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Microfluidics and Nanofluidics (2018) 22:8 https://doi.org/10.1007/s10404-017-2026-0

RESEARCH PAPER

A portable, hand‑powered microfluidic device for sorting of biological particles

Sheng Yan1  · Say Hwa Tan2 · Yuxing Li1 · Shiyang Tang1 · Adrian J. T. Teo2 · Jun Zhang3 · Qianbin Zhao1 · Dan Yuan1 · Ronald Sluyter4,5 · N. T. Nguyen2 · Weihua Li1

Received: 15 October 2017 / Accepted: 5 December 2017 © Springer-Verlag GmbH Germany, part of Springer Nature 2017

AbstractManually hand-powered portable microfluidic devices are cheap alternatives for point-of-care diagnostics. Currently, on-field tests are limited by the use of bulky syringe pumps, pressure controller and equipment. In this work, we present a manually operated microfluidic device incorporated with a groove-based channel. We show that the device is capable to effectively sort particles/cells by manual hand powering. First, the grooved-based channel with differently sized polystyrene particles was characterized using syringe pumps to study their distributions under various flow rate conditions. Afterward, the particle mixtures were sorted manually using hand power to verify the capability of this device. Finally, the manually operated device was used to sort platelets from peripheral blood mononuclear cells (PBMCs). The platelets were collected with a purity of ~ 100%. The purity of PBMCs was enhanced from 0.8 to 10.4% after multiple processes which results in an enrichment ratio of 13.8. During the process of manual hand pumping, the flow fluctuation caused by unstable injection will not influence the sorting performance. Due to its simplicity, this manually operated microfluidic chip is suitable for outfield settings.

1 Introduction

In many rural areas within developing countries, access to quality health care and medical diagnostics tools are impeded by a host of various factors including cost, trans-portation, knowledge, age and language (Chin et al. 2011).

This often results in high death rates as a result of late diag-nosis and delayed treatment. The problem is further com-pounded by the high cost of such clinical tests. On the other hand, in developed countries, the high turn-around time needed for various laboratory tests calls for a portable, low-cost and reliable diagnostic tool. The advent of microfluidics provides solutions to solve and address both situations. For example, a number of microfluidic diagnostic devices have been developed till date. This includes detection of the Ebola and Zika viruses and biomarkers in blood, urine and salvia (Chan et al. 2017; Ryu et al. 2017) for various diseases and disorders.

Microfluidic devices are capable of many key functions. These include sample pre-treatment, fluid manipulation, biosensing, separation and monitoring and signal detec-tion (Chan et al. 2017). Among them, microfluidic separa-tion and sorting of biological targets are an indispensable procedure for biological analysis and clinical diagnostics (Sajeesh and Sen 2013). In microfluidics, microparticle separation can be divided into active and passive methods based on their manipulating forces (Yan et al. 2017). Active methods employ electric (Yan et al. 2014), magnetic (Kang et al. 2012), acoustic (Lin et al. 2012) and optic (MacDon-ald et al. 2003) fields to precisely control the particle posi-tions, whereas passive systems manipulate particles using

Electronic supplementary material The online version of this article (https://doi.org/10.1007/s10404-017-2026-0) contains supplementary material, which is available to authorized users.

* Sheng Yan shengy@uow.edu.au

* Weihua Li weihuali@uow.edu.au

1 School of Mechanical, Materials and Mechatronic Engineering, University of Wollongong, Wollongong, NSW 2522, Australia

2 Queensland Micro- and Nanotechnology Centre, Griffith University, Brisbane, QLD 4111, Australia

3 School of Mechanical Engineering, Nanjing University of Science and Technology, Nanjing 210094, China

4 School of Biological Sciences, University of Wollongong, Wollongong, NSW 2522, Australia

5 Illawarra Health and Medical Research Institute, Wollongong, NSW 2522, Australia

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hydrodynamic force (Zhang et al. 2016). Although active approaches can separate particles with a high selectivity, they require complicated fabrication process and the use of bulky and costly equipment (i.e., function generators, power amplifiers, lasers, optical fibers and magnets). Alternatively, passive techniques do not entail such consideration. Typi-cally, hydrodynamic forces coupled with a suitable channel topography can be applied for effective separations. Unfor-tunately, these approaches currently only work effectively under specific flow conditions. For example, despite deter-ministic lateral displacement being a versatile, robust tech-nique for particle separation, this technology only works at very low flow rates (~ 1 μl min−1), while higher flow speeds will modulate the critical diameter of microfluidic separation and deteriorate the separation efficiency (Huang et al. 2004). Elasto-inertial focusing can separate particles at low flow rates (~ 5 μl min−1) (Yuan et al. 2017). With the increase in flow rate, the inertial effect becomes more dominant and the particles will be unfocused. Inertial microfluidics has attracted great interest due to its high-throughput and simple operation (Xiang et al. 2015). The mechanism of inertial focusing has been investigated experimentally (Xiang et al. 2015; Liu et al. 2015) and numerically (Jiang et al. 2016; Liu et al. 2016). Although inertial device can work at a border range of flow rates (0.4–1.2 ml min−1) (Zhao et al. 2017), the channel cannot work effectively at relatively low flow rates (< 0.4 ml min−1) as the inertial effect is negligible at low Reynolds number (Re). In these methods, expensive and bulky syringe pumps or pressure controllers are still needed to provide a precise and regulated flow for the operation of passive systems.

In order to eliminate the peripheral equipment and to achieve real lab-on-a-chip systems for micrototal analysis applications out of laboratories, a hand-powered microflu-idic device bridges the technological gap. Capillary action is a power-free method that uses the cohesion between the liquid and the surface of the channel to draw reagent (Boyd-Moss et al. 2016). Dimov et al. (2011) developed a cheap, disposable poly(dimethylsiloxane) (PDMS) capillary expansion channel for plasma separation from whole blood pinpricks by the combination of capillary action and hydro-dynamic filtration. Capillary action, however, stops when the pressure between inlet and outlet reservoirs is balanced, limiting their proficiency in large-volume, high-throughput and continuous-flow processing. Human mechanical power can also be harnessed to open up new capabilities, including pressing elastic membrane by a finger (Glynn et al. 2014), pumping fluids by pulling a syringe (Garstecki et al. 2006) and pipetting of blood samples through porous membrane (Liu et al. 2016). Despite simplicity of operation, unstable pumping might cause different results by different users. For example, the vacuum generated by pulling a syringe among different users might differ, which in turn might influence

the liquid mixing performance (Garstecki et al. 2006) or droplet size (Abate and Weitz 2011).

In this work, we propose a high-throughput, continuous-flow, hand-powered microfluidic device which can be used for steadily sorting biological particles regardless of flow rate. The proposed device extends the concept of inertial microfluidics where the separation of particles and biologi-cal targets can be realized using a hand-powered, disposable and low-cost syringe. Specifically, the distributions of par-ticles in the groove-based channel are only slightly different at both low and high flow rates (Fig. 1). This is an impor-tant function to realize manual injection using hand power-ing. The function also allows the microfluidic device to be portable and eliminates the needs of peripheral equipment. However, several key issues pertaining to manual injection should be highlighted. During the different stages of manual injection, there are several kinds of flow fluctuations. These include fluid acceleration at the start of pumping, unstable flow rate caused by hand tremble during pumping and fluid deceleration at the completion of the injection. The proposed groove-based channel eliminates this inconsistency as it bridges hydrophoretic ordering and inertial focusing without any discontinuity. At low flow rates, hydrophoretic ordering predominates the flow, while inertial effects predominate at high flow rates. In order to demonstrate the capability in biological applications, the device is further tested for peripheral blood mononuclear cells (PBMCs) enrichment and purification of platelets. Typically, diagnosis of plate-let disorders requires platelet preparation for downstream immunofluorescence (Greinacher et al. 2017). The device can also be used as a low-cost solution for the preparation of platelets.

2 Theory

2.1 Secondary flow

In the groove-based channel, a secondary flow is generated when the fluid flow invades the groove structure due to the pressure differentiation in the transverse direction (Fig. 2). To generate the strongest secondary flow, the ratio of the height of the groove arrays and main channel is set at ~ 1 (Choi et al. 2009). Drag forces, induced by secondary flows (vortices), will exert on particles suspended in the channel and affect their distribution. Typically, the drag force at low Re can be expressed using the Stokes law (Gerlach 1998):

where vf is the velocity of the fluid, vp is the velocity of the particle, d is particle diameter, and μ is the fluid dynamic viscosity.

(1)FD = 6�d�(

vf − vp)

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2.2 Inertial lift force

When particles travel in a straight channel at a moder-ate Re, the inertia of fluids will cause lateral migration of immersed particles (Segre 1961; Segre and Silberberg 1962). The reason of the particle migration is due to the counteraction of two primary inertial effects: the shear-gradient lift force FLS and the wall lift force FLW (Zhang et al. 2016).

The net inertial lift force FL exerting on the rigid micro-particle can be expressed as (Asmolov 1999; Di Carlo 2009):

where ρ and U are fluid density and flow velocity, respec-tively. Dh is the hydrodynamic diameter. The lift coefficient fL is a function of the lateral position of particles and the Re (Re = ρUDh/μ).

(2)FL = fL�U2d4∕D2

h

2.3 Particle focusing in the groove‑based channel

Particles with diameters larger than half the channel height will have extensive particle–groove interaction, which prevents them entering the grooves by steric hindrance, a process termed hydrophoresis (Choi et al. 2008). The anisotropic groove generates the helical flow, consisting of focusing flow, upward flow, deviation flow and downward flow (Fig. 2b). The particles suspended in the microchan-nel are driven to the left sidewall of the channel, follow-ing the focusing flow. The particles then escape from the rotational flow due to the particle–groove interaction and remain at the left sidewall of the channel. At a high Re, the velocity vectors are deformed by coupling with Dean vor-tices induced by centrifugal forces (Fig. 2c). Meanwhile, inertial effects start to dominate the particles’ behaviors (Zhao et al. 2017). Two inertial forces, wall lift force and

Random distribu�on of par�cles

yz

x

Cross-sec�on A-A’

Low Re

High Re

A-A’

3mm

Re=7.3Re=73

(a)

(c)

(b)

Fig. 1 A manually operated hand-powered microfluidic device for particle/cell sorting. a Schematic showing the structure of the groove-based device. b The particle distributions in the cross section and flu-orescent intensities of 13-μm particles migrating in the groove-based

channel. c Image of the experimental setup. The microfluidic chip is connected to a syringe via the Teflon tubing. The insect shows the groove patterns

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shear-gradient lift force, exert on the particles. Firstly, the wall lift force and shear-gradient lift force are balanced at z direction (Fig. 2c). The drag force pushes a portion of particles to another equilibrium position. At the new equilibrium position, the particles are balanced by iner-tial wall lift force, shear-gradient lift force and drag force. Therefore, two equilibrium positions are formed at high Re. However, particles of smaller size and mass follow the secondary flow and remain unfocused in the channel, as the steric hindrance and inertial effect are negligible.

3 Materials and methods

3.1 Methodology

The particle distributions in the groove-based channel are largely identical but with minor differences at low and high Re. The ability to keep consistency in performance regard-less of flow rate is the key to demonstrate this concept—the plunger of the syringe is pushed manually by hand. In the duration of the hand pumping, there are several aspects that

may cause unstable flow rate. Firstly, at the beginning of hand-enabled pumping (acceleration stage), the fluid veloc-ity is gradually increasing from zero to a high level instead of jumping to a given flow rate. When the flow rate reaches a high level, the hand tremble which is caused by muscle fatigue may lead to unstable liquid perfusion. After comple-tion of injection (deceleration stage), the flow rate decreases from a high level to zero. Due to this wide operational range, these variations in the flow rate will not influence particle movements during a hand-powered injection.

3.2 Design and fabrication

Figure 1a shows a schematic of the 200 μm wide by 1-cm-long channel where each channel has 60 grooves with an angle of 10°. The gap and width of the slanted groove are both 40 μm. The height of the channel and grooves is 23 μm. The output of the inlet channel was connected to an 800-μm-wide expansion region for better observation and to facilitate data analysis. These microfluidic devices were fabricated by two-step photolithography and soft lithography (Yan et al. 2014). Briefly, the first layer of photolithography

1

2

3

4

1 2 3 4yx

z

1

2

3

4

Re=2.9 Re=73.7

Flow

Pa Pa

Upward flow

Focusing flow

Devia�on flow

Downward flow

Drag force Wall li� force Shear gradient li� forcexz

(a)

(c)(b)

Fig. 2 a Three-dimensional model for numerical simulation in the same geometric dimensions with the experimental channel. The flow direction is along the y-axis. Four simulation results are depicted by a combination of the velocity arrow plot and pressure contour plot at

Re = 2.9 (b) and Re = 73.7 (c). The arrows in the insets are veloc-ity vectors. The background color of the insets denotes pressure field (color figure online)

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was defined as the main channel; the second one with the pattern of grooves was aligned to lie on the top of structures in the first layer. After spin-coating, baking, UV exposure and developing, the SU-8 mold was fabricated for PDMS casting. A mixture of PDMS with curing agent (Dow Corn-ing, USA) at a ratio of 10:1 was poured over the silicon master, degassed and baked at 65 °C for 2 h. The devices were peeled from the silicon master, and inlet and outlet holes were punched with a 2-mm biopsy punch. The PDMS was sealed with glass slides after exposure to oxygen plasma (Harrick Plasma, USA) for 3 min.

3.3 Sample preparation

Fluorescent particles of 2, 8, 10 and 13  μm diameter (Thermo Fisher, USA) were used as surrogates for platelets and PBMCs. They were suspended in deionized (DI) water containing 0.1% Tween 20 (Sigma-Aldrich, USA) to impede the beads from sedimentation and aggregation.

Human blood from a healthy male volunteer was col-lected into 10-ml lithium heparin tube. The blood sample was transferred to a 50-ml sterilized tube and was mixed 1:1 with Dulbecco’s phosphate-buffered saline (DPBS, Thermo Fisher, USA). The mixture was underlaid with 10 ml Ficoll-Paque (GE Healthcare Life Science, Sweden) using a glass pipette. The sample was centrifuged at 1500 g for 30 min at room temperature (with low deceleration). The plasma and buffy coat containing platelets and PBMCs were then transferred to a new 50-ml tube.

3.4 Device characterization

Prior to experiments, the chips were sterilized through expo-sure to UV light for 20 min and then rinsed with 1% bovine serum albumin (BSA) (Sigma-Aldrich, USA) in phosphate-buffered saline (PBS) (Sigma-Aldrich, USA) to avoid non-specific adsorption. The microfluidic chip was placed onto an inverted microscope (Olympus, Japan). The images were captured by a CCD camera (Q-imaging, Australia) and then post-processed and analyzed with Q-Capture Pro 7 (Q-imag-ing, Australia) software. Particle and cell separation were measured using an Accuri C6 Plus flow cytometer (BD Bio-sciences, USA).

4 Results and discussion

4.1 Particle movements in a groove‑based channel

Before the implementation of the proposed manual opera-tion, a groove-based channel was characterized using a syringe pump over a range of flow rates (20–500 μl min−1). In this manner, the behavior of polystyrene microbeads can

be systematically investigated in terms of their distribution and position at various flow rates. Fluorescent particles with a diameter of 2 μm were used to mimic the platelets (1–3 μm) (Kersaudy-Kerhoas and Sollier 2013). Since PBMCs had a large distribution in size (8–12 μm), beads in a diameter range of 8–13 μm were used to emulate the PBMCs behavior in the groove-based channel.

The trajectories of fluorescent particles were captured with an exposure time of 50 ms. As depicted in Fig. 3, 2-μm particles remain unfocused regardless of flow rates. At a low flow rate, the small particles that did not meet the hydrophoretic criterion migrated along the rotational flows induced by the grooves. At a high flow rate, inertial effects did not influence particle behavior, as the parti-cle Reynolds number was less than 1 over the flow rate range (Fig. S1). The particle Reynolds number is defined as Rep = Re × d2/Dh

2. When the Rep is over 1, the iner-tial effect will dominate particle behavior (Zhang et al. 2016). For large particles (8, 10 and 13 μm), they satisfied the hydrophoretic ordering and laterally migrated toward the left sidewall. With an increase in flow rate, inertial effect started to dominate particle behavior (Fig. S1) and two streaks of particle focusing were observed (Fig. 3). Although the inertial effect can modulate the particle tra-jectories, the equilibrium positions of particles are over-lapped at low Re and high Re, occupying the half channel width near the left sidewall.

This finding provides a unique opportunity to operate the microfluidic chip using hand power. At the start and end of hand-powered injection, the flow rate is at the low level where large particles are dominated by hydrophoretic ordering. When the flow rate reaches a high level, iner-tial effect plays a dominant role in particle focusing. The small particles distribute evenly at the outlet during the whole process. Based on this, the purified small particles can be collected from Outlet 2 and large particles can also be enriched through this channel regardless of flow rate. Hence, the unstable flow rates induced by manual pumping will not affect the sorting performance. The device, however, has limitation in the maximum operating pressure induced at high flow rate. The slanted grooves in a channel were omitted to simplify the calculation, and the hydrodynamic resistance of the whole channel for water is approximately 6.4 × 10−5 Pa s μm−3 and for plasma is 9.6 × 10−5 Pa s μm−3. The pressure within the channel is ~ 530 kPa (in water) and ~ 795 kPa (in plasma) at the flow rate of 500 μl min−1. Such high pressure may result in the delamination of PDMS or leakage at the connection between the Teflon tubing and PDMS channel. During the experiments, we observed that the Teflon tubing detached from the fluidic port when the flow rate was over 500 μl min−1. Due to the higher viscos-ity of plasma (μ ~ 1.5 × 10−3 Pa s), the flow rate should be further reduced to avoid the aforementioned problems.

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Overall, the working flow rate limit of the device is less than 500 μl min−1.

4.2 Particle sorting in groove‑based channel

After characterization, we verified the capability of the groove-based channel to sort particles from a binary mix-ture by manual operation. Three sets of particle-laden sus-pensions were prepared (2- and 8-μm; 2- and 10-μm; and 2- and 13-μm particle combinations). A 5-ml syringe was connected to the microfluidic device via Teflon tubing. The particle suspension was injected into the channel by pushing the plunger of the syringe. The yielded volume throughputs ranged from 0 to 400 ± 100 μl min−1. After passing through the channel, 2-μm particles were sorted and collected from Outlet 2 and large particles were also enriched from Outlet 1 (Fig. S2). Figure 4a shows the differently sized particle trajectories at the outlet. As expected, all the large particles entered Outlet 1, while small particles were evenly distrib-uted. To quantify the separation performance, particle purity [collected target particle number/collected total number (Zhou et al. 2013)] and recovery rate [collected target parti-cle number/input target particle number (Loutherback et al. 2012)] were measured.

After manually operating the microfluidic sorting, the purity of the 8-μm particles increased from 57.1 ± 2.3 to 87.2 ± 1.7% from Outlet 1. The recovery rate of the 8-μm particles reached 99.8 ± 0.2%. The 2-μm particles were collected from Outlet 2 with a purity of 99.4 ± 0.2% and a recovery rate of 52.3 ± 2.4% (Figs. 4b, c, S3). For 2- and 10-μm particle sorting, the purity of 10-μm particles increased from 46.9 ± 2.7 to 72.3 ± 1.6%, while the medium extracted from Outlet 2 contained 99.5 ± 0.3% 2-μm par-ticles with a recovery rate of 46.4 ± 3.6% (Figs. 4b, c, S3). The sorting results of 2- and 13-μm particles was quite simi-lar to the results above; namely, the purity of 13-μm particles was increased from Outlet 1 and 2-μm particles were sorting from Outlet 2 yielded a high purity. In these experiments, the volume recovery at Outlet 2 was 48.9%, which meant that approximately half of the medium was withdrawn from Outlet 2, containing purified small particles. The volume recovery can be enhanced by infusing the medium collected from Outlet 1 into the same device for multiple sorting pro-cesses. Alternatively, it could also be carried out by con-necting Outlet 1 of the first device to the input of the second one in a cascaded way. As the device can operate normally regardless of flow rate, the downstream chip with a lower flow rate input can still implement cell/particle sorting with-out sheath flow to increase its input.

20 µl min-1

50 µl min-1

100 µl min-1

200 µl min-1

mµ31mµ8 10 µm

500 µl min-1

2 µm

Le�

Right

Fig. 3 The movements of 2-, 8-, 10- and 13-μm fluorescent beads in the groove-based channel under flow rates ranging from 20 to 500 μl min−1. A syringe pump was used to provide the precise control

over the flow rate. Whereas small particles (2-μm particles) remain unfocused, larger particles (8–13-μm particles) were focused in the channel with slight changes under different flow conditions

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4.3 PBMCs enrichment and platelets purification

To demonstrate the capability of this manually operated device in biological sample processing, a mixture of PBMCs and platelets was used to test our device. Cells were manu-ally injected into the groove-based channel with the maxi-mum volume throughput of 300 ± 50 μl min−1 (Fig. 5a). The cell throughput of this hand-powered microfluidic device is about 50,000 cells s−1. After a single round, platelets were collected from Outlet 2 with 100% purity and the volume recovery from Outlet 2 was about 47% compared with the input (Fig. 5b). The PBMCs were enriched by a factor of 2.6 ± 0.8 (Fig. 5c). The enrichment was calculated as the ratio of the numbers of PBMC cells to platelet cells col-lected at the center outlet, divided by the same ratio at the inlet (Song et al. 2015). The enrichment ratio of PBMCs can be enhanced by multiple runs. The cell medium collected from Outlet 1 was reloaded to the inlet. After four rounds of pumping, the purity of PBMCs was enhanced from 0.8 to 10.4%, resulting in an enrichment ratio of 13.8. Besides, the volume recovery of platelets increased to 80 ± 4.2%.

Platelets have been long known to become activated dur-ing isolation (Parker et al. 1984). Since the critical stress for platelet activation is about 1000 dyn cm−2 under a given exposure time of 10 s (Hellums 1994), the maximum shear

stress of 750 dyn cm−2 and retention time of < 1 s are safe for platelets to migrate in the channel. In addition, we did not see any platelet aggregation in the experiments.

This groove-based channel has a simple structure and operation for cell sorting with many distinct advantages. First, the device operates at a continuous-flow, high-through-put mode and provides a physiological environment needed to process cells and minimize apoptosis. Furthermore, since this device is hand-operated, this passive microfluidic chip is portable for outfield settings, especially in developing coun-tries, by eliminating the bulky field generator and pump-ing systems. The current device, however, has fixed particle separable range. The particle manipulative range can be adjusted by changing the channel height for the separation of red blood cells and platelets. Since the intensive cell–cell interaction will heavily affect the inertial (Zhang et al. 2014) and hydrophoretic (Choi et al. 2008) focusing, dilution or sheath flow will be required to address this issue.

5 Conclusions

This work proposed a hand-powered microfluidic device for particle/cell sorting. The groove-based design incorporated into the microchannel seamlessly connects the hydrophoresis

8 µm 13 µm10 µm2 µm

200 µm

(a)

(b) (c)

Fig. 4 a The optical graphs show the distribution of 2-, 8-, 10-, 13-μm particles at the outlet. b The particle purity from inlet, Outlet 1 and Outlet 2 measured by flow cytometry. c Particle recovery for dif-

ferent particle mixtures. The average value was obtained by measur-ing three times, and the error bar represents standard deviation

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ordering and inertial focusing. The overlapped equilibrium positions of particles dominated by hydrophoretic ordering and inertial focusing enable a stable particle sorting per-formance regardless of flow rate. First, the particle distri-bution in the groove-based channel was investigated using differently sized (2, 8, 10, 13 μm in diameter) polystyrene particles under different flow rates (20–500 μl min−1). The particle distributions are largely identical over the large range of flow rate. The device, when powered manually by

hand, was used to sort particle mixture. Over 99% of larger particles (8, 10, 13 μm) were recovered from Outlet 1, and a high separation purity (> 99%) of small particles (2 μm) was achieved. Finally, the manually operated device was used to sort platelets from PBMCs. The platelets were collected with a purity of ~ 100%. The purity of PBMCs was enhanced from 0.8 to 10.4% after multiple processes, resulting in an enrichment ratio of 13.8. Since this simple microfluidic chip can be manually operated, this eliminates the use of bulky

Fig. 5 a A schematic of the manually operated groove-based channel for peripheral blood mononuclear cells (PBMCs) and platelets purification. b Flow cytometry data indicate the percentage of PBMCs and platelet numbers in a (i) unpro-cessed sample, (ii) Outlet 2 collection after the first process, (iii) Outlet 1 collection after the first process and (iv) Outlet 1 collection after fourth process. c PBMCs enrichment ratio and platelets volume recovery of groove-based channel for cell separation with multiple rounds

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syringe pumps and pressure controller; it can potentially be used as a portable tool for outfield diagnostics.

Acknowledgements This work was performed in part at the Queens-land Node of the Australian National Fabrication Facility, a company established under the National Collaborative Research Infrastructure Strategy to provide nano- and microfabrication facilities for Australia’s researchers. S.H Tan and N.T.N. gratefully acknowledge the support of the Australian Research Council Linkage Grant (LP150100153), DECRA Fellowship (DE170100600), Griffith University-Peking Uni-versity Collaboration Grant and Griffith University/Simon Fraser Uni-versity Collaborative Grant.

Authors’ contributions SY, YXL and SYT designed and conducted the experiments, SHT and JTT fabricated the microchannel, and QBZ, DY and RS helped in preparing cell sample. SY wrote the manu-script. WHL, NTN, RS, SYT, JZ and SYT revised and commented the manuscript.

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