Effective Cell and Particle Sorting and Separation in Screen-Printed

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Effective cell/particle sorting and separation in screen-printed continuous flow microfluidic devices with 3D sidewall electrodes Xiaoguang Lin, Jie Yao, Hua Dong, and Xiaodong Cao Ind. Eng. Chem. Res., Just Accepted Manuscript • DOI: 10.1021/acs.iecr.6b03249 • Publication Date (Web): 05 Dec 2016 Downloaded from http://pubs.acs.org on December 5, 2016

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Industrial & Engineering Chemistry Research

Effective cell/particle sorting and separation in screen-printed continuous flow microfluidic devices with 3D sidewall electrodes Xiaoguang Lin†,‡,#, Jie Yao†,‡,#, Hua Dong*,†,‡, Xiaodong Cao*,†,‡,§

†

Department of Biomedical Engineering, School of Materials Science and Engineering, South China

University of Technology, Guangzhou, 510006, China ‡

National Engineering Research Center for Tissue Restoration and Reconstruction (NERC-TRR),

Guangzhou, 510006, China §

Guangdong Province Key Laboratory of Biomedical Engineering, South China University of Technology, Guangzhou, 510641, China

s

Supporting Information

* Corresponding authors: [email protected] (H. Dong), [email protected] (X. Cao) # Lin and Yao contribute equally to this paper

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ABSTRACT: In recent years, microfluidic dielectrophoresis (DEP) devices, as one of the most promising tools for cell/particle sorting and separation, are facing the bottleneck in the development of practical products due to the high-cost yet low-yield device manufacturing via traditional microelectromechanical systems (MEMS) and the challenge to maintain the cell viability during DEP treatment. In this paper, we demonstrate a facile, low-cost and high-throughput method to construct continuous-flow microfluidic DEP devices via screen-printing technology. The new device configuration and operation strategy not only facilitate cell/particle sorting and separation by using 3D electrodes as sidewalls of microchannel, but also improve cell viability by reducing the exposure time of cells to high electrical field gradient. Furthermore, we propose and validate a semi-empirical formula to simplify the complicated calculation and plotting of DEP spectra. As a consequence, the optimal DEP parameters and crossover frequencies can be obtained directly using our devices instead of typical electrorotation method. To evaluate the performance of screen-printed continuous flow microfluidic DEP device, a suspension containing polystyrene (PS) microspheres and erythrocytes is used as the biosample. Our results show that a high sorting efficiency (ca. 93%) with a high cell viability (hemolysis ratio < 4.8%) can be achieved, indicating the excellent performance and promising application of such devices for cell/particle sorting and separation.

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1. INTRODUCTION Scientific research in cellular and molecular biology as well as diagnostics and therapy in clinical practices, often require a large number of single-type cells.1 In nature, however, the vast majority of biosamples collected from organisms and tissues are always composed of multi-type cells. Thus, rapid and high throughput cell sorting and separation from complex, heterogeneous mixtures are strongly desirable. Unfortunately, conventional techniques such as flow cytometry or fluorescence-activated cell sorting suffer restrictions in popularization due to their high cost, bulky instrumentation and complicated operation (for example, immunolabelling).2-4 A promising alternative lies in the lab-on-a-chip systems that work on a microfluidic scale and utilize a multitude of inherent cellular properties to isolate different cell populations. Among all the strategies developed for microchip-based cell sorting and separation such as optical tweezer,5 magnetophoresis,6 acoustic waves 7 and electrical means, dielectrophoresis (DEP),8 which takes advantage of the interaction between the intrinsic dielectric properties of cells and the spatial gradient of non-uniform electric field, exhibits favorable versatility and feasibility. As a label-free technology,9 DEP has proven to separate efficiently and non-invasively various cells including bacteria, cancer cells and stem cells10-12. Despite of the great progress microfluidic DEP devices have achieved in the past decades,13 there are still several problems that need to be addressed before they can meet the end users. The first one is the high-cost yet low-yield device manufacturing induced by traditional microelectromechanical systems (MEMS), which becomes even worse in the case of disposable DEP devices for some applications especially in health care to avoid sample contamination.14, 15 Although other low-cost and easy-going methods like hot embossing, injection molding and laser photoablation can be used to fabricate microfluidic devices, they show poor capability to integrate electronic elements (for example, electrode array) that are necessary for DEP manipulation. The second one is the challenge to construct 3D electrodes in microfluidic DEP devices. Since the DEP forces actuating on the cells are proportional to the electric field gradient squared, the implementation of a strong spatial gradient throughout the microchannel is critical for efficient cell sorting and separation. The common 2D planar microelectrodes obtained via lift-off process generate a fringe-like electric field and make the DEP force only available in the vicinity of the electrodes.16 As a consequence, the dimensions of microchannels, particularly the height and width, are much limited, leading to low throughput and poor sorting performance. A well-recognized solution to this issue is the employment of 3D electrode architecture which is quite 3



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difficult to realize using evaporation or sputtering. In the literature, electroplating,17 ion-implantation,18 micromachining16 and pyrolysis19 were reported to build 3D electrodes in DEP devices20. However, most of these approaches, in our opinion, are either time-consuming or unsuitable for mass production. Recently, we demonstrated for the first time the screen-printing fabrication of microfluidic DEP devices.21 The beauty of our work was the extreme simplicity of the device fabrication procedure, i.e., the predominant components including the microelectrodes and microchannels were constructed via layer-by-layer screen printing process, which made it possible for the low-cost yet high-throughput mass production, especially considering the prevalence of such a technology in industry.22 However, the main drawback of our previous device is the long exposure time of cells of interest to the high electric field gradient and high fluid shear force imposed on cells, leading to the decrease in cell viability (see below detailed discussion).23 In this paper, we improve the DEP device by screen-printing 3D sidewall electrodes so as to realize continuous-flow cell/particle sorting and separation,24-28 and thus enhance the cell viability dramatically by reducing their exposure time to the electric field (Figure 1). In addition to the technical advance, herein we propose and validate a semi-empirical formula to simplify the complicated calculation and plotting of DEP spectra. As a result, the DEP spectra of cells and crossover frequencies can be directly measured in our devices instead of electrorotation chips.29 This theoretical advance facilitates to a great extent the identification process of DEP parameters (voltage, frequency, etc) and potentially enables the sorting and separation of various cell/particle mixtures in our DEP devices.

2. EXPERIMENTAL 2.1. Screen-printing fabrication of continuous-flow microfluidic DEP devices The continuous-flow microfluidic DEP devices were fabricated through screen-printing technology. For the sake of optical observation, transparent glass slides (GL200130-2, Guluoglass Company, Luoyang, China) were used as the substrate. The conductive carbon paste (ED423SS, Acheson, USA) was first diluted evenly and then printed on clean glass surface via a semi-automatic screen printer (F-C4050R, Fufa Company, Shenzhen, China) and a 325 mesh stainless stencil (NaWei Technology Company, Shenzhen, China). After heating at 130 ℃ for 10 min to cure the carbon electrodes, UV curable dielectric paste (YB-1300, Fufa Company, Shenzhen, China) was printed using another stencil and then exposed to UV light (Intelli-Ray 600, Uvitron Company, USA) for 5 min to form the microchannel. The alignment between electrodes and microchannel was realized through the two 4


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microscopes mounted on the screen printer. The thickness of channels and electrodes were controlled at ca. 25-30 µm. The width of main channel (i.e. electrode gap), intermediate subchannel and bilateral subchannels were set as 150 µm, 120 µm and 80 µm, respectively. Finally, a flat polydimethylsiloxane (PDMS) layer (Sylgard 184, Dow Corning, USA) with an inlet port and three outlet ports was sealed onto the glass slide using O2 plasma (PDC-32G, Harrick Plasma, USA).

2.2. Sample preparation Pig erythrocytes (ca. 8 µm in diameter) were isolated from whole blood by centrifugation. Specifically, 3 mL of whole blood was centrifuged at 1000 RPM for 10 min to get a thin layer of erythrocytes, and then the erythrocytes were washed three times with 0.9 wt% of NaCl solution at 1300 RPM for 8min. For measurement of DEP spectra, the erythrocytes and PS microspheres (ca. 9µm, Aladdin, China) were suspended in isosmotic buffer solutions with different conductivity, respectively. The isosmotic buffer solution was a mixture comprising 8.5 wt% of sucrose and 0.3 wt% of glucose, and its conductivity was adjusted to 10 µS/cm, 50 µS/cm, 100 µS/cm, 200 µS/cm, 300 µS/cm and 500 µS/cm by addition of PBS (Double-Helix Biotech, Shanghai, China) with the help of a conductivity meter (DDS-307, Lei-Ci Company, China). The pH values of these isosmotic buffer solutions were all about 7.4. The biosample for continuous-flow sorting and separation was prepared by mixing the above-mentioned erythrocyte and PS microsphere suspension together.

2.3. Measurement of DEP spectra of erythrocytes and PS microspheres A testing chamber (size: 8mm×4mm) printed on a glass slide with wavy electrodes was used to measure the sign and magnitude of DEP force as functions of input frequency and medium conductivity, or namely, DEP spectra. An arbitrary waveform generator (ED1411, Zhongce Electronics, China) was employed to create desirable electric signals. The whole device was mounted on a microscope stage (AO-KV200, Aosvi Company, China), and a 10× objective lens and a color CCD camera attached to the microscope were used to observe and record the motion of erythrocytes and PS microspheres in a bright field.

2.4. Evaluation of erythrocyte viability The erythrocyte viability after DEP treatment under various conditions including medium 5



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conductivity, time, input voltage and frequency was evaluated via hemolysis ratio (Equation 1). % =

× 100%

(1)

Where Ab and Aa are the absorbance values of cell suspensions before and after DEP treatment. Ac is the absorbance value of complete hemolysis. Erythrocytes were suspended in DI water to achieve complete hemolysis, and the absorbance value of cell suspension was measured with Multifunction Microplate Reader (Thermo3001, Thermo Company, USA).

2.5. Investigation of sorting and separation performance in continuous-flow microfluidic DEP device Erythrocyte suspension and PS microsphere suspension were separately injected into the screen-printed microfluidic DEP device using a syringe pump (74900-05, Cole-Parmer, USA) to evaluate the cell/particle sorting performance under various DEP parameters including flow rate (0.06~0.2 µL/min), medium conductivity (5~200 µS/cm), input voltage (6~20 V) and frequency (100~1000 KHz). The optimal conditions were then used for effective separation of erythrocytes and PS microspheres from their mixture suspension. The sorting efficiency (SE) was assessed by counting the cells/particles out of each subchannel after applying input signals (Equation 2 for erythrocytes, Equation 3 for PS microspheres). =

× 100%

= × 100%

(2) (3)

Nl and Ni are the numbers of erythrocytes or PS microspheres collected from lateral outlets and intermediate outlet, respectively. The data for calculation of error bar were collected in different microfluidic devices (n=5), i.e. the screen-printed devices for evaluating the sorting and separation efficiency were used only once.

3.

RESULTS AND DISCUSSION 3.1. Screen-printed continuous flow microfluidic DEP device with 3D sidewall electrodes In our earlier work,21 interdigitated electrodes were printed on the microchannel floor to trap the

cells of interest via positive DEP forces whilst push the other cells away by negative DEP forces. Although high sorting efficiency was achieved in real-time flowing test, this device showed some 6


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problems. Since the trapped cells could only be released from the electrode surface by withdrawal of the input signals after the whole cell mixture suspension passed the electrodes, these cells suffered not only long-term exposure to the high electric field gradient but also high fluid shear force, which inevitably damaged the cell viability. Besides, additional washing step was also required to bring the trapped cells out of the microfluidic device. A possible strategy to solve the above-mentioned problems is the construction of microfluidic DEP devices with 3D sidewall electrodes. Herein 3D wavy and offset castellated electrodes were printed as signal input units as well as the microchannel walls (Figure 1a). This configuration benefits the distribution of non-uniform electric field across the entire main microchannel. When applying an AC electric signal, a long-range electric field gradient can be generated along the direction of the microchannel and thus induce lateral DEP forces on the cells. As a result, lateral displacement normal to the flow direction occurs and different cells can be sorted into different subchannels depending on whether they bear positive or negative DEP forces (Figure 1b), which leads to continuous-flow cell separation. Due to the very short exposure time of cells to the 3D sidewall electrodes, the influence of high electric field gradient on the cell viability can be negligible. Figure 2a and b show screen-printed microelectrodes and microchannels that were aligned using microscopes on the screen-printer and then sealed with a flat PDMS layer with several contrapuntal pores as inlets and outlets. Compared with our previous DEP device, the screen-printed conductive carbon electrodes are thickened to ca. 25-30 µm. Despite of the relatively low resolution of screen-printing technology in comparison to MEMS, the printing qualities of microelectrode and microchannel can meet the requirement of microfluidic DEP device for cell/particle sorting and separation, as illustrated in Figure 2c and d. The distribution of electric field in the DEP device was simulated using COMSOL Multiphysics software. It is obvious from Figure 2e and f that the 3D wavy electrodes can generate a long-range and continuous electric field difference in the entire main microchannel (see electric field gradient in Figure S1, supporting information), in contrast to the weak electric field gradient generated by 2D floor electrode (Figure S2, supporting information) and sharp electric field gradient generated by 3D off castellated electrode (Figure S3, supporting information). The maximum field intensity is found at the arch part of the wavy electrodes whilst the minimum field intensity is located at the intermediate region of the microchannel. Therefore, based on the cell’s movements towards or away from the wavy electrodes, one can judge if the DEP force imposed on cells is positive or negative. It should be noted that the electric field intensity in the vertical direction is almost unchanged and this is especially favorable for 7



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cell/particle sorting and separation because all the cells or particles can experience strong enough DEP forces no matter what height they are flowing at in the microchannel.

3.2. DEP spectra of erythrocytes and PS microspheres

To evaluate the performance of screen-printed continuous flow microfluidic DEP device, the mixture of PS microspheres and erythrocytes were used as the biosample for the convenience of observation (PS microspheres are white and erythrocytes are transparent in bright field image, so additional staining treatment can be avoided.). Actually, PS microspheres were often employed as typical particles to evaluate the performance of DEP devices, although their DEP features are not the same to cells.30-38 Figure 3 displays images of PS microspheres and erythrocytes before and after applying 10 V/300 kHz signals. As can be seen, PS microspheres and erythrocytes are both randomly distributed in the microchannel prior to applying electric signals. Once subjected to non-uniform electric field, PS microspheres move to the center between the wavy electrodes where the electric field is weakest. On the contrary, the erythrocytes shift towards the wavy electrodes where the electric field is strongest. These phenomena prove that PS microspheres hold negative DEP forces while erythrocytes hold positive DEP forces.(Further experiments indicate that only negative DEP forces can be observed on PS microspheres regardless of the applied electric signal and medium conductivity while erythrocytes can be subjected to either positive DEP forces or negative DEP forces under a certain condition.) To optimize the DEP parameters to sort and separate erythrocytes and PS microspheres from their mixture suspension, it is important to plot DEP spectrum, i.e., DEP force as functions of experimental parameters. In our opinion, the main function of DEP spectrum is to illustrate the crossover frequencies of cells at which interconversion between positive DEP force and negative DEP force occurs, and their dependence on medium conductivity. In the literature, DEP spectrum is often calculated using Equation (4).39

= 2

' !" #$%&

' = 2 $%&

()* − ) ,()* + 2) , + 4 ' ()* + 2) ,' + 4 '

'

(* − ,(* ' ( + 2 ,' *

+ 2 ,

(4,

Where * and are the cell and medium permittivity, is the cell radius, ∇%& is the magnitude of 8


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the electric field gradient, Re

23 #

is the real part of the Clausius-Mossotti (CM) factor, )* and ) are

the cell and medium conductivity, and f is the input frequency. However, it is time-consuming to measure the dielectric properties of cells (permittivity and conductivity) via typical electrorotation method,40-43 and the situation becomes more complicated if the variation of medium properties in actual biosamples is taken into account. Instead of Equation 4, a semi-empirical formula (Equation 5) is derived from Newton's second law to simplify the measurement of DEP spectrum (See detailed formula evolution in supporting information). When the flow rate is zero, static cells in the microchannel moves under two kinds of forces FDEP and FR after applying electrical signals, thus =

'4 56

+ 7

(5)

Where m is the mass of cells, L is the average displacement distance of cells in microfluidic chip, t is the average time for cells to shift from initial static state to the equilibrium position with maximum or minimum electric field intensity and FR is the resistance imposed on cells during cell movement. Since the average displacement distance of cells in microfluidic chip is constant, DEP forces are proportional to the reciprocal of time squared. As a result, DEP spectra can be obtained by plotting the reciprocal of time squared (or namely, relative DEP force) as a function of input frequency and medium conductivity. Figure 4a and b show the DEP spectra of PS microspheres and erythrocytes in various medium conductivities. Only negative DEP forces are detected on PS microspheres regardless of the input signal and medium conductivity, and such forces can be increased by either reducing the medium conductivity or applying a higher frequency. Different from PS microspheres, both positive and negative DEP forces can be found on erythrocytes depending on the external conditions. The crossover frequencies at which DEP forces are zero (i.e. no movement of cells can be observed) rise with the increase in medium conductivities. When DEP force is zero, Re

23 #

in Equation 4 equals zero too. Hence, the crossover

frequency of erythrocytes can be calculated theoretically using Equation 6. (8 89 ,(8 '89 ,:;6 < 6 (= =9 ,(= '=9 , (8 '89 ,6 :;6 < 6 (= '=9 ,6

=0

(6)

Obviously, the crossover frequencies at various medium conductivities are only determined by the properties of cell and medium. In another word, the crossover frequencies should be independent on the electrode structure. Figure 4c compares the crossover frequencies of erythrocytes obtained from the DEP spectra that were measured using 3D wavy and offset castellated electrodes. It is clear that the crossover frequencies under various medium conductivities are almost the same for these two kinds of electrodes, 9



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validating the feasibility of the semi-empirical formula in plotting DEP spectra.

3.3. Effects of DEP parameters on cell viability

The presence of non-uniform electric field and the resulting DEP forces imposed on cells are the main reasons for the loss of cell viability in microfluidic DEP device. Hemolysis, the rupture of erythrocytes and the release of their contents into surrounding fluid, was used to investigate the effects of DEP parameters on the viability of erythrocytes.44 Figure 5 shows the relationship between hemolysis ratio of erythrocytes and DEP parameters including medium conductivity, exposure time to electric field, input voltage and frequency. As can be found, the increase in exposure time, input voltage and frequency can accelerate the hemolysis of erythrocytes and thus reduce the cell viability, while a lower hemolysis ratio can be achieved by raising the medium conductivity. Simultaneously, these DEP parameters also play a decisive role in the magnitude or acting time of DEP forces. This means that erythrocytes are more likely to rupture under the conditions inducing larger magnitude or longer acting time of the DEP force. When one selects appropriate DEP parameters for continuous-flow sorting and separation of erythrocytes and PS microspheres, a low hemolysis ratio is essentially required to guarantee the viability of erythrocytes, although the parameters which benefit the sorting of erythrocytes and PS microspheres always increase the hemolysis ratio of erythrocytes at the same time.

3.4. Continuous-flow cell/particle sorting and separation Similar to static experiments, we first studied the dynamic DEP behaviors of erythrocytes and PS microspheres separately to explore the influences of DEP parameters on their sorting performance in the continuous-flow device. Generally speaking, proper statistical design of experiments is needed to identify the optimal parameters for operating the chip. However, since there are so many DEP parameters (medium conductivity, input voltage, input frequency, flow rate) that could affect the sorting efficiencies and cell viability, such statistical data are time-consuming and tedious to collect. In our study, certain priority is considered during the optimization process of DEP parameters. For example, among these four DEP parameters, medium conductivity is, according to our experience, the most important one for generating enough DEP force. The second identified DEP parameters are input frequency and input voltage which affect the signs of DEP force (negative and positive) and hemolysis of erythrocytes to a large extent. It is easy to identify a certain medium conductivity and input frequency from DEP spectra 10


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(Figure 4) under which erythrocytes and PS microspheres are subjected to negative and positive DEP forces, respectively. Considering the magnitude of DEP forces and the hemolysis ratio of erythrocytes, the initial DEP parameters were set as 13 V (input voltage), 500 kHz (input frequency) and 5 µS/cm (medium conductivity). Figure 6a shows the sorting efficiencies of erythrocytes at various flow rates. The maximum sorting efficiency (87.2%) appears at the flow rate of 0.1 µL/min. This can be explained by the synergic effects of DEP forces and fluid forces. When the flow rate is low, fluid forces are too weak to overcome the positive DEP forces which trap erythrocytes to the sidewall electrodes and thus inhibit them from flowing out of the microchannel, leading to a low sorting efficiency. In contrast, strong fluid forces at a high flow rate reduce the retention time of erythrocytes in the main channel, and thus part of cells flow out from the intermediate outlet quickly without a sufficient lateral deviation, resulting in a low sorting efficiency as well. Similar explanation is also suitable to interpret the trend of sorting efficiencies at various input frequencies, input voltages and medium conductivities (Figure 6b-d) where fluid forces are constant and DEP forces are alterable. In Figure 6b， the sorting efficiencies at 600 kHz (88.3%) and 1000 kHz (89.2%) are very close, but 600 kHz is preferable in view of a lower hemolysis ratio. Figure 6c shows that the sorting efficiency reaches the peak value (93.1%) at the input voltage of 15V (Movie S1, Supporting information, SI). When the input signals and flow rate are fixed, the sorting efficiency of erythrocytes decreases with the increase in medium conductivity, as shown in Figure 6d， which can be attributed to the reduced DEP forces in medium with high conductivity. Likewise, the sorting efficiencies of PS microspheres under the same variables are shown in Figure 6e-h (Movie S2, SI). Since PS microspheres are only subjected to negative DEP forces that won’t trap them in microchannel, any strategy aimed to increase the DEP forces can always improve the sorting efficiencies of PS microspheres. It can be learned from Figure 6e-f that high input voltage and frequency as well as low flow rate and medium conductivity are beneficial for sorting and separation of PS microspheres. However, the selection of DEP parameters for erythrocytes becomes relatively complex because of cell trapping and cell viability. After comprehensive consideration of all the influential factors, the DEP parameters for real-time sorting and separation of the mixture suspension containing PS microspheres and erythrocytes were chosen as 0.1 µL/min (flow rate), 15 V (input voltage), 600 kHz (input frequency) and 5 µS/cm (medium conductivity). Under such conditions, the electric field gradient is ca. 1×105V/m in the main microchannel and the time for cells to go through the channel is ca. 9s (the length of microelectrode is 4 11



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mm). Our results show that the separation efficiency is ca.93% and the hemolysis ratio of erythrocytes is as low as ca.4.8% under a device throughput of 3.6×104 cell/particles per hour, indicative of the excellent performance of the screen-printed microfluidic DEP device in cell/particle sorting and separation (Movie S3, SI).

4. CONCLUSION In summary, we demonstrate a facile, low-cost and high-throughput method to construct continuous-flow microfluidic DEP devices via screen-printing technology. The new device configuration and operation strategy not only facilitate cell/particle sorting and separation by using 3D electrodes as sidewalls of microchannel, but also improve cell viability by reducing the exposure time of cells to high electrical field gradient. To evaluate the DEP performance of screen-printed devices, a suspension containing polystyrene (PS) microspheres and erythrocytes is used as the biosample. Prior to continuous-flow cell/particle sorting and separation, DEP spectra of these cells or particles were plotted via a semi-empirical formula in order to identify the optimal DEP parameters. The results show that high sorting efficiency (ca. 93%) can be achieved under optimized conditions. In addition, the low hemolysis ratio of erythrocytes indicates good cell viability after cell/particle sorting and separation in our devices. The future work involves two aspects: one is the separation of two real cells using our microfluidic DEP devices. The other is the combination of DEP and other separation strategies (for example, space-based separation) in order to solve the bottom-neck of DEP devices (only buffer solution with low-conductivity can be used for DEP separation). We believe that our work can further promote the application of screen-printing technology in disposable microfluidic DEP devices for clinical purposes.

■ASSOCIATED CONTENT s

Supporting Information

The Supporting Information is available free of charge on the ACS Publications website. Simulation of electric field distribution in main microchannel with 2D floor electrode and 3D off castellated electrode, and formula evolution for Eq. 5 are shown in supporting information. Real-time sorting and separation of erythrocytes, PS microspheres and their mixture suspensions in screen-printed continuous flow microfluidic DEP devices are shown in Movie S1, S2, S3 respectively. The settlement of some PS microspheres in the microchannel is mainly due to their 12


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non-uniform particle size distribution (or namely, a small part of PS microspheres are larger than 25 µm).

■AUTHOR INFORMATION Corresponding Author

*E-mail: [email protected]; [email protected]. Notes

The authors declare no competing financial interest.

■ACKNOWLEDGEMENT This work was financially sponsored by the National Natural Science Foundation of China (Grant No. 51373056,

51372085,

21574045),

Natural

Science

Foundation

of

Guangdong

Province

(2016A030311010) and Fundamental Research Funds for the Central Universities.

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Figures Figure 1

Figure 1. (a) Layer-by-layer screen-printing fabrication of continuous-flow microfluidic DEP chip with 3D sidewall configuration: (left) schematic illustration of layer-by-layer screen-printing procedure, (1) clean glass slide as substrate； (2) patterned carbon ink as 3D side-wall electrodes and conductive wires； (3) UV curable dielectric paste as microchannel； (4) flat PDMS layer with inlet and outlet ports； (middle) overall configuration of the microfluidic DEP device； (right) enlarged view of 3D microelectrodes and cross-section of the microchannel (5) 3D wavy electrodes； (6) 3D offset castellated electrodes; (7) cross-section of microchannel with 3D microelectrode as side walls; (b) continuous-flow separation of erythrocytes from a mixture with PS microspheres in screen-printed DEP microchip. When suspension solution flows into microdevice, erythrocytes bear positive DEP forces and thus move toward the electrodes, whilst PS microspheres bear negative DEP forces and are repelled from electrodes. Under the combination of lateral DEP forces and fluid forces, erythrocytes and PS microspheres can be collected from bilateral outlets and intermediate outlet, respectively.

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Figure 2

Figure 2. Images of (a) screen-printed microelectrodes, scale bar: 1cm; (b) screen-printed microchannels and microelectrodes, scale bar: 1cm; (c) individual microfluidic DEP device, scale bar: 2mm; (d) enlarged view of 3D wavy electrodes in microfluidic device, scale bar: 300µm. The finite element simulation of electric field intensity in such a device is presented in (e, side view) and (f, top view). The electrical field intensity is indicated by colors, i.e., yellow, red and brown colors represent the high, medium and low electrical field intensity. The thickness of the electrode and the distance between two electrodes were set as 25 and 150 µm, respectively. The input voltage (Vpp) was set as 10V.

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Figure 3

Figure 3. Optical images of (a, b) PS microspheres (size: ~10 µm) and (c, d) erythrocytes (size: ~8 µm) suspended in isosmotic solution (mixture of 8.5% sucrose and 0.3% glucose; conductivity: ~10 µS/cm) before (a, c) and after (b, d) applying AC signals on screen-printed wavy electrodes. Input voltage and frequency were 10V and 300kHz. Scale bar: 150µm.

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Figure 4

Figure 4. Characterization of static DEP behaviors of PS microspheres and erythrocytes in isosmotic solution with different medium conductivity: DEP spectra of (a) PS microspheres and (b) erythrocytes, crossover frequencies of (c) erythrocytes obtained from DEP spectra that were measured using wavy electrodes and offset castellated electrodes. Input voltage (Vpp): 10V.

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Figure 5

Figure 5. Hemolysis ratios of erythrocytes affected by DEP parameters in screen-printed microfluidic device: (a) input voltage (Vpp), (b) input frequency, (c) medium conductivity and (d) exposure time to the non-uniform electric field. Except for the variables in each figure, the other parameters are set as: 5µL/cm (medium conductivity), 1min (exposure time), 300kHz (input frequency), 10V (input voltage).

Figure 6

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Figure 6. Sorting efficiencies of erythrocytes (a-d) and PS microspheres (e-h) measured in continuous-flow microfluidic DEP device as functions of flow rate, signal frequency, input voltage and medium conductivity.

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Effective Cell and Particle Sorting and Separation in Screen-Printed

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