Palaniappan
Sethu
,
Aaron
Sin
and
Mehmet
Toner
Surgical Services and Center for Engineering in Medicine, Massachusetts General Hospital, Harvard Medical School and Shriners Hospital for Children, MA 02114, Boston, USA. E-mail: spalania@sbi.org; asin@partners.org; mtoner@sbi.org; Fax: (617) 724 2999; Tel: (617) 726 6465
First published on 11th November 2005
Apheresis is a procedure used to fractionate whole blood into its individual components. Following fractionation, the desired component is isolated and the remaining blood in many cases is returned to the donor. Leukapheresis is one type of apheresis where leukocytes (white blood cells) are selectively removed. This procedure is commonly used for blood transfusions to remove donor leukocytes from being transferred to the recipient. Apheresis also has several therapeutic applications. In this manuscript we discuss the design, fabrication and testing of a continuous flow diffusive filter, fabricated using simple soft lithographic techniques for depletion of leukocytes. This device employs micro sieves that exploit the size and shape difference between the different cell types to obtain depletion of leukocytes from whole blood. Currently, conventional apheresis methods like centrifugation or fiber mesh filtration are commonly used. A theoretical model was developed to determine the optimal shape of the diffuser to ensure that the volumetric flow through individual sieve elements is equal. This device was designed to serve as a passive device that does not require any external manipulation. Results show that for the given device design, isolation of ∼50% of the inlet erythrocytes (red blood cells), along with depletion of >97% of the inlet leukocytes is possible at a flow rate of 5 µl min−1. Simple modifications to the geometry and dimensions of the sieves can be made to obtain isolation of plasma.
Apheresis is a general term describing removal of blood from a subject; a portion of the blood is fractionated and retained while the rest is returned to the donor. Leukapheresis (or leukodepletion) is a procedure where the leukocytes are isolated or fractionated. Other types of apheresis include plasmapheresis which is a procedure where plasma is separated and manipulated in a variety of ways and peripheral stem cell isolation, a procedure in which stem cells are isolated and retained.
Leukapheresis1,2 is typically performed to decrease a very high white blood cell count in individuals with leukaemia or to remove white blood cells for transfusion. Studies have repeatedly shown that leukocyte contamination during blood transfusion may result in adverse effects3 like febrile transfusion reactions, graft versus host disease, transmission of infectious agents like viruses (cytomegalovirus (CMV), herpes virus, T-cell leukaemia/lymphoma virus), bacteria, toxoplasma gondii, prions, platelet refractoriness and transfusion related immunomodulation. Therefore, the current trend is to ensure that blood used for transfusion is leukodepleted.
Cells in whole blood are predominantly erythrocytes that account for ∼99% of the total cells in blood. Leukocytes make up less than 1% of blood and have three major subpopulations: lymphocytes, monocytes and granulocytes. By taking advantage of the difference in physical properties (Table 1) between the different cell types, selective depletion of leukocytes from whole blood can be accomplished. Erythrocytes are smaller, biconcave in shape and more flexible than the leukocytes. Therefore, the most commonly used technique is filtration using a polymer or wire mesh filter with pore sizes in the order of 3–4 µm. The differences in mass densities of blood cells can also be exploited to achieve separation and depletion of lymphocytes and monocytes from whole blood using methods like differential centrifugation.
Cell type | Concentration/cells ml−1 | Diameter/µm | Surface area/µm2 | Volume/µm3 | Mass density/g cm−3 |
---|---|---|---|---|---|
Erythrocytes (red blood cells) | 4.2 − 5.4 × 109 | 6–9 | 120–163 | 80–100 | 1.089–1100 |
Leukocytes (white blood cells) | 0.4 − 1.1 × 107 | 6–10 | 300–625 | 160–450 | 1.055–1.085 |
Neutrophils | 2 − 6 × 106 | 8–8.6 | 422–511 | 268–333 | 1.075–1.085 |
Eosinophils | 0.4 − 4.8 × 105 | 8–9 | 422–560 | 268–382 | 1.075–1.085 |
Basophils | 0 − 1.1 × 105 | 7.7–8.5 | 391–500 | 239–321 | 1.075–1.085 |
Lymphocytes | 1 − 4.8 × 106 | 6.8–7.3 | 300–372 | 161–207 | 1.055–1.070 |
Monocytes | 1 − 8 × 105 | 9–9.5 | 534–624 | 382–449 | 1.055–1.070 |
Thrombocytes (platelets) | 2.1 − 5 × 108 | 2–4 | 16–35 | 5–10 | 1.04–1.06 |
There have been several attempts to create microfluidic devices for cell sorting applications. Including applications where the focus is on sorting blood cells. Attempts have been made by different groups using techniques like filtration,4,5 obstacles,6 electrophoresis,7 dielectrophoresis,8 immunomagnetic separation,9 margination,10 laminar flow and surface acoustic waves.11 These techniques are either focussed on low throughput processing for isolation of small number of cells or are not as efficient as bulk techniques that use filters or centrifugation (both continuous centrifugation and non-continuous) where ∼99% of the leukocyte population is removed.
We report the fabrication of a diffusive filter for size based depletion of leukocytes from whole blood (leukapheresis). This device exploits the size and shape difference between different cells to obtain the depletion of leukocytes. The sieves were designed to allow the passage of the biconcave erythrocytes, while providing a barrier to the larger spherical leukocytes. To prevent clogging and facilitate continuous separation, the sieves were arranged on the sides of the channels, connecting the channel to a diffuser. In order to ensure equal volumetric flow through every filter element of the sieve, the shape of the diffuser was modified resulting in a flared geometry. The device operates in the continuous flow mode where blood can be continuously fractionated. The leukocyte depleted blood is isolated and the remaining blood can be returned to the source or donor. This device is passive and does not require any external manipulation. Simple modifications to the sieve size and geometry can be made to isolate plasma.
Whole blood was delivered into the inlet of the device using a syringe pump. Blood was introduced at flow rates between 5–12 µl min−1. The blood contains different cells suspended in plasma which enter the main channel and encounter the 2.5 µm × 40 µm sieves that line the walls of the main channel.
The experiments were performed with a sample size of N = 5 and all the data is represented as means ±SE.
A schematic of the designed device is illustrated in Fig. 1. Whole blood enters the device at the inlet and flows through the main channel. The dimensions of the main channel are 2 cm × 200 µm × 50 µm (L × W × H). This channel is connected to a diffuser on both sides through sieves with filter elements as shown in Fig. 1 (insert). The dimensions of each filter element are 40 µm wide × 2.5 µm high (W × H), and each device has a total of 500 filter elements. For filter elements less than 1 µm tall, soft lithographic techniques may be unsuitable. Other materials like silicon or glass may be necessary. The geometry (Fig. 1, insert) of the filter elements accomplishes the design objective that only the biconcave erythrocytes pass through the sieves into the diffuser and the larger spherical leukocytes are contained in the main channel under normal flow conditions. Therefore if a cell encounters a sieve element and cannot pass through it is pushed along the main channel to the outlet due to the flow in the main channel. For this particular device design, at flow rates greater >5 µl min−1, the pressure gradient across the filter elements becomes large enough to deform and force some of the leukocytes through into the diffuser.
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Fig. 1 Device design: (a) Schematic of the diffusive filter for size based continuous flow fractionation of erythrocytes from whole blood. Insert shows the 40 µm × 2.5 µm sieve structure and the arrangement connecting the main channel to the diffuser. (b) Phase contrast microscopic images of the top view of the device at different locations with a magnified image of the 2.5 µm × 40 µm sieves. |
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The diffuser and main channel were broken up into a series of rectangular blocks to simplify the theoretical model as shown in Fig. 2. The device can be represented as a series of independent elements arranged in a network as shown in Fig. 3(a). Each element is associated with a fluidic resistance. The volumetric flow through each element depends on the fluidic resistance of that element and the corresponding pressure difference or pressure gradient across that element (ΔP). The network can be further simplified into discrete repetitive modules as shown in Fig. 3(b). Each module represents the portion of the device between filter elements m and m + 1 (where m = 1 to n
− 1, and n is the total number of filter elements). The module consists of the section of the main channel and the diffuser on either side connected by the filter elements m and m + 1. Since the device is axially symmetrical the module can be further reduced to the closed loop shown in Fig. 3(c). In order to ensure constant volumetric flow ‘Y’ through the each of the filter elements, the fluidic resistance has to be controlled by appropriate diffuser design to provide the necessary resistance. If ‘X’ defines the output flow through the outlet of the main channel (outlet 2) then the ratio of flow through a corresponding set of filter elements to output flow through the main channel ratio can be defined by a dimensionless number ‘’, where:
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Fig. 2 Simplification of device geometry: In order to simplify the design of the device the diffuser was broken down as a series of rectangular blocks such that the fluidic resistance of each block can be easily determined. |
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Fig. 3 Device modelling: (a) Microfluidic circuit represented as a network of resistors with each resistor representing the resistance to fluid flow. X and nY are outputs at the outlets and X + 2nY is the volume of sample through the inlet. (b) This network can be divided into repetitive modules with each module containing a part of the channel, diffuser and two sets of filter elements. (c) Each module can be further reduced to a closed loop where ΔPc, ΔPd, ΔPm and ΔPm+1 are the pressure gradients across the section of the channel, diffuser and filter elements m and m + 1. |
Solving for the condition that the volumetric flow through each and every filter element is a constant ‘Y’, we can then obtain an approximation for the width of the diffuser element wdiffuser for each discrete module using the equation shown below. This breaks down the diffuser into a series of rectangular blocks with increasing widths thereby varying the fluidic resistance appropriately, ensuring uniform volumetric flow through each and every filter element.
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Fig. 4 Simulation results: Finite volume modelling and simulation of the absolute velocity of fluid flow through each sieve element at a height of 1.25 µm from the bottom of the device. The simulation was performed for a device with 100 sieve elements. Normalization of flow velocities to a certain extent of the fluid flow through each sieve element can be achieved using the modified design. |
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Fig. 5 Modified design: Shown are the volumetric flow rates through sieve elements for a device with 100 sieves. The flow through each sieve element depends on the shape of the diffuser. Simulation results show that the using the modified design the volumetric flow rate through individual sieves can be normalized to a certain extent. |
Devices were fabricated and tested. From experimental results it can be seen that for smaller flow rates the filter elements allow only the biconcave erythrocytes oriented parallel to the floor of the channel to pass through unobstructed. The leukocytes are spherical and larger than the cross sectional area of the sieves and can only pass through with significant deformation. For flow rates >5 µl min−1 the pressure gradient across the sieves becomes large enough to deform and force some leukocytes to pass through the sieves.
In the experiments performed, 0.5 ml of whole blood was introduced into the device at different flow rates and after transit through the device the samples were collected at the 3 outlets. Sample from outlets 1 and 3 contain the fraction of sample that passed through the sieve structure into the diffuser whereas the sample from outlet 2 contains the fraction of sample that transited through the main channel alone. The sample from outlets 1 and 3 were combined and compared to the sample from outlet 2. The experiments were performed at sample inlet flow rates of 5, 7 and 12 µl min−1. The samples were collected and cell counts were performed on each sample. Fig. 6 shows the erythrocyte and leukocyte counts obtained for the sample collected from each of the outlets for the three different flow rates. The percentages were obtained by comparing the cell counts at the outlet to the original sample. The device and tubing itself account for a very small source of cell loss (∼1%), but at flow rates of 7 and 12 µl min−1 it can be seen from Fig. 6 that there is loss of cells, possibly due to lysis caused by the high shear experienced by the cells in the filter elements of the sieves.
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Fig. 6 Erythrocyte fractionation results: Erythrocyte fractionation obtained at operating flow rates of 12, 7 and 5 µl min−1. Optimal fractionation occurs at 5 µl min−1; outlets 1 and 3 contain ∼50% of the total erythrocytes while containing only <3% of the total leukocytes. At higher flow rates the pressure difference across the sieves improves the erythrocyte yield through outlets 1 and 3. This happens at the cost of increased leukocyte contamination. |
The fluid flow through the device is controlled using a constant flow rate syringe pump. The pump delivers blood into the inlet of the device and travels through the device into the main channel and through the sieves on either side into the diffuser. The volume of blood flow through the main channel and through the sieves depends on the fluidic resistance and the pressure difference. The pressure difference between the main channel and the diffuser drives fluid through the sieves into the diffuser. The fluidic resistance of the sieve-diffuser combination determines the ratio of the volume of blood obtained at the different outlets. In order to ensure that the volumetric flow through individual filter elements is uniform the diffuser was designed with a flared geometry to create a uniform volumetric flow through each filter element. This model works well for fluids. However, for fluids with particles/cells in suspension the operating conditions have to be sufficiently controlled to ensure proper operation.
The device takes advantage of the geometrical differences between erythrocytes and leukocytes to obtain fractionation. The operating flow rates must be optimized such that the pressure gradient across the filter elements does not support the passage of leukocytes into the diffuser. Erythrocytes that are oriented parallel to the floor of the channel can pass through the filter elements without any obstruction. Leukocytes on the other hand can only pass through the filter elements by deforming and squeezing through the filter elements. At flow rates <5 µl/min the pressure gradients across the sieves is not sufficient to support forced transport of leukocytes across the filter elements. However, at higher flow rates the pressure gradient necessary for forced leukocyte transport is available. Further, in order to ensure that the volumetric flow through all of the filter elements was equal an analytical model was developed. An expression to determine the width and hence the optimal shape of the diffuser was determined to provide the necessary fluidic resistance for the given pressure difference to ensure uniform volumetric flow through the sieves. Further, computational fluid dynamics simulations were performed on CFDRC finite volume software. Navier–Stokes simulations were then carried out using the algebraic multigrid (AMG) solver to at least 4 orders of magnitude residual reduction, within ∼1000 iterations. The results show that optimization to a certain extent can be accomplished using this mathematical model. This compares favorably with a linear diffuser (the flared geometry replaced with a linear geometry).
The expression to determine the width of the diffuser ‘wdiffuser’ depends on: the total number of filter elements ‘n’, the width of the channel ‘wchannel’ and the flow ratio ‘’. ‘
’ is the ratio of flow through the main channel to the flow through a single set of filter elements. This ratio can be fixed to obtain the desired device performance. This way we can control the percentage of fluid obtained through the main channel and percentage of fluid obtained through the outlets of the diffusers. Since our end goal is to filter out all erythrocytes we set this ratio low (as in our case 1∶4). The observed value of ‘
’ is higher than the set value due to the nature of the blood. The flow through the filter elements ‘Y’ becomes smaller due to obstruction or temporary clogging of the sieves due cells at the interface of the main channel and the filter elements (erythrocytes that are improperly oriented and larger leukocytes). The high concentration of cells in blood ensures that erythrocytes are in most cases not oriented parallel to the floor of the channel. The cells near the walls of the main channel are drawn to the sieves due to the pressure gradient that exists across the filter elements. Erythrocytes, if oriented properly travel through the sieves without any delay. Whereas, erythrocytes not oriented properly pause and reorient themselves in order to travel through the sieve. Both leukocyte and erythrocyte interaction with the filter elements strongly depends on the pressure gradient across the filter elements, which is a function of the operating flow rate. At low flow rates, the interaction of leukocytes with the filter elements is weak and the flow through the main channel is sufficient to break this interaction and move them along the main channel. At faster flow rates leukocytes interact strongly with the filter elements and can be forced through into the diffuser and even erythrocytes not oriented properly are forced through the filter elements.
Experiments to determine the fractionation efficiency of the device were performed. Leukocyte depleted blood according to the World Apheresis Association (WAA) is defined as a sample of blood containing <1 × 104 leukocytes ml−1 which corresponds to ∼99% depletion of leukocytes. The experiments were performed at three different flow rates and the recovery of erythrocytes and leukocytes through each of the outlets was documented. As can be seen from the results we can achieve close to 99% depletion of leukocytes and isolation of ∼50% of the total erythrocytes at a flow rate of 5 µl min−1 where our device compares well with commercially available devices (∼99% depletion) for leukocyte depletion used for blood transfusion. Higher flow rates result in isolation of a larger percentage of erythrocytes at the cost of greater leukocyte contamination of the sample.
By reducing the height of the filter elements from 2.5 µm to <0.5 µm this device can be used for fractionation of plasma or plasmapheresis. The use of elastomeric materials might not be suitable for filter elements with sub-micron heights necessary for plasmapheresis. Rigid materials like silicon, glass or polymers with high glass transition temperatures might be necessary. Plasmapheresis applications include plasma exchange (PE), extracorporeal immunoadsorption using columns and low density lipoprotein (LDL) apheresis. Plasmapheresis for therapeutic applications or TPE is commonly performed in patients with autoimmune diseases. TPE can be used to treat both acute and chronic conditions of autoimmune diseases to provide temporary relief by depletion of certain immune factors and complexes.
This journal is © The Royal Society of Chemistry 2006 |