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Technical note
A microfluidic device for continuous manipulation of biological cells using dielectrophoresis Debanjan Das a,∗ , Karabi Biswas a , Soumen Das b a b
Department of Electrical Engineering, Indian Institute of Technology, Kharagpur 721302, India School of Medical Science and Technology, Indian Institute of Technology, Kharagpur 721302, India
a r t i c l e
a b s t r a c t
i n f o
Article history: Received 12 July 2013 Received in revised form 25 November 2013 Accepted 8 December 2013 Keywords: DEP Continuous particles manipulation Microfluidics COMSOL
The present study demonstrates the design, simulation, fabrication and testing of a label-free continuous manipulation and separation micro-device of particles/biological cells suspended on medium based on conventional dielectrophoresis. The current dielectrophoretic device uses three planner electrodes to generate non-uniform electric field and induces both p-DEP and n-DEP force simultaneously depending on the dielectric properties of the particles and thus influencing at least two types of particles at a time. Numerical simulations were performed to predict the distribution of non-uniform electric field, DEP force and particle trajectories. The device is fabricated utilizing the advantage of bonding between PDMS and SU8 polymer. The p-DEP particles move away from the center of the streamline, while the n-DEP particles will follow the central streamline along the channel length. Dielectrophoretic effects were initially tested using polystyrene beads followed by manipulation of HeLa cells. In the experiment, it was observed that polystyrene beads in DI water always response as n-DEP up to 1 MHz frequency, whereas HeLa cells in PBS medium response as n-DEP up to 400 kHz frequency and then it experiences p-DEP up to 1 MHz. Further, the microscopic observations of DEP responses of HeLa cells were verified by performing trapping experiment at static condition. © 2013 IPEM. Published by Elsevier Ltd. All rights reserved.
1. Introduction The recent developments in medical technology and significant improvements in health care has increased life expectancy, but still chronic diseases like cancer, diabetes, heart attack, etc. accounts for majority of deaths around the world, mostly in dense populated countries. In the present scenario, medical practitioners are increasingly focusing on prophylaxis paradigm, which needs rapid diagnostic techniques. Though different staining procedures have been developed to differentiate target cells from other cells, sorting efficiency depends on how well the technicians discriminate the target cells from the background cells manually. During disease characterization if population of target cells is negligible compared to large number of normal cells, distinguishing these target cells from the background using staining techniques are tedious and time consuming. Therefore, we need a technique which can manipulate individual cell from mixed population and characterize them rapidly with greater accuracy. There are several techniques to manipulate and separate biological particles in microfluidic platform. The techniques include Fluorescence
Activated Cell Sorter (FACS) [1], Magnetic Activated Cell Sorter (MACS) [2], optical tweezers [3], acoustic means [4], electrophoresis and dielectrophoresis [5,6]. Among the existing techniques for cell manipulations and separation, dielectrophoresis (DEP) has become extensively promising technique for cell separation because it offers label-free detection, easy to fabricate using well-established microfabrication processes [7] and requires small sample volume. Furthermore, DEP based microfluidic systems are easily interfaced with fluidic connection to load the biological samples, and to provide electrical signal in Lab-on-a-Chip (LoC) [8]. The term dielectrophoresis was first coined by H.A. Pohl [9] to describe the motion of a neutral particle, such as biological cells [10] caused by polarization effects in a non-uniform electric field. The polarization varies from cell to cell and also modifies in the same cell with progression of a disease due to change of its dielectric properties that alter with the biological processes. Therefore, the precise manipulation of the single cell can be realized by controlling electric field in a fluidic channel without using any mechanical parts. The DEP force depends on the volume and the complex dielectric property of the particle as well as medium and the gradient of time averaged applied electric field [11–15] based on Eq. →
∗ Corresponding author. Tel.: +91 9432121661. E-mail address: [email protected] (D. Das).
FDEP = 2εm a3 Re[fcm ]∇ |E|2 , where a is radius of particle, εm is permittivity of suspending medium, E is electric field strength, Re[fcm ] is real part of the Clausius–Mossotti factor. Re[fcm ] describes the
1350-4533/$ – see front matter © 2013 IPEM. Published by Elsevier Ltd. All rights reserved. http://dx.doi.org/10.1016/j.medengphy.2013.12.010
Please cite this article in press as: Das D, et al. A microfluidic device for continuous manipulation of biological cells using dielectrophoresis. Med Eng Phys (2013), http://dx.doi.org/10.1016/j.medengphy.2013.12.010
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Fig. 1. Schematic of dielectrophoretic microfluidic device.
electrode at the same tapering angle. An inhibiter free movement of the biological cells is maintained by providing a continuous 50 m gap between central and side electrodes. Three inlets in the microchannel are designed to guide the mixture and suspension medium over these non-uniform patterned electrodes and corresponding three outlets are created to collect the separated particles. The 50 m height open micro-channel is covered with a transparent material to monitor the movement of particles. The advantages of this planar design are that electrodes do not act as barrier to the cells/particles, leading to lower voltage requirement and separation of smaller particles. In addition, mixtures of multiple cells/particles can be separated in a continuous flow manner due to presence of nonuniform electric field throughout the particle flow in the microchannel. In the principle of operation, the sample (cells/particles suspended in medium) is injected from inlet 2 and is guided along the central streamline by the sheath flow from inlet 1 and 3. When a potential is applied to the electrodes, non-uniform electric field is developed across the central and side electrodes. The electric filed intensity gradient is very high on side electrode, while it is very weak along the central electrode. Therefore, while flowing through the micro-channel the interaction of cells with this non-uniform electric field would induce positive DEP force toward the two side electrodes and negative DEP force along the middle electrode. Thus the induced positive DEP force would push the cells away from the central electrode and they will move along the two side electrodes and will be collected in outlet 1 and 3. The cells experiencing negative DEP force would flow along the middle electrode region and they will be collected in outlet 2. This induced DEP force and directions are dependent on properties of both cells/particles beads and suspending medium. The cells experience DEP forces with different magnitude and direction depending on its size, dielectric properties, electric field and frequency while moving continuously through the microfluidics channel. As a result, they continuously move to different location across the channel and thereby separated into different outlets.
frequency dependent polarization of the particle which is a function of complex permittivity of particle and suspending medium [12]. The sign of Re[fcm ] determines the positive or negative DEP, if Re[fcm ] > 0 then positive DEP occurs otherwise negative [12]. The positive DEP drives the particle to maxima of the electric-field intensity, while negative drives the particle to the minima of the electric-field intensity. The theoretical limit of Re[fcm ] lies in the range of −0.5 < Re[fcm ] < 1 [9,16]. The crossover frequency can be determined only if (m − p )/(εp − εm ) > 0. A non-uniform electric field is required to sustain DEP force in the neighboring region of the cells. This non-uniform field is generated by patterning different shape of electrodes [17] or insulating material altering the electric field lines [18]. Particles are manipulated by discrete [19,20] or continuous flow in a fluidic channel depending upon the design of electrodes. Continuous manipulation has advantages over discrete manipulation in terms of operation time, electrode area, external flow control valve, and direct integration with Lab-on-chip (LoC) device. The existing continuous DEP manipulators are based on field flow fractionation [21–23], trapezoidal electrode arrays [24], 3D electrode [25,26], curved microelectrodes [27], tapered electrodes [28–33], twDEP [34], multiple frequencies [35,36]. In this study, a continuous particle manipulating microfluidic device utilizing conventional DEP have been demonstrated. The microfluidic device is compatible with high flow rate and high speed particle manipulation. The device is fabricated utilizing the advantage of bonding between PDMS and SU8 polymer. It generates both p-DEP and n-DEP force simultaneously depending on the dielectric properties of the particles and thus influencing at least two types of particles at a time. In this study the fabricated device is used to demonstrate manipulation of HeLa cell and polystyrene beads. The study also provides a detailed simulation modeling of dielectrophoresis in COMSOL Multiphysics platform to investigate the different electrical and fluidic parameters.
The proposed dielectrophoretic microchannel for continuous manipulation of particles was designed in COMSOL Multiphysics 4.2a. The device was modeled in 3D using AC/DC and Particle Tracing module of COMSOL for simulating the electric field distribution, quantifying the DEP forces acting on the cells/particles and tracing the movement of cells/particles under DEP force. The 3D geometrical model was built in electric current (ec) interface under the AC/DC module using Eq. (1).
2. Materials and methods
∇ · J = Qj
2.1. Design of dielectrophoretic microfluidic device The primary design task for the dielectrophoretic microfluidic device is to identify an appropriate size and configuration of the micro-electrodes which results non-uniform electric field and corresponding DEP force. DEP force is proportional to the gradient of square of the non-uniform electric field intensity. Different groups [28–32] have designed tapered microelectrodes to generate the non-uniform electric field. In the present design, three planar metal electrodes with different cross sectional area are designed to induce non-uniform electric field which are embedded in the microfluidic channel as shown in Fig. 1. Central electrode is tapered at an angle of 9◦ with a width of 100 m at inlet side, 1.058 mm at outlet side spreading over 3.046 mm along the channel length. Two sideelectrodes having width of 50 m are designed along with central
2.2. COMSOL simulation
E = −∇ V D = ε0 εr E
(1)
J = E + jωD + Je where Qj , , ω, D, V, Je are charge density, conductivity, angular frequency, electric displacement field, applied potential and external generated current density, respectively. Electrical insulating boundary conditions were enforced on all the boundaries, except the electrodes area. All the three electrodes are modeled as perfect conductors by considering as equipotential surfaces. The boundary condition of two side electrodes has been imposed as voltage terminal and a potential of 10 Vpp was applied where as zero potential was applied to central electrode. Particle tracing physics interface was coupled with the electric current interface to predict the DEP force and particles trajectory. COMSOL simulates the particle
Please cite this article in press as: Das D, et al. A microfluidic device for continuous manipulation of biological cells using dielectrophoresis. Med Eng Phys (2013), http://dx.doi.org/10.1016/j.medengphy.2013.12.010
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Fig. 2. Microphotograph of the final fabricated DEP-microfluidic device.
momentum and trajectory by solving the Newton’s second law of motion as given in Eq. (2). d (mp v) = Ft dt
(2)
Ft = 2a3 ε0 εr,m Re[fcm ]∇ |E|2 where mp is the particle mass, v is the velocity of the particle, a is diameter of particle, εr,m is the relative permittivity of the fluid, Re[fcm ] is real part of Clausius–Mossotti factor, and E is the electric field. Freezing wall boundary conditions were imposed on the entire boundary wall, except the inlet where 0.4 mm/s initial velocity along y-axis was applied. Subsequently the properties of particle having mass of 5 mg, diameter of 10 m, and relative permittivity values of particles and fluid were applied to quantify the DEP force on the cells/particles. The model was meshed using fine free tetrahedral configuration having maximum and minimum element size of 0.244 mm and 30.5 m for the discretization of the domain with total 71,653 mesh elements. Initially electric current interface was solved in frequency domain, and then particle tracing interface was computed in time dependent mode to evaluate the distribution of electric field, to predict the DEP force and particle trajectory inside the micro-channel. 2.3. Device fabrication The DEP device consists of inhomogeneous planner bottom electrodes to generate non-uniform electric field and a microchannel to guide the flow. The device is realized by two mask process using the conventional microfabrication technology. The basic steps involved in fabrication of the DEP device are shown in Fig. S1. At first, Cr/Au electrode was patterned, followed by realization of SU-8 microchannel. The SU-8 based open channels is sealed with half cured PDMS. The detailed fabrication processes are described in the supplementary data. Subsequently, the necessary fluidic connections and electrical connections were taken from the complete sealed device. The microphotograph of complete fabricated device with electrodes, inlet and outlet of the fluid channel is shown in Fig. 2. The schematic view of the test setup is shown in supplementary Fig. S2. Electrodes were excited by the electric signal having sine wave with 10 Vpp from a function generator. A micro-syringe pump was used to flow the sample through the channel. The experiment was conducted under an inverted microscope connected to a CCD camera to monitor the movement of particles under electric field. 2.4. Sample preparation Polystyrene beads and human carcinoma (HeLa) cells were selected to test the DEP effect and manipulate them in the fabricated device. Beads sample was prepared by mixing DI water and uncharged polystyrene beads having diameter of 10 m and dielectric constant 2.5 in 1000:10 (v/v) ratio. HeLa cells were cultured by using cell culture facility available in authors’ laboratory. HeLa cells were cultured in its growth medium in a humidified atmosphere containing 5% carbon dioxide at 37 ◦ C. The confluent
Fig. 3. Variation of Clausius–Mossotti (CM) factor with frequency for (i) polystyrene beads in DI water and (ii) HeLa cells in PBS medium.
cells with density of 106 cells in 1 ml were finally re-suspended in 2 ml of fresh PBS medium having pH value 7.4 and conductivity 1.56 S/m and mixed thoroughly by pipetting in and out while storing in a syringe connected to syringe pump for conducting the DEP experiment.
3. Results and discussion 3.1. Analytical results Movement of particles inside the dielectrophoretic based microchannel depends on Clausius–Mossotti factor which varies with conductivity and dielectric property of particles and medium and also with applied frequency of the electric field. The variation of CM factor with the applied AC frequency for polystyrene beads suspended in DI water and HeLa cells in PBS medium is shown in Fig. 3. The real part of CM factor is theoretically simulated using the values as given in Table S1 in supplementary material. The theoretical result shows that Re[fcm (ε, , ω)] value for beads in the frequency range of 100 kHz to 100 MHz is always negative, which indicates that the beads will always experience n-DEP force in this frequency range as shown in Fig. 3(i). The real part of CM factor for HeLa cells under PBS medium is negative below 500 kHz leading to generate n-DEP force to the HeLa cell in this frequency range. However, above 500 kHz the CM factor gradually increases to positive value and thus cells will experience p-DEP force above 500 kHz. Fig. 3(ii)
Please cite this article in press as: Das D, et al. A microfluidic device for continuous manipulation of biological cells using dielectrophoresis. Med Eng Phys (2013), http://dx.doi.org/10.1016/j.medengphy.2013.12.010
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Fig. 4. Trajectory of (a) polystyrene beads in DI water and (b) HeLa cells in PBS medium.
also depicts that theoretically crossover frequency for HeLa cells suspended in PBS medium is at 500 kHz. 3.2. Numerical simulation results The electric field distribution, DEP forces on both polystyrene beads/cells and manipulation of particles for the designed configuration are numerically simulated in COMSOL Multiphysics. The surface electric field distribution across the electrodes in microchannel is shown in Fig. S3(a). It shows that the applied 10 Vpp ac potential to three electrodes configuration with central electrode at ground potential generates a symmetric non-uniform electric field across the micro-channel with higher gradient of electric field near the edges of electrode. It is also observed from electric field distribution that the maximum value of |E| is about 900 V/cm, which is below the threshold value required for electroporation of cells [37]. Thus, 10 Vpp potential will not affect the viability of the cells during the experiment. The numerical simulations of DEP effect to manipulate polystyrene in DI water medium, and HeLa cells suspended in PBS medium were performed with inlet flow velocity of 400 m/s and applied voltage of 10 Vpp at 1 MHz considering the parameters as listed in Table S1. The magnitude of DEP forces were found in pN range for both polystyrene and HeLa cells sample, which produces sufficient force to manipulate the particle by DEP process [38,39] while flowing in microfluidic channel. The simulated results indicate that there is weak electric field region on the central electrode and two strong field regions in the interelectrode gaps as observed in Fig. S3(b). Therefore, it will generate higher DEP force around the side electrodes along the channel length which will push the pDEP particle away from the center of the streamline and trap them in electrode gaps, while the n-DEP particles will follow the central streamline in the micro-fluidic channel. Fig. 4 shows the trajectory of both beads and HeLa cells and their spatial distribution along the channel length. It is apparent from the simulation results of Fig. 4 that at the same conditions the maximum number of beads are mostly located along the central electrode region and moved along with the main streamline of the fluid from inlet to outlet as shown in Fig. 4(a). However, Fig. 4(b) depicts the movement of HeLa cells toward the maximum electric field region and thus it trapped to the side electrode location along the channel length. The results indicate that at 1 MHz electric field, beads experience n-DEP force while HeLa cells experience p-DEP force. The simulation results matches with the theoretically obtained results for cross-over frequency of beads and HeLa cells.
observed their trajectory in the channel under high magnification optical microscope to verify the simulation results. In this process, 10 m diameter uncharged polystyrene beads were mixed in DI water at the density of 103 beads per 1 ml and beads mixture was pumped at 1 l/min into the channel through syringe pump. Particles movement was observed under microscope by applying ac voltage of 10 Vpp with frequency sweep of 100 kHz to 1 MHz. It was found that beads followed through the center streamline, i.e. along central electrode area, indicating n-DEP force for the
3.3. Experimental verification Initial experiment was conducted by flowing beads and HeLa cells separately through fabricated microfluidic device and
Fig. 5. (i) Microscopic observation of movement of HeLa cells (a) at 100 kHz and (b) at 1 MHz. (ii) Microscopic observation of (a) no trapping of HeLa cells at 100 kHz and (b) trapped cells at 1 MHz frequency along the side electrodes at no flow condition.
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entire frequency range. This follows a good agreement with simulated results where beads trajectory was mostly observed over the center electrode region. The experimental observation also matches with the analytical results, which shows the real CM value for this case is in negative range for the applied frequency range indicating n-DEP force. Subsequently, HeLa cells suspended in PBS medium at the density of 106 cells in 2 ml of medium was pumped at 1 l/min under same applied voltage and frequency range as given for the beads movement. It was observed that maximum of the injected cells move along the central electrode and flow over the central streamline in the frequency range of 100 kHz to 400 kHz. A typical microphotograph shown in Fig. 5(i) indicates the movement of HeLa cells along central streamline at 100 kHz frequency. These results indicate that HeLa cells showed n-DEP response in this frequency range. As the frequency increases, at 500 kHz, there was random movement of cells indicating the absence of DEP force which shows the crossover frequency for HeLa cell in PBS media. In 700 kHz to 1 MHz, it was observed that maximum HeLa cells moved along the side electrode having higher electric field and thus experiencing positive DEP response. Fig. 5(ii) shows the optical photograph of the HeLa cell trajectory at 1 MHz frequency. Thus, the experimentally observed DEP responses show a good accordance with the theoretically and numerically observed results. Further to ensure the measured DEP response the same experiment was performed with no-flow condition. In this approach the cells were pushed in the microfluidic channel through syringe pump for a certain period of time. Subsequently the pump was switched off and the electric field was applied to the device. It was observed that at 100 kHz no trapping of cells was occurred at side electrodes while at 1 MHz cells were trapped along the side electrode as shown in Fig. 5(ii). 4. Conclusion In this study, a continuous dielectrophoretic microdevice for dielectric particle manipulation and trapping has been presented. The simulation result shows that the applied electric potential to the designed three electrodes generates a symmetric non-uniform electric field across the micro-channel with weak electric field region on the central electrode and two strong field regions in the interelectrode gaps. This will force the p-DEP particle to move away from the center of the streamline, while the n-DEP particles will follow the central streamline along the channel length. The simulated result also indicates that for applied ac potential of 10 Vpp the maximum value of |E| is about 900 V/cm, which is below the threshold value of electroporation fields required for the cells. The simulated trajectory of polystyrene beads in DI water and HeLa cells in PBS medium shows that for applied voltage of 10 Vpp at 1 MHz HeLa cells are experiencing p-DEP but beads experience nDEP which matches with the analytical results. The microfluidic device was fabricated using standard photolithography technique and exploiting the advantage of bonding between PDMS and SU8 polymer. The test results indicate that beads in DI water always response as n-DEP in the frequency range of 100 kHz to 1 MHz, whereas HeLa cells in PBS medium response as n-DEP in the frequency range of 100 kHz to 400 kHz and as p-DEP in the range of 700 kHz to 1 MHz. Further, the DEP responses of HeLa cells were verified by performing trapping experiment at static condition and a good agreement was achieved with simulation and analytical results. Thus the present investigation may be extended further for separation of biological cells having different DEP response from its mixed population for disease characterization. Acknowledgements The work has been initiated under a project sponsored by National Programme on Micro and Smart Systems (NPMASS), Govt.
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of India. The authors would like to thank NPMASS committee for providing the financial support. The authors would also like to acknowledge the staff members of MEMS & Microelectronics Laboratory of ATDC, cell culture group of SMST, IIT Kharagpur for providing the necessary facilities to fabricate the device and carry out the experiment. Conflict of interest No conflict of interest. Appendix A. Supplementary data Supplementary data associated with this article can be found, in the online version, at http://dx.doi.org/10.1016/j.medengphy. 2013.12.01. References [1] Fu AY, Spence C, Scherer A, Arnold FH, Quake SR. A microfabricated fluorescence-activated cell sorter. Nature Biotechnology 1999;17:1109–11. [2] Handgretinger R, Lang P, Schumm M, Taylor G, Neu S, Koscielnak E, et al. Isolation and transplantation of autologous peripheral CD34+ progenitor cells highly purified by magnetic-activated cell sorting. Bone Marrow Transplantation 1998;21:987–93. [3] Grier DG. A revolution in optical manipulation. Nature Photonics 2003;424:810–6. [4] Hawkes JJ, Coakley WT. Force field particle filter, combining ultrasound standing waves and laminar flow. Sensors and Actuators B: Chemical 2001;75:213–22. [5] Pohl HA. Dielectrophoresis: The Behaviour of Neutral Matter in Non-Uniform Electric Fields. Cambridge: Cambridge University Press; 1978. [6] Washizu M. Electrostatic manipulation of biological objects. Journal of Electrostatics 1990;25:109–23. [7] Gascoyne PRC, Vykoukal J. Particle separation by dielectrophoresis. Electrophoresis 2002;23:1973–83. [8] Morgan H, Green NG. AC Electrokinetics: Colloids and Nanoparticles. Hertfordshire Research Studies Press; 2003. [9] Pohl HA. The motion and precipitation of suspensoids in divergent electric fields. Journal of Applied Physics 1951;22:869–71. [10] Campbell NA, Reece JB. Biology. San Francisco: Pearson; 2005. [11] Jones TB. Basic theory of dielectrophoresis and electrorotation. IEEE Engineering in Medicine and Biology Magazine 2003;22:33–42. [12] Pethig R. Dielectrophoresis: status of the theory, technology, and applications. Biomicrofluidics 2010;4 [Review article]. [13] Regtmeier J, Eichhorn R, Viefhues M, Bogunovic L, Anselmetti D. Electrodeless dielectrophoresis for bioanalysis: theory, devices and applications. Electrophoresis 2011;32:2253–73. [14] C¸etin B, Li D. Dielectrophoresis in microfluidics technology. Electrophoresis 2011;32:2410–27. [15] Gagnon ZR. Cellular dielectrophoresis: applications to the characterization, manipulation, separation and patterning of cells. Electrophoresis 2011;32:2466–87. [16] Jones TB. Electromechanics of Particles. New York: Cambridge University Press; 1995. [17] Nieuwenhuis JH, Jachimowicz A, Svasek P, Vellekoop MJ. Optimization of microfluidic particle sorters based on dielectrophoresis. IEEE Sensors Journal 2005;5:810–6. [18] Lapizco-Encinas BH, Simmons BA, Cummings EB, Fintschenko Y. Insulatorbased dielectrophoresis for the selective concentration and separation of live bacteria in water. Electrophoresis 2004;25:1695–704. [19] Markx GH, Talary MS, Pethig R. Separation of viable and non-viable yeast using dielectrophoresis. Journal of Biotechnology 1994;32:29–37. [20] Huang Y, Joo S, Duhon M, Heller M, Wallace B, Xu X. Dielectrophoretic cell separation and gene expression profiling on microelectronic chip arrays. Analytical Chemistry 2002;74:3362–71. [21] Wang X-B, Yang J, Huang Y, Vykoukal J, Becker FF, Gascoyne PRC. Cell separation by dielectrophoretic field-flow-fractionation. Analytical Chemistry 2000;72:832–9. [22] Gupta V, Jafferji I, Garza M, Melnikova VO, Hasegawa DK, Pethig R, et al. ApoStream, a new dielectrophoretic device for antibody independent isolation and recovery of viable cancer cells from blood. Biomicrofluidics 2012;6. [23] Shim S, Stemke-Hale K, Tsimberidou AM, Noshari J, Anderson TE, Gascoyne PRC. Antibody-independent isolation of circulating tumor cells by continuous-flow dielectrophoresis. Biomicrofluidics 2013;7. [24] Choi S, Park J-K. Microfluidic system for dielectrophoretic separation based on a trapezoidal electrode array. Lab on a Chip 2005;5:1161–7. [25] Wang L, Flanagan LA, Jeon NL, Monuki E, Lee AP. Dielectrophoresis switching with vertical sidewall electrodes for microfluidic flow cytometry. Lab on a Chip 2007;7:1114–20. [26] Ching-Te H, Cheng-Hsin W, Chun-Ping J. Three-dimensional cellular focusing utilizing negative dielectrophoretic force generated by dual-planar electrodes.
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Technical note
A microfluidic device for continuous manipulation of biological cells using dielectrophoresis Debanjan Das a,∗ , Karabi Biswas a , Soumen Das b a b
Department of Electrical Engineering, Indian Institute of Technology, Kharagpur 721302, India School of Medical Science and Technology, Indian Institute of Technology, Kharagpur 721302, India
a r t i c l e
a b s t r a c t
i n f o
Article history: Received 12 July 2013 Received in revised form 25 November 2013 Accepted 8 December 2013 Keywords: DEP Continuous particles manipulation Microfluidics COMSOL
The present study demonstrates the design, simulation, fabrication and testing of a label-free continuous manipulation and separation micro-device of particles/biological cells suspended on medium based on conventional dielectrophoresis. The current dielectrophoretic device uses three planner electrodes to generate non-uniform electric field and induces both p-DEP and n-DEP force simultaneously depending on the dielectric properties of the particles and thus influencing at least two types of particles at a time. Numerical simulations were performed to predict the distribution of non-uniform electric field, DEP force and particle trajectories. The device is fabricated utilizing the advantage of bonding between PDMS and SU8 polymer. The p-DEP particles move away from the center of the streamline, while the n-DEP particles will follow the central streamline along the channel length. Dielectrophoretic effects were initially tested using polystyrene beads followed by manipulation of HeLa cells. In the experiment, it was observed that polystyrene beads in DI water always response as n-DEP up to 1 MHz frequency, whereas HeLa cells in PBS medium response as n-DEP up to 400 kHz frequency and then it experiences p-DEP up to 1 MHz. Further, the microscopic observations of DEP responses of HeLa cells were verified by performing trapping experiment at static condition. © 2013 IPEM. Published by Elsevier Ltd. All rights reserved.
1. Introduction The recent developments in medical technology and significant improvements in health care has increased life expectancy, but still chronic diseases like cancer, diabetes, heart attack, etc. accounts for majority of deaths around the world, mostly in dense populated countries. In the present scenario, medical practitioners are increasingly focusing on prophylaxis paradigm, which needs rapid diagnostic techniques. Though different staining procedures have been developed to differentiate target cells from other cells, sorting efficiency depends on how well the technicians discriminate the target cells from the background cells manually. During disease characterization if population of target cells is negligible compared to large number of normal cells, distinguishing these target cells from the background using staining techniques are tedious and time consuming. Therefore, we need a technique which can manipulate individual cell from mixed population and characterize them rapidly with greater accuracy. There are several techniques to manipulate and separate biological particles in microfluidic platform. The techniques include Fluorescence
Activated Cell Sorter (FACS) [1], Magnetic Activated Cell Sorter (MACS) [2], optical tweezers [3], acoustic means [4], electrophoresis and dielectrophoresis [5,6]. Among the existing techniques for cell manipulations and separation, dielectrophoresis (DEP) has become extensively promising technique for cell separation because it offers label-free detection, easy to fabricate using well-established microfabrication processes [7] and requires small sample volume. Furthermore, DEP based microfluidic systems are easily interfaced with fluidic connection to load the biological samples, and to provide electrical signal in Lab-on-a-Chip (LoC) [8]. The term dielectrophoresis was first coined by H.A. Pohl [9] to describe the motion of a neutral particle, such as biological cells [10] caused by polarization effects in a non-uniform electric field. The polarization varies from cell to cell and also modifies in the same cell with progression of a disease due to change of its dielectric properties that alter with the biological processes. Therefore, the precise manipulation of the single cell can be realized by controlling electric field in a fluidic channel without using any mechanical parts. The DEP force depends on the volume and the complex dielectric property of the particle as well as medium and the gradient of time averaged applied electric field [11–15] based on Eq. →
∗ Corresponding author. Tel.: +91 9432121661. E-mail address: [email protected] (D. Das).
FDEP = 2εm a3 Re[fcm ]∇ |E|2 , where a is radius of particle, εm is permittivity of suspending medium, E is electric field strength, Re[fcm ] is real part of the Clausius–Mossotti factor. Re[fcm ] describes the
1350-4533/$ – see front matter © 2013 IPEM. Published by Elsevier Ltd. All rights reserved. http://dx.doi.org/10.1016/j.medengphy.2013.12.010
Please cite this article in press as: Das D, et al. A microfluidic device for continuous manipulation of biological cells using dielectrophoresis. Med Eng Phys (2013), http://dx.doi.org/10.1016/j.medengphy.2013.12.010
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Fig. 1. Schematic of dielectrophoretic microfluidic device.
electrode at the same tapering angle. An inhibiter free movement of the biological cells is maintained by providing a continuous 50 m gap between central and side electrodes. Three inlets in the microchannel are designed to guide the mixture and suspension medium over these non-uniform patterned electrodes and corresponding three outlets are created to collect the separated particles. The 50 m height open micro-channel is covered with a transparent material to monitor the movement of particles. The advantages of this planar design are that electrodes do not act as barrier to the cells/particles, leading to lower voltage requirement and separation of smaller particles. In addition, mixtures of multiple cells/particles can be separated in a continuous flow manner due to presence of nonuniform electric field throughout the particle flow in the microchannel. In the principle of operation, the sample (cells/particles suspended in medium) is injected from inlet 2 and is guided along the central streamline by the sheath flow from inlet 1 and 3. When a potential is applied to the electrodes, non-uniform electric field is developed across the central and side electrodes. The electric filed intensity gradient is very high on side electrode, while it is very weak along the central electrode. Therefore, while flowing through the micro-channel the interaction of cells with this non-uniform electric field would induce positive DEP force toward the two side electrodes and negative DEP force along the middle electrode. Thus the induced positive DEP force would push the cells away from the central electrode and they will move along the two side electrodes and will be collected in outlet 1 and 3. The cells experiencing negative DEP force would flow along the middle electrode region and they will be collected in outlet 2. This induced DEP force and directions are dependent on properties of both cells/particles beads and suspending medium. The cells experience DEP forces with different magnitude and direction depending on its size, dielectric properties, electric field and frequency while moving continuously through the microfluidics channel. As a result, they continuously move to different location across the channel and thereby separated into different outlets.
frequency dependent polarization of the particle which is a function of complex permittivity of particle and suspending medium [12]. The sign of Re[fcm ] determines the positive or negative DEP, if Re[fcm ] > 0 then positive DEP occurs otherwise negative [12]. The positive DEP drives the particle to maxima of the electric-field intensity, while negative drives the particle to the minima of the electric-field intensity. The theoretical limit of Re[fcm ] lies in the range of −0.5 < Re[fcm ] < 1 [9,16]. The crossover frequency can be determined only if (m − p )/(εp − εm ) > 0. A non-uniform electric field is required to sustain DEP force in the neighboring region of the cells. This non-uniform field is generated by patterning different shape of electrodes [17] or insulating material altering the electric field lines [18]. Particles are manipulated by discrete [19,20] or continuous flow in a fluidic channel depending upon the design of electrodes. Continuous manipulation has advantages over discrete manipulation in terms of operation time, electrode area, external flow control valve, and direct integration with Lab-on-chip (LoC) device. The existing continuous DEP manipulators are based on field flow fractionation [21–23], trapezoidal electrode arrays [24], 3D electrode [25,26], curved microelectrodes [27], tapered electrodes [28–33], twDEP [34], multiple frequencies [35,36]. In this study, a continuous particle manipulating microfluidic device utilizing conventional DEP have been demonstrated. The microfluidic device is compatible with high flow rate and high speed particle manipulation. The device is fabricated utilizing the advantage of bonding between PDMS and SU8 polymer. It generates both p-DEP and n-DEP force simultaneously depending on the dielectric properties of the particles and thus influencing at least two types of particles at a time. In this study the fabricated device is used to demonstrate manipulation of HeLa cell and polystyrene beads. The study also provides a detailed simulation modeling of dielectrophoresis in COMSOL Multiphysics platform to investigate the different electrical and fluidic parameters.
The proposed dielectrophoretic microchannel for continuous manipulation of particles was designed in COMSOL Multiphysics 4.2a. The device was modeled in 3D using AC/DC and Particle Tracing module of COMSOL for simulating the electric field distribution, quantifying the DEP forces acting on the cells/particles and tracing the movement of cells/particles under DEP force. The 3D geometrical model was built in electric current (ec) interface under the AC/DC module using Eq. (1).
2. Materials and methods
∇ · J = Qj
2.1. Design of dielectrophoretic microfluidic device The primary design task for the dielectrophoretic microfluidic device is to identify an appropriate size and configuration of the micro-electrodes which results non-uniform electric field and corresponding DEP force. DEP force is proportional to the gradient of square of the non-uniform electric field intensity. Different groups [28–32] have designed tapered microelectrodes to generate the non-uniform electric field. In the present design, three planar metal electrodes with different cross sectional area are designed to induce non-uniform electric field which are embedded in the microfluidic channel as shown in Fig. 1. Central electrode is tapered at an angle of 9◦ with a width of 100 m at inlet side, 1.058 mm at outlet side spreading over 3.046 mm along the channel length. Two sideelectrodes having width of 50 m are designed along with central
2.2. COMSOL simulation
E = −∇ V D = ε0 εr E
(1)
J = E + jωD + Je where Qj , , ω, D, V, Je are charge density, conductivity, angular frequency, electric displacement field, applied potential and external generated current density, respectively. Electrical insulating boundary conditions were enforced on all the boundaries, except the electrodes area. All the three electrodes are modeled as perfect conductors by considering as equipotential surfaces. The boundary condition of two side electrodes has been imposed as voltage terminal and a potential of 10 Vpp was applied where as zero potential was applied to central electrode. Particle tracing physics interface was coupled with the electric current interface to predict the DEP force and particles trajectory. COMSOL simulates the particle
Please cite this article in press as: Das D, et al. A microfluidic device for continuous manipulation of biological cells using dielectrophoresis. Med Eng Phys (2013), http://dx.doi.org/10.1016/j.medengphy.2013.12.010
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Fig. 2. Microphotograph of the final fabricated DEP-microfluidic device.
momentum and trajectory by solving the Newton’s second law of motion as given in Eq. (2). d (mp v) = Ft dt
(2)
Ft = 2a3 ε0 εr,m Re[fcm ]∇ |E|2 where mp is the particle mass, v is the velocity of the particle, a is diameter of particle, εr,m is the relative permittivity of the fluid, Re[fcm ] is real part of Clausius–Mossotti factor, and E is the electric field. Freezing wall boundary conditions were imposed on the entire boundary wall, except the inlet where 0.4 mm/s initial velocity along y-axis was applied. Subsequently the properties of particle having mass of 5 mg, diameter of 10 m, and relative permittivity values of particles and fluid were applied to quantify the DEP force on the cells/particles. The model was meshed using fine free tetrahedral configuration having maximum and minimum element size of 0.244 mm and 30.5 m for the discretization of the domain with total 71,653 mesh elements. Initially electric current interface was solved in frequency domain, and then particle tracing interface was computed in time dependent mode to evaluate the distribution of electric field, to predict the DEP force and particle trajectory inside the micro-channel. 2.3. Device fabrication The DEP device consists of inhomogeneous planner bottom electrodes to generate non-uniform electric field and a microchannel to guide the flow. The device is realized by two mask process using the conventional microfabrication technology. The basic steps involved in fabrication of the DEP device are shown in Fig. S1. At first, Cr/Au electrode was patterned, followed by realization of SU-8 microchannel. The SU-8 based open channels is sealed with half cured PDMS. The detailed fabrication processes are described in the supplementary data. Subsequently, the necessary fluidic connections and electrical connections were taken from the complete sealed device. The microphotograph of complete fabricated device with electrodes, inlet and outlet of the fluid channel is shown in Fig. 2. The schematic view of the test setup is shown in supplementary Fig. S2. Electrodes were excited by the electric signal having sine wave with 10 Vpp from a function generator. A micro-syringe pump was used to flow the sample through the channel. The experiment was conducted under an inverted microscope connected to a CCD camera to monitor the movement of particles under electric field. 2.4. Sample preparation Polystyrene beads and human carcinoma (HeLa) cells were selected to test the DEP effect and manipulate them in the fabricated device. Beads sample was prepared by mixing DI water and uncharged polystyrene beads having diameter of 10 m and dielectric constant 2.5 in 1000:10 (v/v) ratio. HeLa cells were cultured by using cell culture facility available in authors’ laboratory. HeLa cells were cultured in its growth medium in a humidified atmosphere containing 5% carbon dioxide at 37 ◦ C. The confluent
Fig. 3. Variation of Clausius–Mossotti (CM) factor with frequency for (i) polystyrene beads in DI water and (ii) HeLa cells in PBS medium.
cells with density of 106 cells in 1 ml were finally re-suspended in 2 ml of fresh PBS medium having pH value 7.4 and conductivity 1.56 S/m and mixed thoroughly by pipetting in and out while storing in a syringe connected to syringe pump for conducting the DEP experiment.
3. Results and discussion 3.1. Analytical results Movement of particles inside the dielectrophoretic based microchannel depends on Clausius–Mossotti factor which varies with conductivity and dielectric property of particles and medium and also with applied frequency of the electric field. The variation of CM factor with the applied AC frequency for polystyrene beads suspended in DI water and HeLa cells in PBS medium is shown in Fig. 3. The real part of CM factor is theoretically simulated using the values as given in Table S1 in supplementary material. The theoretical result shows that Re[fcm (ε, , ω)] value for beads in the frequency range of 100 kHz to 100 MHz is always negative, which indicates that the beads will always experience n-DEP force in this frequency range as shown in Fig. 3(i). The real part of CM factor for HeLa cells under PBS medium is negative below 500 kHz leading to generate n-DEP force to the HeLa cell in this frequency range. However, above 500 kHz the CM factor gradually increases to positive value and thus cells will experience p-DEP force above 500 kHz. Fig. 3(ii)
Please cite this article in press as: Das D, et al. A microfluidic device for continuous manipulation of biological cells using dielectrophoresis. Med Eng Phys (2013), http://dx.doi.org/10.1016/j.medengphy.2013.12.010
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Fig. 4. Trajectory of (a) polystyrene beads in DI water and (b) HeLa cells in PBS medium.
also depicts that theoretically crossover frequency for HeLa cells suspended in PBS medium is at 500 kHz. 3.2. Numerical simulation results The electric field distribution, DEP forces on both polystyrene beads/cells and manipulation of particles for the designed configuration are numerically simulated in COMSOL Multiphysics. The surface electric field distribution across the electrodes in microchannel is shown in Fig. S3(a). It shows that the applied 10 Vpp ac potential to three electrodes configuration with central electrode at ground potential generates a symmetric non-uniform electric field across the micro-channel with higher gradient of electric field near the edges of electrode. It is also observed from electric field distribution that the maximum value of |E| is about 900 V/cm, which is below the threshold value required for electroporation of cells [37]. Thus, 10 Vpp potential will not affect the viability of the cells during the experiment. The numerical simulations of DEP effect to manipulate polystyrene in DI water medium, and HeLa cells suspended in PBS medium were performed with inlet flow velocity of 400 m/s and applied voltage of 10 Vpp at 1 MHz considering the parameters as listed in Table S1. The magnitude of DEP forces were found in pN range for both polystyrene and HeLa cells sample, which produces sufficient force to manipulate the particle by DEP process [38,39] while flowing in microfluidic channel. The simulated results indicate that there is weak electric field region on the central electrode and two strong field regions in the interelectrode gaps as observed in Fig. S3(b). Therefore, it will generate higher DEP force around the side electrodes along the channel length which will push the pDEP particle away from the center of the streamline and trap them in electrode gaps, while the n-DEP particles will follow the central streamline in the micro-fluidic channel. Fig. 4 shows the trajectory of both beads and HeLa cells and their spatial distribution along the channel length. It is apparent from the simulation results of Fig. 4 that at the same conditions the maximum number of beads are mostly located along the central electrode region and moved along with the main streamline of the fluid from inlet to outlet as shown in Fig. 4(a). However, Fig. 4(b) depicts the movement of HeLa cells toward the maximum electric field region and thus it trapped to the side electrode location along the channel length. The results indicate that at 1 MHz electric field, beads experience n-DEP force while HeLa cells experience p-DEP force. The simulation results matches with the theoretically obtained results for cross-over frequency of beads and HeLa cells.
observed their trajectory in the channel under high magnification optical microscope to verify the simulation results. In this process, 10 m diameter uncharged polystyrene beads were mixed in DI water at the density of 103 beads per 1 ml and beads mixture was pumped at 1 l/min into the channel through syringe pump. Particles movement was observed under microscope by applying ac voltage of 10 Vpp with frequency sweep of 100 kHz to 1 MHz. It was found that beads followed through the center streamline, i.e. along central electrode area, indicating n-DEP force for the
3.3. Experimental verification Initial experiment was conducted by flowing beads and HeLa cells separately through fabricated microfluidic device and
Fig. 5. (i) Microscopic observation of movement of HeLa cells (a) at 100 kHz and (b) at 1 MHz. (ii) Microscopic observation of (a) no trapping of HeLa cells at 100 kHz and (b) trapped cells at 1 MHz frequency along the side electrodes at no flow condition.
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entire frequency range. This follows a good agreement with simulated results where beads trajectory was mostly observed over the center electrode region. The experimental observation also matches with the analytical results, which shows the real CM value for this case is in negative range for the applied frequency range indicating n-DEP force. Subsequently, HeLa cells suspended in PBS medium at the density of 106 cells in 2 ml of medium was pumped at 1 l/min under same applied voltage and frequency range as given for the beads movement. It was observed that maximum of the injected cells move along the central electrode and flow over the central streamline in the frequency range of 100 kHz to 400 kHz. A typical microphotograph shown in Fig. 5(i) indicates the movement of HeLa cells along central streamline at 100 kHz frequency. These results indicate that HeLa cells showed n-DEP response in this frequency range. As the frequency increases, at 500 kHz, there was random movement of cells indicating the absence of DEP force which shows the crossover frequency for HeLa cell in PBS media. In 700 kHz to 1 MHz, it was observed that maximum HeLa cells moved along the side electrode having higher electric field and thus experiencing positive DEP response. Fig. 5(ii) shows the optical photograph of the HeLa cell trajectory at 1 MHz frequency. Thus, the experimentally observed DEP responses show a good accordance with the theoretically and numerically observed results. Further to ensure the measured DEP response the same experiment was performed with no-flow condition. In this approach the cells were pushed in the microfluidic channel through syringe pump for a certain period of time. Subsequently the pump was switched off and the electric field was applied to the device. It was observed that at 100 kHz no trapping of cells was occurred at side electrodes while at 1 MHz cells were trapped along the side electrode as shown in Fig. 5(ii). 4. Conclusion In this study, a continuous dielectrophoretic microdevice for dielectric particle manipulation and trapping has been presented. The simulation result shows that the applied electric potential to the designed three electrodes generates a symmetric non-uniform electric field across the micro-channel with weak electric field region on the central electrode and two strong field regions in the interelectrode gaps. This will force the p-DEP particle to move away from the center of the streamline, while the n-DEP particles will follow the central streamline along the channel length. The simulated result also indicates that for applied ac potential of 10 Vpp the maximum value of |E| is about 900 V/cm, which is below the threshold value of electroporation fields required for the cells. The simulated trajectory of polystyrene beads in DI water and HeLa cells in PBS medium shows that for applied voltage of 10 Vpp at 1 MHz HeLa cells are experiencing p-DEP but beads experience nDEP which matches with the analytical results. The microfluidic device was fabricated using standard photolithography technique and exploiting the advantage of bonding between PDMS and SU8 polymer. The test results indicate that beads in DI water always response as n-DEP in the frequency range of 100 kHz to 1 MHz, whereas HeLa cells in PBS medium response as n-DEP in the frequency range of 100 kHz to 400 kHz and as p-DEP in the range of 700 kHz to 1 MHz. Further, the DEP responses of HeLa cells were verified by performing trapping experiment at static condition and a good agreement was achieved with simulation and analytical results. Thus the present investigation may be extended further for separation of biological cells having different DEP response from its mixed population for disease characterization. Acknowledgements The work has been initiated under a project sponsored by National Programme on Micro and Smart Systems (NPMASS), Govt.
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of India. The authors would like to thank NPMASS committee for providing the financial support. The authors would also like to acknowledge the staff members of MEMS & Microelectronics Laboratory of ATDC, cell culture group of SMST, IIT Kharagpur for providing the necessary facilities to fabricate the device and carry out the experiment. Conflict of interest No conflict of interest. Appendix A. Supplementary data Supplementary data associated with this article can be found, in the online version, at http://dx.doi.org/10.1016/j.medengphy. 2013.12.01. References [1] Fu AY, Spence C, Scherer A, Arnold FH, Quake SR. A microfabricated fluorescence-activated cell sorter. Nature Biotechnology 1999;17:1109–11. [2] Handgretinger R, Lang P, Schumm M, Taylor G, Neu S, Koscielnak E, et al. Isolation and transplantation of autologous peripheral CD34+ progenitor cells highly purified by magnetic-activated cell sorting. Bone Marrow Transplantation 1998;21:987–93. [3] Grier DG. A revolution in optical manipulation. Nature Photonics 2003;424:810–6. [4] Hawkes JJ, Coakley WT. Force field particle filter, combining ultrasound standing waves and laminar flow. Sensors and Actuators B: Chemical 2001;75:213–22. [5] Pohl HA. Dielectrophoresis: The Behaviour of Neutral Matter in Non-Uniform Electric Fields. Cambridge: Cambridge University Press; 1978. [6] Washizu M. Electrostatic manipulation of biological objects. Journal of Electrostatics 1990;25:109–23. [7] Gascoyne PRC, Vykoukal J. Particle separation by dielectrophoresis. Electrophoresis 2002;23:1973–83. [8] Morgan H, Green NG. AC Electrokinetics: Colloids and Nanoparticles. Hertfordshire Research Studies Press; 2003. [9] Pohl HA. The motion and precipitation of suspensoids in divergent electric fields. Journal of Applied Physics 1951;22:869–71. [10] Campbell NA, Reece JB. Biology. San Francisco: Pearson; 2005. [11] Jones TB. Basic theory of dielectrophoresis and electrorotation. IEEE Engineering in Medicine and Biology Magazine 2003;22:33–42. [12] Pethig R. Dielectrophoresis: status of the theory, technology, and applications. Biomicrofluidics 2010;4 [Review article]. [13] Regtmeier J, Eichhorn R, Viefhues M, Bogunovic L, Anselmetti D. Electrodeless dielectrophoresis for bioanalysis: theory, devices and applications. Electrophoresis 2011;32:2253–73. [14] C¸etin B, Li D. Dielectrophoresis in microfluidics technology. Electrophoresis 2011;32:2410–27. [15] Gagnon ZR. Cellular dielectrophoresis: applications to the characterization, manipulation, separation and patterning of cells. Electrophoresis 2011;32:2466–87. [16] Jones TB. Electromechanics of Particles. New York: Cambridge University Press; 1995. [17] Nieuwenhuis JH, Jachimowicz A, Svasek P, Vellekoop MJ. Optimization of microfluidic particle sorters based on dielectrophoresis. IEEE Sensors Journal 2005;5:810–6. [18] Lapizco-Encinas BH, Simmons BA, Cummings EB, Fintschenko Y. Insulatorbased dielectrophoresis for the selective concentration and separation of live bacteria in water. Electrophoresis 2004;25:1695–704. [19] Markx GH, Talary MS, Pethig R. Separation of viable and non-viable yeast using dielectrophoresis. Journal of Biotechnology 1994;32:29–37. [20] Huang Y, Joo S, Duhon M, Heller M, Wallace B, Xu X. Dielectrophoretic cell separation and gene expression profiling on microelectronic chip arrays. Analytical Chemistry 2002;74:3362–71. [21] Wang X-B, Yang J, Huang Y, Vykoukal J, Becker FF, Gascoyne PRC. Cell separation by dielectrophoretic field-flow-fractionation. Analytical Chemistry 2000;72:832–9. [22] Gupta V, Jafferji I, Garza M, Melnikova VO, Hasegawa DK, Pethig R, et al. ApoStream, a new dielectrophoretic device for antibody independent isolation and recovery of viable cancer cells from blood. Biomicrofluidics 2012;6. [23] Shim S, Stemke-Hale K, Tsimberidou AM, Noshari J, Anderson TE, Gascoyne PRC. Antibody-independent isolation of circulating tumor cells by continuous-flow dielectrophoresis. Biomicrofluidics 2013;7. [24] Choi S, Park J-K. Microfluidic system for dielectrophoretic separation based on a trapezoidal electrode array. Lab on a Chip 2005;5:1161–7. [25] Wang L, Flanagan LA, Jeon NL, Monuki E, Lee AP. Dielectrophoresis switching with vertical sidewall electrodes for microfluidic flow cytometry. Lab on a Chip 2007;7:1114–20. [26] Ching-Te H, Cheng-Hsin W, Chun-Ping J. Three-dimensional cellular focusing utilizing negative dielectrophoretic force generated by dual-planar electrodes.
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