collect cells and particles in tap water, demonstrating the device’s ability to be employed in environmental applications. The second device consisted of interdigitated electrodes to detect cells and particles in suspension using electrochemical impedance spectroscopy and agglutination methods. Small changes in the medium near the working electrode can be interpreted based on the differences in the impedance spectra. Cells and particles have resistive and capacitive properties and will therefore cause changes to the impedance spectra when they are brought close to the electrodes. Two particle and cell agglutination techniques were examined to determine the relative effect on the impedance spectra. Agglutination by electrostatic binding involves changing the surface charge of the cells by increasing or decreasing the media pH beyond the protein or cell’s isoelectric point. The pH agglutinated conglomerates were allowed to settle over the electrodes. Agglutination by biospecific binding uses particles functionalized with molecules that specifically target a particular cell to form conglomerates. ConA-functionalised particles were used to bind to lectin on the yeast cell surface. These conglomerates were either allowed to settle (gold nanoparticles) or driven to the electrodes by applying a magnetic field (magnetic particles). All agglutination methods resulted in distinct differences in the impedance spectra, however the methods using settling to bring the analyte conglomerates to the electrode surface were too slow for practical biosensor applications. The functionalized magnetic particle and yeast cell conglomerates were detectable at 5 minutes, but showed the most significant impedance spectra differences at 15 minutes. Biospecific binding also allows for the impediometric biosensor to be useful in selectively detecting desired cells types in samples containing multiple cells or other debris. 162
Relevance to Inegrated Lab-On-A-Chip Devices The AC electric field driven operations on the microscale discussed in this thesis are amenable to integration with other lab-on-a-chip manipulation, separation and detection techniques. The format of the ACEO particle collection device makes it easy to integrate with optical waveguides as it can be coupled with evanescent waveguide techniques. The use of metal microelectrodes for microfluidic pumping has eluded application in many point- of-care diagnostics or biological applications because they are prone to electrochemical reactions. The electrode coating techniques introduced alleviate this issue allowing microfluidic pumping to be used in devices involving electrolyte-containing suspensions. ACEO pumping on a chip could facilitate studies involving DNA analysis, drug development, and cell signaling studies. Creating gradients of nutrients or toxins over adherent cells for many cell signaling studies is crucial. ACEO devices can be used to finely tune concentration gradients over the cells by inducing predictable flow rates in multiple microfluidic channels at once. To increase the efficacy of using coated electrodes in high- electrolyte suspensions, other coating techniques, such as electrically active coatings should be examined. Lastly, AC electrokinetic pumping can be combined with impediometric biosensors to make an all-in-one electronic cell detection device. Electrical or magnetic means to manipulate cells and particles can facilitate detection in microfluidic devices. For example, when using ConA functionalized magnetic particles in microfluidic flows, the magnetophoretic force should be able to overcome settling and fluid drag and draw the cells to the electrode surface. Once the cells are brought between the electrodes in the presence of microfluidic flows, the impedance spectra from 1 kHz to 1 MHz should emulate the spectra obtained in Chapter 4 in a stagnant chamber. Cells and particles carried by fluid flow allow concentration of the desired analyte on the electrodes and also potentially reduce detection time. For this to be realized, however, the impedance electrode system must be made to be more sensitive to changes caused by cells and particles at the electrodes as opposed to the electrolyte in the media. Therefore, other electrochemical analysis methods such as cyclic voltammetry could be implemented to confirm impedance spectroscopy results. The 163
impedance biosensor electrodes might also benefit from electrically active coatings to reduce any faradaic responses from the electrolyte media. Overall, the AC electric field methods of cell manipulation and detection introduced in this dissertation can lead to improved lab-on-a- chip features in medical diagnostics, food contamination and environmental water sampling applications.