Multi-step microfluidic system for blood plasma separation ...

Microfluid Nanofluid (2014) 17:167–180 DOI 10.1007/s10404-013-1296-4

RESEARCH PAPER

Multi-step microfluidic system for blood plasma separation: architecture and separation efficiency Julien Marchalot • Yves Fouillet • Jean-Luc Achard

Received: 27 May 2013 / Accepted: 15 November 2013 / Published online: 3 December 2013 Ó Springer-Verlag Berlin Heidelberg 2013

Abstract This document presents a multi-step microfluidic system designed for passive and continuous separation of human blood plasma. This system is based on lateral migration of red blood cells, which leads to a cell-free layer close to the wall. The geometry of the plasma separation unit used in this system was determined by research conducted by Sollier et al. in Biomed Microdevices 12:485–497 (2010). It makes the most of geometric singularities to increase the size of the cell-free layer and optimise the quantity of plasma recovered. This article endeavours to show the importance of architectural fluidic connection upstream of the cell on plasma yield. The design of the chip was then modified to remove its connection role. It was then possible to consider installing the yield cell in series. The approach used for the overall optimisation of the system is presented in the article. In the case of two successive patterns, the increase in pure (diluted) plasma yield ranges from 18 to 25 % for 1:20 diluted blood, and the quality of the plasma obtained is compared to traditional separation methods. Keywords Continuous Passive Blood plasma separation Cell-free layer Plasma yield Architectural fluidic connection

J. Marchalot (&) Y. Fouillet CEA-LETI-Minatec, SBSC, DTBS, 17 rue des Martyrs, Grenoble Cedex 9 38054, France e-mail: [email protected] J.-L. Achard Laboratoire des Ecoulements Ge´ophysiques et Industriels (LEGI), Grenoble Cedex 9 38041, France

1 Introduction Diagnosis tests are mainly conducted on blood, which is easily accessible and is representative of the patient’s pathological condition. Plasma contains high levels of protein, which makes it a sample of choice in 90 % of testing cases. The conventional procedure to separate plasma from blood cells usually involves a centrifuging phase followed by a manual extraction phase. Such a macroscopic technique involves several phases requiring a lot of equipment and qualified staff in specialised laboratories. Moreover, the separation process is conducted after variable post-sampling waiting times. This may result in contamination risks and non-reproducibility issues, even for simple diagnoses (Toner and Irimia 2005). Several research teams have already conducted successful blood analyses using a microsystem such as proteins (Browne et al. 2011), RNA and DNA (Sethu et al. 2006), antigens (Cheng et al. 2009) and circulating cancer cells (Gleghorn et al. 2010). However, few have worked on the integration of the upstream stage consisting in extracting plasma, or at least for sample volumes greater than several hundred microlitres (for tens of microlitres Songjaroen et al. 2012). Thus, this stage remains a challenge to be overcome before finalizing the complete integration of on-chip blood analyses. It must be conducted in ideal fashion, with no analyte losses and no blood dilution, despite the presence of a high number of red blood cells. Furthermore, the stress imposed on the blood cells must be managed, since red blood cell lysis is incompatible with most plasma analysis methods. The objective is therefore to develop an efficient sample preparation method that could be integrated in a lab-on-achip quickly and easily, with a view to extracting plasma from non-diluted or slightly diluted whole human blood

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using a reproducible process. Recently, different continuous flow separation methods have been developed to carry out this plasma extraction process in microsystems. Some of these methods rely on the combination of a microfluidic process and a force field, whether electric (Arifin et al. 2007; Nakashima et al. 2010), magnetic (Furlani 2007; Iliescu et al. 2008), or acoustic (Lenshof et al. 2009) or with a rotating CD for centrifuging purposes (Amasia and Madou 2010). Other methods use passive hydrodynamic phenomena in micro-channels, such as sedimentation (Dimov et al. 2011), cross-flow filtration (Homsy et al. 2012), hydrodynamic filtration (Matsuda et al. 2011; Maltezos et al. 2011), cell deviation by a network of spots (Davis et al. 2006) or centrifugation in curved channels (Di Carlo et al. 2007); or spirals (Russom et al. 2009). However, despite some specific advantages, certain existing passive methods also have major drawbacks (Sollier et al. 2009): clogging, relatively low flow rates, and creation of Dean vortices in a curved channel (Berger et al. 1983; Ookawara et al. 2006). The research concerned by this article follows on from previous work conducted by Sollier et al. (2010), who developed a new passive method that limits the previous

Fig. 1 a Schematic view of the separation unit composed of three segments: (1) upstream channel; (2) restriction; (3) downstream expanded channel. QIN, Qe, and Qout represent the respective flow rates in inlet, plasma, and outlet channels. b Picture of the separation unit for 1:20 diluted blood at 100 lL/min

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Microfluid Nanofluid (2014) 17:167–180

drawbacks. The principle implemented in the separation unit used for this method is to extract the plasma from recirculation zones created by a sudden diverging section, because these zones are free from any particles. In this previous work, Sollier et al. present design and performance optimization of this unit for diluted blood (1:10–1:20), and the quality of the extracted plasma has been studied through contamination by platelets and by red and white blood cells depending on flow rate and blood concentration. The obtained plasma has thus been shown of similar quality as plasma obtained by classical centrifugation. The separation unit is described in detail in the first part of the article (Fig. 1a) and then chosen as the benchmark cell. In order to make this unit compatible with industrial integration and constraints, we hereby propose a more efficient overall fluidic architecture around this separation unit leading to a more convenient update of the global device. More precisely to our knowledge, the importance of the upstream channel on the cell extraction efficiency has never been taken into account; in this respect, upstream connections could be likened to this channel. The context of this research is hence the optimisation and integration of a system’s analytical chain, which, in short, comprises three successive channels: an upstream channel with its connections, a reduced intermediate channel (called restriction) leading to a diverging channel in which the circulation zones are located. In the second part, the influence of the upstream channel was studied, and in order to assess its influence easily, it was ‘‘moved outside’’ of the system using different types of connections (connected to the chip by the edge). This helped determine optimum dimensions for the separation unit. This optimum upstream channel was then reintegrated in the chip in the form of a coiled channel. With this configuration, it is possible to limit the surface area taken up by the upstream channel. In the context of this implementation, which is described in the third part of the document, another simplified external connection mode (from the top) is used. In order to increase the volume of plasma extracted, it is obvious that several separation units must be installed in series (Kersaudy-Kerhoas et al. 2009). The downstream channel of a unit then becomes the upstream channel of another. In other words, the upstream conditions of each separation unit became the coupling conditions. This system was optimised on the basis of the work presented in the previous parts. The fourth part presents a typical architecture with two successive extraction units: the upstream channel of each unit guarantees clean extraction, and head losses dictate the flow rate. In this way, pure plasma extraction yield is increased by 18–25 % for 1:20 diluted blood. Furthermore, in order to meet practical requirements,


part 4 will be aimed at transforming the experimental process into a routine, exploitable system that can be used in a medical biology laboratory for example. In this respect, extraction yields can be precisely determined by using several volumetric pumps (syringe pumps), but experimentation remains tricky and long. Yet, since yields are now well documented and controlled thanks to the results of part 3, the ratio between the injection flow rate and the extraction flow rate may be determined definitively. A head loss system is used to control all the flows by maintaining only one injection at the chip inlet. The quality of the plasma obtained this way has been approved by a SELDI analysis. 2 Theoretical background 2.1 Plasma extraction and cell-free layer 2.1.1 Separation unit The reference microfluidic design studied by Sollier et al. (2010) is provided once again in this part. As mentioned in the introduction, three ‘‘cooperative’’ sections can be distinguished that act as the microfluidic system: an upstream channel, a restriction (which is a highly shearing channel), and an expanded channel. These last two sections correspond to the basic pattern, which is used in all the systems presented in this article, denoted ‘‘separation unit’’. Different extraction point couples may be considered in the two recirculation zones of the expanded channel, and the choice made is illustrated in Fig. 1. Regarding the separation unit presented here, the absence of particles in the recirculation zones determines the plasma extraction quality: this absence is due to their vortex movements, which have a tendency to eject red blood cells towards the outside under the effect of centrifugal force, but also to the existence of large cell-free layers near the walls upstream and downstream of these zones. Upstream, the cell-free layer developed in the restriction is used to feed these zones with a particle-free flow, thereby creating a sort of prior separation. Downstream, the cellfree layer guarantees continuous flows without having to reinject particles in these zones. Note that the existence of a cell-free layer is also exploited in various ways in plasma extraction systems, for example to divert this cell-free layer to an appropriate lateral channel (Faivre et al. 2006; Kersaudy-Kerhoas et al. 2009; Tripathi et al. 2013). 2.1.2 Principle The creation of a red blood cell-free layer close to the walls when blood flows through a straight pipe is a physical

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phenomenon commonly known as the Fahraeus or Fahraeus-Lindqvist effect (Fahraeus and Lindqvist 1931). Interpretation prior to this phenomenon is therefore useful to complete the overall optimisation of the system, one of its main objectives being to increase the thickness of this cell-free layer. In this perspective, the experiments of Segre and Silberberg (Segre´ and Silberberg 1962) may be a starting point. They studied the migration of dilute suspensions of neutrally buoyant spheres in pipe flows at Reynolds numbers (Re) between 2 and 700; in the context of the systems described here, it should be noted that for the range of flow rates used (50–400 ll/min), the Reynolds number values calculated with an intermediate diameter Øint (Re ¼ V;int =m ¼ 4Q=pm;int ) range from 4 to 25 (V is the mean velocity of the fluid, Øint the hydraulic diameter of the pipe, m the kinematic viscosity, and Q the volumetric flow rate). The particles migrate away from both the wall and centreline and accumulate at 0.6 of a pipe radius. Shear gradients, due to curvature of the undisturbed flow profile, produce lateral forces. Thus, a particle released at the centre of the pipe will be driven by shear gradients towards the wall. A particle released near the wall will be driven away from the wall by a combination of three mechanisms involving inertial effects and interaction between particles and bounding walls (Feng et al. 1994a, b): wall lubrication repulsion, inertial lift due to shear slip when the particles lag the fluid (Saffman 1965), and a Magnus-type lift associated with the rotation of the particle (Rubinow and Keller 1961). The counter-clockwise rotation of the particle, together with the lag velocity, may produce a lift force directed outwards from both walls to the centre of the channel. For cells, however, such an enlightening scheme may appear inadequate in four respects. Some of them must be taken into account, and others need not be: 1.

2.

Red blood cells are not neutrally buoyant given their higher density compared to the suspending medium (qRBC = 1.09 [ qplasma = 1.02 g/cm3). In the flat horizontal channels in this study, this is of no consequence since the various lift forces considered, which control the cell-free layer development, are perpendicular to the weak buoyant force. Red blood cells are not spherical. At rest, red blood cells can be approximately described as axisymmetric oblate ellipsoids, with major axis, a = 8 lm, and minor axis, b = 2 lm, leading to an aspect ratio of 0.25. Some authors (Keller and Skalak 1982) think that red blood cells move according to two regimes. In the first, the cell rotates as a rigid disc, at low shear rates and at high viscosity ratio associated with inner and outer fluids, respectively. Note that the former is difficult to define since it involves the energy

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dissipated in the membrane. As the shear rate increases, there would be a transition to a tanktreading motion where the cell would maintain a steady tilt with respect to the flow. This migration away from the wall, known to occur for asymmetric deformable particles in laminar viscous creeping flow, is due to a viscous lift force and has long been believed by physiologists to occur in vivo. Some others (Olla 1999) think that due to their stiffness, red blood cells will resist tank-treading motion (Tran-Son-Tay et al. 1984) and will rotate as rigid objects under the effect of the vorticity part of the shear (‘‘flipping’’ motion regime). In the same vein, when the suspending medium viscosity is of the order of 1 Cp, experimental observations (Barthes-Biesel and Sgaier 1985) show

3.

Fig. 2 a The angular distribution function of neutrally buoyant particles (Rb = 75.78) for a solid area fraction (representing concentration) of 25 %. b The solid fraction distributions for buoyant elliptical particles with global solid area fractions of 13, 25, and 40 %. The bulk Reynolds number is defined by Rb = VaW/m, where Va is the average velocity of the suspension, W the channel width, and m is the kinematic viscosity (adapted from Qi et al. 2002, with kind permission from Springer Science and Business Media)

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4.

that the red blood cells behave as flexible elastic solids and have a ‘‘flipping’’ motion even at very high shear rates. This point of view will be adopted here. When isolated, rigid oblate particles exhibit an oscillatory motion towards and away from the wall (flipping regime) and experience several interactions, such as a lift force component, which do not exist for spheres. This has been shown by 2D simulations of ellipsoidal and inertia-less particles placed in a shear flow near a wall, assuming Stokes equations (Gavze and Shapiro 1998). This lift force is maximum for an aspect ratio b/a = 0.35 and is equal to zero when b/a = 1 and 0. Moreover, simulations of neutrally buoyant ellipsoidal particles having b/a = 0.5, placed in a Poiseuille flow in a 2D channel with a confinement ratio of a/Øint = 0.12, have been conducted using Navier– Stokes equations (Qi et al. 2002). It has been shown that the particles slightly lag the local velocity in the undisturbed flow and, when a steady state is reached, have a much higher probability of being oriented along the flow direction than along the cross-flow direction (Fig. 2a). However, note that their results show a slight steady tilt with respect to this direction, result whose importance will be shown later. Such simulations have clearly demonstrated that the Segre–Silberberg phenomenon exists based on wall lubrication repulsion, an inertial lift due to shear slip, and a lift associated with velocity curvature (Fig. 2b). The lift due to rotation disappears in the final configuration. Red blood cells have specific interactions when their concentration increases. In fact, the latter article also investigated such an effect by conducting simulations of 44 and 70 elliptical particles, in which the concentrations are 25 and 40 %, and Rb is 71.81 and 75.78, respectively (Fig. 2b). When the concentration increases, the results show that the two over-concentrated spikes become broader, the central region concentration becomes flat, and the cell-free layer thickness decreases. Note that these results correspond to experimental observation of reference cells. The rotation of a particle is also retarded by the existence of its neighbouring particles, and the probability of being oriented along the flow direction again increases. The velocity distribution of flow becomes blunt. These results suggest that the Segre–Silberberg effect becomes weaker. In our experiments, the concentration range will always be lower (dilution 1:10–1:20) and the Segre–Silberberg effect will be still present. Red blood cells are deformable. In paragraph (2), it was assumed that the red blood cells are not deformable. In fact, their deformation depends on the intensity and duration of mechanical stimulation (shear action) applied to red blood cells (Markle et al. 1983; Mills


et al. 2004). In our case, gradual deformation throughout the flow upstream of the unit is made possible because the shear stress can be applied long enough on the cells that are located in the over-concentrated ring by changing the geometry of upstream channels (length and width). As mentioned in (2), the effect of the vorticity part of the linear shear induces a ‘‘flipping’’ motion regime. The pure elongational part causes a modification of the cell eccentricity via the following process. Let a be the angle between the direction of the flow and the ellipse’s longer axis; clearly, h = p/2-a, h being defined in insert of Fig. 2a. The angular distribution function, presented in Fig. 2a, indicates that the probability to get a value in the region around h = p/4 (i.e. a = p/4) is slightly greater than the probability to get a value in the region around h = 3p/4 (i.e. a = -p/4). Thus, the red blood cells tend to be rather stretched and flattened than compressed. Following this prior interpretation of the phenomena, the hypothesis on which our approach has been based can be introduced. This hypothesis consists in saying that the more the red blood cells are flattened and the stronger the repulsion force applied to them, the larger the cell-free layer will be which is the obvious consequence we are looking to prove. The above-mentioned Segre–Silberberg mechanisms acting in the elliptical particle system reflect the phenomena encountered for blood flows in our channels. In particular, when concentration increases, wall lubrication repulsion plays a fundamental role. Clearly, this force increases when cells become flattened like a wing with an increasing chord. These cells gradually push the dispersed phase away from the wall. Faivre et al. (2006) demonstrated this phenomenon in their experiments by forcing red blood cells into a restriction in such a way that would deform their membrane and their cytoskeleton. The role of this restriction is therefore to ‘‘flatten’’ the red blood cells by shearing them in order to increase the size of the cell-free layer upstream of the unit. For practical reasons (limiting head losses, clogging up, and lysis risks, etc.), this considerable deformation of the red blood cells cannot take place in an excessively long restriction. It is therefore carried out in two stages: in a short-length restriction, but also in the upstream inlet channel, which has a reduced width, without, however, leading to the drawbacks mentioned above. This inlet channel therefore creates a so-called pre-deformation of red blood cells or ‘‘pre-configuration’’ of the cell-free layer. Since the relaxation time for the cytoskeleton following this deformation is greater than the transit time of the red blood cells in the channel (Armstrong et al. 2004), when the red blood cells go through the suddenly diverging section, they stay flattened until the end of the recirculation zone. The

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Fig. 3 Establishing of the cell-free layer (in lm) in a straight channel (50 mm long and 200 lm wide), at the outlet of a 250-lm internal diameter capillary tube

resulting upstream cell-free layer is therefore wider than the cell-free layer that would exist in a channel with the same width for a restriction that would not be as narrow. To conclude, the wise choice of (upstream channel, restriction) couple contributes to thickening of the cell-free layer upstream of the recirculation zone and the cell-free layer downstream. The latter may also be trivially increased by widening the diameter of the downstream channel. Figure 3 highlights two events which will be detailed later. As expected, according to the mechanism explained above, we can see that the formation of the cell-free layer is a progressive phenomenon. Thus, in a straight channel (50 mm long and 200 lm wide), the cell-free layer is established gradually to a maximum value of about 20 lm. Observe that the cell-free layer is already partially formed (width of about 10 lm) at the entrance of the channel: because of their small size, the external connectors typically used in microfluidics have a significant effect on the formation of cell-free layer. In the rest of this document, we will underline the influence of the overall fluidic architecture (part 3), which includes the external connection, as well as the pre-configuration of the cell-free layer. This research makes even more sense in the context of its integration in an analytical chain, and in particular if several separation units are placed in series, where the upstream conditions of each separation unit become coupling conditions that must be characterised to optimise the operation of the successive separation units (part 5). 3 Equipment and method 3.1 Experimental protocol 3.1.1 Microfluidic design and manufacturing The microfluidic channels are made from PDMS (Sylgard 184, Dow Corning), using the traditional soft-lithography

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manufacturing method (Xia and Whitesides 1998). The mould is produced with Ordyl dry, photosensitive film. The channels are 100 lm (h) deep, and their width ranges from 50 to 600 lm. The system has an inlet to feed the blood sample (IN), a main outlet (OUT) for the cells/red blood cells, and, potentially, two plasma outlets to recover the plasma. The associated flow rates are, respectively, denoted QIN, QOUT and QE. QIN and QOUT flow rates are controlled by two syringe pumps (KDScientific) used, respectively, to inject or extract plasma (Fig. 1a). Two types of connections are used: type A connection by the edge (Photon Line), and type B connection from the top (Tygon, MasterFlex). For the connection B, it was decided to use Tygon-type flexible capillary tubes with wider diameters to connect to the chip. These capillary tubes are easier to use, and their connection to the PDMS using a metal tube is more resistant to pressure; furthermore, it is possible to inject samples at higher flow rates and concentrations.

maximum plasma extraction rate can therefore be defined as follows:

3.1.2 Observation and acquisition

3.2 Results and discussion

The flow of the blood sample in the microfluidic channels was observed using a microscope (Olympus B960), and photos and videos were acquired using a video camera (Hamamatsu 1394 ORCA-ERA).

3.2.1 Plasma yield

3.1.3 Blood samples Human blood is taken from healthy volunteers by Etablissements Franc¸ais du Sang (EFS) in EDTA tubes to avoid coagulation (Vacutainer, BD). Each blood sample is diluted before being injected in the microfluidic system at ambient temperature in a phosphate-buffered saline solution (PBS). The dilution (ratio of non-diluted blood volume to total volume) can vary from 1:50 to 1:5 depending on the type of experiment, but it is usually 1:20.

g ¼ QMAX ¼

QEMAX QIN

ð1Þ

This ratio is called the plasma yield and is measured by the maximum possible extraction rate to recover pure plasma. In practice, the contamination of the sample is determined by visually and qualitatively examining the sample flow in the plasma recovery channels. A reference permissible contamination threshold could be taken as being equal to the typical contamination level obtained by centrifuging (approximately 1 % for red blood cells and below 50 % for platelets). The following parts show that quantitative concentration measurements obtained by counting cells (Quick Read, Globe Scientific INC and SCEPTER, Millipore) confirm the accuracy and relevance of this qualitative measurement process.

The performance levels of the reference system are briefly reviewed in this part. For a 1:20 dilution rate, the plasma yield (volumetric extraction yield of clean, diluted plasma) is relatively steady at a value between 15 and 17 % for flow rates ranging from 50 to 175 lL/min. This flexibility is important in the implementation of an integrated system because it shows that a high plasma yield can be obtained without necessarily having to adjust flow rates very precisely. Furthermore, for the dilution interval considered here (from 1:50 to 1:5), plasma yield (g) decreases as sample concentration decreases. In the low dilution zone, a proportionality ratio was observed between the plasma yield and dilution of the injected sample (up to 1:10, when plasma yield amounts to 10 %).

3.1.4 Plasma yield 3.2.2 State of the art The performance levels of the systems are quantified by the plasma yield (g). The volumetric extraction rate in plasma Q* can be defined by the ratio between extracted plasma volume (VE) and the volume of blood injected: Q* = VE/VIN = QE/QIN with QIN = QE ? QOUT. By reducing the discharge rate (QOUT) with the second syringe pump, the plasma flow rate (QE) and thus the extraction rate are increased. It is important to note that beyond the maximum extraction rate (QE), the plasma purity can no longer be guaranteed: the greater the volume of plasma extracted, the more the red blood cells are attracted towards the walls, which can lead to flooding and contamination of the plasma extraction channels. The

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Comparison with previous systems is not necessarily relevant: flow rates in lL/min were much lower than the range of flow rates used for our research. From this point of view, a comparison with work conducted by KersaudyKerhoas et al. (2009) seems to be more relevant. Extraction patterns have similar geometries, the main differences residing in the repetition of the pattern and, more importantly, in much smaller dimensions, in particular the depth of the channels: 20 lm compared to 100 lm in our case. The systems presented by Kersaudy-Kerhoas et al. have interesting performance levels. However, they involve certain clogging (platelet aggregates) and lysis risks and


can lead to major head losses precluding experiments with flow rates of several hundred lL/min. In this respect, with the dimensions presented here, it can be seen that all the devices presented here operate at pressure levels lower than 1 bar. These dimensions are the same as for the system used by Sollier et al. which has not been modified (for the separation unit) throughout this study. As mentioned above in the section entitled ‘‘plasma separation principle’’, an essential performance criterion of this system is related to the control of conditions upstream of the sudden converging section. These conditions include, on one hand, the restriction and, on the other hand, the cellfree pre-configuration channel. This channel will be the focus of our attention for the rest of this article. In the first instance, the connection method originally presented by Sollier et al. was used to perform this study successfully. However, for samples that were not very diluted, difficulties arose, such as leaks at the inlets connected by the edge. This is because the blood becomes viscous and difficult to control: the high-pressure levels required to increase the risks of leaks where the glued points connect Tygon to capillary tubes and capillary tubes to PDMS. A more traditional connection mode, which minimises these issues, is presented later in this document.

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Fig. 4 a Picture of the capillary outlet cell-free layer (Øint = 100 lm and L = 5 cm). b Wall shear stress of the inlet capillary tube (Pa) depending on the travel time(s) in various capillary tube for different plasma yields: (white circle) 19 %; (black square) 17 %; (black uppointing triangle) 13 %; (black circle) 7.2 %; (white square) 2.4 %

4.2 Results and discussion 4 Upstream channel: preparation of the cell-free layer 4.2.1 Influence of upstream connections 4.1 Principle The influence of a cell-free layer pre-configuration on extraction performance levels by the upstream conditions is therefore the starting point of this study. This influence can be highlighted in a simple and indirect way by the connection used. For example, when the capillary tubes are not cut at the exact same length, several millimetres can create a significant bias on extraction performance levels. The existence of this cell-free layer can thus be observed at the outlet of the injection capillary tube (Fig. 4a), and as explained below, the progressive deformation of red blood cells occurs directly after this outlet. As mentioned above, to characterise systematically and easily influence these upstream conditions on the cell-free layer, and by extension on the plasma yield of the separation unit, the upstream channel was ‘‘moved outside’’ of the microfluidic PDMS chip, using different types of connections (connected to the chip by the edge). A variety of upstream capillary tube widths will be used; this will allow us to observe extraction yield variations for the same microfluidic design but for different upstream conditions. The range of possible internal diameters is as follows: Øint = 75, 100, 200, and 250 lm for a connection length L = 5 cm and L = 2, 3, 4, and 5 cm for Øint = 100 lm.

In order to interpret the physical aspect of the results, variables must be introduced that can represent the phenomena involved. These variables must of course be controlled by operational parameters (length, diameter, flow rate). As mentioned before, the width of the cell-free layer depends on the flattening of red blood cells because of the shear stress applied to them. In the vicinity of the wall, this shear stress (s) can be estimated from the following formula, for a given fluid velocity (u) in a circular connection with a given cross section (r): du sð y ¼ r Þ ¼ l ð2Þ dy y¼r For a given flow rate, the shear stress intensity, the cross section of the channel, and the transit time for the length of the channel can be calculated. The red blood cell relaxation time is sufficiently long for the red blood cell deformation imposed by geometric conditions (diameter ? length) to be considered as the dominating parameter regarding the configuration of the cell-free layer. Figure 4b presents the results for capillary tubes with different diameters and for different flow rates. Similar results were obtained by modifying the length, which confirms the relevance of the choice of variables.

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According to Fig. 4, for a given blood sample travel time in the chip, the higher the shear stress applied, the higher the yield. Moreover, for a given shear stress, the yield rises as travel time increases. Overview of the proposed mechanism: (1) the yield increases as the cell-free layer thickens, (2) the cell-free layer becomes thicker as the red blood cells get flatter, and (3) the red blood cells get flatter as the shear stress increases and is applied longer. The link between plasma yield and the dimensions of the connections is therefore clearly shown via the mechanism proposed.

5 Implementation: preparation in situ 5.1 Principle To reach optimum operation regardless of the type of connections used, the influence of these connections must be reproduced in situ: the optimum pre-configuration of the cell-free layer will be integrated in the microfluidic component directly. This will also be beneficial in order to be able to place several extraction cells in series. It was decided to use type B connections by the top, because the in-chip integration of the upstream connection benefits

Fig. 5 Drawing of the PDMS microfluidic chip with the 50 mm 9 100 lm serpentine channel at the inlet and the connection by a top interface


means that it is now possible to use flexible and wider Tygon-type capillary tubes. Furthermore, this choice of connection by the top is more versatile than a connection by the edge, making it possible to join the extraction channels inside the chip. Thus, there is only one plasma outlet and therefore one plasma recovery capillary tube (Fig. 5). Possible disruptions, caused for example by plasma recovery capillary tubes that may not be exactly the same length, are thus eliminated. To conclude, it was found that the in-chip integration of the pre-configuration effect of the cell-free layer created by a channel of reduced section (100 lm) but of greater length (50 mm) can be considered only if the surface area occupied by this channel in the chip is limited. This is why the channel has been compressed into a serpentine shape. 5.2 Results and discussion 5.2.1 ‘‘In-situ’’ pre-configuration of the red blood cell-free layer The flow of the blood sample in the upstream channel is illustrated on Fig. 6. A cell-free layer is gradually formed along the channel; as it reaches the restriction, it attains a width that is equivalent to or even greater than that of the previous design. As far as the impact of a serpentine channel is concerned, questions could be raised about the appearance of secondary currents [Dean vortices (Berger et al. 1983)] that occur in curved channels. For certain flow rate and curvature radius conditions, this recirculation may destabilise the flow and, in particular, modify the configuration of red blood cells and therefore the cell-free layer. For these systems, the Dean numbers range from 0.8 to 10: there is

Fig. 6 Pictures of the blood sample flow along the inlet channel, showing the cell-free layer in situ formation. Lower right inset represents the position of the pictures along the channel

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Fig. 7 Width of the cell-free layer in the inlet serpentine channel for an injection flow rate of 100 ll/min. upper (black square), lower (white square), and mean (hyphen) cell-free layer. The switch between upper and lower layers either side of the average value is caused by the centrifugal effect at the outlet of serpentine channel bends. Upper inset is a picture of the two considered cell-free layers

no reason to make provision for any major negative influence on the cell-free layer configuration at the outlet of each bend. However, in Fig. 7, which shows the width of the cell-free layer along a 50-mm long, 100-lm-wide serpentine channel, it can be seen that the presence of bends destabilises the cell-free layer, as the centrifugal force tends to widen the inner cell-free layer and reduce the outer layer at each exit, creating a fluctuation of the cell-free layer width around its mean value (cf. upper and lower widths on Fig. 7). This point and the break in the slope of the mean cell-free width are probably both a sign of the negative impact of curved channels. In this respect, further research is required to optimise the shape of the serpentine channels (curvature radius, number of bends). Even so, the widths of cell-free layers with an ‘‘integrated pre-configuration’’ (i.e. serpentine channels) are comparable and even slightly higher than cell-free widths obtained by the traditional process. Similar experiments have been performed in 50-mm straight channels, and they have provided similar results, i.e. progressive formation of the cell-free layer along the channel. 5.2.2 Result: plasma yield with an integrated preconfiguration Figure 8 presents the plasma yield [the maximum plasma extraction rate g (Eq. 1)] as a function of flow rate; it is shown that the preparation of the cell-free layer with serpentine channel therefore enables pure plasma to be extracted with yields that are equivalent to our reference optimal device (without serpentine channels and with a 100-lm internal diameter inlet capillary tube) presented by Sollier et al. It is also shown that the range of operational flow rates was significantly improved (plasma yield is

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Fig. 8 Plasma yield depending on flow rate for the system with 50-mm inlet serpentine channel and a wide capillary tube (Øint = 350 lm)

approximately constant up to 350 versus 175 lL/min in previous work). This research was used to maintain the optimum plasma yield values for this extraction unit with an integrated preconfiguration that enables wider and more convenient inlet capillary tubes to be used (Tygon-type with Øint [ 300 lm). Having dealt with connection issues by integrating the upstream channel, the next section discusses the advantages of this type of integration in the context of installing several extraction units in series.

6 Installing units in series 6.1 Principle 6.1.1 Protocol One of the possibilities considered to increase the plasma yield of the system is to install several extraction units in series. In the introduction, emphasis was placed on the importance of switching to a portable system that can be used for routine operations and that will reduce preparation times. Managing flow rates with head losses therefore seems to be an unavoidable option. It will also allow the ratio between injected flow rate and extraction flow rate to be fixed once and for all. More precisely, at the outfall of each extraction unit, preset resistance elements (head losses) are provided to ensure flow ratios that correspond to known yields. In order to prepare the cell-free layer, serpentine channels are installed upstream each extraction units, and this also creates head losses that need to be taken into account. Therefore, installing several extraction units in series implies a precise dimensional design of the microfluidic network.

123

176


Qe1 ¼

g1 Q1 2

and

QIN ¼ Q2 þ 2Qe1

Qe2 ¼ and

ð5Þ

Q2 ¼ Qout þ 2Qe2

Re1 Qe1 ¼ R2 Q2 þ Re2 Qe2

Fig. 9 Equivalent circuit diagram showing how two extraction units are installed in series

g2 Q 2 2

and

Rout Qout ¼ Re2 Qe2

ð6Þ ð7Þ

where QIN, Q2, Qe1, Qe2, and Qout represent the respective flow rates in channels R1, R2, Re1, Re2, and Rout. This system of equations is resolved to provide the following: Re1 ¼ 2

6.1.2 Theoretical yield

1 g1 ½R2 þ ð1 g2 ÞRout g1

ð8Þ

1 g2 Rout g2

ð9Þ

and Non-deformable channels and Newtonian viscous fluid are considered for this case. These two hypotheses are at first sight questionable, with the use of PDMS and blood, but experiments have shown that they are rather acceptable for this level of dimensional design. With these simplified hypotheses, the pressure jump (DP) in a channel is proportional to the flow rate (Q): DP ¼ RQ

ð3Þ

For a rectangular section channel, the following head loss coefficient can be introduced: R¼

8lðW þ hÞ2 L h3 W 3

ð4Þ

where L, w, and h, respectively, represent length, width, and height of the channel and l the fluid viscosity. The viscosity of blood is in practice different to that of water (1 9 10-3 Ps). For 1:20 diluted blood, viscosity is very similar to that of plasma (approximately 1.5 times the viscosity of water). For more concentrated blood, viscosity increases (up to a value of 4 9 10-3 Ps), and for 1:10, 1:5, or even 1:2 diluted blood, about a value of 1.5 9 10-3 Ps may be considered (Whittaker and Winton 1933). In each channel section, flows are managed by linear relations between pressures and flow rates. If we make a comparison with electricity, this is equivalent to writing Ohm’s law and likening the fluidic network to an equivalent simple circuit diagram (Fig. 9) composed of two separation units (E1 and E2) installed in series, i.e. meaning that the main outlet of the first unit is connected to the inlet of the next, allowing extraction of residual plasma. Following previous research, a serpentine channel is placed upstream of each extraction point. These two shapes create an overall known resistance denoted R1 for the first extraction and R2 for the second. The extraction channels must feature the following respective resistances—Re1 and Re2—which must be determined in order to guarantee volumetric extraction ratios g1 (for E1) and g2 (for E2). A linear system of equations is then written to impose flow rate ratios (Eq. 5), flow rate conservation (Eq. 6), and the mesh rule (Eq. 7).

123

Re2 ¼ 2

The g1 = 15 % and g2 = 12 % yields are deduced from research from Sollier et al. (2010), creating a total ratio of 25 %. Note that the result of the dimensional design is independent of the fluid viscosity and flow rate. Accordingly, there is no need for a syringe pump or an imposed flow rate pump; the use of a pressure source or even a manual injection system is conceivable. The only free parameter that can be adjusted is Rout. In order to miniaturise the system, the lowest resistance possible will be taken whilst respecting technological manufacturing and design constraints. Once the resistance values are known, the geometry of the microfluidic network can be dimensioned. Once the section of the channels has been set, their length can then be deduced. 6.2 Results and discussion 6.2.1 Fluidic performance levels In order to assess the constraining aspect of the optimum length of the upstream serpentine channel (implied by Eqs. 6 and 7) on the reconfiguration of the cell-free layer upstream of the second extraction unit, three systems with two extraction channels in series have been designed and tested (Fig. 10). The purpose here is to check the parameters of the 2nd serpentine and their influence on performance, in order to find a compromise between a simple device which does not work well and a ‘‘literally taken’’ optimised device. The first design, ‘‘system-1’’, does not comply with this constraint (R2 does not include a serpentine channel). The second design, ‘‘system-2’’, an optimised design, comprises a 100-lm wide, 5-cm-long serpentine channel in compliance with the results obtained in part 3. The first extraction channels are 14 cm in length (cf. Eqs. 6, 7). The last design, ‘‘system-3’’, is a compromise, with a 200-lm-wide serpentine channel that reduces the length of the first


177

Fig. 10 Different designs tested and pictures of extractions E1 and E2 for various flow rates: a 100 ll/ min; b 400 ll/min. system-1, simple, without 2nd serpentine channel; system-2, optimised, with two 100-lm-wide serpentine channels; system-3, compromise with 2nd serpentine channel 200 lm wide: compromise providing a lower head loss. Information on pollution is given

extraction channel to 4.3 cm whilst also reducing the head loss of the whole component. The surface area of the three designs is almost identical, i.e. 4 9 1.8 cm2. For the three cases, the plasma extraction rates imposed (15 and 12 % for E1 and E2, respectively) lead to a significant increase in the overall volumetric extraction ratio from 17.5 to 25 %. System-1, which was designed to be more compact, shows the relevance of complying with the above-mentioned constraint: indeed, due to the absence of the 2nd serpentine, the 2nd extraction (E2) is significantly

contaminated by RBC (cf. Fig. 10), which is symptomatic of the cell-free layer that has not been prepared appropriately. Systems 2 and 3 provide similar results, which turned out to be satisfactory for the two extraction stages (E1 and E2). By controlling flow rates with head losses, it was possible to simplify the experimental protocol by avoiding the use of outlet syringe pumps and by making the system performance levels (in this case, the flow rate ratio) independent of the injection flow rate. The system (Fig. 11) has therefore been validated as robust for flow rates up to at least 400 lL/min. In practice, a simple inlet pressure

123

178

Fig. 11 Picture of the experimental PDMS device (system-3)

(applied manually for example) guarantees plasma extraction with a volumetric ratio of 25 % for 1:20 diluted blood. 6.2.2 Plasma quality In order to observe the quality of the recovered plasma, the purity of extraction channels (contamination by red blood cells) was first assessed qualitatively using a visual inspection method (see photos in Fig. 10); then, measurements were taken using the Millipore Scepter, which Fig. 12 Proteomic profile by mass spectrometry on the plasma obtained from an identical donor, extracted using our microfluidic device and centrifuging methods for a low masses and b high masses

123


helped quantify red blood cell contamination. The absence of haemolysis was also checked, and lastly, the proteomic profile was compared to that of the plasma obtained by centrifuging. The contamination measurements showed excellent purity of the recovered plasma: for 1:20 diluted blood, the contamination level compared with the injected sample was lower than 1 % (cell concentration in recovered plasma measured with the Scepter was around 102 cells per ll, comparable to concentrations measured in plasma obtained by centrifuging for an identical sample). For a 1:10 diluted injected sample, the design seems to be relatively efficient: contamination compared to the injected sample is also lower than 1 % (concentration of approximately 103 cells per ll), only slightly higher than concentration levels measured in plasma obtained by centrifuging for an identical sample. So as to ensure that our system does not impose higher than reasonable mechanical stress on the red blood cells, the Cripps method was used to assess the percentage of haemolysis of plasma stemming from the system with 1:20 diluted blood and compared it to the plasma from the same samples but obtained by traditional centrifuging. Our system apparently does not lead to any red blood cell lysis, irrespective of the flow rate used. The results are perfectly comparable with those obtained for plasma extracted by centrifuging.


The proteomic profiles of plasma extracted by our microfluidic system and by a centrifuging process (1,500 g, 5 min) were also compared for several donors (1:20 diluted sample). The protein was analysed by mass spectrometry using the SELDI technology with a CM10 chip (positive ion exchanger) for a pH of 4.5. Figure 12 below presents the representative spectra and shows that the two extraction methods are equivalent. Indeed, the spectra obtained for the plasma extracted using the microfluidic chip or by centrifuging are very similar. This proves on one hand that there is no protein absorption in the system and, on the other hand, that there is no haemolysis. Furthermore, it was shown that there is no visible influence of the injection flow rate on the result.

7 Conclusion This research has shown the importance of the choice of fluidic connection system in terms of the performance level of a separation system to prepare blood samples. This connection has an influence on the configuration of flow upstream of a separation unit, more precisely on the thickening of the cell-free layer. And yet, the latter has a direct influence on the performance level of the system. To our knowledge, the role of the connection has not been discussed to any great extent in the existing literature. The attention paid to these upstream conditions has also allowed us to optimise the in-series installation of a microfluidic design to extract plasma from blood. This in-series installation process is not simply a repetition of the same patterns: successive optimum adjustments of the channels upstream of the separation unit and of extraction channels are nonetheless necessary and have been calculated. Thanks to this research, it was possible to extract 25 % of the injected volume (instead of 17 % with only one pattern), with a relatively low contamination rate (approximately 0.1 %). The purity of the recovered plasma is excellent, even for 1:10 diluted blood. The proteomic profile of the recovered plasma was observed using the mass spectrometry method (SELDI), and the results obtained are identical to those with plasma obtained by centrifuging. Furthermore, the absence of haemolysis has been checked in plasma samples recovered with our system. The experimental protocol has also been simplified and improved: the system is compatible with high pressure, and it can then tolerate flow rates (up to about 400 lL/min), allowing large volumes of blood (measured in ml) to be treated. Moreover, flow rate ratios are imposed by head losses, and only one input pressure system is now required (for example an imposed pressure by hand). This was achieved by the choice of minimum channel dimensions

179

(50 lm), which was also required in order to avoid channel clogging and red blood cell lysis risks. Finally, this research provides a concrete contribution to the development of systems based on the physical mechanism of inertial focalisation. Acknowledgments This work was supported by the CEA and more especially the Technologies for Health programs. Special thanks are also dedicated to Myriam Cubizolles for providing proteomic profile by mass spectrometry on the plasma samples as well as helpful advices.

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