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Cavity-induced microstreaming for simultaneous on-chip pumping and size-based separation of cells and particles† Maulik V. Patel,a Imaly A. Nanayakkara,a Melinda G. Simona and Abraham P. Lee*ab We present a microfluidic platform for simultaneous on-chip pumping and size-based separation of cells and particles without external fluidic control systems required for most existing platforms. The device utilizes an array of acoustically actuated air/liquid interfaces generated using dead-end side channels termed Lateral Cavity Acoustic Transducers (LCATs). The oscillating interfaces generate local streaming flow while the angle of the LCATs relative to the main channel generates a global bulk flow from the inlet to the outlet. The interaction of these two competing velocity fields (i.e. global bulk velocity vs. local streaming velocity) is responsible for the observed separation. It is shown that the separation of 5 μm and 10 μm polystyrene beads is dependent on the ratio of these two competing velocity fields. The experimental and simulation results suggest that particle trajectories based only on Stokes drag force cannot fully explain the separation

Received 15th April 2014, Accepted 30th July 2014

behavior and that the impact of additional forces due to the oscillating flow field must be considered to

DOI: 10.1039/c4lc00447g

determine the trajectory of the beads and ultimately the separation behavior of the device. To demonstrate an application of this separation platform with cellular components, smaller red blood cells (7.5 ± 0.8 μm) are separated from larger K562 cells (16.3 ± 2.0 μm) with viabilities comparable to those of controls based

www.rsc.org/loc

on a trypan blue exclusion assay.

Introduction An important function of future MicroTAS systems will be their ability to process samples with heterogeneous cellular and particulate content and separate populations of interest. This is an essential feature in applications ranging from fundamental research on cellular biology to clinical therapeutic and diagnostic applications.1–4 The current bulk lab-based separation of a heterogeneous population is done with the use of magnetic beads tagged with biochemical labels, such as the Dynabeads® or MACS® systems.5,6 Intrinsic properties of cells, such as density, have been used to perform bulk separation via centrifugation with density gradient solutions.7,8 Although these methods are widely adapted within the research and clinical healthcare communities, they have one or more of the following disadvantages: the need for specialized equipment, large sample volumes, time consuming protocols and trained personnel.1–3 Microfluidic-based separation platforms are promising because they can perform separation with small sample a

Department of Biomedical Engineering, University of California, Irvine, CA 92697, USA b Department of Mechanical and Aerospace Engineering, University of California, Irvine, CA 92697, USA. E-mail: [email protected] † Electronic supplementary information (ESI) available. See DOI: 10.1039/ c4lc00447g

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volumes, reduce reagent consumption and have smaller footprints. Additionally, microfluidics platforms are amenable to integration with upstream and downstream processes, which is essential for developing a complete Lab on a Chip (LOC) system. Furthermore, the length scales of microfluidics are on the same order of magnitude as the cells and particles being manipulated. These features make microfluidics-based separation platforms ideally suited for applications in cell and particle manipulation.1–4 Although separation with biochemical labels has been a widely used technique, “label-free” methods based on intrinsic properties of particulates including size, density, electrical properties, compressibility and stiffness are receiving increased attention. Their advantages include minimal perturbation to cells as well as reduced cost and time of analysis.2 Many microfluidic platforms have been developed to separate populations based on size by using physical barriers, DEP, primary acoustic radiation force and inertial forces.1–4 However, these platforms often require external fluidic pumping systems to deliver the sample through the separation device, which increases the device footprint and decreases their portability. The development of Point of Care (POC) devices would benefit significantly from a self-contained microfluidic platform that is capable of simultaneously pumping the sample through the device while separating components by size. Such a platform would be an ideal candidate for integration into a portable POC system.

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A self-contained microfluidic platform based on surface acoustic wave (SAW) techniques has been utilized to separate cells and particles by their differing mechanical properties in discrete sample volumes.9,10 These cells and particles have been shown to maintain their spatial locations after separation. However, the capability to collect the separated cells and particles relies on the evaporation of the droplet in which cells and particles are separated, and it is not clear how the evaporation would affect the viability of the cells. For many applications, the collection of viable cells is essential to downstream analysis. Microfluidic devices that utilize acoustically excited gas/liquid interfaces have recently received a lot of interest as actuators for the manipulation of fluids and particles.11–21 These are simple on-chip actuators that are easy to fabricate and when excited by an external acoustic energy source, generate localized streaming flows that can manipulate the fluid and suspended particles. Recently, aided by external syringe pumps they have been utilized to separate particles of different sizes within microfluidic devices.17 Our lab has developed a novel on-chip pump using an array of dead-end side channels termed Lateral Cavity Acoustic Transducers (LCATs).20,22 When the LCATs are oriented at an angle relative to the main channel, the local streaming flow generates a global bulk flow from the inlet well to the outlet well. In this manuscript, an LCAT separator device for simultaneous on-chip pumping and size-based separation of cells and particles without the need for any external fluidic control systems is demonstrated. The interaction of two competing velocity fields (i.e. global bulk velocity vs. local streaming velocity) is responsible for the observed separation. It is shown that the separation of 5 μm and 10 μm polystyrene beads is dependent on the ratio of these two competing velocity fields. The experimental and simulation results suggest that particle behavior based on Stokes drag force cannot fully explain the separation behavior and that the impact of additional forces due to the oscillating flow field (i.e. the flow due to the oscillations of the air/liquid interface at the applied acoustic frequency) must be considered in determining the trajectory of the beads and ultimately the separation behavior of the device. To demonstrate an application of this separation platform with cellular components, smaller red blood cells (7.5 ± 0.8 μm) are separated from larger K562 leukemia cells (16.3 ± 2.0 μm). The separated cells have viability that is comparable to those of controls based on a trypan blue exclusion assay. To the authors' knowledge, this is the first demonstration of size-based separation and downstream collection of viable cells in which fluid pumping and separation were performed simultaneously on-chip.

Cavity-induced microstreaming flow and the LCAT pump Cavity-induced microstreaming is a form of acoustic streaming in which the streaming flow is formed due to the viscous

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dissipation of acoustic energy by a bubble that is oscillating stably.23 It has recently been used as an actuation mechanism in microfluidic systems for the manipulation of fluids and suspended particulates.11–20,22,24–30 The oscillatory motion of the air/liquid interface leads to a first-order periodic flow within the liquid (Fig. 1A). At the bubble boundary the magnitude of the acoustically excited flow is at a maximum and is given by Uo ∼ dω

(1)

where d is the interface displacement amplitude and ω is the angular frequency of the acoustic field. The velocity of the oscillating flow field, Uo, is defined as the first characteristic velocity of cavity-induced microstreaming. The first-order periodic flow induces a steady second-order streaming flow (Fig. 1A).15,31–34 The magnitude of the steady streaming flow near the surface of the air/liquid interface is: Us ∼ Uo2/ωR

(2)

where Us is the velocity of streaming flow near the surface of the air/liquid interface and is defined as the second characteristic velocity of cavity-induced microstreaming. R is the length scale of the oscillating interface and can be defined as the equivalent radius of the air/liquid interface. Cavity-induced microstreaming flow is characterized by the streaming Reynolds number, which is defined as:34 Res = ρUsR/μ

(3)

where ρ is the density of the fluid phase, U s is the velocity of streaming flow close to the oscillating air/liquid interface, R is the length scale of the oscillating interface and μ is the dynamic viscosity of the fluid. When the streaming Reynolds number, Re s ≪ 1, the streaming flow is characterized by Rayleigh–Nyborg–Westervelt (RNW) streaming and can be analyzed as a Stokes flow.13,34 A flow in which the streaming Reynolds number, Re s > 1, is characterized by a solution to the full Navier–Stokes equation.26 As will be seen in later sections the parameters of the present study result in streaming Reynolds numbers, Re s ≥ 1, indicating that inertial effects will play a role in the behavior of the fluid and the particles within the microchannels. This makes cavity-induced microstreaming one of the few mechanisms by which inertial effects can be implemented within microfluidic systems.34 As previously described, we have developed a novel on-chip pump that utilizes the local streaming flow generated by an array of LCATs to produce a bulk flow within the microchannel from the inlet well to the outlet well (Fig. 1B).20 The total flow rate from the inlet to the outlet divided by the cross-sectional area of the microchannel is the mean velocity of bulk flow, U b, and is the third characteristic velocity pertaining to the work presented in this manuscript.

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Fig. 1 Characteristic velocities of cavity-induced microstreaming flow. (A) Uo (solid line) is the velocity of oscillating flow due to the oscillatory motion of the air/liquid interface in response to the incident acoustic wave. Us (dashed line) is the velocity of streaming flow which arises due to the net displacement of fluid parcels during each air/liquid interface oscillation cycle. (B) An array of angled LCATs generates a bulk flow from the inlet well to the outlet well. This characteristic velocity is Ub, the mean velocity of bulk flow. The line of symmetry (dashed black line) is indicated to show the symmetry of the LCAT pairs along the long axis of the microchannel.

Theoretical mechanism of size based separation Previous work has shown that acoustically excited bubbles or air/liquid interfaces can trap particles within the microstreaming flow.13,24,35,36 Recently however, it has been shown that particles that are trapped within the microstreaming vortices have trajectories dependent on particle properties such as size and density.16,17,26 Wang et al. demonstrated that the particle size influences their trajectory, such that larger particles occupy smaller orbits, while smaller particles occupy larger orbits.17,26 This previous work utilized cavities that are perpendicular to the microchannel and the velocity of the streaming flow, U s, was low enough to assume that the streaming Reynolds number, Res < 1 (i.e. streaming flow is Stokes flow). They explained that since particles are unable to penetrate into the gas phase of the air/liquid interface, particles incident upon the interface will be forced to cross fluid streamlines. Since streaming flow generates closed loop vortices, the center of particles of different sizes will be aligned with different closed loop streamlines resulting in size dependent orbits. For example, the center of a 10 μm particle will be forced into a streamline closer to the center of the vortex compared to that

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of a 5 μm particle.17,26 We have made similar observations of size-dependent trajectories within microstreaming flows with cavities oriented at an angle to the microchannel (ESI† Fig. S1). When Wang et al. introduced a bulk flow into the device using an off-chip syringe pump, they observed that particle release from the streaming flow was based on the ratio between the mean velocity of bulk flow, Ub, and the velocity of streaming flow, Us.17,26 The release or trapping of a given sized particle was determined using a value they defined as dgap. They noted that there is a critical streamline that separates closed streamlines from open streamlines. Based on Stokes flow assumptions, all the fluid below this critical streamline must flow through a narrow gap of width d gap between the bubble surface and the critical streamline. Since the particles cannot penetrate the bubble, particles with a diameter greater than 2 × d gap are pushed away from the bubble such that the center of the particles crosses into a closed streamline, causing it to trap. However, if the diameter of the particle is less than 2 × dgap, then the center of the particle will remain within an open streamline and is released from the streaming flow. Since the streaming Reynolds number, Res, is less than 1, the assumption was made that particles will behave like simple fluid tracers having the same trajectories as the fluid parcels they occupy.

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This separation mechanism can be applied to the LCAT separator device. The advantage of the LCAT separator is the symmetry of the air/liquid interfaces along the long axis of the microchannel. Taking the symmetry of the device into consideration, the threshold U b/U s ratio can be determined in order to separate particles of different sizes. Fig. 2 illustrates a modification to the separation mechanism used by Wang et al. as it applies to the LCAT separator device. Due to the symmetry, an assumption can be made that the total volumetric flow rate will be divided evenly along the long axis of the microchannel such that half of the volumetric flow rate will pass through a narrow gap on the bottom set of air/liquid interfaces and the other half on the top set of interfaces. Based on continuity theory, the total flow rate through half the cross-sectional area of the microchannel must equal the flow rate through a narrow gap at the air/liquid interface. The total volumetric flow rate through half the cross-sectional area (V1) of the microchannel is: W V1  U b   2

 H 

(4)

where W and H are the width and height of the microchannel, respectively. The total volumetric flow rate through a narrow gap at the air/liquid interface (V2) is: V2 = UsdgapH

(5)

By equating V 1 and V 2 and solving for d gap, the following expression is obtained:  U  W  d gap   b     Us  2 

(6)

In order to allow 5 μm diameter particles to pass by the air/liquid interface, while trapping 10 μm particles, a d gap between 2.5 μm and 5 μm would be required. If the total width of the microchannel is ~500 μm, then 5 μm particles

should begin to release from the vortices when Ub/Us = 0.01, and 10 μm beads should begin to release from the vortices when U b/U s = 0.02. Therefore, an U b/U s ratio between 0.01 and 0.02 would allow 5 μm particles to release from the vortices while 10 μm particles are trapped within the vortices resulting in the separation of particles based on size. The separation mechanism proposed by Wang et al. is based on the assumption that the particles behave like simple fluid tracers within the streaming flow field (Us). However, recent theoretical and numerical analysis has shown that the oscillating flow field (Uo) has an impact on the particle trajectories such that the particles deviate from simple fluid tracer behavior.37–39 These studies accounted for the impact of additional force terms of the Maxey–Riley equation on particle trajectories, but were limited to Re s ≤ 1. The Maxey–Riley equation consists of the following terms:40 mp

dVp dt

 FD  FAM  FFA  FH  FB

(7)

where m p is the mass of the particle, dV p/dt is the acceleration of the particle, F D is the force on the particle due to drag, FAM is the force on the particle due to the added mass, F FA is the force on the particle due to the fluid acceleration from the undisturbed flow field, FH is the Basset history force and F B is the force due to buoyancy. Since the solution in which particles are suspended in our system was density matched to the particles' density (~1.05 g mL−1), the buoyancy force can be neglected. Recent theoretical and numerical work has shown that the history term, F H is oriented in the same direction as the force on the particle due to the added mass, FAM.37 Therefore, for the purpose of determining the deviation of particle behavior from simple fluid tracers (i.e. Stokes flow predictions), we only consider one of these two forces, F AM. The modified equation of motion of the particle used in this study includes the drag force, the added mass force due to the disturbance flow field and the fluid acceleration force due to the undisturbed flow field. mp

dVP  FD  Re p   FAM  FFA dt

(8)

FD(Rep) = 3π*μdpφ(Rep)(u − Vp) where φ(Rep) = (1 + 0.15 × Rep0.687) (9) FAM 

1  Du dVP  mf    2  Dt dt 

FFA  mf

Fig. 2 Applying Stokes flow separation mechanism to the LCAT platform. In order to release 5 μm particles while trapping 10 μm particles the Ub/Us ratio must be between 0.01 and 0.02. The width of the microchannels is W = ~500 μm. Dashed black line shows that the line of symmetry in the device is coincident with the long axis of the microchannel.

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Du Dt

(10)

(11)

where μ is the dynamic viscosity of the fluid, dp is the diameter of the particle, φ(Rep) is a correction to the Stokes drag equation due to a finite particle Reynolds number (Rep), u is the velocity of the undisturbed flow field at the particle center, Vp is the velocity of particle, mf is the mass of the fluid which must be displaced due to the presence of the particle and

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Du/Dt is the material derivative of the fluid. By rearranging the terms, the following equation emerges. mp

dVp dt

1 dV  3  du  mf p  3* d p  Re p   u  Vp    mf  dt 2  2  dt

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3    mf   u  Vp u 2 

(12)

The second term on the left hand side of eqn (12) represents the particle having an added mass of 0.5 times the mass of the fluid that it displaces. Also, due to impact of the added mass and the fluid acceleration, there are additional force terms on the particle (2nd and 3rd term on the right hand side). In the next section, computational fluid dynamic (CFD) simulations are conducted in order to determine the impact of these additional forces on particle trajectories and the resulting deviation of these particles from acting like simple fluid tracers.

Computational fluid dynamic simulations In order to determine the impact of the additional force terms due to the oscillating flow field on the trajectory of particles, 2-dimensional CFD simulations were conducted using CFD-ACE+ v2011 (ESI Group, Inc., France).

Simulation setup The simulation was set up using the fluid flow and particle spray modules on a single angled LCAT pair. The microchannel was 250 μm wide and 2200 μm long with a structured grid of 10 μm × 10 μm throughout as shown in ESI† Method 1. The reference pressure at the inlet and the outlet of the control volume were set to 0 kPa. There was a symmetry boundary condition applied to the top wall of the control volume to simulate the symmetry in the flow due to an angled LCAT pair. A no-slip boundary condition was applied to all the other walls within the control volume. The oscillation of the interface was modeled as an inlet with a time varying velocity profile. As shown in ESI† Video 1, the interface oscillates as a traveling wave which originates from the tip of the LCAT to the opposite wall on the other side of the air/liquid interface. This was modeled in the simulation as an inlet with a time varying velocity profile. The details of the model setup are given in ESI† Method 1. Simulation results Fig. 3A shows the measured streaming velocity in the simulations based on the displacement of a 5 μm diameter particle over a single oscillation cycle. The line shows a quadratic fit (R 2 = 0.99) to the streaming velocity as a function of

Fig. 3 LCAT separator simulation results. (A) A quadratic relationship is observed between the velocity of streaming flow, Us, and the amplitude of interface displacement at the time varying inlet (R 2 = 0.99). (B) Sign convention to quantify the deviation of 5 μm particles (green) from the reference trajectory (blue arrow) of 10 nm particles (blue). Particles with an inward trajectory have a negative deviation, while particles with an outward trajectory have a positive deviation. (C) When only the drag force (FD) is activated, 5 μm particles have a slightly inward trajectory. When drag and added mass (F D + F AM) are activated, 5 μm particles begin to spiral outwards. When all three are activated (F D + F AM + F PG) there is an additional deviation outwards due to the pressure gradient term. The deviation as the particle flows past the interface is on the order of several microns, sufficient to allow 5 μm particles to cross from a closed streamline into an open streamline as indicated in Fig. 2. (D) With only the drag force, trajectories of 10 nm and 5 μm particles essentially follow the same path. With the additional force terms, 5 μm particle trajectories deviate outwards within the microstreaming flow. Interface displacement amplitude (d) is 10 μm.

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the interface displacement amplitude, d. To determine the impact of the additional force terms on the trajectories of particles, 10 nm and 5 μm particles were tracked in the vicinity of the LCAT interface. The additional force terms that were incorporated into the solver were the added mass term and the pressure gradient term. As described in eqn (12), the added mass term is the force contribution due to the fluid that must be displaced as the particle accelerates and decelerates in the oscillating flow field, Uo. The pressure gradient term was activated to partially account for the force on the particle due to the fluid acceleration of the undisturbed oscillating flow field.40 The justification is that the force on the particle due to the fluid acceleration term can be separated into the force on the particle due to the pressure gradient and the force on the particle due to viscosity.40 Since the commercial software (CFD-ACE+ v2011) was only able to account for the pressure gradient term, it was only possible to partially account for the force due to the fluid acceleration of the undisturbed oscillating flow field. Although, this may limit the degree of quantification that can be inferred from the simulations, previous theoretical and numerical results have shown that FFA is in the same direction as FAM. Therefore if the particle trajectories continue their trend as each additional force term is incorporated, it can be concluded that fully incorporating the FFA term will result in a larger deviation in the same direction as that observed in the simulations. In one simulation, only the drag force was activated, in a second the drag and added mass force was activated, while in the third simulation, the drag, added mass and the pressure gradient component (FPG) of the fluid acceleration term were activated. Fig. 3B shows the sign convention that was utilized to analyze the results of these simulations. If the 5 μm particle (green) has an inward trajectory compared to the 10 nm particle (blue) then the deviation has a negative value. However, if the 5 μm particle has an outward pointing trajectory then the deviation has a positive value. Using this convention, Fig. 3C shows the resultant particle trajectories from the simulations above. The particle deviation was calculated after the particle had traveled past the air/liquid interface. When only the drag force is activated, 5 μm particles have a negative deviation meaning the particles are traveling towards the center of the vortex. When drag and added mass force are activated, the particles have a positive deviation and have an outward pointing trajectory due to this additional force. When the drag, added mass and the pressure gradient terms are activated, the particles have an additional deviation outwards. The results from Fig. 3C show that as the particle moves past the interface, these additional forces are causing the particle to deviate outwards on the order of several microns, which would be sufficient to allow 5 μm particles to migrate from a closed streamline to an open streamline, allowing them to release from the microstreaming vortices prior to the predictions based on only Stokes drag force separation mechanisms. Fig. 3D shows the trajectory of particles when only 1) drag force and 2) drag, added mass and pressure gradient terms are activated. Due to the small size of the 10 nm particle, it

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behaves as a simple fluid tracer particle (trajectory in blue) in both cases. When only the drag force is activated, the 5 μm particles (trajectory in magenta) essentially follow the same trajectory as the 10 nm particles, indicating that larger particles in the flow will follow the same trajectories as the fluid parcels they occupy. When the additional terms are activated, 5 μm particles are pushed outwards (trajectory in green). These simulation results show that when taking into account the oscillatory flow that gives rise to the streaming flow at Res > 1, larger particles do not follow the trajectories of the fluid parcels they occupy, but rather are pushed outwards by the additional forces exerted on them due to the oscillating flow field. Based on these results it can be concluded that the forces exerted on the particles from the oscillating flow field will cause particles to spiral outwards resulting in their release from the vortices at a lower U b/U s than predicted by Stokes flow assumptions (i.e. separation mechanism proposed by Wang et al.17,26). Therefore 5 μm beads should begin to release from the vortices at U b/U s ratio less than 0.01, while 10 μm beads should release from the vortices at U b/U s ratio less than 0.02.

Materials and methods Device fabrication Devices were fabricated using standard soft-lithography techniques. SU-8 2100 (MicroChem Corp., USA) was spin coated onto a 3" silicon wafer to a thickness of 100 μm according to manufacturer protocol. A transparency mask was used to pattern the SU-8 using a flood exposure UV lamp. The developed SU-8 masters were hard baked on a hot plate for ~30 minutes at a temperature of 150 °C. SU-8 masters were then treated with a hydrophobic silane solution in a vacuum chamber for 2 hours using (tridecafluoro-1,1,2,2-tetrahydrooctyl) trichlorosilane (Gelest Inc., USA). A PDMS mixture (Dow Corning, USA) of 10 : 1 ratio of base to curing agent was thoroughly mixed and degassed in a vacuum chamber for 1 hour before being poured onto the SU-8 master. The PDMS was allowed to cure in a 65 °C oven for at least 24 hours. The devices were cut and inlet and outlet holes were punched using a 5 mm hole puncher. The coverslips (Fisher Scientific, USA) and the cut PDMS devices were first cleaned in acetone, methanol and isopropyl alcohol for 20 minutes each and dried using nitrogen. The cleaned PDMS was then baked for four days at 120 °C in order to fully cure the PDMS. To make the coverslips hydrophobic they were coated with Rain-X® (SOPUS Products, USA) per manufacturer recommendation and were placed at ambient temperature for 24 hours before use.41 The PDMS devices were then contact bonded to the Rain-X® coated coverslips (Fig. 4A). Experimental setup The external acoustic energy source consisted of a piezoelectric transducer (SMD43T105F200S, Steiner and Martins, Inc., USA). A 1" × 1" silicon wafer was glued (Krazy Glue, USA) onto the piezoelectric transducer to make the transducer surface

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Fig. 4 Experimental setup of LCAT separator. (A) Image of completed PDMS/coverslip device. (B) Device is coupled to a piezoelectric transducer using ultrasound gel. Transducer is actuated using a function generator and voltage amplifier. (C) Inlet and outlet bead concentrations are measured to obtain % particle trapping data. Ub and Us are measured for each device under test.

more reflective. This allowed for high quality video capture from an upright microscope for image analysis. The piezoelectric transducer was mounted onto a 3" × 5" glass slide (Fisher Scientific, USA) using double sided foam tape (3M Company, USA).

LCAT separator characterization In order to characterize the trapping curves for particles of different sizes, mixed suspensions of 5 μm and 10 μm polystyrene beads (density = 1.05 g mL−1) (SPI Supplies, USA) at a concentration of ~1000 particles μL−1 each, were suspended in a density matched solution of Ficoll-Paque PLUS (GE Healthcare, USA) & PBS (Life Technologies, USA) with 6mM EDTA (Sigma-Aldrich, USA) (termed F-PE from here onwards). The solution consisted of a 2 : 1 ratio of FicollPaque PLUS to PBS with 6mM EDTA. This F-PE solution was used because the density of this solution is ~1.05 g mL−1. The F-PE solution density was matched to the density of the particles in order to eliminate the experimental bias that would be caused by the sedimentation of 10 μm particles at a faster rate than 5 μm particles. This difference in sedimentation rates would cause more 5 μm particles to enter the microfluidic channel than 10 μm particles, biasing the results. However, changing the suspension fluid may not be favorable in cases where biological particulates are required to be separated. Therefore, to show the versatility of the platform, experiments are conducted for size-based cell/cell separation with culture media and PBS. The dynamic viscosity of the F-PE solution was determined to be μ = 1.28 × 10-3 Pa s using a falling ball viscometer (Thermo Scientific, USA) and the details are described in ESI† Method 3. Initial particle solution concentrations were measured using a hemocytometer. Approximately 200 μL of 5% BSA in DI H 2O (Sigma-Aldrich, USA) solution was incubated at the

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outlet reservoir for 5 minutes in order to eliminate the impact of pressure gradients due to the size of fluid droplets at the inlet and the outlet. The BSA solution was then completely removed from the outlet and 10 μL of F-PE solution was pipetted in the outlet reservoir. A vacuum was pulled at the inlet well using a 1 mL syringe in order to prime the device. During this process, the LCATs form in dead-end side channel structures that trap an air pocket and create an air/liquid interface used to generate the microstreaming flow. 40 μL of the bead suspension was pipetted into the inlet reservoir. The primed devices were coupled to the piezoelectric transducer with the use of AquaSonic Clear (Parker Labs Inc., USA) ultrasound gel (Fig. 4B). The device/piezoelectric transducer platform was placed under an upright microscope. The piezoelectric transducer is supplied with a 50.3 kHz square wave signal from a function generator (Agilent 33220A, Agilent Technologies, USA) and voltage amplifier (Krohn-Hite 7500, Krohn-Hite Corp., USA) with voltages ranging from 3–5 Vpp after amplification. The activated transducer caused the air/liquid interface of the LCATs to oscillate, creating cavity-induced microstreaming flow within the microchannel. Multiple videos were captured using a high speed camera (Phantom v310, Vision Research Inc., USA) at a frame rate of 100 000 fps with an exposure time of 9 μs to measure the velocity of streaming flow, Us of particles near the air/liquid interface. Videos were captured of LCAT pairs 25 and 27 as they were located approximately at the center of the device/transducer setup (Fig. 4C). The total run time of the device was monitored. The total volumetric flow 83 μL ± 2 μL (ESI† Method 2) and the total run time of the device were used in order to determine the mean velocity of bulk flow, U b throughout the experiment. After the inlet reservoir was emptied of the bead suspension solution, the piezoelectric transducer was turned OFF for 5 seconds while 40 μL of wash solution of F-PE was pipetted into the inlet well and the piezoelectric transducer

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was turned back ON again. After the inlet reservoir was emptied of the wash solution, the outlet well was thoroughly mixed with a pipette and the concentration of 5 μm and 10 μm beads in this outlet solution were determined using a hemocytometer. The ratio of the total number of 5 μm and 10 μm particles at the outlet and inlet were used to determine the trapping curve for both 5 μm and 10 μm particles. K562 cell preparation K562 leukemia cells (ATCC, USA) were cultured in media consisting of RPMI 1640 (Life Technologies, USA) supplemented with 10% fetal bovine serum (Life Technologies, USA) and 1% penicillin streptomycin (Life Technologies, USA) under standard tissue culture conditions. K562 cells were then harvested and resuspended in culture media at a concentration of ~1000 cells μL−1. Initial cell counts and viability were taken using a trypan blue exclusion assay. K562 cell viability and yield characterization A solution of 5% BSA in DI H 2O was placed in the inlet and outlet of the device and incubated at room temperature for 5 minutes to prevent non-specific cell adhesion in the wells. Afterwards, the BSA solution was completely removed and the devices were filled with PBS solution. 40 μL of K562 cell suspension was pipetted in the inlet and the device was coupled to the piezoelectric transducer using ultrasound gel. The piezoelectric transducer is supplied with a 50.3 kHz square wave signal from a function generator and voltage amplifier with voltages ranging from 1.0–4.5 Vpp after amplification. Once the inlet well emptied, the waste fluid pumped to the outlet (which did not contain cells due to their trapping in the device) was removed. The inlet well was washed with 40 μL of PBS and this solution was labeled as the “sedimented” population in order to account for the sedimentation of cells in the inlet well. The cells trapped in the device were collected into 40 μL of PBS solution and labeled as the “trapped” population. The viability of the cells in the “trapped” population was determined using a trypan blue exclusion assay. Cell counts were taken from both the “sedimented” and “trapped” populations and summed together to calculate the total number of cells collected from the device after the completion of the experiment. The yield of the device is the total number of cells collected out of the device from the “sedimented” and “trapped” populations

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divided the total number of cells initially pipetted into the device. The yield accounts for the loss of cells due to complete cell lysis which could not be accounted for by the trypan blue exclusion assay.

Dual array for viable cell/cell separation In applications of cell separation where the viability of the cells is important for downstream processing and analysis, the single array of angled LCATs is not capable of separating cells and ensuring sufficient yield and viability. In order to address this concern, a cell separator platform consisting of a dual array of angled LCATs was fabricated as shown in Fig. 5. The dual array device is capable of independent control of Ub and Us such that at lower applied voltages to the piezoelectric transducer, enough bulk flow can be generated in order to separate cells. K562 cells were cultured, harvested and resuspended to a concentration of ~2000 cells μL−1 in RPMI 1640 + 10% FBS + 1% penicillin/streptomycin (Life Technologies, USA). Red blood cells (RBCs) were collected from healthy human donors by a trained nurse and resuspended at a concentration of ~2000 cells μL−1 in phosphate buffered saline (PBS). K562 cell suspension and RBC suspension were mixed at a 1 : 1 ratio to obtain a suspension of ~1000 cells μL−1 for each cell type. Initial cell counts were taken as well as the initial viability using a trypan blue exclusion assay. A solution of 5% BSA in DI H2O was incubated in the cell injection well for 5 minutes to prevent non-specific cell adhesion in the wells. Afterwards, the BSA solution was removed and the devices were filled with PBS. A 40 μL solution of the mixed cell suspension is pipetted into the cell injection well and scotch tape is used to seal the well closed. 80 μL of PBS is pipetted at the pumping fluid well and the collection outlet and piezoelectric transducer #1 is turned ON with an applied voltage of 2.0 V pp and a square wave signal of 50.3 kHz. As the mixed cell suspension flows through the trapping region, piezoelectric transducer #2 is turned ON with an applied voltage of 1.0 V pp. After both cell types are observed to be trapping within the vortices, the applied voltage on piezoelectric transducer #1 is gradually increased until it is visually observed that RBCs are released to the collection outlet and K562 cells stay trapped in the selective trapping region. When array #2 is fully saturated with K562 cells, the initial solution at the collection outlet containing RBCs is collected into a centrifuge tube labeled “outlet”. Then piezoelectric transducer #2 is

Fig. 5 Illustration of the dual array device for cell separation. Array #1 is operated at a higher voltage to generate a large mean velocity of bulk flow, Ub. Array #2 is operated at a lower voltage to generate a low velocity of streaming flow, Us. This maintains cell viability during the separation process while generating enough bulk flow to allow for separation to take place. Array #1 and #2 consisted of serpentine channels with a total of 60 LCAT pairs in each array.

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turned OFF allowing the trapped K562 cells and some RBCs to flow to the collection outlet. This solution is resuspended in 40 μL of PBS solution and collected into a different centrifuge tube labeled “trapped”. The two samples were counted using a hemocytometer and were tested for viability using a trypan blue exclusion assay.

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Experimental results and discussion LCAT separator characterization The particle trapping curve for 5 μm and 10 μm beads based on the voltage applied to the piezoelectric transducer is

Lab on a Chip

shown in Fig. 6A. At 3 V pp, 100% of 5 μm and 10 μm beads are trapped within the microstreaming flow. At 4 V pp, ~50% of 5 μm particles begin to be released to the outlet. At 4.5 Vpp, only about 7% of the 5 μm particles are maintained within the microstreaming flow, leaving the other ~93% to be released to the outlet and at 5 V pp we begin to start observing the release of 10 μm particles to the outlet. Voltages higher than 5 V pp caused cavitation nucleation within the microchannel walls, as previously observed in microfluidic devices.42,43 ESI† Video 2 shows several LCAT pairs within the center of the separator device while the device is in operation.

Fig. 6 Bead separation characterization results. (A) As the voltage applied increases there is a decrease in the trapping of 5 μm beads. At an applied voltage of 5 Vpp, some 10 μm beads are also being released to the outlet reservoir. (n = 3). (B) Measured velocity of streaming flow, Us, at the air/liquid interface. Quadratic relationship (R 2 = 0.99) conforms to theoretical predictions. (n = 3). (C) Measured mean velocity of bulk flow, Ub, from the inlet well to the outlet well. Quadratic relationship (R 2 = 0.99) agrees with previous results.34 (n = 3). (D) Ub versus U s plotted for each device under test (R 2 = 0.98). Black line indicates the U b/U s = 0.01 threshold for 5 μm bead release based on Stokes flow assumptions. (E) Data shows the dependence of particle trapping on the U b/U s ratio. In (D) and (E) (dotted line in blue) shows the parameters at which 5 μm beads should begin to release to the outlet under Stokes flow separation mechanisms. However, experimental results show that 5 μm beads begin to release to the outlet at a lower Ub/Us than predicted (dotted line in magenta). 5 μm beads begin to release at Ub/Us ≈ 0.007, while 10 μm beads begin to release at Ub/Us ≈ 0.01. These experimental results agree with simulation results which show an outward pointing particle trajectory on the order of several microns due to the activation of additional force terms. For (A)–(D) error bars are standard error of the mean (SEM). For (E) error bars are plus = (Ub + SEM)/(Us − SEM) and minus = (Ub − SEM)/(Us + SEM).

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Fig. 6B and C shows the dependence of two characteristic velocities, Us (velocity of streaming flow) and Ub (mean velocity of bulk flow), on the applied voltage to the piezoelectric transducer. Both characteristic velocities increase quadratically as the voltage is increased (R 2 = 0.99). The quadratic relationship between U s and applied voltage (Fig. 6B) conforms with the theoretical predictions for cavitation microstreaming flow.13,17,26 The quadratic relationship between Ub and the applied voltage (Fig. 6C) is consistent with previously published results for an LCAT pump platform.20 With ρ = 1.05 g mL−1, U s = [0.62–1.47 m s−1], R = 25 μm (0.5 × cavity width) and μ = 1.28 × 10−3 Pa s, the streaming Reynolds number, Res, calculated using eqn (3) ranges from 13-30, indicating that Stokes flow based separation mechanisms cannot fully determine the trajectories of the particles and that inertial effects need to be considered. For each device under test, the data from Fig. 6B and C can be plotted to view the relationship between the mean velocity of bulk flow (U b) to the velocity of streaming flow (U s). These results are plotted in Fig. 6D and show a quadratic relationship between U b and U s as illustrated by the best fit line in red (R 2 = 0.98). The black line represents Ub/Us = 0.01, the threshold based on Stokes flow assumptions for 5 μm particles to begin to release while maintaining 10 μm particles within the vortices. At lower values of U s, the devices are operating at Ub/Us < 0.01. As Us increases, the Ub/Us ratio for each device begins to approach the predicted ratio for 5 μm particle release eventually crossing it as indicated by the best fit line (red line) intersecting the U b/U s = 0.01 line (black line). At the intersecting point, the devices under test are operating at a Ub/Us = 0.01 (denoted by the dotted blue line). Experimental results show that we observe the beginning of 5 μm beads releasing from the vortices for devices which are operating well below this intersecting point (denoted by the dotted magenta line). Fig. 6E shows the trapping curve of both 5 μm and 10 μm particles as it depends on U b/U s. The data shows that the release of 5 μm particles to the outlet reservoir begins to occur at Ub/Us ≈ 0.007 (denoted by the dotted magenta line).

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It is also observed that full separation of 5 μm particles occurs just prior to the ratio achieving the predicted 0.01 level. In addition, it should be noted that at a U b/U s ratio of 0.01, 10 μm particles are starting to be released to the outlet well. This ratio is approximately half of what is required for the release of 10 μm particles based on simple Stokes flow assumptions as previously described. These results correlate with the simulation results in that when the additional force terms (i.e. F AM and F PG) are incorporated in addition to the drag force, the particles have an outward pointing trajectory on the order of several microns compared to the trajectory of fluid parcels they originally occupied. The results from simulations and the experimental observations show that in order to more accurately predict the behavior of the LCAT separator platform the Stokes flow assumption of particles following fluid parcel trajectories has to be modified and must take into account the impact of additional force terms on particle behavior due to the oscillatory flow.

K562 cell viability characterization Fig. 7 shows the viability and yield of K562 cells in relation to the voltage applied to the piezoelectric transducer. The data shows that as the voltage applied to the piezoelectric transducer increases from 1.0 V pp to 4.5 V pp there is an approximately linear decrease in the viability of the cells based on trypan blue exclusion assay. A maximum viability of 94% is achieved at a voltage of 1.0 Vpp, while a minimum viability of 79% is observed at an applied voltage of 4.5 Vpp. The total yield of the cells that have been flowed through the device also decreases as the voltage applied to the device is increased. Visual observations showed that as the cells were flowing through the device at 4.5 Vpp they were beginning to clump into large spheres of cellular debris the size of several hundred microns, indicating that the shear forces on the cells at the higher voltages were likely causing the cells to lyse completely and clump together (Fig. 7). In order to determine if the surface temperature of the piezoelectric transducer was the cause of low yield, an infrared temperature gun (Craftsman,

Fig. 7 Cell viability and yield versus applied voltage. The viability and yield of cells as they are flowed through the device decreases linearly as the voltage applied to the piezoelectric transducer is increased. The inset to the left shows trapped cells at 1 Vpp where individual cells are orbiting within the flow. The inset to the right shows trapped cells at 4.5 Vpp where large clumps of cellular debris are trapped in the flow. This is evidence of complete cell lysis at higher voltages which impacts the total number of cells collected from the device. Data is an average of 3 independent experiments and error bars are SEM. The scale bar is 500 μm.

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USA) was used to measure the temperature while operating at 5 Vpp. The average surface temperature before actuation was 22.1 °C, after 5 minutes of actuation it was 22.3 °C and after 30 minutes of actuation it was 22.7 °C, therefore temperature was not the cause of low cell viability and yield.

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Dual array for viable cell/cell separation In Fig. 7 it was demonstrated that at voltages where separation is observed (i.e. 4.5 V pp) a single array of angled LCATs would not be capable of separating cells while maintaining a high rate of viability. In order to address this concern, a cell separator platform consisting of a dual array of angled LCATs was fabricated as shown in Fig. 5. The dual array device is capable of independent control of U b and U s such that at lower applied voltages to the piezoelectric transducer, enough bulk flow can be generated in order to separate cells. Fig. 8 shows the results of the separation of K562 cells (16.3 ± 2.0 μm)

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and red blood cells (7.5 ± 0.8 μm) using the dual array device. As seen from Fig. 8A, the fluid entering array #2 from the cell injection well consists of both K562 cells and RBCs, however, the fluid leaving array #2 and into the collection outlet consists of only RBCs (Fig. 8B). Fig. 8C shows a micrograph of a single LCAT. It is apparent that particulates trapped within the microstreaming vortices are K562 cells while the particulates that are being released from the microstreaming vortices are the smaller RBCs, based on the size of particulates in these locations. A stacked image is also shown in Fig. 8D to show the separation phenomena over several frames from the high-speed video recording. The results of a trypan blue exclusion assay on the cells collected from the trapped solution showed a viability of 94–100% over three different experimental runs while the viability of cells initially pipetted into the cell injection well was 99%. These results show that with the dual array device the viability of the cells in the collected population was

Fig. 8 K562s and RBCs separation using dual array device. (A) K562s and RBCs flowing into array #2 from the cell injection well. (B) Only RBCs flow out of array #2 into the collection outlet. (C) Shows a magnified view of K562s being trapped in the microstreaming flow while RBCs are being released. (D) Stacked images of (C) showing trajectory of K562s trapping and RBCs releasing. (E) Cell suspension in inlet contains both cell types. (F) The trapped solution contains mainly K562s with a few RBCs being trapped. (G) The Outlet solution contains only RBCs. For (E)–(G), at least 120 cells were counted from each solution (inlet, trapped and outlet) for all 3 experimental runs.

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comparable to the viability of the cells in the initial population, demonstrating the capability of this platform to separate two different cell types based on their size while maintaining their viability. Fig. 8E shows the presence of both cell types in the solution that was initially pipetted into the cell injection well, while the sample collected as “Outlet” only consists of RBCs (Fig. 8G) and the sample collected as “trapped” consists of mostly K562 cells with a few RBCs (Fig. 8F). An average enrichment ratio of 7.1 ± 2.7 (SEM) was achieved for multiple runs (n = 3). The enrichment ratio in this case is calculated as the ratio of K562 cells to unselected RBCs of the “trapped” solution divided by the same ratio in the cell injection well.2 The average total number of K562 cells that were trapped within the microstreaming vortices was ~15 800 cells over the 3 experimental runs. Since the separation experiments were completed when all the LCAT pairs were saturated with K562 cells in the selective trapping region of the device, it can be calculated that each LCAT is capable of trapping approximately 130 cells (15 800 cells/120 LCATs) prior to reaching saturation. This platform demonstrates that the LCAT separator device can perform simultaneous on-chip pumping and separation of cells based on size, while ensuring that the cells collected post-separation are intact and viable for downstream analysis. The capability to simultaneously pump the sample and separate particulates on-chip with a simple, passive, low-cost device makes the LCAT separator ideal for integration into a portable POC platform for sample processing.

Conclusions The ability to separate specific populations of cells and other particulates from a complex heterogeneous mixture in a high-throughput manner is an essential step in biological research and in clinical applications. Current bulk lab-based separation methods include separation via the use of antibody-coated magnetic beads and intrinsic properties of cells such as density using centrifugation. However, these separation methods have been challenging to integrate with upstream and downstream sample processing and analysis platforms. Since the beginning of MicroTAS development, the goal has been to integrate multiple sample processing and analysis modules onto a single chip format, one of which is the bulk separation of specific cells and particulates from a complex heterogeneous mixture. A wide variety of microfluidic bulk separation methods have been developed, however many of these modules require the need for external pumping systems in order to deliver the fluid to the device where separation takes place. The continued development of MicroTAS platforms that perform complete sample-to-answer analysis would benefit from a self-contained module that simultaneously performs on-chip pumping and separation of cells and particles of interest that can be collected for downstream analysis.

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This manuscript demonstrates that the LCAT separator device can separate cells and particles based on size while simultaneously performing on-chip pumping. The LCAT separator was first characterized using 5 μm and 10 μm beads to show that the particle trapping curve is dependent on the ratio between the mean velocity of bulk flow, U b and the velocity of streaming flow, Us. As the ratio increases, there is a gradual release of the smaller 5 μm particles while the larger 10 μm particles continue to stay trapped. At the highest ratios of Ub/Us, the larger 10 μm particles also begin to release to the outlet. It has also been shown that the streaming Reynolds number, Re s, of the LCAT separator device lies between 13–30, meaning that the previously suggested mechanism for particle separation based on Stokes flow assumptions cannot fully explain the separation thresholds. Indeed, experimental observations showed that the release of both 5 μm and 10 μm particles begins to occur at a lower U b/U s ratio than predicted based on Stokes flow assumptions. It was shown through simulation results that the additional forces (i.e. the added mass term and the pressure gradient component of the fluid acceleration term) result in an outward pointing particle trajectory on the order of several microns compared to the fluid parcels the particles originally occupied, therefore causing particles to release at a lower Ub/Us ratio than expected based on Stokes flow assumptions. We also demonstrated that larger K562 cells (16.3 ± 2.0 μm) could be separated from smaller RBCs (7.5 ± 0.8 μm) using a dual array device while ensuring the viability of the cells is comparable to that of controls. To the authors' knowledge, this is the first demonstration of size-based viable cell separation in which fluid pumping, separation and downstream collection were performed simultaneously on-chip. Future work on the development of the LCAT separator platform should focus on optimization of the pumping capabilities of the LCAT separator, which will allow for greater control of U b/U s ratio at a lower applied voltage to the piezoelectric transducer giving further control over the range of cell and particulate sizes that can be separated. To enhance understanding of the separation mechanism, a detailed theoretical and numerical model that takes into account the impact of streaming Reynolds numbers, Re s > 1 should be developed in order to account for the range of streaming flows that can be generated by acoustically excited air/liquid interfaces in microfluidic devices. This would further our understanding of the behavior of particles in microstreaming flows, while making it possible to use other intrinsic properties of cells and particles such as density, deformability and stiffness for the purposes of separation.

Acknowledgements The authors thank Dr. Roger H. Rangel (UC Irvine) for insightful discussions and support. This work was partially funded by the Micro/Nano Fluidics Fundamentals Focus (MF3) Center under the DARPA N/MEMS Science and Technology Fundamentals Program, grant no. N66001-10-1-4003.

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