High-Throughput Separation, Trapping, and Manipulation of Single

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High-throughput separation, trapping and manipulation of single cells and particles using combined dielectrophoresis at a bipolar electrode array Yupan Wu, Yukun Ren, Ye Tao, Likai Hou, and Hongyuan Jiang Anal. Chem., Just Accepted Manuscript • DOI: 10.1021/acs.analchem.8b02628 • Publication Date (Web): 07 Sep 2018 Downloaded from http://pubs.acs.org on September 7, 2018

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Analytical Chemistry

High-throughput separation, trapping and manipulation of single cells and particles using combined dielectrophoresis at a bipolar electrode array Yupan Wua, Yukun Ren*ab, Ye Taoa, Likai Houa, and Hongyuan Jiang*ab a b

School of Mechatronics Engineering, Harbin Institute of Technology, Harbin, Heilongjiang, PR China 150001 State Key Laboratory of Robotics and System, Harbin Institute of Technology, Harbin, Heilongjiang, PR China 150001.

ABSTRACT: Microfluidic systems have been developed widely in scaled-down processes of laboratory techniques, but they are usually limited in achieving stand-alone functionalities. It is highly desirable to exploit an integrated microfluidic device with multiple capabilities such as cells separation, single cell trapping, and cells manipulation. Herein, we reported a microfluidic platform integrated with actuation electrodes for separating cells and microbeads and bipolar electrodes for trapping, rotating, and propelling single cells and microbeads. The separation of cells and microbeads can be first achieved by deflective dielectrophoresis (DEP) barriers. Trapping experiments of yeast cells and polystyrene (PS) microbeads suspended in aqueous solutions with different conductivities were then conducted, showing that both of the cells and particles can be trapped at the center of wireless electrodes using negative DEP force. Upon application of a rotating electric field, yeast cells exhibit translational movement along the electrode edges and self-rotation at an array of bipolar electrodes by applying electrorotational torque and travelling wave DEP force on the cells respectively. The current approach allows to switch the propulsion and rotation direction of cells by varying the frequency of the applied electric field. Beyond the achievements of single-cell manipulation, this system permits effective control of several particles or cells simultaneously. The integration of parallel sorting and single trapping stages within microfluidic chip enable the prospect of high-throughput cell separation, single trapping, large-scale cells locomotion and rotation in a non-invasive and disposable format, showing great potential in single cell analysis, targeted drug delivery and surgery.

Cell sorters are indispensable instruments in a variety of applications ranging from cancer diagnostics to cell based therapies, thus playing a significant role in pharmaceutical and biomedical fields. Various cell sorting techniques exist, such as density gradient separation, fluorescence-activated cell sorting (FACS)1 and magnetic-activated cell sorting (MACS)2. However, density based methods exploit differences in density between these cells3, and in MACS and FACS, multiple complicated sampling routines (for example, the addition of labels) are required and they are quite cumbersome and expensive, limiting the widespread use. An alternative potential mechanism for cell separation is dielectrophoresis (DEP). Using DEP, many scholars4 separated cells in terms of a laminate drilled with several electrode-bearing wells. Faraghat et al.5 proposed a DEP based separation device, permitting a processing rate comparable with FACS and MACS. However, those devices failed to allow the real time observation of cell performance, due to the opaque materials. Fiedler et al.6,7 sorted particles and cells based on DEP barriers and Hu et al.8 modulated the dielectrophoretic amplitude response of target cells and separated 1000cells per second, but the complicated fabrication process and the use of chemical labeling are usually required, such as the creating of microfluidic vias in top plate and the accurate alignment of two substrates. Besides DEP can also be used to trap and analyze single cells in the downstream processes. Single cell analysis is of great interest in disease and basic research9,10, such as drug responses of cells and cancer metastasis. Traditional cellular

assays usually measures average properties of a large number of cells, resulting in masking the differences between individual cells. Recently, Lab on a chip systems have emerged as a useful platform for cell analysis. Several techniques have been exploited for spatiotemporal analysis of single cells, such as optical11,12, acoustic, electric13,14, magnetic15 and microwell structures16-19. However, as for optical and acoustic methods, cells are always attracted to the point of highest energy intensity. Many researchers13,20 trapped cells at the minimum electric field to lower the physiological effect on the cells using negative DEP. However those devices are not suitable for high-throughput analysis of cells. Chiou et al11 offered an optical image driven DEP technique on a photoconductive surface, permitting parallel manipulation of large numbers of cells. But optoelectrofluidic system often requires complex fabrication process (preparation of the photoconductive substrate) and illumination components. Folch et al.18 obtained high singlecell occupancies in large arrays of microwells by optimizing the relevant parameters. Nevertheless, microwells may generate shear stress on cell surface interactions and cells can be dislodged from microwells easily during reagent exchange. We previously reported a platform for large scale trapping of single cells in terms of induced charge electroosmosis (ICEO) in a rotating electric field21, but it is not suitable for handling biological samples with high ionic strength. More importantly, it is highly desirable to design an integrated microfluidic system with multiple functions22, such as high throughput cell separation, single cell trapping and manipulation. Robbyn et al. 23,24 presented a high throughput DEP device, only allowing

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both separation of cells and single cell capture. Yet the target cells are captured at the electrode edges where the electric field is strongest, which may damage biological objects. Herein, we reported a microfluidic platform integrated with actuation electrodes for separating cells and microbeads and bipolar electrodes for trapping, rotating, and propelling single cells or microbeads by DEP. The use of a planar electrode structure facilitates the chip integration and improves the chip manipulability by incorporating the PDMS channel, allowing high performance optical detection. The accurate separation of cells or microbeads in suspensions can be first achieved by combing the pressure driven flows with deflective DEP barriers. Single particles or cells then will be trapped, propelled, and rotated bidirectionally in the floating electrode arrays using DEP forces. To our knowledge, we first immobilized large-scale single cells at the center of bipolar electrodes using nDEP forces by choosing the key conditions including solution conductivity and electrical parameters. Compared to previous micromotors whose components are mostly made from synthetic materials, such as magnetic25 particles, metals26,27, and polymers28, the yeast cells can be oriented and steered using a biohybrid approach by exploiting the powering and sensing capabilities of the cells using DEP at bipolar electrode arrays. The microfluidic system offer an unprecedented level of flexibility with which particles or cells can be separated, trapped, rotated, and propelled by fuel free electrically driven locomotion, holding particular promise for various biomedical applications, such as surgery, targeted therapy, and cell manipulation. THEORETICAL BACKGROUND Dielectrophoresis. Integrating multiple steps on a chip for sample preparation and cell manipulations is of paramount importance to fully exploit microfluidic possibilities, thus making tests cheaper, faster and more accurate29. The chip is composed of a separation channel embedded with actuation electrodes and three single cell trapping chambers equipped with wireless electrodes at the central area (Figure 1). The maximum force (perpendicular to the electrode pair) exerted by the DEP barrier on a spherical particle is FD = π r 3ε m Re  K ( w )  ∇ ( E% ⋅ E% ∗ ) − 2π r 3ε m Im  K ( w )  ( ∇ × Re( E% ) × Im( E% ) ) (1) The DEP force vector depends on the non-uniformity degree of the electric field and the magnitude of the ClausiusMossotti factor:

K( w) =(εp∗ −εm∗ ) (εp∗ +2εm∗ ) with ε ∗ = ε − j (σ where ε p and ∗

εm∗ denote

ω) .

(2)

the frequency-dependent complex

permittivities of the particles and its suspending medium,

ε

is the permittivity of liquid, σ is the conductivity, r is the particle radius. Im(K(w)) and Re(K(w)) are the imaginary and real parts of the factor K(w). The first term in Eq (1) is the conventional DEP (cDEP) force. The particle will be attracted to or propelled from the dense electric filed region depending on whether Re(K(w)) is positive or negative. The former and the latter are termed positive DEP (pDEP) and negative DEP (nDEP) force respectively .The second term in Eq (1) denotes the traveling wave DEP (twDEP) force. Depending on the polarity of the factor Im(K(w)), the particle will be directed

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towards the regions where the phases of the field component are smaller (Im(K(w))0). The transition from positive to negative DEP or vice versa is called cross-over frequency. As the medium conductivity increases, the cross over frequency of yeast cells would rise shown in Figure 2c and 2d. By adjusting the appropriate experimental parameters, such as applied frequency and medium conductivity, cells or particles of interest can be deflected in an electrode arrangement. The separation efficiency29 is evaluated by output and input concentration values achieved by manual counting in the haemocytometer. Empirical studies reveal the losses are largely in the tubing. Cell (particle) loss can be achieved from the proportion of lost yeast cells (PS microbeads) to the initial number of yeast cells (PS microbeads). Details can be found in supplementary information. Bipolar electrodes. The effect of a floating (bipolar) electrode30-32 in the electrolyte solution has been reported by many scholars with an emphasis on induced charge electroosmosis21,33,34, electrochemical reactions35,36 and dielectrophoresis21,23,24. Crooks et al.37 reported wireless electrochemical bipolar electrode microarray. The principles of bipolar electrochemistry are shown in Figure 1c. A linear potential gradient through the solution above the bipolar electrodes can be generated by applying a voltage between driving electrodes. Each bipolar electrode floats to an equilibrium potential which is controlled by the potential gradient in solution. The potential difference at the electrode/solution interface will change across the length of the electrode in terms of the potential gradient. The potential difference across every electrode will be identical when the potential gradient is linear. For single cell trapping and manipulation, two designs of the central wireless electrode arrays were offered and investigated by using BPEs to change the distribution of an AC electric field (Figure 4a). In the circle electrode design, a maximum electric field occurs at the edges of the floating electrodes while minimum field intensity is generated at the center of the floating electrodes (Figure 4b). Special attention is also paid to the influence of AC electro-osmosis (ACEO) flow, induced charged electroosmosis flow, and AC electro-thermal (ACET) flow38. The details of the above physical mechanisms can be found in the supplementary information. MATERIAL AND METHODS PDMS channel is widely used for chip fabrication since it is convenient to assemble and produce sample inlet and outlet. The use of a planar electrode structure will incorporate with PDMS channel easily and improve the chip manipulability39. The microchip with one separation area and three trapping areas was fabricated using soft lithography, presented in Figure 1. Miniaturized electrode arrays which are made from indium tin oxide (ITO) on glass substrates, are housed in the microchannel. Figure 1d displays a schematic representation of the actuation electrode arrangement in separation area. The experimental details can be obtained in the supporting information and the details of fabrication process could be found in our previous work40.

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Figure 1 (a) Schematic view of the integrated device. (b) Schematic illustration of the trapping principle in the three trapping areas; random distributed particles or cells are propelled to the center of the wireless (floating) electrodes in terms of the interactions among different forces (ICEO flow, DEP force, gravitational force and ACET flow). Each microarray contains 625 floating electrodes of the same size (D=20 µm) in each of the three trapping regions. (c) Diagram showing the electrical potential dropped linearly along the microchannel which leads to potential differences between the BPE and the solution. (d) Representation of DEP separation strategy; black circles represent the microbeads which are deflected into the central outlets since they experience negative DEP. White circles represent the yeast cells which would follow the streamlines and pass into the up or down stream at the outlets by adjusting operating conditions due to the insufficient DEP forces. (e) Microscope image of actuation electrodes in the separation area. Scale bar, 700 µm. (f) Microscope image of the bipolar electrode array in the trapping area. Scale bar, 700 µm. RESULTS AND DISCUSSION DEP separation of cells at an array of actuation electrodes. The differences in PS microbeads, nonviable and viable yeast cells based on dielectrophoresis at different medium conductivities, were exploited as the basis for separating cells and particles41. Figure 1d shows the DEP separation strategy: When the actuation electrodes are energized, the PS microbeads were deflected into the buffer stream due to the strong nDEP force near the edges of the electrodes while yeast cells would follow the streamlines and pass into the up or down stream at the outlets by adjusting operating conditions including the fluid velocity, and applied frequency to generate small DEP forces which was insufficient in deflecting or trapping yeast cells near the electrode edges. As shown in Figure2a, two designs of the actuating electrode arrays were exploited with electrode width/spacing ratios of 40µm:170µm and 40µm:400µm along the axis of symmetry and the electrodes are angled at 15° to the flow so that they converge from the full width of the channel to a gap of 70 µm and 270µm respectively. We evaluated the DEP responses of yeast cells and PS microbeads in aqueous solutions with different con-

ductivities (0.004 and 0.1S/m) using COMSOL software (Multiphysics 5.3), since different conditions faced in cell analysis and detection. Figure 2a displays the resulting electric field distribution in each design at 1MHz and A=10V in a low conductivity medium (0.004S/m). Figure 2b shows plots of electric field strength along the cut-lines in Figure 2a. The electric field strength near the electrodes edges in the device design with small electrode spacing (170µm) is higher than that in another device design with large electrode spacing (400µm). Thus using the actuating electrode array designed with small electrode spacing, particles experiencing nDEP can be expected to be efficiently deflected. Figure 2c and 2d show the plots of the real and imaginary parts of Clausius-Mossotti factors for yeast cells and PS microbeads for different suspending medium conductivities. We investigate the electric field distribution and the dielectrophoretic forces numerically. Special attention is also paid to the influence of unavoidable AC electro-osmosis (ACEO) flow and AC electro-thermal (ACET) flow. The results in Figure 2e and 2f demonstrate the DEP force plays a dominant role over the effect of ACEO and ACET during the whole process of separation. The effects of temperature rises in the separation area were discussed in Figure S2.

Figure 2(a) Numerical simulations of the electric field strength in different designs with electrode width/spacing ratios of 40µm:170µm (top row) and 40µm:400µm (bottom row) along the axis of symmetry. (b) Plots of the electric field strength along the brown cut lines shown in (a). (c) and (d) the real and imaginary parts of the CM factors as a function of frequency for yeast cells and PS microbeads at two suspending medium conductivities (0.004S/m (c) and 0.1S/m (d) respectively). (e) and (f) Frequency dependent surface maximum velocity of ACEO flow, particle cDEP motion, and ACET flow for PS microbeads and yeast cells at A=10V in aqueous solutions with different conductivities (4 mS/m (e) and 100 mS/m (f)).



When the applied frequency was adjusted from 10 kHz to 100 kHz with the voltage of 20 V peak to peak, the repelling velocity was about 0.13mm/s for PS microbeads (Figure 2e) and the absolute value of the repelling or absorbing velocity was less than 0.07mm/s for yeast cells (Figure 2e) at a medium conductivity of 4mS/m. The dielectrophoretic forces that yeast cells or PS microbeads experienced are less than or greater than the hydrodynamic force due to the streaming at flow rates ranging from 0.08mm/s to 0.12mm/s. When the flow rate is 0.1mm/s and the applied voltage is 20 V peak to peak at 80kHz, PS microbead stream follows the edges of the electrodes (Figure 3a) till hydrodynamic force exceeds the negative dielectrophoretic force (Figure2c) and yeast cells would follow the streamlines and pass into the up or down stream at the outlets (Figure 3a) due to small positive DEP force (Figure2c and 2e) which was insufficient in deflecting or trapping yeast cells near the electrode edges. Figure 3b displays superimposed images, taken from Movie S1, for the movements of yeast cells and PS microbeads over a 10 s period. 93% of PS microbeads were recovered in the central outlet and 91% of yeast cells were recovered in the up and down outlets as determined manual counting in the haemocytometer. We found that cell and particle losses were about 3.1% and 4.3% respectively. To minimize cell loss and reduce the adverse effects on the separation efficiency, a second or three pass through the device using fresh medium can be performed to collect lost cells. Thus, the yeast cells and PS mcirobeads can be effectively separated. It was found that the yeast cells were trapped at the edges of electrodes due to large pDEP force when the applied frequency was greater than 500 kHz. Yeast cells were trapped at the electrode edges by large pDEP force and PS microbeads were deflected by nDEP force into the central region at 1MHz with the voltage of 20 V peak to peak.

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Figure 3 The yeast cells follow the streamline and enter the up and down outlets and the PS microbeads are deflected into the center outlet at a medium conductivity of 4mS/m. Total flow rate is 1mm/s. (a) Sequentially captured images of PS microbeads and yeast cells moving under the interaction of DEP deflection and the sample stream at 80kHz with the voltage of 20 V peak to peak. Scale bar, 100µm. (b) The overlay images, taken from Movie S1 in the supporting information, show the movements of yeast cells and PS microbeads at 10s. Scale bar, 100µm. The separation experiments were also carried out in a medium with conductivity of 0.1 S/m. When the medium conductivity was changed, the separation parameters such as inlet velocity and applied frequency have to be adjusted for separation of yeast cells from PS microbeads. As the medium conductivity increases, the cross over frequency of yeast cells would rise shown in Figure 2c and 2d. When the applied frequency was greater than 1MHz with the voltage of 20 V peak to peak, the repelling velocity was about 0.17mm/s for PS microbeads (Figure 2f) and the absolute value of the repelling or absorbing velocity was less than 0.03mm/s for yeast cells (Figure 2f) at a medium conductivity of 0.1 S/m. The dielectrophoretic forces that yeast cells or PS microbeads experienced are less than or greater than the hydrodynamic force due to the streaming at flow rates ranging from 0.04mm/s to 0.16mm/s. Thus microbeads showing strong nDEP were funneled by the angled electrodes into the central region and yeast cells followed the streamlines and passed into the up or down trapping area in terms of small DEP force when the flow rate is 0.1mm/s and the applied voltage is 20 V peak to peak at 1MHz. The separation efficiency was larger than 90% for both yeast cells and PS microbeads by manual counting in the haemocytometer. At high flows and low amplitude of excitation, particles and cells were forced through the barriers. Single cells or particles trapping at an array of wireless electrodes. After the PS microbeads were deflected into the central trapping area due to the strong nDEP force and yeast cells followed the streamlines and passed into the up or down trapping area in terms of small DEP forces, single particles or cells then would be trapped in the wireless electrode arrays using ICEO flow or DEP forces. Two designs of the central wireless electrode arrays were provided and investigated by using BPEs to change the distribution of an AC electric field (Figure 4a). In both designs, cells experiencing nDEP were expected to be trapped at the electric field minima. The distance between the driving electrodes was 2mm. At a medium conductivity of 100mS/m, results of the simulation of the electric field in the circular and annular electrode designs are displayed in Figure 4a when the applied voltage is 10 V peak to peak at 100 kHz. The circle wireless electrodes are designed with a diameter of 50 µm and the inside and outside diameters of the annular electrodes are 25 and 50µm respectively. The distance between two adjacent electrodes was 75µm. The plots of the electric field strength along cut-lines are shown in Figure 4b. It is clear that the center regions of the wireless electrodes in the circle electrode design show a local minimum electric field and the cells undergoing an nDEP response can be trapped in the center of the wireless electrodes. As shown in Figure 4c and 4d, yeast cells located near the circle electrode rim are transported by the nDEP force toward the center of the wireless electrode where the electric field has a local


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minimum. But, for the annular electrode design, minimum field intensity is not located at the center of the wireless electrodes and it will be difficult to trap cells using nDEP in the center area of the wireless electrodes. So the circle electrode design should be selected as an ideal electrode design for trapping single cells. For single cells trapping, the diameter of the wireless electrodes and the distance between two adjacent electrodes are set as 20 and 40µm respectively. Figure 4e and 4f show that the variation of the surface averaged velocity with frequency is almost identical for the designs with different numbers (n=1, 4, and 9) of central wireless electrodes. The electric filed distributions over wireless electrodes are independent of each other, allowing us to massively trap single cells by wirelessly arranging multiple electrode arrays in the center area. Besides, for an AC field, the steady state flow decays above the RC frequency33,42 and the ICEO flow velocity will decrease by half at the double layer relaxation frequency. As shown in Figure 4f, the flow velocity is reduced by a factor of two at f=50 kHz.

Figure 4. (a) Numerical simulation of the electric field strength on logarithmic scale in the circular design (top row) and annular design (bottom row). (b) Plots of the electric field strength along cut lines in the circular and annular designs. (c) and (d) DEP velocity (in m/s) of a yeast in the x-y (c) and x-z (d) plane at 100 kHz and A=5V on logarithmic scale. (e) and (f) Surface averaged velocity of the DEP velocity of yeast cells (e) and ROT-ICEO flow (f) on the wireless electrode surface plotted against the applied frequency at A=5V for different wireless electrode numbers in a low conductivity medium (4mS/m). To test the ability of the device to capture single cells by DEP force, we conducted trapping experiments using yeast cells and PS microbeads which were suspended in KCl aqueous solutions with different conductivities of 0.1 and 0.004

S/m. Trapping experiments are first carried out by stably trapping PS microbeads and yeast cells in a low conductivity medium (4mS/m). The results can be found in supporting information (Figure S3), showing that it is impossible to trap yeast cells using DEP force in aqueous solutions with a conductivity of 4m S/m. To capture single cells using nDEP, we then conducted trapping experiments in KCl aqueous solutions with a conductivity of 0.1 S/m. When the applied frequency was 10 kHz, far lower than the characteristic frequency, the ROT-ICEO flow played a major role (Figure 5a and 5b). The PS microbeads and yeast cells would be trapped by the strong ROT-ICEO flow at the center of the flow vortex near the rim of each wireless electrode at A = 5V (Figure 5c(I) and 5d(I)). When the frequency was increased to 50 kHz, the ROT-ICEO flow could generate the stagnant region for trapping cells and particles shown in Figure 5c(II) and 5d(II). At f = 250kHz and 300kHz, the yeast cells and PS microbeads were difficult to be trapped respectively (Figure 5c(III) and 5d(III)) due to the small DEP force and weak TOR-ICEO flow (Figure 5a and 5b). As a high f far beyond the characteristic frequency was applied, the effect of ROT-ICEO flow would weaken dramatically and the DEP force dominated over other forces. At f=1MHz, the PS microbeads would be trapped stably at the center of wireless electrodes due to the strong nDEP force (Figure 5c(IV)). It is noteworthy that, at f =500 kHz, the yeast cells still experienced negative DEP force which prevailed over the ROT-ICEO flow in aqueous solutions with σ m =0.1 S/m and single yeast cells could be stably trapped using nDEP force (Figure 5d(IV)). The single cell occupancy of over 72% can be achieved at a cell density of 800 cells/µL. When f was raised to around 5MHz larger than the cross-over frequency which indicates the transition from positive to negative DEP or vice versa (Figure 2d), the yeast cells would be propelled to the edges of the wireless electrodes (Figure 5d(V and VI)) due to the positive Re[K(w)] (Figure 2d). Effects of AC electro-thermal flow are also discussed and it is clear that the influence of local joule heating is of minor importance under the current conditions (Figure 5a and 5b). The effects of temperature rises in the trapping area were discussed in Figure S2. Noted that the maximum temperature rises in all cases are less than 2 °C, a value that is very safe for biomedical applications. Thus, by adjusting the applied frequency, it is possible to trap yeast cells using nDEP force in aqueous solutions with σ m =0.1S/m (Figure 5d (IV)). By pumping aqueous solutions (0.1S/m) in the inlets of the side channels (s-I1 and s-I3), the nDEP force could hold the yeast cells against a fluid flow with a velocity of 15µm/s at f = 250 kHz and A=5V. The current platform shows great potential in cell immobilization, exchange of reagents and live imaging of cell morphology43. The common trap microfluidic designs often suffer from low single cell trapping efficiencies, such as hydrodynamic sieve-like traps(~50%)44, micropatterns only (~40%)45 or micro-wells (~85%)18 as well as the approach we reported here (72%). So capture efficiency of cells into microfluidic chambers is not very high. However the trapping efficiency can be compensated for by using a large array of trapping sites to achieve sufficient single cells for experimental analysis. Moreover, to further enhance the single cell trapping efficiency, an optimization study should be conducted by thoroughly investigating the key parameters includ-



ing, cell densities, wireless microelectrode dimensions, and the applied signals in terms of cell types.

Figure 5 Investigation of yeast cells and PS microbeads in in aqueous solutions with a conductivity of 0.1 S/m. (a) Frequency dependent surface averaged velocity of ACET flow, ICEO flow, cDEP motion, and twDEP motion for a PS microbead with a diameter of 5µm at A=5V. (b) Frequency dependent surface averaged velocity of ICEO flow, cDEP motion, and twDEP motion for yeast cells at A=5V. (c) Images showing the motion of PS microbeads in KCl aqueous solutions at different frequencies (from Movie S2). Scale bar, 20µm. (d) Images displaying the rotation and propulsion of yeast cells at different frequencies (from Movie S3). Scale bar, 20µm. Furthermore, the current chip can also enable cell pairing by adjusting the concentrations of cells and particles. Single cellcell pairing is a crucial step for in vitro analysis of intercellular interactions which is significantly important for immune system, tissue regeneration, and cell differentiation39,46-48. We could achieve 2-particle capture efficiency up to 55.5% at a particle density of 1000 particles/µL in aqueous solutions with σ m =0.1S/m upon application of a traveling wave signal of f = 1MHz and A = 5V (Figure S4a). At a cell density of 1200 cells/µL, nDEP force yielded yeast cell pairing efficiency of over 60% (Figure S4c) by applying a traveling wave signal of f = 500 kHz and A = 5V. We further conducted multiple-cell capture at a high cell density (2000 cells/µL), 3-cell capture efficiency of up to 50% were obtained at f = 500 kHz and A = 5V in aqueous solutions with σ m =0.1S/m (Figure S4d). The approach also shows great application potential in cell-cell fusion where two or more cell types are merged into a hybrid cell. DEP manipulation of single cells at an array of wireless electrodes. The current platform also permits single cells or particles to be manipulated non-invasively using DEP force at a wireless electrode array, showing greater potential in single cell analysis, targeted drug delivery and surgery. Previously, the components of micromotors are mostly made from synthetic materials, such as magnetic particles25, metals, and polymers28. Using a biohybrid approach49, the current system

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exploited the powering and sensing capabilities of the cells by incorporating nDEP with bipolar electrode arrays. For the specific application of automatic, wirelessly controlled manipulation of cells or micro-objects, control of both rotation and translation is required. Rotational motion is of importance where the chip is designed with a nonsymmetrical geometry for engagement and trapping. Our research is focused on the bidirectional propulsion and rotation of yeast cells which are actuated by DEP forces at a bipolar electrode array. By employing a rotating electric field, electrorotational torque is produced to control rotational motion and the twDEP force is used to produce bidirectional propulsion. The use of wireless electrode array enables parallel and distributed operations to be conducted by swarms of yeast cells, showing great potential in remotely powering cells by external electric fields to undergo navigational steering along different paths. DEP manipulation of single yeast cells were carried out in KCl aqueous solutions with different conductivities of 0.1 and 0.004 S/m. When f was increased to much higher than the characteristic frequency (around 50 kHz), the effect of ROTICEO flow weakened dramatically and the DEP force dominated over the drag force arising from ACET flow and ROTICEO flow in a low conductivity medium (0.004S/m). The yeast cells would be attracted to the edges of wireless electrodes (Figure S3d (III and IV)) due to the positive Re[K(w)] when f was raised to much higher than the cross-over frequency (around 80 kHz). At f=500 kHz, the yeast cells were propelled around the edges of the wireless electrodes in the opposite direction to the field travel (Figure S3d (III)) due to the twDEP force originating from the positive Im[K(w)], whereas the yeast cells would be propelled in the same direction as the field travel around the wireless electrode edges (Figure S3d(IV)) because of the negative Im[K(w)] at 20 MHz. The electrorotational torque still made the cell revolve on its axis and yeast cells can rotate in both a clockwise and counterclockwise direction. The rotational speeds and the propulsion velocities of cells are obtained according to the values determined for ten different cells. Ten tests were conducted using ten different cells instead of the same cell and the error bars represent the standard deviation of ten tests. Thus the yeast cells can be rotated and propelled bi-directionally around the wireless electrode edges by adjusting the applied frequency. To demonstrate the applicability and generality of the control approach, yeast cells are then oriented and steered in aqueous solutions with a conductivity of 0.1 S/m. At f=50kHz, the electrorotational torque forced yeast cells into clockwise rotation counter to the field rotation direction due to positive Im[K(w)] (Figure 2d) and the dominant ROT-ICEO vortices propelled yeast cells to the center of the wireless electrodes shown in Figure6a. Despite the electrorotational torque still triggered the clockwise rotation of the yeast cell on its axis at f=500 kHz, yeast cells would be propelled to the center of the wireless electrodes by the dominant nDEP forces instead of ROT-ICEO flow (Figure 5b and 6b). The rotational speeds and the propulsion velocities of cells in KCl aqueous solutions with different conductivities of 0.1 and 0.004 S/m are shown in Figure 7a and 7b. The rotation performance of yeast cells in aqueous solutions with σ m =0.1S/m indicates that the rotation speed of yeast cells (7.5 r/min) on the electrode surface is lower than that (22.5 r/min) in the center area where bipolar elec-


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trodes are absent at f=500 kHz. The reason for this is that the electric field strength in the center region where floating electrodes are absent is higher than that in the center of the floating electrode as shown in Figure 4b. At f=5 MHz, yeast cells were propelled around the edges of the wireless electrodes in the opposite direction to the field travel (Figure 6c) due to the twDEP force originating from the positive Im[K(w)], whereas the yeast cells would be propelled in the same direction as the field travel around the wireless electrode edges (Figure 6d) because of the negative Im[K(w)] at 40 MHz. The electrorotational torque simultaneously triggered the rotation of yeast cells on its axis and yeast cells could rotate in both a clockwise and counterclockwise direction upon switching the frequency of the electric field (Figure 6c and 6d). Yeast cells could be propelled at a relatively high speed of about 94.2µm/min around the edges of the wireless electrodes in the same direction to the field travel at A=5V and f=40 MHz. One other thing to note here is that yeast cells experiencing pDEP forces were rotated around the electrode edges at almost the same speed as that in the center area where floating electrodes are absent. Thus the high electric field strength distributed above the rim of the floating electrode has slightly effect on the rotation performance of yeast cells since the center region of the wireless electrodes shows a local minimum electric field.

Figure 6 Time-lapse images of motions of yeast cells in an aqueous solution with a conductivity of 0.1 S/m (from Movie S3). (a) Sequence of images illustrating the rotational motion of yeast cells at the center of wireless electrodes (A=5V, f=50 kHz). The electrorotational torque forced yeast cells into clockwise rotation and the dominant ROT-ICEO vortices propelled yeast cells to the center of the wireless electrodes. Scale bar is 20µm. (b) Sequence of images illustrating the rotational motion of yeast cells at the center of wireless electrodes and in the center area where floating electrodes are absent (A=5V, f=500 kHz). The electrorotational torque forced yeast cells into clockwise rotation and the dominant nDEP forces propelled yeast cells to the center of the wireless electrodes. (c)

Sequence of images showing the clockwise rotation of yeast cells on its axis and propulsion around the wireless electrode edges (A=5V, f=5 MHz). The electrorotational torque forced yeast cells into clockwise rotation and the dominant pDEP forces attracted yeast cells to the center of the wireless electrodes. (d) Sequence of images showing the counterclockwise rotation of yeast cells on its axis and propulsion around the wireless electrode edges (A=5V, f=40 MHz) (arrows indicate the direction of the motion). Furthermore, the trajectories of yeast cells can be changed by designing floating electrodes with different shapes, such as square, and triangular. The time-lapse images in Figure 7c, 7d and the corresponding movie (Movie S4) show that the yeast was propelled around the square electrode edges in the opposite direction to the field travel at a speed of 100µm/min due to twDEP force and pDEP force in KCl aqueous solutions with a conductivity of 0.1 S/m (A=5V, f=2 MHz). Several yeast cells could also be propelled around the triangular floating electrode edges and time lapse images of the propulsion of yeast cells were represented in Figure S5 and Movie S5. We guided the yeast cells along the electrode edges by applying electric field at an average speed of and the yeast cells were found able to maintain its initial speed as long as we switched on the power. The self-propelled, electrically activated yeast cells’ behaviors show good durability. It is often clear that one cell type can be distinguished from others in terms of the associated behaviors, showing great potential in identifying cell types. Besides, cells can be used as smart drug delivery agents by encapsulation of drugs inside cells or attached to the surface and subsequent transportation through the body. The cells can be navigated and perform specific tasks inside enclosed organ on a chip microfluidic devices.

Figure 7 Moving performance of yeast cells in KCl aqueous solutions with different conductivities of 0.004 S/m (a) and 0.1 S/m (b) at different frequencies. The negative velocities denote the counterclockwise propulsion in the opposite direction to the field travel, while the negative roation speeds represent the rotational motion counter to the field rotation direction. The error bars represent the standard deviation of ten tests. (c) Trajectories of individual yeast cell showing the propulsion of yeast cells around the square electrode edges in the opposite direction to the field travel in KCl aqueous solutions with a conductivity of 0.1 S/m (A=5V, f=2 MHz). Scale bar, 10µm. (d) Superimposed sequence of images showing the movements of yeast cells over 1 min (from Movie S4) around the electrode edges. Scale bar, 10µm Our experimental investigation on the behavior of yeast cell by varying salt concentration and the frequency of the applied



voltage demonstrates that yeast cells exhibit translational movement along the electrode edges and self-rotation in terms of twDEP force and electrorotational torque. The concept of manipulating cells or particles by DEP at a wireless electrode array paves the way to micromotor systems that are capable of bidirectional locomotion and rotation of large-scale single cells in a convenient and ingenious way. By manipulating single cells in a continual, automated manner in organ-on-achip systems, the performance of current organ-on-a-chip models in drug screening can be further improved. The multifunctional cell based manipulation approach could be useful for medical applications in organ-on-a-chip systems. CONCLUSIONS To develop an integrated microfluidic system with multiple functionalities such as cells separation, single cell trapping, and cells manipulation, we developed a microfluidic platform integrated with actuation electrodes for separating cells and microbeads and bipolar electrodes for trapping, rotating, and propelling single cells and microbeads. The high-throughput separation of cells and microbeads can be first implemented by combing the pressure driven flows with deflective DEP barriers. The separation efficiency was larger than 90% for both yeast cells and PS microbeads by manual counting in the haemocytometer. By combining the bipolar electrodes, a BPE array can be capable of trapping single cells by wirelessly controlling the electric field distribution. Trapping experiments were conducted using yeast cells and PS microbeads suspended in aqueous solutions with different conductivities and the single yeast cell occupancy of over 72% can be achieved at a cell density of 800 cells/µL at f =500 kHz and A=5V in aqueous solutions with σ m =0.1 S/m, indicating that yeast cells could be trapped using nDEP force by adjusting the applied frequency. Yeast cells were also oriented and steered bidirectionally using a biohybrid approach by exploiting the powering and sensing capabilities of the cells using DEP at bipolar electrode arrays. Yeast cells could be propelled at a speed of about 94.2µm/min around the edges of the wireless electrodes in the same direction to the field travel at A=5V and f=40 MHz. The multifunctional cell based manipulation approach could be useful for medical applications and takes a step toward in vitro wireless manipulation of large-scale cells. The integrated microfluidic platform is easy to fabricate by conventional photolithographic techniques and shows excellent optical quality, permitting complete miniaturized cells processing and analysis, surgery, and targeted therapy on a chip.

ASSOCIATED CONTENT Supporting Information The Supporting Information is available free of charge on the ACS Publications website.

AUTHOR INFORMATION Corresponding Author * Email: [email protected] and [email protected], Phone.: +86 451 86418028; Fax: +86 451 86402658.

Notes The authors declare no competing financial interest.

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ACKNOWLEDGMENT This work was supported by the National Natural Science Foundation of China (Grant No. 11672095, and 11702075), and the Foundation for Innovative Research Groups of the National Natural Science Foundation of China (Grant No. 51521003). Y. Wu’s greatly appreciates Y. Meng’s continuous support and encouragement.

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High-Throughput Separation, Trapping, and Manipulation of Single

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