Automated Microfluidic Instrument for Label-Free and High

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An automated microfluidic instrument for label-free and high-throughput cell separation Xinjie Zhang, Zhixian Zhu, Nan Xiang, Feifei Long, and Zhonghua Ni Anal. Chem., Just Accepted Manuscript • DOI: 10.1021/acs.analchem.8b00539 • Publication Date (Web): 01 Mar 2018 Downloaded from http://pubs.acs.org on March 3, 2018

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Analytical Chemistry

An automated microfluidic instrument for label-free and high-throughput cell separation Xinjie Zhang,a Zhixian Zhu,a Nan Xiang,*a Feifei Long,b and Zhonghua Ni*a

a

School of Mechanical Engineering, and Jiangsu Key Laboratory for Design

and Manufacture of Micro-Nano Biomedical Instruments, Southeast University, Nanjing, 211189, China

b

Nanjing Foreign Languages School, Nanjing, 210096, China

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Analytical Chemistry 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60

ABSTRACT Microfluidic technologies for cell separation were reported frequently in recent years. However, compact microfluidic instrument enabling thoroughly automated cell separation is still rarely reported until today due to difficult hybrid between macro-sized fluidic control system and micro-sized microfluidic device. In this work, we propose a novel and automated microfluidic instrument to realize size-based separation of cancer cells in label-free and high-throughput manner. Briefly, the instrument is equipped with a fully integrated microfluidic device and a set of robust fluid-driven and control units, and the instrument functions of precise fluid infusion and high-throughput cell separation are guaranteed by a flow regulatory chip and two cell separation chips which are the key components of the microfluidic device. With optimized control programs, the instrument is successfully applied to automatically sort human breast adenocarcinoma cell line MCF-7 from 5mL diluted human blood with a high recovery ratio of ~ 85% within a rapid processing time of ~ 23 min. We envision that our microfluidic instrument will be potentially useful in many biomedical applications especially cell separation, enrichment, and concentration for the purpose of cell culture and analysis.


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INTRODUCTION Microfluidic technologies for cell separation in the biomedical field are booming in the recent 15 years.1 Benefitted from the microscale features approaching to a single cell, microfluidic cell sorting devices have shown many attractive advantages comparing to their forerunners (e.g., fluorescence-assisted cell sorting (FACS)2,3 and magnetic-assisted cell sorting (MACS)4,5), such as excellent temporal and spiral precision,6 high-throughput cell processing,7,8 low-cost material for disposable use,9,10 small footprint for compact integration,11,12 etc. With the advance in diagnostic and therapeutic medicine for point-of-care test, the need for cell separation has been greatly expanded nowadays, and different microfluidic devices were developed to perform blood fractionation,13,14 rare cell capture,15,16 bacteria and protein isolation,17,18 etc. Although many microfluidic cell sorting devices have been reported for leading researches and applications, there still exists a wide gap between the laboratory devices and the commercialization of microfluidic instruments, and the number of proposed microfluidic device is far higher than the number of instrument undergoing commercialization. The first microfluidic-based cell sorter was proposed in 1972 by Kamentsky and Melamed,19 after 45 years of development in microfluidics, the compact microfluidic instrument enabling automated cell separation is still rarely reported until today, and only a small amount of microfluidic devices are one the way to be commercialized. Technically, the impediment in commercialization is mainly due to the limitation of microfluidic device in mass production and the difficult hybrid between instrument and device from macro to micro systems.20 Therefore, it is vital for researchers to design a certain device suitable for mass production, while they also need to develop a set of macro fluidic control systems to support the device so as to achieve good cell separation efficiency. According to the working principle of reported cell sorting devices, current cell sorters can be divided into two categories, which are biochemical method and biophysical method.21 The biochemical instrument usually sorts cells through affinity capture in combination with special surface antigens of target cells, and the detection and identification of cells are realized through fluorescent labeling. Notably, Toner’s group developed a microfluidic platform, which is capable of capturing human tumor cells from ACS Paragon Plus Environment


whole blood using antibody-coated microposts inside of chip.16 As an optimization for improving chip production, they also proposed a tumor antigen-independent technology, which applies deterministic lateral displacement, inertial focusing and magnetophoresis techniques to sort cells.22,23 A key challenge of the biochemical method is that some biological cells are inherently heterogeneous, and they don’t express tumor-specific markers such as epithelial cell adhesion molecules (EpCAM) and cytokeratins (CKs), thus they couldn’t be detected and sorted efficiently. Instead, the biophysical method doesn’t rely on antigen expression but realizing cell separation based on physical characteristics, such as cell size,24 deformability,25 shape,26 electrical27 and acoustic properties,28 etc. Therefore, the biophysical method is totally label-free and can greatly benefit sample preparation and postanalysis processes.21 A representative instrument of biophysical method is the commercialized ClearCell® FX system developed by Clearbridge Biomedics Pte Ltd. By applying an inertial microfluidics-based technique, the system can automatically sorts tumor cells from human blood based on size difference between tumor cells and blood cells.29,30 In this work, we aim to develop a new instrument which employs a disposable microfluidic device and a robust fluidic control system, and the instrument can be applied for automated cell separation in label-free and high-throughput manner. At first, the structure and working principle of microfluidic device are introduced. Then, the instrument architecture and operational procedures under separation mode, back flow mode, and enrichment mode are demonstrated. To understand the flow rate for cell separation in the microfluidic device, we investigate the optimal flow rates of cell separation chip and the flow characteristics of flow regulatory chip systematically, and the system pressure and flow rate of the finished instrument are also studied. In the end, the instrument is applied to sort human breast adenocarcinoma cell line MCF-7 from human blood.

MATERIALS AND METHODS Design of microfluidic device. The fully integrated microfluidic device, illustrated in Figure 1(a), is composed of five components: (#1) a flow inlet chip for sample and sheath ACS Paragon Plus Environment

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inputs, (#2) a flow regulatory chip for precise flow rate autoregulation, (#3 and #4) two cell separation chips for high-throughput cell isolation, and (#5) a flow outlet chip for waste (i.e., isolated blood cells) and product (i.e., isolated cancer cells) outputs. Among the chips, four layers of adhesives are applied to bond the chips to be a compact device. To connect the device to microfluidic instrument, a custom-made fixture is used to load and clamp the device. The functional flow chart of device is illustrated in Figure S-1. Briefly, sample and sheath fluids are infused into the flow inlet chip (#1) under pressurized gas. Then, both the fluids are introduced into the flow regulatory chip (#2) and are passively regulated by the chip to be constant. Next, the constant fluids flow into two cell separation chips (#3 & #4), and blood cells and cancer cells in the sample are continuously isolated. In the end, the flow outlet chip (#5) outputs the product and waste.

Figure 1. Schematics of integrated microfluidic device. (a) Explosive view of device; (b) Working principle of flow regulatory chip. Images of A-A section show membrane deformation when fluidic pressures in the control channels are increased to push two elastic membranes towards the fluidic channel; (c) Working principle of cell separation chip. Large and small particles are separated in the outlets of spiral channels due to the inertial lift force and Dean drag force.

The device was designed to have two critical functions of flow rate autoregulation and



cell isolation, which are guaranteed by the flow regulatory chip and cell separation chip, respectfully. To regulate the flow rates of sample and sheath, nine passive microvalves were employed in the flow regulatory chip, as shown in Figure 1(b). Specifically, one single valve was applied for sample control, and the other eight valves (i.e., the integrated valve in Figure 1(b)) were applied for sheath control. The valve structure has been reported previously by our group.31,32 Briefly, the valve has three channels which are two control channels and a fluidic channel. The two control channels are respectfully located at the upper and lower sides of a fluidic channel. Two elastic membranes are used to separate and seal the three channels, which produces two dead zones in the terminals of the two control channels (see Figure S-2(a)). Since the three channels share the same inlet, when the inlet pressure (P+∆P) is increased to be higher than the threshold value, the flow resistance (R+∆R) of the valve is also increased to compensate the pressure variation due to the membrane deformation under the increased pressures from the control channels (see Figure 1(b)). Finally, the valve outputs a constant flow rate (Q=(P+∆P)/(R+∆R)). For the realization of cell label-free separation, an inertial microfluidics-based technique termed Dean Flow Fractionation (DFF) was applied to design the cell separation chip.29,33 The chip structure, shown in Figure 1(c), includes four spiral channels which are patterned in parallel, and each spiral channel has two inlets and two outlets. In the spiral channel, sample and sheath are introduced into the outer inlet and inner inlet, respectfully. Due to the influence of size-dependent inertial lift force (FL) and Dean drag force (FD), larger particles (ap/h > 0.07, ap is particle diameter, h is channel height) undergo dominant inertial lift force and focus near the inner wall region, while smaller particles (ap/h < 0.07) experience stronger Dean drag force and migrate along the Dean vortex streams which are two symmetrical counter-rotating fluid vortices in the channel cross-section.34 When the smaller particles move to the outer wall region near the outer outlet, the larger and smaller particles are separated effectively. In order to separate particles with high-throughput manner, four spiral channels were designed in one chip, and two stacked chips were used to double the throughput again. To match the flow rates between the flow regulatory chip and cell separation chip, we first designed the dimension of cell separation chip according to the DFF mechanism, and ACS Paragon Plus Environment

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then we determined the dimension of flow regulatory chip according to the design principle of passive valve.31 The width and height of spiral channel of cell separation chip were designed to be 500 µm and 160 µm, respectfully, and the gap between two spiral channels was designed to be 2 mm so as to increase the chip strength. The radii of spiral channel at inlet and outlet are 3.7 mm and 10.65 mm, respectfully, and the total channel length is 120 mm. In the valve of flow regulatory chip, shown in Figure S-2(a), the width (L) of control channel is 325 µm, and the width (W) and height (H) of fluidic channel are 150 µm and 105 µm, respectfully. The thickness of membrane is 20 µm. Fabrication of microfluidic device. Flow inlet chip, inertial separation chip and flow outlet chip were fabricated using UV laser cut and thermal lamination methods described in our previous work.9 Briefly, transparent polyvinyl chloride film (PVC) was applied as the channel layer (thickness of 160 µm film was used for inertial separation chip, and thickness of 200 µm film was used for flow inlet and outlet chips), and transparent polyethylene terephthalate films coated with thermal sensitive ethylene vinyl acetate copolymer (PET/EVA, total thickness of 100 µm) were applied as the lamination layer. To fabricate the chip, circuit pattern of the channel layer was firstly cut through using a UV laser machine (TH-UV200A, Tianhong Laser, China), and then the layer was bonded and sealed inside two lamination layers using a commercial grade laminator (LM8-330, Rayson, China). Flow autoregulatory chip was fabricated using the UV laser cut and oxygen plasma bonding methods. Firstly, silicon membranes were applied to fabricate the control channel layers (thickness of 200 µm) and fluidic channel layer (thickness of 105 µm) through laser cutting, and the protection film of silicon membrane was used as the cover and bottom layers.

Then,

elastic

membrane

was prepared

through

spinning

and

curing

polydimethylsiloxane (PDMS, Sylgard 184, Dow Corning, USA) on a sheet of CAPTON film. Finally, every layer was bonded and assembled through oxygen plasma treatment. The device assembly was performed through stacking all the above chips using double-sided adhesive tap layer by layer, and the finished device has a compact size of 51 x 53 mm2 with a thickness of 2.7 mm. Architecture of microfluidic instrument. A photograph of microfluidic instrument is ACS Paragon Plus Environment


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shown in Figure 2(a), and an additional schematic in Figure S-2(b) shows the side view of instrument. The instrument was designed as portable and compact specification, and the unit dimension is 300 × 280 × 270 mm with the weight of 10 kg including reagent. The instrument contains two centrifugal tubes of 50 ml volume to hold sample and product in the front side, and two reagent bottles of 500 ml volume are located in the left side to hold waste and sheath, respectfully. The microfluidic device is sandwiched by a fixture and is connected to the instrument on the top. To show operational message, a LCD displayer is used in the front shell accompanied with four buttons for setting parameters. Inside the instrument (see Figure 2(b)), a diaphragm pump (KVP04-1.1-12, Kamoer, China) producing pressurized gas up to 90 kPa for fluid-driven, four solenoid valves (VT307V-5G-01,

SMC,

Japan)

for

switching

gas

on/off,

a

gas

sensor

(PSAN-C01CV-Rc1/8, Autonics, Korea) for pressure detection, and six electromagnets (JYE-2620Z4.8, Jin CloudCN, China) for controlling fluid on/off are employed. In addition, a custom-made MCU (Microcontroller Unit) with ARM program is used to control the above electrical components to perform the preset actions. The transportation of sample and sheath from the instrument to the microfluidic device is realized via several elastic silicone tubes for medical use. The instrument was designed to be totally automated for cancer cell separation, and the procedure of instrument operation includes three working modes which are separation mode, back flow mode and enrichment mode. In the separation mode, shown in Figure 2(c), all the solenoid valves are switched on to connect port P(1) and A(2) at first, and all the electromagnets are also switched on to stop the fluids in the liquid circuit. Next, the diaphragm pump is turned on to draw atmosphere into the sample tube and sheath tank continuously via the gas circuit. When gas pressure is detected to be higher than a start-up value set by the pressure sensor, the electromagnets a, c, d and e are switched off to pass the fluids, and the sample and sheath flow into the microfluidic device under the gas pressure to start cell separation. Finally, the cancer cell suspension flows into the product tube, and the blood cell suspension flows into the waste tank. Figure 2(d) shows the fluid control in the liquid circuit. When the electromagnet is switched on, the electromagnetic bar is lowered quickly to press the silicon tube due to the magnetic actuation, and the tube ACS Paragon Plus Environment

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is compressed tightly to stop the fluid. When the electromagnet is switched off, the electromagnetic bar is lifted to the initial position again to enable the fluid to flow smoothly.

Figure 2. Photograph of compact microfluidic instrument and schematic of instrument operation. (a) Photograph of instrument appearance. Enlarged pictures show the (i) close and (ii) open status of microfluidic device assembled in a custom-made fixture; (b) Interior architecture of instrument; (c) Operational procedure of instrument under cell separation mode. Components A, B, C, and D represent solenoid valves, Components a, b, c, d, e, and f represent electromagnets. Dashed and solid lines represent gas and liquid circuits, respectfully. (d) Fluid control in the liquid circuit. Fluid stops or flows when the electromagnet bar lowers or lifts.

In order to sort cancer cells with lower blood cell pollution, 2nd round separation is necessary, which means the product obtained from 1st round separation should returns to



the sample tube to experience separation process again. However, the product is high in volume, and it is vital to concentrate the product to reduce its volume before the following separation. Therefore, we designed back flow mode and enrichment mode, respectfully, to withdraw the product from the product tube to the sample tube at first and then to perform 2nd round separation. The details of the two modes are shown in Figure S-3. Microbead specification. Four kinds of polystyrene microbeads with different diameters of 4.8 µm (Cat. No.: G0500, 1% solids), 7 µm (Cat. No.: 4207A, 0.3% solids), 15 µm (Cat. No.: 4215A, 0.3% solids) and 20 µm (Cat. No.: 4220A, 0.3% solids), respectfully, were used to mimic human blood cells and MCF-7 cells. To prepare mixed microbead suspension for separation experiments, 4.8 µm microbead with 0.005% volume fraction and 7 µm, 15 µm, and 20 µm microbeads with a higher volume fraction of 0.01% each were added together into deionized water to be acted as the sample fluid, and deionized water was also used as the sheath fluid. All microbeads were purchased from Thermo Scientific Inc., Fremont, CA, USA. Cell preparation. Blood sample was obtained from healthy donors with informed consent, and human breast adenocarcinoma cell line MCF-7 was purchased from KeyGEN BioTECH, Jiangsu, China. In the cell separation experiments, MCF-7 cells were stained with calcein AM (acetoxymethyl ester) (C3100MP, Invitrogen, USA) for clear identification. We also added a high concentration of MCF-7 cells (~ 105 cells/mL) into blood sample diluted by phosphate buffered saline (PBS) to facilitate analysis, and the total sample volume was 5 mL with the blood cell concentration of ~ 108 cells/mL. In addition, in order to prevent cell adsorption in the microchannels and tubes, 0.5% volume fraction bovine serum albumin (BSA) (Miltenyi Biotec, Germany) was added into the sample suspension. Experimental setup. In the microbead separation experiments, infusions of microbead suspension and deionized water were provided using two syringe pumps (LEGATO 270, KD Scientific, USA), and microbead motions near the channel outlets of cell separation chip were captured using an inverted fluorescent microscope (IX71, Olympus, Japan) which was equipped with a high-speed CCD camera (Retiga EXi, QImaging, USA). In the flow characterization experiments, the flow rate of single valve was measured using a set ACS Paragon Plus Environment

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of gas-driven flow setup. The setup uses compressed air and a pressure controller (OB1 Base MkIII, Elveflow, France) to produce the fluid driving force in the inlet of valve, and the flow rate in the valve outlet was measured by a flow sensor (MFS 5, Elveflow, France). In addition, the flow rate of integrated valve was measured using an analytical balance (AX523ZH, OHAUS, USA). In the cell separation experiments, an automated cell counter (Countess II FL, Life technologies, USA) was used to distinguish the fluorescent MCF-7 cells and the nonfluorescent blood cells, and the concentrations of MCF-7 cells and blood cells were also calculated.

RESULTS AND DISCUSSION Microbead separation characterization of cell separation chip. The operational flow rate for mixed microbead separation in a DFF chip had already been studied in our previous work, and it demonstrates that a total flow rate of 1.65 ml min-1 (sample and sheath flow rates are 0.2 ml min-1 and 1.45 ml min-1, respectfully) is the best choice for high-efficiency separation.9 To further confirm the optimal flow rate for high-throughput separation in this work, we investigated the microbead separation through stacking two cell separation chips. To hold the chips for experiments, a custom-made fixture was prepared and connected with two syringe pumps which were used to provide precise sample and sheath infusions. Since the two chips included eight spiral channels, the sample flow rate was fixed to be 1.25 ml min-1 to make the tight microbead band, and the sheath flow rate was set to be varied from 10 ml min-1 to 12.5 ml min-1. Microbead separations at outlet #1 under different sheath flow rates are shown in Figure S-4. It was found that the mixed microbeads are successfully separated under all flow rate ranges. Specifically, the large microbeads of 15 and 20 µm diameters always focus tightly near the inner wall, and the small microbeads of 4.8 and 7 µm diameters move to the outer wall with different lateral positions across the microchannel. To further analyze the separation performance, two dashed lines of red and yellow colors were marked at the borders of the small and large microbead bands, respectfully, and the microbead-free gaps between the two lines were calculated. It was noticed that when the sheath flow rate is increased from 10 ml min-1 to 11.5 ml min-1, the gap also increases from 139 µm to 250 µm. However, ACS Paragon Plus Environment


when the flow rate is increased again from 11.5 ml min-1 to 12.5 ml min-1, the gap decreases to be 175 µm at last. Therefore, the best separation effect was achieved under the sheath flow rate of 11.5 ml min-1. Next, we compared the microbead separations at four outlets of two cell separation chips under the sheath flow rate of 11.5 ml min-1, as illustrated in Figure 3(a-d). It was found that the microbeads achieve good separations at all outlets, and most of the small microbeads move to the outer wall forming a tight band, while all the large microbeads focus near the inner wall. We calculated the microbead-free gap between the small and large microbead bands, and it shows that even the minimum gap at the outlet #4 (see Figure 3(d)) is wider than 200 µm, which is a safe distance for avoiding the interactive pollution of the small and large microbeads. To characterize the separation efficiency, the fluids from the inner and outer outlets were collected and observed. It was noticed that the fluid from the inner outlet only includes the large microbeads of 15 and 20 µm (see Figure 3(e)), while the fluid from the outer outlet includes the microbeads of 4.8 and 7 µm (see Figure 3(f)), which demonstrates the 100% separation of large and small microbeads under the determined flow rates.

Figure 3. Microbead separations at four outlets of cell separation chips under the sheath flow rate of 11.5 ml min-1. (a-d) Stacked composite images illustrate microbead-free gaps between the large and small microbeads. The calculated gaps are 250 µm (outlet #1), 230 µm (outlet #2), 232 µm (outlet #3), and 208 µm (outlet #4); (e) Large microbeads (diameters of 15 and 20 µm) collected from inner outlets; (f) Small microbeads (diameters ACS Paragon Plus Environment

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of 4.8 and 7 µm) collected from outer outlets. All scale bars are 100 µm.

Flow characterization of flow regulatory chip. The flow rate for microbead separation was investigated systematically in the last section, and the optimal sheath and sample flow rates were determined to be 11.5 ml min-1 and 1.25 ml min-1, respectfully. Since sample contributes minimum on the total Dean flow for microbead separation in the DFF chip (ratio of sample to sheath flow rate is ~ 0.11), precise sheath infusion is necessary to maintain a nearly constant Dean flow. For precise fluid infusion, syringe pump is applied in most microfluidic systems. However, limited capacity of syringe cartridge is not suitable for the pump to be continuously used in high-throughput applications. To solve this problem, we designed a flow regulatory chip which was integrated with nine passive valves to realize the high-throughput and constant fluid autoregulation in a gas-driven flow system. The flow characterization of chip was studied through measuring the flow rates of single valve and integrated valve under varied test pressures. Since the valve was designed with low threshold pressure, the test pressures were set to be varied from 10 kPa to 70 kPa in this work. The obtained pressure and flow rate curves are shown in Figure 4. It was found that when the pressure is increased from 10 kPa to 70 kPa, the flow rate of single valve also increases at first and then saturates to achieve a constant value of 1.24 ± 0.05 ml min-1 when the pressure is higher than a low threshold of 20 kPa. It was also found that the flow rate of integrated valve increases linearly at the beginning when the pressure is increased from 10 kPa to 30 kPa, and constant flow rate of 11.58 ± 0.3 ml min-1 is also achieved when the pressure is higher than 30 kPa. Since the integrated valve was composed of eight valves which were used to control the sheath in eight spiral channels individually, every valve should have the nearly equal flow rate so as to ensure the high-efficiency separation. To test the flow rate of every valve, we blocked the outlets of seven valves and measured the flow rate of the open outlet of the rest valve. The flow performance of the rest valve was found to be similar to the single valve, and the measured flow rate is equal to be one eighth the total flow rate of integrated valve. From the above experimental results, we conclude that both the single valve and integrated ACS Paragon Plus Environment


valve output the constant flow rates under low threshold pressures, and the valve flow rates are nearly equal to be the optimal flow rates for cell separation tested by syringe pumps. In addition, every valve shows the similar flow characteristic, which ensures the precise fluid infusion in every spiral channel when the flow regulatory chip works. Furthermore, since the threshold pressure of integrated valve is ~ 30kPa, which is a little higher than the threshold of single valve (~ 20 kPa), the working pressure of flow regulatory chip should be set higher than 30 kPa so as to maintain the constant flow rate outputs of both valves.

Figure 4. Measured flow rates of single valve and integrated valve under test pressures varied from 10 to 70 kPa. The black line, green line and red line represent the total flow rate of integrated valve, the flow rate of one valve, and the flow rate of single valve, respectfully.

System pressure and flow characterization of microfluidic instrument. Our microfluidic instrument was designed for automated cell separation, and its function is guaranteed through the programmed actions of every electrical component inside of instrument. The accurate program control ensures that each component performs within the preset parameters including system pressure, trigger time, sample volume, etc. An implied but important parameter here is the system flow rate, especially the fluids inside of microfluidic device in the duration of cell separation, which directly determines the recovery of cancer cells. Since the flow rate of microfluidic device is directly dependent


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on the flow regulatory chip which is totally passive in flow regulation, the external components for flow control and detection are not required. However, the system flow rate should be studied carefully so as to ensure the precise fluid infusion in the microfluidic device when the instrument works. As illustrated in the operational procedure of instrument, the instrument was designed with three working modes. Besides the back flow mode, both the system flow rates under the separation mode and enrichment mode should be studied. As the enrichment mode only refers sheath infusion, we just need to investigate the system flow rate under the separation mode. However, the system flow rate has a complicated relationship with the preset parameters. For example, the system pressure has a direct influence on the flow rate because of flow autoregulation characteristic of microfluidic device. Due to threshold limitation of microfluidic device, the system pressure should be set higher than the working pressure of device so as to ensure constant fluid control. The other parameters influenced by the system flow rate are the sample volume and separation time. Since separation flow rate is totally passive and constant, the processing volume of sample per unit time is also constant. Therefore, for fully separation, the preset sample volume and separation time should be calculated in advance. We then set the start-up pressure of system in the separation mode to be 35 kPa, and the rotational speed of diaphragm pump was set to be 30 rpm so as to produce working pressure higher than the start-up value. The sample volume was set to be 5 ml, and the separation time was determined to be 250 s according to the sample flow rate of 1.25 ml min-1. The measured system pressure and flow rate of instrument are shown in Figure 5. It was found that the system pressure increases gradually to achieve a constant value of ~ 56.3 kPa at first, and then it decreases directly to be zero in the end. In comparison to the pressure curve, the flow rate curves of sheath and sample show the significantly different performances. As shown in the figure, the two curves can be divided into four sections, which accordingly indicate four stages of instrument operation under the separation mode. To analyze the flow performances, we define the four stages to be preparation stage (#1), start-up stage (#2), separation stage (#3), and terminal stage (#4), respectfully. In the preparation stage, the diaphragm pump was switched on to produce pressurized gas into the sheath tank and sample tube. Since the inlets of microfluidic device were blocked by ACS Paragon Plus Environment


the electromagnets a, b and c, no sheath and sample flowed in the liquid circuit. When the gas pressure was increased to achieve the start-up value (i.e., 35 kPa), the instrument entered into the start-up stage. In this stage, all the device inlets were opened, and the sheath and sample flowed into the device in order. Due to the existence of air bubbles inside of tubes and device, both the flow rates of sheath and sample fluctuated dramatically. After all the bubbles were removed, the flow rates achieved to be constant soon, and the instrument entered into the separation stage. The recorded flow rates of sheath and sample were 11.7 ± 0.34 ml min-1 and 1.26 ± 0.03 ml min-1, respectfully, which are nearly same as the measured flow rates of flow regulation chip. We also found the constant sample was maintained to be ~ 240s, and the constant sheath was sustained for additional 10s than the sample so as to ensure complete exhaustion of sample in the tube. In the terminal stage, the diaphragm pump was switched off and no fluids flowed in the tube. The above experiments well validate the stable flow autoregulation characteristic of our microfluidic device under the automated separation mode, and the system pressure and flow rate are shown successfully applicable for precise fluid transportation which is strictly required by inertial separation.

Figure 5. System pressure, sheath and sample flow rates of instrument under separation mode. The two flow rate curves indicate four different stages of instrument operation which are preparation stage (#1), start-up stage (#2), separation stage (#3), and terminal stage (#4).

Automated cell separation and enrichment. To validate the applicable capability of


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finished instrument for automated cell separation and enrichment, the instrument was applied to sort MCF-7 cells from human blood. The instrument was embedded two control programs to compare the cell separation efficiency, one was used for one-time cell separation, and another was for two-time cell separation, and each program was performed three independent experiments. To study the separation efficiency, an easy and intuitive method is to observe the color of product. It was found that the product obtained from 1st round separation still shows a little pink in appearance (see Figure 6(a)), which demonstrates the possible pollution of blood cells. When the blood sample experienced 2nd round separation, it was no surprise to find that the product turned to be nearly transparent with the further reduction of blood cells. In comparison to the product, the waste involved most blood cells, which shows red in color (see Figure 6(b)). We next observed the cell distributions of the two products using a fluorescent microscope, and comparisons of bright and fluorescent images are shown in Figure 6 (c) and Figure S-5. It was noticed that the MCF-7 cells were successfully captured after 1st round separation, and most of the blood cells were removed (see Figure S-5). However, the quantity of the blood cells was still much higher than the quantity of MCF-7 cells. In comparison to 1st round separation, the quantity of blood cells was further reduced after 2nd round separation, and the MCF-7 cells were significantly enriched (see Figure 6(c)). To quantitatively study the separation efficiency, the quantities of MCF-7 cells and blood cells were calculated using a cell counter. The quantitative analysis of the ratio of blood cells to MCF-7 cells in Figure 6 (d) shows that the blood cells were removed significantly from the sample (ratio of ~ 1000) to the 2nd round product (ratio of ~ 1), which leads to enrichment ratio of ~ 1000 times and purity of ~ 50% of MCF-7 cells (see Figure 6(e)). We also calculated the removing ratio of blood cells and the recovery of MCF-7 cells, and the comparison results between the 1st round and 2nd round separations are illustrated in Figure 6(f) and 6(g). It was found that the removing ratio of blood cells of 1st round product has already higher than 95%, and the 2nd round product shows much higher blood removing ratio (~ 99.9%) than the 1st round (see Figure 6(f)). Thus, lower pollution of blood cells was achieved after 2nd round separation. It was also noticed that the instrument captured the MCF-7 cells with a high recovery ratio of ~ 90% in the 1st round ACS Paragon Plus Environment


separation, and recovery of ~ 85% was still maintained after 2nd round separation (see Figure 6(g)), which indicates a few MCF-7 cells were lost in the automated operational procedure. Therefore, two-time separation is acceptable and necessary for achieving better cell separation efficiency. In addition, due to the high-throughput characteristic of microfluidic device, the total processing time of two-time separation is only ~ 23 min. Our microfluidic instrument was proved to automatically sort MCF-7 cells from human blood with high-throughput and good efficiency. However, it still needs to continuously improve the technical parameters of instrument in the future. For example, the existence of air bubbles inside of tubes and microfluidic device is a very serious problem for preventing the instrument from achieving higher cell recovery. As blood cells are overwhelming in quantity than cancer cells, product may be severely polluted even a small quantity of blood cells accompanied with air bubbles flowing into the product tube. Therefore, complete exhaustion of air bubbles is necessary for not only our instrument but also for other microfluidic platforms with automated operations.

Figure 6. Automated cell separation using microfluidic instrument. (a) Blood sample and collected products; (b) Waste and sheath; (c) Bright and fluorescent images illustrating captured MCF-7 cells after 2nd round separation; (d) Quantitative analysis of ratio of blood cells to MCF-7 cells; (e) Quantitative analysis of MCF-7 cell purity and enrichment ratio. Here, the enrichment ratio means the enhancement of MCF-7 cell to blood cell ratio at the input sample to the output product; (f) Quantitative analysis of blood cell removing


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ratio; (g) Quantitative analysis of MCF-7 cell recovery ratio. All scale bars are 50 µm. All error bars are standard deviations from three independent experiments.

CONCLUSIONS In this work, we propose a novel microfluidic instrument to sort biological cells in label-free and high-throughput manner. The instrument is equipped with a disposable and fully integrated microfluidic device which enables the constant flow autoregulation and high-throughput cell separation, and a set of electrical fluid-driven and control units are employed to perform the totally automated operations. The flow rates of sample and sheath for high-efficiency separation in the microfluidic device are investigated systematically, and precise flow control under automated separation mode is realized. The instrument is then applied to sort MCF-7 cells from human blood, and good separation efficiency with high cancer cell recovery is achieved after two round separations. The entire operational process from sample to product is finished within 23 min without any human intervention. The successful development of our microfluidic instrument is a good supplement for many biomedical applications which rapid and high-throughput cell processing are required, especially for the application purpose of cell separation, enrichment, and concentration.

ACKNOWLEDEMENTS This research work is supported by the National Natural Science Foundation of China (51505082, 81727801 and 51775111), the Natural Science Foundation of Jiangsu Province (BK20150606), and the Fundamental Research Funds for the Central Universities (2242017K41031).

SUPPORTING INFORMATION The supporting information is available free of charge on the ACS publication website: functional flow chart of microfluidic device; schematics of microvalve and instrument; schematic operational procedures of instrument; microbead separations under different sheath flow rates; captured MCF-7 cell images. ACS Paragon Plus Environment


Corresponding Authors: *Nan Xiang, [email protected]. Tel.: +86-025-52090518. Fax: +86-025-52090501. ORCID Nan Xiang: 0000-0001-9803-4783 *Zhonghua Ni, [email protected]

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For TOC only


Automated Microfluidic Instrument for Label-Free and High

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