University of Groningen

New strategies for (biological) particle handling and separation in microfluidic devicesJellema, Laurens-Jan Cornelis

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25

Separating Biological Particlesin Flowing Microsystems

L.C. Jellema and E. VerpoorteIn preparation as invited paper for Electrophoresis

Pharmaceutical Analysis, Department of Pharmacy, University of Groningen,

A. Deusinglaan 1, PO-box 196, 9700 AD Groningen, The Netherlands

Chapter 2

Chapter 2

26

Microfluidic devices have become increasingly multifunctional tools, applicable to

many different areas of research. The well-defined flows within the channels are use-

ful when working with particles, as they follow a predictable path through the de-

vice. This report focuses on the use of laminar flows to separate particles in

microchannels. Specially designed channels are used to generate flow patterns which

facilitate particle separation by allowing the capture and transport of particles in clo-

sed streamlines or, alternatively, enabling particles to easily switch from one stream-

line to another. External forces can also be used to influence susceptible particles and

induce separation. The particles may be the actual analytes that have to be separated

from other particles or the fluids in which they are suspended. On the other hand,

particles may also be used as reagents to interact with the analyte under investigation.

In this review, we present studies describing the separation not only of well-charac-

terized inorganic nano- or microspheres but also biologically relevant particles, in-

cluding (blood) cells, bacteria and DNA. These biological particles deserve special

attention with respect to their behaviour and separation in microfluidic channels, as

they often have shapes deviating from the spherical, are deformable, and have the

ability in many cases to change conformation.

Abstract

Separating Biological Particles in Flowing Microsystems

27

2.1 Introduction

In the fields of chemical, biochemical and clinical analysis, sample preparation ge-

nerally remains a crucial step in the analysis process, as few methods allow the ana-

lysis of analytes directly in complex matrices. Sample preparation often includes the

manipulation of particles in fluids or the removal of particulate matter from fluids.

The particles may be the actual analyte, interact with the analyte, or interfere with ana-

lysis of the analyte, in which case they need to be removed from the sample. The

term “particles” is used in the broadest sense of the word here, and includes larger (bi-

ological) particles, like polymer microspheres, cells, bacteria or DNA. In all cases, the

method chosen for particle handling and separation strongly depends on the compo-

sition of the sample and the final goal of the analysis. Microfluidic channels can be

used to integrate analytical procedures like sampling, reactions, separations and de-

tection onto one automated device. Due to the micro- or nanometer-scale dimensions

of the channels, small aliquots of fluid can precisely be handled and manipulated in

short times. The smallest dimensions of the microchannel also determine the size of

the analyte and other species that can be manipulated. The first on-chip separation was

performed by Manz et al., namely a capillary electrophoresis separation of small fluo-

rescent molecules.1, 2 Shortly after these publications, Wilding and Kricka examined

sperm motility through microfabricated barriers in a microchannel, in one of the first

examples of cell analysis on a chip.3 Furthermore, they demonstrated that red blood

cells and microparticles could be separated from samples using filters with 5-µm-

wide gaps. However, whole blood was not processed because of clogging problems.4

These first publications showed the potential of lab-on-a-chip technologies for a wide

range of applications, from analysis of molecular species to cell analysis. Since then,

many other different separation mechanisms in microchannels for molecular and par-

ticulate species have been studied, many of which can be and have been used for ana-

lysis of biological samples.

The geometry, fluid flowrate, and fluid properties all dictate whether flow in a given

flow system is laminar, turbulent or transitional between these flow types. The type

of flow, on both the macro- and micro-scale, can be characterized with a dimension-

less parameter, the Reynolds number, Re.5 The Re represents the balance between

inertial and viscous forces on the fluid.5, 6 The inertial forces are represented by the

average velocity, v, and the density, ρ, of the fluid. The viscous forces are represen-

Chapter 2

28

ted by the dynamic viscosity, η, and the characteristic length, d, where d reflects the

effective diameter of the channel.5, 6

In microchannels, the Reynolds numbers are small and thus flows are well-defined

and predictable. When an analyte is present in the flow, it will follow the streamline

in which it finds itself. However, the analyte will also diffuse while being transpor-

ted by the flow, either in the direction of the flow or perpendicularly to it. Diffusion

is a time-dependent process,7 and diffusion rate decreases as analyte size increases.8,

9 Generally speaking, diffusion is a relatively slow process when compared to con-

vective transport of the analyte in the channel, especially where micrometer-sized

particles are concerned. Particles will thus tend to follow the streamlines in which

they find themselves, with little diffusion occurring between adjacent streamlines.

Hence, predictable flows in microchannels provide an extremely interesting route to

controlled particle transport and manipulation in microfluidic devices. Different me-

chanisms for particle separation and/or removal from samples in microfluidic devi-

ces have been reported in recent years, based on well-defined laminar flow profiles.

Separation of particle samples can be performed in two ways, namely as batch pro-

cesses or as continuous processes. In batch processes, a small aliquot of sample (nL

or less) is injected into a main channel, where separation is then performed (e.g. chro-

matographic or electrophoretic separation) and the different particle fractions detec-

ted. This type of separation is primarily analytical in nature and is not really suited to

procedures requiring the removal of particles from larger sample volumes. Though

this process could be repeated numerous times to process larger sample volumes and

collect larger particle amounts for preparative purposes, it would prove to be quite

cumbersome and time-consuming. Continuous separation methods provide an alter-

native when larger numbers of particles in larger volumes need to be processed. In this

case, the sample is continuously introduced into the separation channel, to be sorted

into different streams and subsequently leave the channel at different positions after

separation is performed. This can be achieved in liquid flows in branched channel

networks or in flows around obstacle arrays in microchannels. Other approaches in-

volve the use of externally applied forces, like acoustic or magnetic forces, to influ-

ence the behaviour of particles susceptible to these forces in the flow. Continuous

separation mechanisms may often also be executed with a small aliquot of sample,

Re = ���

� (1)

thus serving as batch separation procedures. Particle separation mechanisms in con-

tinuously flowing systems have been reviewed thoroughly by Pamme.10 A look at the

relevant literature will reveal that the number of applications for separation of biolo-

gical particles, based on various separation mechanisms, has recently grown sub-

stantially.

In this paper, different particle separation mechanisms inspired by the laminar flow

in micrometer-sized channels are reviewed. The different techniques are categorized

into three different sections. The first section includes all those mechanisms in which

particles are separated in laminar flows through channel networks, and around ob-

stacles placed in the channels. Separation in channels designed to induce secondary

flow effects are discussed in the second section. The third and last section focuses on

the use of external forces to separate particles susceptible to these forces, based on dif-

ferences in their susceptibility. Examples are given for the application of each sepa-

ration mechanism to biological particles, to highlight the usefulness of many of these

approaches for real-world samples. The Tables in the text give overviews of the dif-

ferent techniques and focus on applications utilizing well-defined microspheres, bi-

ological particles like cells, bacteria, and blood.

2.2 Separation by laminar flow through channels

A number of microfluidic separation techniques benefit from the exquisite control of

laminar solution flows possible only at the nano- and microscale. Several of these

are based on the use of pressure-driven flow velocity profiles. Channel networks are

designed to generate specific routes for the flow through the device. This section des-

cribes different techniques for separation based on laminar flow. Table 1 at the end

of the section summarizes the different separation mechanisms and their application

to the separation of biological particles.

2.2.1 Hydrodynamic Chromatography

The use of microfluidic devices has led to new particle separation approaches based

both on totally new laminar-flow enabled mechanisms and on the miniaturization of

existing methods. Hydrodynamic chromatography (HDC) is an example of an exis-

ting method first described by Dimarzio & Guttman11-13 and Small14 in the early

1970’s. HDC is reminiscent of Taylor-Aris dispersion,15-17 a phenomenon by which a

Separating Biological Particles in Flowing Microsystems

29

soluble compound is dispersed in a pressure-driven laminar flow (PF) due to the com-

bination of velocity variation across the flow profile and molecular diffusion between

streamlines. However, where Taylor-Aris dispersion addresses the problem of point

particles, HDC involves the separation by flow of particles of finite size.11 The par-

ticles are separated due to the fact that particles in a fluid flow follow the streamli-

nes in which they are present. Because the centre of a particle cannot get any closer

to the walls than its radius,11 larger particles occupy a slightly smaller region of the

channel, as shown in Fig. 1. Therefore, larger particles cannot sample the low-velo-

city streamlines near the wall, and thus gain a higher average velocity than small par-

ticles.14 Particles can be separated based on this phenomenon if the separation channel

is long enough, as larger particles will reach the end of the channel before smaller par-

ticles do. A schematic representation of the mechanism is given in Fig. 1.

HDC was performed initially in packed columns in order to obtain the narrow flui-

dic conduits required for size separation of polymers11 and colloidal particles ranging

in size from tens of nm to a few μm.14 The introduction of precision-made micro-

channels produced using chip-based technologies and having depths of 1 μm or less

has represented a major advance for this technique. It has become possible, for in-

stance, to perform separations of nm-sized particles and proteins in channels having

lengths of just a few cm.18-20

The first demonstrations of HDC separations on a chip involved the separation of

samples containing fluorescent 110-, 44- and 26-nm particles, injected onto an 80-

mm-long, 1-mm-wide and 1-µm-deep channel.18, 20 The larger 110-nm-diameter par-

ticles, as expected from theory, were carried faster by the flow and eluted first,

Chapter 2

30

Figure 1 Schematic overview of hydrodynamic chromatography separation. Left to right: (A)definition of the area of the channel occupied by the particles with respect to their diameters;(B) sample plug with small and large particles at the beginning of the channel; the parabo-lic flowprofile; (C) small and large particles are separated at the end of the channel; arrowsindicate the average velocities.

A B C

Separating Biological Particles in Flowing Microsystems

31

followed by 44- and 26-nm particles. However, it was observed that the elution of

both small molecules and particles often differed from theoretical predictions. It was

concluded that effects such as analyte adsorption to the wall (in the case of small

fluorescent molecules), ionic strength and differences in particle surface properties

such as charge, can all play a significant additional role in these HDC separations.

Comparison of separations performed in 1.0-mm-wide and 0.5-mm-wide channel re-

vealed that dispersion due to the contribution of the sidewalls was reduced by using

wider channels.20

The application of chip-based HDC to the analysis of proteins has been reported.

Blom et al. demonstrated the separation of two proteins, bovine serum albumin (BSA)

and bovine eye α-crystallin,20 in about 1 minute, rather than the 8 minutes required

for the same separation in a packed column HDC.20, 21 A report by Stein et al., inves-

tigating the mobility of DNA strands in shallow microchannels, also suggests that

this chip technique could be used for DNA separation.22 Linear DNA strands have an

equilibrium coil size which increases with length. In microchannels having a height

larger than 2 µm, the pressure-driven mobility of DNA is observed to increase as a

function of strand length.22 The cause of this behaviour is analogous to that of the

solid nanospheres discussed above, in that the DNA molecules cannot get any closer

to the walls than their radii allow. Therefore, larger DNA molecules are more locali-

zed towards the center of the channel than smaller DNA molecules, exhibiting a hig-

her mobility as a result. In fact, it should be possible to separate 8.8 kbp DNA from

48.5 kbp DNA in a 50-µm-wide, 3-µm-deep channel, as 8.8 kbp DNA exhibits a 10%

lower mobility in this system than 48.5 kbp DNA.22 However, the separation of two

different lengths of DNA would require a complete device with injection element and

separation channel and was not tested in ref. 22.

2.2.2 Pinched flow fractionation

Pinched flow fractionation (PFF), introduced by Yamada and Seki,23 is another mi-

crofluidic separation technique exploiting the laminar flow velocity profile. In PFF,

the sample stream containing particles is continuously introduced into a pinched chan-

nel segment via one inlet, as shown in Fig. 2A, and pinched against the wall with a

second carrier stream, introduced via a second inlet. In this way, the particles are alig-

ned against the wall. The center of a smaller particle ends up closer to the wall than

that of a larger particle, so that particles in the pinched stream are forced to sample

different streamlines depending on their size. When the flow leaves the pinched seg-

ment, the flow streamlines containing the different-sized particles fan out, and the

distance between streamlines is amplified. As a result, the particles are found at dif-

ferent positions across the wide segment perpendicular to the flow, with the smallest

particles closest to the walls and the largest located more towards the center of the

channel, as shown in Fig. 2. Because the streamlines containing different particles

are spread out over the width of the channel after exiting the pinched segment, this

technique is also referred to as hydrodynamic spreading.24

Separations were first performed with 15- and 30-µm-diameter particles using a 50-

µm-wide pinched segment.23 Multiple outlet channels located after the pinched seg-

ment could be used to collect the separated particle types, Fig 2B. The amount of

Chapter 2

32

Figure 2 (A) Schematic overview of pinched flow fractionation. Particles are pinched againstthe wall by a carrier fluid. When the flow enters the wide segment, it fans out and the sepa-ration between streamlines is amplified, with a resulting separation of particles perpendicu-lar to the flow. Microspheres and erythrocytes are colored dark and light grey, respectively.�ote that the erythrocytes are aligned vertically in the flow in the pinched segment. The re-lative flowrate distribution is indicated with the dashed line. (B) Channel design with multi-ple outlet channels to collect particle fractions separately. Outlet numbers correspond to theoutlets mentioned in the text. (C) Schematic overview of separation enhancement using a phy-sical barrier to separate the flow, followed by a second fan-out step.

A B

1

5

2

4

3

C

pinched segment

enhancing segment

fluid entering each outlet is determined by the relative flowrate distribution between

the outlets. Changing the flowrate distribution was used as a means to tune the sepa-

ration and collect 0.5 to 5.0-µm-diameter particles in different outlets.25-27 Different

approaches were chosen to change the flowrate distribution. In one case, the dimen-

sions (width, depth and length) of one of the outlet channels were altered so that this

channel acted as drain (e.g. outlet 5 in Fig. 2B could be changed). As a result, the

largest portion of the liquid flowed into this outlet. This approach was named Asym-

metric Pinched Flow Fractionation (AsPFF), as the outlet positioned closest to car-

rier flow acted as the drain and design of the outlet became asymmetric.25 The second

approach used a poly(dimethylsiloxane) (PDMS) microvalve, based on the valving

mechanism described by Unger et al.,28 to increase the resistance of one outlet and

decrease the amount of fluid running into the channel.26 A PDMS membrane placed

on top of the channel (outlet 1 in Fig. 2B) could be actuated with pneumatic pressure

to change channel resistance. A third method used electro-osmotic flows (EOF) in-

stead of PF to generate fluid flows. Electrodes were connected to the inlets (positive)

and outlets (negative), and different potentials were applied to the electrodes in the

outlets to tune the flowrate distribution precisely and thereby enhance separation per-

formance.27 The group of Hjort developed a fourth method, in which the pressure-

driven carrier flow was tuned by adding EOF. However, this was performed for 1.9-

and 9.9-µm-diameter particles in a channel where the outlet consisted of one wide

channel segment, as shown in Fig. 2A. Separated particles were therefore not col-

lected as different fractions.24 The separation in a channel with one wide outlet could

be improved by introducing a so-called ‘enhancing segment’, which physically se-

parates the particle-containing-flow from the non-particle-containing-flow in the wide

channel segment. The enhancing segment itself has a unique geometry, widening after

a short distance. This allows the separation of particle-containing streamlines to be

amplified a second time, improving separation as shown in Fig. 2C.29 These resear-

chers found that separation resolution depends on the microchannel aspect ratio, par-

ticle size difference and wall roughness. Others have noted that particles with

diameters on the order of the wall roughness could not be separated , due to the dis-

persion caused by the wall roughness.30 In contrast to HDC, the separation in PFF is

not dependent on the dispersion of particles between different streamlines in the la-

minar flow. Rather, the particles are pinched by the flow to end up in defined stre-

amlines, making PFF more suitable for larger particles than HDC.

Separating Biological Particles in Flowing Microsystems

33

The first application of PFF for separation of biological particles was done in an

AsPFF-device.25 Erythrocytes (red blood cells) were pinched with a carrier stream

and collected in one of thirteen outlets, using an outlet design similar to that in Fig.

2B. Importantly, erythrocytes are not spherical, having a biconcave discoid geome-

try instead with diameters of 6-8 µm and a thickness of 2 µm. Interestingly, 80% of

the erythrocytes were collected in the outlet where spherical particles of 2.6- to 3.9-

µm-diameter particles normally ended up. The remaining portion was collected in

the outlet where normally even smaller spherical particles were found. It was post-

ulated that the erythrocytes were aligned vertically in the flow towards the wall in the

pinched segment and beyond, as shown in Fig. 2A. Particle behaviour is thus domi-

nated by the minimum length, e.a. the thickness of the erythrocyte.25

The Hjort group enhanced the separation by PFF of E. coli and yeast by tuning the

carrier flow with EOF.24 The carrier and particle flow were operated at the same pres-

sure-driven flowrates, which resulted in an equal distribution of the flows over the

channel width. Particle pinching was then performed by tuning the carrier flow with

EOF. This is in contrast to the work from the group of Seki where different flowra-

tes were used to pinch the particles. In the Hjort group, neural cells were separated

from glial cells by Wu et al..31 Non-Newtonian fluids were used to control the width

of the particle-containing- and carrier-flows. One of the properties of a non-Newto-

nian fluids is that the viscosity changes when shear stress is applied. At low flowrate,

that is, low shear stress, the fluid exhibits a large viscosity. When the flowrate, and

thus the shear stress, increases, the viscosity decreases. At a constant flowrate ratio,

the viscosity will determine the width of the carrier and particle flow, in essence tu-

ning these flows viscoelastically. At high viscosity (low flowrate), the carrier stream

will pinch the particles or cells against the wall and hence separation of particles can

be performed.

The surfaces of particles can be specially coated to capture an analyte of interest or

extract species from solutions. One reported PFF application utilized different-sized

streptavidin-coated particles functionalized with oligonucleotides for the detection

of point-mutations or single-nucleotide polymorphisms (SNPs). After DNA hybridi-

zation, the larger ‘wild-type’ particles were separated with PFF from smaller ‘mu-

tant-type’ particles.32 A last example of PFF is the size separation of oil droplets in

water using PFF.33 This could be very useful, as droplets can be used to encapsulate

Chapter 2

34

biological particles like cells and therefore act as a tool in biological particle hand-

ling.34

Separating Biological Particles in Flowing Microsystems

35

Part

icle

type

M

echa

nism

(B

io-)

Parti

cle

diam

eter

C

omm

ents

C

ontin

uous

/B

atch

R

ef.

• Po

lyst

yren

e na

nopa

rticl

es

Hyd

rody

nam

ic

Chr

omat

ogra

phy

•

110,

44

and

26 n

m

Proo

f of p

rinci

ple

Bat

ch

18, 2

0

• Po

ly(s

tyre

ne/d

ivin

ylbe

nzen

e)

Pinc

hed

flow

frac

tiona

tion

• 15

& 3

0 μm

Pr

oof o

f prin

cipl

e C

ontin

uous

23

•

Poly

mer

•

Poly

styr

ene

Pinc

hed

flow

frac

tiona

tion

• 1.

0 –

5.0

μm

1.

9 –

9.9

μm

C

hang

e of

cha

nnel

resi

stan

ce

EOF

tuni

ng /

visc

oela

stic

ally

tuni

ng

Con

tinuo

us

25, 2

6 24

, 31

• Po

lyst

yren

e Pi

nche

d flo

w fr

actio

natio

n •

0.25

– 2

.5 μ

m

Extra

spre

adin

g ar

ea

Con

tinuo

us

29

• Po

lyst

yren

e Pi

nche

d flo

w fr

actio

natio

n •

0.5

– 3.

0 μm

Fl

ow d

istri

butio

n co

ntro

l by

EOF

Con

tinuo

us

27

Bio

logi

cal p

artic

les

•

Bov

ine

seru

m a

lbum

in a

nd

• B

ovin

e ey

e α-

crys

talli

n H

ydro

dyna

mic

C

hrom

atog

raph

y •

~10

nm

~ N

.A.

Sepa

ratio

n ba

sed

on si

ze

Bat

ch

20

• Li

near

DN

A

Hyd

rody

nam

ic

Chr

omat

ogra

phy

• 48

.5 k

bp (1

.46

μm)

20.3

kbp

(0.9

2 μm

) 8.

8 kb

p (0

.58

μm)

Sepa

ratio

n ba

sed

on le

ngth

/coi

l dia

met

er

Bat

ch

22

• Er

ythr

ocyt

es (d

isco

id)

Pinc

hed

flow

frac

tiona

tion

• 2

x 6

– 8

μm

Sepa

rate

d fr

om b

lood

con

stitu

ents

C

ontin

uous

25

•

E. c

oli

• Y

east

Pi

nche

d flo

w fr

actio

natio

n •

Smal

l la

rge

(< 9

.9 μ

m)

EOF

tuni

ng

C

ontin

uous

24

• G

lial c

ells

•

Neu

ron

cells

Pi

nche

d flo

w fr

actio

natio

n

• 4.

9 μm

20

μm

Se

para

tion

usin

g vi

scoe

last

ical

ly tu

ning

of

flow

stre

am w

idth

C

ontin

uous

31

• Fu

nctio

naliz

ed p

artic

les

Pi

nche

d flo

w fr

actio

natio

n •

3.09

μm

5.

6 μm

M

utan

t DN

A h

ybrid

izat

ion

on su

rfac

e W

ild-ty

pe D

NA

hyb

ridiz

atio

n on

surf

ace

Con

tinuo

us

32

Tab

le 1

. O

verv

iew

of

(bio

logi

cal)

par

ticl

es s

epar

ated

by

lam

inar

flo

w t

hrou

gh c

hann

els

2.3 Separation based on filtration and flow around obstacles

The previous section described particle separations enabled by laminar flow profiles

in channels. However, flow and particle transport in microchannels can also be con-

trolled and manipulated through the placement of obstacle arrays in mirochannels.

Micromachined pillars can serve as physical barriers for actual filtering, or can be

used to generate specific flow pathlines. Particle separation approaches exploiting

these types of concepts are discussed in this section and summarized in Table 2, which

is located at the end of the section.

2.3.1 Filtration techniques

Particle separation and removal from solutions is commonly accomplished in the con-

ventional lab by filtration, using materials with controlled pore sizes to retain parti-

cles larger than the pores and allow smaller particles to pass through. Key to a good

filter is the precision with which pores of a certain size are formed. Micromachined

filters often take the form of arrays of posts positioned across a microchannel, ex-

ploiting the micrometer size resolution and precision of photolithographic processes

to achieve very small and well-defined gaps between posts. Kricka and Wilding de-

monstrated several early examples of such systems micromachined in silicon. An

array of posts with complex geometry, separated by short, 5-μm-wide periodic ser-

pentine channels was used to separate erythrocytes and 5.78-μm latex particles in se-

parate experiments from small fluid samples. While erythrocytes were sometimes

able to pass through the gaps because of their ability to deform, the particles were too

large and remained on the injection side.4 In other examples from the same group, post

arrays were used to separate erythrocytes from leukocytes prior to polymerase chain

reaction (PCR) to amplify leukocyte DNA.35-37

More recently, the Groisman group used a filter-like concept to construct a micro-che-

mostat for the cultivation of microorganisms. The biological particles were retained

in 100-μm-wide, 6-μm-deep chambers, which were connected with shallow (0.6 μm)

capillaries to adjacent 150-μm-wide, 6-μm-deep channels. The capillaries were small

enough to prevent escape of micro-organisms from the chambers, but large enough

to allow for diffusive exchange of nutrients and metabolites between the channels

and chambers. Fresh medium was continuously passed through the channels to ensure

that medium was refreshed every 40 s next to the chambers. This method was used

Chapter 2

36

to culture E. coli and yeast and expose them to other media.38 This filtering approach

was also applied by the same group to the separation of plasma from blood, using a

device with a deep main channel with arrays of 1920 shallow channels aligned along-

side each side. Cross-flow filtration was used to extract ~8% of the blood volume as

plasma through these side channels as the blood sample, including all cellular com-

ponents, flowed past. Again, the shallow channels had a depth of only 0.5 μm, pre-

venting the co-extraction of blood cells along with plasma. Dilute blood was pumped

with a pulsatile average flowrate of 0.65 µL/min to prevent the device from clog-

ging.39 To clean the device, a flow was applied through the shallow side channels to

wash away the small components from blood and exchange the medium. The amount

of red blood cells in whole blood were reduced by a factor of ~4000 while leaving

98% of the white blood cells behind.40

Cross-flow filtration for blood separation has also been performed by other resear-

chers in straight channels which were separated into three regions by two arrays of

pillars along the length of the channel. The flow containing blood was introduced

into the central region in between the pillar arrays. Small components were able to

pass the array and reach the outer channel areas, leaving larger particles behind.41-43

This was done in a PDMS device to remove leukocytes from blood, a process known

as leukapheresis,41 with as a result the isolation of 50 % of the erythrocytes and a gre-

ater than 97 % depletion of the leukocytes. Other examples also involve the separa-

tion of erythrocytes from leukocytes,42, 43 with subsequent cell lysis integrated for

PCR.43 The same approach was used to separate large myocytes from small non-my-

ocytes, with a viability test afterwards. It was noted by the author that the small 7-9

µm non-myocyte cells had to deform as they passed through 5-µm-wide gaps.44

Hydrodynamic filtration (HDF), developed by the group of Seki, resembles the cross-

flow filtration devices presented above, except that the extraction side channels have

the same depth as the main channel itself. There are thus no physical barriers pre-

venting extraction of large blood components into these channels, rather, selective

extraction of different-sized components is controlled by flowrate distribution. In

HDF, a continuous flow with randomly spread particles enters a main channel with

multiple side branches.45 Only the fluid flow near the wall enters the side branches,

with the amount of fluid leaving the main channel being dictated by the relative flo-

wrate distribution between the main and side channels. This will thus also define the

Separating Biological Particles in Flowing Microsystems

37

width of the flow stream near the walls which enters the side branch. The relative

flowrate distribution is determined by the flow resistances in the various channels,

which are in turn determined by channel dimensions. When the flowrates into the

side branches are sufficiently low, only a small portion of the flow near the wall goes

into the side branches, as shown in Fig 3A. Particles with radii larger than the width

of this flow will not enter the side branch, but will be carried past the channel ope-

ning. The particles remain in the main channel even if present near the wall or if their

diameter is smaller than the cross-sectional area of the side branch. Interestingly, the

particles are aligned against the wall as the fluid near the wall is extracted from the

main channel, and thus after each branch the particles are shifted towards the walls

of the channel, shown in Fig. 3B. This operation mode is called the “flow state”. Par-

ticles with radii small enough to be contained within the width of the flow portion en-

tering the side branches, enter the side branch with the flow, schematically shown in

Fig 3B-D. This technique was utilized to concentrate and align 1 to 3-µm-diameter

particles by extracting fluid using multiple side branches. Further downstream, the re-

lative flowrate distribution changes and the portion of fluid that enters the side

branche increases. Particles that fit into the fluid portion are collected and separation

Chapter 2

38

Figure 3 Schematic overview of hydrodynamic filtration. The relative flowrate distrubtion isindicated with the dashed line. (A) The relative flow rate distribution into the side channelsis low, only fluid leaves the main channel. (B) After repeated operation with relative low flowrate distribution particles are aligned against the wall. Relative flow rate distribution changed,therefore aligned erythrocytes go into the side branches. When the relative flow rate distri-bution is further increased, small (C) and large (D) particles can enter the side branches.

A B

C D

is performed according to size.45, 46 The power of this technique is its dual capability

for particle concentration and separation.

The flow state mode of HDF can be exploited to concentrate and align particles not

only at the walls but also in the center of the channel. To do this, fluid is extracted at

one side of the main channel to align the particles at that wall. Further downstream,

the fluid from the side branches is recombined with the fluid in the main channel,

and the aligned particles are shifted towards the center.47 Side branches on the oppo-

site side of the channel operated in the flow state mode enhance particle concentra-

tion, alignment, separation selectivity and recovery.46 Particles could also be

concentrated and aligned in the center of the channel with side branches on both sides

of the channel. Fluid was first removed from the channel on both sides, to be recom-

bined further downstream.47

HDF was applied to blood samples to enrich the number of leukocytes compared to

erythrocytes. Erythrocytes were able to enter the side channels, as the disc-thickness

was smaller than the width of the fluid stream entering the side branches, as shown

in Fig. 3B. The alignment of these cells with respect to the flow was the same as in

the PFF case, shown in Fig 2. Prior to HDF, 780 times more erythrocytes then leu-

kocytes were present in the sample. After enrichment, an increased number of leu-

kocytes was measured, with erythocytes outnumbering leukocytes by a factor of 29

rather than 780.45

Separation of liver cells, hepatocytes and nonparenchymal cells is another example

of the application of HDF to biological particles.48 Viability tests before and after se-

paration yielded the same results, indicating that shear stress didn’t influenced the

cells. After filtration, off-chip anti-albumin antibody staining was used to confirm

that hepatocytes were separated from non-parenchymal cells.

The flow state mode of HDF can also by employed to exchange cell medium.49 This

was done to study the time-dependent exposure of HeLa cells to Triton X-100, a sur-

factant used to solubilize the cellular membrane. The device in this case had a se-

cond inlet located on one side. On the other side and further downstream, multiple

side branches were located to replace the original medium with Triton X-100 using

the flow state mode. The “stimulation area”, so called because the cells were meant

Separating Biological Particles in Flowing Microsystems

39

to interact with the Triton X-100 in this region, was located downstream from these

channels. A second inlet and set of side branches were located after this region to re-

move the Triton X-100 and wash the cells in the same manner as the Triton X-100 was

introduced. The residence time in the “stimulation area” could be changed by vary-

ing the length of the “stimulation area” or changing the flowrates. Cells were expo-

sed 17 to 210 ms and collected in the outlet to study their viability.

HDF is thus a proven technique for the concentration, alignment and enrichment pro-

cedures of (biological) particles. Furthermore, the technique can be applied to toxi-

city studies, biochemical assays and other studies, where controlled exposure to fluids

or reagents is required. It can also be used for encapsulated biological material, as it

has been shown that this technique is applicable to the separation of droplets.50

2.3.2 Obstacle array for sorting by diffusion

The intelligent placement of obstacles in a channel can be used to size-separate par-

ticles by diffusion. Particles migrate not only under the influence of the flow, but dif-

fuse at the same time. Rectangular obstacles placed at a 45-degree angle with respect

to the direction of flow create an asymmetric obstacle array. Slowly diffusing parti-

cles are likely to travel straight through the obstacle array without being deflected

Chapter 2

40

Figure 4 Schematic representation of the separation mechanism underlying the Brownian ra-chet. (A) Particles pass gap A between the obstacles and are transported in the direction ofgap B by the electric field, E. Between gap A and B the particles diffuse perpendicularly tothe direction of transport. Small particles, which diffuse faster, are deflected by the obstacleand pass through gap B+, whereas larger particles diffuse slowly and pass through gap B. Inthis way, particles are separate based on diffusion. (B) Schematic overview of the separationof a sample containing small and large particles in the obstacle array. (Figures adapted fromDuke et al. ref. 51)

injection

small

large

A B

y

x

B

B+

B-

A

while passing the obstacles on the short side,51 as shown in Fig. 4A. Diffusion is a

size-dependent process and small particles, with larger diffusion coefficients, diffuse

faster than large particles. Small and therefore faster diffusing particles will also tra-

vel straight down through the obstacle array. However, at the same time small parti-

cles diffuse perpendicular to the direction of the flow to such an extent that they may

be deflected by the obstacles and pass along the long side of the obstacle, as shown

in Fig. 4A. Particles with different sizes and therefore different diffusion coefficients

will be deflected by differing amounts and find themselves at different lateral positi-

ons, resulting in size separation based on diffusion (Fig. 4B). This approach is refer-

red to as the Brownian rachet.51-55 The time to diffuse in directions perpendicular to

the flow can be altered by modifying the distance between the obstacles and the trans-

port velocity of the analyte.

This diffusion-based method for size separation was used to separate 15-kbp and

33.5-kbp DNA fragments with radii of gyration of 0.31 µm and 0.43 µm, respectively.

After passing through a 10-cm-long array consisting of 1.5 x 6 µm obstacles separa-

ted by a 1.5 µm gap, the two DNA fragment bands were separated by 6.4 mm from

each other in the direction perpendicular to the flow, with each band exhibiting a nar-

row bandwidth.52 An array of different dimensions was used to separate T2 and T7

DNA (radii of gyration of 2.3 µm and 1.1 µm, respectively). Based on theoretical

considerations and observed results, the authors expect this technique to be applica-

ble to the separation of 50-500 bp ladders within 10 minutes, and mixtures of 65 kDa

and 68 kDa proteins within 2 minutes.53 Huang et al. improved the separation reso-

lution with a factor ~3.8 and increased the separation speed by a factor 10 simply by

tilting the direction of flow at a small angle with respect to the original direction of

flow (the y-axis in Fig. 4A).55 The device works better at low than at high flowrates,

as the particles have more time to diffuse at the lower flowrates.

2.3.3 Deterministic lateral displacement

Besides filters and Brownian ratchets, an obstacle array in a microchannel may also

be used to generate a pattern of flow streamlines to separate particles. Huang et al.placed obstacles in a channel with a gap larger than the particles under investigation

to generate flow streamlines, mentioned with G in Fig. 5.56 This concept is based on

the fact that the flow splits into defined lanes (defined by flow streamlines) when it

travels around the obstacles and will recombine after passing. In a periodic array of

Separating Biological Particles in Flowing Microsystems

41

obstacles, with a horizontal shift, δ, of the obstacles each row with respect to the pre-

vious row, this process is repeated each row when flow passes the obstacle array,

shown in Fig. 5. However, the relative positions of the lanes change each row which

is determined by the shift of the obstacles and rejoins its original position in a perio-

dic way. Now imagine two types of particles present in this flow, with radii smaller

and larger than the lane width. Particles with radii smaller than the lane width will re-

main within the lane, and return towards the same relative position one spatial period

downstream, via a so-called “zigzag mode” around the obstacles, shown in Fig. 5. If

its radius is larger than the lane width, the particle transfers into the neighbouring

lane when its centre of mass is within that lane. This is repeated every time a parti-

cle passes a row in the periodic array of obstacles, resulting in the deflection of its po-

sition within the obstacle array relative to the direction of flow. This mode is called

the “displacement mode”. Depending on their size, particles are transported in the

Chapter 2

42

Figure 5 Size-separation modes for deterministic lateral displacement. G indicates the gapwidth between posts, λ the distance between rows, and δ the displacement of posts in one rowwith respect to posts in adjacent rows. At the far left, the flow lanes around the obstacles andtheir pathway through the post array are depicted. Towards the right, the trajectories for smallparticles and large particles through the post array are given. Small particles remain in theirrespective lanes in the so-called ‘zigzag-mode’, and follow a relatively straight path throughthe post array as a result. Large particles are transferred into the neighbouring lane at eachrow and are thus deflected with respect to the flow when passing through the array. (Figureadapted from Morton et al. ref. 58)

G

δ

λ

small particle large particle

zigzag-mode displacement mode

Separating Biological Particles in Flowing Microsystems

43

• Pa

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•

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•

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• T2

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ge D

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•

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• 48

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Con

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• 61

arti

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lace

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t •

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. Se

para

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•

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st

Det

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tera

l di

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• 17

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13

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card

iom

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nd n

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Con

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Tab

le 2

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l pa

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les

sepa

rate

d by

fil

trat

ion

or f

low

aro

und

obst

acle

s pl

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in

the

chan

nels

“zigzag” or in the “displacement mode” and can thus be separated by size, shown in

Fig. 5. The main difference between this technique and the Brownian rachet51-55 is

that this method does not depend on diffusion.56 In fact, increasing the flowrate would

decrease the effect of diffusion and give sharper separation bands. In contrast, an inc-

reased flowrate in a Brownian rachet would decrease the time for diffusion would

decrease and thus the separation would be less efficient.

Size separation with deterministic lateral displacement was demonstrated by Huang

et al. with 0.80 – 1.03-µm-diameter, fluorescently-labelled particles, and 61 kbp and

158 kbp artificial chromosomes from E. coli. Furthermore it was shown that DNA

could be preconcentrated when a broad stream was directed to one side of a channel

exploiting the displacement mode.56 The gap size and horizontal shift of the row in

the array determine the particle size that is transported according to the displacement

mode. A model for the critical particle size depended on the gap between the posts (Gin Fig. 5), distances between the rows (λ in Fig. 5) and fluid transport mechanism was

presented and supported with data from 2.3-µm to 22-µm-diameter particles.57

Arrays with different characteristics can be used to focus or guide particles. Separa-

tion of lymphocytes from blood platelets was shown with an obstacle array using nu-

cleic-acid-staining dye to follow the trajectories of the different cell types. The array

was designed such that lymphocytes are transported in the displacement mode and

platelets and other staining residues are subjected to the zigzag mode.58 In an array

where platelets are transported in the displacement model it was demonstrated that

this technique is useful for downstream analysis as the platelets were stained with

immunofluorescent dye (phycoerythrin conjugated CD-41 antigen). This was done by

parallel introducing a running buffer with the cells, the antigen solution and washing

buffer. The array runs the cells from the running buffer to the washing buffer through

the antigen solution. Similar to this was the chromosomal separation of E. coli bac-

teria using lysis buffer.59

Different arrays in series was used to size blood components.60, 61 In the first array, the

smallest particles follow the direction of the flow and larger particles are displaced.

In the next array the stream with the smallest particles continuous in the direction of

the flow. The fraction containing larger particles is then again separated an array, with

smaller particles continuing in the direction of the flow and larger ones which are

Chapter 2

44

displaced by the array. With several arrays placed in series different lateral displace-

ment routes are generated that particles follow according to their size and particles end

up on different lateral positions perpendicular to the flow direction. Inglis et al. se-

parated monocytes and lymphocytes for 99% from red blood cells without the need

to lyse red blood cells prior to separation, which is required for flow cytometry. They

reported that red blood cells behaved as particles with diameters less than 5 µm whe-

reas the diameter is 6 to 8 µm and thickness is 2 µm. This reminds us of the way

erythrocytes behave in PFF25 and HDF.45 This device was used to differentiate he-

althy lymphocytes from larger malignant lymphocytes after activation.60 In another

study they first separated large white blood cells from cells smaller than 5 µm in one

array. The platelets were then separated in an array with different characteristics to

study the effect of heat and thrombin on the size.61 Combining these studies would re-

sult in one device with different arrays capable to size separate whole blood in its

different components. Davis et al. took blood apart with deterministic lateral displa-

cement arrays. First removing the large white blood cells would prevent the device

from clogging when separating the smaller red blood cells and platelets in an array

with smaller gaps. However, they were also not able to do it in one device jet.62

Deterministic lateral displacement was also used to select cells for tissue enginee-

ring. A heterogeneous cell suspension containing epithelial and fibroblast cells as a

model for large cardiomyocytes (17 ± 4 µm diameters) and small non-myocytes (12

± 3 µm diameters), respectively, was run over an array. Large cells were purified

(>97%) and viable after only a single run through the device. However, high input

concentrations (1 x106 cells mL-1) of cells caused clogging of the device.63

2.4 Separation based on secondary flow effects

The flow of a solution through appropriately designed channels can induce secon-

dary flow effects that significantly decrease the time required to fully mix two solu-

tions. So-called Dean flows can be generated in solutions flowing at higher flowrates

through curved channels64-67, and microstructured grooves on the bottom of the chan-

nel can generate helical flow along the length of the channel.68-70 Both effects were

first used for mixing fluids, but have now also been used for particle separation. The

reader is referred to Table 3 for an overview of the different particle types, both in-

organic and biological, separated using secondary flow effects.

Separating Biological Particles in Flowing Microsystems

45

2.4.1 Dean flows

It was shown that mixing can be enhanced in curved channels where secondary flow

effects, so-called Dean flows, are present.64-67 Besides mixing, Dean flows can also

be exploited for separating particles, as follows. In curved channels, the fluid at the

center of the channel has a higher velocity than fluid at the walls, resulting in the cen-

tral fluid elements experiencing centrifugal forces towards the outside of the curva-

ture. The generated radial pressure gradient induces a recirculation pattern across the

cross-section of the channel, as the slower fluid at the walls is forced inward towards

the channel center. In the mid-plane of the channel, the fluid moves towards the out-

side of the curve and along the top and bottom back towards the inside of the curve,

resulting in two symmetric vortices in the top and bottom halves of the channel. This

is the secondary flow effect known as Dean flow. Particles in the fluid are influenced

not just by the Dean flow, but also by a centrifugal force and the tubular pinch effect

(TP effect).71, 72 The TP effect is a balance of three lateral forces, one of which is the

wall effect, where an asymmetric wake of a particle near the wall leads to a lift force

away from the wall,73 another a shear-gradient-induced lift towards the wall.74 While

being transported through a curved channel, the Dean flow, centrifugal force and tu-

bular pinch effect on the particles are all balanced, which leads to longitudinal

alignment of particles according to size and mass, shown in Fig. 6.71, 72, 75-79 Di Carlo

recently reviewed these fluid inertia and provided fundamental explanations about

the fluid dynamics and practical design rules.80

Chapter 2

46

Figure 6 (A) Schematic representation of the mechanism for separation by Dean flow, with amixture of randomly distributed particles introduced at the inlet. When particles arrive at theoutlet, they are aligned longitudinally, due to a balance of different forces as described in thetext. (B) When the Dean flow effect is negligible, the red blood cells are randomly distributedin the channel. (C) Schematic overview of the alignment of red blood cells due to Dean flowsin the x-y plane. (D) Alignment of red blood cells due to the Dean flow in the z-y plane, withimages taken in the midplane and on the bottom of the channel. (Figure B-D adapted from DiCarlo et al. ref. 78)

Inlet Outlet

A

B

C D

Particle filtration and separation based on the Dean effect was demonstrated with 2-

to-20 µm diameter particles in curved,75 spiral,71, 72 arc76 and asymmetrical curved

channels.77, 78 The flow velocity and radius of the curve are important factors deter-

mining the particle cut-off size. As the Dean flow effect increases, particles will be

located more towards the inside of the curvature. Di Carlo et al. used this method to

align polystyrene beads, silicone oil droplets, trypsinized H1650 cells and red blood

cells.78 Furthermore, they were able to separate platelets from whole blood. It was

shown that the relative number of platelets was enriched by approximately a factor

of 100.77 Interestingly, the disc-shaped red blood cells aligned rotationally so that the

disk axis was directed to the nearest walls, as shown in Fig. 6B.78

Dean flows have been exploited for three-dimensional focusing of particles. In a cur-

ved channel, particles are vertically focused by the presence of the Dean vortices.

Three-dimensional focusing can be achieved by introducing two fluid streams via

side channels located after the curve to pinch and thereby horizontally focus the sam-

ple stream.81, 82 This was done with particles having sizes and densities comparable

to human CD4+ T-lymphocytes used for e.g. HIV-diagnosis. However, results with

the human cells had not yet been reported at the time of this review.

Dean vortices have been used to separate neuroblastoma and glioma cells with 80%

efficiency and relatively high viability (>90%).83 The throughput of 1 million

cells/min is comparable to conventional flow cytometry techniques and 100-fold hig-

her than in PFF.25

In a serpentine channel, the inside and outside of the curvature switches at each turn.

An analyte will thus be subjected to a back-and-forth transfer across the width of the

channel while moving downstream. The influence of the forces is larger on double-

stranded DNA (dsDNA) than on single-stranded DNA (ssDNA) and thus these DNA

strands will be separated perpendicular to the flow.67, 84, 85 Wu et al. designed a chan-

nel with three inlets meeting at a single junction, where a sample stream from the

central inlet is pinched in between an acting flow and a protecting sheath flow coming

from two side inlets and operated at high and low velocities, respectively.86 This re-

sulted in a curved flow trajectory of the centered sample stream at the junction, which

then entered a main channel and flowed away from the junction. The authors claimed

that at high flowrates Dean vortices are present and large particles get enough mo-

Separating Biological Particles in Flowing Microsystems

47

mentum to escape the flow trajectory in the main channel. Upon a rapid change of the

fluids’ and particles’ momentums, a mismatch would cause them to separate. This

method was used to separate E. coli bacteria from a sample containing blood cells and

resulted in a 300-fold enrichment of bacteria. The authors demonstrate a separation

which they claim is based on soft inertial force-induced migration. However, Di Carlo

is of the opinion that a kinetic separation mechanism is the most likely cause of the

observed separation, whereby hydrodynamically focusing cells/particles of different

sizes against a wall led to faster migration of larger particles away from the wall.80 In

fact, it is our impression as well that the separation mechanism in this example shows

similarities to the mechanism described for PFF where particles are pinched against

the wall with a carrier flow. The only difference here is that the particle-containing

flow is pinched in between two fluid flows in this case; however, particles are still for-

ced to sample different streamlines depending on their size.

2.4.2 Hydrophoretic separation

Choi & Park defined hydrophoresis as the movement of suspended particles under the

influence of a microstructure-induced pressure field.87 Their work is based on the se-

condary flow effect first introduced by Stroock et al. to enhance mixing in micro-

channels.68-70 Mixing is performed in a channel with slanted ridges, which generate a

transverse pressure gradient resulting in a helical recirculation in the direction of the

flow, as shown in Fig. 7A. Particles present in the flow will also be influenced, and

are transported towards the walls due to the direction of the secondary flow effect,

schematically shown in Fig. 7B-C. An alternating pattern of slanted ridges on the bot-

tom and top of the channel is used to focus particles on one side of the channel (Fig.

7B). In a region with only slanted ridges on the bottom, particles align along one side

of the channel, due to the rotation of the helical flow (Fig. 7C). After alignment, par-

ticles can then be separated according to their sizes relative to the gap between slan-

ted obstacles at the top or bottom of the channel. Large particles align at the center

of the z-axis and follow the flow going up and down over the obstacles. Small parti-

cles are exposed to lateral pressure gradients along the width of the channel and fol-

low an oscillating path through the channel, as shown in Fig. 7D. Importantly, the

gap height limits the particle sizes which can be separated.87-89

During the cell cycle, cells pass through different stages representing, for example,

cell growth (G), DNA synthesis and mitosis (M). Cell size is indicative of where in

Chapter 2

48

the cycle a cell finds itself, and can therefore be used to select cells in the same stage.

Hydrophoresis was used to sort human leukemic monocyte lymphoma cells from the

U937 cell-line based on size. Leukemic cells in the early G0/G1- and late G2/M-pha-

ses, with diameters between 11 and 22 µm, were sorted in a channel with a 22-µm-

gap-height. Results showed that synchronization of early- and late-stage cells could

be performed. This method could be an advantageous approach for tumor cell detec-

tion and interaction studies with drugs.90

As in the work of Stroock,68, 69 a herringbone ridge structure in a microchannel was

used to generate a pressure gradient from the sides of the channel towards the center,

Separating Biological Particles in Flowing Microsystems

49

Figure 7 (A) Induction of a secondary helical flow pattern in a microchannel with an arrayof slanted ridges. (Figure adapted from Stroock et al. ref. 69) (B)-(D) Overview of the sepa-ration principle of hydrophoresis. Diagrams of channel cross-sections at different locationsdescribe the motions of large and small particles leading to their separation. Dark grey areasare obstacles (ridges) on the floor of the channel, light grey areas are obstacles (ridges) onthe ceiling of the channel. (B) Particles are aligned to one side of the channel by the lateralforces of the flow. (C) Particles are focused by the clockwise helical flow. (D) Large particlesremain focused in the area where no lateral pressure gradients are present. Small particles areexposed to lateral pressure gradients and are therefore transported in an oscillating patternas they follow the focusing and deviating flow. (Figure adapted from Choi et al. ref. 87).

Upwardflow

Diviationflow

Downwardflow

Focusingflow

Slanted obstacles

B C D

A

resulting in double helical flows in which hydrophoresis could be performed. When

using the herringbone structure, particles are focused in the center of the channel by

the helical flow and simultaneously centered on the z-axis by the obstacles. This ef-

fect was demonstrated with Jurkat cells.91 To study the behaviour of biconcave dis-

coid particles, samples containing red blood cells were introduced to a channel

containing a herringbone ridge array. Decreasing the height of the herringbone struc-

ture from 15.4 to 7.4 µm resulted in an increase in the number of red blood cells pre-

sent in the center of channel from ~57% to ~72% of the total cells present in the

channel.89 Focusing particles into parallel lines using parallel double helical flows

was accomplished by placing multiple herringbone arrays side by side in a channel.92

Particles which were were randomly distributed across the channel were rapidly fo-

cused into multiple lines in the direction of the flow. Since particles focus at the apex

of the herringbone structures, herringbone ridge arrays positioned laterally with res-

pect to one another in the direction of flow could be used to guide particles across the

channel in a controlled fashion.92

Hydrophoresis was also used in two examples in combination with other mechanisms

to perform separation. In the first example, λ-phage and micrococcus DNA were se-

parated in a device with a gap height of 1.2 µm.93 This gap height is between the 0.86-

and 1.45-µm radii of gyration of λ-phage and micrococcus DNA, respectively. Ho-

wever, DNA molecules need to be considered as spherical particles which can change

their conformation to form long strands. The small λ-phage could freely pass the gaps

and follow hydrophoretic alignment along one side of the channel. Large micrococ-

cus DNA first moved to the other side of channel along the slanted obstacle, before

changing conformation and passing the obstacles. To pass the obstacle, then, the DNA

has to change conformation, which costs energy. It was assumed that the obstacles

work as an energy barrier which can be exploited to separate DNA molecules, as the

conformational change of large DNA molecules costs more energy than that of small

DNA molecules.93 This is reminiscent of the entropic separation of DNA performed

by Han and Craighead in a microtrapping array.94

The second example used hydrophoresis to isolate and enrich white blood cells from

rat blood. Blood cells were focused on one one side of a channel with slanted obsta-

cles whose height was equal to half the channel height. These obstacles were located

on the top and bottom of the channel. The gap height (13 µm) in this region of the

Chapter 2

50

Separating Biological Particles in Flowing Microsystems

51

Part

icle

type

Fl

ow e

ffec

t Pa

rtic

le si

ze

Com

men

ts

Con

tinuo

us/

batc

h R

efer

ence

(s)

• Po

lym

er m

icro

sphe

res

Dea

n flo

w

• 1

- 20

μm

Test

ed w

ith d

iffer

ent g

eom

etrie

s C

ontin

uous

73

-79,

81,

86

• Po

lym

er m

icro

sphe

re

Hyd

roph

oret

ic se

para

tion

• 0.

5 –

20

μm

C

ontin

uous

87

-90,

92,

93

Bio

logi

cal p

artic

les

•

Neu

robl

asto

ma

cells

•

Glio

ma

cells

D

ean

flow

•

15 ±

5 μ

m d

iam

eter

•

8 ±

3 μm

Se

para

tion

Con

tinuo

us

83

• Si

ngle

stra

nded

DN

A

• D

oubl

e st

rand

ed D

NA

D

ean

flow

•

20 m

er

U

neve

n di

strib

uted

C

ontin

uous

84

, 85

• H

uman

leuk

emic

mon

ocyt

e ly

mph

oma

cell

line

(U93

7)

Hyd

roph

ores

is

• 11

– 2

2 μm

Ea

rly a

nd la

te p

hase

cel

l sep

arat

ion

Con

tinuo

us

90

• Ju

rkat

cel

ls

Hyd

roph

ores

is

• 11

.0 ±

1.4

μm

Fo

cusi

ng o

f cel

ls

Con

tinuo

us

91

• 48

.5 k

bp λ

-pha

ge D

NA

•

115

kbp

mic

roco

ccus

DN

A

Hyd

roph

ores

is a

nd

entro

pic

trapp

ing

• 0.

86 μ

m ra

dius

•

1.43

μm

radi

us

Sepa

ratio

n C

ontin

uous

93

Tab

le 3

.O

verv

iew

of

(bio

logi

cal)

par

ticl

es s

epar

ated

in

devi

ces

whe

re s

econ

dary

flo

w e

ffec

ts a

re p

rese

nt

channel was larger than the size of red and white blood cells. The placement of the

obstacles resulted in anisotropic fluidic resistance that generated lateral pressure gra-

dients to induce the helical flow recirculation for focusing the cells. Actual separation

was performed using a filtration principle similar to that described above, using slan-

ted obstacles on top and bottom of the channel which partially cover the width of the

channel. The resulting gap on one side of the channel could serve as filtration pore.

The 4-µm-high gap in this region had a dimension that fell between the sizes of red

and white blood cells, as red blood cells have diameters of 6-8-µm, a thickness of 2

µm, and are deformable. Therefore, small red blood cells passed through the gap and

remained focused, whereas large white blood cells were blocked by the obstacles and

passed through the filtration pore (20-µm wide and 7.8-µm high) on the other side of

the channel. A 210-fold enrichment of white blood cells was achieved with a proces-

sing rate of 4000 cells/s.88

2.5 Separation based on external forces influencing particles

Previously discussed separation mechanisms were based on solution flows in micro-

channels; however, there are several types of external forces that can be applied across

a microfluidic channel to influence and control particles susceptible to them. These

forces can be used to transfer particles from their original streamline into other stre-

amlines. If particles exhibit differences in susceptibility, external forces can often be

used to separate them. This section discusses separation of particles in laminar flows

by external forces and covers free-flow electrophoresis, acoustophoresis and magne-

tophoresis. Table 4 reviews the different inorganic and biological particles separated

using the approaches discussed in this section.

2.5.1 Free-flow Electrophoresis

The technique of free-flow electrophoresis (FFE) predates the introduction of Lab-on-

a-chip technology.95 However, FFE was one of the first techniques performed in a

microfluidic channel, exploiting the fact that well-defined flows of fluids can be ge-

nerated side by side without mixing. The first examples of FFE on a chip described

the analysis of labelled amino acids and proteins.96, 97 Recent papers by Kohlheyer etal.98 and Turgeon & Bowser99 thoroughly review the theory and applications of FFE,

and are referred to for more detailed information. Briefly, in FFE the continuously in-

troduced sample flow runs down a shallow, ribbon-like channel in between carrier

Chapter 2

52

buffer solution. An electric field is applied perpendicular to the flow to induce elec-

trophoretic movement of the analytes. The result is deflection of charged species from

the direction of flow, with the angle of deflection increasing as the electrophoretic mo-

bility and/or electric field strength increases. The original sample stream is effectively

split into several streams, each containing analyte having a different electrophoretic

mobility.

Three different operation modes of FFE have been reported on chip. The use of car-

rier buffers with a constant composition with respect to pH and electrical conducti-

vity results in separation of the sample analytes according to their mobility. This mode

is called free-flow zone electrophoresis (FFZE)100, and several examples exist on

chip.96, 97, 101, 102 Isoelectric focusing (IEF) was achieved in a microchannel with FFE

by using carrier electrolyte solutions with ampholytes, which led to a pH gradient

when an electric field was applied across the channel. The analyte migrated through

the pH gradient until reaching the pH at which it had no net charge, at which point it

stopped migrating. This mode is called free-flow isoelectric focussing (FFIEF).102, 103

The presence of membrane proteins gives rise to cell organelles with different iso-

electric points. Mitochondria from lysed HT-29 cells were focused at a pI between 4

and 5, while intact cells were attracted to the anode due to their negative surface char-

ges. Fluorescently tagged cell organelles, peroxisomes and mitochondria from HeLa

cell lysate were focused at pI values between 4 and 5 but were not separated, indica-

ting that they may have the same pI’s.104

In free-flow isotachophoresis (FFITP)105, 106 the sample flows in between a leading and

terminating carrier buffer containing ions with a higher and lower mobility than the

analyte, respectively. Rearrangement of samples containing several analytes into se-

parate bands in microfluidic devices was demonstrated using a mixture of fluorescein,

Eosin G and acetylsalicylic acid; the separation took less than 1 minute. The separa-

tion of a reaction mixture of myoglobin and fluoresceinisothiocyanate was also re-

ported, demonstrating the technique’s potential for sample preconcentration.105

2.5.2 Acoustic particle manipulation

Lining up particles and transferring them into other streamlines can also be done with

acoustic forces.107 Particles suspended in a fluid are influenced by acoustic waves di-

rected perpendicular to the flow direction. Acoustic or ultrasonic waves cause axial

Separating Biological Particles in Flowing Microsystems

53

acoustic radiation forces, Fr, influencing the particles and is given in Equation (2).

Where p0 is the acoustic pressure amplitude, Vc is the volume of the particle, λ is the

wavelength and k defined by 2π/λ and x is the distance from a pressure node. The

acoustic contrast factor, Equation 3, represented by ɸ depends on both the compres-

sibility of the particle (βc) and medium (βw) and the density of the particle (ρc) and

medium (ρw).

Particles in continuous flows influenced by acoustic forces move to the standing wave

pressure node or anti-node, which most notably depends on the contrast factor, i.e.

density and compressibility of the particle. If a particle has a higher density and/or is

less compressible compared to the medium it is suspended in, it will move towards

the nearest pressure node. If the medium has a higher density and/or is less com-

pressible than the particle, the particle moves towards the antinode, shown schema-

tically in Fig. 8A.108-110 Aligning particles on the nodes and antinodes was used to

translate particles from one solution into another while the two flows run side by side

Chapter 2

54

Figure 8 (A) Schematic representation of particles with different contrast factors aligned alongthe walls and in the center of the channel, respectively. (B) Cross-section of channel witherythrocytes and lipid particles collected on the pressure node (center) and antinodes (sides)of the acoustic wave standing in the channel, respectively. (Figures adapted from Laurell etal., ref. 107)

Flow

B

A

�� = − ( �02��

2� ) �(�, �) sin(2� � ) (2)

�(�, �) =5�� − 2�

2�� + � −��

� (3)

through the channel.111 Particles with the same acoustic contrast factors can be sepa-

rated based on sizes as the particle volume influences the acoustic forces, as shown

in Equation (2). Larger particles obtain a larger acoustic force and will thus move

faster in a field of standing waves.112, 113 The amount of wavelengths generated in the

channel determines the number of lines with focused (biological) particles, as parti-

cles are focused on nodes and anti-nodes of the waves. This obviously depends on the

wavelength of the acoustics but also on the width of the channel.108, 114 Employing

the amount of lines was exploited for raw milk lipid enrichment and depletion in a

channel splitting in three outlets. In 750-µm-wide and 1125-µm-wide channels two

and three half-wavelengths were generated applying the same excitation frequency

and input power. In the narrow-channel case lipids were focused on the anti-node and

lipid enriched milk was collected via the central outlet. In the wider-channel case the

central flow was depleted of lipids, however, small proteins, lactose and casein par-

ticles were still present. A beneficial effect from the lipids being transported on the

antinodes is that they don’t reach the walls and start clogging of the channel via ads-

option to the walls.115

One of the first publications of the Laurell group directly showed that this approach

was applicable to biological particles as erythrocytes and lipid vesicles suspended in

blood plasma were separated from each other, shown in Fig. 8B.116, 117 This method

was successfully used during open-heart surgery to separate lipid micro-emboli from

blood. The process was continuously performed with flowrates up to 0.5 mL min-1

whereas in conventional methods 400-500 mL needs to be collected to process in a

batch.117-119 Three streams of particles were generated using acoustic forces as erythro-

cytes were located on the pressure node in the center of the channel and lipid vesi-

cles were located on the anti-node near the walls. In a channel ending in three outlets

70% of the erythrocytes was collected in 1/3 of the original fluid while 80% of the

lipid vesicles were removed. Erythrocytes focussed in the center of the channel by

acoustic forces could also be washed.110 On one side of the channel washing buffer

was introduced via a side channel and via the opposing side channel buffer extracted.

The erythrocytes were shifted towards the side where fluid was extracted. However,

acoustic focussing was used to realign the cells in the center. Repeating this opera-

tion resulted in 58% recovery of the cells and removing 98.3% of the Evans Blue

present prior to washing. Recent work showed that the developed acoustic technique

can be used to process plasma fraction out of whole blood.120 The blood focussed in

Separating Biological Particles in Flowing Microsystems

55

the middle of the channel could be redrawn from the channel via outlets in the co-

verplate along the length of the channel. Different concentrations of blood are col-

lected via these outlets and plasma was collected at the end of the channel. The plasma

was of high enough quality for plasma transfusion and screened for prostate specific

antigen by coupling to an antibody microarray. Integration of the microarray on-chip

would generate a tool for clinical diagnostics.121, 122

Red blood cells and E. coli could be aligned and patterned by placing two transdu-

cers parallel or orthogonal to the channel, respectively. Interference of the two gene-

rated acoustic waves of orthogonal placed transducers resulted in nodes and antinodes

in a two dimensional pattern. In other words the acoustic waves act as tweezers to hold

cells on one position.123

Washing with acoustics can also be performed when introducing clean medium via

the central inlet of three inlets. The contaminated medium, introduced via the two

side inlets, remains at both sides due to the well-defined laminar flows. Antibody-

based affinity beads can be extracted from the contaminate medium employing the

acoustic focussing of the beads in the center of the channel. When placing more chips

in series beads can be extracted and washed when the bead suspension is guided to-

wards the outside inlets of the channel. This technique was used to extract viral pha-

ges bound to beads.124, 125

Acoustic forces have proven not only to be able to align particles but are also capa-

ble of trapping cells and avoid them to contact the walls.126 This was done with in a

channel with the height of half a wavelength and a reflector on top. Miniaturized

transducers were placed on the cross-section of the main channel, introducing the hy-

drodynamically focussed sample and side channels for perfusion of the trapped cells

with analyte for assays. First experiments demonstrated that rat spleen cells could be

trapped and hold on with acoustic trapping. Cell culturing experiments were perfor-

med with yeast strains, Saccharomyces cerevisiae, capable of producing yellow fluo-

rescent protein under influence of the LEU2 gene. Trapped cells were perfused with

cell medium and successfully cultured for 6 hours. A genetically modified neural stem

cell line from embryonic rat hippocampus, HiB5-GFP, was used for a viability assay.

Trapped cells were perfused with phosphate-buffered saline solution for 15 minutes

prior to viability test based on the acridine orange marker. The increased fluorescence

Chapter 2

56

signal revealed that cells were still viable, indicating that the acoustic technique is ap-

plicable to samples containing living cell and not harmful to them.126

Extracting particles with acoustic trapping was recently performed for forensic ana-

lysis of sexual assault evidence. Sperm cells were separated from fluid containing

female epithelial cell lysate, as acoustic forces are strong enough to retain sperm cells

while the lysate is unretained. The use of acoustic trapping was beneficial for further

DNA analysis (extraction, quantization, amplification and separation) as an enriched

sperm cell sample was obtained.127

Jung et al. demonstrated that eukaryotic yeast cells could be separated from viruses,

MS2 bacteriophage, as the yeast cells were focused by the acoustic forces and bac-

teriophage remained unaffected under the applied conditions. Yeast cells, 4-6 µm,

showed similar focusing conditions as 2-3 µm polystyrene particles, with the same

density, however, yeast obtains a lower compressibility.128

2.5.3 Magnetic particle manipulation

Besides acoustic and electrokinetic forces, particles can also be susceptible to hand-

ling with magnetic fields. Magnetic susceptible materials can be manipulated with

magnetic forces, this can be used for pumps, valves, mixing and manipulating parti-

cles as reviewed by Gijs129 and Pamme.130 Like other particles the surface of magne-

tic particles can be functionalized to perform for example dynamic DNA

hybridization131 or immunoassays.132, 133 Magnetic forces are in these examples used

to retain particles on a specific location in the channel. The analyte could interact

with the surface while the particles were retained by a magnet placed under the chip.

The first examples of separating magnetic particles involved simply separating par-

ticles from solution. This was done by retaining the particles on an electromagnet

and releasing the current when fresh buffer close to the particles.134 Another method

used an H-channel design where particles were introduced via one of the parallel

channels and focused on the wall with a magnet in the same plane as the channel.

Via the horizontal-connection channel a sample as large as the volume of the chan-

nel junction is transported towards the other parallel channel by placing the magnet

on the other side of the channel. The horizontal channel can be filled with reagents

to interact with the surface of the particles.135 With this method the particles are se-

parated from the original solution they were suspended in.

Separating Biological Particles in Flowing Microsystems

57

The differences in magnetic forces, Fmag, experienced by the particles can be used

to separate them. The Fmag depends on the externally applied flux density (B) and

its gradient in the field (∇.Β) induced by the magnet, the difference in magnetic sus-

ceptibility between the particle and the fluid (Δχ), the particle volume (V) and the

permeability of a vacuum (μ0) and can be written as Equation (4).

Magnetic particles can thus be separated on either size or magnetic susceptibility dif-

ferences.136-138 This was performed in continuous flow with a magnetic field in so cal-

led free-flow magnetophoresis. In a microchannel the clean buffer solution flows

parallel with the sample solution containing 2.0-µm- and 4.5-µm-diameter magnetic

particles, with susceptibilities of 1.12x10-4 and 1.6x10-4 m3 kg-1, respectively. A mag-

netic field perpendicular to the direction of flow was used to separate particles from

each other and from the sample solution also containing the non-magnetic particles,

shown in Fig. 9. The particles with the largest susceptibility were deflected over the

largest distance. It was observed that large agglomerates of magnetic particles were

deflected in a larger extent than separate particles, due to the larger size of the ag-

glomerate.136 Not only particles but also superparamagnetic droplets could be mani-

Chapter 2

58

Magnetic field

Buffer/reagents

Samplemixture

non-M, M(a), M(b)

M(b)

M(a)

non-M

Figure 9. Schematic representation of the separation of magnetic particles by free-flow mag-netophoresis. At the bottom right, the mixture of (non-) magnetic particles enters the channel,with the magnetic susceptibility of particles increasing in the order non-M<M(a)<M(b). Buf-fer solution is also introduced from the right. Separated particles leave the channel on the leftwith particles with the largest susceptibility deflected over the largest distance into the y-di-rection. The buffer was in more advance research replaced for different reagents to performbioanalytical procedures. (Figure adapted from Pamme et al., ref. 136)

���� =∆�(∇ ∙ �)�

�0 (4)

pulated with magnetic forces. It was shown that these magnetic droplets, which can

be used to carry biological materials, were deflected perpendicular to the flow and di-

rected in to specific outlet channels.139

In two parallel channels connected by a diagonal junction channel instead of a hori-

zontal channel in the H-shaped channel size separation of magnetic particles was per-

formed together with separation from the medium.138 In contrast to the H-shaped

channel135 the magnetic field is induced in the plane of the channel but with ring sha-

ped wires placed on top and under the channel. The generated force in the direction

of the junction channel separated magnetic particles on their sizes, with the largest

particles obtaining the largest deflection as expected from Equation 4.138

Improving the reproducibility of the separation was done by changing geometry of the

outlet into a tapered structure and increasing the flowrate.137 This movement of mag-

netic particles perpendicular to the direction of flow was exploited to perform bio-

analytical procedures on the particle surface. Magnetic particles cross different

reagents streams running side by side introduced through different inlets. This was

performed with a similar setup to the one shown in Fig. 9, however the buffer was re-

placed with the different reagents. A model binding assay was performed with strep-

tavidin-coated particles running sequentially through buffer, fluorescently-labelled

biotin solution and buffer.140 A mouse IgG-sandwich immunoassay could be perfor-

med when particles crossed two fluid steams for binding and were washed with buf-

fer after each binding step.141

The surface of magnetic-susceptible particles can be coated with antibodies which

bind to specific receptors of cells. After binding, the cells became magnetically sus-

ceptible and can be separated, this process is called immunomagnetic cell separation

and was used to separate leukocytes from whole blood.142 Integration of micropatterns

from magnetic material under a slight angle with the direction of the flow in a mi-

crofluidic channel was used to generate magnetic field gradients. The coated leuko-

cytes were deflected from their streamline by the magnetic field while other blood

components were transported in the direction of the flow.142 Xia et al. separated E.Coli bound to magnetic nanoparticles from red blood cells processing 10.000 cells s-

1 and was able to obtain separation efficiencies up to 80%. They used two inlets to in-

troduce the sample and clean running buffer, while placing the magnet in plane with

Separating Biological Particles in Flowing Microsystems

59

the channel.143

The native magnetic properties of blood components can be used to separate different

type of blood cells from whole blood.144 Separation of red and white blood cells was

performed in a channel where after the inlet a ferromagnetic wire is placed along the

length, dividing a part of the channel into two parallel channels as a wall. A uniform

magnetic field perpendicular to flow direction near the wire generates a high magnetic

field gradient. When the magnetic susceptibility of a particle is larger than the buf-

fer, the particle is attracted towards the wire and called a paramagnetic particle. If a

particle bears a smaller susceptibility than the buffer, the particle is repelled from the

wire and named diamagnetic particle. In the case of red and white blood cells, red

blood cells are attracted towards the wire and white blood cells repelled. At the end

of the channel three outlets were located to collect the separated red blood cells and

the depleted whole blood. This method used only “stage 1” and the outlet geometry

shown in Fig. 10A. Via the middle outlet 91.1% of the red blood cells collected after

being separated from diluted whole blood, shown in Fig. 10B.

It was shown that separation could be improved in a chip containing three areas (se-

paration stages) containing wires. The first area is an exact copy from the above sys-

tem. After this first separation stage the attracted cells are guided towards the center

o f

Chapter 2

60

Inlet

Outlets 1

2

3

Stage 1 Stage 2 Stage 3

A

Figure 10 (A) Schematic representation of a three-stage magnetic blood separator using em-bedded ferromagnetic wires. (B) Stream of red blood cells (RBC) leaving via outlet 2 usingonly stage 1. (C) Improved separation of red blood cells from whole blood using three-stageseparation. Red blood cells leaving via outlet 2 and white blood cells (WBC) entering outlets1 and 3. (Figures adapted from Han et al., ref. 143)

CB

Separating Biological Particles in Flowing Microsystems

61

Mic

ro p

artic

les

Sepa

ratio

n m

echa

nism

Pa

rtic

le si

ze

Com

men

ts

Con

tinuo

us/

batc

h R

efer

ence

(s)

• O

rgas

ol/p

olya

mid

e sp

here

s A

cous

toph

ores

is

• 5

μm d

iam

eter

•

Sepa

ratio

n fr

om m

ediu

m

Con

tinuo

us

108,

110

, 111

, 11

4 •

poly

amid

e sp

here

s

• 1.

9 μm

dia

met

er

• Fo

cusi

ng

Con

tinuo

us

109

• Su

perp

aram

agne

tic p

artic

les

Mag

neto

phor

esis

•

0.8

- 4.5

μm

di

amet

er

• Se

para

tion

from

med

ium

and

no

n-su

scep

tible

par

ticle

s C

ontin

uous

13

2, 1

34-1

38

Bio

logi

cal p

artic

les

•

HT-

29 c

ells

(Hum

an c

olon

ca

rcin

oma)

Fr

ee-f

low

isoe

lect

ric

focu

sing

•

N.A

. Fo

cusi

ng o

f cel

ls o

n IP

C

ontin

uous

10

4

• H

eLa

cells

Fr

ee-f

low

isoe

lect

ric

focu

sing

•

N.A

. Fo

cusi

ng o

f cel

ls o

n IP

C

ontin

uous

10

4

• R

aw m

ilk li

pids

A

cous

toph

ores

is

• N

.A.

Lipi

d en

richm

ent o

r dep

letio

n C

ontin

uous

11

5 •

Stre

ptav

idin

-coa

ted

bead

s A

cous

toph

ores

is

• 2.

8 μm

dia

met

er

Sepa

ratio

n un

boun

d m

ater

ial

Con

tinuo

us

124

• Su

perp

aram

agne

tic M

OA

C b

eads

A

cous

toph

ores

is

•

Sepa

ratio

n fr

om u

nbou

nd p

eptid

es

Bat

ch

125

• Sp

leen

cel

ls fr

om ra

ts

Aco

usto

phor

esis

•

N.A

. Tr

appi

ng

Bat

ch

126

• Y

east

(Sac

caro

myc

es c

erev

isia

e)

Aco

usto

phor

esis

•

N.A

. Tr

appi

ng fo

r cul

turin

g te

sts

Bat

ch

126

• R

at e

mbr

yoni

c hi

ppoc

ampu

s neu

ral

stem

cel

l lin

e H

iB5-

GFP

A

cous

toph

ores

is

• N

.A.

Trap

ping

for v

iabi

lity

test

s B

atch

12

6

• Sp

erm

cel

ls

• Fe

mal

e ep

ithel

ial c

ell l

ysat

e A

cous

toph

ores

is

• N

.A.

Trap

ping

, was

hing

and

sepa

ratio

n C

ontin

uous

12

7

• Y

east

(Sac

caro

myc

es c

erev

isia

e)

• M

S2 b

acte

rioph

age

Aco

usto

phor

esis

•

4-6

μm

• 30

nm

dia

met

er

Sepa

ratio

n C

ontin

uous

12

8

• E.

Col

i M

agne

toph

ores

is

•

Sepa

ratio

n fr

om re

d bl

ood

cells

C

ontin

uous

14

3

Tab

le 4

.O

verv

iew

of

(bio

logi

cal)

par

ticl

es s

epar

ated

usi

ng e

xter

nal

forc

es

the channel to flow in between two wires. The remaining fluid will flow around these

two wires for a second extraction step using the same principle, “stage 2” shown in

Fig. 10A. Before repeating this operation one more time the attracted cells in the

outer channels are guided towards the center stream of cells. The complete setup is

shown in Fig. 10A. The separation of red blood cells increased to 93.5% and cells

were collected in the center outlet, whereas 97.4% of the white blood cells were col-

lected in the outermost outlets, shown in Fig. 10B and C.144

The density of red and white blood cells is larger than the density of the medium they

are suspended in. Therefore, cells sediment in time, which becomes a problem when

running assays for longer period of time. This effect could be reduced by adding bo-

vine serum albumin to the medium.145

When ferromagnetic wires are embedded in the bottom of the channel the magnetic

susceptible (biological-) particle solution flows over the wires. However, the high

magnetic field gradients are still present and will thus affect the particles. With the

wire placed under an angle with the direction of flow white blood cells were separa-

ted from whole blood. White blood cells were deflected in the direction of the angle

of the wires when crossing the wires and are thus separated by lateral-driven mag-

netophoresis.146 Deposition of ferromagnetic dots on the bottom of a channel can be

used to extract red blood cells from blood running over the dots. With an external

applied magnetic field perpendicular to the flow a magnetic field gradient is induced

and red blood cells were trapped on the dots while other components continued with

the flow.147

2.6 Conclusion

As this review reveals, microfluidics now offers a wide variety of particle separation

techniques for situations where particles are the analyte of interest or interact with the

analyte. Separation can in both cases be performed to separate particles from other

particles and/or from media. The well-defined laminar flow patterns in specially de-

signed microfluidic devices can be exploited to guide different particles into different

streamlines. Furthermore, external forces can be used to influence particles to actively

switch streamlines. To study the separation mechanism involved, proof-of-principal

studies are generally first performed with well-defined nano- or microspheres. More

Chapter 2

62

advanced research will then focus on the application of a separation technique for

more biologically relevant particles, such as blood cells, bacteria and DNA. Pamme

also noted in her 2007 review that applications involving biological particles were be-

coming increasingly important for continuous flow separation approaches.10

Biological particles like (blood) cells and DNA are non-spherical and/or deformable.

In several studies, it has been observed – not surprisingly - that these types of parti-

cles do not behave like solid nano- or microspheres. In PFF25 and HDF,45 biconcave

discoid-shaped red blood cells aligned in the flow according to their shortest axis,

thickness. Red blood cells can pass through small gaps due to their thickness and de-

formability,88 and DNA strands can change conformation from coiled- to stretched-

form to pass through nanogaps.93 Table 5 gives an overview of the separation

mechanisms which were used to separated the different types of blood cells from

each other or from the plasma to obtain e.g. depleted or enriched samples. One con-

Separating Biological Particles in Flowing Microsystems

63

Cell type / size Separation technique Process* Reference Erythrocytes Pinched flow fractionation Blood constituents 25 2 x 6 – 8 μm Filtration Sample fluid 4 (biconcave discoids) Filtration Leukocytes 35

Hydrodynamic filtration Depletion for leukocyte enrichment 45 Hydrophoretic separation Depletion for leukocyte enrichment 88 Hydrophoresis Focussing 89 Acoustophoresis Washing/extracting from medium 110, 111, 114 Acoustophoresis Separation from phospholipids and enrichment 116, 117 Magnetophoresis Separation of 1.6 μm beads from sample 143 Magnetophoresis E. Coli 143 Magnetophoresis Whole blood 144 Magnetophoresis Trapping on ferromagnetic dots 147

Leukocytes Filtration Erythrocytes 35 7 – 15 μm diameter (Cross-flow) filtration Blood constituents 36, 37, 41, 43

Cross-flow filtration Exchange of medium 40 Hydrodynamic filtration Enrichment 45 Deterministic lateral displacement Blood 60, 61 Deterministic lateral displacement Healthy from malignant 60 Hydrophoretic separation Enrichment 88 Magnetophoresis Immunomagnetic separation from whole blood 142 Magnetophoresis Whole blood 146

Lymphocytes Deterministic lateral displacement Red blood cells and platelets 62 Monocytes Deterministic lateral displacement Red blood cells and platelets 62 Plasma Cross-flow filtration Whole blood 39

Deterministic lateral displacement Whole blood 62 Acoustophoresis Whole blood 120

Platelets Deterministic lateral displacement Leukocytes 58 2-4 μm diameter Deterministic lateral displacement Immunofluorescent labelling 59

Deterministic lateral displacement Study effect of heat & thrombin 61 Dean flows Depletion and enrichment 77

Whole blood Dean flows E. Coli 86

Table 5 Overview of the different separation mechanisms used to process blood. * Cell typeswere separated from the blood components given except when mentioned otherwise.

sideration is the reduction of cluster formation, as these tend to cause channel clog-

ging and do not behave like individual particles. This was observed, for instance,

with large agglomerates of magnetic particles, which were deflected to a larger ex-

tent than individual magnetic particles.136

The different separation mechanisms inspired by laminar flows can be used as tools

for research in a wide range of applications. For example, acoustic forces have pro-

ven themselves useful for clinical applications like intra- and postoperative blood

washing111 and the separation of sperm cells from sexual assault evidence during fo-

rensic analysis.127 Deterministic lateral displacement will likely find application in

the near future for the selection of specific cells for tissue engineering.63 These de-

velopments are excellent examples of the potential of microfluidic particle separation

techniques and their incorporation into multifunctional micro total analysis systems.148

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