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WO2008036082A1 - Appareil microfluidique de diélectrophorèse à courant continu et ses applications - Google Patents

Appareil microfluidique de diélectrophorèse à courant continu et ses applications Download PDF

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Publication number
WO2008036082A1
WO2008036082A1 PCT/US2006/036446 US2006036446W WO2008036082A1 WO 2008036082 A1 WO2008036082 A1 WO 2008036082A1 US 2006036446 W US2006036446 W US 2006036446W WO 2008036082 A1 WO2008036082 A1 WO 2008036082A1
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Prior art keywords
channel
cells
particles
microchannel structure
sidewall
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English (en)
Inventor
Dongqing Li
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Vanderbilt University
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Vanderbilt University
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Priority to PCT/US2006/036446 priority Critical patent/WO2008036082A1/fr
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    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01N—INVESTIGATING OR ANALYSING MATERIALS BY DETERMINING THEIR CHEMICAL OR PHYSICAL PROPERTIES
    • G01N27/00—Investigating or analysing materials by the use of electric, electrochemical, or magnetic means
    • G01N27/26—Investigating or analysing materials by the use of electric, electrochemical, or magnetic means by investigating electrochemical variables; by using electrolysis or electrophoresis
    • G01N27/416—Systems
    • G01N27/447—Systems using electrophoresis
    • G01N27/44756—Apparatus specially adapted therefor
    • G01N27/44791—Microapparatus

Definitions

  • the present invention generally relates to microfluidics and in particular to DC- Dielectrophoresis microfluidic apparatus, and applications of same including separation of biological cells according to their sizes.
  • Microfluidics deals with the behavior, precise control and manipulation of microliter and nanoliter volume of fluids. It is a multidisciplinary field comprising physics, chemistry, engineering and biotechnology, with practical applications to the design of systems in which such a small volumes of fluids will be used. Microfluidics is used in the development of DNA chips, micro-propulsion, micro-thermal technologies, and lab-on-a chip technology.
  • Lab-on-a-chip devices are miniaturized bio-medical laboratories on a credit card sized glass/plastic plate. These lab chips can duplicate the specialized functions of their room-sized counterparts in clinical diagnoses and tests.
  • the advantages of these lab-on- a-chip devices include significantly reduced sample/reagent consumption, very short analysis time, high throughput and portability.
  • a lab-chip should be able to directly take a drop of whole blood and start the analysis.
  • all lab-on- a-chip devices require purified DNA sample, because these devices do not have the capability to separate white blood cells from the whole blood to extract DNA.
  • the lab-on-a-chip devices still rely on conventional room-sized laboratories for blood sample pretreatment. This is a major limitation to the development and applications of lab-on-a- chip technology.
  • a whole blood sample contains plasma, erythrocytes or red blood cells (RBC), leukocytes or white blood cells (WBC), and thrombocytes or platelets. Only 3% of the blood cells are WBC.
  • the size of RBC is about 6 to 8 ⁇ m, the size of WBC ranges from 10 to 15 ⁇ m.
  • separating the WBC requires using centrifuge, which is not suitable for lab-on-a-chip devices.
  • Magnetic cell separation (MACS) method was used to sort out cancer cells.
  • MCS Magnetic cell separation
  • this method requires using nanometer-sized magnetic beads coated with specific antigens/antibodies that attract certain cells. After the desired cells are attached to the magnetic beads, an external magnetic field is applied to separate these cells from the rest. This method may not work with normal blood cells, and is expansive (the cost of the nano magnetic beads).
  • the magnetic beads attached to cells have to be separated from the cells for the subsequent analyses and DNA amplification (polymerase chain reaction (PCR)) processes.
  • Dielectrophoretic field-flow-fractionation was applied to cancer cell separation.
  • Cell separations were achieved in a thin chamber equipped with a microfabricated, interdigitated electrode array on its bottom wall that was energized with AC electric signals.
  • Cells were levitated by the balance between dielectrophoresis (DEP) and sedimentation forces to different equilibrium heights and were transported at differing velocities and thereby separated when a velocity profile was established in the chamber.
  • This method requires complicated, microfabricated, interdigitated electrode array on the chamber wall, and hence the cost of the device is high and the electronic operation control is sophisticated.
  • pressure-driven flow must be used in this method to generate a parabolic velocity profile in the chamber; the cell separation efficiency is therefore dependent on the flow control as well. Additionally, this requires relatively large, external pump, tubing and valves and thus limits the portability of the device.
  • DC-DEP was employed for particle trapping and concentration in microsystems.
  • An insulator based DEP device was developed with an array of insulating rods in a microchannel, DEP trapping of 200 nm polystyrene particles was realized. Selective trapping of polystyrene particles, live E. coli, and dead E. coli in arrays of insulating posts using DC electric fields was demonstrated. However, no one has shown the separation of particles or cells by size in DC electrokinetic flow by DC-DEP.
  • NCI National Cancer Institute
  • tumor cell enrichment methods can be grouped into the following two broad categories: mechanical (e.g., centrifugation, cytospin, sucrose gradients); and antibody-based selection with mechanical separation (e.g., flow-assisted cell sorting (FACS), magnetic-assisted cell sorting (MACS)). All of these methods have good but not adequate sensitivity or specificity required for detecting precancerous cells in body fluids.
  • mechanical e.g., centrifugation, cytospin, sucrose gradients
  • antibody-based selection with mechanical separation e.g., flow-assisted cell sorting (FACS), magnetic-assisted cell sorting (MACS)
  • the CellSearchTM technique involves mixing a blood sample with iron particles coated with an antibody that attaches to epithelial cells like those found in breast tissue.
  • the cells are further characterized with other antibodies that have been tagged with a fluorescent dye, so that the cancer cells can be easily distinguished and counted. Since epithelial cells are not typically found in blood, their presence suggests they are cancerous cells from the breast tissue.
  • ISET a filtering method that isolate epithelial tumor cells by size. This study demonstrated isolation by size of circulating tumorous cells in patients with carcinoma.
  • ISET uses a module of filtration (Biocom company, Les UHs, France) and a polycarbonate Track-Etch-type membrane (Cyclotron Technology) with calibrated, 8- ⁇ m-diameter, cylindrical pores. Sample is filtered through a 0.6-cm- diameter surface area in the membrane.
  • the present invention relates to a microchannel structure that can be utilized to separate particles or cells in a liquid medium according to their sizes.
  • the microchannel structure includes: a substrate having a first end, and an opposite, second end defining a body portion therebetween, wherein the body portion has a first surface and an opposite, second surface; a first channel formed on the first surface of the substrate with a width, W 1 , defined by a first sidewall and a second, opposite sidewall; a second channel formed on the first surface of the substrate with a width, W 2 , defined by a first sidewall and a second, opposite sidewall, wherein the second channel is in fluid communication with the first channel at a first at least three-way intersection; and an insulating hurdle member having a top portion and protruding from the first sidewall of the first channel, wherein the top portion of the insulating hurdle member and the second sidewall of the first channel defines a width, W la , therebetween, and wherein Wi and W la satisfy the relationship OfW
  • the top portion of the hurdle member has at least one corner with a corresponding angle ⁇ , wherein the angle ⁇ is in the range of 0 to 90°.
  • the hurdle member can have a cross-section in a variety of geometric shapes.
  • the hurdle member in one embodiment, is substantially rectangular cross-sectionally, and the top portion of the hurdle member has two corners.
  • the hurdle member in another embodiment, is substantially triangular cross-sectionally, and the top portion of the hurdle member has one corner.
  • the top portion of the hurdle member in a further embodiment, has a surface characterized by a curvature.
  • the hurdle member in yet another embodiment, is in the form of an insulating liquid droplet, and the surface of the top portion of the hurdle member is at least partially spherical.
  • the microchannel structure further has a third channel formed on the first surface of the substrate with a width, W 3 , defined by a first sidewall and a second, opposite sidewall, wherein the third channel is in fluid communication with the first channel at a second at least three-way intersection, wherein the third channel is formed on the first surface of the substrate such that the hurdle member is positioned between the first at least three-way intersection and the second at least three-way intersection, wherein the width, W 3 , is same or different from at least one of the width, W 1 , and the width, W 2 .
  • the width, W 1 can be same or different from width, W 2 .
  • the substrate in one embodiment, is formed with at least one insulating polymeric material.
  • the insulating polymeric material can be PDMA.
  • the hurdle member can be made from a material different from or substantially same as the at least one insulating polymeric material.
  • the first channel is formed with a first portion with a width, Wi, and a second portion with a width, W ⁇ t> , which is defined by a first sidewall portion and a second sidewall portion, wherein the first sidewall portion is located between the top portion of the hurdle member and the first at least three-way intersection, and the second sidewall portion is located between the top portion of the hurdle member and the first at least three-way intersection, respectively, and wherein the width, W 1 b , is varied at least for a portion along the first channel between the top portion of the hurdle member and the first at least at least three-way intersection, and wherein W ⁇ > W 1 .
  • the width Wi b of the second portion of the first channel proximate to the first at least three- way intersection
  • Each of the first at least three-way intersection and the second at least three-way intersection can be an N-way intersection, where N is an integer no smaller than 3.
  • the first at least three-way intersection is a three-way intersection that is substantially T-shaped.
  • the second at least three-way intersection is a three-way intersection that is substantially T-shaped.
  • the first at least three-way intersection and the second at least three-way intersection can have other shapes, and can be same or different from each other.
  • the niicrochannel structure further has an insulating base member, wherein the insulating base member is bonded with the substrate to form a sealed microchannel structure.
  • the insulating base member is a glass plate.
  • the microchannel structure further has a first well in fluid communication with the first channel at a first end of the first channel, a second well in fluid communication with the second channel at a first end of the second channel, a third well in fluid communication with the second channel at a second end of the second channel, which is apart from the first end of the second channel, and a fourth well in fluid communication with the third channel at a first end of the third channel, respectively.
  • the microchannel structure further has a first electrode configured to be positioned in the first well and to be electrically connectable to a power source, a second electrode configured to be positioned in the second well and to be electrically connectable to a power source, a third electrode configured to be positioned in the third well and to be electrically connectable to a power source, and a fourth electrode configured to be positioned in the fourth well and to be electrically connectable to a power source, respectively.
  • a microfluidic chip can be formed with one or more microchannel structures as set forth above.
  • the one or more microchannel structures can be arranged in an array.
  • a device can be made with one or more such microfluidic chips.
  • the present invention relates to a method of separating particles or cells according to their sizes, wherein the size of each of the particles or cells is characterized by a corresponding diameter.
  • the method includes the steps of: providing a microchannel structure having a substrate having a first end, and an opposite, second end defining a body portion therebetween, wherein the body portion having a first surface and an opposite, second surface, a first channel formed on the first surface of the substrate with a width, W 1 , defined by a first sidewall and a second, opposite sidewall, a second channel formed on the first surface of the substrate with a width, W 2 , defined by a first sidewall and a second, opposite sidewall, wherein the second channel is in fluid communication with the first channel at a first at least three- way intersection, and an insulating hurdle member having a top portion and protruding from the first sidewall of the first channel, wherein the top portion of the insulating hurdle member and the second sidewall of the first channel defines a width, W la , therebetween, and wherein W 1 and W la satisfy the relationship of Wi > Wi a ; introducing a plurality of particles or cells in a liquid medium into the microchannel
  • the step of applying a direct current (DC) electrical field further includes the step of generating a second voltage difference along the second channel such that at the first at least three-way intersection, a first group of the plurality of particles or cells moves to the second channel along a first direction, Y 1 , and a second group of the plurality of particles or cells moves to the second channel along a second direction, Y 2 , that is different from the first direction, respectively, wherein each of the first group of the plurality of particles or cells has a diameter that is larger than a predetermined diameter threshold, and each of the second group of the plurality of particles or cells has a diameter that is not larger than the predetermined diameter threshold.
  • DC direct current
  • the method further has the step of collecting particles or cells after the separation of particles or cells according to their sizes.
  • the plurality of particles or cells can contain white blood cells and red blood cells.
  • the plurality of particles or cells can also contain normal cells and tumor cells.
  • the microchannel structure further has a third channel formed on the first surface of the substrate with a width, W 3 , defined by a first sidewall and a second, opposite sidewall, wherein the third channel is in fluid communication with the first channel at a second at least three-way intersection, wherein the third channel is formed on the first surface of the substrate such that the hurdle member is positioned between the first at least three-way intersection and the second at least three-way intersection.
  • the present invention relates to a method of separating particles or cells according to their sizes, wherein the size of each of the particles or cells is characterized by a corresponding diameter.
  • the method includes the steps of providing a microchannel structure having at least one channel that is defined by a first sidewall and a second, opposite sidewall and has an insulating protrusion formed on one of the first sidewall and the second, opposite sidewall, introducing a plurality of particles or cells in a liquid medium into the at least one channel, and generating a non- uniform electrical field in the at least one channel such that when the plurality of particles or cells passes by the insulating protrusion, the plurality of particles or cells each receives a dielectrophoretic force proportional to its diameters, thereby being separable according to their sizes.
  • the method further has the step of collecting particles or cells after the separation of particles or cells.
  • the present invention relates to an apparatus of separating particles or cells according to their sizes, wherein the size of each of the particles or cells is characterized by a corresponding diameter.
  • the apparatus has a microchannel structure having at least one channel that is defined by a first sidewall and a second, opposite sidewall and has an insulating protrusion formed on one of the first sidewall and the second, opposite sidewall, and means for generating a non-uniform electrical field in the at least one channel such that when the plurality of particles or cells passes by the insulating protrusion, the plurality of particles or cells each receives a dielectrophoretic force proportional to its diameters, thereby being separable according to their sizes.
  • the apparatus has a plurality of microchannel structures in an array.
  • the means for generating a non-uniform electrical field includes a DC power source, a conducting liquid containable in the at least one channel and a plurality of electrodes configured to be electrically connectable to the DC power source and when connected to the DC power source, a non-uniform electrical field is generated at least in the at least one channel.
  • the means for generating a non-uniform electrical field further has a voltage controller electrically coupled to the DC power source and the plurality of electrodes, wherein the voltage controller is capable of controlling the voltage output of each of the plurality of electrodes independently.
  • the apparatus further has means for receiving the plurality of particles or cells, and means for collecting the plurality of particles or cells after the separation of the plurality of particles or cells.
  • the particles or cells are normally provided in a liquid medium of interest, which may comprise a biological fluid of a living subject.
  • the biological fluid includes blood or urine.
  • the blood or urine comprises one or more types of particles or cells.
  • the one or more types of cells are differentiable by their sizes, functions or a combination of them.
  • the one or more types of cells may comprise red blood cells, white blood cells, CD4+ cells, and/or CD3+ cells.
  • the one or more types of cells may be associated with a disease, which may be then detected and/or treated through the cells.
  • Fig. 1 shows contour of a DC electric field E and electric-field lines around an insulating hurdle in a microchannel.
  • the darkness level indicates the magnitude of E.
  • the x-direction is the channel length (the flow) direction, and the y-direction is the channel width direction.
  • the coordinates are normalized by channel width h.
  • (B) shows an enlarged view of a particle moving around the edge region of the hurdle.
  • the darkness level indicates the magnitude of the electric field strength.
  • Fig. 2 schematically illustrates a microchannel structure according to one embodiment of the invention: (a) a perspective view; (b) a partial, perspective view of a microchannel structure formed on a substrate; and (c) a schematic, top view of (b).
  • Fig. 3 shows superimposed sequential images of separation of white blood cells in lysed blood solution by DC-DEP.
  • the channel width is about 300 ⁇ m.
  • the gap width between the channel and the hurdle is 40 ⁇ m.
  • the applied voltages at different branches are indicated, (a) 10 ⁇ m threshold separation; and (b) 5 ⁇ m threshold separation.
  • Fig. 4 is a schematic drawing of a microchannel structure formed on a substrate according to one embodiment of the invention with a rectangular hurdle.
  • Fig. 5 shows: (a) a microchannel structure formed on a substrate according to one embodiment of the invention with a triangle hurdle; and (b) a microchannel structure of (a) but with a diverging channel section.
  • Fig. 6 shows a chip using a microchannel structure according to one embodiment of the invention: (a) a photo image of the chip; and (b) dimension of the chip and corresponding inner structure (the inset).
  • Fig. 8 shows separation of 5.7 and 15.7 ⁇ m particles at different voltage levels. Voltage level increases from 500 to 900 V. Applied voltages at different electrodes are specified in Table 1. (a) Superposed particle trajectories; and (b) comparison of the simulation results and the experimental data.
  • Fig. 11 shows an example of superimposed sequential microscope images of the separation of 6 ⁇ m and 15 ⁇ m polystyrene particles by DC-DEP.
  • the maximum electrical field in the channel is about 80V/cm.
  • Fig. 12 schematically shows a PDMS micro fluidic chip for particle separation according to one embodiment of the invention, where oil from branch 4 forms a droplet in the main channel 1-3.
  • Teflon tubing for oil delivery is press fitted in the PDMS lid of the oil reservoir.
  • Fig. 13 shows superimposed trajectories of the 5.7 ⁇ m and 15.7 ⁇ m particles separated under 400 V applied at reservoir 1 and 160 V at reservoir 2 of Fig 12.
  • the gap between the droplet and the channel wall was 46 ⁇ m wide.
  • Fig. 14 shows the effect of the applied voltage at reservoir 1 of Fig 12 on the separation of 5.7 ⁇ m and 15.7 ⁇ m particles with the 46 mm wide gap. In all three cases the large and small particles were separated by approximately 240 ⁇ m.
  • Fig. 15 shows the effect of the gap width of (a) 95 ⁇ m and (b) 197 ⁇ m on (c) the separation of 5.7 ⁇ m and 15.7 ⁇ m particles under the applied 600 V at reservoir 1 of Fig 12. The downstream separation of particles decreased with the increasing gap width.
  • Fig. 16 shows separation image of 1 ⁇ m and 5.7 ⁇ m particles at the gap width of
  • LOC lab-on-a-chip
  • the LOC is capable of handling substantially small fluid volumes down to less than picoliters to perform desired biological and/or chemical analysis.
  • microchannel refers to a channel structure having a cross-sectional dimension, e.g., a width, a depth or a diameter, in a microscale range from about 0.1 ⁇ m to about 1 mm.
  • the microchannels preferably have a cross-sectional dimension between about 0.1 ⁇ m and 500 ⁇ m, more preferably between about 0.1 ⁇ m and 300 ⁇ m.
  • a device referred to as being microscale includes at least one structural element or feature having such a dimension.
  • microfluidics refers to the science of designing, manufacturing, and formulating devices and processes that deal with volumes of fluid on the order of nanoliters (nl) or picoliters (pi).
  • a microfluidic device has one or more channels with a cross-sectional dimension less than 1 mm.
  • Common fluids used in microfluidic devices include whole blood samples, bacterial cell suspensions, protein or antibody solutions and various buffers.
  • microfluidic devices include, but not limited to, capillary electrophoresis, isoelectric focusing, immunoassays, flow cytometry, sample injection of proteins for analysis via mass spectrometry, PCR (polymerase chain reaction) amplification, DNA (deoxyribonucleic acid) analysis, cell manipulation, cell separation, cell patterning and chemical gradient formation. Many of these applications have utility for clinical diagnostics.
  • electrokinetics refers to the science of electrical charges in moving substances, such as water or blood, which studies particle motion that is the direct result of applied electric fields. Electrokinetics includes electroosmosis, electrophoresis, dielectrophoresis and electrorotation.
  • Electroosmosis also called electroendosmosis, is the motion of polar liquid through a membrane or other porous structure (generally, along charged surfaces of any shape and also through non-macroporous materials which have ionic sites and allow for water uptake, the latter sometimes referred to as "chemical porosity" under the influence of an applied electric field.
  • electrostatic charge will be established at the surface.
  • These surface charges in turn attract the counter ions in the liquid to the region close to the solid-liquid interface to form the electrical double layer. In the electrical double layer region, there are excess counter ions (net charge). If the solid surface is negatively charged, the counter ions are the positive ions.
  • electroosmosis electrophoresis
  • electroosmosis electrophoresis
  • electrophoresis electrophoresis
  • an external electrical field is applied tangentially to the solid surface
  • the excess counter ions will move under the influence of the applied electrical field, pulling the liquid with them and resulting in electroosmotic flow.
  • the liquid movement is carried through to the rest of the liquid in the microchannel by the viscous effect.
  • electroosmotic flow is preferred over pressure driven flow.
  • One of the reasons is the plug-like velocity profile of electroosmotic flow. This means that fluid samples can be transported without dispersion caused by flow shear.
  • pumping a liquid through a small microchannel requires applying a very large pressure difference depending on the flow rate.
  • Electroosmotic flow can generate the required flow rate in very small microchannels without any applied pressure difference cross the channel. Additionally, using electroosmotic flow to transport liquids in complicated microchannel networks does not require any external mechanical pump or moving parts, it can be easily realized by controlling the applied electrical fields via electrodes.
  • Electrophoresis is the motion of a charged particle relative to the surrounding liquid under an applied electrical field.
  • the net velocity of a charged particle is determined by the electroosmotic velocity of the liquid and the electrophoretic velocity of the particle. If the surface charge of the particle is not strong or the ionic concentration of the liquid (e.g., typical buffer solutions) is high, the particle will move with the liquid.
  • Using electrical fields to manipulate and transport particles and biological cells in microchannels is particularly suitable for LOC applications.
  • the applied electrical field has negligible effects on the cells, other than generating the flow and the cell motion. This can be appreciated by comparing the applied electrical field strength with the electrical field strength of the cells' electrical double layer (EDL) field, i.e., the field around each cell generated by the natural surface electrostatic charge.
  • EDL electrical double layer
  • Dielectrophoresis or its acronym "DEP" refers to a phenomenon in which a force is exerted on a dielectric particle when it is subjected to a non-uniform electric field. This force does not require the particle to be charged.
  • this invention in one aspect, relates to a method and device for separating particles or cells of different sizes in a liquid medium according to their sizes.
  • the microchannel structure 200 has a substrate 210 that has a first end 212, and an opposite, second end 214 defining a body portion 220 therebetween, wherein the body portion 220 has a first surface 222 and an opposite, second surface 224.
  • a first channel 230 is formed on the first surface 222 of the substrate 210 with a width, W 1 , defined by a first sidewall 232 and a second, opposite sidewall 234.
  • a second channel 240 is formed on the first surface 222 of the substrate 210 with a width, W 2 , defined by a first sidewall 242 and a second, opposite sidewall 244, wherein the second channel 240 is in fluid communication with the first channel 230 at a first three-way intersection 241.
  • an insulating hurdle member 250 which has a top portion 251, protrudes from the first sidewall 232 of the first channel 230, wherein the top portion 251 of the insulating hurdle member 250 and the second sidewall 234 of the first channel 230 defines a width, W la , therebetween, and wherein Wj and W la satisfy the relationship of W 1 > W 1 a .
  • the hurdle member can be formed with different geometric shapes as well as different materials.
  • the hurdle member 250 is substantially rectangular cross-sectionally, and the top portion 251 of the hurdle member 250 has two corners 252, 254, each with a corresponding angle ⁇ , wherein the angle ⁇ is in the range of 0 to 90°.
  • the hurdle member 550 is substantially triangular cross- sectionally, and the top portion 551 of the hurdle member 550 has one corner 552 with a corresponding angle ⁇ , wherein the angle ⁇ is in the range of 0 to 180°.
  • the top portion 1251 of the hurdle member 1250 has a surface characterized by a curvature 1252.
  • the hurdle member 1250 can be an insulating liquid droplet, and the surface of the top portion of the hurdle member 1250 is at least partially spherical.
  • the microchannel structure 200 further has a third channel 260 formed on the first surface 222 of the substrate 210 with a width, W 3 , defined by a first sidewall 262 and a second, opposite sidewall 264, wherein the third channel 260 is in fluid communication with the first channel 230 at a second three-way intersection 261, wherein the third channel 260 is formed on the first surface 222 of the substrate 210 such that the hurdle member 250 is positioned between the first three-way intersection 241 and the second three-way intersection 261.
  • the width, W 3 can be same or different from at least one of the width, W 1 , and the width, W 2 , wherein the width, Wi, can also be same or different from width, W 2 .
  • the substrate 210 is formed with at least one insulating material such as a polymeric material, glass or other types of materials.
  • the insulating polymeric material comprises PDMA.
  • the hurdle member 250 is made from a material substantially same as the at least one insulating material of the substrate 210.
  • the microchannel structure 500 has a first channel 430, a second channel 440, a third channel 460, a first three-way intersection 441, a second three-way intersection 461 and a hurdle member 450, which is made from a material different from the insulating material that forms the substrate.
  • a first channel 530 is formed with a first portion 530a with a width, W 1 , and a second portion 530b with a width, Wi b , which is defined by a first sidewall portion 532b and a second sidewall portion 534b, wherein the first sidewall portion 532b is located between the top portion 551 of the hurdle member 550 and the first three-way intersection 541, and the second sidewall portion 534b is located between the top portion 551 of the hurdle member 550 and the first three-way intersection 541, respectively, and wherein the width, Wit,, is varied at least for a portion along the first channel 530 between the top portion 551 of the hurdle member 550 and the first three-way intersection 541.
  • the microchannel structure 200 further has an insulating base member 290, where the insulating base member 290 is bonded with the substrate 210 to form a sealed microchannel structure 200.
  • the insulating base member 290 can be a glass plate.
  • the first three-way intersection 241 is substantially T-shaped
  • the second three-way intersection 261 is substantially T-shaped. They can, of course, have other shapes.
  • microchannel structure 600 has a first well 633 in fluid communication with a first channel 630 at a first end 630a of the first channel 630, a second well 643 a in fluid communication with a second channel 640 at a first end 640a of the second channel 640, a third well 643b in fluid communication with the second channel 640 at a second end 640b of the second channel 640, which is apart from the first end 640a of the second channel 640, and a fourth well 663 in fluid communication with the third channel 660 at a first end 660a of the third channel 660.
  • These wells, or ports or reservoirs can be utilized for inputting and outputting functions.
  • wells 633 and 663 can be used for inputting a solution, such as a buffer solution, and a particle mixture, respectively
  • wells 643a and 643b can be used for collecting the separated small and large particles/cells, respectively.
  • microchannel structure 600 further has a first electrode 635 configured to be positioned in the first well 633 and to be electrically connectable to a power source (not shown), a second electrode 645a configured to be positioned in the second well 643a and to be electrically connectable to a power source (not shown), a third electrode 645b configured to be positioned in the third well 643b and to be electrically connectable to a power source (not shown), and a fourth electrode 665 configured to be positioned in the fourth well 663 and to be electrically connectable to a power source (not shown), respectively.
  • These electrodes are adapted for establishing electrical potentials or voltages in the microchannel structure.
  • microchannel structures can be used to form a microfluidic chip, which in turn can be used to form a device that, for example, can separate particles or cells according to their sizes.
  • a plurality of microchannel structures can be arranged in an array (not shown) in one embodiment of the present invention.
  • microchannel structure 200 can be used as follows. A plurality of particles or cells 205 in a liquid medium is introduced into the microchannel structure 200, for example, through well 263 and/or well 233.
  • a direct current (DC) electrical field is applied within the microchannel structure 200 to generate a non-uniform electrical field at least around the insulating hurdle member 250 and a first voltage difference along the first channel 230 such that the plurality of particles or cells 205 is driven by the direct current (DC) electrical field along the first channel 230 and separated according to their diameters by a dielectrophoretic force corresponding to the non-uniform electrical field when the plurality of particles or cells 205 passes by the insulating hurdle member 250.
  • DC direct current
  • a second voltage difference is also generated along the second channel 240 such that at the first three-way intersection 241, a first group 205b of the plurality of particles or cells 205 moves to the second channel 240 along a first direction, Y 1 , and a second group 205a of the plurality of particles or cells 205 moves to the second channel 240 along a second direction, Y 2 , that is different from the first direction, respectively.
  • Y 2 is opposite to the first direction, Y 1 .
  • Each of the first group 205b of the plurality of particles or cells 205 has a diameter that is larger than a predetermined diameter threshold, and each of the second group 205a of the plurality of particles or cells 205 has a diameter that is not larger than the predetermined diameter threshold.
  • the plurality of particles or cells 205 is separated according to their sizes. Note that, approximately speaking, the size of a cell is proportional to (d) 3 , where d is the diameter of the cell. The separated particles or cells can be further collected for processing.
  • the present invention can be practiced to separate a plurality of particles or cells having a mix of white blood cells and red blood cells.
  • the present invention can also be practiced to separate a plurality of particles or cells having a mix of normal cells and tumor cells.
  • the present invention can be utilized for other applications as well.
  • the present invention relates to a method of separating particles or cells according to their sizes, wherein the size of each of the particles or cells is characterized by a corresponding diameter.
  • the method includes the steps of providing a microchannel structure having at least one channel that is defined by a first sidewall and a second, opposite sidewall and has an insulating protrusion formed on one of the first sidewall and the second, opposite sidewall, introducing a plurality of particles or cells with a conducting liquid into the at least one channel, and generating a non-uniform electrical field in the at least one channel such that when the plurality of particles or cells passes by the insulating protrusion, the plurality of particles or cells each receives a dielectrophoretic force proportional to its diameters, thereby being separable according to their sizes.
  • the method further has the step of collecting particles or cells after the separation of particles or cells.
  • the present invention relates to an apparatus of separating particles or cells according to their sizes, wherein the size of each of the particles or cells is characterized by a corresponding diameter.
  • the apparatus has a microchannel structure having at least one channel that is defined by a first sidewall and a second, opposite sidewall and has an insulating protrusion formed on one of the first sidewall and the second, opposite sidewall, and means for generating a non-uniform electrical field in the at least one channel such that when the plurality of particles or cells passes by the insulating protrusion, the plurality of particles or cells each receives a dielectrophoretic force proportional to its diameters, thereby being separable according to their sizes.
  • the apparatus may have a plurality of microchannel structures in an array.
  • the means for generating a non-uniform electrical field includes a DC power source, a conducting liquid containable in the at least one channel and a plurality of electrodes configured to be electrically connectable to the DC power source and when connected to the DC power source, a non-uniform electrical field is generated at least in the at least one channel.
  • the means for generating a non-uniform electrical field further has a voltage controller electrically coupled to the DC power source and the plurality of electrodes, wherein the voltage controller is capable of controlling the voltage output of each of the plurality of electrodes individually.
  • the apparatus further has means for receiving the plurality of particles or cells, and means for collecting the plurality of particles or cells after the separation of the plurality of particles or cells.
  • the present invention utilizes the principle of dielectrophoretic force.
  • a suspension of dielectric particles In a dielectric fluid.
  • the particle and the surrounding medium are electrically polarized and the surface charge accumulates at the interfaces due to the difference in electric properties.
  • the distribution of the surface charge of the particle gives rise to an induced dipole moment.
  • the dipole tends to align in parallel with the local electric field.
  • the forces acting on the opposite charges of a dipole become asymmetric.
  • DEP dielectrophoretic
  • the induced motion of the particle due to the DEP force is known as dielectrophoresis.
  • DEP Using DEP, manipulation of particles can be realized by controlling the electric field without any mechanical moving part. Furthermore, different from the conventional electrophoresis that works only on the charged particles, dielectrophoresis force also acts on the electrically neutral particles, which greatly increases its biological applicability. The magnitude of the DEP force is dependent on the size and dielectric property of the particle.
  • the chips with microchannels and wells are typically made of Polydimethylsiloxane (PDMS) and glass by soft lithography method at very low cost.
  • DEP does not necessarily require an AC field; it requires only a non-uniform electrical field.
  • a non-uniform DC field in a microchannel can generate DEP.
  • the sizes of WBC and RBC are different, and the DEP force is proportional to the volume of the particle/cell. Therefore, it is possible to separate WBC from RBC by DC-DEP, or to separate the larger tumor cells by DC-DEP in a simple DC electrokinetic microfiuidic chip, which is provided by the present invention.
  • l(a) and (b) illustrate a particle 105 moving under an applied DC field in a microchannel with an insulating hurdle 150.
  • the hurdle 150 is attached on one side of the microchannel to form an abruptly narrow section 103. Since only the liquid (an aqueous solution) conducts electrical field, the narrow section 103 of the microchannel generates a spatially nonuniform DC electrical field 101 in the space near the hurdle 150.
  • Fig. Ib shows an enlarged view of the local electrical field 101 near the up-stream corner 152 of the hurdle 150. Under the combined effect of the electroosmotic flow (EOF) and the electrophoresis (EP), a particle 105 moves towards the entrance region of the narrow section 103 of the channel. As shown in Fig.
  • EEF electroosmotic flow
  • EP electrophoresis
  • the electric field 101 is stronger close to the corner 152 of the hurdle 150. Since the negative DEP force directs to the region of lower electric-field strength, the particle experiences a repulsive force from the corner 152 of the hurdle 150.
  • the magnitude of the repulsive DEP force is proportional to the volume of the particle 105 and the local value of (E • V) E, as indicated by Eq. (1).
  • E • V the repulsive DEP force on a 15 ⁇ m particle is 27 times of that on a 5 ⁇ m particle under the same conditions. Therefore a larger particle is subject to a stronger DEP force and tends to be pushed further away from the corner 152 compared with a smaller particle.
  • the similar DEP repulsion occurs when the particle 105 passes by the other corner 154 of the hurdle 150. As a result, the trajectory shift (in y-direction) will be different for particles of different sizes and hence particle separation by size can be expected.
  • the DC-DEP method according to the invention may be used for separation of biological cells of different sizes, such as for separation of white blood cells from red blood cells.
  • a DC-DEP blood cell separation chip 200 has a PDMS plate 220 and a glass slide, i.e., a PDMS plate 210 with the microchannel-well structure is bonded with a glass slide 290 to form a sealed microchannel structure.
  • a PDMS plate 210 with the microchannel-well structure is bonded with a glass slide 290 to form a sealed microchannel structure.
  • DC electrical field is applied via four electrodes (not shown) inserted in these wells.
  • a hurdle 250 is formed on one side of microchannel 230 to form an abruptly narrow section. Since only the liquid (an aqueous solution) conducts electrical field, the narrow section of the microchannel 230 generates a spatially non-uniform DC electrical field in the liquid near the hurdle 250.
  • the mixed larger and smaller cells 205 in a liquid medium is introduced into channel 230 (inputting channel) from channel 260 (i.e., inputting branch).
  • the negative DC-DEP force at the corners 252, 254 of the hurdle 250 pushes larger cells 205b more than smaller cells 205a, and hence generates different trajectories for larger and smaller cells once they pass the hurdle 250.
  • a T-shaped channel structure 241 may be used so that the separated larger cells 205b and the separated smaller cells 205a are drawn into separate cell collection wells by electrokinetic flows. Since the particles' trajectories after the hurdle 250 and hence the final separation efficiency are also coupled with the electrokinetic flows of the liquids in different channel branches, voltage output may be adjusted to obtain the optimal voltages applied at the four electrodes.
  • this design may be integrated to existing lab-on-a-chip devices that require separated white blood cells (WBC), and extend this method into a high throughput technology to handle a large volume of blood sample. For instance, using micro-fabrication method one can build hundreds of parallel microchannels on a single chip, so that separation of blood cells from a larger volume of blood sample can be achieved.
  • WBC white blood cells
  • the optimal design parameters e.g., the hurdle size and position
  • the optimal controlling parameters e.g., applied voltages
  • the cell transport in a microchannel involves both the electroosmotic flow (liquid motion) and the electrophoresis (particle motion relative to the liquid).
  • the non-uniform electrical field and hence the DC-DEP force has to be considered.
  • a theoretical model for the cell transport processes in the microchannel under an applied DC electrical field includes: (1) the Laplace equation for the applied electrical field; (2) 3D equations of motion to describe the flow field of the liquid; and (3) Newton's 2 nd law equation, including the electrophoresis force, the dielectrophoresis force, the flow friction force, to determine the motion (velocity and trajectory) of the particle.
  • the system used here involves buffer solutions with high ionic strength, the thickness of the electric double layer fields around the cells and channel walls are less than 10 nm. Therefore, the thin double layer treatment may be used, i.e., the double layer causing electroosmotic flow may be considered as a slip velocity boundary condition for the equation of motion.
  • the complexity and the coupling effects involved in these equations and the boundary conditions require developing an efficient numerical method.
  • an ALE arbitrary Lagrangian-Eulerian
  • a generalized Galerkin finite element formulation may be used to solve over a 3D unstructured tetrahedral mesh.
  • the mesh becomes distorted and new meshes are required at specific time steps in order to capture this motion.
  • These re-meshing steps are very time-consuming, rendering this method unsuitable for tracking the particle over large distances.
  • Employing the Chimera or overlapping grid scheme may solve the electrophoretic motion of a particle in a microchannel. The advantages of this method are that it eliminates the need for computationally expensive re-meshing steps, and it simplifies the procedures for solving the discretized equations by using structured rather than unstructured grids.
  • the PDMS (polydimethylsiloxane) microchannel can be fabricated following the soft lithography protocol.
  • Fluorescent (carboxylate-modif ⁇ ed) polystyrene particles of different sizes 6 ⁇ m, 10 ⁇ m and 15 ⁇ m in diameter (Bangs Laboratory Inc.) may be used as sample particles for fundamental studies. These particle sizes are similar to the size of typical biological cells such as the red blood cells and the white blood cells.
  • the channel and all the wells may be primed with the 1 mM sodium carbonate buffer solutions. Then the cells or particle mixture may be introduced into the well with a syringe.
  • a high- voltage DC power supply (Labsmith HVS448) may be used to drive the liquid flow though the microchannel structure by platinum electrodes submerged in each well. This power supply can provide and control the voltage outputs of each of the four electrodes independently. Following the results of the numerical simulations as guidance, the voltage applied to electrodes may be carefully adjusted to realize that the liquids and the cells/particles in the inputting branches always move towards the hurdle and eventually flow into the two separation branches.
  • the effect of the pressure-driven flow can be minimized by using sufficiently large well size and by carefully balancing the liquid level in four wells before each experimental run.
  • the cell/particle motion may be monitored by an inverted optical microscope (TE2000-U, Nikon Inc.) and recorded by a progressive CCD camera (Qlmaging, British Columbia, Canada).
  • the parameters that have effects on the cell/particle separation include: (1) design parameters, such as the channel's dimensions, the hurdle's size and position; and (2) operation parameters, i.e., the applied voltages at different electrodes.
  • design parameters such as the channel's dimensions, the hurdle's size and position
  • operation parameters i.e., the applied voltages at different electrodes.
  • the experimentally measured cells' trajectories with different design parameters and operation parameters may be compared with the predictions of numerical simulations. Several sets of optimal design parameters and operation parameters may be determined.
  • the microfluidic chip of the invention may include a plate with hundreds of parallel microchannels, making it feasible to perform high-throughput of on-chip blood cell separation. Conventionally, most lab-on-chip applications are interested only in the treatment of very small amount of samples.
  • the typical speed of the cell motion in the microchannels is about 1000 ⁇ m/s.
  • the invention may make it feasible to treat 200 ⁇ 300 cells/second or 12000-18000 cells/min in one microchannel. For instance, one can use the design provided in the present invention to build a chip with 10 parallel channels for performing high-throughout of on-chip blood cell separation.
  • Example 1
  • the present invention has been used to separate different white blood cells.
  • a 50 ⁇ l volume of blood was mixed with 50 ⁇ l of a Red Blood Cell Lysis Buffer (Caltag, Burlingame, CA) to lyse the red blood cells, and then diluted with 500 ⁇ l of de-ionized water (this protocol fixes WBC in the sample, and lyses RBC).
  • 10 ⁇ l of this sample solution was loaded to the sample well on the chip by a micro-pipette.
  • 10 ⁇ l of this sample solution contains approximately 8,000 cells (granulocytes, monocytes, and lymphocytes) and approximately 100,000 small components (platelets, RBC debris, etc).
  • the inventor By adjusting the applied voltages at different electrodes inserted in the wells at the ends of the microchannels, the inventor separated the white blood cells at a specified cell size, which is corresponding to a predetermined diameter for the cell such as 10 ⁇ m as shown in Fig. 3 as an example.
  • the applied electrical field has negligible effects on the cells, other than generating the cell motion. This can be appreciated by comparing the applied electrical field strength with the electrical field strength of the cells' electrical double layer (EDL) field (i.e., the field around each cell generated by the natural surface electrostatic charge).
  • EDL electrical double layer
  • a blood sample solution 205 containing a few larger tumor cells 205b and many smaller normal blood cells 205a may be introduced into a sample well.
  • the mixed larger and smaller cells will be introduced with the flow into the main channel 230 from a channel 260 (inputting branch) through a three-way intersection 261.
  • the negative DC-DEP force at the corners 252, 254 of the hurdle 250 will push the larger tumor cell 205b more than the smaller normal cells 205a, and hence generate different trajectories for the larger tumor cell and the smaller normal cells once they pass the hurdle 250.
  • a T- shaped three-way intersection 241 will allow the separated larger tumor cells 240 and the separated smaller normal cells 240 to be draw into separate cell collection wells by electrokinetic flows.
  • the method(s) of the invention can be used to conduct extensive numerical simulations of the on-chip processes of separating the larger tumor cells from the rest, under various conditions.
  • the results and findings of these numerical experimental studies can allow us to determine the optimal design parameters (e.g., the hurdle size and position) and the optimal controlling parameters (e.g., applied voltages) for such a tumor cell separation chip.
  • a circulating tumor cell separation chip is similar to the one shown in Fig. 2. There are four reservoirs (wells) connected at the ends of the four microchannel branches.
  • the PDMS (poly-dimethylsiloxane) microchannel may be fabricated following the soft lithography protocol.
  • the channel and all the wells may be primed with the 1 mM sodium carbonate buffer solutions.
  • a mixture of normal blood cells and tumor cells after being processed appropriately, may be introduced into the sample well with a digital micro-pipette and a Nikon cell injector.
  • a high- voltage DC power supply (Labsmith HVS448) may be used to drive the liquid flow though the microchannel network by platinum electrodes submerged in each well. This power supply can provide and control the voltage outputs of the four electrodes independently. Following the results of the numerical simulations as guidance, the voltage applied to the electrodes may be carefully adjusted to realize that the liquids and the cells in the inputting branches 260 will move towards the hurdle 250 and eventually flow into the two separation branches 243a, 243b.
  • the effect of the pressure-driven flow can be minimized by using sufficiently large well size and by carefully balancing the liquid level in four wells before each experimental run.
  • the cell motion may be monitored by an inverted microscope (TE2000-U, Nikon Inc.) and recorded by a progressive CCD camera (Qlmaging, British Columbia, Canada) and a digital imaging system.
  • the parameters that may have effects on the tumor cell separation include: (1) design parameters, such as the channel's dimensions, the hurdle's size and position; and (2) operation parameters, i.e., the applied voltages at different electrodes.
  • design parameters such as the channel's dimensions, the hurdle's size and position
  • operation parameters i.e., the applied voltages at different electrodes.
  • the experimentally measured cells' trajectories with different design parameters and operation parameters may be compared with the predictions of numerical simulations. Several sets of optimal design parameters and operation parameters may be determined.
  • an electrokinetic based microfiuidic chip to separate circulating tumor cells from blood with high sensitivity, one can analyze the cells in each collection well using the microscope to examine the number of the isolated tumor cells.
  • the preferred separation efficiency is to reach about 100% isolation/separation of the circulating tumor cells.
  • the microfiuidic chip of the present invention is the first device developed that have applicability in the DC-DEP separation of circulating tumor cells, and may prove its ability of separating one tumor cell out of 10,000 or more normal cells.
  • the method disclosed in the invention is not limited to separate tumor cells in blood; it can be applied to separate tumor cells in other bio-fluids as well, such as sputum and urine. It may be used to separate various types of tumor cells.
  • the device and the method of the invention may be used for high throughput, for example, by using multiple parallel microchannels in an array on a single chip. Further, the invention utilized be made to be a fully-automatic, practical and effective tool for biomedical research.
  • the invention provides a new electrokinetic based microfiuidic method to separate the larger tumor cells in blood samples.
  • the separation is performed on a chip with a size of a microscope glass slide, is a single step process and operated in a continue flow mode.
  • the device according to the invention has no mechanical moving parts and no filters. No any labeling is required.
  • This method can also be applied to separate circulating tumor cells in other forms of body fluids.
  • the DC-DEP method of the invention may further include a step of providing a diverging microchannel section immediately after the hurdle, the diverging microchannel section being located on the opposite side of the hurdle and connecting the narrow section to the outputting channel.
  • the DC-DEP particle/cell separation chip 400 uses a rectangular hurdle 450.
  • the design in Fig. 5 uses a triangle shaped hurdle 550, instead of a rectangular hurdle.
  • the DC-DEP force at the first corner (upstream) 452 of the rectangular hurdle 450 will push larger cells more and push smaller cells less, and hence create the initial separation.
  • the larger cells will move away from the second corner (downstream) 454 of the hurdle 450, in comparison with the smaller cells.
  • the larger cells will receive much less interaction from the second corner 454.
  • Fig. 5B uses a diverging microchannel section immediately after the hurdle 550, which will create diverging stream lines, and help the separated lager cells move further away from the smaller cells.
  • Fig. 6 there are four branches connected to four different reservoirs.
  • Wells or reservoirs B (663) and C (633) are for inputting the particle mixture and the buffer solution, respectively.
  • Wells or reservoirs A (643a) and D (643b) are for collecting the separated small and large particles, respectively.
  • Branches A (640), C (630), and D (640) are 300 ⁇ m in width.
  • Branch B (660) is 90 ⁇ m in width. All of the branch channels are 45 ⁇ m in depth (in z-direction).
  • the kernel structure is a rectangular block 650 (240 ⁇ m x 130 ⁇ m) located between the first three-way intersection 641 and the second three- way- intersection 661.
  • the polydimethylsiloxane (PDMS) microchannel was fabricated following the soft lithography protocol. Fluorescent (carboxylate-modified) polystyrene particles of three different sizes,
  • 5.7, 10.35, and 15.7 ⁇ m in diameter were used as sample particles. These particle sizes are similar to the size of typical biological cells such as the red blood cells and the white blood cells.
  • the particles were supplied in the form of 1% suspension in pure water. These particle solutions were further diluted with the ImM sodium carbonate buffer (Na 2 CO 3 /NaHCO 3 ) solutions. The number density of particle was normally about 10 5 AnL. Since the mass density of the particles was slightly greater than that of water (nominal density is 1.05 g/mL), the particle solutions were gently vibrated prior to use to prevent sedimentation. Before the experiment, the channel and all the wells or reservoirs were primed with the ImM sodium carbonate buffer solutions.
  • the particle mixture was introduced into reservoir B (663) with a 1-mL plastic syringe.
  • a high-voltage DC power supply (Glassman High Voltage, High Bridge, NJ) was used to drive the fluid flow through the microchannel structure by platinum electrodes submerged in each reservoir.
  • a custom-made voltage controller was used to adjust independently the voltage output of each of the four electrodes.
  • electrode D (645b) was always grounded.
  • the voltage outputs to electrodes A (645a), B (665), and C (635) were carefully adjusted to realize that the fluids in the inputting branches B (660)and C (630) always moved to the block 650 and flowed into the separation branches A (640a) and D (640b).
  • the pressure-driven flow was minimized by carefully balancing the liquid level in four reservoirs before each experimental run.
  • the particle motion was monitored by an inverted optical microscope (Nikon Canada) and recorded by a progressive CCD camera (Q Imaging, Burnaby, British Columbia, Canada).
  • the camera was operated in video mode at a frame rate of 11.4 frames per second.
  • the acquired images (viewed from the top) had a resolution of 640 x 484 pixels.
  • the reading error to determine the particle positions is about ⁇ 2 pixels which corresponds to actual dimension of ⁇ 5.4 mm.
  • the zeta potential of the PDMS channel wall was set to 22OmV.
  • the electrophoretic mobilities of the 5.7 and 15.7 ⁇ m particles were fixed as 3.3 x 10 "s and 3.7 x 10 "9 In 2 S -1 V “1 , respectively, which were based on an independent measurement in a straight channel using the same buffer solution. Because the ionic concentration of the working solution is very low, the liquid properties are not different from that of deionized (DI) water, that is, dynamic viscosity 1.0 x 10 '3 kg “ m ' 's "1 , density 998 kg/m 3 , and electrical permittivity 6.96 x 10 "10 C-V " " 1 V " '
  • a typical case The magnitude of the particle trajectory deviation is proportional to the DEP force acting on the particle, and hence the particle volume. Therefore, the trajectories of the particles of different sizes can be diverted into different streams after they pass the block 650.
  • a typical case of separation of 5.7 ⁇ m particles and 15.7 ⁇ m particles is shown in Fig. 7(a), which was obtained by superposing a series of consecutive images of the moving particles. Initially the particle mixture came out as a single stream from the inputting branch B 660. Then the main stream of the buffer solution from branch C 630 squeezed the mixture and forces the particles to move closely to the block comer. After the particles passed through the gap between the block 650 and the channel wall, their trajectories were changed.
  • the trajectory deviation for a larger particle was greater than that for a smaller particle because of the different magnitude of the DEP force they experience at the block corners.
  • the single mixture stream was separated into two.
  • the larger particle moved into the separation branch D 696, while the smaller particle moves into the other separation branch A 698.
  • the particle mixture is continuously separated into two different reservoirs.
  • the particles in reservoirs D 643b and A 643a were pure 15.7 and 5.7 mm particles, respectively.
  • Fig. 7b demonstrates the comparison between the simulation results and the experimental results for the separation of 5.7 and 15.7 ⁇ m particles.
  • the dotted symbols are digitized positions of the particles based on Fig. 7a.
  • Fig. 8 shows the separation of 5.7 and 15.7 ⁇ m particles at different voltage levels from about 500 to 900 V.
  • Fig. 8A shows the superposed particle trajectories and
  • Fig. 8B shows the comparison of the experimental and simulation results. It was demonstrated that the magnitude of trajectory deviation for both larger and smaller particles increased with increasing the voltage level. This was because the DEP force is proportional to the gradient of the electric-field intensity, V
  • the major function of the electrode B in the inputting reservoir was for driving the particle mixture into the block region, so that VQ should not be very small. Otherwise the EOF will be directed to flow back into branch B 660 and the particle mixture cannot be successfully introduced into the channel network. In this experiment, we found it wise to keep VQ around 50% of VQ, as shown in Table 1.
  • Fig. 9B where smaller cells 905a move to branch 943a after passing the block or hurdle member 950 from main channel 930 and larger cells 905b move to branch 943b.
  • the values of the above two threshold voltages were based on the numerical simulation for 5.7 and 15.7 ⁇ m particles at the voltage level of 500 V. The real values may be slightly different in the experiments. It can be reasonably inferred that the threshold voltages are dependent on the channel configuration and the sizes of the particles. They can be determined by experimental calibration.
  • Fig. 10a shows the streamlines of the EOF near the block 650 which was obtained by using 1 ⁇ m fluorescent polystyrene particles suspended in ImM sodium carbonate buffer solution. This image was taken by a CCD camera at an exposure time of 4 s. It could be seen that fluid from the particle input branch B 660 was squeezed to move closely near the wall of the block 650 by the main flow from the buffer input branch C 630. This is the purpose of branch C 630. This design can ensure that the particles move sufficiently near the block corners where they will experience strong DEP force. Thus, the particles will have the greatest trajectory changes after they pass the block region 650.
  • the electrostatic field and the EOF field in the microchannel were simulated using FEMLAB® (Comsol).
  • the simulated streamlines are shown in Fig. 10b.
  • the simulation results fit reasonably well with the experimental results and were used for the subsequent simulation of the particle trajectory.
  • Fig. 11 shows trajectories of 6 ⁇ m and 15 ⁇ m polystyrene particles by superimposing a series of sequential microscopy images obtained from an experiment.
  • the microchannel (made of PDMS and glass by soft lithography) in this case was 300 ⁇ m in width and 40 ⁇ m in depth (perpendicular to the paper).
  • the narrow section of the microchannel was 60 ⁇ m in width.
  • a microfiuidic chip illustrated in Fig. 12 which is composed of a PDMS upper substrate 1210 and a glass lower substrate 1290.
  • This microchamiel structure 1200 has four 40 ⁇ m deep branch channels (correspondingly, four reservoirs filled with particle solution, distilled water or oil): 515 ⁇ m wide and 12.5 mm long channel 1 reservoir supplies the system with the distilled water; 131 ⁇ m wide and 25 mm long channel 2 reservoir transports particles suspended in distilled water to the droplet region (see the inset of the zoom-in view in Fig.
  • the microchip was fabricated in PDMS using a standard single-layer soft lithography technique. A detailed procedure is described in one of the reference, except for the oil connection shown in Fig. 12, which required few additional fabrication steps.
  • a square piece of a thin PDMS layer with a punched through opening for the Teflon tube was used as a lid for the oil reservoir. Both the PDMS chip and the lid were treated in oxygen plasma (plasma cleaner PDC-32G, Harrick Scientific, Ossining, NY) for 30 s and then bonded to each other.
  • oxygen plasma plasma
  • a high- voltage DC power supply Glassman High Voltage Inc., High Bridge, NJ
  • Particle motion was visualized using a Leica DMLM fluorescence microscope with IQx and 16 ⁇ objectives, the appropriate filter set and a 100 W broadband mercury lamp, Images were recorded with a 12 bit Retiga-1300 cooled digital CCD camera (Pulnix America Inc., Sunnyvale, CA) and OpenLab 3.1.5 image acquisition software. All images were recorded with 80 ms exposure time. The acquired images have the resolution of 640 x 512 pixels, with each pixel representing 2 ⁇ m and 3.3 ⁇ m square in object plane for 16x and 1Ox objective, respectively.
  • Fig. 14 compares the y-coordinates of 5.7 ⁇ m and 15.7 ⁇ m particles when their trajectories became steady downstream i.e. parallel to the channel sidewalls. The y-coordinates of both particles before the droplet are also included for reference. It is apparent that all particles approached the droplet at its base, at approximately zero y-coordinate (3 ⁇ m for the 5.7 ⁇ m particles and 13 ⁇ m for the 15.7 ⁇ m particles).
  • the size of the insulating obstacle can be varied, so the electric field gradient was easily controlled by changing the droplet size.
  • the droplet size was increased, the electric field was more compressed between the droplet and the channel wall, inducing a higher DEP force.
  • the applied voltages at reservoirs 1 and 2 were fixed at 600 V and 235 V, respectively.
  • Fig. 15C shows that as the gap size decreased four-fold from 197 ⁇ m to 46 ⁇ m, the distance between the trajectories of the small and big particles increased 5.5-fold from 43 ⁇ m to 236 ⁇ m. With this large distance, as large as half of the channel width, the big and small particles could be easily separated downstream into two branches/reservoirs .
  • the optimal particle separation would occur when the gap is as small as possible, i.e. just large enough to allow continuous particle flow.
  • the highest possible field gradient and thus the highest DEP force would be achieved for the given voltages. Therefore, the gap was reduced to about 16 ⁇ m while the voltage at reservoir 1 was kept at 600 V.
  • no particle separation or trajectory shift was observed due to a flow circulation region formed just before the gap. 15.7 ⁇ m particles seemed to be repulsed back from the gap region along with 5.7 ⁇ m particles and formed a so-called pearl chain, which is a known DEP phenomenon originated from particle-particle interaction. The chain circulated and trapped more incoming particles.
  • Fig. 16 shows that the critical parameter in the DC-DEP particle separation is the gap size, i.e. the field gradient.
  • the electrical field strength can have a very low value if the field gradient is sufficient to discriminate particles of different sizes.
  • Choosing lower voltages for the DC-DEP separation is also preferable when dealing with biological samples since the biological cells are prone to lysis at a strong electrical field.
  • High voltage also can cause significant Joule heating and gas bubble generation at metal electrodes due to the electrolysis.
  • the only drawback of the low voltage separation is that the velocity of the particles and therefore separation process becomes slower.
  • the present invention discloses a new particle or cell separation technology with high separation efficiency by a DC-DEP microfluidic device and method.
  • T cell levels results of 6 years of follow-up. The Los Angeles Center, Multicenter AIDS Cohort Study. J Acquir Immune Defic Syndr. 1993;6:904-12.
  • CD4 lymphocyte percentage predicts disease progression in HIV-infected patients initiating highly active antiretro viral therapy with CD4 lymphocyte counts >350 Iymphocytes/mm3. J Infect Dis. 2005;192:950-957.
  • Kang KH Li D. Dielectric Force and Relative Motion between Two Spherical Particles in Electrophoresis. Langmuir. 2006;22:1602-8.

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Abstract

La présente invention a trait à un appareil et à des procédés de séparation des particules ou cellules en fonction de leurs tailles, la taille de chacune des particules ou cellules étant caractérisée par un diamètre correspondant. Selon un mode de réalisation, le procédé inclut les étapes consistant à fournir une structure microcanal dotée d'au moins un canal qui est défini par une première paroi latérale et une seconde paroi latérale opposée et qui présente une protubérance isolante formée sur la première paroi latérale ou la seconde paroi latérale opposée ; introduire une pluralité de particules ou cellules dans un milieu liquide dans le ou les canaux ; et générer un champ électrique non uniforme dans le ou les canaux de telle sorte que, lorsque la pluralité de particules ou cellules passent par la protubérance isolante, la pluralité de particules ou cellules reçoivent chacune une force diélectrophorétique proportionnelle à leurs diamètres, ce qui permet ainsi de les séparer en fonction de leurs tailles. Le procédé inclut en outre l'étape consistant à collecter les particules ou cellules après la séparation des particules ou cellules.
PCT/US2006/036446 2006-09-19 2006-09-19 Appareil microfluidique de diélectrophorèse à courant continu et ses applications Ceased WO2008036082A1 (fr)

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WO2019003230A1 (fr) * 2017-06-29 2019-01-03 Technion Research & Development Foundation Limited Dispositifs et procédés pour commander un flux à l'aide d'un flux électro-osmotique
CN115074240A (zh) * 2022-06-15 2022-09-20 大连海事大学 一种基于可变形微液滴的介电泳微颗粒多级分选装置及方法
CN117483018A (zh) * 2023-10-17 2024-02-02 清华大学 生物粒子分离装置、加工方法及微流控芯片

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Cited By (4)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
WO2019003230A1 (fr) * 2017-06-29 2019-01-03 Technion Research & Development Foundation Limited Dispositifs et procédés pour commander un flux à l'aide d'un flux électro-osmotique
CN115074240A (zh) * 2022-06-15 2022-09-20 大连海事大学 一种基于可变形微液滴的介电泳微颗粒多级分选装置及方法
CN115074240B (zh) * 2022-06-15 2024-11-15 大连海事大学 一种基于可变形微液滴的介电泳微颗粒多级分选装置及方法
CN117483018A (zh) * 2023-10-17 2024-02-02 清华大学 生物粒子分离装置、加工方法及微流控芯片

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