Sequential Field-Flow Cell Separation Method in a ... - IEEE Xplore

1120

JOURNAL OF MICROELECTROMECHANICAL SYSTEMS, VOL. 16, NO. 5, OCTOBER 2007

Sequential Field-Flow Cell Separation Method in a Dielectrophoretic Chip With 3-D Electrodes Liming Yu, Ciprian Iliescu, Guolin Xu, and Francis E. H. Tay

Abstract—This paper presents a sequential dielectrophoretic field-flow separation method for particle populations using a chip with a 3-D electrode structure. A unique characteristic of our chip is that the walls of the microfluidic channels also constitute the device’s electrodes. This property confers the opportunity to use the electrodes’ shape to generate not only the electric field gradient required for dielectrophoretic force but also a fluid velocity gradient. This interesting combination gives rise to a new solution for the dielectrophoretic separation of two particle populations. The proposed sequential field-flow separation method consists of four steps. First, the microchannel is filled with the mixture of the two populations of particle. Second, the particle populations are trapped in different locations of the microfluidic channels. The population, which exhibits positive dielectrophoresis (DEP), is trapped in the area where the distance between the electrodes is the minimum, while the other population that exhibits negative DEP is trapped in locations of maximum distance between electrodes. In the next step, increasing the flow in the microchannels will result in an increased hydrodynamic force that sweeps the cell population trapped by positive DEP out of the chip. In the last step, the electric field is removed, and the second population is swept out and collected at the outlet. For theoretical and experimental exemplification of the separation method, a population of viable and nonviable yeast cells was considered. [2006-0157] Index Terms—Cell separation, dielectrophoresis (DEP), microfluidic device, 3-D silicon electrodes.

I. I NTRODUCTION

D

IELECTROPHORESIS (DEP), which is the manipulation of neutral particles in a nonuniform electric field, is one application in which microfabrication can play an important role. This technique gives the opportunity of fabrication, on the same structure, of elements such as microchannels, valves, filters, reaction chambers, inlet/outlet holes, etc. As a result, lab-on-a-chip structures can be developed on a small area for particle manipulation. The first report of a microfabricated DEP device was made by Masunda et al. [1]. Particle manipulation based on the DEP force requires a suitable gradient of the electric field to be generated. This can be achieved using various solutions, such as micropatManuscript received July 31, 2006; revised March 12, 2007. Subject Editor A. Ricco. L. Yu and F. E. H. Tay are with the Institute of Bioengineering and Nanotechnology, Singapore 138669, and also with the National University of Singapore, Singapore 119077 (e-mail: [email protected]; [email protected]). C. Iliescu is with the Institute of Bioengineering and Nanotechnology, Singapore 138669 (e-mail: [email protected]). G. Xu is with the Institute of Bioengineering and Nanotechnology, Singapore 138669, and also with Nanyang Technological University, Singapore 639798 (e-mail: [email protected]). Color versions of one or more of the figures in this paper are available online at http://ieeexplore.ieee.org. Digital Object Identifier 10.1109/JMEMS.2007.901136

terned electrodes that can be thin films [2], [3], 3-D extruded electrodes [4]–[6], or even a combination of thin films with extruded electrodes [7]. Another method used in producing electric field gradients uses the changing of the phase of the applied electric field [8], [9]—a method known as traveling wave DEP. Another method—isolating DEP or iDEP—uses the nonuniformities generated by a dielectric material placed in a uniform electric field [10], [11]. The applications of DEP include manipulation of viruses [12], bacteria [13], or nanoparticles [14], and even trapping of DNA by DEP was achieved [15]. However, most of the research has been focused on cell concentration [16] or separation [17]. In a DEP device, movement of the particles toward the regions with higher electric field strength is called positive DEP, whereas movement toward regions with lower electric field strength is called negative DEP. The response of the particle to the electric field depends on the dielectric properties of the particle relative to the medium. Here, we report a new sequential separation field-flow technique using a DEP chip with 3-D silicon electrodes. The technique was made possible by the unique design of the device, where the electrodes are also the microfluidic channel walls, thus serving a double function. The first function is to generate positive and negative DEP forces to trap two populations of cells in different locations. The second function is to produce a fluid velocity gradient, i.e., to have zones where the velocity is maximum and, also, to have fluidic dead zones. As a consequence, the resultant hydrodynamic force will first flush out the population trapped by positive DEP. After the removal of the electric field, the second population is collected at the outlet.

II. S EPARATION M ETHODS U SING DEP The DEP-based filtration technique offers general advantages such as the ability to process complex liquid suspensions with minimum risk of clogging when compared with the membrane filters. A detailed description of DEP separation methods in lab-on-a-chip devices is presented elsewhere [18]. In addition, Gascoyne and Vykoukal reviewed particle separation by DEP in [19]. These techniques can be summarized as follows: flow separation, field-flow fractionation, stepped flow separation, travel wave DEP, and the ratcheting mechanism. The flow separation method consists of flowing a particle suspension solution over an electrode array. When there are multiple populations that exhibit positive and negative DEP, one population will be trapped near the electrode, while the others will be repelled into the center of the chamber to be subsequently pushed by the flow toward the outlet. Flow

1057-7157/$25.00 © 2007 IEEE

YU et al.: FIELD-FLOW CELL SEPARATION METHOD IN A DIELECTROPHORETIC CHIP WITH 3-D ELECTRODES

separators have been reported and demonstrated in [20] and [21]. A microfluidic device, with 3-D arrays of electrodes embedded in microchannels and with the use of a so-called “deflector” structure (electrodes oriented at certain angle compared with the flow direction), is presented in [22]. Another method uses a fluid velocity gradient to separate particles. Using an applied dielectrophoretic force field, different particles will be located at different regions within the fluid velocity gradient and will travel with different velocities. This separation method is known as field-flow fraction [23]. Another method that can be used for separation of bacteria, yeast, and plant cells uses castellated electrodes for cell trapping, two ports for outlet, and one port for inlet [24]. Particles flow from the center of the array to one of the ports. When the electric field is applied, positive and negative DEP forces are experienced by each population of particles. Similar to field-flow fraction, when the liquid is pumped, a hydrodynamic force is applied to the cells, and if this force is larger than the DEP force, the cells will be swept out. The populations can be separated and driven in opposite directions toward the two ports by selecting the correct frequency and electrical properties of the medium. Traveling wave DEP can be also used as a separation technique. A traveling electric field is generated by interdigitated parallel electrodes usually connected in 3–4 periodic intervals with different phases (0◦ , 120◦ , and 240◦ or 0◦ , 90◦ , 180◦ , and 270◦ ). Related work was reported by Huang et al. [25] (separation of yeast cells according with size), by Hughes et al. [26] (separation by changing the frequency of the electric field), and by Fuhr et al. [27] (“spiral electrode arrays”). The ratcheting mechanism has been reported in two configurations. The first system uses a “Christmas tree” electrode and, as source for particle motion—thermal motion (Brownian motion). Early work in this area was performed by Rousselet et al. [28]. The second method employs stacked ratchets, which consists of two pairs of electrodes that are stacked one over the other. The populations are inserted between these electrodes, and alternating the potential between the upper and lower electrodes moves the particles. The method was presented also by Gorre-Talini et al. [29]. III. P RINCIPLE OF S EQUENTIAL F IELD -F LOW P ARTICLE S EPARATION M ETHOD IN A DEP C HIP W ITH 3-D E LECTRODES The objective of the design is to achieve efficient separation of particle populations. The efficiency is achieved by creating a structure where separation takes place over the entire cross section of the fluidic flow. This ensures that particles flowing at any height within the channel are separated. The two distinct features of the separation structure are periodic fluidic dead volumes that are created by an undulating microchannel wall shape and an extruded electrode design that exerts a DEP force parallel to the channel floor and ceiling, and normal to the fluidic dead zones. A DEP force that is parallel to the channel floor and ceiling ensures that there are no unnecessary friction forces on the particles to impede their motion and reduce the sorting efficiency. The force exerted on particles with

1121

Fig. 1. Separation method: (a) insertion of the particles in the DEP chip, (b) cells separation using positive and negative DEP, (c) removing the first population by increasing the velocity of the fluid, and (d) the second population is released after removing the electric field.

different dielectric properties can be a positive or negative DEP force. Correct selection of the operating frequency allows this difference to be used in separating the particle populations. The separation sequence consists of four steps presented in Fig. 1. Initially, the channel is filled with the particle mixture [Fig. 1(a)]. At the optimal frequency, one population will experience a negative DEP force that drives particles into the dead fluidic zones where they remain trapped [Fig. 1(b)—zone B]. Simultaneously, the other population will experience a positive DEP force that focuses them in the region of the channel cross section where maximum fluid velocity occurs during the flow [Fig. 1(b)—zone A]. After the particles have segregated into the two regions within the channel, fresh buffer solution is pumped through the channel. The population that was focused at the center of the channel where the velocity is greatest is swept out by the drag force exerted by the fluidic flow [Fig. 1(c)]. The population trapped in the fluidic dead zones remains trapped under flow or no-flow conditions. The DEP force capturing the population in the dead zones is then reversed, and the captured population can be swept out [Fig. 1(d)]. In the following section, we will discuss the formation of the fluidic dead zones by numerical (ANSYS) simulation of different wall structures. We will then discuss the extruded electrode design using electric field analysis (Maxwell simulation) and analyze the DEP force and the gradient of the electric field that are generated across the channel for different electrode designs. Finally, we analyze the cumulative effect of DEP and hydrodynamic forces for different electrode profiles. IV. T HEORETICAL A NALYSIS AND S IMULATIONS A. Hydrodynamic Force The flow of the fluid in the microchannel can play an important role in the separation performance of the dielectrophoretic device. In our special case, the shape of the channel walls generates a specific velocity gradient in the fluid flowing in the microchannel that is essential for separation. Fig. 2 shows an ANSYS simulation of the flow in microchannels defined by electrodes that extruded the walls with different geometries (semicircular, triangular, and rectangular), with a minimal

1122


Fig. 2. Velocity gradient of the fluid flow in a microfluidic channel with semicircular, triangular, and square electrodes.

channel width of 100 µm and a maximal opening of 300 µm. The simulations were performed for a flow rate of 0.05 mL/min. These geometries lead to a huge gradient in the flow velocity between the maximum and minimum wall separation regions, neglecting the boundary layer values (which are zero by definition in the nonslip continuum model case). The low velocity regions will, from here on, be referred to as dead zones (let us approximate with those areas where the fluid velocity is less than 10% of the average velocity at the center of the channel). For this reason, the hydrodynamic force F = 6πηrν

(1)

exerted on the particles situated in the dead zones is more than one order of magnitude lower than in the channel at its narrowest region. In (1), η is the viscosity of the fluid, r is the

radius of the particle, and ν is the velocity of the fluid. The hydrodynamic force is directly proportional to the velocity, and radius of the particle, hence decreasing the velocity by one order of magnitude, correspondingly decreases the hydrodynamic force with the same amount. The calculated hydrodynamic force for a yeast cell, which is approximated by a sphere with a diameter of 7 µm, for a medium viscosity of 10−3 N · s/m2 (water) and a velocity of 100 µm/s, is 6.6 × 10−12 N. B. DEP Force The DEP force acting on a spherical particle with radius r is given by [30] F = 2πr3 εm Re [K(ω)] ∇E 2

(2)


Fig. 3. Frequency variation of the Re[K] for viable and nonviable yeast cells in a suspending medium with a conductivity of 1 mS/m.

where εm is the absolute permittivity of the suspending medium, and ∇E is the electric field gradient intensity. Re[K(ω)] is the real part of the Clausius-Mossotti factor, which is defined as K(ω) = ε∗p − ε∗m / ε∗p + 2ε∗m ε∗ = ε − j(σ/ω)

(3)

where ε∗p and ε∗m are the complex permittivity of the particle and medium, respectively. The complex permittivity for a dielectric material can be described by its permittivity ε and conductivity σ, where ω is the angular frequency of the applied electrical field E. In the expression of the dielectrophoretic force (2), the term Re[K(ω)] (Re stands for “the real part of”) plays an important role. In (3), the difference of the real part of (ε∗p − ε∗m ) can be positive or negative (Fig. 3), giving either a positive or negative DEP force. As a result, the movement of the particles toward the areas of high (positive DEP) or low electric field strength (negative DEP) is determined by the dielectric properties of the particles and medium. Particle populations that exhibit DEP forces of opposite polarity can thus be separated in this way. The complex permittivity of the particle is strongly dependent on the frequency of the generated electric field. We consider, for our experiment, viable and nonviable yeast cells. According to the model elaborated by Huang et al. [31] and applied also by Talary et al. [32] and Hughes [18], we consider the values of 60, 6, and 50, respectively, for the relative permittivities of the yeast cell wall, cytoplasmic membrane, and cell interior, respectively. For viable yeast cells, the conductivities of the cell wall, membrane, and interior were 14 mS/m, 0.25 µS/m, and 0.2 S/m, with thicknesses of the cell wall and membrane of 0.22 µm and 8 nm, respectively. In contrast, the corresponding conductivity values were 1.5 mS/m, 160 µS/m,

1123

and 7 mS/m, respectively, for nonviable yeast cells (with the same dimension of the cell wall and membrane). Employing a cell wall thickness of 0.22 µm, a cell membrane thickness of 8 nm, and values of 7 and 6 µm for the radii of viable and nonviable yeast cells, respectively, the dependence of Re[K(ω)] on frequency was derived for a suspending medium with a conductivity of 1 mS/m, as shown in Fig. 3. It can be noticed that, for low frequencies up to 50 KHz, viable yeast cells present a negative value of Re[K(ω)] which will result in a negative dielectrophoretic force, while nonviable yeast cells present a constant positive value. This range of frequencies will allow separation of these populations. Another window of frequency can be between 10 and 200 MHz where the situation is opposite: the viable yeast cells will be trapped by positive DEP, while nonviable yeast cells will experience negative DEP. For different conductivities of the medium, different frequency bands of cell separation can be achieved. The electric field gradient and the DEP force calculated for the same yeast cell, with a diameter of 6 µm positioned near the edge and in the center of the channel for all the three structure studied, were presented in Table I. The calculations were made for a permittivity of the medium of εm = 8.1 × 10−10 F/m and a Re[K(ω)] of 0.15 (maximum value for nonviable yeast cells) for a yeast cell suspended in aqueous solution with a conductivity of 1 mS/m. These data show that the hydrodynamic force and DEP force are in the same range. A comparison between the electric field generated in a DEP structure with planar electrodes and a device with extruded electrodes is presented in Fig. 4. The extruded electrode design will generate an almost uniform electric field in the vertical direction with a very small variation at the interface between the corner of the electrodes and the glass [Fig. 4(b)]. The uniformity of the electric field will determine a null DEP force in the direction perpendicular to the electrode plane (the electric field gradient is zero). A gradient of the electric field along the vertical direction—characteristic of the planar electrode structure [Fig. 4(a)]—will generate a DEP force—FZ−DEP —in this direction. For positive DEP, this force will be in the same direction with the gravity force and will increase the sedimentation and the trapping of the particle to the floor (where the hydrodynamic force is null). C. Considerations About Electrothermal Forces In addition to the DEP force and hydrodynamic force, as aforementioned, there are also other electrohydrodynamic forces that act on the particles in a DEP device. An analysis of these forces is presented by Ramos et al. [33]. Most important forces are electrothermal forces generated by high electrical fields that are used to manipulate small particles. The high electric field can generate a large power density in the fluid surrounding the electrode. Moreover, the nonuniform electric field can generate a nonuniform power density which can further generate thermal gradient which is equivalent with local changes in the density, viscosity, permittivity, and conductivity within the medium. These nonuniformities give rise to forces on the fluid. Characteristic for our structure with 3-D electrodes, as compared with the classical structure with coplanar electrodes,

1124


TABLE I ELECTRIC FIELD GRADIENT AND DIELECTROPHORETIC FORCE

Fig. 4. Simulation of the electric field in vertical plane (perpendicular to electrode) for (a) a planar DEP structure and (b) extruded DEP structure.

is a low Joule effect. At the same applied voltage, the gradient of temperature is around ten times lower (we present a detail analysis in [34]). While in a DEP device with coplanar electrodes the DEP force decreases with the distance from the electrodes, this force is constant in vertical plane in our structure. As a result, for a particle that flows at a certain distance from the floor of the microfluidic channel, an increase DEP force is required [34]. Since the relationship between the DEP force and the applied voltage is FDEP ≈ V 2

(4)

it results that an increased voltage is required for the particle trapping. Once the relationship between the gradient of temperature and the applied voltage (presented in [33]) is ∆T ≈

σV 2 k

(5)

where k is the thermal conductivity, and σ is the electrical conductivity of the medium, it became very clear that the requirement for an increased force will generate an increased gradient of temperature. We can conclude that, for the DEP structure with 3-D electrodes, the gradient of temperature and the temperature are considerately reduced, and for this reason, the forces associated with the electrothermal effect could be ignored. Nevertheless, low variation of temperature in the structure with 3-D electrodes makes it suitable for biological applications (the temperature can affect the viability of biological samples).

D. 3-D Electrode Design Consideration The theoretical selection of electrode design requires an analysis of positive and negative DEP forces exhibited on the particles (in our case, yeast cells) and also the generated hydrodynamic force. Two zones are analyzed for each electrode type: one where the positive DEP is experienced (in the area where the space of the channel is minimal) and the second where the opening of microfluidic channel is maximal (the zone where negative DEP is generated). For positive DEP, the variation of DEP force (for all of the mentioned types of electrodes) and hydrodynamic force is presented in Fig. 5. For simplification, only one graph was presented for the hydrodynamic force. We can observe that the triangular shape of the electrode gives a stronger DEP force, with a maximal value near the tip of the electrode. The cells population, which experienced positive DEP, will be trapped between the tips of the electrodes; therefore, the area where the particles will be in contact with the electrodes (and at the same time with the microchannel walls) will be reduced. The triangularly shaped electrodes provide a wider zone where the modulus of DEP force is larger than the modulus of hydrodynamic force, as compared to the semicircular or squareshaped ones. However, a comparison between the absolute values of DEP and hydrodynamic forces is not relevant due to the fact that their direction is different. It is therefore much more relevant to analyze the resultant force (which is composed of DEP and hydrodynamic forces). For this purpose, we studied three cases: when the direction of the resulting force R is inside the electrode area [Fig. 6(a)] or is parallel with the triangle edge [Fig. 6(b)] and when the direction of the resulting force R is outside the electrode area [Fig. 6(c)]. In the former two cases, the particles cannot be released. The particles can be only


1125

Fig. 7. Directions of the resulted force for semicircular and square electrodes for particle that exhibits positive DEP.

Fig. 5. Variation of hydrodynamic force and positive dielectrophoretic force for different electrode profiles between electrodes tips for 100-µm channel width.

Fig. 6.

Fig. 8. Variation of hydrodynamic force and positive dielectrophoretic force (different electrode profiles) for population that experiences negative DEP up to 100-µm distance for the channel wall.

Typical cases for triangularly shaped electrodes.

moved or rolled along the edge of the electrode in the regions where the hydrodynamic force is increased (high flow velocity) and, at the same time, where DEP force is weak. As a result, the particles can be released. Improving the release of the particles can be achieved in three ways: by decreasing the angle of the triangle apex (in this way, the probability that the direction of the resulting force to be outside the electrode area increases), by decreasing the DEP force, or by increasing the velocity of the fluid (in effect, increasing the drag—hydrodynamic force). In Fig. 6(b), the following condition for the maximal value of the triangle apex (α) can be extracted: tg(α/2) =

FHD . FDEP

(6)

By using the values from the graph presented in Fig. 5, at a distance of 5 µm from the apex of the electrode, the maximal value of α is around 50◦ . However, decreasing the angle of the profile in the dead regions will be more akin to having a square electrode. The second solution, decreasing the DEP force, can be achieved by a careful selection of some parameters such as conductivity of the solution (σ) or working frequency (that is equivalent to imposing a smaller value of Re[K(ω)]). Increasing the velocity can be performed within some experimental limits due to the increase of the velocity and force simultaneously in the area where negative DEP is experienced. For the semicircular

Fig. 9. Direction of the resulting force for semicircular for particle that exhibits negative DEP.

electrode, the positive DEP force presents a lower value, but in this case, the resulting force will keep the particles in the region situated near the electrode edge (where the velocity and also the drag force are reduced)—as shown in Fig. 7(a). For the square electrode, the resultant force will keep the particle in contact with the electrode edge—as shown in Fig. 7(b). In addition, for these cases, a small positive DEP force is recommended. For the population that experiences negative DEP, the variation of forces is presented in Fig. 8. The values of the dielectrophoretic force are presented in the absolute value. It can be observed that, for the dead zone, the DEP force and also the hydrodynamic force are very weak. Moreover, for the example presented in Fig. 9 for the semicircular electrode, it can be observed that the resultant force will keep the particles in the “dead zone” region.

1126


Fig. 10. Schematic drawing of the DEP chip.

The cells population, which experienced positive DEP, will be trapped between the tips of the electrodes; therefore, the area where the particles will be in contact with the electrodes (and at the same time with the microchannel walls) will be reduced. From the velocity point of view, the maximal velocity is achieved in the channel with rectangular electrodes, and in this case, the vertical plane that contains the maximal velocity value is not identical to the plane where most of the population is trapped by positive DEP. For the rectangular design, this plane is on the electrode corners. V. 3-D D IELECTROPHORETIC C HIP D ESCRIPTION The separation method proposed in this paper is based on a dielectrophoretic chip with 3-D electrodes. We described the fabrication process of the device and its application to yeast cell trapping in [6], and electrical and thermal characterization is presented in [34], while a similar device, but with bulk and thin electrodes, was presented in [7]. An improved version of this chip is presented in Fig. 10. As can be observed, the thick electrodes made from bulk silicon are sandwiched between two glass dies. The electrodes’ surface forms the walls of the microfluidic channels, and the glass die forms the ceiling and the floor of the microchannels. Via holes were created in the bottom glass die, and a metallization connects the electrodes to the contact pads. The inlet and outlet connections to the microfluidic channels are on the lateral sides of the chip, and the sample is injected through classical syringe needles with a diameter of 0.41 mm. The main steps of the fabrication process are presented in Fig. 11. A 4 conductive silicon wafer, with a resistivity of 0.005–0.01 Ω · cm and 300-µm thick, was anodically bonded to a glass wafer (Corning 7740) at 305 ◦ C with an applied voltage of 1000 V for 15 min [Fig. 11(a)]. In the next step [Fig. 11(b)], the patterning of electrodes was carried out. For the masking layer, a 2-µm-thick photoresist mask AZ7220 (from Clariant) was applied on top of a 2-µm-thick PECVDSiO2 layer (deposited on STS ProCVD system). The patterning of SiO2 layer was performed in a deep reactive ion etching (RIE) inductive coupling plasma (ICP) system (Adixen AMS100_DE) using CHF3 /He, while the final patterning of silicon electrodes was performed in a deep RIE ICP (Adixen AMS100_Si) using a classical Bosch process (in SF6 /C4 F8 ).

A small notching effect was noticed in the large expose area (inlet/outlet regions), but this phenomenon cannot affect the functionality of the device. A second wafer-to-wafer anodic bonding process was performed at 380 ◦ C using an applied voltage of 1200 V and an applied pressure of 500 N [Fig. 11(c)]. Prior to the previous process step, in the top glass wafer, two channels (400-µm wide and 200-µm deep) were etched using an amorphous silicon mask in an HF 49%/HCl 37% solution (10/1) [35], [36]. The bottom glass wafer was thinned up to 100 µm by wet etching process in the same solution [Fig. 11(d)]. The uniformity of the process was in an acceptable range (5%). The roughness (Ra ) of the generated surface after the wet etching process was 10 nm. Via holes were created by wet etching in the same solution through a Cr/Au (50 nm/1 µm) mask—as shown in Fig. 11(e) [37]. After removing the mask, another wet etching process of 1.5 min (equivalent to an etched depth of 10 µm) was performed mainly to remove the sharpness of the edges and also to remove the nonuniform effects of the wet etching process. In this way, the risk of metallization step coverage issues over a sharp edge is eliminated. The metallization was performed using Cr/Au deposition. The patterning of the metal layer [Fig. 11(f)] was performed using an optimized spray-coating process described in [38], with a mixture of positive photoresist AZ4620, methyl-ethyl ketone, and metoxypropyl acetate. An image with the fabricated DEP chip for sequential separation of cells is presented in Fig. 12. The dimensions of the chip are 6 × 15 × 0.9 mm. VI. T ESTING For testing of the system performance, two populations of viable yeast cells and dead yeast cells were used. 100 mg of yeast, 100 mg of sugar, and 2 mL of DI water were incubated in an Eppendorf tube at 37 ◦ C for 2 h. The cells were then concentrated by centrifugation at 1000 r/min for 1 min. The supernatant solution was removed, and the cell pellet was washed by adding 2 mL of DI water into the tube. The centrifugation and washing process was repeated three times. The cell culture was divided into two, and one population was boiled for several minutes in 5-mL phosphate buffered saline (PBS) (dead cells). Then, the cells were recollected by centrifugation. Both populations were mixed and resuspended in the separation buffer, which was a mixture of PBS and DI water. The conductivity of the separation buffer was adjusted to about 20 µS/cm−1 using NaOH. The final concentration of the cells was 107 cells/mL. A function generator and a linear amplifier were used for drive signal generation of the dielectrophoretic chip. The suspension with cell populations was injected into the chip [Fig. 13(a)], and another buffer solution was prepared for removal of the population that will be trapped. The drive signal was increased from 0 to 25 V peak to peak gradually. The signal frequency was in the range of 20–100 kHz. The trapping of the two populations by positive and negative DEP was achieved for 20 V (peak to peak), as shown in Fig. 13(b). A stable equilibrium cell concentration pattern was achieved in around 30 s. After separation of the


1127

Fig. 11. Main steps of the fabrication process: (a) anodic bonding, (b) deep RIE process, (c) SU8 bonding, (d) glass thinning, (e) via holes etching, and (f) metallization.

Fig. 12. DEP chip for sequential separation of cell populations.

Fig. 14. Optical image with the ratio between dead (red color) and living (green color) yeast cells (a) before insertion of the solution in the DEP device and (b) after the separation process.

Fig. 13. Yeast cell separation using sequential field-flow separation method.

population, a fresh buffer solution was flown through to collect the population that expresses positive DEP. The flow rate of 0.05 mL/min was assured by syringe pump (Cole Palmer

749000). The image in Fig. 13(c) shows the microchannel after removing the population trapped by positive DEP. After the release of the electric field, the second population was also removed, injecting a fresh buffer at a speed of 0.5 mL/min. In order to test the efficiency of the DEP device, yeast viability was determined using the live/dead yeast cell viability kit from Molecular Probes (Invitrogen) following the method as described in the instruction manual. Briefly, yeast cells (∼106 cells/mL) were stained with 20 µm of FUN 1 cell stain in sterile solution containing 4% D-(+)-glucose and 10 mM of Na-HEPES (pH 7.2). The cell suspension was then mixed thoroughly and incubated in the dark for approximately 1 h to allow sufficient amount of stain to diffuse into the cytoplasm and nucleus of the cells. After this, 10 µL of the stained yeast suspension was loaded into the cell counting chamber of the hemacytometer. The cells were then observed with an Olympus IX71 microscope using the blue fluorescent cube with an excitation wavelength of about 480 nm and emission wavelength of > 520 nm for the FUN 1 cell stain. The ratio of live to dead yeast cells was then determined before passing the cell suspension into the DEP chip. The initial ratio between the percentage of live and dead yeast was 42%/58% [an image is presented in Fig. 14(a)]. With the device powered at an ac voltage of 20 V, at a frequency of 20 kHz, the cell suspension was then pumped into it. After the separation

1128


process, the proportions of live to dead yeast cells were noted using the fluorescence imaging technique together with the hemacytometer. As observed in Fig. 14(a), the ratio between living and dead cells changes drastically in such a way that the percentage of the dead cells increases (up to 80%), while the percentage of living cells decreases (around 20%). This proved clearly that the separation of the two populations can be achieved. Surely, the geometry of the electrodes and the geometry of the microfluidic channel can be further optimized. VII. C ONCLUSION A new sequential field-flow separation method in a DEP chip with 3-D electrodes has been presented. A characteristic of the DEP device presented in this paper is that the geometry of the defined electrodes is also the geometry of the microfluidic channel’s lateral walls. This method successfully uses the variation of the geometry of the microfluidic channel that, at the same time, generates the gradient of the gradient of the electric field to remove the population trapped by positive DEP, while the second population remained trapped in the velocity “dead zone” of the microfluidic channel. The removal of the electric field enables to the second population to be collected at the outlet. R EFERENCES [1] S. Masunda, M. Washizu, and T. Nanba, “Novel method for cell fusion in field constriction area in fluid integrated circuit,” IEEE Trans. Ind. Appl., vol. 25, no. 4, pp. 732–737, Jul./Aug. 1989. [2] J. Yang, Y. Huang, X. B. Wang, F. F. Becker, and P. Gascoyne, “Cell separation on microfabricated electrodes using dielectrophoretic/gravitational field flow fractionation,” Anal. Chem., vol. 71, no. 5, pp. 911–918, Mar. 1999. [3] Y. Huang, J. M. Yang, P. J. Hopkins, S. Kassegne, M. Tirado, A. H. Foster, and H. Resse, “Separation of simulants of biological warfare agents from blood by a miniaturized dielectrophoresis device,” Biomed. Microdevices, vol. 5, no. 3, pp. 217–225, Sep. 2003. [4] J. Voldman, M. Toner, M. L. Gray, and M. A. Schmidt, “Design and analysis of extruded quadrupolar dielectrophoretic traps,” J. Electrostat., vol. 57, no. 1, pp. 69–90, Jan. 2003. [5] B. Y. Park and M. J. Madou, “3-D electrode designs for flowthrough dielectrophoretic systems,” Electrophoresis, vol. 26, no. 19, pp. 3745–3757, Oct. 2005. [6] C. Iliescu, G. L. Xu, V. Samper, and F. E. H. Tay, “Fabrication of a dielectrophoretic chip with 3D silicon electrodes,” J. Micromech. Microeng., vol. 15, no. 3, pp. 494–500, Mar. 2005. [7] C. Iliescu, L. Yu, G. L. Xu, and F. E. H. Tay, “A dielectrophoretic chip with a 3D electric field gradient,” J. Microelectromech. Syst., vol. 15, no. 5, pp. 1506–1513, Dec. 2006. [8] R. Pethig, M. S. Talary, and R. S. Lee, “Enhancing traveling-wave dielectrophoresis with signal superposition,” IEEE Eng. Med. Biol. Mag., vol. 22, no. 6, pp. 43–50, Nov./Dec. 2003. [9] L. Cui, D. Holmes, and H. Morgan, “The dielectrophoretic levitation and separation of latex beads in microchips,” Electrophoresis, vol. 22, no. 18, pp. 3893–3901, Oct. 2001. [10] E. B. Cummings, “Streaming dielectrophoresis for continuous-flow microfluidic devices,” IEEE Eng. Med. Biol. Mag., vol. 22, no. 6, pp. 75–84, Nov./Dec. 2003. [11] B. H. Lapizco-Encinas, B. A. Simmons, E. B. Cummings, and Y. Fintschenko, Electrophoresis, vol. 25, pp. 1704–1965, 2004. [12] M. P. Hughes, H. Morgan, F. J. Rixon, J. P. H. Burt, and R. Pethig, “Manipulation of herpes simplex virus type 1 by dielectrophoresis,” Biochim. Biophys. Acta, vol. 1425, no. 1, pp. 119–126, Sep. 1998. [13] J. Johari, Y. Hübner, J. C. Hull, J. W. Dale, and M. P. Hughes, “Dielectrophoretic assay of bacterial resistance to antibiotics,” Phys. Med. Biol., vol. 48, no. 14, pp. 193–196, Jul. 2003. [14] A. T. J. Kadaksham, P. Singh, and N. Aubry, “Dielectrophoresis of nanoparticles,” Electrophoresis, vol. 25, no. 21/22, pp. 3625–3632, Nov. 2004.

[15] C. L. Asbury, A. H. Diercks, and G. van den Engh, “Trapping of DNA by dielectrophoresis,” Electrophoresis, vol. 23, no. 16, pp. 2658–2666, Aug. 2002. [16] Y. Huang, E. L. Mather, J. L. Bell, and M. Madou, “MEMS-based sample preparation for molecular diagnostics,” Anal. Bioanal. Chem., vol. 372, no. 1, pp. 49–65, Jan. 2002. [17] Y. Li and K. Kaler, “DEP based cell separation utilizing planar electrode array,” in Annu. Report Conf. Elect. Issulation and Dielectric Phenomena, 2002, pp. 680–684. [18] M. P. Hughes, “Strategies for dielectrophoretic separation in laboratoryon-a-chip systems,” Electrophoresis, vol. 23, no. 16, pp. 2569–2582, Aug. 2002. [19] P. R. C. Gascoyne and J. Vykoukal, “Particle separation by dielectrophoresis,” Electrophoresis, vol. 23, no. 13, pp. 1973–1983, Jul. 2002. [20] H. Morgan, A. G. Izquierdo, D. Bakewell, N. G. Green, and A. Ramos, “The dielectrophoretic and travelling wave forces for interdigitated electrode arrays: Analytical solution using Fourier series,” J. Phys. D, Appl. Phys., vol. 34, pp. 1553–1561, 2001. [21] R. Pethig, J. P. H. Burt, A. Parton, N. Rizvi, M. S. Talary, and J. A. Tame, “Development of biofactory-on-a-chip technology using excimer laser micromachining,” J. Micromech. Microeng., vol. 8, no. 2, pp. 57–63, Jun. 1998. [22] J. Kentsch, M. Durr, T. Schnelle, G. Gradl, T. Muller, M. Jager, A. Normann, and M. Stelzle, “Microdevices for separation, accumulation, and analysis of biological micro- and nanoparticles,” Proc. Inst. Electr. Eng.—Nanobiotechnol., vol. 150, no. 2, pp. 82–89, 2003. [23] G. H. Markx, R. Pethig, and J. Rousselet, “The dielectrophoretic levitation of latex beads, with reference to field-flow fractionation,” J. Phys. D, Appl. Phys., vol. 30, no. 17, pp. 2470–2477, Sep. 1997. [24] G. H. Markx and R. Pethig, “Dielectrophoretic separation of cells: Continuous separation,” Biotechnol. Bioeng., vol. 45, no. 4, pp. 337–343, Feb. 1995. [25] Y. Huang, X.-B. Wang, J. Tame, and R. Pethig, “Electrokinetic behavior of colloidal particles in travelling electric fields: Studies using yeast cells,” J. Phys. D, Appl. Phys., vol. 26, pp. 1528–1535, 1993. [26] M. P. Hughes, R. Pethig, and X. B. Wang, “Dielectrophoretic forces on particles in traveling electric fields,” J. Phys. D, Appl. Phys., vol. 29, no. 2, pp. 474–482, Feb. 1996. [27] G. Fuhr, S. Fiedler, T. Müller, T. Schnelle, H. Glasser, T. Lis, and B. Wagner, “Particle micromanipulator consisting of two orthogonal channels with travelling-wave electrode structures,” Sens. Actuators A, Phys., vol. 41, no. 1–3, pp. 230–239, Apr. 1994. [28] J. Rousselet, L. Salome, A. Ajdari, and J. Prost, “Directional motion of brownian particles induced by a periodic asymmetric potential,” Nature, vol. 370, no. 6489, pp. 446–447, Aug. 1994. [29] L. Gorre-Talini, J. P. Spatz, and P. Silberzan, “Dielectrophoretic ratchets,” Chaos, vol. 8, no. 3, pp. 650–656, Sep. 1998. [30] T. B. Jones, Electromechanics of Particle. Cambridge, U.K.: Cambridge Univ. Press, 1995. [31] Y. Huang, R. Hölzel, R. Pethig, and X. B. Wang, “Differences in the AC electrodynamics of viable and non-viable yeast cells determined through combined dielectrophoresis and electrorotation studies,” Phys. Med. Biol., vol. 37, no. 7, pp. 1499–1518, Jul. 1992. [32] M. S. Talary, J. P. H. Burt, J. A. Tame, and R. Pethig, “Electromanipulation and separation of cells using travelling electric fields,” J. Phys. D, Appl. Phys., vol. 29, no. 8, pp. 2198–2203, Aug. 1996. [33] A. Ramos, H. Morgan, N. G. Green, and A. Castellannos, “AC electrokinetics: A review of forces in microelectrode structures,” J. Phys. D, Appl. Phys., vol. 31, no. 18, pp. 2338–2353, Sep. 1998. [34] F. E. H. Tay, L. Yu, A. J. Pang, and C. Iliescu, “Electrical and thermal characterization of a dielectrophoretic chip with 3D electrodes for cells manipulation,” Electrochim. Acta, 2006. in press. [35] C. Iliescu, J. Jing, F. E. H. Tay, J. M. Miao, and T. T. Sun, “Characterization of masking layers for deep wet etching of glass in an improved HF/HCl solution,” Surf. Coat. Technol., vol. 198, no. 1–3, pp. 314–318, Aug. 2005. [36] C. Iliescu, J. M. Miao, and F. E. H. Tay, “Optimization of PECVD amorphous silicon process for deep wet etching of Pyrex glass,” Surf. Coat. Technol., vol. 192, no. 1, pp. 43–47, 2005. [37] F. E. H. Tay, C. Iliescu, J. Jing, and J. M. Miao, “Defect-free wet etching through Pyrex glass using Cr/Au mask,” Microsyst. Technol., vol. 12, no. 10/11, pp. 935–939, Aug. 2006. [38] L. Yu, F. E. H. Tay, G. Xu, B. Chen, M. Avram, and C. Iliescu, “Spray coating of photoresist for 3D microstructures with different geometries,” J. Phys., Conf. Ser., vol. 34, no. 1, pp. 776–782, Apr. 2006.


Liming Yu was born in Zhejiang Province, China, in 1976. He received the B.S. degree and M.S. degree from Tsinghua University, Beijing, China, in 1999 and 2002, respectively, and the Ph.D. degree from the National University of Singapore, Singapore, in 2007. He is currently with the Institute of Bioengineering and Nanotechnology, Singapore, and also with the National University of Singapore. His research interests include MEMS technologies, bio-MEMS, and dielectrophoresis.

Ciprian Iliescu was born in Bucharest, Romania, in 1965. He received the B.S. and Ph.D. degrees from the Polytechnic University of Bucharest, Bucharest, in 1989 and 1999, respectively. He has more than 17 years of working experience in microfabrication. While pursuing his Ph.D. degree, he was with Baneasa S.A. (IC company), Bucharest, where he was involved in the design and fabrication of pressure sensors. From 1997 to 2000, he collaborated with the Institute for Microtechnologies, Bucharest, on projects related to magnetic sensors. From 2001 to 2003, he was a Postdoctoral Fellow with the Micromachines Center, Nanyang Technological University, Singapore, where he was involved in projects related to microphone, wafer level packaging of MEMS devices, and RF microrelay. Currently, he is a Senior Research Scientist with the Institute of Bioengineering and Nanotechnology, Singapore. He is the author or coauthor of more than 140 papers published in journals and conference proceedings. His current research projects are related to dielectrophoresis, electrical characterization of cells by impedance spectroscopy, transdermal drug delivery using microneedles array, and microfabricated dialysis system.

1129

Guolin Xu was born in Guangxi Province, China. He received the B.S. degree in mechatronics from Tsinghua University, Beijing, China, in 1992 and the M.S. degree in MEMS from the National University of Singapore, Singapore, in 1999. He is currently working toward the Ph.D. degree at Nanyang Technological University, Singapore. From 1992 to 1998, he was with Tsinghua University, working on automation control. From 2000 to 2002, he was with PBA Systems Pte. Ltd., Singapore. He joined the Institute of Bioengineering and Nanotechnology, Singapore, in 2002 as a Research Officer. His research interest includes microfluidic base rare cell isolation, biosample preparation, and lab-on-a-chip system using MEMS technology.

Francis E. H. Tay received the Ph.D. degree from the Massachusetts Institute of Technology, Cambridge, in 1995. He is currently an Associate Professor with the Department of Mechanical Engineering, Faculty of Engineering, National University of Singapore. He is the Deputy Director (Industry) of the Centre of Intelligent Products and Manufacturing Systems, where he takes charge of research projects involving the industry and the Centre. He was the Founding Director of the Microsystems Technology Initiative and had established the microsystems technology specialization. He is also the Medical Device Group Leader with the Institute of Bioengineering and Nanotechnology, Singapore. He was the Technical Advisor in the Micro and Nano Systems Laboratory, Institute of Materials Research Engineering, Singapore. He is also the Principal Investigator for several Agency for Science, Technology, and Research projects. The most recent one is the “MEMSWear: Incorporating MEMS Technology Into the Smart Shirt for Geriatric Healthcare,” which was widely published by the local and overseas media and well received by the public. His research areas are MEMS, biotechnology, nanotechnology, and wearable devices.