Single-cell manipulation on microfluidic chip by dielectrophoretic actuation and impedance detection

Sensors and Actuators B 150 (2010) 167–173

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Sensors and Actuators B: Chemical journal homepage: www.elsevier.com/locate/snb

Single-cell manipulation on microfluidic chip by dielectrophoretic actuation and impedance detection Hyunjin Park, Dongil Kim, Kwang-Seok Yun ∗ Department of Electronic Engineering, Sogang University, 1 Shinsu-dong, Mapo-gu, Seoul 121-742, Republic of Korea

a r t i c l e

i n f o

Article history: Received 3 June 2010 Received in revised form 7 July 2010 Accepted 13 July 2010 Available online 21 July 2010 Keywords: Dielectrophoresis Single-cell Impedance Microfluidic device

a b s t r a c t This paper presents the design, fabrication, and characterization of a microfluidic biochip with integrated actuation electrodes used to manipulate a cell and a microbead by dielectrophoresis and sensing electrodes to detect the trapping by using the impedance detection method. Combining deflective dielectrophoretic barriers with controlled pressure-driven liquid flows allows the accurate control of a cell/microbead in suspensions. The threshold voltage for microbead trapping was experimentally verified at various flow rates. The impedance change caused by the blockage of the electrical conducting path between sensing electrodes with the trapping of an MCF7 cell and a polystyrene microbead was measured. The impedance before the trapping of an MCF7 cell was 10.9 M at 1 kHz and increased to 12 M when the cell was placed between sensing electrodes. © 2010 Elsevier B.V. All rights reserved.

1. Introduction Analysis of single cells allows the characterization of rare or unusual cells without the averaging effect or cell–cell interaction through junctional proteins [1,2], lipid nanotubes [3], and membrane-bound receptors and ligands [4,5]. Microfabrication technology is expected to provide a microfluidic platform by solving the problems found in conventional systems, which require expensive equipment and a large sample volume for cell assay experiments [6–8]. Thus far, several technologies have been evaluated for single-cell assay on microfluidic chips, including optical tweezers [9–11], physical capturing [6,7,12], magnetic tweezers [13], and dielectrophoresis (DEP) [14–16,31–33]. Among them, active manipulation techniques using DEP have been widely used on current microfluidic chips because these techniques are suitable for integrated microfluidic systems [17] and do not require tagging materials such as magnetic materials [18]. DEP arises from the interaction of a nonuniform electric field with the dipoles induced in the particles. Depending on the relative polarizability of the particles with respect to the surrounding medium, DEP is categorized into two methodologies. In positive DEP, the cells are attracted to the high field region, which results in a cell lysis problem [19]. Conversely, in negative DEP the cells are forced to move away from the high field region and are less affected by the electric field [14–16]. Recently, active positioning control of single cells using negative DEP has been developed

∗ Corresponding author. Tel.: +82 2 705 8915; fax: +82 2 705 8915. E-mail address: [email protected] (K.-S. Yun). 0925-4005/$ – see front matter © 2010 Elsevier B.V. All rights reserved. doi:10.1016/j.snb.2010.07.020

for high-throughput single-cell isolation and microfluidic chips for cell-based assays. However, the need for visual monitoring using an optical microscope and the manual control of electrical signals hinder the automatic manipulation of the microfluidic chips [20]. Previously, the systematic control of microfluidic chips has been accomplished with the optical recognition of cell position and movement, but these methods required an optical microscope, image analysis software, and sometimes the biochemical modification of cells for fluorescence detection [10,21]. In addition to these issues, the light energy associated with optical detection can generate a local heating problem in the cells and the cell medium. In this paper, we propose a microfluidic chip that employs an electrical measurement technique for single-cell monitoring and active manipulation of cells by using DEP. This electrical monitoring requires only simple electrodes in a microfluidic channel. For the purpose of electrical characterization and sorting of cells, there are several reports about on-chip integration of microfluidics and electrodes [22–25]. Recently, Ferrier et al. demonstrated a microfluidic chip with interdigitated electrodes in a microfluidic channel that simultaneously actuates and detects individual biological cells [26,27]. In their system, a single-cell is trapped using impedancebased detection on electrodes within a uniform flow channel. This restricts both the expansion of the trapping location to a one- or two-dimensional array and the specific chemical excitation of the trapped cell because the trapped cell is not physically isolated. In our proposed device, we integrated a microchannel with a trapping chamber, sensing electrodes and actuation electrodes to trap single cells in a specific chamber positioned to the side of the main flow channel using impedance measurement and DEP application.

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Fig. 1. Schematic diagram of proposed microfluidic chip.

2. Design Fig. 1 shows the schematic view and operation principles of the proposed chip. The chip is composed of an inlet, an outlet, and trapping chambers connected to the main channel. Each trapping chamber is equipped with sensing electrodes for monitoring of cell trapping and actuation electrodes for DEP application. First, cells are introduced into the inlet and move through the main channel by external pressure. When a cell arrives at the first actuation electrode, it is forced into the trapping chamber by negative DEP (a → b → c → d → e). Once the cell is positioned in the trapping channel (position “e” in Fig. 1), the impedance between the sensing electrodes increases because the cell blocks the electrical conducting path between the two sensing electrodes. Thus, the trapping of the cell can be electrically monitored by measuring the impedance change at the sensing electrodes. Then, the DEP signal on first actuation electrode is turned off and the subsequent cell passes the first trapping chamber (a → b → f), from where it is captured in the following second trapping chamber by the same method. A cell arriving at the actuation electrodes is forced into the trapping chamber by DEP force and hydrodynamic force, as described in Fig. 2(a), when an electrical potential is applied on the actuation electrodes. The strength of the force depends on the permittivity of the medium and the particles, the size of the particles, and the frequency of the electric field. The time-averaged DEP force in the dipole is given by



2

FDEP = 2εm r 3 Re[fCM ] · ∇  Erms 

(1)

where r denotes the radius of the particles; εm , the permittivity of the medium; fCM , the Clausius–Mossotti factor; and Erms , the root mean square value of the electric field [28]. The Clausius–Mossotti factor describes the frequency dependence of the effective polarizability and is given by Eq. (2), with complex permittivity of particle (ε∗P ) and buffer medium (ε∗m ) for a spherical and homogeneous particle. fCM =

ε∗P − ε∗m

ε∗P + 2ε∗m

(2)

The sign of the Clausius–Mossotti factor fCM determines whether a particle is being repelled from (negative DEP) or attracted to (positive DEP) regions of high field strength. In addition to the DEP force (FDEP ), the particle is forced into the direction of liquid flow by drag force, which is given by FHD = 6˛

(3)

where ˛ is the radius of the particle,  is the viscosity of the fluid, and v is the velocity of the flow. Then, the particle will be deflected

Fig. 2. (a) Schematic diagram of actuation electrode in microchannel and (b) electrical equivalent models without (left) and with (right) cell between sensing electrodes.

by the action of the DEP force as long as FDEP > FHD sin  = 6˛ sin 

(4)

where  is the angle between the direction of liquid flow and the electrodes. Because the planar actuation electrodes are deposited on the bottom surface of the microchannel, the particle experiences both lateral deflection and vertical lifting force as a result of the combination of the negative DEP and hydrodynamic forces. When a particle is pushed to the side of the microchannel and arrives at the end region of the actuation electrodes (position “d” in Fig. 1), it is pushed into the trapping chamber by the negative DEP force. The trapping of a cell is monitored by measuring the impedance between the sensing electrodes. The impedance detection is related to the electrode response to the application of a periodic alternating current (AC) signal. Generally, these measurements are carried out at different frequencies. Impedance detection that uses sensing electrodes in an aqueous solution can be modeled as a simple electrical circuit [29]. Fig. 2(b) shows the electrical equivalent models before and after trapping of a cell between the sensing electrodes. Without the cell, the impedance between the two electrodes is formed by electrical double-layer capacitance on both electrodes (Cdl1 and Cdl2 ), resistance of the bulk medium in main channel (Rb1 ) and in drain channel (Rb2 ), capacitance of the bulk medium in main channel (Cb1 ) and in drain channel (Cb2 ), and resistance (Rm ) and

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capacitance (Cm ) for the trapping zone medium, as shown in Eq. (5): ZNC =

1 1 Rb1 Rb2 + + + jωCdl1 jωCdl2 jωCb1 Rb1 + 1 jωCb2 Rb2 + 1 +

Rm jωCm Rm + 1

(5)

When a cell is placed between electrodes, the cell membrane and the intercellular medium create additional impedance, which is much larger than the impedance of the culture medium or Phosphate buffered saline (PBS). The impedance for this case becomes ZC =

Rb2 1 1 Rb1 + + + jωCdl1 jωCdl2 jωCb1 Rb1 + 1 jωCb2 Rb2 + 1 +

Rmem + Rcyt jωCmem Rmem + 1

(6)

where Cmem is the capacitance of the cell membrane, Rmem is the resistance of the cell membrane, and Rcyt is the resistance of the cell cytoplasm. Thus, the cell trapping can be electrically monitored by measuring the impedance between the sensing electrodes. 3. Experiments 3.1. Microchip implementation The proposed microfluidic chip was implemented using a polydimethylsiloxane (PDMS) layer and a glass substrate that were bonded together. Fig. 3(a) shows the fabrication procedures. For ˚ layer was deposited and the glass substrate, a Cr/Pt (200/2000 a) patterned by a conventional wet etching process using positive photoresist (AZ1512, AZ electronic materials) as a masking layer to form electrodes. The Cr and Pt layers were etched using a commercial chrome etchant (CR-7) at room temperature and a HCl:HNO3 3:1 solution at 50 ◦ C, respectively. For the upper PDMS structures, a photosensitive epoxy, SU-8 (Microchem Co.), was patterned on a silicon substrate to be utilized as a replica mold for the PDMS. To form two different channel thicknesses, we used the double exposure and single development technique [30]. First, the SU-8 2007 was coated to a thickness of 7 ␮m and exposed to ultraviolet (UV) to define the thin drain channel region. Without development, a second coating of SU-8 2050 (thickness: 50 ␮m) was applied, followed by UV exposure, to define the thick channel region. The UV exposure energies were 135 and 190 mJ for the shallow and deep exposures, respectively. Then, an unexposed region of SU-8 was developed in a commercial SU-8 developer (Microchem Co.) for 6 min, forming a PDMS mold. Next, a self-assembled monolayer (SAM) of tridecafluoro-(1,1,2,2-tetrahydrooctyl)-1-trichlorosilane (United Chemical Technology Inc.) was coated on both the silicon and SU-8 surface in a vacuum desiccator for 20 min to ensure that the PDMS layer can be easily peeled off. Then, a PDMS prepolymer (Sylgard 184, Dow Corning Co.) was poured onto the mold structure and inherent bubbles were removed in a vacuum chamber. Finally, the PDMS was cured at 80 ◦ C for 1 h and peeled off from the substrate mold. The access holes for the sample inlet, outlet, and drain port were formed by manual punching. Finally, the completed PDMS structure was bonded with the glass substrate after surface treatment by using oxygen plasma treatment for 30 s at an RF power of 100 W in a plasma exposure apparatus (COVANCE, FemtoScience, Inc.). Fig. 3(b) shows the fabricated microfluidic chip and a magnified view of a microchamber. The width and height of the main channel were 50 and 50 ␮m, respectively. The drain channel was located at the bottom of the trapping chamber, which had dimensions of 10 ␮m width and 7 ␮m height. The actuation electrodes were placed at the entrance of the trapping chamber to guide the

Fig. 3. (a) Fabrication process of proposed device and (b) fabricated microfluidic chip and magnified view.

cell into the chamber. For sensing electrodes, one electrode was placed in the trapping chamber and the other in the drain channel to maximize the impedance change when a cell was positioned between these electrodes.

3.2. Materials for cell experiment For the cell experiment, cancerous human breast epithelial cells, MCF7, were used. The cell suspensions were prepared in Dulbecco’s modified Eagle’s medium (GibcoBRL, Grand Island, NY, USA) containing 10% fetal bovine serum (FBS; PAA Laboratories, Exton, PA, USA) and penicillin–streptomycin (Sigma–Aldrich). Polystyrene microbeads (Polysciences, Inc., USA) with diameter of 15 ␮m were also prepared in cell culture medium to provide the same condition with the cell experiment. A DEP signal was applied by using a function generator (33220A, Agilent), and the impedance was measured with an LCR meter (4284A, Agilent).

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4. Results and discussion Before the experiment, the fabricated chip was fully primed using culture medium for both of the microbead and cell experiment to avoid air trapping during the experiment. Then, the microbeads or cells were loaded in the inlet port. In this experiment, the fluid stream was controlled by the height difference of each port. From Eq. (4), it is clear that a cell or microbead will be captured in the trapping chamber when a DEP force is strong enough to overcome the hydrodynamic friction force exerted by the buffer flow. In this work, we experimentally determined the threshold voltage required to capture a microbead in the trapping chamber at various flow velocities, as displayed in Fig. 4. In this experiment, the conductivity of culture medium was 1.4 S/m and the frequency of applied DEP signal was set at 10 MHz. First, the flow rate was controlled by the differences in the liquid levels between the inlet, outlet, and drain port. The flow rate in the drain channel was maintained at less than 7% of the flow rate in the main channel to prevent the autonomous capturing of a particle without DEP actuation. Then, the voltage was increased until a microbead was deflected and trapped in the trapping chamber by the DEP force. The velocity of each microbead at a given flow rate was deter-

Fig. 4. Threshold voltage required to capture microbead as a function of flow velocity.

mined using video microscopy by monitoring the movement of the microbeads in the absence of the DEP force. In this experiment, most of the microbeads flowed in the middle of the microchannel without hydrodynamic focusing at the inlet. Infrequently, a few microbeads that flowed slowly along the side of the microchannel

Fig. 5. Sequence of capturing (a) single-cell (MCF7) and (b) microbead in trapping chamber.

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Fig. 6. (a) Average impedance values before and after trapping of cell (n = 5) and microbead (n = 7), (b) incremental rate of impedance after trapping, (c) impedance variation at 1 kHz when MCF7 cell is captured in trapping chamber, (d) impedance variation for microbead experiment.

were excluded from the experiment and measurement. Because the experiment would have been influenced by any blockage in the thin drain channel, every experiment was repeated after removing the trapped microbead to clean the trapping chamber. In this experiment, the data were obtained after repeating each experiment 3–5 times for each voltage except 4, 9 and 10 V, for each of which only one experiment was performed. The threshold flow velocity was expected to be proportional to the square of the amplitude of the DEP voltage from Eq. (1). However, in our experiment, a linear relationship was found between the flow velocity and DEP voltage. As the voltage increased, the force pushing the microbead up increased and pushed the microbead into a high velocity region in the microchannel having a parabolic flow profile. This then produced a higher hydrodynamic force, that could explain a linear increase in the threshold trapping force with the voltage increase. Fig. 5 shows the sequence of capturing a single-cell and a microbead in the trapping chamber. In this experiment, MCF7 cells and polystyrene microbeads in culture medium were tested. The flow velocity was 500 ␮m/s and the voltage applied on the actuation electrodes was set to 10 Vpp at 10 MHz to apply enough force to trap a cell or a microbead. We were able to clearly monitor that a cell and a microbead were carried toward the trapping chamber and exactly positioned between the sensing electrodes when the negative DEP force prevailed over drag force. Fig. 6 shows the impedance measurement results of the MCF7 cell and polystyrene microbead experiments. The impedance was measured at 1 Vpp with the LCR meter. In this experiment, the impedance on sensing electrodes was continuously monitored while a cell (or a microbead) was positioned into trapping chamber by DEP. According to the exposed area of sensing electrode in the trapping chamber, the initial impedance varies from 6.6 to 14.7 M with the average value of 10.9 M at 1 kHz. After a cell was positioned between the sensing electrodes, the average impedance increased to 12 M because the cell blocked the electrical conducting path between the sensing electrodes. For the microbead experiment, the impedance was increased by 0.4 M after trapping. As shown in Fig. 6(b), average 10 and 5% of increment of impedance was monitored for cells and microbeads experiments, respectively. The impedance change after microbead trapping was small compared to the cell experiment.

For the cell experiment, the drain channel had been tightly blocked by the cell because of the flexible property of the cell membrane. However, the sealing was not perfect for the hard polystyrene microbead experiment, resulting in an electrical conducting path through the media around the microbead. Fig. 6(c) and (d) show examples of impedance variation while a cell and a microbead in the main channel were captured in the trapping chamber. For the cell experiment, the impedance was measured at 1 kHz and was approximately 14.7 M before trapping. After the cell was positioned between the sensing electrodes, the impedance increased to 16 M. As shown in these figures, the impedance increased sharply when a cell or microbead arrived at the entrance of the thin drain channel between sensing electrodes. After the initial change, a continuous increase in impedance during the trapping period was observed in both the cell and microbead experiments. The impedance increase can be explained by a gradual blockage of the electrical path between the sensing electrodes, which is not included in our equivalent model in Fig. 2(b). The impedance might have increased as the cell membrane blocked the drain channel tightly and the microbead settled down because the drain channel was formed on the bottom surface of the trapping chamber. In this experiment, we also measured the impedance after releasing the cell into the main channel by applying positive pressure at the drain port. It is clear from Fig. 6(c) that the impedance decreased to its initial value after release of the cell into the main channel. The procedure for the microbead experiment was similar to the previous cell experiment. The impedance was approximately 6.3 M at 1 kHz before microbead trapping, as shown in this figure. After positioning of the microbead between sensing electrodes, the impedance increased to 6.8 M. Figs. 7 and 8 show the results of impedance measurements at various frequency ranges before trapping, after trapping, and after release of the cell and the microbead. The impedance measurement over wide frequency range revealed that the impedance decreased as the frequency increased because of the double-layer capacitance on each electrode. In addition, the increment ratio was more evident for the cell experiment and the low frequency range, as shown in Fig. 8. The solid line and dotted line without symbols in Fig. 7(a) are calculated results using Eq. (5) and Eq. (6), respec-

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H. Park et al. / Sensors and Actuators B 150 (2010) 167–173 Table 1 Values for the model parameter used in electrical model. Parameter Resistivity of DMEM Capacitance of DMEM MCF-7 cytoplasmic conductivity MCF-7 surface capacitance MCF-7 surface conductance

Value 33.8 24.0 0.23 12.4 11

± ± ± ± ±

Reference 2

0.7  cm 0.3 ␮F cm−2 0.01 S/m 1.8 mF/m2 2 pS

[34] [34] [35] [35] [35]

surement data shows quite good agreement with equivalent circuit model in Fig. 7(a), even though further verification is required after development of improved measurement procedure as future work. 5. Conclusions In this experiment, a microfluidic chip to isolate a single-cell or a microbead was developed. The cell and microbead were electrically manipulated by DEP actuation applied on parallel planar electrodes in a fluidic channel. The drain channel design and sensing electrodes placed in the trapping chamber allowed electrical monitoring of single-cell isolation into desired locations. An MCF7 cell and a polystyrene microbead were successfully isolated in the trapping chamber at the proper DEP conditions, which were verified experimentally at various flow rates. The impedance change was clearly monitored during the trapping and releasing of the cell and microbead. It is expected that this type of microfluidic chip can be used for high-throughput single-cell manipulation systems with the rapid detection of the trapping of single cells in specific chambers, but would require further improvements such as increasing the number of trapping chambers and integrating the feedback control electronics. Acknowledgements Fig. 7. (a) Impedance measured and calculated at various frequency ranges before trapping, after trapping, and after releasing of cell, and (b) impedance measurement results of polystyrene microbead experiments.

tively. The model parameters used in this calculation are listed in Table 1. There are several possible reasons for the difference between measured and calculated values including inaccuracy of model parameters and problem of measurement procedure. In our experiments, the impedance was measured using LCR meter while the DEP signal was applied on actuation electrodes. There was interference between DEP actuation and impedance detection and the impedance was increased by about 2 M when the DEP signal was turned on. Considering the influence from DEP signal, the mea-

Fig. 8. Impedance change ratio measured at various frequency ranges before and after trapping of cell and microbead.

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Biographies Hyunjin Park received his B.S. degree in Electronic Engineering from Sogang University in 2008. He is currently pursuing his M.S. degree in Electronic Engineering from Sogang University. His research area includes Bio-MEMS, and microscale biological analysis systems. Dongil Kim received his B.S. and M.S. degrees in Electronic Engineering from Sogang University in 2007 and 2009, respectively. He is currently pursuing his Ph.D. degree in Electronic Engineering from Sogang University. His current research area includes MEMS, Bio-sensors. Kwang-Seok Yun received his B.S. degree in Electronics Engineering from Kyungpook National University in 1996, M.S. and Ph.D. degrees in Electrical Engineering and Computer Science from Korea Advanced Institute of Science and Technology (KAIST) in 1999 and 2003, respectively. He was a post-doctorial researcher at University of California, Los Angeles from 2005 to 2007. He joined the Department of electronic Engineering at Sogang University, Korea in 2007, where he is now an Assistant Professor. His current research area includes micro total analysis systems, Lab-on-a-chip, MEMS, and micro sensors and actuators.