Continuous flow separations in microfluidic devices - ResearchGate
Transcript of Continuous flow separations in microfluidic devices - ResearchGate
Continuous flow separations in microfluidic devices
Nicole Pamme
Received 20th August 2007, Accepted 2nd October 2007
First published as an Advance Article on the web 2nd November 2007
DOI: 10.1039/b712784g
Biochemical sample mixtures are commonly separated in batch processes, such as filtration,
centrifugation, chromatography or electrophoresis. In recent years, however, many research
groups have demonstrated continuous flow separation methods in microfluidic devices. Such
separation methods are characterised by continuous injection, real-time monitoring, as well as
continuous collection, which makes them ideal for combination with upstream and downstream
applications. Importantly, in continuous flow separation the sample components are deflected
from the main direction of flow, either by means of a force field (electric, magnetic, acoustic,
optical etc.), or by intelligent positioning of obstacles in combination with laminar flow profiles.
Sample components susceptible to deflection can be spatially separated. A large variety of
methods has been reported, some of these are miniaturised versions of larger scale methods,
others are only possible in microfluidic regimes. Researchers now have a diverse toolbox to
choose from and it is likely that continuous flow methods will play an important role in future
point-of-care or in-the-field analysis devices.
Introduction
The development of microfluidic devices for analytical and
bioanalytical chemistry has, in the last two decades, led to
ground breaking advances in terms of the speed of analyses,
the resolution of separations and the automation of proce-
dures.1–3 Microfluidic chips have proven ideal tools to
precisely handle small volumes of samples, such as proteins
or DNA solutions, as well as cell suspensions.4,5 Additionally,
microfluidic devices can form part of portable systems for
point-of-care or in-the-field detection. In such micro total
analysis systems (microTAS) an entire analytical procedure
can be performed, including sample pre-treatment, labelling
reactions, separation, downstream reactions and detection.
However, to date this ideal case has seldom been realised.
One notable challenge is certainly the fact that on-chip separa-
tions are usually carried out as batch procedures, requiring the
precise injection of a very small amount of sample (nL or less)
into a separation channel for chromatography or electro-
phoresis (Fig. 1a).
An exciting development within the microfluidics commu-
nity has been the investigation of continuous flow separation
methods (Fig. 1b). This concept has gained momentum in
The University of Hull, Department of Chemistry, Cottingham Road,Hull, UK HU6 7RX. E-mail: [email protected];Fax: +44 1482 466416; Tel: +44 1482 465027
Nicole Pamme obtained aDiploma in Chemistry fromthe University of Marburg,Germany, in 1999. For herPhD she went to ImperialCollege London, where sheworked on single particleanalysis in microfluidic chipsuntil 2004. This was followedby a stay in Tsukuba, Japan asan independent research fellowin the International Centre forYoung Scientists (ICYS) atthe National Institute forMaterials Science (NIMS).In December 2005 she took
up a lectureship position at the University of Hull (UK). Hercurrent research interests include continuous flow separation inmicrofluidic devices based on magnetic forces.
Nicole Pamme
Fig. 1 (a) A typical batch separation procedure entails the injection
of a small amount of sample into a separation column and detection
after the separation has been completed, often followed by repeats
to optimise separation parameters. (b) A typical continuous flow
separation procedure: sample is injected continuously together with a
carrier liquid into a wide separation chamber, a force acts at an angle
to the direction of flow and sample components are deflected from
their flow path and thus spatially separated. Separation efficiency can
be monitored in real-time and separation parameters can be varied to
optimise conditions.
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recent years. Here, the sample is fed into a separation chamber
continuously and subjected to a force at an angle, often
perpendicular, to the direction of flow. The sample compo-
nents respond differently to this force and are thus deflected
from the direction of flow and can be collected at different
outlets. Researchers have tried a wide variety of forces to
fractionate sample components. These include the application
of electric or magnetic fields, as well as standing ultrasonic
waves and/or the intelligent design of flow obstacles. Some of
these methods are miniaturised versions of pre-existing larger
scale procedures, although with the added advantage of
favourable scaling laws. Other continuous flow separation
techniques, however, are only possible on the microfluidic
scale. In this review, the various microfluidic based continuous
flow separation methods are described and compared in terms
of performance and applicability.
Motivation for continuous flow separation
The separation of sample components is an important part of
any chemical or biochemical analysis. Common separation
processes include filtration and centrifugation and most
prominently, forms of chromatography and electrophoresis.
All these methods can be classified as batch procedures, since
in each case a particular volume of sample undergoes a
separation process and separation time is plotted versus
signal (see Fig. 1a). The efficiency is evaluated after the
separation has finished. More often than not, the separation
conditions, such as flow speed, carrier liquid composition
or instrumental parameters are not ideal and the sample
injection and separation has to be repeated several times before
optimal conditions have been found. Not only is this rather
labour and time intensive, it also poses challenges for the
combination of the separation step with upstream or down-
stream applications.
The general principle of a continuous flow separation is
shown in Fig. 1b. One of the key features is the fact that a
sample solution is introduced into the separation channel
continuously. It is not necessary to inject a small and precise
amount of sample, as in chromatography or electrophoresis. A
force acts on the sample components at an angle to the
direction of flow, often perpendicular, resulting in the sample
components taking different paths through the separation
channel and thus being spatially separated. Signal is detected
versus position. This allows for both continuous feedback for
the separation parameters and for continuous collection of
sample components.
The advantageous or characteristic features of continuous
flow separation can be summarised as follows:
(a) Continuous introduction of sample
Precise injection of small sample plugs is not required. A
relatively high throughput of sample can be achieved,
although, given the miniaturised nature of the channels,
volume throughput rarely exceeds mL min21. Another aspect
of continuous injection is that there is no capacity limit for
the separation column; in principle any volume of sample can
be processed.
(b) Continuous readout of separation efficiency
Since the separation can be monitored continuously, changes
to the separation parameters can be implemented and their
effects observed in real-time—enabling on-line feedback. Often
parameters for the fluid flow and the induced force can be
changed independently from each other.
(c) Lateral separation of sample components
The sample components are spread out orthogonally to the
direction of flow and can be continuously and independently
collected. Devices may have two outlets or multiple outlet
channels. In principle, it is possible to guide a desired sample
component for further processing downstream, whilst discard-
ing all other components. Alternatively, a sample can be
purified of undesired components and then used for down-
stream applications.
(d) Potential for integration
Since the sample is introduced continuously and the sample
components are collected continuously, separation methods
can be integrated relatively easily with upstream or down-
stream analysis steps. In other words, separation can be
performed ‘‘in-line’’ with other continuous flow processes,
such as continuous flow reactions. There is no large increase in
resistance due to a filter.
(e) Label-free
Separation is often based on intrinsic properties of the sample
components, such as their size, charge or their polarisability.
In such cases, no labelling of sample components is required.
In the literature, the term ‘‘continuous flow separation’’ is
sometimes used in different contexts. ‘‘Separation’’ often
appears in the context of ‘‘trapping’’ or ‘‘holding’’ a sample
component in a fixed position against a continuing fluid flow.
For example, magnetic particles can be held within a channel
by means of an external magnet, or a sample component can
become immobilised within a channel packing. Although these
methods are carried out under a continuous flow of carrier
fluid, the sample components themselves are not collected
continuously, and hence these methods are not reviewed here.
The term ‘‘continuous flow separation’’ is also sometimes used
for liquid–liquid phase extraction6,7 or for two-phase parti-
tioning.8,9 In these methods the separation is not based on a
specific force applied at an angle to the direction of flow.
Instead, sample components diffuse from one phase to the
other and remain in their preferred phase. Such phase
extraction methods are also not covered here.
A list of many of the on-chip continuous flow separation
methods that are covered in this review is featured in Table 1.
The large variety of forces that can be applied becomes clear
immediately. Some methods are very simple, others require
more sophisticated equipment. Examples of more simple
methods include pinched flow injection, hydrodynamic filtra-
tion or bifurcation, which make use of laminar flow behaviour
in microfluidic channels. Microfabricated obstacles, such as
pillars or dams, have been applied for controlled size
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separations, sometimes in connection with electric fields,
such as in the ‘‘DNA prism’’. Free-flow separation methods,
such as electrophoresis, isoelectric focussing, magnetophoresis,
dielectrophoresis and acoustophoresis, are based on homo-
geneous electric fields, pH gradients, inhomogeneous magnetic
fields, inhomogeneous electric fields or acoustic standing
waves, respectively. Other examples of applied forces include
gravity and optical pressure. The force utilised dictates the
type of sample that can be separated. The different separation
methods are discussed in detail below, arranged by type of
separation force.
Channel design and flow rates
Flow regimes in microfluidic devices are predominantly
laminar. With an appropriate microfluidic design and control
of hydrodynamic flow, it is possible to force microparticles
or cells into specific flow stream lines and to separate
these objects.
Pinched flow fractionation
Pinched flow fractionation (Fig. 2), developed by Seki and
Yamada,10 allows for separation of particles and cells based on
their size. Two laminar streams are pumped through a narrow
channel section before entering a widening channel. The
sample stream contains a particle suspension; the carrier
stream contains a buffer solution. The flow rate of the carrier
stream is higher than that of the sample stream, and hence the
microparticles are pushed against the wall in the pinched
channel segment. Small particles find themselves in a flow
stream close to the wall; larger particles have their centre of
gravity further away from the wall, it might even be outside the
original sample stream and instead within the carrier stream.
Table 1 Listing of continuous flow separation methods with details of the forces utilised, the basis of separation and the applicable samples. Thedata for flow rates and throughput were taken from the selected reference, if provided
Method Separation induced by Separation based on Sample Flow rate/throughput Ref.
Pinched flow fractionation Laminar flow regime Size Microparticles, cells 70 cm s21 1140 000 particles h21
Hydrodynamic filtration Laminar flow regime Size Microparticles 1 mm s21 14100 000s particles h21
Bifucation Laminar flow Size, shape Blood 5 mL h21 161–3 mm s21
Filtration obstacles Diffusion and obstacles Size Blood 5–12 mL min21 1910–20 mm s21
Lateral displacement Obstacle array Size Particles (0.8, 0.9, 1.0 mm) 20 mm s21 (DNA) 24DNA (61 kbp, 158 kbp) 100 pg h21 (DNA)
Brownian ratchet Diffusion and obstacles Size DNA 2 mm s21 2610 pg h21
Hydrophoreticseparation
Pressure gradient fromslanted obstacles
Size Microparticles 1 to 3 cm s21 3010 000s beads h21
DNA prism Obstacle array andelectric field
Size DNA 10 ng h21 31
Entropic trap array Dams and electric field Size DNA 10 pg h21 32
Repulsion array Dams, electric double layers Charge to size ratio Proteins 300 pg h21 32
Free-flow electrophoresis Homogeneous electric field Charge to size ratio Proteins, amino acids 6 mm s21 44
Free-flow isoelectricfocussing
pH gradient Isoelectric point Proteins and cells 9 mm s21 45
Magnetophoresis Inhomogeneousmagnetic field
Size, magnetisation Magnetic particles, cells up to 1.2 mm s21 8110 000s cells h21
Dielectrophoresis Inhomogeneous electric field Size, polarisability Microparticles, cells up to 4 mm s21 63100 000s cells h21
Acoustophoresis Acoustic pressure Size, density,compressibility
Microparticles, cells 11 cm s21 90
Optical lattice Optical force Size, refractive index Microparticles, cells 30 mm s21 9290 000 particles h21
Sedimentation Gravity Density, size Microparticles, droplets 6 mm s21 slowedto 1 mm s21
94
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When the fluid enters the widening channel, the laminar flow
carries the two particle populations along different stream
lines. The particles are thus separated across the width of
the wide channel. If required, they can be collected further
downstream.
Initially, the separation of 15 mm and 30 mm polymer
particles was demonstrated through a pinched segment of
about 50 mm width.10 Subsequently, the method was modified
to improve performance. Outlet channels were added and
negative pressure was applied to one outlet.11 Most of the
carrier liquid was forced to leave via this outlet, whereas the
remaining flow streams could fan out more over the remaining
exits. This enabled the separation of 1 mm and 2 mm particles.
This device was also applied to the focussing of red blood cells.
In another design, flow into the outlet channels was controlled
by addition of a valve,12 which enabled the separation of 1 mm,
2 mm and 3 mm diameter particles from each other, as well as
the separation of 0.5 mm and 0.8 mm particles.
A similar concept was also published by Zhang et al.;13 their
design featured three inlet channels leading into a bent and
asymmetrically widening separation channel. By varying the
ratio of sample and buffer flow rates, microparticles could be
forced onto a specific flow stream. Any differences in initial
alignment were then enhanced by the flow profile within the
bent channel. 10 mm and 25 mm particles were separated from
each other.
Pinched flow fractionation is based on laminar flow and
channel geometry, and is suitable for the separation of
polymer particles of micrometre and sub-micrometre12
diameters as well as for biological cells.11 The separation is
based on particle size, no labelling is required and throughput
(see Table 1) can be 10 000s particles per hour. The dimensions
of the pinched segment have implications for the applicable
size range, and a significant amount of fine tuning has been
necessary for the separation of sub-micron particles
Hydrodynamic filtration
Hydrodynamic filtration is based on channel design and flow
control and can be utilised for the sorting of micrometre sized
particles.14,15 The microfluidic design features a main channel
with multiple narrow channels branching to the sides (Fig. 3).
Fluid is pumped along the main channel, so that some liquid
exits via the branch outlets. The amount of liquid leaving the
main channel depends on the flow resistance of these side
channels. When a mixture of particles is pushed through the
main channel it is important to note that smaller particles can
occupy flow streams much closer to the channel wall than
larger particles, simply because the centre of a particle cannot
come closer to the wall than its radius. In Fig. 3a, a situation is
shown where none of the particles can reach a flow stream
close enough to the wall to enter a side channel. Due to the
constant removal of liquid into the side channels, all particles
eventually become aligned against the wall of the main
channel, but they cannot enter the side channels. In Fig. 3b,
the flow resistance of the side channels is reduced and more
fluid can enter. Now, the centre of a small particle can be
positioned within a stream line that enters the side channel,
whereas the centre of a larger particle is still too far away from
Fig. 2 The principle of pinched flow fractionation for the separation
of microparticles or cells, based on laminar flow behaviour and
negligible diffusion of the particles. A buffer stream and a particle
carrying stream are forced through a narrow, pinched channel segment
before entering a widening channel. Due to the higher flow rate of the
buffer stream, the particles are pushed against the wall of the pinched
segment and are thus transported into specific laminar flow streams
based on their size. This difference is enhanced as the particles enter
the widening channel. Redrawn with permission from ref. 10,
copyright 2004, American Chemical Society.
Fig. 3 The principle of hydrodynamic filtration: (a) only a small
portion of flow enters the branching channel and neither large nor
small particles can be in a flow stream close enough to the wall of the
main channel to enter the side channel. However, after passing a series
of such branching channels, more and more fluid is removed from the
main channel and the particles are gradually aligned against the main
channel wall. (b) With more flow entering the side channels, the small
particles can now be close enough to the wall of the main channel to be
in a flow stream that enters the side channel. Large particles cannot
enter the branch channels. However, after passing several branch
points they become aligned against the main channel wall. (c) Finally,
if a relatively large portion of the flow enters the side channels, both
small and large particles can enter the side channels. (d) Design of
hydrodynamic filtration channel network for enrichment of small and
large particles. Redrawn with permission from ref. 14, copyright 2005,
Royal Society of Chemistry.
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the wall and hence larger particles are forced to flow straight
on. Any particles that are initially far away from the channel
wall will be brought closer to the wall each time they pass an
outlet branch, until eventually they are aligned along the wall.
Finally, if the flow into the side channels is increased further
(Fig. 3c), both large and small particles can enter the branches.
Using such a channel network, particles of different sizes can
be aligned against the main channel wall and collected into
different side channels (Fig. 3d).
In a first demonstration,14 a suspension of polymer particles
of 1 mm and 3 mm diameter was pumped through a 20 mm wide
and 7 mm deep channel featuring 50 side channels at a flow rate
of 1 mL min21. The particles were collected in different outlets
and could be concentrated 20-fold and 50-fold. Different
channel network designs were utilised for the concentration
of 1 mm and 2 mm particles and for leukocyte enrichment,
respectively. The design was further improved by splitting and
re-combining flows15 for more efficient aligning of particles
against the main channel wall. Particles of 1 mm, 2 mm and 3 mm
diameter were collected into different outlet reservoirs and
their concentrations were increased 60-fold to 80-fold.
Although this method does not allow for 100% separation
efficiency, it is a very effective way of enriching particle
concentrations according to size. Flow rates are fast (ca.
10 cm s21), and throughput can be 100 000s particles per hour.
The separation is determined purely by channel geometry,
minimal tuning is required after microfabrication and the
devices are thus simple to operate.
Hydrodynamic blood separation
Blood circulation in microcapillaries shows an intriguing
behaviour which can be utilised in microfluidic devices for
blood separation in continuous flow. Red blood cells have a
tendency to flow in the centre of the stream, blood at the edge
of the channel is thus depleted in erythrocytes and this can be
withdrawn continuously to obtain plasma with a low red blood
cell concentration (hematocrit).16 When blood encounters a
channel junction, the red blood cells have a tendency to flow
into the branch with the faster flow rate; this channel will thus
exhibit a higher red blood cell concentration, whereas the
channel with the lower flow rate will feature a lower hematocrit.
This phenomenon is sometimes referred to as ‘‘bifurcation law’’
or the ‘‘Zweifach–Fung effect’’ and has been demonstrated for
plasma separation in continuous flow.17 Another phenomenon
is referred to as ‘‘margination’’. Collisions between red blood
cells that are predominantly found in the microchannel centre
and white blood cells lead to the white blood cells rolling along
the channel wall. This process has been used for leukocyte
enrichment in continuous flow.18
Continuous flow filtration
Filtration usually means that a sample is pumped through a
membrane or a pillar array. This method is prone to clogging.
An alternative is to localise filter components, such as pillars
or posts, in-line with the direction of flow. If a particle
suspension is pumped through a channel with pillars on either
side, sample components that are sufficiently small can pass
through the gaps between the pillars into a neighbouring
receiving stream, whereas larger sample components remain in
the original sample stream. Such a diffusive filter is not prone
to clogging, since the larger size particles can continue to flow
through the device in a direction parallel to the filter.
Diffusive filters have been employed for differentiation
between red and white blood cells.19,20 Leukocytes are
spherical (6–10 mm diameter), erythrocytes are biconcave with
a diameter of about 8 mm and a thickness of about 2 mm. By
fabricating posts with appropriate gaps on either side of a
blood carrying channel, the red blood cells passed through the
gap and entered a parallel channel beyond the posts, whereas
the white blood cells remained in the central channel. Such
separations have been carried out at flow rates of 5 to
10 mL min21, equalling speeds of about 10 to 20 mm s21.
Continuous flow filtration has also been demonstrated with
an asymmetric pillar array at a channel junction featuring
narrow gaps (32 nm) for flow along the straight main channel,
but wider gaps for flow into a channel branching off at 45u.This has been applied to the separation of DNA molecules of
300 kbp (57 nm radius of gyration) and 100 kbp (22 nm
radius).21 Again, this type of filter is less prone to clogging
than a head-on filter, since larger components can flow away
into the side channel.
Another option of continuous flow filtration has been the
application of a cross-flow through shallow (3 mm) channels
running perpendicular to a 30 mm deep sample carrying
channel.22 The orthogonal flow led to continuous exchange of
medium in the main channel, which was demonstrated for a
suspension of 10 mm polystyrene beads. The device was also
tested for blood; red blood cells could pass into the side
channels and their concentration in the main channel was
reduced by a factor of 4000, whereas almost all white blood
cells were retained. A cross flow device with even shallower
side channels of 0.5 mm depth was used to obtain plasma from
blood, albeit only at 8% efficiency.23
Movement around obstacles
Obstacles can not only be used as filters to prevent or block
access. Flow around obstacles can also be used in clever ways
to separate particles or DNA molecules.
One such method has been termed ‘‘deterministic lateral
displacement’’ (Fig. 4)24,25 and has been applied to the size
separation of particles and DNA molecules by pumping
through an array of obstacles in the y-direction. The array
consisted of rows of posts with a consistent gap between the
posts in each row. The posts, however, were slightly shifted
from one row to the next, until after a few rows they aligned
again; in the example in Fig. 4, this is after three rows. Liquid
pumped through this array flowed around the posts in laminar
streams, as indicated by the different grey shades in Fig. 4. If
the liquid contained particles that were small in comparison to
the gap between the posts, then the particles followed the
direction of laminar flow. After three rows of obstacles the
particles were in the same position in the x-direction as in
the first row. If however, particles with a diameter only a little
smaller than the gap were pumped through the array of posts,
the particles had to position themselves in the middle of the
gap. The particle’s centre of gravity was thus pushed out of the
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original flow stream into the neighbouring flow stream. After
three rows of obstacles the position of the particle had been
shifted in the x-direction, i.e. it had been laterally displaced.
With a gap between posts of 1.6 mm it was possible to
separate fluorescent particles with diameters of 0.6 mm, 0.8 mm
and 1.0 mm as well as particles 0.7 mm and 0.9 mm in diameter.24
Pressure driven flow was applied at a rate of 400 mm s21 and
the separation was detected after a running time of 40 s. By
gradually increasing the gap size from 1.4 mm to 2.2 mm
along the array, particles with diameters of 0.8 mm, 0.9 mm and
1.0 mm were separated with high resolution. The same tech-
nique was also used to separate DNA molecules of 61 kbp
and 158 kbp. The DNA containing solution was pumped via
EOF at about 20 mm s21 through an array with 3 mm gaps. The
two sizes of DNA molecules were separated from each other
within minutes.
The performance of this displacement method is certainly
impressive in terms of speed and particle size resolution. The
separation is based on inherent properties of the particles or
DNA molecules, no labelling is required. The path through the
obstacle array is pre-determined by the geometry of the array.
Each particle can only take one possible path. There is no
averaging of flow paths. Only diffusion could throw the
particle off-path. Hence, resolution of the separation is better
at faster flow rates, which makes this separation method
conceptually different from most other separation methods.
Another separation method based on an asymmetric array
of obstacles, a so-called ‘‘Brownian ratchet’’, has been
employed for the continuous flow separation of DNA
molecules based on their diffusion coefficients (Fig. 5).26–29
The DNA molecules were sufficiently small to pass through
the gaps in the array. They were pumped at relatively slow flow
rates, so that they had some time to diffuse before encounter-
ing the next row of obstacles. A large molecule (low diffusion
coefficient) would on average diffuse a much shorter distance
than a small molecule (high diffusion coefficient). Hence,
the larger DNA molecules followed the laminar flow in the
y-direction and passed almost straight through the array (solid
line). A smaller DNA molecule, however, could diffuse further
and could indeed travel far enough in the x-direction to be
forced around an obstacle (dotted line). Smaller DNA
molecules were thus likely to take a flow path shifted in the
positive x-direction and after a number of obstacle rows the
two populations of molecules would be separated. Of course,
diffusion also occurred in the negative x-direction. However,
the obstacle array was designed such, that the molecules would
have to travel a much larger distance to be displaced in this
direction (see grey cone in Fig. 5). The concept was initially
demonstrated by showing different flow paths taken by 15 kbp
and 33.5 kbp DNA.27 After improving the sample injection
method, the separation of lambda DNA (49 kbp) from T2
DNA (167 kbp) was achieved,28 as well as the separation of T2
and T7 DNA.29
Separation through a Brownian ratchet is a quite elegant
method, molecules are differentiated based on their diffusion
coefficient, which is directly related to their size. No labelling is
required. Once the obstacle array has been designed and an
optimum flow rate has been identified, the operation is very
simple. Fine tuning of the array design and flow rate for the
diffusion coefficients of the sample components however are
not trivial and a particular array design might only be suitable
for a limited range of DNA sizes. The method has two
drawbacks in comparison to lateral displacement separation
(Fig. 4). (i) Flow rates for the Brownian ratchet must be very
slow (few mm s21) to give molecules enough time to diffuse.
With an array length of 12 mm, the separation can take more
than an hour to complete,28 whereas lateral displacement
separations were completed within a few minutes. (ii) The flow
path through the Brownian ratchet is based on probability,
and it is not precisely pre-determined. Molecules within a
Fig. 5 The principle of diffusion based separation through a
Brownian ratchet. Molecules are pumped slowly in the y-direction
through an asymmetric array of obstacles. The gaps are sufficiently
large to let all molecules pass. Due to the slow flow rate, the molecules
are given some time for diffusion before encountering the next row of
obstacles. Whilst a molecule with a large diffusion coefficient might
migrate far enough to move around the obstacle (dotted line), a larger
molecule with low diffusion coefficient will be largely following the
laminar flow in y-direction (solid line). It should be noted that the
obstacles are arranged asymmetrically, the distance around an obstacle
in the negative x-direction is much larger than in the positive
x-direction (see grey cone), and hence there is a very low probability
for a molecule to take a path in the negative x-direction. Left part of
figure redrawn with permission from ref. 26, copyright 1998 by the
American Chemical Society; right part of figure redrawn with
permission from ref. 29, copyright Wiley VCH Verlag.
Fig. 4 The principle of particle separation via lateral displacement
around obstacles. Small particles follow the laminar flow streams
(shaded in grey) in the y-direction, whereas large particles are
continuously forced to change the laminar flow stream and are thus
continuously displaced in the x-direction. Redrawn from ref. 24, with
permission of AAAS.
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population will take a range of paths, leading to relatively wide
bands. The flow path in the deterministic lateral displacement
experiments is pre-determined and bands are narrow.
Recently, ‘‘hydrophoretic separation’’ of 9 mm and 12 mm
particles was demonstrated by implementing slanted obstacles
at the bottom and top of a microfluidic channel.30 These
obstacles generated local flow perturbations, in the form of
helical re-circulations, perpendicular to the main direction of
flow. The obstacles were arranged such that initially particles
were focussed at the bottom of the microchannel and
subsequently deviated from their direction of flow depending
on the size difference between the particle and the obstacle
gap. Large particles were found to flow in the centre of the
channel, whereas smaller particles were directed onto an
oscillating path.
Similar to the other obstacle based separation methods
mentioned above, hydrophoretic separation does not require
any labelling of sample components and the devices are
relatively easy to fabricate and to operate. Remarkably, the
sorting performance is independent of the applied flow rate
in the range of 0.7 to 3.3 cm s21. Only the geometry of the
obstacles determines the separation efficiency. This, however,
poses limitations for the range of particle sizes. For example,
with a gap of 25 mm, sorting was limited to particles between
8 mm and 12 mm in diameter.
Separations involving electric fields
Electric fields have long been popular for microfluidic
separations. Due to the large surface to volume ratios, Joule
heating can be dissipated rapidly, allowing for the application
of high electric fields and thus efficient separation of amino
acids, DNA molecules, proteins and peptides. It is not
surprising then, that electric fields have been used extensively
for continuous flow separation. For the purpose of this
review, these methods have been divided into four groups: (i)
separation involving electric fields in combination with
obstacles, (ii) free-flow electrophoresis based on charge to size
ratio, (iii) free-flow isoelectric focussing based on isoelectric
point (pI) and (iv) dielectrophoretic separation based on
polarisability.
Electric fields and obstacles
Arrays of obstacles can be combined with electric fields for the
separation of charged molecules, such as DNA or proteins. A
‘‘DNA prism’’ was developed by Huang et al.,31 consisting of a
regular array of posts over which an electric field could be
applied parallel or diagonally to the posts. By continuously
switching between two fields, E1 and E2, the separation of
DNA molecules based on their molecular mass could be
achieved as follows (Fig. 6): at the time t0 a relatively large
electric field E1 was applied diagonally to the pillars and both
DNA molecules experienced a strong electrophoretic force in
direction of the field. Then at t1 the field direction was
switched and a smaller electric field E2 was applied parallel to
the array in the x-direction. The DNA molecules experienced a
relatively weak electrophoretic force in direction of E2. The
short DNA molecule could pass through the obstacle array
easily. The larger DNA molecule however, became entangled
in the pillars. Before this long DNA molecule could
disentangle itself, the electric field was switched back to E1
and both molecules were pulled along diagonally again. The
long DNA molecule thus moved back onto its original path,
the short DNA molecule however, had now been displaced
in the x-direction, or, in other words: long DNA molecules
moved along the direction of E1 and short DNA molecules
moved along the average direction of the two electric field
vectors. Using an array of 2 mm posts with 2 mm gaps, Huang
et al. demonstrated the separation of 61 kbp, 114 kbp, 158 kbp
and 209 kpb DNA within 15 s. Applied electric fields were of
the order of 20 to 250 V cm21 for pulse durations of a few tens
of milliseconds. This separation method is rather fast and
robust and the authors report a throughput of 104 molecules
per second or 10 ng of DNA per hour.
Obstacles do not have to be in the form of pillars. Shallow
gaps can also be used very effectively as demonstrated by
Fu et al.,32 who fabricated arrays of deep parallel channels
separated by shallow nanometre gaps. Orthogonal electric
fields were applied simultaneously over this array, pulling
molecules in the x- as well as in the y-direction. Movement
in the x-direction however, was hindered by the nanogaps.
Depending on the size of the gap, the thickness of the electric
double layer (Debye layer), and the size of the sample
components, three separation regimes could be differentiated:
Ogston sieving, entropic trapping and electrostatic sieving
(Fig. 7).
An array consisting of deep trenches in the y-direction (1 mm
wide, 300 nm deep), and shallow passages in the x-direction
(1 mm wide but only 55 nm deep) was used to demonstrate
the three regimes.32 The thickness of the Debye layer was
controlled via the buffer concentration: at 130 mM the layer
Fig. 6 Separation in a DNA prism, consisting of a regular array of
posts. A large electric field is applied diagonally (E1), a small electric
field (E2) is applied horizontally. By switching between the two fields,
DNA molecules of different length are separated. Depicted are a long
(light grey) and short (dark grey) strand of DNA. When E1 is applied
at the time t0, both molecules experience a strong electrophoretic force
and migrate along the field direction. When the smaller electric field E2
is applied at t1, the molecules experience a weaker electrophoretic force
in the direction of E2. Whilst the small DNA molecule can pass
through the post array easily, the long DNA molecule becomes
entangled. When switching back to E1 at t2, the large DNA molecule is
pulled back into its original path, whereas the short DNA molecule has
been displaced in the x-direction. Figure redrawn with permission from
Macmillan Publishers Ltd.: Nature Biotechnology,31 copyright 2002.
1650 | Lab Chip, 2007, 7, 1644–1659 This journal is � The Royal Society of Chemistry 2007
was 0.8 nm thick, whereas at 1.3 mM the thickness increased
to 8.4 nm.
Ogston sieving is based on steric hindrance. This regime
occurs when the double layer is thin in comparison to the gap
and the sample components are small enough to fit easily
through the gap. DNA strands of 50 bp to 766 bp, corres-
ponding to a strand length of 16 nm to 150 nm, were
suspended in a 130 mM buffer. When introduced into the
array, the larger DNA strand would only fit through the gaps
if correctly aligned. The smaller DNA stands, however, could
pass the gaps easily and would on average travel further
into the x-direction. By applying 35 V cm21 in the x-direction
and 75 V cm21 in the y-direction, the five different strand
lengths were separated.
Entropic trapping had previously been shown as a batch
procedure in a straight channel33 and has now been adapted to
continuous flow. For this regime, the DNA molecules were
considerably larger than the gap, ranging from 2.3 kbp to
23 kbp, corresponding to radii of gyration between 140 nm
and 520 nm. The electric field in the x-direction pulled the
DNA molecules into the trap, they became stuck for some
time, until, eventually, a long enough part of the strand, a
hernia, had ventured into the gap and the entire molecule
could get drawn through the narrow slit. Interestingly, the
larger DNA molecules passed through the gaps more easily.
This was explained by that fact that a large molecule is in
contact with a large surface area of the gap region and hence
there is a higher probability for hernia formation and
thus being pulled through the gap. Electric fields of Ex =
185 V cm21 and Ey = 100 V cm21 were used to separate the
five DNA sizes in less than 1 min. DNA throughput was about
130 pg h21.
Electrostatic repulsion is dominant if the Debye layer is of
considerable thickness with respect to the gap size. This was
demonstrated for two proteins of similar size but different
charge: lectin (49 kDa, pI 8–9) and streptavidin (53 kDa,
pI 5–6). Using a 1.3 mM buffer (pH 9.4), the more negatively
charged streptavidin molecules were repelled from the gap to a
larger extent than the less charged lectin molecules. Hence,
lectin was deflected further into the x-direction and the two
proteins could be separated within tens of seconds.
The nanogap array for continuous flow separation has only
been published recently and is still in its infancy. A striking
advantage is the versatility of the method, the same chip could
be used for a variety of samples and separation regimes. The
separation speed could be tuned by two independent para-
meters, Ex and Ey. Separations could be performed within
minutes, with throughputs of 100s pg h21.
Free-flow electrophoresis and free-flow isoelectric focussing
Free-flow electrophoresis (FFE) is performed in a shallow
chamber, as shown in Fig. 8a. Buffer and sample solutions are
continuously pumped into this chamber through numerous
inlet channels and collected via outlet channels. Perpendicular
to the direction of flow, a homogeneous electric field is
applied. Charged molecules are thus subjected to two flow
vectors: the hydrodynamic flow in the y-direction and
the electrophoretically induced flow in the x-direction. The
observed flow path is the sum of these two vectors. The
direction of deflection depends on whether the molecule is an
anion or a cation; the extent of deflection depends on the
charge to size ratio. Molecules with zero net charge are not
deflected and flow straight through the chamber. Free-flow
electrophoresis was initially developed on the larger scale.34,35
Miniaturisation of the separation chamber, typically to a width
and length of several mm and a depth of a few mm, brings
similar advantages to those for capillary electrophoresis: Joule
heating can dissipate quickly due to the larger surface to
volume ratio, hence, high electric fields can be applied, and
thus separation efficiency and speed are improved.
Fig. 7 Anisotropic sieving through shallow nanometre sized gaps.
Depending on the size of the gap and the thickness of the Debye layer,
three separation regimes can be distinguished. (a) In Ogston sieving,
the electric double layer is thin compared to the gap and sample
molecules are also smaller than the gap. Passage through the gap is
based on steric hindrance, the smaller molecules (light grey) pass more
easily than the larger molecules (dark grey). (b) Entropic trapping: if
the sample contains DNA molecules that are larger than the gap, the
DNA becomes trapped at the gap, but eventually a sufficient length of
DNA will have been pulled into the gap by the electric field, that the
entire molecule squeezes itself through the gap. Counter-intuitively,
larger DNA molecules move through the gap more easily. (c)
Electrostatic sieving: when the Debye layer is thick in comparison to
the gap size, electrostatic forces become dominant. A molecule of
higher negative charge is repelled more by the gap than a molecule of
low negative charge. Figure redrawn with permission from Macmillan
Publishers Ltd.: Nature Nanotechnology,32 copyright 2007.
This journal is � The Royal Society of Chemistry 2007 Lab Chip, 2007, 7, 1644–1659 | 1651
Free-flow isoelectric focussing (FF-IEF) is based on a
similar setup to free-flow electrophoresis, and hence many
researchers have demonstrated both applications with the
same device layout. The principle of FF-IEF is shown in
Fig. 8b. A pH gradient is generated in the x-direction, fluid is
pumped in the y-direction. A sample mixture, containing
amphoteric molecules, such as proteins, is introduced over the
entire width of the separation chamber. A sample protein
entering at a location where the pH is different from its
isoelectric point (pI) exhibits a net charge, and thus experiences
an electrophoretic force. The protein migrates along the
x-direction until it reaches a point in the pH gradient where
the pH equals its pI. The protein no longer has a net charge
and hence ceases to migrate in the x-direction. The protein is
focussed at its isoelectric point and now only follows the
direction of flow. FF-IEF allows for the focussing and
concentration of proteins based on their pI. Hence, proteins
with different pIs can be separated from each other.
On-chip free-flow electrophoresis and isoelectric focussing
were first demonstrated in the late 1990s and there have been a
relatively large number of publications since then. Differences
in performance result from the chip material and most notably
from the connection between the electrode reservoirs and the
separation chamber.
The first free-flow electrophoretic devices were fabricated in
silicon.36,37 However, the maximum electric field that could be
applied to this material was 100 V cm21. Later, glass devices38–41
and hybrid devices made from glass and silicon42,43 or glass
and PDMS44,45 were introduced, in which higher electric fields
could be applied, typically between 200 V cm21 and 300 V cm21.
In one case43 the field was as high as 580 V cm21.
The most challenging aspect of free-flow electrophoresis
concerns the generation of a high electric field across the
separation chamber. This must be undertaken whilst minimis-
ing fluid exchange between the chamber and electrode
reservoirs and also minimising the detrimental effect of the
products of electrolysis (local pH changes, bubble formation).
Researchers who opted for direct contact of the electrodes with
the separation chamber fluid were forced to limit their electric
fields to a few tens of V cm21.46–48 Most researchers chose to
separate the electrode reservoirs from the separation chamber
by some form of membrane. The disadvantage here is the
voltage drop across the membrane; only a fraction of the
applied field is present over the chamber itself. A simple way of
introducing a membrane is to connect the electrode reservoir
with the separation chamber via a number of microfluidic
channels. With small cross section channels, as used by Zhang
et al.44 and Xu et al.,45 only 4.5% of the applied electric field
was transferred across the chamber. The large cross section
channels, as reported by Fonslow et al.,42 would transfer as
much as 50% of the applied field, however, at the cost of a
facilitated fluid exchange. By removing connection channels
altogether and introducing trenches with high speed laminar
flow along the electrodes in the y-direction, as much as 91%
of the applied field could be transferred to the separation
section.43 Membranes made from UV-polymerised acryl
amide,41 nitrocellulose,38 or agarose gel49 have also been
applied successfully. Perhaps the most elegant way to separate
the electrode reservoir from the separation chamber is via a
dielectric barrier, such as a thin wall of glass, as demonstrated
by Janasek et al.:50 50% of the applied field could be
transferred across the chamber and liquid interchange could
be totally avoided.
Isoelectric focussing requires the generation of a pH
gradient across the separation chamber. This is conventionally
achieved by a mixture of ampholytes that establish the
gradient as soon as an electric field is applied. The challenge
again is that the electrodes are usually positioned outside the
separation chamber. The pH gradient should ideally not
extend beyond the separation chamber. Unfortunately, this is
sometimes unavoidable, for example when channels are used
to connect the electrode reservoirs and chamber. Xu et al.45
used a wide pH gradient (pH 3 to 10), of which only a fraction
would have been over the separation chamber, in order to
focus two similarly charged proteins, angiotensin I (pI 6.69)
and angiotensin II (pI 7.75). Kohlheyer et al.41 made more
effective use of a wider pH gradient and were able to focus four
proteins with quite different pIs (4.5, 5.5, 7.6 and 8.7). A
totally different approach was taken by Cabrera and Yager.47
Having the electrodes in direct contact with the separation
medium led to electrolysis and local pH changes in the vicinity
of the electrodes. These pH changes resulted in a ‘‘natural’’ pH
gradient that could be used for isoelectric focussing. Song et al.
took yet another approach;51 they surrounded their sample
containing stream (pH 6) by two buffer streams at pH 8.
Differences in ionic strength and pH between the streams led
to a diffusion based pH gradient, which again was used for
isoelectric focussing.
FFE and FF-IEF have been applied to a variety of samples.
Free-flow electrophoresis is suitable for the separation of
molecules with different charge to size ratios. For example,
fluorescent dyes, such as fluorescein or rhodamine 110,
Fig. 8 Free-flow separation methods usually feature a shallow
separation chamber with flow pumped in the y-direction and a field
applied in the x-direction. (a) In free-flow electrophoresis, a homo-
geneous electric field is applied over the chamber orthogonal to the
direction of flow. A sample mixture is fed into the chamber through a
narrow channel, surrounded by carrier liquid. Charged sample
components are deflected from the direction of laminar flow depending
on their charge to size ratio, whereas neutral components follow the
direction of flow. (b) In free-flow isoelectric focussing, a pH gradient is
generated over the separation chamber. Amphoteric samples, such as
proteins, are injected over the entire width of the chamber. They
migrate until they exhibit zero net charge, i.e. until they have reached
the pH that equals their isoelectric point.
1652 | Lab Chip, 2007, 7, 1644–1659 This journal is � The Royal Society of Chemistry 2007
fluorescently labelled amino acids or proteins have been
separated. The throughput depends on the size of the
separation chamber and the strength of the applied electric
field. Reported separation times vary greatly, ranging from
sub-second,44,45 to several seconds,39,41,42 to minutes.36–38
Free-flow isoelectric focussing is applicable to components
with different pIs. Samples reported in the literature include
the separation of fluorescent IEF marker molecules,41
proteins,45,46,51 but also mitochondria48 and bacteria.47 Due
to their large size, the separation of the cells took minutes.
Reported focussing times for smaller sample components,
however, vary significantly, again ranging from sub-second,45
to several seconds,41 to minutes.46
Another variation of free-flow separation based on electric
fields is free-flow isotachophoresis, as demonstrated by
Janasek et al.50 It should also be noted that the term ‘‘free-
flow’’ has been adopted for many other separation methods,
for example those based on magnetic gradients,52 acoustic
pressure53 or, recently, thermal gradients.54
Continuous flow dielectrophoretic separations
Dielectrophoresis (DEP) occurs in inhomogeneous electric
fields and can be utilised for the sorting of particles or cells.
When a particle is subjected to an electric field, charges are
induced and the particle becomes polarised along the direction
of the electric field. In a homogeneous electric field, the
electrostatic forces on opposite ends of the dipole are equal
and there is no net movement. In an inhomogeneous electric
field (Fig. 9), the forces on either side are different and the
net force results in movement of the particle. The movement
can be towards or away from the high intensity field. If
the electrical polarisability of the particle is larger than the
polarisability of the surrounding buffer medium, then the
particle moves into the electric field, this is called positive
dielectrophoresis (pDEP). If, on the other hand, the polarisa-
bility of the particle is lower than that of the surrounding
medium, the particle is forced away from the electric field
towards a low intensity region. This is referred to as negative
DEP (nDEP). The dielectrophoretic force thus depends on the
difference in polarisability between particle and surrounding
medium. The force is also proportional to the volume of the
particle, as well as to the gradient of the electric field. DEP
separation can thus be based on differences in polarisability or
size or a combination of the two.
One method used to generate an inhomogeneous electric
field in a microfluidic channel has been to apply a voltage
along a channel that contained obstacles, such as ridges,55
narrowing wedges56,57 or an array of posts.58 In the vicinity of
these obstacles, the applied electric field became non-uniform
and thus dielectrophoretic forces could be generated on
particles or cells that flowed along the edges of these obstacles.
Pumping of particle suspensions was achieved via electro-
osmotic flow and particle concentrations ranged from 105 to
107 particles mL21.
With ridges restricting access to part of the microchannel,55
bacterial cells could be continuously enriched from a mixture
of 200 nm polymer particles. When an electric field of at least
3 V cm21 was applied over the channel network, the bacteria
cells were prevented from crossing the 5 mm ridge gaps due to
nDEP, while the 200 nm diameter polymer particles could
pass. In another example, a 300 mm wide channel was
narrowed to 60 mm by a rectangular obstacle.56 Polymeric
microparticles experienced nDEP when flowing through this
narrowed channel section, the DEP forces depending on
their size and the applied voltage. The authors demonstrated
controlled sorting of 5 mm, 10 mm and 15 mm diameter particles
into two downstream exits. Using the same principle, an oil
droplet can also be used as a narrowing obstacle.57 Since the
size of the droplet can be altered, the DEP forces can be tuned.
This has been utilised for the separation of a mixture of 1 mm
and 5.7 mm polymer particles, as well as a mixture 5.7 mm and
15.7 mm particles. An array of posts was used for dielectro-
phoretic focussing of 200 nm latex particles.58 The posts were
diamond shaped, with a 36 mm edge, leaving 30 mm wide and
7 mm deep channels. When an electric field of 8 V cm21 was
applied, the latex particles experienced nDEP in the vicinity of
the posts and were found to concentrate in streamlines. The
shape of the posts and orientation of the array were found to
be crucial to achieve this DEP based focussing. Very recently, a
device for three-dimensional continuous dielectrophoresis was
theoretically evaluated and demonstrated for the separation of
2 mm and 3 mm particles.59
Inhomogeneous electric fields for DEP can also be generated
by microelectrodes embedded in the microfluidic channel.
Such microelectrodes could be a single stripe or many parallel
stripes. When switched on, they generate strong gradients,
albeit usually only over a limited volume. Often electrodes are
integrated above and below a microchannel to maximise the
DEP force. When particles are pumped over the electrodes, the
non-uniform electric field induces a dielectrophoretic force
orthogonal to the direction of flow. This can be used for the
sorting of particles or cell populations. Such continuous
flow sorters are usually operated under alternating current
in order to minimise electrolysis at the microelectrodes inside
the channel.
Sorting of 4 mm and 6 mm polymer particles was achieved in
a 500 mm wide and 28 mm deep channel, with embedded Pt
Fig. 9 The principle of dielectrophoresis (DEP): when subjected to an
electric field, a particle or cell becomes polarised. If the electric field is
inhomogeneous, then the electrostatic forces on the two ends of the
dipole are not equal and a movement is induced. (a) Positive DEP
occurs when the particle exhibits a larger polarisability than the
surrounding medium. Positive DEP is directed towards the stronger
electric field. (b) Negative DEP occurs if the particle is less polarisable
than the surrounding buffer medium. Negative DEP forces the particle
away from areas of high field intensity towards areas of low intensity.
Figure redrawn from ref. 63 with permission from Elsevier.
This journal is � The Royal Society of Chemistry 2007 Lab Chip, 2007, 7, 1644–1659 | 1653
stripe electrodes that were 50 mm wide and 50 mm apart and
arranged at a 45u angle to the direction of flow.60 When 10 V
peak to peak were applied to the electrodes at a frequency of
1 MHz, the particles were subjected to a nDEP force transverse
to the direction of flow. The 6 mm particles were deflected
further from the direction of flow than the 4 mm particles and
thus the two populations could be separated at a flow rate of
about 100 mm s21. A series of gold electrodes in the shape of
trapezoids were integrated at the bottom of a microchannel
and used to sort polystyrene particles of 6 mm and 15 mm
diameter via nDEP into two different outlet channels.61 A
voltage of 8 V at 50 kHz was applied to each electrode and
particles were sorted at a flow rate of 2 mm s21. In another
device, nDEP was used to block particles from entering a
channel.62 A gate electrode was integrated above and below a
100 mm central outlet channel. When 20 V was applied at MHz
frequencies to this gate electrode, 1 mm diameter polymer
particles were prevented from entering this channel and forced
to take adjacent channels. This valve was found to be 97%
efficient at a flow rate of 500 mm s21.
Viable and non-viable cells are often found to exhibit
different polarisabilities and can be sorted using DEP.63,64 At
low frequencies, the viable cells exhibit nDEP and the non-
viable cells pDEP. At high frequencies however, these forces
are reversed. Continuous flow sorting of live and dead yeast
cells was demonstrated in a 400 mm wide and 50 mm
deep channel with embedded gold electrodes.63 When 8 V
was applied at a frequency of 5 MHz, the viable yeast cells
exhibited pDEP, whereas the non-viable yeast cells exhibited
nDEP. Thus the two cell populations could be separated at a
flow rate as fast as 4 mm s21, with a cell throughput of
1000 cells s21. In another example, viable and non-viable
yeast cells were separated using 10 mm wide stripes of Pt/Au
electrodes, with a spacing of 10 mm that was integrated in
the bottom of a 600 mm wide and 100 mm deep separation
channel.64 Again, cells were deflected from the direction of
flow based on positive or negative DEP forces. The viable cells
were found to exhibit pDEP at 14 V and an applied frequency
of 100 kHz, whereas at the higher frequency of 2 MHz, the
viable cells were found to experience nDEP forces.
Cell separation via dielectrophoresis can be enhanced using
polymer particles as labels.65,66 The DEP force on an E. coli
cell attached to a polymer particle was found to be about
100 times larger than the force on an unlabelled cell.65 With a
20 mm deep microchannel featuring stripes of gold electrodes at
the top and bottom and an ac voltage of 20 V at 0.5 MHz, it
was possible to enrich 95% of the labelled E. coli cells. The cell
suspension was pumped at a flow rate of 3 mm s21 and thus a
throughput of 10 000 cells s21 was obtained. This device was
later improved to a two stage sorting system.66 Cells labelled
with 5.6 mm polymer beads had to pass two electrode arrays
to be collected. This device was operated for a period of
three hours with a throughput of 5 6 108 cells, equalling
about 50 000 cells s21.
Continuous flow dielectrophoretic sorting has so far only
been applied to a small range of samples, mainly polymer
particles and yeast cells. The method is mild enough to prevent
damage to biological cell samples and throughput can be tens
of thousands of cells s21.
Another interesting application involving inhomogeneous
electric fields was demonstrated by Leineweber and Eijkel.67
They integrated a regular array of electrodes into a micro-
channel. When applying a voltage, each electrode influenced
the local electrolyte concentration and molecular enrichment
against a concentration gradient was obtained. This was
used for continuous flow demixing: components of a
homogeneously distributed electrolyte could be focussed
into streams.
Separations involving magnetic fields
Continuous flow separation of magnetically susceptible
materials can be achieved by applying an inhomogeneous
magnetic field perpendicular to the direction of flow (Fig. 10).
A magnetically susceptible object, such as a magnetic
microparticle or a magnetically labelled cell, is pulled into
the magnetic field. The magnetic force depends on the gradient
of the applied magnetic field and also on the volume and
magnetisation of the particle. Particles of different sizes and/or
different magnetisation can be separated. The observed
particle trajectory is the sum of two flow vectors: the
hydrodynamic velocity arising from pumping the liquid
through the microfluidic system and the magnetically induced
velocity. Balancing of these two flow vectors is required
to generate sufficient deflection. General applications of
magnetism and magnetic particles in microfluidics have been
reviewed recently.68,69
Continuous flow magnetic separation methods were initially
developed on larger scales,70–73 but have been frequently
applied to microfluidics as well. Separators either feature two
outlets (Fig. 10a) or multiple outlets (Fig. 10b). The design
with two outlets is sometimes also termed split flow thin
fractionation (see below), whereas the design with multiple
outlets has been termed free-flow magnetophoresis, in analogy
to free-flow electrophoresis (see above).
Fig. 10 Continuous flow magnetic separations: an inhomogeneous
magnetic field is applied perpendicular to the direction of flow.
Magnetic particles or magnetically labelled cells are attracted into the
field and thus deflected from the direction of laminar flow. (a) Sorting
of magnetic particles is often carried out via two outlets. (b) In free-
flow magnetophoresis, multiple outlets are used, allowing for the
separation of different magnetic particles from each other, as well as
from non-magnetic material.
1654 | Lab Chip, 2007, 7, 1644–1659 This journal is � The Royal Society of Chemistry 2007
Separation of magnetic particles from a sample suspension
using a microchannel network with two outlets have been
reported, for example by Blankenstein et al.74 and by Luke
Lee’s group.75 In these earlier examples, researchers have
utilised small permanent magnets for particle deflection.
More recently, Kim and Park used a similar setup for an
immunoassay based on magnetic nanoparticles as labels.76
Microfabricated magnets in close proximity to the channel
can generate high magnetic field gradients, but have a limited
reach and might lead to extensive heating. Recently, a very
simple method for fabricating a micro-electromagnet has
been introduced by Siegel et al.,77 who filled a channel
adjacent to the particle carrying channel with liquid solder.
After solidification, the solder acted as wire and thus as an
electromagnet for continuous sorting of magnetic particles.
Most efficient for continuous flow separations are hybrid
magnets, a combination of a larger external magnet and a
small magnetisable feature in close proximity to the micro-
channel. For example, with a microfabricated NiFe comb and
an external magnet, a three times higher gradient could be
achieved in comparison to the external magnet alone; this in
turn allowed for the application of a three times higher flow
rate for the separation of magnetic microparticles or magne-
tically labelled E. coli cells.78 In another example, nickel wires
in close proximity to the microchannel were magnetised by an
external permanent magnet for the separation of red blood
cells in continuous flow.79
Free-flow magnetophoresis devices, as shown in Fig. 10b,
feature multiple outlet channels. Thus it is not only possible to
separate magnetic from non-magnetic particles but also to
separate different magnetic particles from each other. Using a
small permanent magnet, this has been demonstrated for
magnetic microparticles,52,80 as well as for cells labelled
with magnetic nanoparticles.81 In another example, micro-
fabricated magnetic Ni-stripes were embedded at the bottom
of the separation chamber at an angle to the direction of flow.
These stripes served to deflect magnetically labelled cells from
the direction of laminar flow.82
Magnetic manipulation has some distinct advantages over
electric manipulation. The magnetic field is applied externally;
there is no need for physical contact between the magnet
and any liquid. Unlike electrically induced forces, magnetic
forces are not significantly influenced by ionic strength, pH or
surface charges. Magnetic manipulation is mild and not
usually damaging to biological materials. However, magnetic
manipulation is rarely based on intrinsic properties. Most
materials do not exhibit intrinsic magnetism and hence
must be labelled with magnetic micro- or nanoparticles.
Exceptions include de-oxygenated red blood cells79,83 and
magnetotactic bacteria,84 which can be manipulated without
prior labelling. Magnetic forces on individual nanoparticles
are too small, given that the force is proportional to the cube
of the particle radius. Individual nanoparticles can generally
not be manipulated, agglomerates of thousands of nano-
particles are usually required. Even when using magnetic
microparticles, the applied flow rates tend to be somewhat
slow, hundreds of mm s21 or at most a few mm s21. This
imposes limitations on throughput in comparison to other
continuous flow methods.
It is less commonly known that non-magnetic materials,
or more specifically diamagnetic materials, are repelled by
magnetic fields. This effect is several orders of magnitude
weaker than magnetic attraction exhibited by ferromagnetic
materials. However, it has recently been demonstrated that
6 mm polystyrene particles can be deflected in continuous
flow using the high magnetic gradients generated by super-
conducting magnets.85
Separations involving sound pressure
Continuous flow separation of particles and biological cells
can also be performed with acoustic forces generated from
ultrasonic waves. For this, a standing sound wave must be
generated over the cross section of a microchannel, i.e.
orthogonal to the direction of flow (Fig. 11). Usually the
wave is tuned, such that the node is in the centre of the channel
and two anti-nodes at the edges. Such waves can be generated,
for example, with piezoelectric transducers. Particles or cells
subjected to the sound wave experience a force, either towards
the node or towards the anti-node.
A detailed description of the theory is beyond the scope of
this review and the interested reader is referred to reviews on
this topic by Coakley,86 and recently by Laurell et al.87 Briefly,
the acoustic force on a particle depends on the properties of
the acoustic field, as well as on the properties of the particle
and its surrounding medium. The acoustic field is charac-
terised by its frequency and amplitude: the higher the
frequency and the larger the applied voltage, the higher the
acoustic force. The required frequency is predetermined by
the width of the microchannel. For example, a channel of
Fig. 11 Continuous flow separation based on sound pressure. (a) A
standing ultrasonic wave is generated over the width of a microchannel
by means of an external transducer. Particles experience a force, either
toward the pressure node or the anti-node. The magnitude of the force
is influenced, amongst other parameters, by the particle size. The
direction of the force is determined by the ratio of the particle’s density
and compressibility in relation to the density and compressibility of the
surrounding buffer medium. (b) Once aligned in the sound field, the
different particles can be collected into separate outlets.
This journal is � The Royal Society of Chemistry 2007 Lab Chip, 2007, 7, 1644–1659 | 1655
1 mm width requires a wave frequency of about 100 kHz,
whereas a channel of 10 mm width requires 10 MHz. In short,
the smaller the channel, the higher the required frequency and
in turn the larger the acoustic force generated. The properties
of the particle and the medium in which it is suspended also
have a pronounced influence. The acoustic force is propor-
tional to the volume of the particle. Furthermore, it depends
on the density and compressibility of the particle in relation to
the density and compressibility of the surrounding medium.
This ratio determines the direction of the acoustic force
towards a node or anti-node. In general, a solid particle
suspended in an aqueous medium will be pushed towards a
pressure node, whereas a gas bubble will be pushed towards
an anti-node. For a given acoustic field, particles can be
separated if they differ in size, density or compressibility or a
combination of these factors. As for all continuous flow
separation methods, the flow rate must be adjusted such that a
particle spends sufficient time in the pressure field.
Microfluidic devices with ultrasonic transducers have been
used to focus polymer particles or cells into a stream and thus
to separate these from the bulk of the suspension. Hawkes
and Coakley88 fabricated a stainless steel device with a channel
10 mm wide and 250 mm deep, over which a standing
ultrasonic wave was generated with a frequency of about
3 MHz. Yeast cells of 5 mm diameter were introduced at a
flow rate of 6 mL min21 (40 mm s21). The cells were focussed
into the centre of the channel and collected in a central outlet
channel, whereas the clarified medium was collected in the two
outer outlet channels. A throughput of 2 6 108 cells min21
was achieved.
Laurell’s group have reported on a number of ultrasonic
separation devices. In a 750 mm wide and 250 mm deep silicon
channel, a first harmonic ultrasonic wave with two nodes was
generated at a frequency of 2 MHz.89 The device featured three
outlet channels. A suspension of 5 mm diameter polyamide
particles was focussed into the two nodes and collected via the
two outer outlet channels with 90% efficiency in 2/3 of the
original volume at a flow rate of 0.1 mL min21 (9 mm s21). In
a subsequent design, a 350 mm wide and 125 mm deep channel
was fabricated and a fundamental ultrasonic wave with one
node was generated at 2 MHz.90 In this device, almost 100% of
the 5 mm polyamide particles could be collected in the central
outlet channel, i.e. into 1/3 of the original volume at a flow rate
of 0.3 mL min21 (110 mm s21). More recently, a fluidic design
with a similar separation channel cross section, but with a
more sophisticated design of inlet and collection channels,53
was utilised to separate a mixture of polystyrene particles
of 2 mm, 5 mm, 8 mm and 10 mm diameter at efficiencies
between 62% and 94%. Due to the multiple outlets, the authors
termed their method ‘‘free-flow acoustophoresis’’. The same
device was also used to separate polystyrene particles of 3 mm,
7 mm and 10 mm with efficiencies between 76% and 96%.
The separation of polyamide microparticles and of red
blood cells from blood plasma was also demonstrated by
Kapishnikov et al.91
As mentioned above, particles of the same size can be
separated if they differ in density from each other and the
surrounding medium. This was demonstrated with 3 mm dia-
meter polystyrene and PMMA particles that were suspended
in an aqueous solution of 0.22 g mL21 cesium chloride.53 The
same CsCl solution was also applied to the ultrasound based
separation of red blood cells, leukocytes and platelets.
An interesting application of continuous flow ultrasonic
separation is the purification of blood from lipid droplets
during open heart surgery. The sonic force on red blood cells
suspended in plasma pushes the cells towards pressure nodes,
whereas lipid droplets are pushed towards pressure wave anti-
nodes. Petersson et al. demonstrated how on-chip continuous
ultrasonic separation could be used for this application.90
The elegance of acoustic separations stems from fact that no
physical contact is required between the ultrasonic transducer
and the liquid. A standing wave is generated inside the channel
by an external device. Ionic strength and surface charges are
not important, and according to the literature, the ultrasonic
treatment has not been shown to damage cells or biological
material. Notably, throughput can be quite high with flow
rates of tens of mm s21. The method, however, is not
applicable to small nanoparticles or small molecules, given
that the acoustic force is proportional to the cube of the
particle radius.
Continuous flow separation via optical forces
Optical forces have been applied extensively in optical tweezers
where a focussed laser beam is utilised to trap and move
microparticles or cells, and also for particle by particle sorting
in flow cytometry. Continuous flow separation, as indicated in
Fig. 1b, can be achieved by means of an optical interference
pattern, a so-called optical lattice. A three-dimensional optical
field can be generated, such that the optical force is directed
perpendicular to the direction of flow. The optical force on a
particle or cell depends on optical polarisability. Sorting can be
achieved based on differences in particle size or refractive
index. MacDonald et al. demonstrated the sorting of two types
of particles based on refractive index.92 In an H-shaped
channel device, they introduced a buffer solution, as well as a
suspension with a mixture 2 mm silica particles (n = 1.73) and
2 mm polymer particles (n = 1.58). When passing through the
optical field, the polymer particles were deflected into the
buffer carrying stream, whereas the silica particles flowed
straight on. Sorting was achieved with almost 100% efficiency,
at a flow rate of 30 mm s21, equalling a throughput of
about 25 particles s21. The same device was also utilised for
deflecting and separating protein capsules of 2 mm from a
mixture of 4 mm diameter capsules at a flow rate of 20 mm s21.
More recently, the same group demonstrated the separation
of a mixture of four differently sized silica particles (2.3 mm,
3.0 mm, 5.3 mm and 6.8 mm).93
Optical sorting is another example of a non-invasive
method. Separation is based on intrinsic properties, no
labelling is required. The optical force can be tuned
independently to the flow rate. The method can be applied
to inorganic, polymeric and biological particles.
Separation based on gravity
Gravity is an obvious force to utilise for continuous flow
separations. However, gravity forces on polymer particles or
1656 | Lab Chip, 2007, 7, 1644–1659 This journal is � The Royal Society of Chemistry 2007
cells suspended in buffers tend to be several orders of
magnitude smaller than, for example, dielectrophoretic or
magnetic forces. The gravity force is proportional to the
particle volume and also to the difference in density between
a particle and the surrounding buffer medium. Recently,
Huh et al. demonstrated the separation of 20 mm polystyrene
particles from 1 mm particles and also from 3 mm particles
in a microfluidic device (Fig. 12) by combining gravity
forces with hydrodynamic flow.94 A sheath flow was applied
to initially focus the particles parallel to the gravitational
field. The particle stream was then directed perpendicular to
gravity into a separation channel. By gradually widening this
channel, the flow rate slowed down and sedimentation started
to take effect. Differences in particle position were then further
enhanced by the hydrodynamic flow profile and channel
geometry, similar to pinched flow fractionation10 or asym-
metric cavity separation.13 Separation was achieved in less
than 1 min, with almost 100% efficiency. The method was also
utilised to separate perfluorocarbon droplets in water.
Split flow thin fractionation
Split flow thin (SPLITT) fractionation was originally
developed by Giddings95 and has been extensively used in
millilitre volume flow cells. The method is based on forcing
sample components out of a sample stream into a neighbour-
ing buffer stream by a force applied perpendicular to the
direction of flow. This force can be based on gravity,96
electrophoresis,97 or magnetism.98 In contrast to a conven-
tional H-shaped flow cell design, as for example shown in
Fig. 10a, SPLITT devices feature two flow splitters at the inlet
and outlet of the flow cell (Fig. 13). The carrier stream is often
introduced at a faster flow rate than the sample stream, thus
focussing the sample stream into a relatively narrow band, also
referred to as the inlet splitting plane. With no force applied,
all sample components leave via the outlet opposite the inlet.
When a force is applied, any susceptible material must be
pulled beyond the so-called outer splitting plane to be collected
in the second outlet channel.
Split flow thin fractionation has also been demonstrated
in microfluidic devices.99 For example, an electrophoresis
SPLITT cell was fabricated by Narayanan et al.,100 featuring
gold electrodes at the top and bottom of a 1 mm wide
and 40 mm deep channel. By applying an electric field of
300 V cm21, a mixture of 108 nm and 220 nm polymer
particles could be separated at a flow rate of 14 mm s21.
Conclusion
There are now an immense variety of continuous flow
separation methods available to the bioanalytical chemist for
the separation of particles, cells or DNA molecules, proteins
and smaller molecules.
Comparison of performance is not always straightforward:
some methods offer high throughput, whereas others offer
high resolution; several of the microfluidic devices are simple
in terms of operation, other techniques might require a
specialist; a number of separation principles require labelling
of sample components, whilst some processes are based on
intrinsic sample properties. As always, the optimum method
will depend on the sample and the analytical task at hand.
Most publications on continuous flow separation methods
so far have been concerned with proof of principle experiments
or characterisation of performance. As yet, there are very few
cases where continuous separation methods have been
integrated within a microTAS. One exception is the pillar
based leukocyte enrichment filter by Chen et al.,20 that was
combined with downstream cell lysis and DNA purification.
However, researchers now have an extensive toolbox of
methods available, and in the near future we are likely to see
more of these continuous flow techniques integrated within
miniaturised total analysis systems, such as continuous flow
blood separation for point-of-care systems, continuous flow
DNA separation for at-the-scene forensic testing and con-
tinuous flow molecule separation for in-the-field environ-
mental detection. It is quite possible that these continuous flow
separation methods will contribute significantly to a wider
breakthrough of microfluidic technology in the market place.
Fig. 12 Separation of two particle populations based on gravity.
Initially, particles are hydrodynamically focussed parallel to the
direction of gravity and then guided around a 90u bend into a
gradually widening separation channel. As particles enter this
separation channel, flow rates slow down and sedimentation starts to
take effect. Any differences in particle position caused by sedimenta-
tion are further enhanced by the widening flow streams, thus allowing
for the separation of the two particle populations. Figure redrawn
from ref. 94 with permission from the ACS, copyright 2007.
Fig. 13 The principle of split flow thin (SPLITT) fractionation. A
particle suspension and a carrier stream are pumped into a separation
chamber, often at different velocities. Flow splitters at the entrance
and exit of the chamber smooth the flow and define an inner and outer
splitting plane. A force field, such as gravity, an electric field or a
magnetic field, is applied perpendicular to the direction of flow,
deflecting some of the particles into the carrier stream.
This journal is � The Royal Society of Chemistry 2007 Lab Chip, 2007, 7, 1644–1659 | 1657
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