AC Electrothermal Circulatory Pumping Chip for Cell Culture - ACS

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AC Electrothermal Circulatory Pumping Chip for Cell Culture Qi Lang,† Yanshuang Wu,§ Yukun Ren,*,† Ye Tao,† Lei Lei,*,§ and Hongyuan Jiang*,†,‡ †

School of Mechatronics Engineering and ‡State Key Laboratory of Robotics and System (HIT), Harbin Institute of Technology, West Da-zhi Street 92, Harbin, Heilongjiang P. R. China 150001 § Department of Histology and Embryology, Harbin Medical University, Xuefu Road 194, Harbin, Heilongjiang, P. R. China 150081 ABSTRACT: Herein we describe a novel AC electrothermal (ACET) fluidic circulatory pumping chip to overcome the challenge of fluid-to-tissue ratio for “human-on-a-chip” cell culture systems. To avoid the deleterious effects of Joule heating and electric current on sample cells, a rectangular microchannel was designed with distantly separated regions for pumping and cell culture. Temperature variations were examined using a commercial thermocouple sensor to detect temperature values in both pumping and culture regions. To generate a sufficient ACET circulatory pumping rate, 30 pairs of asymmetrical electrodes were employed in the pumping region; generated ACET velocity was measured by fluorescent microparticle image velocimetry. The benefits of our pumping chip were demonstrated by culturing human embryonic kidney cells (HEK293T) and human colon carcinoma cells (SW620) for 72 h with an energized voltage of 3 V and 10 MHz. Cells grew and proliferated well, implying our ACET circulatory pumping chip has great potential for cell culture and tissue engineering applications. KEYWORDS: AC electrothermal, circulatory pumping chip, cell culture, human-on-a-chip, microfluidics

1. INTRODUCTION One key reason for the high cost of pharmaceutical development is the lack of a standardized miniature platform allowing for drug testing on human tissue samples.1,2 “Humanon-a-chip” systems, which incorporate circulatory systemintegrated multiple organ chips capable of mimicking human physiology, have vast potential to improve the efficiency of pharmaceutical exploitation. These systems have been recognized as a novel approach to avoid potential dangers to clinical trial participants3 and overcome poorly predictive evaluations caused by phylogenetic distance between animals and humans.4 An advanced human-on-a-chip system should satisfy two strict requirements: (i) cells and tissues should be cultured at the same fluid-to-tissue ratio as found in human body,5 and (ii) automated culture systems with integrated fluid control and real-time information acquisition should be miniaturized such that they can be placed inside an incubator.6 The requirement for suitable fluid-to-tissue ratio ensures metabolic products from cells are maintained at appropriate levels within the culture medium. Otherwise, the dead volume caused by use of external, commercially available conventional syringes and peristaltic pumps becomes orders of magnitude larger than the total blood volume, breaking the fluid-to-tissue ratio. In this manner, the level of drugs or metabolic products from tissue cells is diluted during drug testing or other experiments.7,8 Presently, human-on-a-chip systems utilize a mass of culture chambers, valves, sensor chips, tubing, medium reservoirs, and external peristaltic pumps.8−10 An overwhelming amount of medium is brought into the system by external reservoirs and peristaltic pumps, whereas the small number of cells in the system allows for only a tiny fraction of overall circulating medium volume. To reduce the ratio of fluid-to-tissue to its correct value, the channels, reservoirs, and pumps contained © 2015 American Chemical Society

within the system should be significantly minimized; this calls for integration of a circulation system onto the chip. Chip-based micropumps have been extensively studied. Among them, mechanical reciprocating displacement micropumps have been commonly adopted for organ-on-a-chip research. Reciprocating pumping is based on exertion of reciprocate pressure forces on fluid by displacing polydimethylsiloxane (PDMS) elastic membranes. Reciprocating micropumps have been integrated in microdevices by Schimek,11 Wu,12 et al.; however, complex pneumatic configurations, difficult fabrication, and limited lifespan of membranes are continuing obstacles to the flexibility and stability of human-ona-chip systems. Additionally, these systems are unusable if the culture medium contains suspended particles13 such as red blood cells. Electrohydrodynamic (EHD) micropumps are a desirable solution to this problem, as they involve no moving parts or valves, are easy to fabricate, and are compatible with microsystem integration.14−16 While DC EHD micropumps present the shortcomings of high voltage, reaction, biofouling, etc., the recent emergence of AC EHD micropumps has shown great promise. AC EHD micropumps include AC electroosmosis (ACEO)17−20 and AC electrothermal (ACET)21,22 mechanisms. Arising from the interaction of a tangential electric field and induced charge, ACEO is only effective for lowconductivity fluids and, thus, is unsuitable for human body fluids. However, ACET emerges from the interaction of an electric field and temperature gradient, and is especially suited for driving a wide range of high conductivity fluids.13,23 Moreover, ACET presents many advantages such as low Received: September 19, 2015 Accepted: November 12, 2015 Published: November 12, 2015 26792

DOI: 10.1021/acsami.5b08863 ACS Appl. Mater. Interfaces 2015, 7, 26792−26801

Research Article

ACS Applied Materials & Interfaces

The electric field in the fluid, generated by electrodes with AC electric voltage, can be determined by Laplace’s equation:38,39

voltages, low power consumption, and high frequencies, which are appropriate for biofluids.13,24 In this paper, an ACET circulatory pumping chip (CPC) with asymmetric electrode array (Figure 1) was designed and

∇2 V = 0

(2)

in which V denotes root-mean-square voltage, and electric field strength is calculated as15 (3)

E = −∇V

The Navier−Stokes equation is applied for an incompressible fluid at low Reynolds number in the electrothermal body force for a steady flow as follows:34 −∇p + η∇2 u ⃗ + ⟨Fet⟩ = 0

(4)

where p is pressure, η is viscosity of the fluid, u⃗ is the fluid velocity vector, and ⟨Fet⟩ is the time-averaged electrothermal force per unit volume, which can be expressed as27

Figure 1. Schematic of ACET circulatory pumping chip. The region within the dotted line contains 30 pairs of electrodes is the pump area, while chamber “C” is the region for cell culture.

⟨Fet⟩ =

ε(α − β ) 1 1 · (∇T ·E)E − εα|E|2 ∇T 2 1 + (2πfε /σ )2 4

(5)

where α and β are calculated as

fabricated for cell culture. The chip is composed of a cell culture chamber, a medium reservoir chamber, and microchannels integrated with pumping electrodes. The culture region (“C” in Figure 1) and pump region with medium reservoir chamber (“R” in Figure 1) have been placed at either side of the chip to avoid potential thermal stress on sample cells caused by the Joule heating effect at the pumping region. We investigated ACET pumping performance using cell culture medium with a conductivity of 1.5−2 S/m, which is higher than the fluid conductivities reported in traditional ACET experimental studies incorporating a salt medium (usually KCl with a conductivity of 1 S/m25,26). High-conductivity medium was effectively driven, and two representative types of cells were successfully cultured in the ACET CPC. Both human embryonic kidney cells (HEK293T) and human colon carcinoma cells (SW620), that is, normal and cancer cells, grew well for 72 h in the ACET CPC. Use of an ACET pump has not been previously reported in cell culture or tissue engineering applications.

22

α=

∂ε /∂T = −0.004(K −1) ε

(6)

β=

∂σ /∂T = 0.02(K −1) σ

(7)

2.2. Numerical Modeling. Mathematical simulations were performed using COMSOL Multiphysics 4.4 (Comsol, Inc.) to evaluate the performance of the ACET CPC. A 3D model of the same size as the experimental chip was established. Appropriate electrostatics, heat transfer, and laminar flow modules were chosen to solve equations for electric, thermal, and velocity fields, respectively. Electric field distribution along the microchannel was obtained by applying a sinusoidal signal (amplitude of 2.5 to 5 V) to 30 pairs of asymmetric microelectrodes; simulations were performed at a frequency of f = 10 MHz. As the double layer chip was acutely suppressed in high conductivity medium, the effect of ACEO can be neglected; the PDMS layer and glass slide are electrically insulated. Fluids with conductivities of 1.5, 1.7, and 2.0 S/m were used for simulations. Heat transfer occurred throughout the whole system, which contained a glass slide (kglass = 1.1 W/mK34), electrolyte solution (kfluid = 0.6 W/mK40), and PDMS layer (kPDMS = 0.2 W/mK30). Temperatures on the external surface of the PDMS cover, glass substrate, and medium had the same value of 37 °C (310.15 K). Laminar flow with electrothermal force as an external body force was solved for electrolyte solution with a temperaturedependent viscosity. The air/liquid interface in the large chamber was given zero pressure. For the incompressible Navier−Stokes simulation, a no-slip boundary condition was applied at fluidic microchannel walls. Specific parameters used in the numerical study are shown in Table 1. Figure 2a (not drawn to scale) shows a sectional view of one microchannel with reservoir chamber located at A−A in Figure 1, indicating the boundary conditions for electrical, thermal, and fluidic settings. Shown in Figure 2b−d are simulation results for electrical potential, temperature, and velocity fields of an electrode pair located at B−B in Figure 1. As shown in Figure 2c, temperature is nonuniform; the highest temperature was observed ∼150 μm

2. THEORETICAL BASIS 2.1. Alternating Current Electrothermal. When a nonuniform electric field is applied to the medium, a nonuniform temperature rise (i.e., temperature gradient) is generated by Joule heating,27−30 thereby creating fluid permittivity and conductivity gradients in ACET micropumps. As electric properties of fluids are nonuniform, induced free charges move in the nonuniform electric field. Finally, this nonuniform electric field exerts body force on the fluid to induce ACET flow. Nonuniform electric fields for ACET pumps can be generated by an asymmetric electrode array,25,31−33 asymmetric coplanar electrodes array,34 T-shaped electrode array,13 and three-dimensional (3D) electrode structures.35 Electric field distribution throughout the fluid region and Joule heating can be expressed by the energy balance equation as22,34,36,37 1 k∇2 T + ⟨σE2⟩ = 0 (1) 2 where k is the thermal diffusivity coefficient, T is temperature of the fluid, σ is electrical conductivity of the fluid, and E is the applied electric field. 26793


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pump areas are located on opposing sides of the chip to reduce temperature rises within the cell culture chamber. A small piece of PDMS sheet was used to cover the culture chamber to reduce medium evaporation. The electrodes used were 4 mm long, 0.2 μm (200 nm) thick, and 120 μm/500 μm wide, distributed with a 120 μm in-pair separation and 80 μm between-pair gap separation, as shown in Figure 2b. The glass slide with ITO electrodes was fabricated using photolithography,42 whereas the PDMS layer was generated using a conventional soft-lithography procedure. Microfabrication of the glass slide with ITO electrodes required the following steps. First, negative dry film photoresist (RistonSD238, Dupont) was laminated onto a glass slide with 200 nm thick ITO film (ITO8008−1, Kaivo). Next, the photoresist was selectively exposed by a photomask designed with AutoCAD software (Autodesk), and developed in 1% Na2CO3 water solution (w/w). The ITO glass slide with patterned photomask was then soaked in 20% HCl acid for 45 min. Finally, photoresist was stripped using acetone. For the PDMS layer, a poly(methyl methacrylate) (PMMA) sheet (1 mm thick) was cut using a Laser cutter (K3020, Huitian) to form the channel and chamber mold, which was pasted onto a glass slide. Next, an elastomer base and PDMS curing agent (Dow Corning) were mixed at a ratio of 10:1 (v/v), then poured onto the mold, and degassed in a vacuum drying chamber. The PDMS mixture was cured at 80 °C in an oven for 90 min, and then chambers were mechanically punched by biopsy punch (PSAM14, Acuderm). Finally, the glass slide with ITO electrodes and PDMS layer were aligned and bonded together by treatment with oxygen plasma (ZEPTO, Diener). Before culturing cells in the CPC, the CPC unit was sterilized by soaking in 70% ethanol for 45 min, then subsequently applying UV irradiation for 45 min. 3.2. Fluid Temperature Measurement. A commercial thermocouple temperature sensor (8620, TASI) was used to measure realtime temperature within the cell culture chamber and medium reservoir. Sensor electrodes were submerged into the chambers, which were filled with KCl solution of 2.0 S/m. A voltage was applied to the device and solution in a 37 °C incubator for 60 min, and temperatures within the two chambers were recorded every 10 min. AC peak voltages of 2.5, 3, 3.5, 4, 4.5, and 5 V were energized in individual experiments. 3.3. Evaluation of Fluid Velocity. The electrical conductivity of a cell culture primarily depends on chloride saline solution; the conductivity of most culture medium ranges from 1.5 to 2.0 S/m. Therefore, aqueous KCl solutions with conductivities of 1.5, 1.7, and 2.0 S/m were utilized as fluids to be actuated by the ACET pump. AC peak voltages of 2.5, 3, 3.5, 4, 4.5, and 5 V were generated at a frequency of 10 MHz on electrodes using a function generator (TGA12104, TTI) and adjusted by an oscilloscope (TDS2024,

Table 1. Property Values Used in the Numerical Study parameters σ(T0) ε(T0) ε0

value

T0 f Vp phase

1.5−2 (S/m) 80ε0 8.85 × 10−12 (F/m)40 310.15 (K) 1 × 107 (Hz) 2.5−5 (V) π

ρm η(T0) kfluid kglass kPDMS H W d1/d2 x1/x2 n d1 d2

0.96 (g/mL)41 0.89 (mPa s)41 0.6 (W/mK) 1.1 (W/mK) 0.2 (W/mK) 1 (mm) 3 (mm) 10:3 (mm) 120:500 (μm) 30 120 (μm) 800 (μm)

implication fluid conductivity dielectric permittivity vacuum permittivity ambient/reference temperature frequency amplitude of the drive voltage phase difference of signal on asymmetric electrodes mass density of fluid sample dynamic viscosity at reference temperature thermal conductivity of fluid sample thermal conductivity of glass slide thermal conductivity of PDMS height of the microchannel width of microchannel diameters of circles width of asymmetric electrodes number of electrode pairs gap between asymmetric electrodes gap between electrode pairs

above the narrow electrode. Figure 2d illustrates the magnitude of electrical potential decreases rapidly as distance from the electrode increases. Observed velocity distribution and vectors suggest electrothermal flow moves from the narrow electrode toward the broad one; further, the largest flow velocity appears at the region with the largest temperature gradient. A vortex occurs in the middle of the channel, as determined by temperature distribution.

3. MATERIALS AND METHODS 3.1. Design and Fabrication of ACET Circulatory Pumping Chip. The primary design requirement for the CPC is to reduce temperature rises in the cell culture area to avoid detrimental effects to cells. Figure 3a shows the design of the CPC. The chip consists of a glass slide with 30 pairs of indium−tin−oxide (ITO) electrodes and a 6 mm thick PDMS layer containing the fluid channel (3 mm width, 1 mm height), cell culture chamber (10 mm diameter), and medium reservoir (3 mm diameter). The medium reservoir chamber was designed to maintain the pressure balance of medium. Cell culture and

Figure 2. Numerical simulation of CPC. (A) Boundary conditions of electrical, thermal, and fluidic domains, where U is electric potential, v is velocity of the fluid, and T is temperature. (B−D) Two-dimensional (2D) sectional plots of 3D simulation results for a micropump at Vp = 3 V, σ = 1.7 S/m, and f = 10 MHz. (B) Electric potential distribution and isopotential lines, (C) temperature distribution and isothermal lines, and (D) surface and arrow plot of ACET velocity. 26794


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Figure 3. Chip design and fabrication. (A) Photograph of CPC in a plastic culture dish. (B) Microscopic view of fabricated microelectrode array. (C) CPC fabrication process of (I) etching electrode in lithography, (II) casting PDMS channel, and (III) bonding. (a) ITO glass slide was coated with an unexposed photoresist layer. (b−d) Photoresist was selectively exposed by photomask and developed. (e) ITO glass slide was etched with HCl acid. (f) Photoresist was stripped off using acetone. (g−h) PMMA sheet (1 mm thick) was cut to form the channel and chamber mold, then pasted onto a glass slide on which PDMS was poured and cured. (i) Holes were punched. (j) PDMS channel was bonded to glass slide with ITO electrodes. Tektronix). Suspended within the KCl solution, 0.5 μm yellow fluorescent microspheres (69A1−2, Molecular Probes) facilitated tracing and calculation of fluid velocity. Fluid-suspended microspheres were captured under a fluorescent microscope (BX53, Olympus) and recorded at a rate of one frame per second using a CCD camera (RETIGA2000R, Qimaging). Fluid velocity experiments were conducted at ambient temperature of 20 °C (293 K). The microscope was focused 300 μm above the bottom of the glass surface (position “F” in Figure 1) to acquire images and record videos. 3.4. Cell Culture and Culture Medium. Human embryonic kidney cells, HEK293T (ATCC), were maintained at 37 °C with 5% CO2 in Dulbecco’s Modified Eagle Medium (DMEM, 11965−084, Life Technologies) containing 10% (v/v) heat-inactivated fetal bovine serum (FBS, 04−001−1A, Life Technologies) and antibiotics (100 U/ ml penicillin and 100 U/mL streptomycin, P/S, Life Technologies). Human colon carcinoma cells (SW620, ATCC) were cultured under the same conditions as HEK293T. 3.5. Measurement of Electrical Conductivity of Medium. The culture medium in our study (DMEM, 10% FBS, 1% P/S) was used to culture HEK293T at a density of 100 000 cells per 400 μL of culture medium in 48-well plates. Samples of culture supernatant were collected at 24, 48, and 72 h, during which time culture medium went unchanged. The electrical conductivity of each medium sample was tested using a commercial conductivity sensor (DDSJ-308A, Leici). A preliminary analysis of additional medium types included conductivity measurements of DMEM with 15% FBS (and 1% P/S) and DMEM/ F-12 (11320−033, Life Technologies) with 10% FBS (and 1% P/S). Experiments were conducted in four replicates. 3.6. Cell Culture in the Circulatory Pumping Chip. First, 400 μL of culture medium was injected into the CPC. Next, a high density of cells (100 000 cells in 10 μL) was gently seeded into the CPC cell

culture chamber. To ensure an even distribution of cells within the culture chamber, the cell suspension was gently stirred with a pipet. After seeding, the device was gently placed in a plastic culture dish (430167, Corning) and put into an incubator for 4 h, at which time a peak voltage of 3 V was applied with an AC function generator. The AC function generator was stopped to perform culture medium changes; medium was carefully changed in the sterile laminar flow cabinet using a pipettor. After the CPC was replaced into the incubator, the AC function generator was reactivated. Culture medium was changed every 24 h; at this time, culture supernatant was also collected for subsequent measurement of pH values using a commercial pH sensor (PB20, Sartorius). For control and comparison, cells were seeded at the same concentration into the CPC without use of AC signal or cultured in 48-well plates (CLS3338, Corning). Growth status of HEK293T and SW620 cells in both experimental and control groups were examined under a microscope (TE2000-U, Nikon) every 24 h. Experiments were conducted in three replicates. Pumping performance of the CPC was verified after 72 h of continued use. Cell culture medium was removed and replaced by KCl electrolyte with a conductivity of 1.7 S/m for pumping experiments. Electrical signal with peak voltages of 2.5 to 5 V and a frequency of 10 MHz were used to drive the fluid. Fluorescent polystyrene microspheres with a diameter of 0.5 μm were chosen to reveal electrothermal flow rate. 3.7. Cell Viability. HEK293T and SW620 cells were seeded at equal densities into CPC with and without AC signal, as well as 48well plates. A cell viability imaging kit (R37610, ThermoFisher) was used according to manufacturer’s instructions to examine cell viability following 24, 48, and 72 h in each culture condition. To stain cells, medium was replaced with 400 μL of viability dye solution (25 μL of NucBlue Live reagent and 25 μL of propidium iodide in 1 mL of 26795


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Figure 4. (A) Electrical conductivity detection of culture medium. (B) 3D simulation results for temperature distribution at Vp = 3 V, σ = 1.7 S/m, f = 10 MHz. (C) Simulation results for average temperature at the bottom of the cell culture chamber and medium reservoir chamber incorporating different electrical conductivities and AC voltages. (D) Measured and simulated temperature results of cell culture and medium reservoir chambers at σ = 2.0 S/m and peak voltage of 0 V, 2.5 to 5 V. Dulbecco’s phosphate-buffered saline) for 30 min in the dark at room temperature. After incubation, cells were imaged using a Nikon fluorescence microscope. Experiments were conducted in three replicates. Total numbers of live and dead cells were quantified using NIH ImageJ software, and cell viability was determined as the ratio of live cells to the total cell number.43

m, respectively). The concentration of FBS changes the volume of medium, which in turn affects the concentration of salt ions. 4.2. Temperature Rise of the Circulatory Pumping Chip. The CPC is developed to provide a dynamic culture environment for cells. As the mechanism of this pump is based on electrothermal flow, the temperature rise (ΔT) is not neglected. Indeed, too high a temperature will damage cells; thus, a mechanism to control temperature within a reasonable range was of utmost priority to our research. As Figure 4b shows, cell culture regions are separated far from hot working electrodes to decrease potential negative effects of temperature on cells. Results from our simulation of temperature distribution at Vp = 3 V, σ = 1.7 S/m, and f = 10 MHz revealed the working electrode region is the hottest area, whereas the cell culture region does not exhibit an obvious temperature rise. To quantitatively analyze the effect of temperature, average temperature distribution at the bottom of the cell culture chamber and medium reservoir at different conductivities were calculated using finite software (COMSOL Multiphysics 4.4). As Figure 4c shows, only a 0.0001−0.0008 °C temperature rise occurs at the bottom of culture medium for electrical conductivities ranging from 1.5 to 2.0 S/m (inspired by actual experimental medium conditions) and peak voltages of 2.5 to 5 V. However, an obvious rise in temperature occurs at the bottom of medium reservoir, where ΔT varies from 0.68 °C (at Vp = 2.5 V, σ = 1.5 S/m) to 3.60 °C (at Vp = 5 V, σ = 2.0 S/m), increasing with a rise in electrical conductivity of medium. Thus, temperature fluctuations in the medium reservoir are far greater than for the area of cell culture. From our simulation results, we concluded that the highest temperature rise occurred in the medium with a conductivity of 2.0 S/m, which is also the actual conductivity value of

4. RESULTS AND DISSCUSSIONS 4.1. Electrical Conductivity Detection of Medium. Electrical conductivity is one of the key factors to ACET flow, which affects the driven velocity of culture medium; thus, the electrical conductivity of culture medium should be measured. Figure 4a shows the results of our analyses of culture medium and supernatant electrical conductivity after different exposure times. DMEM supplemented with 10% FBS and 1% P/S (DMEM, 10% FBS) was the primary medium used in culture experiments; its electrical conductivity before cell culture is 1.680 S/m. After HEK293T was cultured for 24, 48, and 72 h in 48-well plates at a density of 100 000 cells per 400 μL culture medium without changing, the electrical conductivities of supernatant samples were measured with a commercial conductivity sensor. Measured electrical conductivities were 1.728, 1.828, and 1.976 S/m after culture for 24, 48, and 72 h, respectively. The main reason a change in conductivity is observed is variation of salt ion concentration, which affects the electrical conductivity of medium. Cell metabolites and fluid evaporation may result in rising concentrations of salt ions; however, if medium is changed every 24 h, electrical conductivity is maintained at ∼1.7 S/m. To explore additional types of medium, the electrical conductivity of DMEM with 15% FBS and 1% P/S (DMEM, 15%FBS) and DMEM-F12 with 10% FBS and 1% P/S (DMEM-F12, 10% FBS) were measured (1.548 and 1.694 S/ 26796


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Figure 5. Fluid velocities at 300 μm height as a function of electrical conductivities and applied voltages at position “F” in Figure 1. (A) Superimposed frames at σ = 1.7, 2.0 S/m and Vp = 3, 4, or 5 V. (B) Experimental and simulated results for velocity at different applied voltages and electrical conductivities. (C) Verification of pumping performance at σ = 1.7 S/m and Vp = 2.5−5 V after culturing cells for 72 h.

experimental medium after working 72 h, as shown in Figure 4a. Figure 4d shows simulated and experimental temperature rises in the cell culture region and medium reservoir. The experimental result indicates nearly no ΔT occurs in the CPC culture area at voltages below 4 V, whereas voltages of 4.5 and 5 V result in a small increase in ΔT. Simulated boundary conditions of temperature on the top surface of PDMS and bottom surface of glass slide were set to a constant temperature of 37 °C; however, it is hard to dissipate heat to maintain surface temperature at 37 °C. Therefore, simulated ΔT in the culture chamber was lower than experimental results at voltages of 4.5 and 5 V. Furthermore, ΔT ranged from 0.18−2.56 °C in the reservoir chamber, which is ∼1 °C less than the simulated result. One reason for this variation is that the reservoir chamber is not a closed system, and heat loss through liquid evaporation was ignored in the simulation process. Furthermore, experimental data reflect the temperature around the sensor electrode placed on the bottom of reservoir chamber, whereas simulated results reflect the average temperature of the bottom surface of the reservoir chamber. An error in approximation may be another reason for an observed difference between simulated and measured results. Therefore, for cells demonstrating a high degree of sensitivity to elevated temperatures, it is safe to use voltages under 4 V with medium exhibiting electrical conductivity of no more than 2.0 S/m. 4.3. Pumping Velocity of the Circulatory Pumping Chip. The velocity of ACET flow in the CPC was evaluated by capturing the motion of 0.5 μm yellow fluorescent microspheres in the channel during video recordings. Individual video frames were extracted with Premiere Pro CS6 software (Adobe). Thirty adjacent frames were superimposed using ImageJ. Figure 5a shows superimposed frames from video of velocity experiments at electrical conductivities of 1.7 and 2.0

S/m, with applied peak voltages of 3, 4, and 5 V, respectively. The velocity of fluid flow increases as medium conductivity or applied voltage amplitude increases. The moving trajectories of microspheres are not straight at an applied voltage of 3 V in solutions with electrical conductivities of 1.7 and 2.0 S/m. A possible reason for this may be the Brownian motion of the 0.5 μm fluorescent microspheres.44 Although the moving trajectories of fluorescent microspheres are straighter at applied voltages of 4 and 5 V in solutions with electrical conductivities of 1.7 and 2.0 S/m, they still demonstrate some fluctuations across the microchannel. In addition, the concentration of microspheres is high; thus, some microspheres may stick together and appear to be bigger in superimposed photos (Figure 5a). Experimental results of fluid velocity with electrical conductivities of 1.5, 1.7, and 2.0 S/m at different applied voltages are shown in Figure 5b. Velocity increases with increasing voltage. With contribution from the 30 electrode pair-ACET pump, 400 μL of fluid was circularly driven at a velocity up to 26.17 μm/s at Vp = 5 V, σ = 2.0 S/m. Figure 5b also shows a comparison of experimental and simulated velocity results. Obviously, velocities in simulation results were approximately twice that of the results from experiments conducted with the same conditions. The main reason for the observed difference between experimental and simulated results is the lower temperature gradient that occurred during experiments. The temperature of the outside surface maintained a constant 37 °C in simulations, whereas it was difficult to dissipate heat to maintain surface temperatures at 37 °C in experimental conditions. The temperature gradient was lower as a result of nonideal heat dissipation; thus, ACET flow velocity was lower than simulated results. Additional reasons potentially contributing to deviation of experiments from simulated results include electrochemical 26797


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Figure 6. Results from cell culture. (A) Images of HEK293T and SW620 after culturing 24, 48, and 72 h in CPC with activated pumping. Cell proliferation of HEK293T (B) SW620 (C) in CPC with and without activated pumping, as well as in 48-well plates.

between 1.68 and 1.728 S/m; thus, fluid conductivity can be approximated at 1.7 S/m. At 300 μm above the bottom of the glass surface, the velocity is measured as 4.15 μm/s, which is close to the average speed of the whole channel section; hence, flow rate in the CPC was established to be 44.82 μL/h. Shear stress present in the flow chamber was calculated using the formula16

reactions and other electrothermal effects such as buoyancy force38,42,45 and strong electrothermal coupling of fluid properties (fluid conductivity and permittivity). Buoyancy force was ignored in our simulation, even though it exists in experiments. Additionally, loading effect may play a part in observed velocity differences, that is, the actual voltage drop on electrodes caused by numerous electrodes pairs and electrochemical reactions on electrodes will change with applied voltages. 4.4. HEK293T and SW620 culture in the Circulatory Pumping Chip. Compared with static methods, culture of cells or tissues in a dynamic circulatory chip is an outstanding method to mimic human physiological environments. Dynamic flow, like blood, removes cell waste immediately and provides nutrients; thus, metabolites and toxins can be diluted and washed away. Culture medium can be changed every several hours or several days, corresponding to human intake of foods and detoxification performed by the liver. As a result, the widely concerned human-on-a-chip system uses a large and dynamic circulatory system. To verify the practicality of our CPC, a normal cell line (HEK293T) and cancer cell line (SW620) were used as model cells in our experiments. Difficulties in experiments arose from seeding of cells into the culture chamber. If 100 000 cells were suspended in 400 μL of medium, seeding occurred everywhere within the CPC, both in the desired culture area and undesired pump area. Henceforth, 400 μL of medium was injected into the CPC first, before a 10 μL drop containing a high density of cells was gently seeded into the culture chamber; this method resulted in most cells being seeded within the culture chamber. Culture medium was changed every 24 h. According to Figure 4a, the electrical conductivity of culture medium ranges

τ=

6Qη bh2

(8)

where τ is wall shear stress, Q is flow rate, η is fluid viscosity, h is the flow chamber height, and b is the flow chamber width. As flow rate is low and the size of the culture is large, the small amount shear stress exhibited on cells could be neglected. The images in Figure 6a record the status, form, and concentration of HEK293T and SW620 cultured in the CPC after 24, 48, and 72 h of culturing with activated pumping. Compared with culture results from cells in 48-well plates and CPC without pumping, the status, form, and concentration are almost identical at every time point. Figure 6b,c shows the statistical results of cell proliferation of both cell types at each time point, respectively. Cell proliferation observed in the CPC with activated pumping is consistent with 48-well plates and CPC without pumping, confirming the practicality of our CPC. Most cell types have the capacity to adapt their environment. Dynamic culture with low shear stress provides a similar environment to static culture; thus, hardly any difference in cell morphology and proliferation exists between static and dynamic culture. In a dynamic human-on-a-chip system, however, several organ chips constitute a large dynamic circulatory system. It is of great significance to transport metabolites from one organ to another in an orderly fashion through a dynamic culture 26798


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Figure 7. Results of cell viability analysis. (A) Fluorescent images of HEK293T and SW620 after culturing for 24, 48, and 72 h in CPC with activated pumping. Live cells are blue, and dead cells are purple. (B) Cell viability of HEK293T in CPC with activated pumping, no pumping, and 48-well plates. (C) Cell viability of SW620 in CPC with activated pumping, no pumping, and 48-well plates.

5. CONCLUSIONS This paper introduces ACET as a novel circulatory pumping approach for cell culture applications with great potential for human-on-a-chip systems. The measured conductivity of medium in our experiments, which replaced fresh culture medium (DMEM with 10% FBS) every 24 h, was maintained at ∼1.7 S/m, ensuring the arising possibility of ACET. We successfully avoided the Joule heating effect and electrical damage to sample cells by properly positioning distinct pumping and culture regions. The flow rate in our CPC was ∼44.82 μL/h at a voltage of 3 V and frequency of 10 MHz, which was efficient for cell culture. Satisfactory performance of our device was demonstrated by successful culture of HEK293T and SW620 cells for 72 h. Moreover, our device was experimentally verified to be still effective after 72 h of use.

environment. Moreover, we measured pH values of cell supernatants from each device at 24, 48, and 72 h. Our results indicated sample pH values ranging from 7.22 to 7.31, which are within the normal range (7.2−7.6) of cell culture. The ability of cells to grow in the chip is fundamental to establishing the effectiveness of the device. We evaluated the viability of HEK293T and SW620 by quantifying live and dead cells cultured in the CPC with and without AC signal, as well as in 48-well plates. Figure 7a shows fluorescent staining of HEK293T and SW620 cultured in the CPC with AC signal for 24, 48, and 72 h. NucBlue Live reagent (blue) stains the nuclei of all cells, while propidium iodide (red) only stains the nuclei of cells with compromised plasma membrane integrity. Therefore, in the merged photo, nuclei with blue fluorescent color indicate live cells, while nuclei with purple fluorescent color (merged by red and blue color) indicate dead cells. The quantified viability of HEK293T and SW620 are shown in Figure 7b,c, respectively. Cell viabilities were observed to be higher than 95% at 24, 48, and 72 h for all devices, demonstrating effectiveness of the CPC. Moreover, cell viabilities of HEK293T and SW620 increased to their highest point at 72 h, indicating cells were at their best status at this time point. Pumping ability of the CPC was verified after 72 h of cell culture. To verify the device was still effective, Figure 5c presents the velocity result at σ = 1.7 S/m and Vp = 2.5−5 V, both before and after 72 h in culture. A calculated velocity decrease of 5−10% was deemed acceptable and within a reasonable range; thus, the pump effectively functioned throughout this period.

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AUTHOR INFORMATION

Corresponding Authors

*E-mail: [email protected]. (Y. R.) *E-mail: [email protected]. (L. L.) *E-mail: [email protected]. (H. J.) Notes

The authors declare no competing financial interest.

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ACKNOWLEDGMENTS Yukun Ren and Hongyuan Jiang gratefully acknowledge the National Natural Science Foundation of China (Nos. 51305106 and 11372093), the Fundamental Research Funds for the Central Universities (Grant Nos. HIT. NSRIF. 2014058 and HIT. IBRSEM. 201319), Self-Planned Task (No. 201510B) of State Key Laboratory of Robotics and System (HIT) and the 26799


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Programme of Introducing Talents of Discipline to Universities (Grant No. B07018). Lei Lei acknowledges the State Key Development program of basic research of China (No. 2012CBA01303).

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ABBREVIATIONS AC, alternating current ACEO, electroosmosis ACET, AC electrothermal CPC, circulatory pumping chip DC, direct current DMEM, Dulbecco’s modified eagle medium EHD, electrohydrodynamic FBS, fetal bovine serum ITO, indium−tin−oxide PDMS, polydimethylsiloxane PMMA, poly(methyl methacrylate) P/S, penicillin and streptomycin 2D, two-dimensional 3D, three-dimensional

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AC Electrothermal Circulatory Pumping Chip for Cell Culture - ACS

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