Microfluidic Hydrodynamic Cell Separation: A Review

Micro and Nanosystems, 2009, 1, ????-????

1

Microfluidic Hydrodynamic Cell Separation: A Review Zhigang Wu* and Klas Hjort Microsystem Technology, Department of Engineering Sciences, Uppsala University, Box 534, The Angstrom Laboratory, 751 21 Uppsala, Sweden Abstract: Microfluidic continuous cell separation based on hydrodynamic interaction in a microfluidic channel has attracted attention because of its robustness, high throughput and cell viability. This paper systematically gives an overview on recent advances in hydrodynamic particle and cell separation in microfluidic devices. It presents the basic ideas and fluid mechanics for the hydrodynamic interaction of a particle in a microfluidic system. Secondly, different kinds of devices are introduced with detailed descriptions of their mechanisms, designs and performances. Finally, the review addresses some practical issues of microfluidic sorting devices for use in biological or medical studies.

INTRODUCTION In the late 20th century, with the development of micro engineering, a new interdisciplinary field – microfluidics – appeared [1-2]. Today the largest use of microfluidics is in ink jet printheads but in future it may revolutionize the present experimental techniques in the biological, chemical and clinical analyses, and even the daily lives of people – just as microelectronics has done for the last sixty years. During the last twenty years, microfluidics has witnessed rapid advancements in different kinds of areas [3-4]. It has shown high potential in clinical applications, such as diagnostics [5] and portable point-of-care applications [6]. Examples of high potential impact are microfluidic platforms for analyses of cells, with their many typical processes for cell studies, including cell culture, treatment, selection, lysis, separation and analysis [7]. Cell sorting and selection is one of the most important processes for these platforms [8]. Various approaches have been investigated in the past decade. Depending on the developing strategy, they could be categorized into two groups: miniaturization of the traditional capillary based fluorescent activated cell sorting (FACS) [9-12]; and utilization of unique microfluidic geometries to interact with cells in the suspension, e. g. [13-14]. Although, the chip version of FACS has been an intensively studied device from academia and industry, today its performance cannot compete with the traditional capillary instrument. Most new cell sorting assays are demonstrated on traditional instruments rather than on microfluidic devices. Thus, the researchers’ eyes have been changed to exploit the unique merits from microfluidics. The cell sorting techniques can also be categorized in two other groups according to their working mode: discrete or continuous cell separation. For the microfluidic community, continuous cell separation is very interesting because of its potential of high throughput and that it is easy to add to other upstream and downstream applications, integrating it into more complex microfluidic systems and achieving true micro total analysis systems (µTAS) [15]. Here, hydrodynamic based continuous separation has recently attracted large interest. The reason is that it doesn’t introduce any external

*Address correspondence to this author at the Microsystem Technology, Department of Engineering Sciences, Uppsala University, Box 534, The Angstrom Laboratory, 751 21 Uppsala, Sweden; E-mail: [email protected] 1876-4029/09 $55.00+.00

field or force for the separation. Only fluid flow is used, which increases the cell viability. Secondly, it potentially has high throughput – especially when it is based on inertial effects. Finally, the design is often more simple than for other techniques, and hence it is easy to integrate and more robust. In this paper, we intend to give a systematic review of the microfluidic hydrodynamic cell separation. Comparing to other microfluidic separation techniques [15-16], e.g. microFACS [11], it has not been intensely reviewed or highlighted. We start by introducing its basic laws and principles. A presentation of various design examples according to their interaction mechanisms for separation is followed. Their typical designs and working parameters are summarized. Furthermore, some other merits of the development of hydrodynamic cell separation are discussed. BASIC LAWS, PRINCIPLES AND DIMENSIONLESS NUMBERS In a microfluidic flow system with cells, or other small particles, various physical phenomena occur and compete against each other in the channel. Using dimensionless numbers, we can observe some basic principles for particle motion inside the microfluidic devices. Here, for the convenience of the reader we introduce some basic laws and the dimensionless numbers for the flow and particles. For a further study, the readers can refer to the review by Squires and Quake [4] or the recent text book on theoretical microfluidics by Bruus [17]. Similar to the solid mechanics, when mass conservation is satisfied a Newtonian fluid obeys a continuous version of his second law, F=ma. That leads to the Navier-Stokes equation,

r % "v r r( r ur ! f ' + v # $ v * = +$P + µ$ 2 v + F & "t )

(

)

(1)

where ρf is the fluid density, v is the velocity, t is the time, P is the pressure, µ is the fluid viscosity and F is the body force. From the equation, only the inertial term is non-linear. When this term disappears or can be neglected the remaining expression becomes the Stokes equation:

r r ur "v !f = #$P + µ$ 2 v + F "t

© 2009 Bentham Science Publishers Ltd.

(2)

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Micro and Nanosystems, 2009 Vol. 1, No. 3

Wu and Hjort

Especially, when the body force is negligible and the flow is in stable state, which is often found in microfluidic devices, it becomes:

r !P = µ! v 2

(3)

which means that the flow is reversible and can be predicted precisely. In all three cases, mass conservation should be satisfied, i.e., Normally, the Reynolds number (Re) is used to describe the relation between inertial force and viscous force, to judge which one of the two that is dominant. It is defined by

Re =

! f VDh µ

=

VDh "

(4)

where V is the magnitude of the velocity, Dh is the hydraulic diameter of the channel and υ is the kinetic viscosity of the fluid. As Reynolds number increase, the inertia force becomes apparent. Especially, when the fluid flow along a curved channel, the wall force the fluid changes the flow direction. The corresponding centrifugal force on the fluid becomes larger, which induces a secondary flow vertical to the major flow. This special centrifuge acceleration is described by a dimensionless number, the so-called Dean number (De). It is

Dh R

De = Re

(5)

with R, the curvature radius of the curved channel. When De is larger enough, the secondary flow can form Dean Circulation vertical to the major flow inside the channel. When a particle is suspended in the fluid, naturally it is influenced by the inertial and viscous forces from the fluid. The particle Reynolds number (Rep) can be defined as:

Re p = Re Table 1.

Flow

Particle

a 2 Va 2 = Dh2 Dh!

(6)

where Re is the flow Reynolds number, and a is the diameter of the particle. For analysis of mass transport in the fluid, the particle behaviour can be observed through its Péclet number (Pep), which describes the ratio between the mass transport due to convection and diffusion.

Pe p =

Va D

(7)

where D is the diffusion coefficient of the particle. For a sphere, it can be estimated by D=kBT/3πµa, with the Boltzmann constant, kB, and absolute temperature, T. Changing the diameter of the particle to the hydraulic diameter of the channel, the Péclet number (Pe) of the flow can be defined similarly as Pe = VDh / D . For a particle in the accelerating flow its Stokes number (St) is used to describe how quickly the particle adjusts to the changes in the surrounding flow, which is defined as the ratio between the particle relaxation time (τr) to the characteristic time (τf) of the flow as below:

St =

2 ! r " p a / 18µ # a 2V = = !f Dh / V 18$ D h

(8)

where β=ρp/ρf and ρp is the particle density. A larger St means that the particle has a stronger preference to continue with its original velocity direction instead of following the fluid trajectory when the fluid is in accelerating motion. For convenience, the dimensionless numbers used in this review are summarized in Table 1. Using the dimensionless numbers above, we can observe some basic principles for particle motion inside the microfluidic devices by sorting them as (a) size determined distribution; (b) size determined filtering; (c) size determined displacement; and (d) size determined guiding, respectively: (a)

Size determined distribution follows from that when Re is small and St of the particle is small the particle

Summary of the Dimensionless Number Used in the Review

VDh !

Re

Reynolds number

Re =

De

Dean number

De = Re

Pe

Péclet number

Pe =

Rep

Reynolds number

Re p =

St

Stokes number

Pep

Péclet number

St =

Dh R

VDh D

Va 2 Dh!

#r "a 2V = # f 18!D h

Pe p =

Va D

inertial/viscous

Eq. (4)

secondary circuiting/major flow

Eq. (5)

convection/diffusion

Eq. (7)

inertial/viscous

Eq. (6)

particle relaxation time/flow characteristic time

Eq. (8)

convection/diffusion

Eq. (7)

Microfluidic Hydrodynamic Cell Separation

will follow the streamline of the flow. Hence, the particle behaviour can be expected as well as that of the flow in the channel, Fig. (1a). At the same time, since the particle is relatively rigid and its size is not ignorable, the centres of particle can only appear in the area where the distance between the particle centre and wall is larger than its radius. For example, the large particle cannot appear in the streamline 1 in Fig. (1a). (b)

Size determined filtering is based on the rigidity of the wall and particle; uses the fact that rigid particles cannot pass a rigid gap smaller than its size, Fig. (1b). This is the basic principle of sieving. Both principles benefit from the predictable Stokes flow in the microfluidic channel (eq. 3).

The last two employ the non-linear behaviour of big particle while the small ones are still behaving as in a linear form Stokes fluid: (c)

(d)

Size determined displacement comes from that when a particle meets an obstacle on its way multiple modes will appear. Here, two fundamental examples are shown in Fig. (1c). When the particle has a small momentum, i.e. its St is small, it will follow its streamline. When it has a large momentum, it will bump away and deflected away without following the streamline where its centre was. Size determined guiding, which is that due to its rigidity a wall can guide the flow with its special geometry or channel shape. Since the particle will follow the streamline of the flow when Re and St is small, the particle motion can be guided with the flow in the channel. For example, for a slanted ridge on the surface which has a similar height of the (small) par-


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ticle, the particle will be guided by the ridge while the large one will pass over the ridge, Fig. (1d). The four principles can be combined and modified further in various conditions. We have to point out that the number of principle is not limited to these four. They are summarised here just for convenience of the discussion for our later examples of design and devices. Hydrodynamic separation is highly dependant on the interaction between the particle and the wall or the fluid. Hence, it is classified by two basic types: particle wall interaction and particle fluid interaction, Fig. (2). The particle wall interaction can be further divided into continuous wall, discrete wall (obstacles) and biomarker on wall, according to the different types of wall. For the discrete wall, the obstacles could be further divided to posts and ridges. The posts are usually used to bump the large particle while ridges are used to guide the small ones. The hydrodynamic separation by particle fluid interaction can be further classified into laminar flow based, inertial flow based and biomimetic flow based, according to the dominant force (influence). Hydrodynamic chromatography is also a separation technique based on particle fluid interaction [18]. However, it works at a discrete mode while this review focuses on continuous methods. Hence, it will not be further discussed. DEVICES BASED ON PARTICLE-WALL INTERACTIONS Field flow fraction (FFF) is a conventional technique used to separate particles or molecules [19-20]. However, most of the FFF devices work in a discrete mode, where a discrete sample is injected at the beginning. Combining flow field flow fraction (FlFFF) and micro technology, Seki’s group proposed a continuous separation technique – pinched

Fig. (1). Basic principles of particle motion in a laminar continuum fluid: (a) size determined distribution; (b) size determined filtering; (c) size determined displacement; and (d) size determined guiding.

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Wu and Hjort

Fig. (2). Classification of hydrodynamic separation.

Fig. (3). Schematics of the principles for continuous wall based hydrodynamic separation: (a) pinched flow fractionation; (b) pneumatic controlled pinched flow fractionation; (c) pinched flow fractionation with asymmetric collectors; (d) electroosmotic flow tuned pinched flow fractionation; and (e) viscoleastic flow tuned pinched flow fractionation.

flow fractionation [21], Fig. (3a). It is based on the rigidity of the wall and particle and that the particle cannot appear in any position in the channel as illustrated in Fig. (1a). The idea was that one fluid stream with different particles (particle flow) and another fluid without particles (elution flow) converge together in a pinched channel. With proper control of the hydrodynamic spreading of the particle and elution flow, the elution flow occupies a broader channel width and the sample flow occupies a narrower channel width. Subsequently, the larger particles are aligned against the channel wall and their centres are immersed in the carrier fluid but the smaller particles are still in the original sample fluid. Via an abrupt channel amplification mechanism, the larger particle will be forced into the elution flow while the small particles are kept in the original sample flow. The particles will be separated and be collected via different outlets. The various collections are also studied by assuming that the flow is laminar and St of particle is small. Under these conditions, the flow behaviour can be studied using equivalent electric circuit model. The flow rate Qp (volume/time) is

linear to the pressure difference Δp between its outlet and its inlet in the channel: (9)

Q p = (!"p) / Rp

where Rp is the hydrodynamic resistance of the channel. For a straight rectangular channel with channel width w, channel width h and a length of L, the resistance Rp is about 12µL /wh3 when channel width is far smaller than channel height. When channel width w, channel width h is closer, the resistance Rp can be obtained by: !1

# n" w &&. 12µ L + h # 192 ) 1 Rp = 1! tanh % % (((0 * % wh3 -, w $ " 5 n=1 n5 $ h ''0/

(10)

Using this formula the flow distribution and the particles trajectories along the channel network can be predicted by analyzing the flow distribution in an electric equivalent circuit based on Kirchhoff’s 2nd law. The use of electric equivalent analysis was demonstrated by Seki’s group in designing asymmetric outlets and predict-



5

Fig. (4). Schematics of principle of size determined displacement particle separation (a) and various modifications: (b) modification of post size and spatial post distribution; (c) horizontal symmetrically modification of spatial post distribution; (d) modification of post shape to rectangle; and (e) modification of post shape to triangle.

ing the pneumatic control of the outlet [22-24]. Later, when introducing electroosmotic flow control, a more precisely tuning of the separation was demonstrated by Wu et al. [25]. Recently, the same group found another tuning method using visocoelastic flow [26]. It is making the tuning easier while it is still keeping the cell at high viability. Miniaturized step-split-flow [27] is also FFF, which is equipped with a mechanically controlled flow splitting component. Still, it is not a purely hydrodynamic separation device and will not be further discussed here. Besides hydrodynamic particle separation, electroosmotic flow pumping also can be used [28]. For the discrete wall methods, the simplest way of hydrodynamic cell separation is by filtering or sieving, Fig. (1b), which is derived from the membrane filtering with micro fabricated discrete posts (wall) or a dam [29-33]. Further, the posts can be arranged as an array to separate different particles as shown as Fig. (4). The principle was derived from earlier work on array cell separation [34] and DNA [35]. As mentioned in the section on basic principle, St describes much of the particle behaviour when they meet an obstacle affront. With its larger St, the large particle will be bumped away while the small particle has just follow the flow around the obstacle, referring to size determined displacement separation, Fig. (1c). If another obstacle appears at the centre of the two particle trajectories down stream, the particles can be further separated and collected [36], Fig. (4a). In an array the particle can be separated at very precise size resolution through iterated particle trajectory control by the spatial distribution of the posts [37-43]. The distribution of the array can be modified according to the different application and requirement, Fig. (4b and c). Furthermore, the cross sectional shape of the posts is not necessarily round. It can be any shape, e.g., rectangular or triangular, Fig. (4d and e). The discrete wall may also have another form – ridges, as shown in Fig. (1d). The ridges will introduce secondary flow

patterns vertical to the major flow along the channel direction. This has been intensively studied for mixing in microfluidic devices [43]. At the same time, the ridges can serve as sliding guides for the particle following the secondary flow. Using this concept, Park’s group demonstrated particle and blood separation with a similar structure as shown in Fig. (5a) [45-47]. Toner’s and Gao’s groups demonstrate another ridge design, Fig. (5b) [48-49]. However, when comparing to the designs of ridges investigated for mixing in micro scale, Fig. (5a-d), designs like those in Fig. (5c and d) have not been demonstrated for hydrodynamic separation [50]. A specific biomarker also can be applied on the surface wall, on both continuous and discrete walls. The binding between the biomarker and the cell will immobilize the target cells or impede their movement to obtain good separation [51-53]. Table 2 summarizes the presented devices based on particle-flow interaction. DEVICES BASED INTERACTIONS

ON

PARTICLE-FLOW

The particle-wall interaction is strong and at high flow rates it may be harmful to sensitive cells. Particle-flow interaction is regarded as a more gentle approach and hence could be advantageous when high throughput is needed. There are three basic mechanisms for particle separation: streamline following, inertial drifting and biometric interaction. Assuming a low Reynolds number and negligible gravity, in a stable flow the particles would follow the streamline. Fig. (6a) illustrates this principle with particles flowing near a channel wall: The flow in the Zone 1 will come to subbranch A, Zone 2 to sub-branch B and Zone 3 to sub-branch C, obeying the fluid distribution in a laminar flow. Considering that the particle is near or close to the wall and according the basic principle of size determined distribution in a laminar flow: if its radius is larger than the largest distance from the external layer of zone 1 to the wall and smaller than that

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Wu and Hjort

Fig. (5). Various designs of ridges: (a) slanted ribs; (b) staggered-herringbone grooves; (c) mixed slanted ribs and staggered-herringbone grooves; and (d) slanted grooves. Table 2.

Summary of the Separation Approaches Based on Particle Wall Interaction

Ref.

methods

sub-approach

sample

separation principle

typical flow rate/velocity

materials

Yamada 2004

continuous wall

pinched fractionation

particle

size

20 µl/h

PDMS

Takagi 2005

continuous wall

asymmetric collection

particle

size

20 µl/h

PDMS

Sai 2006

continuous wall

pneumatic tuning

particle

size

30 µl/h

PDMS

Wu 2007

continuous wall

EOF tuning

particle/ cell

size

not reported

PDMS/ glass

Wu 2008a

continuous wall

viscoelastic tuning

particle/ cell

size

5 µl/h

PDMS/ glass

Zhu 2004

discrete wall

dam-filter

Bactria

size

2.0 mm/s

silicon/ glass

Li 2007

discrete wall

dam-filter

diluted Blood

size

10 µl/min

PMMA

Sethu 2005

discrete wall

post-filter

whole blood

size

5 µl/min

PDMS/ glass

Murthy 2006

discrete wall

post-filter

rat cardiac cell

size

20 µl/min

PDMS/ glass

Ji 2008

discrete wall

dam, post, membrane

whole blood

size

20 µl/min

silicon/ glass

Huang 2004

discrete wall

post array

particle

size

0.4 mm/s

silicon/ glass

Davis 2006

discrete wall

post array

particle/blood

size

1.0 mm/s

silicon/ glass

Inglis 2006

discrete wall

post array

particle

size

1.5 mm/s

PDMS/ glass

Choi 2007

discrete wall

ridges

particle

size

30 mm/s

PDMS/ PDMS

Hsu 2008

discrete wall

ridges

particle

density

0.25 mm/s

PDMS/ PDMS

Nagrath 2007

biomarker

marker on post

tumour cell

marker

1 ml/h

silicon/ glass

Plouffe 2008

biomarker

ligand on surface

cells

ligand

9.3 µl/min

PDMS/ glass

Karnik 2008

biomarker

ligand on surface

HL-60 cell

ligand

0.3 ul/min

PDMS/ glass



7

Fig. (6). The separation methods based on streamline following: (a) principle and two design examples (b) symmetrical design and (c) asymmetrical design.

of zone 2, the particle cannot go into sub-branch A but will flow into sub-branch B; if the particle radius is larger than the largest distance from the external layer of the zone 2 to the wall and smaller than that of zone 2, the particle will flow into sub-branch C; and larger particles near to the wall will stay in the major channel. This concept was used to filter particles with a precisely designed numerous subbranches as shown as in Fig. (6b) [54]. However, this method has an inevitable problem; the small particles cannot be filtered out from the large particles by 100 %. To make a higher purity particle separation, many narrow sub-channels can be adopted to extract particles inside the flow for the larger particles and sending them back to the major channel well before the separation point, to once again aligning the particles against the wall [55], Fig. (6c). This structure can also be used to self-focus the particle with a symmetric design. Refereeing to Fig. (6a), in the biomechanics there is a famous bifurcation law, the so-called Zweifach–Fung effect. In the microcirculation, when erythrocytes flow through a bifurcating region of a capillary blood vessel, they have a tendency to travel into the daughter vessel which has the higher flow rate leaving very few cells flowing into the lower flow rate vessel. We can regard this effect as a typical example of particle behaviour in laminar flow inside a microfluidic device. The microfluidic version of this study can be found in Yang’s work [56] and similar works [57-58]. To keep the particle following the flow, the most critical parameter is to keep a small Re, or more precisely, to keep

the flow static and St small. However, a small Re will limit the throughput. When increasing the flow rate to increase the throughput, inertial force cannot be ignored. Since the centrifuge is widely used to separate different kinds of particles, cells or biological macromolecules, naturally it was among the first technique to be miniaturized in inertial force separation [59-60], Fig. (7a). As expected, the Dean circulation in the channel plays a significant role for separation [61-63]. With an asymmetric design of a curved channel, Toner’s group demonstrated very fast separation [64], Fig. (7b). With the competitive forces of the lift force and Dean drag and by using an oscillating structure, the particles were selectively focused and collected in the centre of the channel. A slightly modified structure was used to separate palettes from human blood [65]. Alternatively, selective focusing can be achieved by using a multiple-orifice channel based on tubular pinch effect [66]. In addition, the shear-induced lift force can also be used for droplet based cell separation, e.g. [67]. However at high flow rates, due to the “hard” inertial interaction, and hence high shear stress near the walls, there is a large risk of stressing or even damaging sensitive cells without sheath flow protection. Therefore, a “soft” handling of cells should be preferred at high flow rates. Wu et al. developed a new approach to separate particles/cells based on soft inertial force induced migration with flow defined curved and focused sample flow inside a microfluidic device [68]. This approach relies on a combination of an asymmetrical sheath flow and proper channel geometry to generate a soft inertial force on the sample fluid in the curved and fo-

Fig. (7). The example designs of geometry induced inertial force separation: (a) Archimedes spiral; (b) asymmetrical curved channel; and (c) multiple-orifice channel.

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Wu and Hjort

Fig. (8). Principle of soft inertial separation with photos: (upper left) fluorescent visualization of the soft interactions among the sample flow, protecting sheath and acting flow and (upper right) particle visualization of the particle separation (bright band for larger particle and darker band for small particle)

cused sample flow segment to deflect larger particles away while the smaller ones are kept on or near the original flow streamline. With an asymmetric focusing structure, a rapid change in momentum when the three flows meet, i.e. St is large enough, a mismatch between the fluid and particles will appear. The subsequent force deflects the particles away from the fluid streamline and finally the particles escape from their original carrier fluid, Fig. (8). Different properties of cells can also be used to separate them from each other. It is well-known that non-polar molecules prefer to non-polar environment. In the same way, cells also have their favourite environments and they tend to migrate to the fluid they prefer to. Similar to extraction, Seki and co-workers demonstrated a continuous plant cell migration from a PEG-enriched solution to a Dextran-enriched solution using two phase extraction [69]. This approached was modified to fractionation of living CHO-K1 cells from the total population [70], Fig. (9). Cell deformability also can be used to separate cells. Since the red blood cells easily are deformed and have a tendency to appear in the centre of the capillary channel. This feature was used to enrich white blood cell in blood [71]. In another work, Whiteside and coworkers observed that E. coli has a specific self rotation direction on an agars surface [72]. Later, this discovery was used to separate E. coli based by length [73]. Table 3 summarizes the presented devices based on particle-flow interaction. DISCUSSIONS The performance of a cell separator is evaluated by throughput, recovery and purity. For the continuous approaches, the most interested parameter has been considered to be the throughput or flowrate or velocity, as presented in the previous tables. It is clearly one of the most important advantages to adopt hydrodynamic cell separation. The tech-

niques based on inertial separation undoubtedly have the highest flow rates or velocities, while the other approaches have shown other merits. For example, the approaches based on particle and post array interaction have separation with the highest size resolution and the biomarker based ones present separation based on biological selectivity apart from their inertia. However, the flowrate or velocity is not equal to the throughput, which is highly dependant on the concentration of the cells. Furthermore, recovery and purity are important factors of the cell sorter’s performance. Unfortunately, we rarely find such data in published works. Hence, it is difficult to give a thorough evaluation of the designs presented.

Fig. (9). Cell separation with two aqueous phase separation. White circles indicate one kind of cells, which prefer to stay in the flow 1 and gray circles indicate another kind of cells which prefer to stay in the flow 2.

The hydrodynamic separation approaches are most often based on size separation. Most hydrodynamic separation methods actually are relying on more general mechanisms, such as inertia, and other features than size contribute. For


Table 3.


9

Summary of the Separation Approaches Based on Particle Fluid Interaction

Ref.

methods

sub-approach

sample

separation principle

typical flow rate / velocity

materials

Yamada 2005

laminar flow

streamline following

particle/blood

size

1.0 µl/min

PDMS/ glass

Yamada 2006

laminar flow


particle

size

0.5 µl/min

PDMS/ glass

Yang 2006

laminar flow


particle/blood

size

18.5 mm/s

PDMS

Jaggi 2007

laminar flow


blood

size

2 ml/min

PMMA

Seo 2007

inertial separation

centrifugal/dean drag

particle

size

92 mm/s

PDMS/ glass

Re=13.8

Di Carlo 2007

inertial separation

inertial focusing

particle

size

0.5 m/s

PDMS

Re=7.5

Di Carlo 2008

inertial separation

inertial focusing

particle/blood

size

0.9 ml/min

PDMS

Rp=1.53

Chabert 2008

inertial separation

shear induced migration

cell

size (droplet)

7 µl/h

PDMS/ glass

Wu 2008b

inertial separation

Dean drag

particle

size

1.4 m/s

PDMS/ glass

Re=31.8

Bhagat 2008

inertial separation

centrifugal/dean drag

particle

size

10 µl/min

PDMS/ glass

Re=5

Park 2009

inertial separation

shear induced migration

particle

size

80 µl/min

PDMS/ glass

Rp=1.94

Wu 2009

inertial separation

soft inertial deflection

particle/cell

size

15 µl/min

PDMS/ glass

Re=62

Hulme 2008

Biomimetic

motile of cells near agar

E. coli

length of E. coli

31 µm/s

Agar/ PDMS

Ratchets combined

Shevkoplyas 2005

Biomimetic

relative rigidity of leukocyte

blood

rigidity

6.6 nl/s

PDMS/ glass

Nam 2005

Biomimetic

Aqueous two-phase extraction

CHO-K1

affinity

0.04 µl/min

PDMS

example, a few works are based on density, rigidity (compressibility), length and so on. However, all these methods differ from the widely used biomarker based separation in biological and clinical applications. With the large variety in size of the same kind of cells due to their environment or where they are in their cell cycle and with so many cells of similar sizes, cell sorting by size will always have limited application although it have demonstrated very high separation resolution or efficiency in specific applications, e.g. [36, 65, 68]. In this perspective, the biomarker interactive hydrodynamic cell sorting may have a good chance to reach the required selectivity and still enable high throughput and simple integration in microfluidic systems. Although it is convenient for the proof of concept to demonstrate artificial particle separation this is not nearly enough to evaluate its use in cell separation. Since the cells are living dynamic systems, they are much more complex than a simple particle. Thus, it is very important to know the size of the cells, their size variations, and other relevant biological variations, to know when it is meaningful to apply a separation technique for biological or medical studies, e.g., when separating different cells in blood. In addition, most of the demonstrated applications were focused on mammalian

observation

cell separation for biological or clinical studies, due to their medical relevance and relatively large size (which eases both the fabrication and separation control). However, there are many interesting applications that need sorting of bacteria at high throughput but still only few works have been presented, e.g. biopsies of infections [74], environmental monitoring [75] and food safety [76]. In the microfluidic community PDMS is undoubtedly the most popular material for fabrication. Most of the presented works are based on PDMS on PDMS or glass since PDMS has a lot of advantages for research e.g. ease of fabrication, rapid process and low cost. Few works are based on the silicon [32, 36] and glass or hard plastics e.g. PMMA [33]. Since most of these materials are hydrophobic, cell adherence becomes an important issue for the device fabrication. Many efforts have been put on surface passivation. The most widely used method is to introduce non-ionic amphiphilic molecules into the buffer for dynamic coating, e.g., Tween 20 [22, 26, 45], Pluronic F108 [42] and Hydroxypropyl cellulose [27]. Also, silane based chemistry has been adapted for PDMS surface modification [71]. All these methods enhance the reliability of the device and benefit the future device or whole system development.

10 Micro and Nanosystems, 2009 Vol. 1, No. 3

Wu and Hjort

CONCLUSIONS

[2]

Hydrodynamic cell separation has been advancing rapidly in many directions in the past few years, especially in the mechanism understanding and device design. It has many advantages when comparing to other microfluidic separation techniques. The robust, simple and continuous format indicates its high possibility for practical applications, production and further for commercialization [3, 77]. Today, most of the works have been performed by researchers from the engineering sector. With more players in the field having a biological or clinical background, the future of microfluidic hydrodynamic cell separation is very promising.

[3]

NOMENCLATURE

[4] [5] [6] [7] [8] [9]

ρf

= fluid density

[10]

v

=

[11]

t

= time

P

= pressure

µ

=

fluid viscosity

F

=

body force

Re

=

Reynolds number

V

= magnitude of the velocity,

[15]

Dh

=

[16]

υ

= kinetic viscosity of the fluid

De

=

Dean number

Rep

=

Particle Reynolds number

a

= diameter of the particle.

Qp

= flow rate (volume/time)

k B,

= Boltzmann constant

T

= absolute temperature

St

=

Stokes number

τr

=

particle relaxation time

τf

=

characteristic time of flow

ρp

= particle density

Pe

=

Péclet number of the flow

Pep

=

Péclet number of the particle

D

= diffusion coefficient of the particle

Rp

= hydrodynamic resistance of the channel

w

=

channel width

h

=

channel height

L

= channel length

R

= radius of curvature

velocity

hydraulic diameter of the channel

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