Patent Application
20170291172
Accurate and high throughput cell sorting technologies are critical for applications in molecular and cellular biology, biotechnology, and medicine (Shields, et al. (2015) Lab Chip, 15: 1230-1249). Biological, chemical, or medical processes involving complex fluids with embedded particles (e.g., blood) often require preparative separation of particles, cells, or even molecules that are needed for subsequent procedures. In conventional macroscale separation processes, centrifugation and membrane filtration approaches have been commonly used for decades, whereas more sophisticated methods such as fluorescence-activated cell sorting (FACS) and magnetically activated cell separation (MACS) were rapidly established as the standard methods for high quality cell and particle separation. While conventional methods can provide highly efficient label-based sorting in short timescales, advances in microfluidics have enabled miniature devices not only offering similar capabilities (Chen et al. (2013) Analyst, 138: 7308-7315; Fan et al. (2013) Biomicrofluidics, 7: 044121), but also unprecedented label-free sorting functions by exploiting a variety of physical parameters as biomarkers, such as cell size, deformability, compressibility, shape, density, size, surface properties, electrical polarizability, magnetic susceptibility and refractive index. Among these, cell size is the most straight-forward feature and sorting based on size can be easily accessible through microfiltration (Mohamed et al. (2004) IEEE Transactions on Nanobioscience, 3: 251-256), pinched flow fractionation (Yamada et al. (2004) Analyt. Chem. 76: 5465-5471), inertial microfluidics (Amini et al. (2014) Lab Chip, 14: 2739-2761), acoustophoresis (Petersson et al. (2007) Analyt. Chem. 79: 5117-5123), and dielectrophoresis (Kim et al. (2007) Proc. Natl. Acad. Sci. USA, 104: 20708-20712). However, a high throughput and reliable approach is still lacking for high purity sorting of particles with small size difference.
Particles in different streamlines in a microfluidic channel flow at different speeds due to the parabolic flow velocity distribution resulting from the zero-slip boundary conditions and viscous fluid environment. Focusing randomly distributed particles into a narrow stream in a continuous flow allows all particles to flow at the same speed and in the same cross-section location. This function is critical for applications that need accurate synchronization and coordination of particles in both space and time domains, for example, flow cytometer, one of the most efficient and effective approaches for single cell analysis. One of the core features of a flow cytometer is the ability to three-dimensionally focus cells and particles into a single stream. This allows all particles and cells to travel at an identical speed through an optical detection zone such that all particles receive the same illumination intensity and time for reliable and consistent optical detection (Cram (2002) Meth. Cell. Sci. 24(1): 1-9). Fluorescence activated cell sorter (FACS) is a special type of flow cytometer that adds a downstream sorting function to select target particles detected in the upstream. Tight particle focusing is extremely important to synchronize the detection and switching event. Without particle focusing, particle flow speed varies and the arrival time of particles into the switching zone is difficult to predict, which results in failure of sorting out target particles or cells. For many other particle sorting techniques that are based on magnetic forces, dielectrophoresis, acoustics, inertial forces, tight particle focusing is also important to provide particles and cells of different properties a common reference in space for high efficiency and high purity sorting.
Particle focusing can be achieved by many prior approaches. Traditionally, sheath fluid is used to sandwich the sample fluid in a co-flow manner. The laminar flow nature in microfluidics allows tight focusing to be achieved by using a large sheath-to-sample fluid ratio. However, serious dilution of sample is a potential concern for many applications. Tuning particle focusing location can be realized by changing the input sheath flows, yet it takes some time before the new equilibrium focusing location becomes stabilized. To eliminate the need of sheath flows for focusing, some other approaches have been developed.
Inertia effects of particles flowing in high-speed flows have recently been found significant in microfluidic channels. Particles can be focused into few equilibrium streams where hydrodynamic forces on particles resulting from flow velocity gradient and wall effects are equal. Single stream microparticle focusing can also be achieved by using secondary flows induced by periodic structures (Chen et al. (2014) Small 10(9): 1746-1751; Chung et al. (2013) Small, 9(5): 685-690; Chung et al. (2014) Lab Chip, 13(15): 2942-2949; Lee et al. (2009) Lab Chip, 9(21): 3155-3160). Some limitations of inertial focusing include size-dependent focusing location, insensitive to small sized particles, focusing stability after passing through focusing structures, unprecise focusing, and the need of high-speed flows. Tuning focusing stream location in real-time might be challenge since the equilibrium positions are dependent upon channel geometry and flow speed.
Acoustic focusing is another broadly applied mechanism. By forming standing acoustic waves in microfluidic channels, either through bulk acoustic waves (BAW) (Antfolk et al. (2014) Lab Chip, 14(15): 2791-2799; Grenvall et al. (2014) Lab Chip, 14(24): 4629-4637; Chen et al. (2014) Lab Chip, 14(5): 916-923; Shi et al. (2008) Lab Chip, 8(2): 221-223; Shi, et al. (2011) Lab Chip, 11(14): 2319-2324) or surface acoustic waves (SAW) (Chen et al. (2014) Lab Chip, 14(5): 916-923; Shi et al. (2008) Lab Chip, 8(2): 221-223; Shi, et al. (2011) Lab Chip, 11(14): 2319-2324), particles and cells can be focused to nodal points. Acoustic focusing approaches can provide tunable focusing functions by either adding an additional echo channel in the BAW cases (Fong et al. (2014) Analyst, 139(5): 1192-1200; Jung et al. (2015) Lab Chip, 15(4): 1000-1003) or using chirped inter-digital transducers (IDT) as in the SAW-based approaches (Ding et al. (2012) Lab Chip, 12(21): 4228-4231; Li et al. (2013) Analyt. Chem., 85(11): 5468-5474). However, such tuning is currently limited to one dimension across the channel, unprecise focusing, and multiple focusing positions across the channel cross section.
Dielectrophoretic forces have also been utilized to provide focusing functions in microfluidics. This electrical based mechanism provides a great tuning flexibility since the applied voltages can be continuously and easily adjusted in real-time. However, limited DEP forces, typically in the orders of tens of pN, prevent most current DEP devices from providing tight focusing functions in high-speed flows (Haandbæk et al. (2014) Lab Chip, 14(17): 3313-3324; Holmes et al. (2006) Biosensors and Bioelectronics, 21(8): 1621-1630; Morgan et al. (2003) IEEE Proc.-Nanobiotechnol, 150(2): 76-81), which also limit their throughputs. Furthermore, DEP devices usually require the manipulated particles and cells to be suspended in isotonic buffers with low ionic strength to increase DEP forces and responsibility of cells. Such low ionic buffers, different from regular physiological buffers, may impact cells' physiological conditions and viability.
Various embodiments contemplated herein may include, but need not be limited to, one or more of the following:
A device for focusing cells, viruses, particles, molecules or molecular complexes in a microfluidic channel, said device comprising:
The device of embodiment 1, wherein said device comprises:
The device according to any one of embodiments 1-2, wherein said device is configured to apply voltages independently to each of said electrodes.
The device according to any one of embodiments 1-3, wherein said device comprises two pairs of electrodes disposed parallel to each other around the microfluidic channel.
The device according to any one of embodiments 1-4, wherein said plurality of electrodes comprises electrodes disposed along each side of said microfluidic channel at or near the top of said channel and electrodes disposed along each side of said microfluidic channel at or near the bottom of said channel.
The device according to any one of embodiments 1-4, wherein said plurality of electrodes comprises electrodes disposed along the midline of each side of said microfluidic channel and along the midline of the top and bottom of said channel.
The device according to any one of embodiments 1-6, said device applies an ac voltage to said electrodes.
The device of embodiment 7, wherein said ac voltage applied to said electrodes is independently at a frequency ranging from about 0 Hz, or from about 1 Hz, or from about 100 Hz, or from about 1 kHz, or from about 10 kHz, or from about 50 kHz, or from about 100 kHz, or from about 500 kHz, up to about 5 MHz, or up to about 10 MHz, or up to about 15 MHz, or up to about 20 MHz, or up to about 50 MHz, or up to about 100 MHz, or up to about 500 MHz, or ranging from about 10 kHz, or from about 50 kHz, or from about 100 kHz, or from about 500 kHz, up to about 5 MHz, or up to about 10 MHz, or up to about 15 MHz, or up to about 20 MHz.
The device according to any one of embodiments 7-8, wherein said voltage applied to said electrodes independently ranges from about close to 0V, or from about 0.001 mV, or from about 0.01 mV, or from about 0.1 mV, or from about 1 mV, or from about 100 mV, or from about 500 mV, or from about 1V, or from about 5V, or from about 10V, up to about 500V, or up to about 100V, or up to about 80V, or up to about 50V, or up to about 40V, or up to maximum voltage above which a fluid in said channel will undergo electrolysis, or ranges from about 1V, or from about 5V, or from about 10V, up to about 100V, or up to about 80V, or up to about 50V, or up to about 40V.
The device according to any one of embodiments 1-9, wherein said electrodes are configured to provide a field minimum at or near a lower or upper corner (diagonal region) of said channel.
The device according to any one of embodiments 1-9, wherein said electrodes are configured to provide a field minimum at or near one side of said channel and/or at or near the top or bottom of said channel.
The device according to any one of embodiments 1-11, wherein said microfluidic channel length is at least about 10 μm, or at least about 100 μm, or at least about 500 μm, or at least about 1 cm, or at least about 2 cm, or at least about 3 cm, or at least about 4 cm, or at least about 5 cm, or at least about 6 cm, or at least about 7 cm, or at least about 8 cm, or at least about 9 cm, or at least about 10 cm, or at least about 25 cm, or at least about 50 cm, or at least about 80 cm, or at least about 100 cm.
The device of embodiment 12, wherein said channel is linear.
The device of embodiment 12, wherein said channel is serpentine.
The device of embodiment 14, wherein said channel is serpentine and has a length greater than a linear length of the substrate in which said channel is disposed.
The device according to any one of embodiments 1-15, wherein the average depth of said microfluidic channel ranges from about 0.1 μm, or from about 0.5 μm, or from about 1 μm, or from about 10 μm, or from about 20 μm, or from about 30 μm, up to about 100 μm, or up to about 80 μm, or up to about 60 μm, or up to about 50 μm, or up to about 40 μm.
The device according to any one of embodiments 1-16, wherein the average width of said microfluidic channel ranges from about 0.1 μm, or from about 0.5 μm, or from about 1 μm, or from about 10 μm, or from about 20 μm, or from about 30 μm, or from about 40 μm, or from about 50 μm, or from about 80 μm, or from about 100 μm up to about 500 μm, or up to about 400 μm, or up to about 300 μm, or up to about 200 μm, or up to about 400 μm, or up to about 500 μm, or up to about 1 mm.
The device according to any one of embodiments 1-17, wherein said fluid has a conductivity that ranges from about 10−6 S/m, or from about 10−5 S/m, or from about 10−4 S/m, or from about 10−3 S/m, or from about 10−2 S/m up to about 10 S/m, or up to about 5 S/m, or up to about 2 S/m, or up to about 1.5 S/m, or up to about 1 S/m.
The device according to any one of embodiments 1-18, wherein said fluid comprises a physiological buffer.
The device of embodiment 19, wherein said buffer comprises a mammalian ringer's solution.
The device of embodiment 18, wherein said fluid comprises PBS.
The device according to any one of embodiments 18-21, wherein the conductivity of said fluid is about 1 S/m.
The device according to any one of embodiments 1-22, wherein said hydrodynamic flows are at a rate ranging up to about 10 m/s, or up to about 5 m/s, or up to about 1 m/s, or up to about 50 cm/s, or up to about 20 cm/s, or up to about 15 cm/s, or up to about 11 cm/s, or up to about 10 cm/s, or up to about 8 cm/s, or up to about 5 cm/s, or up to about 3 cm/s, or up to about 1 cm/s, or up to about 500 μm/s, or up to about 250 μm/s, or up to about 100 μm/s, or up to about 50 μm/s, or up to about 30 μm/s, or up to about 20 μm/s, or up to about 10 μm/s.
The device according to any one of embodiments 1-23, wherein channel is fabricated from a material selected from the group consisting of silicon, a plastic, and an elastomeric material.
The device of embodiment 24, wherein said elastomeric material is selected from the group consisting of polydimethylsiloxane (PDMS), polyolefin plastomers (POPs), perfluoropolyethylene (a-PFPE), polyurethane, polyimides, and cross-linked NOVOLAC® (phenol formaldehyde polymer) resin.
The device of embodiment 24, wherein said channel is fabricated from PDMS.
A method of focusing cells, organelles, viruses, particles, molecules or molecular complexes to an off-center location in a microchannel, said method comprising:
introducing said cells, organelles, viruses, particles, molecules or molecular complexes into a device according to any one of embodiments 1-26, wherein said electrodes provide an electric field minimum that is not centered in said microfluidic channel; and
flowing said cells, organelles, viruses, particles, molecules or molecular complexes along a length of the channel sufficient to permit said cells, organelles, viruses, particles, molecules or molecular complexes to focus in said channel at an off-center location wherein said off-center location is the location of an electric field minimum.
The method of embodiment 27, wherein said flowing comprises flowing said cells or particles along at least about 100 μm, or at least about 500 μm, or at least about 1 cm, or at least about 2 cm, or at least about 3 cm, or at least about 4 cm, or at least about 5 cm, at least about 10 cm, or at least about 25 cm, or at least about 50 cm, or at least about 80 cm, or at least about 100 cm of said channel.
The method according to any one of embodiments 27-28, wherein said flowing comprises flowing said cells or particle at a rate ranging up to about 10 m/s, or up to about 5 m/s, or up to about 3 m/s, or up to about 1 m/s, or up to about 50 cm/s, or up to about 20 cm/s, or up to about 15 cm/s, or up to about 11 cm/s, or up to about 10 cm/s, or up to about 8 cm/s, or up to about 5 cm/s, or up to about 3 cm/s, or up to about 1 cm/s, or up to about 500 μm/s, or up to about 250 μm/s, or up to about 100 μm/s, or up to about 50 μm/s, or up to about 25 μm/s, or up to about 10 μm/s.
The method according to any one of embodiments 27-29. wherein said cells, viruses, particles, molecules or molecular complexes comprise a moiety selected from the group consisting of a particle, a biological molecule, a biological complex, an immune complex, a liposome, a protoplast, a platelet, a bacterium, a virus, and a prokaryotic cell, and a eukaryotic cell.
The method according to any one of embodiments 27-30, wherein said cells, viruses, particles, molecules or molecular complexes comprise a particle.
The method according to any one of embodiments 27-30, wherein said cells, viruses, particles, molecules or molecular complexes comprise a cell.
The method of embodiment 32, wherein said cell comprises a prokaryotic cell.
The method of embodiment 32, wherein said cell comprises a eukaryotic cell.
The method of embodiment 34, wherein said cell comprises a mammalian cell.
The method according to any one of embodiments 31-35, wherein said cell is in a physiological buffer.
The method of embodiment 32, wherein said cell is in an isotonic buffer.
A device for sorting cells, organelles, viruses, particles, molecules or molecular complexes, said device comprising:
The device of embodiment 38, wherein said first location is at or near a wall of said channel and said second location is at or near the opposite wall of said channel.
The device of embodiment 38, wherein said first location is at or near a corner of said channel and said second location is diagonally opposite at or near a corner of said channel.
The device according to any one of embodiments 38-40, wherein the second region of said channel diverges into a plurality of channels whereby different size particle are diverted into different channels providing particle having different size or size distribution in each different channel of said plurality of channels.
The device of embodiment 41, wherein said second region diverges into two different channels.
The device of embodiment 41, wherein said second region diverges into 3, 4, 5, 6, 7, 8, 9, or 10 or more channels.
The device according to any one of embodiments 38-43, wherein said device comprises a port or channel for introducing said cells, viruses, particles, molecules or molecular complexes into the first region of said channel.
The device according to any one of embodiments 38-44, wherein said device comprises a port or channel for introducing a sheath flow into said microfluidic channel.
The device according to any one of embodiments 38-45, wherein said first plurality of electrodes and said second plurality of electrodes independently each comprise two pairs of electrodes disposed parallel to each other around that region of the microfluidic channel.
The device according to any one of embodiments 38-46, wherein said first plurality of electrodes and said second plurality of electrodes each comprises electrodes disposed along each side of said microfluidic channel at or near the top of said channel and electrodes disposed along each side of said microfluidic channel at or near the bottom of said channel.
The device according to any one of embodiments 38-46, wherein said first plurality of electrodes and said second plurality of electrodes each comprises electrodes disposed along the midline of each side of said microfluidic channel and along the midline of the top and bottom of said channel.
The device according to any one of embodiments 38-48, said device applies an ac voltage to first plurality of electrodes and to said second plurality electrodes.
The device of embodiment 49, wherein said ac voltage applied to said first plurality of electrodes and to said second plurality of electrodes is independently at a frequency from about 0 Hz, or from about 1 Hz, or from about 100 Hz, or from about 1 kHz, or from about 10 kHz, or from about 50 kHz, or from about 100 kHz, or from about 500 kHz, up to about 5 MHz, or up to about 10 MHz, or up to about 15 MHz, or up to about 20 MHz, or up to about 50 MHz, or up to about 100 MHz, or up to about 500 MHz, or ranging from about 10 kHz, or from about 50 kHz, or from about 100 kHz, or from about 500 kHz, up to about 5 MHz, or up to about 10 MHz, or up to about 15 MHz, or up to about 20 MHz.
The device according to any one of embodiments 49-50, wherein said voltage applied to said first plurality of electrodes and to said second plurality of electrodes independently ranges from about close to 0V, or from about 0.001 mV, or from about 0.01 mV, or from about 0.1 mV, or from about 1 mV, or from about 100 mV, or from about 500 mV, or from about 1V, or from about 5V, or from about 10V, up to about 500V, or up to about 100V, or up to about 80V, or up to about 50V, or up to about 40V, or up to maximum voltage above which a fluid in said channel will undergo electrolysis, or ranges from about 1V, or from about 5V, or from about 10V, up to about 100V, or up to about 80V, or up to about 50V, or up to about 40V.
The device according to any one of embodiments 38-51, wherein said first region and said second region each independently range in length up to about 3 cm, or up to about 4 cm, or up to about 5 cm, or up to about 6 cm, or up to about 7 cm, or up to about 8 cm, or up to about 9 cm, or up to about 10 cm, or at least about 25 cm, or at least about 50 cm, or at least about 80 cm, or at least about 100 cm.
The device of embodiment 52, wherein at least a portion of said channel is linear.
The device of embodiment 52, wherein all of said channel is linear.
The device of embodiment 52, wherein at least a portion of said channel is serpentine.
The device of embodiment 55, wherein said first region is serpentine.
The device according to any one of embodiments 55-56, wherein said second region is serpentine.
The according to any one of embodiments 55-57, wherein at least a portion of said channel is serpentine and said channel has a length greater than a linear length of the substrate in which said channel is disposed.
The device according to any one of embodiments 38-58, wherein the average depth of said microfluidic channel ranges from about 0.1 μm, or from about 0.5 μm, or from about 1 μm, or from about 10 μm, or from about 20 μm, or from about 30 μm, up to about 100 μm, or up to about 80 μm, or up to about 60 μm, or up to about 50 μm, or up to about 40 μm.
The device according to any one of embodiments 38-59, wherein the average width of said microfluidic channel ranges from about 0.1 μm, or from about 0.5 μm, or from about 1 μm, or from about 10 μm, or from about 20 μm, or from about 30 μm, or from about 40 μm, or from about 50 μm, or from about 80 μm, or from about 100 μm up to about 500 μm, or up to about 400 μm, or up to about 300 μm, or up to about 200 μm, or up to about 400 μm, or up to about 500 μm, or up to about 1 mm.
The device according to any one of embodiments 38-60, wherein said fluid has a conductivity that ranges from about 10−6 S/m, or from about 10−5 S/m, or from about 10−4 S/m, or from about 10−3 S/m, or from about 10−2 S/m up to about 10 S/m, or up to about 5 S/m, or up to about 2 S/m, or up to about 1.5 S/m, or up to about 1 S/m.
The device according to any one of embodiments 38-61, wherein said fluid comprises a physiological buffer.
The device of embodiment 19, wherein said buffer comprises a mammalian ringer's solution.
The device of embodiment 61, wherein said fluid comprises PBS.
The device according to any one of embodiments 61-64, wherein the conductivity of said fluid is about 1 S/m.
The device according to any one of embodiments 38-65, wherein said hydrodynamic flows are at a rate ranging up to about 10 m/s, or up to about 5 m/s, or up to about 1 m/s, or up to about 50 cm/s, or up to about 20 cm/s, or up to about 15 cm/s, or up to about 11 cm/s, or up to about 10 cm/s, or up to about 8 cm/s, or up to about 5 cm/s, or up to about 3 cm/s, or up to about 1 cm/s, or up to about 500 μm/s, or up to about 250 μm/s, or up to about 100 μm/s, or up to about 50 μm/s, or up to about 30 μm/s, or up to about 20 μm/s, or up to about 10 μm/s.
The device according to any one of embodiments 38-66, wherein said channel is fabricated from a material selected from the group consisting of silicon, a plastic, and an elastomeric material.
The device of embodiment 67, wherein said elastomeric material is selected from the group consisting of polydimethylsiloxane (PDMS), polyolefin plastomers (POPs), perfluoropolyethylene (a-PFPE), polyurethane, polyimides, and cross-linked NOVOLAC® (phenol formaldehyde polymer) resin.
The device of embodiment 67, wherein said channel is fabricated from PDMS.
The device according to any one of embodiments 38-69, wherein said device can separate a 9 μm particle from a 10 μm particle.
The device of embodiment 70, wherein said device can separate a 9 μm particle from a 10 μm particle at a flow rate of 3 cm/s.
The device according to any one of embodiments 38-71, wherein said first region provides a 3D tunable, size-independent, single-stream focusing having sub-micron precision.
The device of embodiment 72, wherein said focusing precision of said first region is less than about 0.2 μm.
The device according to any one of embodiments 38-73, wherein said second region provides a 3D tunable, size-independent, single-stream focusing having sub-micron precision.
The device of embodiment 74, wherein said focusing precision of said second region is less than about 0.2 μm.
The device according to any one of embodiments 72-75, wherein said focusing precision is at a flow rate of about 3 cm/s.
The device according to any one of embodiments 38-76, wherein said device provides sorting purity of greater than about 90%, or greater than about 94%, or greater than about 98%, or greater than about 99%.
The device according to any one of embodiments 38-77, wherein said device is a component of a lab on a chip.
A method of sorting cells, organelles, viruses, particles, molecules, or molecular complexes, said method comprising:
introducing said cells, organelles, viruses, particles, molecules or molecular complexes into a device according to any one of embodiments 38-77; and
capturing said cells, organelles, viruses, particles, molecules, or molecular complexes from said device that have been sorted by size.
In various embodiments a novel dielectrophoretic (DEP) mechanism for tunable, sheathless, three dimensional, and single-stream microparticle and cell focusing in high-speed flows is provided. In certain embodiments, it is realized by fabricating a 3D microfluidic device with two substrates (e.g. glass substrates sandwiching a thin and open microfluidic channel (e.g., a PDMS channel). Electrodes are laid out to provide DEP forces completely perpendicular to hydrodynamic flows along the channel (see, e.g.,
In certain embodiments a device is provided for focusing cells, viruses, organelles, particles, molecules. molecular complexes, and the lie in a microfluidic channel, where the device comprises a microfluidic channel comprising a plurality of electrodes disposed on surfaces of the channel to provide three-dimensional spatially tunable tunnel electric field minimum for dielectrophoretic (DEP) forces that are perpendicular to hydrodynamic flows along the channel. In certain embodiments the device comprises a fluid within the channel providing hydrodynamic flow along the channel. In certain embodiments the device is configured to apply voltages to the electrodes to provide an spatially adjustable electric field minimum or electric field pattern that is programmable by the voltage combinations on the electrodes. In certain embodiments the focusing is obtained without sheath flow. In certain embodiments the microfluidic channel comprises a plurality of electrodes disposed to provide dielectrophoretic (DEP) forces that are perpendicular to hydrodynamic flows along the channel; and the device is configured to apply voltages to the electrodes to provide an electric field minimum that is not centered in said microfluidic channel.
In certain embodiments the device is device is configured to apply voltages independently to each of the electrodes. In certain embodiments the device comprises two pairs of electrodes disposed parallel to each other around the microfluidic channel. In certain embodiments the plurality of electrodes comprises electrodes disposed along each side of the microfluidic channel at or near the top of the channel and electrodes disposed along each side of the microfluidic channel at or near the bottom of said channel. In certain embodiments the plurality of electrodes comprises electrodes disposed along the midline of each side of the microfluidic channel and along the midline of the top and bottom of the channel.
In certain embodiments the device is configured to provide, and/or applies an a.c. voltage to one or more, or to two or more, or to three or more, or to all of the electrodes.
In certain embodiments the device is configured to provide the voltages described above by integration of a voltage source (e.g., one or more power supplies, signal generators, etc.). In certain embodiments the device is configured to provide the voltages described above by integration of voltage regulators that can adjust one or more externally applied voltages. In certain embodiments the device is configured to provide the voltages described above by electrical coupling to one or more external voltage sources (e.g., power supplies).
In certain embodiments the device is configured to apply to the electrodes and/or the voltage applied to the electrodes is independently (e.g., independently regulated voltage(s)) at a frequency ranging from about 0 Hz, or from about 1 Hz, or from about 100 Hz, or from about 1 kHz, or from about 10 kHz, or from about 50 kHz, or from about 100 kHz, or from about 500 kHz, up to about 5 MHz, or up to about 10 MHz, or up to about 15 MHz, or up to about 20 MHz, or up to about 50 MHz, or up to about 100 MHz, or up to about 500 MHz, or ranging from about 10 kHz, or from about 50 kHz, or from about 100 kHz, or from about 500 kHz, up to about 5 MHz, or up to about 10 MHz, or up to about 15 MHz, or up to about 20 MHz.
In certain embodiments the device is configured to apply to the electrodes and/or the voltage applied to the electrodes is independently (e.g., independently regulated voltage(s)) that range from about 0V, or from about 0.001 mV, or from about 0.01 mV, or from about 0.1 mV, or from about 1 mV, or from about 100 mV, or from about 500 mV, or from about 1V, or from about 5V, or from about 10V, up to about 500V, or up to about 100V, or up to about 80V, or up to about 50V, or up to about 40V, or up to maximum voltage above which a fluid in said channel will undergo electrolysis, or ranges from about 1V, or from about 5V, or from about 10V, up to about 100V, or up to about 80V, or up to about 50V, or up to about 40V.
In certain embodiments the electrodes are configured (e.g., the voltages applied to the electrodes are selected) to provide a field minimum at or near a lower or upper corner (diagonal region) of the channel. In certain embodiments the electrodes are configured (e.g., the voltages applied to the electrodes are selected) to provide a field minimum at or near one side or near the top or bottom of the channel.
In certain embodiments the microfluidic channel length is at least about 1 μm, or at least about 10 μm, or at least about 100 μm, or at least about 500 μm, or at least about 1 cm, or at least about 2 cm, or at least about 3 cm, or at least about 4 cm, or at least about 5 cm, or at least about 6 cm, or at least about 7 cm, or at least about 8 cm, or at least about 9 cm, or at least about 10 cm, or at least about 25 cm, or at least about 50 cm, or at least about 80 cm, or at least about 100 cm. In certain embodiments the microfluidic channel is linear. In certain embodiments the microfluidic channel or a portion of the channel is serpentine. In certain embodiments the microfluidic channel is serpentine and has a length greater than a linear length of the substrate in which said channel is disposed.
In certain embodiments the wherein the average depth of the microfluidic channel ranges from about 0.1 μm, or from about 0.5 μm, or from about 1 μm, or from about 10 μm, or from about 20 μm, or from about 30 μm, up to about 100 μm, or up to about 80 μm, or up to about 60 μm, or up to about 50 μm, or up to about 40 μm. In certain embodiments the average width of said microfluidic channel ranges from about 0.1 μm, or from about 0.5 μm, or from about 1 μm, or from about 10 μm, or from about 20 μm, or from about 30 μm, or from about 40 μm, or from about 50 μm, or from about 80 μm, or from about 100 μm up to about 500 μm, or up to about 400 μm, or up to about 300 μm, or up to about 200 μm, or up to about 400 μm, or up to about 500 μm, or up to about 1 mm.
In certain embodiments devices for sorting cells, organelles, viruses, particles, molecules, or molecular complexes are provided. In certain embodiments the device comprises a microfluidic channel comprising a first region comprising a first plurality of electrodes disposed to provide dielectrophoretic (DEP) forces that are perpendicular to hydrodynamic flows along the first region of the channel (e.g., as described above); and a second region downstream from the first region comprising a second plurality of electrodes disposed to provide dielectrophoretic (DEP) forces that are perpendicular to hydrodynamic flows along the second region of the channel (e.g., as described above; and where the device is configured to apply voltages to first plurality of electrodes to provide an electric field minimum at a first location in the cross-section of the channel and to apply voltages to the second plurality of electrodes to provide an electric field minimum at a second location in the cross-section of the channel, where the first location and the second location are different locations in the cross-section of the channel (see, e.g.,
In certain embodiments the device comprises a port or channel for introducing the cells, organelles, viruses, particles, molecules or molecular complexes into the first region of the channel. In certain embodiments the device comprises a port or channel for introducing a sheath flow into the microfluidic channel. In certain embodiments the device has sufficient resolution to separate a 9 μm particle from a 10 μm particle, e.g., at a flow rate of 3 cm/s.
In certain embodiments the first region and/or the second region provides a 3D tunable, size-independent, single-stream focusing having sub-micron precision. In certain embodiments the focusing precision is less than about 0.2 μm (e.g., at a flow rate of about 3 cm/s). In certain embodiments the device provides sorting purity of greater than about 90%, or greater than about 94%, or greater than about 98%, or greater than about 99%.
In various embodiments the device can be incorporated with other components. In certain embodiments the device is a component of a lab on a chip.
In certain embodiments methods of sorting cells, organelles, viruses, particles, molecules, or molecular complexes are provided where the methods involve introducing the cells, organelles, viruses, particles, molecules, or molecular complexes into a sorting device as described and claimed herein, and capturing the cells, organelles, viruses, particles, molecules, or molecular complexes from the device that have been sorted by size.
To fabricate the device, utilize a new fabrication method previously reported by our group (see, e.g., Kung et al. (2015) Lab Chip, 15: 1861-1868) was used. A large area (6 cm×2.2 cm) PDMS thin film with an open microchannel (H W, 83 μm 80 μm) is sandwiched between a glass slide and a coverslip with strip electrodes aligned to the four corners of the channel. This allows the creation of a three-dimensional electric field profile through the entire focusing channel.
The focusing positions of microparticles can be precisely predicted by numerical simulation in COMSOL (
The snapshot images (
In another illustrative, but non-limiting embodiment the DEP mechanism described above, is exploited for ultra-high precision microparticle and cell focusing and separation in high-speed flows. For the first time, particle size differences as small as 1 μm can be separated with high purity (>90%). This is realized by a 3D tunable, size-independent, single-stream sub-micron precision (variation <0.2 μm) focusing function. As illustrated in
One of skill in the art would recognize that the embodiments described herein are illustrative and non-limiting. For example, in certain embodiments, electrodes could be disposed the middle of the top and bottom and at the middle of each side of the microfluidic channel. Similarly, the microfluidic channel can be fabricated in or from any of a number of materials including, but not limited to silicon/glass, a plastic, an elastomeric material (e.g., polydimethylsiloxane (PDMS), polyolefin plastomers (POPs), perfluoropolyethylene (a-PFPE), polyurethane, polyimides, and cross-linked NOVOLAC® (phenol formaldehyde polymer), and the like). Using the teachings provided herein, numerous variations of the illustrated and described devices and methods will be available to one of skill in the art.
The following examples are offered to illustrate, but not to limit the claimed invention.
Here, we demonstrate a novel tunnel dielectrophoresis (TDEP) mechanism for tunable, sheathless, three dimensional, and single-stream microparticle and cell focusing in high-speed flows. It is realized by fabricating a 3D microfluidic device with two glass substrates sandwiching a thin and open PDMS channel (see, e.g., Kung et al. (2015) Lab Chip, 15: 1861-1868 for an illustrative fabrication protocol). Electrodes are laid out to provide DEP forces completely perpendicular to the hydrodynamic flow along a channel that can range in length up to several centimeters (e.g., up to about 1 cm, or up to about 2 cm, or up to about 3 cm, or up to about 4 cm, or up to about 5 cm, or up to about 6 cm, or up to about 7 cm, or up to about 8 cm, or up to about 9 cm, or up to about 10 cm. This provides a long DEP interaction such that microparticles and cells have sufficient time to migrate to the focused stream even in high-speed flows.
Dielectrophoresis refers to the interaction force between a non-uniform electric field and the dipole moment it induces on a polarizable object. The magnitude of DEP force on a spherical particle can be expressed by the following formula derived based on a diploe approximation:
F
DEP
(1)=2π∈mR3Re[CM(ω)×∇Ē2] (1)
where FDEP(1) refers to the DEP force, £m the permittivity of the medium surrounding the sphere, R the radius of the particle, w the radian frequency of the applied field, and Ē the applied electric field. CM is the Clausius-Mossotti (CM) factor given by
where and ∈p and ∈m are the complex permittivities of the particle and the medium, respectively, and ∈=∈+σ/(jω), where ∈ is the permittivity, and σ the conductivity. The magnitude of DEP force is linearly proportional to the gradient of electric field strength (E2) and the volume of particles. For particles more polarizable than the medium, Re [CM]>0, they experience positive DEP forces that move them toward the strong electric field region. On the other hand, if Re[CM]<0, particles migrate to the weak electric field region.
DEP manipulation on mammalian cells is usually conducted in low ionic isotonic buffers for several reasons. One is that different types of mammalian cells suspended in low ionic buffers (0.01 S/m˜0.1 S/m) can show very distinct dielectric signatures and CM curves, which makes cell sorting easier to perform. Second, higher voltage can be applied to electrodes to generate larger DEP forces on cells without inducing electrolysis on electrodes or causing significant heating. Yet, suspending cells in isotonic buffers with low ionic strength over a long time may impact cells' viability.
The challenge of manipulating mammalian cells in regular physiological buffers is that only negative DEP forces can be induced on cells since the Re[CM] is negative and small over the entire frequency spectrum. In addition, high frequency, typically 1 MHz to 10 MHz, and low voltage operation is required to prevent electrolysis on electrodes. This limits the DEP forces that can be induced on mammalian cells in physiological buffers. The design of TDEP has several unique features to solve these challenges. TDEP is for focusing particles and cells with negative DEP forces. The four independently voltage controlled quadro-electrodes ensure that there is only one potential minimum in the cross section of a microfluidic channel. The location of potential minimum can be real-time adjusted by changing the voltage combinations applied to electrodes. In TDEP, there is only one focused position for all particles, regardless of their sizes and types, as long as they show negative DEP responses. This is an important feature for applications, such as flow cytometers, that need all particles to flow at the same speed, pass through the same location in a channel for light detection or imaging, and be at the same reference position for downstream applications. The extremely long DEP interaction distance throughout the entire channel allows cells in physiological buffers with weak negative DEP forces to have sufficient time to migrate to the focused location in high-speed flows. For example, if the DEP induced cell migration speed is 80 μm/sec, and the channel width and height are both 80 μm. It takes an average of 0.5 second for cells near the channel wall to migrate to the center of the channel. If the DEP interaction distance is 6 cm, the length used in TDEP devices presented in this paper, single stream focusing can be realized at an average flow speed of 12 cm/sec, three orders of magnitude higher than prior DEP focusing devices that operate at a flow speed<100 μm/sec in regular physiological buffer (Gao et al. (2012) Analyst, 137(22): 5215-5221). Long DEP channel can be easily fabricated by utilizing a serpentine channel design. However, running cells at a particle Renold's number (Rp)>0.5 (Carlo et al. (2007) Proc. Natl. Acad. Sci. USA, 104(48): 18892-18897; Hur et al. (2010) Lab Chip, 10(3): 274-280) (corresponding flow velocity ˜40 cm/s in this design) in a microfluidic channel may encounter inertial effects that affect the DEP focusing functions. More studies are required to understand the coupling between inertial forces and DEP forces in this regime.
Another unique feature of TDEP is that the DEP forces on cells and particles only exist in the direction perpendicular to the channel and completely decouple from the hydrodynamic forces that carry particles to flow along a streamline in the channel. This means that particles' transverse migration driven by DEP forces is not affected by the flow speeds in channels. Unlike prior DEP devices that utilized titled electrode design for cell focusing, DEP forces and hydrodynamic forces are not decoupled. Particle focusing behavior is highly dependent upon flow speeds, particle sizes, and particle types. (Han et al. (2008) Lab Chip, 8(7): 1079-1086; Hu et al. (2005) Proc. Natl. Acad. Sci. USA, 102(44): 15757-15761; Kim et al. (2008) Anal. Chem., 80(22): 8656-8661; Kim et al. (2007) Proc. Natl. Acad. Sci. USA, 104(52): 20708-20712).
3D Tunable Focusing
The four independently controlled quadro-electrodes provide the real-time tuning function to adjust the location of potential minimum and particle focused stream.
These simulation results are confirmed by experimental data shown in
Size Independent Focusing
In
Mammalian Cell Focusing and Viability Study
The snapshot images in
In order to verify the viability effect before and after DEP focusing operation, short-term and long-term viability tests were carried out on HeLa cells. In the short-term viability test, HeLa cells in condition of
To achieve continuous and high size precision sorting two stages of particle manipulation methodologies can be utilized. In the first stage, all particles, regardless of their different sizes, are three-dimensionally focused into a single-stream in a continuous flow such that different sizes of particles have exactly the same reference position. In the second stage, particles migrate to a new focusing position under a new set of boundary conditions. Due to different forces acting on particles of different sizes, particles of different migration speeds can be sorted out and collected. A size-independent, tight upstream particle focusing is key for the downstream high purity sorting of particles with minor size differences.
Here, we demonstrate a novel DEP device that can provide a microfluidic device capable of providing tunable and particle size-independent, sub-micron precision single-stream focusing in the upstream (Kung et al. (2015) Tunable, Sheathless, and Three Dimensional Single-Stream Cell Focusing in High Speed Flows, in The 19th International Conference on Miniaturized Systems for Chemistry and Life Sciences, Gyeongju, Korea), followed by high purity sorting of particles with size difference smaller than 1 μm in the downstream. Furthermore, such sorting was achieved at flow speeds up to 3 cm/s in regular physiological buffers, without the need to swap the medium to a low ionic isotonic buffer, which may affect cells' viability and physiological conditions.
Dielectrophoresis (DEP) is a phenomenon in which a particle in a non-uniform electric field can experience an electrostatic force moving it towards a stronger electric field region if it is more polarizable than the medium, or to a weaker electric field region if the particle is less polarizable than the medium.
Step 1: Fabrication of master molds. SU-8 mold masters on silicon wafers were fabricated using photolithography (
Step 2: Fabrication of hybrid stamps. It starts from preparing the Sylgard 184 silicone elastomer mixture (Dow Corning Corporation, Miland, USA). The weight ratio of Base:Curing agent is 10:1. Few drops of this mixture are poured into a petri dish. A suitable size of polystyrene plastic plate is cut and pressed against the bottom of the petri dish under a pressure of 3 psi. A thin layer of polydimethylsiloxane (PDMS) with a thickness of roughly 30 μm is formed between the petri dish and the plastic plate. Additional uncured PDMS is poured to fill up the petri dish, and followed by a curing step at 60° C. in an oven for 12 hours. A hybrid stamp is formed when the plastic plate together with a thin PDMS layer on its surface is peeled off from the petri dish (
Step 3: Demolding PDMS films from the master mold. During the demolding process, the cured PDMS thin film has stronger adhesion to the hybrid stamp than the master mold since more PFOCTS is coated on the master mold due to a longer treatment time (
Step 4: Transfer the PDMS thin film. Oxygen plasma treatment is performed on both the PDMS thin film on the hybrid stamp and the substrate with strip electrodes to be bonded. The alignment between channel and strip electrode is needed. (
Step 5: Removing the hybrid stamp. It starts from peeling off the bulk PDMS part on the plastic plate (
Step 6: Align and cover the device with a top coverslip with strip electrodes by oxygen plasma bonding to finish the device fabrication. (
The histograms of particle positions at locations in
To demonstrate the bio-compatibility of this platform, we spiked rare GFP-HeLa cells into lysed human whole blood, and did the high purity size-based sorting. The regular sizes of GFP-HeLa cells and human white blood cells fall in the range of 15 μm˜20 μm and 8 μm˜12 μm, respectively. The sorting results are shown in
It is understood that the examples and embodiments described herein are for illustrative purposes only and that various modifications or changes in light thereof will be suggested to persons skilled in the art and are to be included within the spirit and purview of this application and scope of the appended claims. All publications, patents, and patent applications cited herein are hereby incorporated by reference in their entirety for all purposes.
This application claims priority to and benefit of U.S. Ser. No. 62/321,133, filed on Apr. 11, 2016, which is incorporated herein by reference in its entirety for all purposes.
This invention was made with government support under Grant Nos. DBI 1256178 and ECCS1232279 awarded by the National Science Foundation. The Government has certain rights in this invention.
| Number | Date | Country | |
|---|---|---|---|
| 62321133 | Apr 2016 | US |
