Continuous flow separations in microfluidic devices - Semantic Scholar

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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 procedures.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 separations 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 electrophoresis (Fig. 1a). An exciting development within the microfluidics community 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 a Diploma in Chemistry from the University of Marburg, Germany, in 1999. For her PhD she went to Imperial College London, where she worked on single particle analysis in microfluidic chips until 2004. This was followed by a stay in Tsukuba, Japan as an independent research fellow in the International Centre for Young Scientists (ICYS) at the National Institute for Materials Science (NIMS). Nicole Pamme In December 2005 she took up a lectureship position at the University of Hull (UK). Her current research interests include continuous flow separation in microfluidic devices based on magnetic forces. 1644 | Lab Chip, 2007, 7, 1644–1659

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 components 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 downstream 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. This journal is ß The Royal Society of Chemistry 2007

(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 discarding all other components. Alternatively, a sample can be purified of undesired components and then used for downstream 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 downstream 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 partitioning.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 filtration 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 Lab Chip, 2007, 7, 1644–1659 | 1645

Table 1 Listing of continuous flow separation methods with details of the forces utilised, the basis of separation and the applicable samples. The data 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 40 000 particles h21

11

Hydrodynamic filtration

Laminar flow regime

Size

Microparticles

1 mm s21 100 000s particles h21

14

Bifucation

Laminar flow

Size, shape

Blood

5 mL h21 1–3 mm s21

16

Filtration obstacles

Diffusion and obstacles

Size

Blood

5–12 mL min21 10–20 mm s21

19

Lateral displacement

Obstacle array

Size

Particles (0.8, 0.9, 1.0 mm) DNA (61 kbp, 158 kbp)

20 mm s21 (DNA) 100 pg h21 (DNA)

24

Brownian ratchet

Diffusion and obstacles

Size

DNA

2 mm s21 10 pg h21

26

Hydrophoretic separation

Pressure gradient from slanted obstacles

Size

Microparticles

1 to 3 cm s21 10 000s beads h21

30

DNA prism

Obstacle array and electric field

Size

DNA

10 ng h21

31

Entropic trap array

Dams and electric field

Size

DNA

10 pg h21

32

21

Repulsion array

Dams, electric double layers

Charge to size ratio

Proteins

300 pg h

Free-flow electrophoresis

Homogeneous electric field

Charge to size ratio

Proteins, amino acids

6 mm s21

44

21

45

32

Free-flow isoelectric focussing

pH gradient

Isoelectric point

Proteins and cells

9 mm s

Magnetophoresis

Inhomogeneous magnetic field

Size, magnetisation

Magnetic particles, cells

up to 1.2 mm s21 10 000s cells h21

81

Dielectrophoresis

Inhomogeneous electric field

Size, polarisability

Microparticles, cells

up to 4 mm s21 100 000s cells h21

63

Acoustophoresis

Acoustic pressure

Size, density, compressibility

Microparticles, cells

11 cm s21

90

Optical lattice

Optical force

Size, refractive index

Microparticles, cells

30 mm s21 90 000 particles h21

92

Sedimentation

Gravity

Density, size

Microparticles, droplets

6 mm s21 slowed to 1 mm s21

94

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 homogeneous 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 1646 | Lab Chip, 2007, 7, 1644–1659

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. This journal is ß The Royal Society of Chemistry 2007

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

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.

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 This journal is ß The Royal Society of Chemistry 2007

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. 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 1648 | Lab Chip, 2007, 7, 1644–1659

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 This journal is ß The Royal Society of Chemistry 2007

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.

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 technique 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 encountering 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 This journal is ß The Royal Society of Chemistry 2007

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.

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 Lab Chip, 2007, 7, 1644–1659 | 1649

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.

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.

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 1650 | Lab Chip, 2007, 7, 1644–1659

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 This journal is ß The Royal Society of Chemistry 2007

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.

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, corresponding 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 This journal is ß The Royal Society of Chemistry 2007

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 parameters, 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. Lab Chip, 2007, 7, 1644–1659 | 1651

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 homogeneous 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.

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 1652 | Lab Chip, 2007, 7, 1644–1659

separation chamber. This must be undertaken whilst minimising 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, This journal is ß 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 ‘‘freeflow’’ 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 polarisability 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

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.

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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 electroosmotic 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 dielectrophoretic 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 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 nonviable 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. 1654 | Lab Chip, 2007, 7, 1644–1659

Another interesting application involving inhomogeneous electric fields was demonstrated by Leineweber and Eijkel.67 They integrated a regular array of electrodes into a microchannel. 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 freeflow magnetophoresis, multiple outlets are used, allowing for the separation of different magnetic particles from each other, as well as from non-magnetic material.

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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 microchannel. 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 magnetically 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, microfabricated 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 nanoparticles 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. This journal is ß The Royal Society of Chemistry 2007

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 superconducting 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 characterised 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.

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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 proportional 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 diameter polystyrene and PMMA particles that were suspended 1656 | Lab Chip, 2007, 7, 1644–1659

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 antinodes. 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 This journal is ß The Royal Society of Chemistry 2007

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 sedimentation 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.

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 asymmetric 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 neighbouring 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 conventional 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 This journal is ß The Royal Society of Chemistry 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.

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 continuous flow molecule separation for in-the-field environmental 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. Lab Chip, 2007, 7, 1644–1659 | 1657

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