Instituto Tecnológico y de Estudios Superiores de Monterrey
Campus Monterrey
School of Engineering and Sciences
DC- Voltage reduction for electrokinetic particle trapping in PDMS-based microfluidics
A thesis presented by
Cinthia Janet Ramírez Murillo
Submitted to the
School of Engineering and Sciences
in partial fulfillment of the requirements for the degree of Master of Science
In
Nanotechnology
Monterrey Nuevo León, Nov 29th, 2020
4
Dedication
My first version of this was becoming too long. But I figured a shorter message might be more significant since there is no way I will express all my gratitude in one page. I will just say this:
To my family, including friends who have become part of my family, thank you for believing in me and for your unconditional support and encouragement. I love you and I thank God for the opportunity to have you in my life, and for the strength and ability to enjoy the journey of the past two years.
5
Acknowledgements
I would like to express my deepest gratitude to Dr. Víctor Hugo Pérez González for all the support and guidance he provided for me during the entire master’s program and for giving me the opportunity to be a part of the Interface Science Research Group. Thank you for all your patience, encouragement and advise in the process of brainstorming, selecting, and developing this thesis research project and for introducing and guiding me through the publication process in the form of a review article.
I would also like to thank Dr. Blanca H. Lapizco Encinas for her support and guidance in this project and for allowing the collaboration with the Microscale BioSeparations Laboratory at Rochester Institute of Technology on this project.
I would also like to acknowledge MSc. Braulio Cárdenas Benitez for being willing to share details and tips from his previous experience working with microfabrication and trapping experiments. Thanks for always being kind to answer all of the ‘quick questions’ and clarifying concepts.
Finally, I want to thank Tecnológico de Monterrey for the support with tuition expenses through a scholarship during the 2 years of this program. Also, the Nanosensors and Devices group of Tecnológico de Monterrey for their sponsorship on this project and to Consejo Nacional de Ciencia y Tecnología de México (Conacyt) for the support with living expenses.
6
DC-voltage reduction through constriction geometry optimization for electrokinetic particle trapping in PDMS-based microfluidics
by
Cinthia Janet Ramírez Murillo Abstract
The objective of this work is to reduce the voltage requirement for particle manipulation and trapping in an insulator-based microfluidic channel. Insulator-based microfluidic devices have been used in the past for particle analysis, separation, and concentration.
Although some efforts have been successfully carried out to manipulate particles in microfluidic channels of this type, the electric fields required for particle movement and trapping are generally higher than 100 V cm-1 [1], limiting the possibility of creating an integrated, portable device that is suitable for point-of-care applications. Starting with a two-post geometry for the insulating feature in our channel, we amplify the electric field at the center of the channel through dimensional optimization of the constriction and the post diameter, lowering the voltage required to be applied across the channel in order to achieve particle displacement and trapping [2]. The present work includes the fabrication and experimental trapping results obtained in the channel designs produced after a modelling and optimization process to select the most efficient geometries to be created.
7
List of Figures
Figure 1. DEP concept. A Alignment of an effective electric dipole moment in the presence of a non-uniform electric field. B Electric field distortion in an electrode- based device and its effect on a particle that is more polarizable than its suspending fluid medium (positive DEP – pDEP). C Electric field streamlines and their distortion in a device with insulating posts, resulting in higher electric fields at the post constrictions. ... 14 Figure 2. Different reported applications for DEP based microdevices. A iDEP device used for the extraction of plasma from blood. Reprinted with permission from [15] copyright (2015) Springer-Verlag Berlin Heidelberg. B eDEP device used for the extraction of plasma from blood. Reprinted with permission from [16]
copyright (2015) AIP Publishing LLC. C eDEP device used to pattern cells into a hexagonal matrix for skin tissue engineering. Reprinted with permission from [17]
copyright (2019) Springer-Verlag London Ltd., part of Springer Nature. D DNA sorting in an iDEP device. Here, DNA molecules of different size are sorted and redirected to different outlets to posterior collection. Reprinted with permission from [18] copyright (2016) American Chemical Society. E eDEP device used to separate particles with different properties; in this case live and dead yeast cells.
Reprinted with permission from [19] copyright (2019) American Chemical Society.
F contactless dielectrophoresis device used to sort three different particles based on their size. Reprinted with permission from [20] copyright (2016) Chinese Mechanical Engineering Society and Springer-Verlag Berlin Heidelberg. ... 15 Figure 4. Fabrication process for dielectrophoretic microdevices. A Fabrication process for a carbon electrode-based device using photolithography and pyrolysis. Reprinted from [42], copyright (2016) AIP Publishing. B Standard fabrication process using photolithography (i-iv) and soft lithography (v-vi) to produce an insulator-based device. Reprinted from [43], copyright (2007) Nature Publishing Group. ... 22 Figure 5. Examples of electrode-based microdevices. A 3D electrode array (i) SEM image of the electrodes (ii)-(iv) show the device performing a separation of HCT116 (green) and lymphocytes (red). Reprinted with permission from [60] copyright (2018) The Royal Society of Chemistry. B eDEP device used to concentrate HeLa cells. Reprinted with permission from [61] copyright (2014) The Korean BioChip Society and Springer. C interdigitated electrode array in combination with a layer of atom-thin graphene to capture different particles such as polystyrene beads, nanodiamonds and DNA. Reproduced from [48]. D device with interdigitated- irregular electrodes used for the concentration and detection of CA 19-9, a pancreatic cancer biomarker. Reproduced from [62]. E serpentine-shaped carbon electrode used to separate live from dead yeast cells. Reprinted with permission from [63] copyright (2018) Springer Science+Business Media, LLC, part of Springer Nature. F array of optically induced (or virtual) electrodes, projected on a photoconductive layer over an ITO surface used to manipulate the trajectory of a
8
cancer cell cluster. Reprinted with permission from [29] copyright (2017) Elsevier
B.V. ... 28
Figure 6. Insulator-based devices examples. A DEP Trapping with optimized (i) diamond-shaped posts, (ii) circular posts, and (iii) square posts. Reprinted with permission from [98], copyright (2015) Elsevier. B DEP trapping in a device with (i) 21 columns of posts and (ii) 1 column of posts. Reprinted with permission from [1], copyright (2018) American Chemical Society. C Trapped cells at 300 Vrms at 30 kHz in (i) a device with 100 µm diameter insulating posts showing pearl formation and (ii) in a device with 20 µm posts where only one or two cells can trap on each post. Reprinted with permission from [103], copyright (2016) AIP Publishing LLC. D Separation of oil droplets by size: (i) separation of 25 µm and 50 µm diameter oil droplets, (ii) separation of 50 µm and 75 µm droplets. Reprinted with permission from [33], copyright (2017) Elsevier. E (3D)rDEP device schematic and inset showing 5 µm particle trapping (enclosed in red) at reservoir-channel junction at an applied voltage of 25 VDC +375 V at 1kHz. Reprinted from [32], copyright (2018) MDPI. F g-iDEP device (i) schematic of the sawtooth geometry of the device with the region observed enclosed, RBCs trapped in g-iDEP device at 400 V with (ii) MyO protein solution and (iii) H-FABP protein solution. Reprinted with permission from [34], copyright (2017) Springer. ... 39
Figure 7. Examples of hybrid microdevices. A OπDEP device used to separate polystyrene beads from E. coli cells. Reprinted with permission from [38] copyright (2013) the authors. B EπDEP device used to capture E. coli and S. aureus cells. Reprinted with permission from [24] copyright (2013) Springer-Verlag Berlin Heidelberg. C DEP and hydrophoretic sorter device used to separate 3 and 10 μm polystyrene particles. Reprinted with permission from [113] copyright (2014) WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim. D Device combining eDEP and inertia to separate 5 and 13 μm polystyrene particles. Reprinted with permission from [25] copyright (2018) Elsevier B.V. E eDEP+iDEP device used to manipulate MCF-7 cells. Reproduced from [14]. ... 44
Figure 8. AutoCAD mask designs for geometries B1 and B7. ... 51
Figure 9. Printed mask. ... 52
Figure 10. PDMS device molding. ... 53
Figure 11. Two PDMS replicas inside the plasma cleaner chamber. ... 54
Figure 12. Finished PDMS device. ... 54
Figure 13. Experimental setup for particle trapping. ... 55
Figure 14. 20/200 µm geometry. Sample of measured channel device... 57
Figure 15. Particles trapped at channel collapse (white curves added to delineate collapse section). ... 59
9
Figure 16. Channel with impurities at the base due to use of suboptimal materials and improper manipulation during the fabrication process. ... 59 Figure 17. Total trapping occurred at 800 V in the 200/2000 µm geometry. ... 63 Figure 18. Geometry 20/200 µm. At 1050 V, some particles exhibit reverse flow and vortices have formed in the space between each post and the channel wall. ... 63 Figure 19. Polystyrene particles of 2.0 µm channel B1 trial 1-Total trapping in the geometry with gap 40 µm and post diameter 2000 µm was observed at approximately 300 V. ... 65 Figure 20. Polystyrene particles of 2.0 µm channel B7 trial 3- Total trapping in the geometry with gap 30 µm and triangle base 600 µm was observed approximately at 250 V. ... 65 Figure 21. Polystyrene particles of 6.8 µm design B1 trial 6- Total trapping in the geometry with gap 40 µm and post diameter 2000 µm was observed at approximately 275 V. ... 66 Figure 22. Polystyrene particles of 6.8 µm, design B7 trial 2- Total trapping in the geometry with gap 30 µm and triangle base 600 µm was observed at approximately 125 V. ... 67 Figure 23. 3D printed phone holder for video capture in microfabrication lab. ... 67
10
List of Tables
Table 1. Stimulation voltage requirements in current electrode-based devices ... 29 Table 2. Stimulation voltage requirements in current insulator-based devices .... 40 Table 3. Stimulation voltage requirements in current hybrid devices ... 45 Table 4. Bicylindrical post geometries used before optimization process ... 49 Table 5. Bicylindrical post geometries selected after optimization [25]. ... 50 Table 6. Two-post semi-triangular geometries selected after optimization [25]. ... 50 Table 7. Fabricated dimensions for devices fabricated before geometry optimization. ... 57 Table 8. Comparison of target design dimensions and fabricated device dimensions. ... 58 Table 9. Bicylindrical post geometries selected after optimization [25]. ... 60 Table 10. Two-post semi-triangular geometries selected after optimization [25]. . 60 Table 11. EK regimes for a particle of size 1.0 µm and an EEEC of 942.3 ± 77.3 V/cm determined from Cardenas-Benitez et al. [2]. ... 62 Table 12. Trapping experiments results for one cylindrical geometry (B1) and one triangular (B7) for 2.0 µm polystyrene particles. ... 64 Table 13. Trapping experiments results for one cylindrical geometry (B1) and one triangular (B7) for 6.8 µm polystyrene particles. ... 66
11
Contents
Abstract ... 6
List of Figures ... 7
List of Tables ... 10
Chapter 1: Introduction ... 12
Introduction ... 12
General Objective ... 16
Specific Objectives ... 16
Hypothesis ... 16
Chapter 2: Insulator-based DEP and EK Theory ... 17
Chapter 3: Background and State of the Art ... 20
Dielectrophoretic systems ... 20
Voltage requirements in eDEP devices ... 23
Voltage reduction strategies in insulator-based devices ... 33
Hybrid DEP systems ... 43
Concluding remarks and future perspectives ... 47
Chapter 4: Materials and Experimental Methodology ... 49
Channel Geometries Before Optimization ... 49
Optimized Geometries ... 49
Fabrication ... 50
Experimental Setup ... 54
Experimental Procedure ... 55
Chapter 5: Results and Discussion ... 57
Fabrication Results ... 57
PIV High Voltage Experiments Results... 60
Particle Trapping Experiments Results ... 61
Chapter 6: Conclusions and Future Work ... 68
Appendix A: Abbreviations and acronyms ... 70
Bibliography ... 72
Published papers ... 83
12
Chapter 1: Introduction
Introduction
Microfluidics is the science and technology of fluid manipulation in channels with dimensions in the micrometer scale [3]. Although the definition of microfluidics does not constrain the overall size of a microfluidic device, having a small portable device to perform different analyses is critical to the field of application of these devices. One of the main applications of this technology is in the development of Point-of-care (POC) devices [4]. Technically, the small scale of microfluidic channels allows to harness the effects of different phenomena (e.g., electrokinetics) for the mixing and pumping of fluid or for particle manipulation. In practice, these devices offer advantages such as high- throughput, small sample volume, low processing time per analysis, low reactant consumption, portability, high test resolution for sensing applications, and not requiring specialized personnel for their operation. These characteristics result in inexpensive and quick diagnostic tests that can make healthcare accessible to more communities [5].
Electrokinetics is the general term used to describe the effects of an electric field on material (i.e. fluids and particles suspended in those fluids) movement [6]. This term describes a wide range of phenomena, including electrophoresis (EP), electroosmosis (EO), and dielectrophoresis (DEP). EP and EO exist in the presence of any electric field.
However, DEP exists only in the presence of spatially non-uniform electric fields. While EP and EO act on charged particles and ions in fluid, respectively, DEP acts on polarizable particles. When a polarizable sphere is placed in a non-uniform electric field, an effective electric dipole aligned with the field is induced, as illustrated in Figure 1A.
Since the magnitude of the field is greater in one side of the sphere than the other, a net DEP force will cause a displacement of the particle [7].
The first work reporting dielectrophoresis was done by Herbert Pohl in 1951 [8], when he observed the movement of particles suspended in fluid as a result of the forces induced by a non-uniform electric field. The initial work of Pohl utilized a setup with a tin-foil-made circular outer electrode and a tungsten wire inner electrode. It was until later, in 1989, that Masuda et al. introduced a different DEP modality [9]. They used insulating structures made of silicone rubber to create the non-uniform field required for cell trapping at the channel constriction. These two works did set the basis for the two main approaches used in DEP-based microfluidics: electrode-based dielectrophoresis (eDEP) and insulator- based dielectrophoresis (iDEP). Figure 1B-C shows the electric field distortion in an electrode-based device and an insulator-based device, respectively.
To this day, DEP is one of the most widely used driving forces for particle (e.g., cells, proteins, DNA, virus, etc.) manipulation in microfluidic devices. An advantage of DEP is that it depends on the intrinsic electric properties of the particles and fluids, which allows for particle separation without the need for labelling. It is also a versatile tool as the options for patterning and channel designs are only limited by the constraints of the fabrication equipment available [10]. Some of the specific applications of DEP include particle and
13
cell trapping and characterization, biological extraction operations such as the extraction of plasma from blood in both insulator-based (Figure 2A) and electrode-based devices (Figure 2B); cell patterning for tissue engineering which uses different electrode geometries such as the hexagonal geometry shown in Figure 2C; sorting of DNA by size (Figure 2D); separation of live and dead yeast cells(Figure 2E) an particle separation by size (Figure 2F); single-cell 3D manipulation; and affordable and portable diagnostic tools [11].
Because of its many attractive features, some commercial laboratory benchtop systems based on DEP have been developed recently. Among these we can list the 3DEP system produced by DEPtech to measure the resistance and capacitance of cells, the bacterial counter developed by Panasonic applying dielectrophoretic impedance measurements (DEPIM), Menarini-Silicon Biosystems DEPArray to trap cells in DEP cages for their identification and recovery, or the ApoStream instrument commercialized by Precision for Medicine to isolate circulating tumor cells (CTC). Although these instruments have a reasonable size for laboratory benchtop equipment (comparable to a benchtop centrifuge or a desktop PC), they are still not fully portable devices.
One of the most critical aspects in designing a functional POC device is the voltage required for its operation. Although it is the electric field that drives particle manipulation in DEP-based devices, the input voltage generates that field. A high voltage requirement implies more sophisticated electronic instrumentation, therefore negatively affecting portability. Hence, it is of utmost importance to devise strategies to reduce input voltage requirements. Recently, a portable power supply with an output in the range of 200 V has been developed [12]. However, a voltage requirement of thousands of volts still prevents the miniaturization of the voltage source [13]. Moreover, it has been reported that high voltages in microfluidic channels filled with conductive solutions generate significant Joule heating [14], [15]. This can cause temperatures to rise above the optimal levels for cell viability and generate electrothermal flow (ETF). We note that, although there have been efforts to effectively use ETF for particle/cell manipulation [15], it remains generally viewed as a negative effect for cell trapping. Also, the application of high voltages, independently of the heating effects they produce, may damage the biological sample, decreasing cell viability [16]. Thus, high input-voltages represent an obstacle towards creating fully integrated, stand-alone platforms that can be used outside of a laboratory.
14
Figure 1. DEP concept. A Alignment of an effective electric dipole moment in the presence of a non-uniform electric field. B Electric field distortion in an electrode-based device and its effect on a particle that is more polarizable than its suspending fluid medium (positive DEP – pDEP). C Electric field streamlines and their distortion in a device with insulating posts, resulting in higher electric fields at the post constrictions.
In this contribution, we review the efforts made during recent years to lower the voltage requirement for DEP-based devices. The different approaches to lower this operational requirement depend on the type of device and application being considered, with control parameters ranging from application of electrical stimulation directly on the channel or outside, the channel and electrode materials being used, the geometry and dimensions of insulating structures and electrodes. As it will be discussed in the next section, several DEP-variants branch out from the two main types (i.e. eDEP and iDEP), depending on the device feature creating the electric field non-uniformity. In this review we will group them as electrode based, insulator based, and hybrid devices.
15
Figure 2. Different reported applications for DEP based microdevices. A iDEP device used for the extraction of plasma from blood. Reprinted with permission from [17] copyright (2015) Springer-Verlag Berlin Heidelberg. B eDEP device used for the extraction of plasma from blood.
Reprinted with permission from [18] copyright (2015) AIP Publishing LLC. C eDEP device used to pattern cells into a hexagonal matrix for skin tissue engineering. Reprinted with permission from [19] copyright (2019) Springer-Verlag London Ltd., part of Springer Nature. D DNA sorting in an iDEP device. Here, DNA molecules of different size are sorted and redirected to different outlets to posterior collection. Reprinted with permission from [20] copyright (2016) American Chemical Society. E eDEP device used to separate particles with different properties; in this case live and dead yeast cells. Reprinted with permission from [21] copyright (2019) American Chemical Society. F contactless dielectrophoresis device used to sort three different particles
16
based on their size. Reprinted with permission from [22] copyright (2016) Chinese Mechanical Engineering Society and Springer-Verlag Berlin Heidelberg.
General Objective
The objective of this work is to minimize the voltage requirement for particle manipulation and trapping in an insulator-based microfluidic channel. The motivation for this work is to integrate it with modeling to develop a process for voltage requirement reduction in microfluidic devices in order to make the design of these devices a more objective, rational process in which the specific needs of the application can be taken into account while defining the channel geometry.
Specific Objectives
• To complete modelling to identify the optimal geometry in each of two stages:
using bicylindrical stages and semi triangular posts.
• To fabricate the channel geometries with the precise dimensions obtained from the optimization process.
• To observe first trapping, total trapping, and reverse flow of polystyrene particles in a microfluidic channel and determine the specific voltages for each of these stages.
Hypothesis
The hypothesis of this work is that through constriction geometry optimization starting with a bicylindrical geometry design it is possible to lower the voltage requirement to perform particle trapping in a PDMS microfluidic device. Two insulating structure geometries will be analyzed and tested for polystyrene bead trapping. By changing the insulating feature dimensions such as post size and gap between posts, there will be an optimal set of dimensions that will amplify the electric field at the center of the channel to a maximum point, minimizing the voltage required to be applied across the channel in order to achieve particle displacement and trapping. This will bring us one step closer to the development of portable, electrokinetically-driven, lab-on-a-chip and point of care devices.
17
Chapter 2: Insulator-based DEP and EK Theory
Dielectrophoresis is the phenomenon that creates a net force on a polarizable particle in the presence of a non-uniform electric field [23]. The particle can be either attracted (when the polarizability of the particle is larger than that of the suspending solution) or repelled (when the polarizability of the particle is smaller than that of the suspending solution) from the region of strong electric field non-uniformity, leading to either positive dielectrophoresis (pDEP) or negative dielectrophoresis (nDEP), respectively. The dielectrophoretic force on a spherical particle is expressed as
𝐅𝐃𝐄𝐏= 𝟐𝝅𝒂𝟑𝜺𝒎𝐑𝐞[𝑲]𝛁(𝐄 ∙ 𝐄) (1) where 𝒂 represents particle radius, 𝛆𝐦 is the medium permittivity, 𝐑𝐞[𝑲] is the real part of the complex Clausius-Mossotti (CM) factor, 𝐄 = −𝛁𝝋 is the electric field vector, and 𝝋 is the electric potential due to the input voltage. The CM factor, 𝑲, is defined by
𝑲 = 𝜺𝒑
∗−𝜺𝒎∗
𝜺𝒑∗+𝟐𝜺𝒎∗ (2) with ε* = ε - jσ/ω. The subscripts p and m refer to particle and medium, respectively, j is the imaginary unit, σ is the conductivity and ω is the angular frequency of the voltage signal. In iDEP theory, the condition for trapping of particles was modelled in the following equation:
𝝁𝐄𝐊𝐄 ⋅ 𝐄 + 𝐂𝝁𝐃𝐄𝐏𝛁(𝐄 ⋅ 𝐄) ⋅ 𝐄 = 𝟎 (3) where 𝝁𝐄𝐊 and 𝝁𝐃𝐄𝐏 are the electrokinetic mobility and dielectrophoretic mobility, respectively, and 𝐂 is an empirical correction factor that is necessary to account for the behavior observed in experiments. Correction factors as high as 600 have been used in previous works to balance the DEP contribution and the EK (EP, EOF) contribution to trapping [24].
Regardless of the need for this correction, DEP was considered the dominant phenomena taking place in the channel considering that the field nonuniformity was necessary in order to observe particle trapping. It has been reported, however, that particle flow reaches a point of equilibrium and even reverses in a channel without posts, that is, with a uniform electric field [2]. This finding renders the DEP contribution negligible in these devices. In the model applied in this work, there is a linear and a nonlinear electrophoretic component. The addition of these two components gives the electrophoretic term in the general equation for particle velocity. The total particle velocity is then a sum of the electrophoretic and electroosmotic velocity components. Figure 3 illustrates an example of this. Suppose a negatively charged microsphere is suspended in fluid with a lower conductivity than that of the particle. The plasma treated PDMS channel walls are
18
negatively charged. As a conductive solution fills the channel, positive charges will line up along the channel wall and on the negatively charged particle surface. An electrical double layer (EDL) forms at the interface between the particle and the fluid, with the innermost charges being positive and the outermost layer having both positive and negative charges interchanged. When an electric field is applied across the channel, the EDL deforms as shown on the diagram, until EP and EOF reach equilibrium, and if a higher voltage is applied it results in net particle movement in the direction opposite to the field direction. The electrophoretic effect is opposing electroosmotic flow, while EOF continues to be in the direction of the electric field.
Figure 3. Diagram of electrokinetic phenomena experienced by a negatively charged particle in a channel with an applied uniform electric field.
In this new interpretation of insulator-based phenomena in a microfluidic channel presented by Cardenas-Benitez et. al [2], the equation modeling electrophoretic velocity is:
𝒖𝐄𝐏= 𝝁𝐄𝐏(𝟏)𝑬𝟎+ 𝝁𝐄𝐏(𝟑)𝑬𝟎𝟑 (4) Where 𝑬𝟎 is the uniform electric field applied across the channel while 𝝁𝐄𝐏(𝟏) and 𝝁𝐄𝐏(𝟑) are the first and third-order electrophoretic mobilities, respectively, defined by equations 5 and 6:
𝝁𝐄𝐏(𝟏)= −𝝐𝒎𝛟𝐓
𝜼 (𝜻𝟎+𝐃𝐮∙𝐥𝐧(𝟏𝟔)
𝟏+𝟐𝐃𝐮 ) (5) 𝝁𝐄𝐏(𝟑)= −𝒂𝟐𝝐𝒎
𝜼𝝓𝑻 𝒇(𝑫𝒖, 𝜻𝟎, 𝜶, 𝜶̀) (6) where 𝝐𝒎 is the medium permittivity, 𝝓𝑻 is the thermal voltage, 𝜼 is the medium viscosity 𝜻𝟎 is the zeta potential of the particle, Du is the Dukhin number, a is the radius of a particle suspended in the medium and 𝜶, 𝜶̀ are dimensionless coefficients related to ionic drag
19
coefficients. The function 𝒇(𝑫𝒖, 𝜻𝟎, 𝜶, 𝜶̀) in eq. 6 is nonlinear and depends on the electrokinetic properties of the particle. Under this model, electroosmotic flow is assumed to be linear and described by equation 7:
𝒖𝐄𝐎= −𝝐𝒎𝜻𝒘
𝜼 𝑬𝟎 (7) where 𝜻𝒘 is the zeta potential of the PDMS channel wall. Total particle velocity is then described as the sum of the electrophoretic velocity and the electroosmotic velocity as shown on equation 8.
𝒖𝐩= 𝒖𝐄𝐏+ 𝒖𝐄𝐎 (8) Substituting the variable definitions into equation 8, setting 𝒖𝐩 equal to zero and solving for 𝑬𝟎 gives the electric field condition necessary for electrokinetic equilibrium
constrained to first trapping of particles. The electrokinetic equilibrium condition 𝑬𝐄𝐄𝐂 at which 𝒖𝐩= 𝟎 is defined as follows:
𝑬𝐄𝐄𝐂= √−𝝁𝐄𝐏
(𝟏)+𝝁𝐄𝐎
𝝁𝐄𝐏(𝟑) (9) This condition allows to predict the expected particle trapping field in channels with or without posts. When a channel has posts, the field applied (which is uniform in the case of a plain channel), is amplified at the constriction between the posts. To calculate the amplification factor for a specific channel geometry design, the electric field at the constriction of a channel with posts (𝑬𝐯) is divided by the uniform electric field value that would be present in a flat channel (𝑬𝟎) according to eq. 10:
𝐀𝐦𝐩𝐥𝐢𝐟𝐢𝐜𝐚𝐭𝐢𝐨𝐧 𝐟𝐚𝐜𝐭𝐨𝐫 = 𝑬𝐯
𝑬𝟎 (10) Finally, to obtain a voltage trapping prediction (𝑬𝐭𝐫𝐚𝐩), the 𝑬𝐄𝐄𝐂 is divided by the
amplification factor (AF):
𝑬𝐭𝐫𝐚𝐩= 𝑬𝐄𝐄𝐂
𝐀𝐅 (11) The ratios obtained in equations 10 and 11 are utilized in the next section to provide predictions for the experimental trapping results in our device geometries. A more detailed explanation of the modeling and optimization approach used in this work is discussed in a master’s thesis presented by De los Santos Ramírez [25].
20
Chapter 3: Background and State of the Art
Dielectrophoretically-driven microfluidic devices have demonstrated great applicability in biomedical engineering, diagnostic medicine, and biological research. One of the potential fields of application for this technology is in Point-of-Care (POC) devices [4], ideally allowing for portable, fully integrated, easy to use, low cost diagnostic platforms.
Two main approaches exist to induce Dielectrophoresis (DEP) on suspended particles, i.e., electrode-based DEP and insulator-based DEP, each featuring different advantages and disadvantages. However, a shared concern lies in the input voltage used to generate the electric field necessary for DEP to take place. Therefore, input voltage can determine portability of a microfluidic device. This chapter outlines the recent advances in reducing stimulation voltage requirements in DEP-driven microfluidics.
Dielectrophoretic systems
Dielectrophoretic systems can be grouped into two big categories: 1) electrode based dielectrophoresis (eDEP), and 2) insulator based dielectrophoresis (iDEP). However, we introduce category 3) hybrid DEP, for systems that are variations and/or combinations of either eDEP or iDEP with themselves or another manipulation method [26]–[28].
In eDEP, electrode arrays are used to create a non-uniform electric field distribution.
These arrays are commonly designed as interdigitated [29], castellated electrodes [30] or quadrupole electrodes [31] and the preferred materials are usually metals (e.g., gold).
However, they are relatively expensive to produce, and prone to electrode fouling and react with the medium they are in contact with [10]. Recently, optical (or virtual) electrodes—created when a beam of light is projected onto a photoconductive surface—
have been used in eDEP systems [32]–[34].
In iDEP, insulator structures are used to create the non-uniform electric field distribution.
These insulating structures can be as varied in geometry as one wishes. During the last two decades, iDEP techniques have proliferated and evolved, introducing many subclassifications such as reservoir based dielectrophoresis (rDEP) [35], wall induced dielectrophoresis (w-iDEP) [36], gradient insulator based dielectrophoresis (g-iDEP) [37], curvature induced dielectrophoresis (c-iDEP) [38], contactless dielectrophoresis (cDEP) [39] and others that will surely arise in the future. The main advantage of iDEP is the simplicity to design and fabricate low-cost devices [10]. Nonetheless, current iDEP systems require high input-voltages to operate, which in some cases can reach thousands of volts [40]. This requirement has therefore made it increasingly difficult to integrate iDEP systems into POC devices. Figure 4 shows general eDEP and iDEP fabrication processes.
Finally, there are the hybrid DEP systems, which are variations and/or combinations of eDEP and iDEP with themselves or another particle manipulation method. Some examples are: off-chip passivated-electrode, insulator based dielectrophoresis (OπDEP)
21
[41]; embedded passivated-electrode, insulator based dielectrophoresis (EπDEP) [26];
DEP+gravity [28]; DEP+inertial forces [27]. These hybrid systems are designed to take the best properties of each technique and utilize them to improve the overall performance of the devices.
It is important to note that using DEP for effective particle manipulation usually implies generating high magnitude electric fields. Given the wide range of applications of DEP in cell manipulation, there have been studies aiming at reducing the damage of high fields to cells by limiting the time of stimulation [42], reducing the sources of excitation in multi- purpose platforms (e.g. combined trapping and electroporation) [43], and thoroughly characterizing thermal and conductivity effects due to high fields to identify irreversible electroporation effects [44].
22
Figure 4. Fabrication process for dielectrophoretic microdevices. A Fabrication process for a carbon electrode-based device using photolithography and pyrolysis. Reprinted from [45],
23
copyright (2016) AIP Publishing. B Standard fabrication process using photolithography (i-iv) and soft lithography (v-vi) to produce an insulator-based device. Reprinted from [46], copyright (2007) Nature Publishing Group.
Voltage requirements in eDEP devices
Electrode based dielectrophoresis devices are commonly known to operate at what would be considered as low voltages, making this their principal advantage. Additionally, eDEP devices are capable of producing strong electric fields and, consequently, significant non- uniformities of which the dielectrophoretic force depends on. These features are mainly due to the short distances that exist between electrodes. Therefore, since the electric stimulation requirements are low in eDEP based systems, just a few works have aimed at further reducing the input voltage. Nonetheless, eDEP brings disadvantages that include electrode fouling, possible sample contamination, and high fabrication costs [47], [48]. Figure 5 shows examples of eDEP devices.
2D electrodes
In eDEP devices, flat electrodes are fabricated on the floor or ceiling of the channel. This kind of electrode design is known as 2D electrodes and it is the most widely used approach to DEP. Figure 5B-E show examples of 2D electrodes.
In 2015, polystyrene beads (20 μm) and yeast cells (5-6 μm) were captured using a castellated carbon electrodes device [30]. Two electrode designs were used, off-set and non-off-set with an electrode gap between 100-150 μm. Two AC signals were used, 20 Vpp and 10 Vpp both at 100 kHz. 20 Vpp were applied to study polystyrene beads behavior, exhibiting nDEP while 10 Vpp were used for yeast cells, which showed pDEP.
The latter allowed for separation of a mixture of polystyrene beads and yeast cells with separation efficiencies up to 96% with a flow rate of 0.1 ml/h. The same year, Song et al.
[49] conducted a study focused on separation of stem cells from differentiation products (osteoblasts). Stem cells are of interest for regenerative medicine due to their ability to differentiate into other cells with specialized functions. The device consisted of an array of gold interdigitated electrodes with a 45° inclination angle. It was able to continuously separate human stem cells from osteoblasts. At a flow rate of 1.8 μl/min, an AC signal of 7.2 Vpp oscillating at 3 MHz was used, achieving a collection efficiency of 92% for stem cells and 61% for osteoblasts. However, when increasing the flow rate to 5.4 μl/min, the voltage had to be raised to 15.4 Vpp (same frequency) to achieve collection efficiencies up to 88% and 69% for stem cells and osteoblasts, respectively. We stress that the voltage was applied in an on-off switching mode to allow particles to flow, instead of getting trapped within electrodes.
The next year, a tunnel dielectrophoresis device was developed [50]. The purpose of this device was to focus particles, in high-speed flows, into a single focal stream. The device
24
achieved that by using an 80 μm wide, 83 μm high and 6 cm long microchannel. This length is necessary to give particles enough time to migrate to the focal stream. When viewed from a cross-section, electrodes are placed on the four corners of the microchannel and they continue along the channel. Such configuration allows, when particles experience nDEP, to only have one focal stream. Polystyrene beads (9, 15 and 20 μm) were successfully focused when applying 10 Vpp at 1 MHz at an average flow speed of 5 cm/s. Similar experiments were conducted on HeLa cells. Here, AC signals were 15.4 Vpp and 13.8 Vpp, both oscillating at 10 MHz and flow speed was set to 8.7 cm/s and 11 cm/s, respectively. The viability of HeLa cells was estimated at 85.3% after experiments. That year, a device that intended to trap bioparticles when bounded to polystyrene beads (5 μm) was studied [29]. An interdigitated electrode design was used where fingers of each comb had different widths (20 μm and 80 μm, respectively) with a 20 μm inter-electrode gap. First, a 10 V and 5 kHz signal was used, showing no particle trapping effect. Once the voltage was increased to 15 V (same frequency) polystyrene beads were successfully trapped on the edge of slim electrodes. After that, the frequency was increased to 10 MHz, showing a reduction in the trapping effect and highlighting the importance of careful electrode design and voltage selection.
In 2017, Barik el at. [51] were able to trap DNA molecules and particles at the nanoscale using voltages with peak amplitude as low as 0.45 V. Polystyrene beads (190 nm) were trapped by pDEP with a peak amplitude of 0.75 V oscillating at 1 MHz. When increased to 10 MHz, however, particles were released due to nDEP. Nanodiamonds (40 nm) were also used, first trapping was observed at 0.45 V and 1MHz. However, 2 V at 100 kHz were the conditions where most particles were trapped. Finally, 10 kbp and 500 bp DNA molecules were successfully captured at maximum voltages of 3 and 2.5 V, respectively, both at 1 MHz. All this was achievable due to the device’s design. It consisted of an interdigitated electrode array, which incorporates an atom-thin graphene sheet. This graphene sheet is connected to the electrodes, producing extremely large electric field non-uniformities at low voltages.
The following year, a device capable of manipulating RBCs (5 μm) along both channel height and width (3D switching) using voltages between 1 and 10 Vpp was reported [52].
To achieve 3D switching, two independent sets (top and bottom) of interdigitated electrodes were used. Each featuring 50 μm wide electrodes with 50 μm inter-electrode gaps. When applying 1 Vpp to the top set and 10 Vpp to the bottom set, RBCs were deflected toward the top set of electrodes. When swapping voltages, RBCs were deflected toward the bottom set. Finally, when voltages were 10 Vpp in both sets, RBCs flowed in the center of the channel. In all cases, the frequency to induce nDEP on RBCs was 10 kHz and a flow rate of 5 μl/h was used.
In 2019, Kim et at. [53] reported a DEP device capable of detecting Alzheimer’s disease biomarkers, Aβ42 and tau-441, by capturing them using voltages between 0.5 and 0.9 Vpp. Two designs were used. The first design was an interdigitated electrode array with 10 μm gaps. It was able to capture Aβ42 with 0.8 Vpp and tau-441 with 0.9 Vpp. The
25
second design incorporated patterned square (3 μm) microstructures within the electrode gaps in the first design. When including the microstructures, the electric field gradient increased, which translates into a lower capturing voltage. With the second design, Aβ42 and tau-441 were captured with 0.5 Vpp and 0.6 Vpp, respectively. A frequency of 50 MHz was used in all cases to induce nDEP on the target particles.
3D electrodes
3D electrodes aim to make the DEP force effective in the entire channel volume, lowering the required input voltage for affecting particles that would otherwise be located far away from the 2D electrode set (Figure 5A). The most common 3D electrode fabrication method is photolithography + pyrolysis [10]. This technique uses a polymer, usually SU-8, as a base material to produce carbon electrodes [54].
In 2015, Puttaswamy et al. [55] developed a 3D interdigitated electrode DEP device. The electrodes (200 μm wide and 40 μm tall) were fabricated from a mixture of silver conductive adhesive and carbon nanopowder. Polystyrene beads (10 μm) were used to test the device by directing them to three outlets. When no voltage was applied, beads headed to the first outlet. After applying a 40 Vpp at 5 MHz signal, particles experienced nDEP and were deflected to the second outlet. Increasing the voltage further to 60 Vpp (same frequency), particles deflected toward the third outlet. A flow rate of 0.2 μl/min was used to introduce the particles into the channel and a flow rate of 0.6 μl/min for the sheath flow.
A couple of years later, a 3D carbon electrode device was used to study separation of live and dead U937 (human myeloid leukemia) monocytes (23 μm [live] and 22 μm [dead]) under lower than 40 Vpp stimulation voltages [56]. 3D carbon electrodes were fabricated via pyrolysis and their geometry consisted of 100 μm tall cylinders with 50 μm in diameter, placed over an array of interdigitated electrodes. Separation of live and dead cells was achieved with an AC signal of 20 Vpp at 300 kHz and a 1 μl/min flow. At that frequency, live cells experienced pDEP while dead cells showed no response. After separation, 90%
of dead cell were successfully removed from the mixture.
Optically induced (virtual) electrodes
Virtual electrodes (VE) can be defined as electrodes that manifest when a light beam is projected onto a photoconductive surface (Figure 5F). The method reported in [32]–[34], [57] consists in using ITO slides as electrodes in the ceiling and floor of the microchannel.
One of the ITO slides was covered by a photoconductive layer. Therefore, when light meets this layer its impedance is reduced, creating the VE and generating a non-uniform electric field. With this technique, a great degree of flexibility is achieved in the geometry of the electrodes (even making them dynamic) and therefore, it provides enhanced particle manipulation capabilities.
26
In 2016, Chiu et al. [34] used an ODEP device to isolate circulating tumor cells (CTCs, 23±2.1 μm). The device consisted of a T-shaped microchannel composed of two ITO glass slides as floor and ceiling, one of which was coated with a photoconductive layer.
The isolation area was focused on the T intersection, where CTCs were separated from leucocytes. 8 V at 100 kHz were applied to the ITO slides, the CTCs were visually identified, and the image of a ring was projected on the photoconductive layer, creating a non-uniform electric field, trapping CTCs inside the ring. Afterwards, a light bar was projected on the T intersection to remove non-target particles. Then, CTCs were moved to a collection area. Next year, a similar device was used to trap and isolate CTCs (18.3±3 μm) from a mixture with leucocytes (8.4±1.2 μm) [33]. However, it was possible to achieve particle trapping with an AC signal of 5 V at 100 kHz. A light bar width of 40 μm was chosen for producing the maximum velocities when moving CTCs and leucocytes, around 148 and 63.8 μm/s, respectively. Up to 100% cell purity was achieved in this experiment through the presence of multiple trapping zones along the microchannel, which allowed reducing the trapping voltage. One year later, Chiu et al. [32] developed another ODEP device to manipulate CTC clusters (individual CTCs 14-25 μm). In this case the VE pattern was a matrix conformed of square elements. The optimal conditions for CTCs cluster manipulation were experimentally determined to be a matrix element of 100 μm by 100 μm, flow rate of 0.5 μl/min, measured velocity of 110 μm/s and an AC signal of 5 V at 100 kHz. We note the voltage requirement did not change from the previous work.
However, an improvement in the degree of particle manipulation took place, especially due to the dynamic properties of VE—something that is not possible with conventional electrode technologies, either 2D or 3D.
Liquid metal electrodes
Gallium based liquid metal alloys have high electrical conductivity and have been used as electrodes [58]. This material presents low viscosity and high surface tension, allowing for fast fabrication and flexibility in the type of shapes that can be generated. One example of this work is the creation of liquid metal droplet shaped Galinstan electrodes (68.5%
gallium, 21.5% indium, and 10% tin) using dielectrophoresis. In 2015, Tang et al.
demonstrated trapping of 80 nm tungsten trioxide nanoparticles at 15 V and 1 MHz [59].
Another work used liquid metal electrode channels consisting of 67% Ga, 20.5% In, and 12.5% Sn to stretch red blood cells from 6 µm to 8 µm using voltages from 1 Vpp-10 Vpp at 1.5 MHz, providing evidence that this system can be used for cell characterization [60].
Electrohydrodynamic mixing was explored at 100 VDC using eutectic gallium indium liquid channel electrodes contained by PDMS posts. Liquid electrodes are in contact with the main channel through the openings between posts [61]. A microdroplet generation system with electrodes composed of 75.5% Ga and 24.5% In operating between 2600 V and 3000 V achieved the formation of 600 µm long droplets [62]. Studies such as these suggest that the use of liquid electrodes is promising for its versatility and ease of fabrication. Table 1 shows a list of the characteristics of eDEP devices.
27
28
Figure 5. Examples of electrode-based microdevices. A 3D electrode array (i) SEM image of the electrodes (ii)-(iv) show the device performing a separation of HCT116 (green) and lymphocytes (red). Reprinted with permission from [63] copyright (2018) The Royal Society of Chemistry. B eDEP device used to concentrate HeLa cells. Reprinted with permission from [64] copyright (2014) The Korean BioChip Society and Springer. C interdigitated electrode array in combination with a layer of atom-thin graphene to capture different particles such as polystyrene beads, nanodiamonds and DNA. Reproduced from [51]. D device with interdigitated-irregular electrodes used for the concentration and detection of CA 19-9, a pancreatic cancer biomarker. Reproduced from [65]. E serpentine-shaped carbon electrode used to separate live from dead yeast cells.
Reprinted with permission from [66] copyright (2018) Springer Science+Business Media, LLC, part of Springer Nature. F array of optically induced (or virtual) electrodes, projected on a photoconductive layer over an ITO surface used to manipulate the trajectory of a cancer cell cluster. Reprinted with permission from [32] copyright (2017) Elsevier B.V.
29 Table 1. Stimulation voltage requirements in current electrode-based devices
DEP type and
application Particle type+size Electrode material Electrode geometry Electrode
dimensions Voltage
eDEP, cell
stretching NB4 leukemia cells Ø
14.04 μm ITO Two rectangular
electrodes 20 µm gap AC: 2-9 Vpp, 1 MHz
[67]
eDEP, patterning Yeast cells Ø2.15 μm PCB standard 16 concentric circles 700 µm surrounding circle Ø; 400 µm center Ø; 100 µm gap
AC: 2-10 V*, 6 MHz [68]
eDEP, trapping dsDNA 10 kbp Quartz with gold coating Nanopipette O.D. 1.0 mm, I.D. 0.5
mm AC: 10-20 Vpp, 0.5-4
MHz [69]
eDEP, patterning Yeast cells PCB (copper) 16 concentric circles 700 µm surrounding circle Ø; 400 µm center Ø; 100 µm gap
AC: 60 Vpp, 6 MHz [70]
eDEP, trapping Polystyrene
beads Ø 5 μm Cr-Au Interdigitated ratios between +/-
20:20 µm, 20:40 µm, 20:60 µm, 20:80 µm
AC: 10-15 V*, 5 kHz- 10 MHz
[29]
eDEP, separation Polystyrene beads Ø
5-10 μm Au Interdigitated 40 μm electrode
length AC: 5.71 Vpp, 60 kHz
[71]
eDEP, separation Polystyrene (10 μm) and white blood cells (12.4 μm)
Au-Cr Interdigitated ** AC: 7.5 Vpp, 800
MHz [72]
eDEP, separation Polystyrene (5 μm and 15 μm), yeast cells
Ti-Au Triangle H=1213 μm, Cx=1153
μm & Cy=375 μm AC: 8 Vpp, 16 Vpp, 24 Vpp, 100 kHz (PS) and 1050 kHz (yeast) [21]
eDEP, separation Polystyrene (7 μm and 20 μm), yeast cells (5-6 μm)
Carbon paste Interdigitated, castellated 100 μm –150 μm AC: 10-20 Vpp, 30 kHz-3 MHz [30]
eDEP, sorting Gold nanoparticles
(10 nm) Gold Arrow-like (with rounded
point) 10 μm gap AC: 18-26 Vpp, 1
kHz-20 MHz [73]
eDEP, separation
‘liquid electrode’
HT-29 (13.2 μm), PLT (1.8 μm), T- Lymph (7 μm), RBC
Ti-Pt Rectangular, comb-like
structure length varied from 20
μm to 40 μm AC: 6-10 Vpp, 100 kHz [74]
30 (5 μm), MD 231 (12.4
eDEP, trappping μm) Polystyrene (190 nm), nanodiamond (40 nm), DNA (10 kbp, 500 bp)
Ti-Pd gate electrodes, covered by insulator and graphene electrode layer with Cr/ Au ohmic contacts and Cr/Al electrical leads.
Graphene sheet edge Gate and graphene electrode:
interdigitated with 5 μm thick fingers
AC: 0.9-6 Vpp, 1 MHz, 10 MHz [51]
eDEP, isolation HCT116 cells Doped silicon crystals 3D comb array with
electrode digits consisting of castellated blocks
15 pairs of electrode digits, 10 identical pores per digit, each pore 60 μm in diameter
AC: 10-25 Vpp, 60- 400 kHz
[63]
eDEP, patterning MC3T3-E1 bone cells
(7.5 μm) Stainless steel Honey-comb multilayer L:140 μm W:50 μm, mean pore size:140 μm. Thickness: 50 μm
AC: 10 V*, 500 kHz [75]
eDEP, trapping Polystyrene (1 μm) Carbon 3D posts 100 μm height, 50 μm
diameter AC: 28 Vpp, 5-50 kHz [76]
eDEP,
concentration Polystyrene (1 μm), E. coli, MS2 virus vs Troponin I
ITO Coplanar electrodes 100 nm thick ITO layer
and 25 μm gap between electrodes.
AC: 5 Vpp, 10 Vpp, 10 kHz, 100 kHz, 2 MHz [77]
eDEP,
concentration Yeast cells (5 μm) Glass-like carbon Cylindrical 50 μm diameter, 100 μm height, center to center 115 μm
AC: 20 Vpp, 100 kHz [45]
eDEP,
characterization Candida strains - albicans, parapsilosis, tropicalis (5-7 μm)
Carbon Cylindrical 50 μm diameter, 100
μm height, distance between posts 58 μm in all directions
AC: 20 Vpp, 10 kHz - 1 MHz [78]
eDEP, isolation AS2-GFP Au (ground electrode) and
ITO (voltage electrode) Serpentine with “V” shaped
cross-section 700 μm width and 200
μm depth AC: 18 Vpp, 10 kHz [79]
eDEP, sorting Stem cells (hMSCs) Au-Cr Interdigitated 50 μm width and 50
μm gap AC: 7.2 Vpp, 3MHz
[49]
eDEP, detection
and quantification Biotin functionalized polystyrene beads (0.74 μm)
Au Interdigitated pearl shaped ** AC: 10 Vpp, 500 kHz
– 2 MHz [80]
eDEP, sorting RBCs Au-Cr Interdigitated (2 sets: top
and bottom) 50 μm width and 50
μm gap AC: 1 Vpp and 10
Vpp (switching sets), 10kHz (both) [52]
31 ODEP,
purification PC-3 ITO + photoconductive layer Rectangle and “O” shaped Rectangle width: 150, 200 and 250 μm; “O”
shaped: 40 μm
AC: 8 V*, 100kHz [34]
ODEP,
concentration 5-μm particles ITO + photoconductive layer Square lattice ** AC: 20 Vpp, 250 kHz [57]
ODEP, isolation
and purification PC-3 (18.3 ± 3 μm) ITO + photoconductive layer Interdigitated Width: 40 μm; gap: 80
μm AC: 5 V*, 100 kHz
[33]
ODEP, isolation H209 ITO + photoconductive layer Square matrix 100 x 100μm (by
matrix element) AC: 5 V*, 100 kHz [32]
eDEP, patterning HFF Stainless steel Hexagonal ** AC: 14-56 Vpp,
300kHz [19]
eDEP, focusing PS beads (9, 15, 20
μm), HeLa cells Au Parallel along channel ** AC: 10,13.8, 15.4
Vpp, 1-10MHz [50]
eDEP, sorting PS beads (10 μm) Ag-conductive adhesive and
carbon nanopowder Interdigitated 5000x100x40 μm AC: 40, 60 Vpp, 5MHz [55]
eDEP, separation RBCs Ti-Au Castellated and serpentine ** AC: 8 Vpp, 1MHz [18]
eDEP, counting
and detection Shewanella
oneidensis Au Interdigitated ** AC: 1, 3, 5 Vpp, 0.1-1
MHz [81]
eDEP,
concentration PS beads (20 and 40 nm), BSA
(14x4x4nm)
Au on ITO Cone in matrix array Cone: H=100nm, base
diameter= 100nm AC: 10 V*, 2.5 MHz [82]
eDEP,
concentration AuNP (150 nm) Au Planar 10 x 25 μm AC: 17 Vpp, 100kHz
[83]
eDEP, trapping Yeast Au and C Serpentine ** AC: 10 Vpp, 100kHz-
1MHZ [66]
eDEP, separation U937 monocyte cells C Interdigitated with 3D
structures on top 3D structures:100 μm height and 50 μm diameter
AC: 20 Vpp, 50kHz- 1MHz [56]
eDEP,
concentration Aβ42 and tau-441 Ta-Pt Interdigitated with
microstructures in between ** AC: 0.5-0.9 Vpp, 50MHz [53]
eDEP, detection Biotinylated DNA +
PS beads (750 nm) Au Irregular interdigitated ** AC: 10 Vpp, 0.5-
2MHz [84]
eDEP, separation K562 cells Cu Interdigitated Height: 25-30 μm,
Thickness: 30 μm, width: 40 μm
AC: 9 Vpp, 48.64 MHz [85]
eDEP, detection Biotinylated CA 19-9 antibody + PS beads (750 nm)
Au Irregular interdigitated ** AC: 10 Vpp, 0.5-2
MHz [65]
32 eDEP, separation PS beads (2, 5, 8
µm) Cr-Au Insulator microchannel
Electrodes: two linear electrodes with varying distance along channel
Electrodes: width 50 µm,
gap 50 µm (increasing over length of
channel), height 100 nm
AC: 10 Vpp, 200 kHz [86]
eDEP, trapping Tungsten trioxide (WO3)nanoparticle (80 nm diameter)
Galinstan liquid metal electrodes (68.5% gallium, 21.5% indium, and 10% tin)
Circular droplet-shape 3D electrodes(half- sphere)
Electrodes: Cr/Au pad diameter 80 µm, height 30 µm
AC: 15 V*, 1 MHz [59]
eDEP, cell
stretching RBCs Liquid metal consisting of 67% Ga, 20.5% In, and 12.5
% Sn and ITO electrode.
Liquid electrode channel and triangular ITO electrode
Liquid metal electrode:
1000 µm width, 3 cm length
AC: 1 Vpp-10 Vpp at 1.5 MHz
[60]
eDEP, mixing N/A Eutectic gallium indium liquid electrode-EGaIn (Ga 75% In 25% by weight)
Liquid electrode channel walls contained by PDMS posts
1000 µm width, 50 µm
height DC: 100 V [61]
eDEP, microdroplet generation
Water droplets formed in silicon oil (660 µm droplet length)
Ga 75.5In24.5 (75.5 wt.%
Ga, 24.5 wt.% In) Liquid electrode squared
channels 70 µm electrode width DC: 2600 V-3000 V [62]
LFFF-DEP,
isolation CTCs Au Interdigitated electrodes with
channel bonded on top Width of each
electrode ~60 μm and the width of the main channel is ~200 μm
Top set of electrodes at 10 Vpp, 10 kHz while the bottom set of electrodes is held at 15 Vpp, 40 kHz [87]
* the author(s) did not specify whether the voltage was measured as Vrms, Vp or Vpp.
** the author(s) did not provide detailed information about the electrode dimensions
33
Voltage reduction strategies in insulator-based devices
In comparison with electrode-based devices, some of the advantages of insulator-based devices are easy and low-cost fabrication [10], and no direct contact between electrodes and sample at the trapping region, which avoids contamination of the sample due to electrolysis or corrosion, or of the electrode due to electrode fouling [20], [48]. A disadvantage of introducing insulating materials in these devices is that the voltage required to achieve particle or cell manipulation is much higher than that required in devices where the electrodes are directly in contact with the sample [40].
We note that, although insulator-based dielectrophoresis (iDEP) has been reported in the literature for two decades, a study published recently by our group [2] provided evidence that DEP is generally negligible in devices stimulated with DC voltages. Findings suggest that the main contributors to particle movement in such devices are linear and non-linear electrophoresis and electroosmosis. Therefore, insulator-based electrokinetics (iEK) is a more adequate term for this technology. Nonetheless, two decades of work in the ‘DC- iDEP’ field to decrease input voltage requirements deserve mention in this review.
iEK
In 2000, Cummings and Singh [88] introduced the use of an array of insulating posts rather than only using a single insulator constriction, as had been done in previous works.
In the following years, they were also the first to present streaming DEP [89], [90] using insulating structures as an alternative to embedded metal electrodes to produce nonuniform electric fields. Their initial work presented streaming DEP at a field of 800 V/cm and trapping DEP at a field of 1000 V/cm using carboxylated latex spheres. An example of the array of insulating posts design used in some iDEP devices is presented in Figure 1C. The setup for these devices typically includes one reservoir on each end of the main channel, where wire electrodes are placed so that the voltage is applied for particle manipulation in the channel. In 2004, the first application of iDEP for bacteria and polystyrene beads separation and concentration at fields up to 2000 V/cm was presented [91].
A more recent study presented a device that assessed the DEP response of submicron (100 nm - 1 µm) aminated and carboxylated particles. The study considered the response as a function of size and surface charge at low frequency (<1000 Hz) with AC signals of 2800 Vpp and 3200 Vpp, respectively [40]. The voltages in this work were in the range initially used in iDEP. Although they were manipulating nm-sized particles, their insulating structures were not optimized for this particle size as both the post diameter and the gap were in the µm size range. Furthermore, the array of insulating posts consisted of 18 columns of posts in the channel center, adding more resistance to the channel and thus requiring a higher voltage. In 2015, an iDEP device for trapping yeast cells (~6 µm) and 2 µm polystyrene particles in low concentration from a diluted sample with smaller particles was developed, working at a DC voltage of 600 V with 99% trapping efficiency
