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Journal of Chromatography A, 1406 (2015) 244–250

Contents lists available at ScienceDirect

Journal of Chromatography A

jo ur nal ho me pag e: www.elsev ier .com/ locate /chroma

icrofluidic device for sheathless particle focusing and separationsing a viscoelastic fluid

eonghun Nama,1, Bumseok Namgunga,1, Chwee Teck Lima,b, Jung-Eun Baec,wa Liang Leoa, Kwang Soo Choc, Sangho Kima,d,∗

Department of Biomedical Engineering, National University of Singapore, 9 Engineering Drive 1, Singapore 117575, SingaporeMechanobiology Institute, National University of Singapore, Singapore 117411, SingaporeDepartment of Polymer Science and Engineering, Kyungpook National University, Daegu, South KoreaDepartment of Surgery, National University of Singapore, 1E Kent Ridge Road, Singapore 119228, Singapore

r t i c l e i n f o

rticle history:eceived 20 March 2015eceived in revised form 10 June 2015ccepted 12 June 2015vailable online 19 June 2015

eywords:

a b s t r a c t

Continuous sheathless particle separation with high efficiency is essential for various applications suchas biochemical analyses and clinical diagnosis. Here, a novel microfluidic device for highly efficient,sheathless particle separation using an elasticity-dominant non-Newtonian fluid is proposed. Our deviceconsists of two stages: sheathless three-dimensional focusing (1 st stage) and separation (2nd stage). Itis designed based on the principle of a viscoelasticity-induced particle lateral migration, which promisesprecise separation of particles in a microfluidic device. Particles of 5- and 10-�m diameters were all

icrofluidicson-Newtonian fluidiscoelastic separationlood component separation

focused at the centerline of a circular channel at the 1st stage and successfully separated at the 2nd stagewith an efficiency of ∼99.9% using size-based lateral migration of particles induced by the viscoelasticity ofthe medium. We also demonstrated the capability of our device for separation of blood cells into multiplefractions. The tunability of separable particle size could be achieved by changing the viscoelastic propertyof the medium and flow rate.

© 2015 Elsevier B.V. All rights reserved.

. Introduction

Cell separation is an essential pre-processing step for down-tream application such as biochemical analysis, disease diagnosisnd therapeutics [1–6]. Owing to recent advancements inicrofluidic technology, microfluidic-based particle/cell separa-

ion techniques have attracted much attention for their prominentdvantages in lowering fabrication cost and sample/reagent vol-me needed [7–9]. Microfluidic separation techniques can beategorized into two operating modes in general-batch mode andontinuous flow mode. In batch mode, a prolonged operation timend complicated fluidic control are required [10,11]. These draw-acks can be circumvented in the continuous flow mode separation.

n particular, passive separation with no external force has gar-ered great interest owing to its relative ease of fabrication andse [12–14]. Particle/cell manipulation in a viscoelastic fluid can

∗ Corresponding author at: Department of Biomedical Engineering, National Uni-ersity of Singapore, 9 Engineering Drive 1, 117575, Singapore. Tel.: +65 6516 6713;ax: +65 6872 3069.

E-mail address: bieks@nus.edu.sg (S. Kim).1 These authors contributed equally to this work.

ttp://dx.doi.org/10.1016/j.chroma.2015.06.029021-9673/© 2015 Elsevier B.V. All rights reserved.

provide more flexible manipulation than other passive methodsthat rely on flow rate variation in channels with fixed dimensions[15–27]. In viscoelastic flows, the heterogeneous distribution of thefirst normal stress difference (N1) enhances the lateral migrationof suspended particles in a microfluidic channel [15]. Based on thisprinciple, previous studies have attempted to achieve 2-D [16] or3-D particle focusing [17–21], ordering [22], and separation usinga viscoelastic fluid [23–27].

Despite the various attempts for viscoelastic particle manip-ulation, only few studies have been conducted for particleseparation. Furthermore, achieving high efficiency in separationremains a challenge. In a previous study by Yang et al. [24], theydemonstrated sheathless viscoelastic 3-D focusing and continu-ous size-dependent separation of particles by using the synergisticeffect of elasticity and inertia in a square microchannel. How-ever, a narrow range of flow rates examined in their studylimit the optimal synergistic effect of elasticity and inertia. Inaddition, their approach may be limited by a low separationefficiency caused by the random initial distribution of particles

[23–25]. Therefore, it would be difficult to apply these previousapproaches to separation of multiple particles with various sizes,which is often required for clinical applications such as blood cellseparation.

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Recently, other studies have attempted to improve the sepa-ation efficiency using sheath flows for enhanced initialization ofarticle position before separation [28]. However, such an approachequires additional injection channels and controls, making theevice design more complicated. To overcome the previous lim-

tations, we propose a novel device for high-efficiency particleeparation using a viscoelastic non-Newtonian fluid. To the bestf our knowledge, no microfluidic device has been reported forighly efficient, sheathless particle initialization and continuous

abel-free separation in viscoelastic flow. Additionally, we exam-ned the potential of separation performance in our device withlood cells as well as polystyrene beads.

. Material and methods

.1. Dimensionless number

The migration of particles suspended in a viscoelastic fluid istrongly dependent on flow conditions. Thus it is necessary to

ig. 1. (a) Schematic illustration of the two-stage device for sheathless particle focusingocusing in the 1st stage; (C) bifurcation region for the initialization; (D) particle separa

icrochannel.

1406 (2015) 244–250 245

utilize non-dimensional numbers to compare the effect of differ-ent forces on the particle migration. The Reynolds number (Re)describes the ratio of the inertial force to the viscous force whilethe elasticity of a non-Newtonian fluid can be characterized usingthe Weissenberg number (Wi = � �̇c , �: relaxation time, �̇c: charac-teristic shear rate). Since these two non-dimensional parametersrepresent the inertia and elasticity of flow, the relative impor-tance of elasticity to inertia in flow can be estimated by usingthe ratio between the two parameters. Thus, the elasticity num-ber (El = Wi/Re) can be defined as the ratio of the fluid elasticity toinertia.

2.2. Design principle

A schematic of the proposed device for sheathless particle focus-

ing and separation using a viscoelastic fluid is depicted in Fig. 1.The device can be divided into two parts—the initialization part(1st stage) and separation part (2nd stage). The 1st stage includesa circular channel that was fabricated by rounding the corners of a

and separation using a viscoelastic fluid. (A) Entrance region; (B) viscoelastic 3-Dtion in the 2nd stage. (b) Photographic images of experimental setup and PDMS

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46 J. Nam et al. / J. Chroma

quare channel (Fig. 1(a), lower left inset). At the entrance regionf the 1st stage, all particles are randomly introduced without anyheath flows (Fig. 1(a) (A)). During the 1st stage, particles of differ-nt sizes migrate toward the equilibrium position at the center ofhe circular microchannel induced by non-uniform N1 [24]. There-ore, 3-D focusing of particles can be achieved along the centerlineFig. 1(a) (B)) by utilizing a viscoelastic fluid and by eliminating theorner effect (particles being trapped in the corners of a rectan-ular channel). At the end of the 1st stage, there is a bifurcatingoint which is connected to the 2nd stage. Thus, the location of allarticles can be initialized at the inner wall (Fig. 1(a) (C)). The 2ndtage is designed to have a symmetric bifurcation, followed by audden expansion region that can maximize the difference in lat-ral displacements of the particles. In the viscoelastic flow, particlesndergo different extents of the lateral displacement depending onheir sizes since the elastic force acting on the particles is propor-ional to the particle size (Fig. 1(a) (D)) [16].

For the device fabrication, the required channel length for eachtage was determined by employing an empirical design ruleeported previously and can be simplified as follows [18]:

A ∝ R4

ˇ2(1)

here ̌ (=a/R) is the blockage ratio, a is the particle radius, and R ishe radius of the channel cross-section. The design recipe can onlyonsider the device dimensions since the flow rate and the elasticroperty of the suspending medium can easily be varied during thexperiment. Accordingly, the required length is proportional to thehannel radius and inversely related to the blockage ratio. Hence,arger particles require a shorter length to be focused. In the presenttudy, the required length (LA) for particle initialization in the 1sttage was determined using the smaller particle (5-�m diameter)hile the length of the 2nd stage was selected such that only largerarticles (10-�m) can be focused along the centerline, which couldaximize the difference in the lateral displacement between 5- and

0-�m particles. The focusing length of the 2nd stage could also bestimated although a rectangular channel was used in the 2nd stagehannel. At the bifurcating point of the 2nd stage, all the particlesould be focused and placed at the mid-height on the inner wall.

.3. Device fabrication

The standard soft-lithography technique was adopted to fab-icate the polydimethylsiloxane (PDMS) microfluidic device withne inlet and two outlets as shown in Fig. 1(b). Since it has onlyne inlet, it is capable of easy parallelized processing by havingeveral layers of microchannels stacked together [29]. The ini-ial cross-sectional dimensions of the 1st and 2nd stages were5 �m × 45 �m (width × height) and 40 �m × 45 �m, respectively.he channel length in the 1st and 2nd stages was 3 and 1 cm,espectively. To avoid the trapping of particles at the corners underiscoelastic flow in a rectangular channel, the cross-section of thehannel in the 1st stage was changed to a circular shape by utiliz-ng the liquid/bubble injection method [30,31]. Briefly, the channel

odification was accomplished by coating its inner surface withncured liquid PDMS. The use of PDMS has become popular in var-

ous microfluidic fields due mainly to its high reproducibility using silicon master and good suitability for analytical and biologicaltudies [32–34]. The square microchannel was first filled with liq-id PDMS from the inlet of the microchannel. The liquid PDMS inhe microchannel was then removed by suction using a vacuumump at about −60 kPa. Several circular channel devices could be

abricated simultaneously by connecting multiple devices to theacuum pump through T-shaped valves (30600-25, Cole-Parmer,SA), which forms a manifold connection. Thus, the overall fabrica-

ion time per device could be greatly shortened. After removing the

1406 (2015) 244–250

liquid PDMS in the channel, the devices were baked and cured on ahot plate at 100 ◦C. To fully cure the PDMS coating, the devices wereleft on the hot plate at least for 15 min. We found that this methodwas highly reliable and the standard deviation of the fabricatedcircular channel diameter was less than 1 �m.

The effect of circular-shaped channel on particle separation wasvalidated on the basis of the particle distribution at the bifurca-tion in circular and square microchannels (Supporting informationFig. 1). With the multi-coating process, the blockage ratio (ˇ) canalso be controlled easily to reduce the required channel length orto focus even smaller particles in the 1st stage. In this study, afterthe coating process, the mean diameter of the circular channel inthe 1st stage was ∼43 ± 1 �m (mean ± SD), which was determinedat four different locations along its length.

2.4. Sample preparation

A viscoelastic solution (8 wt% of Polyvinylpyrrolidone (PVP))was used as a suspending medium. The addition of PVP has beenwidely used for various rheological measurements of red blood cells(RBCs) and does not significantly affect the physical characteristicsof RBCs [22,35,36]. The rheological properties of the viscoelasticfluid were measured with a rotational rheometer (AR2000-ex, TAInstrument, USA). The viscosity and relaxation time of the fluidwere ∼0.14 Pa s at shear rates 1 < � < 103 s−1 and 0.0013 s respec-tively, which were in good agreement with the values reported inprevious studies [21]. To verify the effect of the viscoelastic fluid onparticle migration, particles were first suspended in a Newtonianfluid (88% glycerol aqueous solution) that has the same viscosity(∼0.14 Pa s) as that of 8 wt% PVP aqueous solution. Then, the parti-cle focusing results by the viscoelastic and Newtonian fluids werecompared (Supporting information Fig. 2).

Polystyrene particles with two different diameters (5- and 10-�m) were suspended in the viscoelastic fluid at a final particleconcentration of 0.1% (v/v). The particle sample was sonicatedfor 30 min before suspension to prevent particle aggregation. Todemonstrate the potential of our device for biomedical applica-tions, finger prick blood samples were taken and transferred intoan anti-coagulant (ethylenediamine tetraacetic acid, EDTA) coatedmicrocentrifuge tube. Blood cells were separated from the bloodsamples by centrifugation and re-suspended in the prepared PVPsolution at a fixed hematocrit of 0.1%. The sample was gently mixedwith a vortex stirrer for ∼30 s prior to the experiment. Polystyreneparticles (15-�m in diameter) were then added to the dilutedblood-PVP solution at a final particle concentration of 0.1% (v/v) tosimulate a disease-related rare cell that is larger than blood cells.During the blood cell separation experiment, a blue filter (B-390,HOYA, Japan) with peak transmittance at 394 ± 4 nm and spectralbandpass of 310–510 nm was used to enhance the contrast betweenRBCs and others.

2.5. Experimental procedure and data analysis

The viscoelastic solution was injected using a syringe pump(KDS210, KD Scientific, Holliston, MA) with a 500-�L glasssyringe (1005TTL SYR, Hamilton, Reno, NV) as shown in Fig. 1(b).Non-dimensional numbers, Re and Wi, were defined in the fully-developed flow region. Lateral migration of particles inducedby the fluid viscoelasticity in 8 wt% PVP solution was investi-gated for a range of flow rates (0.001–0.5 �L/min). Under thepresent experimental condition, Re was approximately in a range of3.57 × 10−6 to 1.78 × 10−3 while Wi was in 0.002–0.93. Thus, elas-

ticity number (El = Wi/Re) was much greater than 1, confirming theelasticity-dominant flow in our experimental conditions. Particlemigration in the channel was monitored using an inverted micro-scope (IX-71, Olympus) and recorded using a high-speed camera

J. Nam et al. / J. Chromatogr. A 1406 (2015) 244–250 247

Fig. 2. Probability distribution functions for off-center distances of (a) 5-�m and (b) 10-�m particles at different flow rates. From the movie clips obtained, a total of 1000p arious Q, where empty (�) and solid (�) down-pointing triangles indicate 5- and 10-�mp

(pm

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3

bflptftwtapc

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Fig. 3. Lateral distribution of particles in the expansion region of the 2nd stage. (a)Normalized lateral migration in the 2nd stage at different flow rates. The dashed

articles were analyzed. (c) Typical microscopic images of particles during flow at varticles, respectively (Supporting information, Video 1).

FASTCOM-1024PCI, Photron, USA). An image processing softwareackage (MATLAB, Mathworks, Natick, MA) was used for auto-ated particle counting and evaluation of separation efficiency.

. Results and discussion

.1. Polystyrene particles

To verify the particle initialization in the 1st stage, the distri-ution of 5- and 10-�m particles was observed by varying theow rate (Q) ranging from 0.001 to 0.5 �L/min. Fig. 2 shows therobability distribution functions for the off-center distance ofhe two different sized particles at the bifurcation region at dif-erent Q (Fig. 1(c)). As Q increased, all the particles showed aendency to migrate towards the flow center. As shown in Fig. 2(c),hen Q < 0.01 �L/min, 10-�m particles were slightly deviated from

he walls, whereas 5-�m particles remained randomly dispersedcross the channel. When Q ≥ 0.01 �L/min, 3-D focusing of 10-�marticles was achieved. All the particles were focused along theenterline of the channel in the 1st stage when Q > 0.05 �L/min.

It is of note that particles may stagnate at the apex of bifurca-ion. Thus, the time delay at the apex may cause collisions betweenront-loaded and trailing particles in a concentrated suspension.he particle collision at the stagnation point might potentially leado the disruption of the particle initialization. However, we didot observe any discernible particle collisions at the given particleoncentration (0.1%).

Once the initialized particles were introduced to the 2nd stage,hey started to migrate away from the wall of the channel. Thus,

articles could be separated in the 2nd stage by the principle ofhe size-dependent particle migration. The difference in the lat-ral displacement for particles with different sizes became moreronounced by the sudden expansion. Fig. 3(a) shows the lateral

black line at y/Y = 0.8 indicates the separation boundary. N = 5 for each Q. (b) Lat-eral position of particles at Q = 0.05 �L/min. Solid-white and empty-white trianglesindicate 5- and 10-�m particle, respectively (Supporting information Video 2).

248 J. Nam et al. / J. Chromatogr. A 1406 (2015) 244–250

Fig. 4. Microscope images of separation of particles including red blood cells (RBCs), platelets, and 15-�m polystyrene particles at 0.1 �L/min (a) at inlet, (b) at bifurcation,and (c) in expansion outlet region. Solid white, solid yellow, and empty yellow triangles indicate 15-�m particle, RBCs, and platelets, respectively. The dashed white lined nsionl

dmmldmAtaptttfrFmtapd

escribes the separation boundary. (d) Lateral position of cells/particles in the expaine at 120 �m indicates the separation boundary.

istribution of 5- and 10-�m particles at different Q. The lateraligration distance from the inner wall (y in Fig. 3(a)) was nor-alized by the half channel width (Y) in the expansion region. At

ow flow rates (Q < 0.01 �L/min), most particles appeared to be ran-omly distributed over the channel width, which would be dueainly to the poor initialization in the 1st stage as shown in Fig. 2(c).s Q increased, 10-�m particles showed an apparent confinement

o a narrow region near the flow centerline. However, there was still considerable amount of 5-�m particles flowing near the 10-�marticle region, which is depicted with the large standard devia-ion at Q = 0.025 �L/min. When Q was increased to 0.05 �L/min,he two different particles were distinctly separated by size sincehe elasto-migration of 10-�m particles toward the flow center isaster than that of 5-�m particles. As depicted in Fig. 3(a), the flowate range favorable for the separation was 0.05 ≤ Q < 0.14 �L/min.ig. 3(b) shows the lateral displacements of particles in the opti-al flow condition (Q = 0.05 �L/min) for the separation, in which

he maximum difference in lateral migration displacement of 5-nd 10-�m particles could be achieved. The lateral displacement ofarticles passing through the expansion region was quantified asescribed in a previous study [37]. Briefly, the channel width of the

region of the 2nd stage (n = 5) (Supporting information Video 3). The dashed black

expansion region (300 �m) was divided into twenty virtual seg-ments. The number of particles found in each segment was thennormalized by the total number of particles flowing the entireregion. In contrast, as Q increased further (Q ≥ 0.14 �L/min), parti-cle separation was unachievable since most particles seemed to befocused near the center, which limits the optimal flow rate range ofthe device, leading to a constraint on the device throughput. How-ever, it is of note that this limit can be modulated by varying thepolymer concentration of the viscoelastic fluid.

To evaluate the performance of our device, the separationboundary was determined and the separation efficiency was calcu-lated [38]. The separation efficiency (�) was defined as the ratio ofthe number of cells migrated over the separation boundary (ntarget)to the total number of cells at the outlet (ntotal).

� = ntarget

ntotal× 100 (2)

When 0.05 ≤ Q < 0.14 �L/min, most 10-�m particles migratedlaterally >120 �m, whereas 5-�m particles < 120 �m. Based onthese results, the particles with different sizes could be separatedwith the 120-�m line as a separation boundary, which was shown

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n Fig. 3(a). Our separation efficiency for the 5- and 10-�m parti-les was ∼99.9% at 0.05 ≤ Q < 0.14 �L/min based on 1000 countedarticles.

.2. Blood cells

By varying Q from 0.001 to 1.0 �L/min, the capability of ourevice for the separation of blood cells into multiple fractions was

llustrated to demonstrate its potential for flow cytometry. The flowate of 0.1 �L/min was used to initialize (1st stage) and separate theells depending on the different lateral migration displacements2nd stage). The shear rate at the 1st stage of the channel wasalculated as 213.5 s−1 which does not exceed the critical value2200 s−1) for shear-induced activation of platelets [39–41].

As shown in Fig. 4(a), all the particles, including white bloodells (WBCs), red blood cells (RBCs), platelets, and 15-�m parti-les were randomly distributed at the inlet, and then they wereocused along the centerline. The initial location of all the particlesefore the separation process was homogenized at the centerlineithin 3 cm of travel distance in the 1st stage (Fig. 4(b)). As shown

n Figs. 4(c) and (d), the lateral migration of particles was quanti-atively analyzed. For a quantitative analysis, each particle (bloodells and 15-�m polystyrene particles) flowing in the microchannelas counted from recorded videos. A total of 1000 particles were

ounted as 15-�m particles and RBCs while 50 as platelets.All the particles of different sizes could successfully be sepa-

ated in our device. The 15-�m polystyrene particles that simulated diseased cell in this study apparently accomplished the migra-ion towards the center (∼150 �m), whereas the majority oflatelets (2–4 �m in diameter) showed small lateral displacements<50 �m) from the sidewall. As expected, RBCs (6–8 �m in diame-er) were observed in the region of 80 ± 15 �m from the sidewalletween the 15-�m particles and platelets (Fig. 4(c)). In contrast,

ocation of WBCs (7–20 �m in diameter) overlapped those of the5-�m particles and RBCs due likely to its wide range of size.owever, since the fraction of WBCs (<0.1%) is in general much

maller than that of RBCs (∼93%) in the total population of bloodells, it was infeasible to conduct a quantitative analysis for WBCs.owever, we observed size-dependent migrations of WBCs in ourevice. Nonetheless, our results confirmed the excellent versatilityf our device in separating platelets, red blood cells and larger cellsimultaneously.

. Conclusion

We proposed a novel two-stage microfluidic device for sheath-ess particle focusing and continuous separation using a viscoelasticuid. The circular channel in the 1st stage could be fabricated by aimple uncured PDMS coating process, which eliminates the cor-er effect found in a rectangular channel and thus enhances thefficiency of particle focusing, resulting in enhanced separation.olystyrene particles of different sizes (5- and 10-�m) were 3-D ini-ialized at the bifurcation and separated by size with an efficiency of99.9%. Also, we demonstrated the potential of the device in sepa-

ating blood cells (including 15-�m particles) from a diluted bloodample into multiple fractions while considering shear-sensitiveells for minimal shear-induced activation. The tunability of sep-rable particle size would be attained simply by changing theiscoelastic property of the suspending medium and flow rate.dditionally, the throughput of our device can be enhanced bysing a viscoelastic fluid with low viscosity but high elasticity, such

s DNA [19] or hyaluronic acid (HA) solutions [20]. Therefore, theevice could potentially be further developed as a clinical tool forigh-throughput separation of targeted rare cells such as circulat-

ng tumor cells [42] and/or bacteria [43] from a biological sample.

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1406 (2015) 244–250 249

In addition, it is of note that a post-analysis platform can easily beintegrated for the analysis of separated samples.

Acknowledgments

This work was supported by National Medical Research Council(NMRC)/Cooperative Basic Research Grant (CBRG)/0019/2012.

Appendix A. Supplementary data

Supplementary data associated with this article can be found, inthe online version, at http://dx.doi.org/10.1016/j.chroma.2015.06.029

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