Microfluidic platforms have the great potential to change the way medicine and biology is conducted in hospitals and laboratories. Such platforms have attracted considerable research interest due to the opportunity to fabricate a highly integrated system that is able to perform parallel sample handling and analysis on a single chip (i.e., lab-on-a-chip). Although microfluidics promises to have an impact in many research fields, one of the more attractive applications of microfluidics has been toward biomedical and life science diagnostics.1–3 Microfluidic devices are attractive because they offer many advantages such as smaller reagent volume consumption, shorter reaction times, and the possibility of parallel operation. These advantages result not only in time and cost savings for diagnostic tests, but can also be life-saving in time-critical environments such as critical medical diagnostics. For instance, when performing blood analysis in medical laboratories, the blood cells are separated from whole blood by centrifugation, and plasma is analyzed for electrolyte concentration, glucose, lactate, and total cholesterol, among other parameters. These sequential procedures can take up to several hours. However, there is a need for a system that can separate blood plasma from whole blood and measure the concentration of the clinically relevant proteins in real time. For example, several studies have clearly shown that cardiac surgery induces systemic inflammatory responses, particularly when cardiopulmonary bypass (CPB) is used.4–13 These systemic responses are attributed to several factors, including exposure of blood to nonphysiologic surfaces of the heart-lung circuit, ischemia-reperfusion of the involved tissues, surgical trauma, and hypothermia.4,13,14 Currently, there is no effective method to prevent this systemic inflammatory response syndrome in patients undergoing CPB. The ability to clinically intervene in inflammation, or even study the inflammatory response to CPB, is limited by the lack of timely measurements of inflammatory responses. Current technology provides measurements of the effects of CPB on activation of complements, neutrophils, platelets, and cytokines hours or days after surgery. However, more immediate measurements would help us to understand the mechanisms of cellular activation, and to modify surgical and perfusion protocols that would minimize the adverse effects of CPB. Currently, there are several reports related to blood-handling microdevices.15,16 In this study, a microfluidic device for continuous, real-time blood plasma separation, which may be integrated with a downstream plasma analysis device, is presented. The device is made out of polydimethylsiloxane (PDMS), which is considered to be a hemocompatible material. Even though no truly hemocompatible biomaterial has yet been found, PDMS is assumed to be a suitable biomaterial for the experimental devices because it causes minimal endotoxin contamination, leukocyte activation, and complement activation.17,18 It is also assumed that the activation of inflammatory responses caused by the PDMS (foreign surface) itself can be ignored when the blood exposure time to PDMS is short (i.e., seconds).
Principle of the Blood Plasma Separation Microdevice
The principle of the plasma separation device is based on the bifurcation law,19,20 also called the Zweifach-Fung effect. The bifurcation law describes that, in the microcirculation, when erythrocytes flow through a bifurcating region of a capillary blood vessel, they have a tendency to travel into the daughter vessel that has the faster flow rate, leaving very few cells traveling into the slower-flow-rate channel. The critical flow rate ratio between the daughter branches for this cell separation is on the order of 2.5:1 when the cell-to-vessel diameter ratio is on the order of 1. The reason for this apportioning is that cells are drawn into the higher-flow-rate vessel because they are subjected to a higher pressure gradient. In addition, the asymmetric distribution of shear forces on the surface of a particle produces a torque on the cell, pulling it toward the faster-flow-rate vessel. In a previous report,21 particle separation based on the bifurcation law was successfully demonstrated using a simple bifurcation channel by controlling flow rate ratio between two daughter microchannels. It was found that all particles travel into a faster daughter channel when the flow rate ratio between two daughter channels is more than 6:1 for both 16-μm-diameter fluorescent particles and 8–10-μm-diameter human C8161 melanoma cells in a 35 × 35 μm2 channel.
For this study, a blood plasma separation device is designed to have a whole blood inlet, a purified plasma outlet, and a concentrated blood cell outlet as shown in Figure 1. Each channel is designed to be 5 mm long. The main blood channel is designed to be 15 μm wide and all plasma-skimming channels are designed to be 9.6 μm wide. The channel depth for the entire device is 10 μm. This device is designed to separate plasma from whole blood with up to 45% hematocrit of the inlet blood, and a 14:1 flow rate ratio is assumed as a required flow rate ratio between main channel and plasma channel to form a critical separation stream-line, which occurs 1 μm from the main channel wall for a 15-μm-wide channel. Also, to increase the plasma-skimming volume for whole blood, five parallel plasma channels are considered in design. The bifurcation positions are indexed from 1 to 5.
The microfluidic device was fabricated in PDMS using the conventional soft-lithography process.22,23 In soft lithography, a master mold is first made by lithographic techniques and followed by silicon etching to form a negative imprint of the channels on a silicon wafer. Next, an elastomeric material (PDMS) is cast from this master (Si) mold to obtain a positive pattern. After forming a positive PDMS structure, the PDMS microchannels are bonded to another piece of PDMS to form a closed capillary structure. After preparing the device, it is mounted on an inverted microscope for visualization and defibrinated sheep blood is infused through the device using syringe pump (KDS210; KD Scientific Inc., Holliston, MA). In all experiments, defibrinated sheep (hematocrit = 39%) blood (Hemostat Labs, Inc.) was used as a test fluid. The blood was used within 3 weeks of harvest and stored at 4ºC. All experiments are conducted for 30 minutes at a 10-μl/h blood inlet flow rate. The blood cells in microchannels are imaged using two different types of CCD cameras. The first is a conventional CCD camera, which has a 640 ×1 480 pixels resolution and a frame rate of a 30 frames/s. The other one is with a high-resolution (1,376 × 1,040 pixels) CCD camera (Cooke SensiCam QE, The Cooke Corp., Romulus, MI) at a 10 frames/s, a shutter time of 100 μs/frame.
Results and Discussion
Electrical Circuit Analysis of the Blood Plasma Separation Microdevice
In a manner analogous to an electrical current, fluidic flow rate (Q) can be defined by the following formula:
where R is the fluid flow resistance and ΔP is a pressure drop. Thus, the flow rate ratio at each bifurcation can be determined by controlling flow resistance ratio between main channel and plasma channel at each bifurcation. For calculating flow resistance of each channel, an analogous electrical circuit is analyzed where the flow rate is modulated by the current and the flow resistance is modulated by an electrical resistance as described in Figure 2. The flow rate of whole blood at the inlet can be assumed as a current source. Plasma outlet and concentrated blood cell outlet can be assumed as grounds because they are at atmospheric pressure. Also, individual fluidic channels can be modulated as electrical resistors. Figure 3 shows an analogous electrical circuit simulation result. When the resistances of plasma channels are 210 times larger than that of individual segments of main channel, a 14:1 current ratio (flow rate ratio) is obtained.
Analytical Study of the Blood Plasma Separation Microdevice
Once the optimal flow resistance ratios are obtained from the electrical circuit analysis, an analytical study was conducted to determine optimal fluid channel dimensions that meet the flow resistance ratio requirements. Equation 2 describes an analytical solution of flow rate for rectangular duct24:
Where μ is a fluid viscosity, ΔP is a pressure drop across a channel length L, a is a half width of a channel, and b is a half depth of a channel. By combining Equations 1 and 2, the flow resistance relationship is,
By calculating the ratio of the plasma channel resistance to the main channel resistance, the optimal plasma channel dimensions, which produce the required flow rate ratio, is obtained. Figure 4 shows the flow resistance ratio between the plasma channel and the main channel with respect to the plasma channel width, when the main channel width and length are fixed as 15 μm wide and 20 μm long, respectively. A 9.6-μm-wide plasma channel is required to get a 210-times larger resistance in the plasma channel than that in the main channel.
A Numerical Study of the Blood Plasma Separation Microdevice
Once the blood plasma separation device is designed to skim plasma from up to 45% hematocrit of the inlet blood with 14:1 flow rate ratio at each bifurcation, a numerical simulation was conducted to evaluate performances of the designed blood plasma separation device. For this numerical study, CFDACE (CFD Research Corporation, Huntsville, AL) is used. In all simulations, the blood is assumed as a homogeneous non-Newtonian fluid which follows the Walburn-Schneck blood model.25 As measures of the performance of the designed blood plasma separation microdevice, flow rate ratios and plasma skimming volumes at each bifurcation are obtained by calculating mass flow rates at each bifurcation for 45% and 39% of whole blood inlet hematocrits, respectively. Figure 5a plots the mass flow rate at the main channel and the plasma channel at each bifurcation. The mass flow rate for both the main channel and plasma channels decreased as blood flowed through a separation region because of plasma skimming through the plasma channels. The mass flow rate in the main channel also increased with decreasing hematocrit because the flow rate is inversely proportional to hematocrit. Another important value that has to be determined is the performance of the blood plasma separation device. The volume percent of the plasma skimming at each bifurcation was determined. As shown in Figure 5b, the volume percent of the plasma skimming decreased as blood traveled through the device because the total amount of blood remaining in the main channel after passing a bifurcation decreases as blood travels through the device. The volume percent also decreased as blood hematocrit decreased because the volume percent of the plasma skimming is directly proportional to mass flow rate in plasma channels. By summing the plasma-skimming volume percent of each bifurcation, 27% and 25% of total plasma skimming is expected with the infusion of 45% and 39% hematocrit blood at the inlet, respectively.
An Experimental Study of the Blood Plasma Separation Microdevice
As a final stage of this study, a blood plasma separation experiment was conducted using experimental setup as described. Figure 6a shows a photograph (20× magnification, 640 × 480 pixel resolution) of a device showing the blood plasma separation region after infusing the defibrinated sheep blood (39% hematocrit) through the whole blood inlet at a flow rate of 10 μl/h. The flow rate corresponds to the average velocity of 18.5 mm/s in the main channel. Because the longest traveling length of the blood in the microdevice is about 10 mm (data not shown), the blood infused through the inlet has a residence time within the device of less than 1 second. Thus, the additional activation of the inflammatory response caused by the exposure of blood to PDMS (microdevice) can be ignored because it is improbable that the activation of markers (e.g., cytokines, complements) of inflammatory response will be seen within such a short time (<1 second). The movement of blood cells is imaged using a conventional CCD camera. Because of the high cell velocity, the images appear as streak lines at 30 frame/s. The streak lines show the cell flow profile. It is clear that all blood cells move into the concentrated blood cell outlet without any cells traveling into a purified plasma bifurcation. To observe the motion of the individual blood cells, images were taken using a higher-resolution (1,376 × 1,040 pixels) CCD camera with a shorter shutter open time (100 μs) at 40× magnification. As shown in Figure 6b, blood cells are more clearly observed, whereas no cells are moving into the purified plasma outlet. A severe deformation of red blood cells was not observed at the given flow rate. During 30 minutes of continuous infusion of blood, only one platelet flowed into the purified plasma outlet. Therefore only the plasma is skimmed into the purified plasma outlet. In addition, no clogging or lysis of cells was observed. These results demonstrate that continuous, real-time blood plasma separation could be accomplished by controlling the flow rate ratio at each bifurcation.
A microfluidic device for blood plasma separation was successfully designed and demonstrated. To find the optimal flow resistance at each bifurcation, an analogous electrical circuit analysis was successfully applied. A 210-times larger flow resistance is required in the plasma channel to obtain a 14:1 flow rate ratio between the main channel and the plasma channel. Based on the flow resistance ratios obtained by the electrical circuit analysis, analytical and numerical studies were successfully conducted. From the analytical study, the dimensions of the blood plasma separation microdevice were determined. From the numerical study, we determined that 27% and 25% of the plasma volume could be collected from total volume of blood with the infusion of 45% and 39%, respectively, hematocrit blood at the whole blood inlet. The device’s functionality was demonstrated using defibrinated sheep blood (hematocrit = 39%). During the 30-minute experiment, all blood cells traveled through the device toward the concentrated blood outlet whereas only the plasma flowed toward the plasma outlet without showing any clogging or lysis of cells. Future work will include a quantification of blood hematocrit at the inlet and outlet using image processing to validate computation models and to determine skimming volume percent. Also, blood flow profiles at bifurcation will be studied using microparticle image velocimetry. Future work will focus on determining the hemocompatibility and separation efficiency of sampling anticoagulated/heparinized human blood from a mock CPB circuit through the separation devices. Because of its simple structure and control mechanism, this microdevice is expected to be used for highly efficient, real-time, continuous cell-free plasma separation.
This project is funded, in part, under a grant with the Pennsylvania Department of Health using Tobacco Settlement Funds. The Department specifically disclams responsibility for any analyses, interpretations or conclusions.
1.Manz A, Graber N, Widmer HM: Miniaturized total chemical analysis systems: A Novel concept for chemical sensing. Sens Actuat B
1: 244–248, 1990.
2.Polla DL, Erdman AG, Robbins WP, et al: Microdevices in Medicine. Annu Rev Biomed Eng
02: 551–576, 2000.
3.Voldman J, Gray ML, Schmit MA: Microfabrication in biology and medicine. Annu Rev Biomed Eng
1: 401–425, 1999.
4.Asimakopoulos G: Mechanisms of the systemic inflammatory response. Perfusion
14: 269–277, 1999.
5.Birdi I, Caputo M, Underwood M, et al: The effects of cardiopulmonary bypass temperature on inflammatory response following cardiopulmonary bypass. Eur J Cardiothorac Surg
16: 540–545, 1999.
6.Fosse E, Mollnes T, Ingvalden B: Complement activation during major operations with or without cardiopulmonary bypass. J Thorac Cardiovasc Surg
93: 860–866, 1987.
7.Kirklin J, Barratt-Boyes B (ed): Cardiac Surgery. New York: Elsevier, 1993.
8.Kirklin J, Chenoweth D, Naftel D: Effects of protamine administration after CPB on complement, blood elements, and the hemodynamic state. Ann Thorac Surg
41: 193–199, 1986.
9.Kirklin J, Westaby S, Blacstone E, et al: Complement and the damaging effects of cardiopulmonary bypass. J Thorac Cardiovasc Surg
86: 845–857, 1983.
10.Moore F Jr, Warner K, Assousa S: The effects of complement activation during cardiopulmonary bypass: attenuation by hypothermia, heparin, and hemodilution. Ann Surg
208: 95–103, 1988.
11.Schlag G, Redl H, Hallstrom S: The cell in shock: The origin of multiple organ failure. Resuscitation
21: 137, 1991.
12.Steinberg JB, Kapelanski DP, Olson JD: Cytokine and complement levels in patients undergoing cardiopulmonary bypass. J Thorac Cardiovasc Surg
106: 1008–1016, 1993.
13.Westaby S: Organ dysfunction after cardiopulmonary bypass: a systemic inflammatory reaction initiated by the extracorporeal circuit. Intensive Care Med
13: 89–95, 1987.
14.Utley JR: Pathophysiology of cardiopulmonary bypass: current issues. J Card Surg
5: 177–189, 1990.
15.Carlson RH, Gabel CV, Chan SS, et al: Self-sorting of white blood cells in a lattice. Phys Rev Lett
79: 2149–2152, 1997.
16.Gifford SC, Frank MG, Derganc J, et al: Parallel microchannel-based measurements of individual erythrocyte areas and volumes. Biophys J
84: 623–633, 2003.
17.Bélanger M-C, Marois Y, Raynald R, et al: Selection of a polyurethane membrane for the manufacture of ventricles for a totally implantable artificial heart: Blood compatibility and biocompatibility studies. Artif Organs
24: 879–888, 2000.
18.Gorbet MB, Yeo EL, Sefton MV: Flow cytometric study of in vitro neutrophil activation by biomaterials. J Biomed Mater Res
44: 289–297, 1999.
19.Fung YC: Stochastic flow in capillary blood vessels. Microvasc Res
5: 34–48, 1973.
20.Yen RT, Fung YC: Effects of velocity distribution on red cell distribution in capillary blood vessel. Am J Physiol
235: H251–H257, 1978.
21.Yang S, Zahn JD: Particle separation in microfluidic channels using flow rate control, in Proceedings of ASME Conference, International Mechanical Engineering Conference Exp Fluids Engineering IMECE
22.Duffy D, McDonald J, Schueller O, Whitesides G: Rapid prototyping of microfluidic systems in poly(dimethylsiloxane). Anal Chem
70: 4974–4984, 1998.
23.Xia Y, Whitesides GM: Soft lithography. Annu Rev Mater Sci
28: 153–184, 1998.
24.Shah RK, London AL (ed): Laminar Flow Forced Convection in Ducts.
New York, Academic Press, 1978.
25.Walburn FJ, Schneck DJ: A constitutive equation for whole human blood. Biorheology
13: 201–210, 1976.