HK1242783B - Sorting particles in a microfluidic device - Google Patents

Description

Sorting particles in microfluidic devices

CROSS-REFERENCE TO RELATED APPLICATIONS

This application claims the benefit of U.S. Provisional Patent Application No. 62/074,213, filed November 3, 2014, and U.S. Provisional Patent Application No. 62/074,315, filed November 3, 2014, each of which is incorporated herein by reference in its entirety.

Technical Field

The present disclosure relates to particle sorting across microfluidic flow lines.

Background Art

Particle separation and filtration are used in numerous applications across diverse industries and fields. Examples include chemical process and fermentation filtration, water purification/wastewater treatment, separation and filtration of blood components, concentration of colloidal solutions, and purification and concentration of environmental samples. Various macroscale techniques have been developed for these applications, including methods such as centrifugation and filtration-based techniques. These techniques typically require large, bulky, and expensive systems with complex moving parts.

In some cases, microscale technologies offer advantages over macroscale technologies in that the reduction in scale allows for the use of unique fluid dynamic effects for particle sorting and filtration, thereby eliminating the need for large systems with complex moving parts. Furthermore, microscale technologies enable portable devices that can perform sorting and filtration at a much lower cost than larger macroscale systems. However, typical microscale sorting and filtration devices may be limited in the amount of fluid that can be processed in a given period of time (i.e., low throughput), which may place such devices at a disadvantage compared to their macroscale counterparts.

Summary of the Invention

The present disclosure is based, at least in part, on the discovery that if the geometry and dimensions of a microfluidic device are carefully controlled, fluid extraction and inertial lift can be combined to sort and/or move particles within or between fluids. In particular, the microfluidic devices disclosed herein can be used to transport fluids to and across different fluid channels of the device, without accompanying particle movement, so that particles can be indirectly transferred to another fluid. Alternatively or additionally, the techniques disclosed herein can be used in certain embodiments to manipulate not only the transfer of fluids across microchannels, but also the position of particles suspended within a fluid sample by moving the particles across fluid streamlines.

For example, a first fluid containing particles can be introduced into a first microfluidic channel having an array of rigid island structures that separate the channel from two adjacent microfluidic channels. The fluid is extracted from the first microfluidic channel to one of the adjacent microfluidic channels through the gaps between the island structures in the first array, so that the particles are drawn closer to the island structures. When the particles are closer to the island structures, the particles are subjected to inertial lift forces that deviate from the direction of fluid extraction, causing the particles to remain in the first channel. At the same time, a second fluid enters the first microfluidic channel from another adjacent microfluidic channel through the gaps between the island structures in the second array. When the fluid enters the first channel from another adjacent channel, the particles in the first channel cross the fluid streamlines, causing the particles to shift from the first fluid to the second fluid. If the amount of the first fluid extracted from the first channel at each gap in the first array is equal to the amount of the second fluid entering the first channel at each gap in the second array, a constant particle concentration can be maintained.

In addition to moving particles between fluids, the combination of fluid extraction and inertial lift allows for a variety of different ways of manipulating fluids and particles. For example, in some embodiments, different types of particles can be divided into different channels, such as large particles can be separated from small particles, to achieve micro-scale particle sorting and/or filter particles from a fluid. Alternatively, in some embodiments, the combination of fluid extraction and inertial lift can be used for mixing different types of particles. In some cases, particle separation and displacement between fluids (or particle mixing and displacement between fluids) can be carried out together. In other examples, the combination of fluid extraction and inertial lift can be used for particles to be gathered into desired positions in a microfluidic channel. These and other applications can be expanded to a large number of microfluidic channels, to achieve fluid sorting/filtration of high throughput in a system with low device manufacturing cost.

Generally speaking, in one aspect, the subject matter of the present disclosure can be embodied as a microfluidic device comprising a particle sorting region having a first microfluidic channel, a second microfluidic channel extending along the first microfluidic channel, and a first array of islands separating the first microfluidic channel from the second microfluidic channel, wherein each island in the array is separated from an adjacent island in the array by an opening that fluidically connects the first microfluidic channel to the second microfluidic channel, and wherein the first microfluidic channel, the second microfluidic channel, and the first array of islands are arranged such that a fluid resistance of the first microfluidic channel varies relative to a fluid resistance of the second microfluidic channel along a longitudinal cross-section of the particle sorting region such that a portion of fluid from a fluid sample in either the first microfluidic channel or the second microfluidic channel passes through the opening.

In general, in another aspect, the subject matter of the present disclosure can be embodied as a microfluidic device comprising: a particle sorting region having a first microfluidic channel, a second microfluidic channel extending along a first side of the first microfluidic channel, a first array of islands separating the first microfluidic channel from the second microfluidic channel, wherein each island in the first array is separated from an adjacent island in the first array by an opening fluidically connecting the first microfluidic channel to the second microfluidic channel, a third microfluidic channel extending along a second side of the first microfluidic channel, the second side being opposite the first side of the first microfluidic channel, a second array separating the first microfluidic channel from the third microfluidic channel, wherein each island in the second array is fluidically connected to an adjacent island in the second array by an opening fluidically connecting the first microfluidic channel to the third microfluidic channel. The openings of the channels are separated, wherein the first microfluidic channel, the second microfluidic channel and the first array of islands are arranged so that the fluid resistance of the second microfluidic channel decreases relative to the fluid resistance of the first microfluidic channel along the longitudinal direction of the particle sorting region, so that during operation of the microfluidic device, a portion of the fluid from the fluid sample in the first microfluidic channel passes through the first array into the second microfluidic channel, and wherein the first microfluidic channel, the third microfluidic channel and the second array of islands are arranged so that the fluid resistance of the third microfluidic channel increases relative to the fluid resistance of the first microfluidic channel along the longitudinal direction of the particle sorting region, so that during operation of the microfluidic device, a portion of the fluid from the fluid sample in the third microfluidic channel passes through the second array into the first microfluidic channel.

Embodiments of these devices can have one or more of the following features. For example, in certain embodiments, the change in fluid resistance is a function of the increase in cross-sectional area of the first microfluidic channel or the second microfluidic channel along the longitudinal direction of the particle sorting region. For example, the width of one of the first microfluidic channel or the second microfluidic channel can increase along the longitudinal direction. The width of the other of the first microfluidic channel or the second microfluidic channel can be substantially constant along the longitudinal direction. Alternatively, the width of the other of the first microfluidic channel or the second microfluidic channel can decrease along the longitudinal direction.

In certain embodiments, the reduction in fluidic resistance of the second microfluidic channel relative to the fluidic resistance of the first microfluidic channel is a function of the increasing cross-sectional area of the second microfluidic channel along the longitudinal direction of the particle sorting region.The width of the first microfluidic channel can be substantially constant along the longitudinal direction.

In certain embodiments, the cross-section of the gaps between the islands in the first array increases along the longitudinal direction, with the cross-section of each gap in the first array being defined along a plane transverse to the fluid flowing through the gap. The increase in the fluidic resistance of the third microfluidic channel relative to the fluidic resistance of the first microfluidic channel can be a function of the decreasing cross-sectional area of the third microfluidic channel along the longitudinal direction of the particle sorting region. The width of the first microfluidic channel can be substantially constant along the longitudinal direction.

In certain embodiments, the cross-section of the gaps between the islands in the second array increases in the longitudinal direction, and the cross-section of each gap in the second array is defined along a plane transverse to the fluid flowing through the gap.

In certain embodiments, the microfluidic device further comprises: a first inlet channel; and a second inlet channel, wherein each of the first inlet channel and the second inlet channel is fluidly connected to the particle sorting region.

In certain embodiments, along a longitudinal cross-section, a dimension of each opening in a first array is greater than a dimension of a previous opening in the array.

In certain embodiments, the particle and fluid movement area further includes a third microfluidic channel extending along the first microfluidic channel, and a second array of islands separating the first microfluidic channel and the third microfluidic channel, wherein the first microfluidic channel is located between the second and third microfluidic channels. The varying relative fluid resistance can be a function of the cross-sectional area that the second microfluidic channel or the third microfluidic channel increases in the longitudinal direction of the particle sorting area. For example, the width of one of the second microfluidic channel or the third microfluidic channel can increase in the longitudinal direction. Alternatively, the width of the other of the second microfluidic channel or the third microfluidic channel can decrease in the longitudinal direction. Sometimes, the width of the first microfluidic channel can be substantially constant in the longitudinal direction. The varying relative fluid resistance can be a function of the cross-sectional area that the second microfluidic channel and the third microfluidic channel increase in the longitudinal direction of the particle sorting area.

In certain embodiments, the device further comprises a first inlet channel and a second inlet channel, wherein each of the first inlet channel and the second inlet channel is fluidly connected to the particle sorting region.

In another aspect, the subject matter of the present disclosure can be embodied as a method for sorting particles in a fluid sample, wherein the method includes passing a first fluid sample comprising a set of first particles into a particle sorting region of a microfluidic device, wherein the particle sorting region includes a first microfluidic channel, a second microfluidic channel extending along the first microfluidic channel, and a first array of islands separating the first microfluidic channel from the second microfluidic channel. The method can also include passing a second fluid sample into the particle sorting region, wherein the fluid resistance between the first microfluidic channel and the second microfluidic channel varies along a longitudinal cross-section of the particle sorting region such that a portion of the first fluid sample passes from the first microfluidic channel into the second microfluidic channel through openings between the islands in the first array, and wherein the first microfluidic channel, the second microfluidic channel, and the first array of islands are further arranged to generate an inertial lift force that substantially prevents the set of first particles from propagating with the portion of the fluid siphoned through the openings in the first array.

In another aspect, the disclosed subject matter can be embodied as a method for displacing particles between fluid samples in a microfluidic device, the method comprising flowing a first fluid sample comprising a plurality of first particles into a particle sorting region of the microfluidic device, wherein the particle sorting region comprises a first microfluidic channel, a second microfluidic channel extending along a first side of the first microfluidic channel, a first array of islands separating the first microfluidic channel from the second microfluidic channel, a third microfluidic channel extending along a second side of the first microfluidic channel, and a second array of islands separating the first microfluidic channel from the third microfluidic channel, the second side being opposite to the first side of the first microfluidic channel, each island in the first array being separated from an adjacent island in the first array by an opening fluidically connecting the first microfluidic channel to the second microfluidic channel, and each island in the second array being fluidically connected to an adjacent island in the second array. The first microfluidic channel is separated from the opening of the third microfluidic channel; the second fluid sample is then allowed to flow into the particle sorting area, wherein the fluid resistance of the second microfluidic channel varies along the longitudinal direction of the particle sorting area relative to the fluid resistance of the first microfluidic channel, so that a portion of the first fluid sample enters the second microfluidic channel from the first microfluidic channel through the openings between the islands in the first array, the fluid resistance of the third microfluidic channel varies along the longitudinal direction of the particle sorting area relative to the fluid resistance of the first microfluidic channel, so that a portion of the second fluid sample enters the first microfluidic channel from the third microfluidic channel through the openings between the islands in the second array, and the first array of the first microfluidic channel, the second microfluidic channel and the island are further arranged to generate an inertial lift force, which basically prevents the plurality of first particles from propagating along with the portion of the first fluid sample passing through the opening of the first array.

Embodiments of these methods may have one or more of the following features. For example, in some embodiments, a first fluid sample and a second fluid sample are both delivered to a first microfluidic channel. The second fluid sample can flow continuously through the first microfluidic channel without being siphoned into the second microfluidic channel through an opening. Inertial lift can cause the group of first particles to displace across the fluid flow line so that the group of first particles is transferred to the second fluid sample.

In certain embodiments, a first fluid sample is delivered to a first microfluidic channel and a second fluid sample is delivered to a third microfluidic channel.

In certain embodiments, inertial lift forces displace the plurality of first particles across the fluid streamlines such that the plurality of first particles are transported from the first fluid sample to the second fluid sample within the first microfluidic channel.

In certain embodiments, the first fluid sample includes a plurality of second particles, wherein the plurality of second particles propagate with the portion of the first fluid sample that enters the second microfluidic channel. The first particles can be larger than the second particles.

In certain embodiments, the first fluid sample comprises a different fluid than the second fluid sample.

In certain embodiments, the amount of the first fluid sample entering the second microfluidic channel from the first microfluidic channel is substantially equal to the amount of the second fluid sample entering the first microfluidic channel from the third microfluidic channel, thereby maintaining a substantially constant concentration of the first particles within the first microfluidic channel.

In certain embodiments, the change in the fluidic resistance of the second microfluidic channel relative to the fluidic resistance of the first microfluidic channel comprises an increase in the cross-sectional area of the second microfluidic channel along the longitudinal direction.

In certain embodiments, the change in the fluidic resistance of the third microfluidic channel relative to the fluidic resistance of the microfluidic channel comprises an increase in the cross-sectional area of the third microfluidic channel along the longitudinal direction.

In certain embodiments, the first fluid sample includes a plurality of second particles, and the plurality of second particles propagate with the portion of the first fluid sample that enters the second microfluidic channel. The first particles can be larger than the second particles.

The first particles may have an average diameter between about 1 μm and about 100 μm.

On the other hand, the subject matter of the present disclosure can be embodied as a method for sorting particles in a fluid sample, wherein the method includes allowing a fluid sample comprising a group of first particles and a group of second particles to flow into a particle sorting region of a microfluidic device, wherein the particle sorting region includes a microfluidic device, a second microfluidic device extending along the first microfluidic device, and a first array of islands separating the first microfluidic device from the second microfluidic device, wherein the fluid resistance of the first microfluidic device varies along a section of the particle sorting region relative to the fluid resistance of the second microfluidic device, so that a first portion of the fluid sample is siphoned from a first microfluidic channel into a second microfluidic channel through openings between the islands in the first array, and wherein the first microfluidic channel, the second microfluidic channel, and the first array of islands are further arranged to generate an inertial lift force that substantially prevents the group of first particles from propagating with the portion of the fluid siphoned through the openings of the first array, while allowing the group of second particles to propagate with the portion of the fluid siphoned into the second microfluidic channel.

Embodiments of these methods can have one or more of the following features. For example, in some embodiments, inertial lift causes the group of first particles to move across the fluid streamline so that the group of first particles continues to propagate with the second portion of the fluid sample remaining in the first microfluidic channel. In some embodiments, the first particles can be larger than the second particles. In some embodiments, the first portion of the fluid sample passes through openings in the first array of islands in response to a change in fluid resistance between the first microfluidic channel and the second microfluidic channel. The change in fluid resistance can include a change in the cross-sectional area of one of the first microfluidic channel or the second microfluidic channel along the direction of fluid flow. The change in fluid resistance can include a change in the size of the openings between the islands in the array.

The embodiments of the subject matter described herein have several advantages. For example, in certain embodiments, the subject matter described herein can be used to separate particles in a fluid and/or aggregate particles in a fluid. In certain embodiments, the subject matter described herein can be used to filter particles from a fluid and/or move particles from one fluid to another. High volume capacity and handling capacity, significant and adjustable fluid volume reduction, and high particle yield can be achieved with inexpensive and simple equipment using the technology described herein. In certain embodiments, the currently described technology can also provide improved processing and simple integration with other microfluidic modules. For clinical applications, the system described herein can be configured to be self-contained and disposable. In contrast, for bioprocessing/industrial applications, the device can be configured to be continuous flow/processing.

For purposes of this disclosure, a channel refers to a structure through which a fluid can flow.

For purposes of this disclosure, microfluidics refers to a fluidic system, device, channel, or chamber that typically has at least one cross-sectional dimension between about 10 nm and about 10 mm.

For the purposes of this disclosure, a gap refers to an area in which a fluid or particle can flow. For example, a gap can be a space between two obstacles in which a fluid flows.

For the purposes of this disclosure, a rigid island structure refers to a physical structure that particles cannot generally penetrate.

For the purposes of this disclosure, fluidic resistance refers to the ratio of the pressure drop across a channel (eg, a microfluidic channel) to the fluid flow rate through the channel.

The particles within the sample can have any size that allows the sample to be transported within the microfluidic channel. For example, the particles can have an average hydraulic size between 1 μm and 100 μm. The particle size is limited only by the channel geometry; therefore, particles larger or smaller than the above particles can be used. The size of particles (e.g., cells, eggs, bacteria, fungi, viruses, algae, any prokaryotic or eukaryotic cell, organelles, exosomes, droplets, bubbles, contaminants, sediments, organic and inorganic particles, magnetic beads, and/or magnetically labeled analytes) can be determined using standard techniques known in the art, such as average hydraulic particle size or average diameter.

Unless otherwise defined, all technical and scientific terms used herein have the same meaning as commonly understood by those of ordinary skill in the art to which the invention belongs. Although methods, materials, and apparatus similar or equivalent to those described herein can be used in the practice or testing of the present invention, suitable methods, materials, and apparatus are described below. All publications, patent applications, patents, and other references mentioned herein are incorporated herein by reference in their entirety. In the event of a conflict, the present specification, including definitions, will prevail. In addition, the materials, methods, and examples are illustrative only and are not intended to be limiting.

The details of one or more embodiments are set forth in the accompanying drawings and the description below. Other features and objects will be apparent from the description, drawings, and claims.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG1 is a schematic diagram illustrating a top view of one example of a microfluidic device capable of shifting the position of particles within and across fluid flow lines.

2 is a schematic diagram illustrating a top view of an example of a particle and fluid moving device in which a fluid is allowed to travel across a propagation channel of a particle.

3 is a schematic diagram illustrating a top view of one example of a device having two inlet microfluidic channels coupled to a confluence channel, which in turn is coupled to a particle movement region.

4 is a schematic diagram illustrating a top view of one example of an apparatus having a particle movement zone for sorting particles based on size.

5 is a schematic diagram illustrating a top view of one example of an apparatus having a particle movement zone for sorting particles based on size.

FIG6A is a schematic diagram illustrating a top view of one example of a microfluidic system including a particle sorting region.

FIG6B is an enlarged view of the particle sorting region of FIG6A .

7 and 8 are schematic diagrams showing top views of examples of particle sorting regions.

FIG. 9 is a series of graphs (FIGS. 9A-9D) illustrating the volume reduction characteristics of a device that is staged according to inertial lift.

FIG. 10 is a series of graphs (FIGS. 10A-10B) of white blood cell (WBC) yield versus fluid flow rate.

FIG. 11 is a graph of the yield of fluorescent beads of different sizes across different flow rates.

FIG. 12 is a graph of white blood cell yield versus fluid movement.

13 is a photograph of a microscope slide housing a multiplexed array of fluid movement devices.

DETAILED DESCRIPTION

The interaction between particles (e.g., cells, such as blood cells in general, as well as fetal blood cells in maternal blood, bone marrow cells, and circulating tumor cells (CTCs), sperm, eggs, bacteria, fungi, viruses, algae, any prokaryotic or eukaryotic cell, cell population, organelle, exosome, droplet, gas bubble, contaminant, sediment, organic and inorganic particles, beads, bead-labeled analytes, magnetic beads and/or magnetically labeled analytes), the fluid in which the particles travel (e.g., blood, aqueous solution, oil, or gas), and rigid structures can be used to move the particles across fluid flow lines in a microfluidic device in a controlled manner. In particular, the forces experienced by particles traveling in a microfluidic device can be used to precisely position the particles, thereby enabling a variety of useful microfluidic operations. Examples of microfluidic operations that can be performed using these forces include, but are not limited to, concentrating particles in a carrier fluid, moving particles from one carrier fluid to another, separating particles in a fluid based on particle size (e.g., average diameter), converging particles in a carrier fluid onto a single streamline (or multiple distinct streamlines), precisely positioning particles at any location within a microchannel, and mixing (or dispersing) particles. Furthermore, any of the above operations can be performed simultaneously with other techniques (e.g., magnetic separation) to enhance the effectiveness of the operation.

Several different mechanisms can be used to generate the forces that can move particles across fluid streamlines. The first force is called the "impact force" (also known as the decisive lateral displacement (DLD)). The impact force is a direct interaction between the rigid walls of a structure and the particle due to the size of the particle relative to the walls. Since the center of a particle with radius _rp cannot be closer than _rp to an adjacent structure, if the center of the particle lies on a streamline that is closer than _rp to the structure, the particle will be dislodged by the structure to a distance of at least _rp . This impact can cause the particle to move across the fluid streamline.

Another force is called inertial lift (also called wall force or wall-induced inertia). In contrast to collision force, inertial lift is a fluid force on the particle that is not caused by contact with a rigid structure. Although not fully understood, inertial lift is a repulsive force caused by the flow disturbance generated by the particle when it approaches the wall. Particles flowing close to the microchannel wall are subject to inertial lift forces normal to the wall. At high flow rates, inertial lift forces are strong and can cause particles to move across streamlines. Furthermore, because this force is highly size-dependent (larger particles are subject to greater forces), it can be exploited to group particles according to size. More details on inertial flows can be found in D. Di Carlo, D. Irimia, R. G. Tompkins, and M. Toner, “Continuous inertial focusing, ordering, and separation of particles in microchannels,” Proceedings of the National Academy of Sciences of the United States of America, vol. 104, no. 48, pp. 18892–18897, November 2007; D. Di Carlo, J. F. Edd, K. J. Humphry, H. A. Stone, and M. Toner, “Particle segregation and dynamics inconfined flows,” Physical Review Letters, vol. 102, no. 9, p. 094503, March 2009; and D. Di Carlo, “Inertial microfluidics,” Lab on a Chip, vol. 9, no. 21, p. 3038, 2009, each of which is incorporated herein in its entirety.

Another force is the result of differential pressure drag from Dean flow. Microfluidic channels with curvature can create additional drag on particles. When curvature is introduced into rectangular channels, secondary flows (i.e., Dean flow) can be generated perpendicular to the flow direction due to the uneven inertia of the fluid. Therefore, faster-moving fluid elements in the center of a curved channel can generate greater inertia than fluid elements near the channel edges. Due to high Dean flow, the drag on suspended particles in the fluid can become significant.

Another force occurs in high Stokes number flows. The Stokes number (Stk) describes how fast the particle trajectory changes with the change of fluid trajectory. If Stk is greater than 1, there is a hysteresis between the change of fluid trajectory and the change of particle trajectory. Under high Stokes flow conditions (for example, Stokes number is greater than about 0.01), changing the direction of fluid flow can be used to force particles to cross streamlines. More details about Dean flow and high Stokes number can be found in, for example, U.S. Patent No. 8,186,913, which is incorporated herein by reference in its entirety. In high Stokes flow applications and Dean flow applications, fluid displacement causes particles to cross fluid streamlines.

Other techniques for moving particles include viscoelasticity and inertial elastic aggregation. Details on these methods can be found in "Sheathless elasto-inertial particle focusing and continuous separation in astraight rectangular microchannel" (Yang et al., Lab Chip (11), 266-273, 2011), "Single line particle focusing induced by viscoelasticity of the suspending liquid: theory, experiments and simulations to design a micropipe flow=focuser" (D'Avino et al., Lab Chip (12), 1638-1645, 2012), and "Inertio-elastic focusing of bioparticles in microchannels at high throughput" (Lim et al., Nature Communications, 5(5120), 1-9, 2014), each of which is incorporated herein by reference in its entirety.

The above techniques are "internal" because they use the structure of the fluid flow and/or the microfluidic channel itself to generate the force necessary to move the particles across the streamline. Sometimes, other external mechanisms can also be used to change the route of the particles in the fluid together with one or more internal forces. For example, sometimes externally applied magnetic forces, gravity/centrifugal forces, electric field forces or acoustic field forces can be used to cause the particle position to move across the fluid streamline. Detailed information about how to apply this force can be found in, for example, WO2014/004577, "Sorting particles using high gradient magnetic fields", U.S. Patent No. 7,837,040, "Acoustic focusing", WO2004/074814, "Dielectrophoretic focusing", and "Microfluidic, Label-Free Enrichment of Prostate Cancer Cells in Blood Based on Acoustophoresis" (Augustsson et al., Analytical Chemistry 84 (18), September 18, 2012).

The present disclosure focuses on combining inertial lift with periodic fluid extraction to move particles across fluid streamlines for use in sorting particles between fluids and/or separating particles from fluids. In particular, a first fluid containing particles is introduced into a first microfluidic channel having an array of rigid island structures that separate the first channel from adjacent microfluidic channels. When the first fluid is extracted from the first microfluidic channel into the second microfluidic channel through the gaps between the island structures, the particles are drawn closer to the island structures. When the particles are closer to the island structures, they are subjected to inertial lift that deviates from the direction of fluid extraction, causing the particles to cross the fluid streamlines and remain within the first channel. At the same time, a second fluid enters the first microfluidic channel from a third microfluidic channel through the second island structure array, causing the particles in the first channel to be transferred to the second fluid in the first channel. In certain embodiments, the amount of the first fluid entering the second channel from the first channel is the same as the amount of the second fluid entering the first channel from the third channel. Therefore, particles can move from one fluid to another while maintaining the same particle concentration. The combination of fluid extraction and inertial lift can be used to perform particle positioning, particle filtering, particle mixing, fluid mixing, and/or moving fluid across a particle stream, among other operations.

However, it should be noted that the techniques described herein for particle and/or fluid sorting are not limited to the use of inertial lift forces. Rather, periodic fluid extraction can also be combined with one or more of the aforementioned (internal and external) forces to control the position of particles in a fluid propagating in a microfluidic device.

The mechanism disclosed herein for making particles move can also be based on size, and therefore can be used to perform size-based particle operations (e.g., based on the average diameter of the particles). By allowing the particles to move repeatedly across streamlines, both the fluid and the particles in the microfluidic device can be manipulated to perform the following operations, such as aggregating the particles to one or more fluid streamlines, filtering the particles from the fluid, mixing different particles from different fluid streams, and/or sorting the particles according to size. Typically, "aggregating" particles refers to repositioning the particles in a width less than the channel width across the lateral extent of the channel. For example, the technology disclosed herein can position particles suspended in a fluid into a fluid stream, wherein the ratio of channel width to fluid stream width is about 1.05, 2, 4, 6, 8, 10, 20, 30, 40, 50, 60, 70, 80, 90, or 100. The particles can have various sizes, including but not limited to an average diameter between about 1 μm and about 100 μm.

The mechanisms disclosed herein for moving particles can also be used to manipulate and sort particles differentially based on their shape (e.g., spherical vs. cylindrical) and deformability (e.g., rigid vs. flexible).

Particle movement/sorting using inertial lift

Before discussing how to move particles from one fluid to another using one or more of the devices disclosed herein, it is helpful to first review fluid extraction and inertial forces in the context of a more basic device structure, such as the device shown in FIG1 . FIG1 is a schematic diagram illustrating a top view of an example of a microfluidic device 100 that is capable of moving/sorting the positions of particles 102 across fluid streamlines as the fluid propagates through the microfluidic device 100. As described below, the movement of particles across fluid streamlines relies on the inertial lift forces experienced by the particles as fluid is periodically extracted from the microfluidic channel. A Cartesian coordinate system is shown for reference, with the X direction extending into and out of the page.

During operation of the device 100, a fluid carrying particles 102 is introduced through the inlet microfluidic channel 104. In this and other embodiments of the particle movement device, the fluid can be introduced by using a pump or other fluid-driven mechanism. The inlet channel 104 is divided into a particle sorting area having two different fluid flow channels (a second microfluidic channel 106 and a first microfluidic channel 108 substantially parallel to the second microfluidic channel 106), and the two channels are separated by a one-dimensional array of rigid island structures 110. The one-dimensional array of island structures 110 extends substantially in the same direction as the direction of fluid flow through the second and first microfluidic channels. Each island structure 110 in the array is separated from the adjacent islands 110 by an opening or gap 114 through which fluid can flow. In the example of Figure 1, each gap 114 between adjacent islands 110 has the same distance. In other embodiments, different gaps between adjacent islands 110 can have different distances. For example, in some embodiments, the length of each subsequent opening in the first array (e.g., measured along the direction of fluid propagation, i.e., the Z direction in Figure 1) is greater than the size of the previous opening in the array. Alternatively, in some embodiments, the distances can alternate between larger and smaller for subsequent openings. Furthermore, while a one-dimensional array is shown in FIG1 , islands 110 can be arranged in various configurations, including, for example, a two-dimensional array of islands. The boundaries of the fluid flow region within the microfluidic channel are defined by device walls 112 and the walls of islands 110.

As the fluid propagates generally along the z-direction from the inlet channel 104 toward the channels (106, 108), the particles 102 are subjected to forces (inertial lift forces in this example) that cause the particles 102 to move across the fluid streamlines and along the first microfluidic channel 108. These inertial lift forces are in the negative y-direction (see the short arrows near each particle 102 in FIG1 ).

For example, when the particle 102 is within the inlet channel 104 and approaches the top wall 112, the inertial lift force applied to the particle pushes the particle downward toward the first microfluidic channel 108. Once within the first microfluidic channel 108, the particle 102 may approach the wall of the first island 110, causing it to again be subjected to the inertial lift force that pushes the particle 102 downward, which retains the particle within the first microfluidic channel 108. Thus, within each "particle movement" region shown in FIG1 , the inertial lift force repeatedly applied to the particle 102 serves to separate/filter the particle from the fluid propagating through the second microfluidic channel 106.

At the same time, a portion of the fluid traveling in the first microfluidic channel 108 is drawn into or flows into the second microfluidic channel at one or more "fluid movement" or "fluid extraction" regions in the device 100 (see FIG. 1 ). In the example of FIG. 1 , each fluid movement region corresponds to an opening or gap extending between the first microfluidic channel 108 and the second microfluidic channel 106. Each "fluid movement" region primarily allows fluid to be extracted from the first microfluidic channel 108 into the second microfluidic channel 106. Because particles follow fluid streamlines, the movement of fluid into the gap tends to pull particles 102 toward the gap as well. However, as particles move closer to the gap 114, they approach the island structure 112, which exerts an inertial lift force that causes the incoming particles to cross the fluid streamline in a direction away from the gap 114. In other words, particles 102 move from the fluid streamline entering the second microfluidic channel 106 to the fluid streamline continuing to flow within the first microfluidic channel 108. Consequently, particles 102 continue to propagate within the first microfluidic channel 108 rather than moving with the fluid into the second microfluidic channel 106. If no fluid were to move from the first microfluidic channel 108 to the second microfluidic channel 106, the particles would migrate due to inertial aggregation. However, by moving the fluid across the channel, the particles 102 tend to follow the fluid flow toward areas where the inertial lift force is much stronger than the shear gradient force, allowing the particles to move across the streamline in a very efficient and controlled manner.

In the present embodiment, fluid is extracted through the fluid movement area due to the decrease in fluid resistance. That is to say, for fluids of constant viscosity, the gap 114 between adjacent islands 110 increases the channel area through which the fluid can flow, resulting in a decrease in fluid resistance. When the fluid propagates through the channel 108 of the device 100 and reaches the gap 114, a portion of the fluid will flow into the gap 114 and subsequently flow into the second microfluidic channel 106 (i.e., the portion of the fluid is extracted into the channel 106). Increasing the channel width of the second microfluidic channel 106 will also reduce fluid resistance. In some embodiments, the width of the channel 106 can be understood as the distance between a position on the second channel wall 112 and a position on the surface of the island 110 that is directly opposite to the above-mentioned position on the second channel wall 112 and facing the above-mentioned position on the second channel wall 112. For the gap area between the islands 110, the width of the channel 106 can be understood in some embodiments as the distance between a position on the second channel wall 112 and a position on a virtual surface that passes through the gap between adjacent islands 110 and is located directly opposite and facing the above-mentioned position on the channel wall 112, where the virtual surface is collinear with the side of the adjacent island closest to and facing the wall 112.

In a particular example, such as shown in FIG1 , the second microfluidic channel wall 112 is tilted away from the island at an angle (oriented relative to the island) so that the width of the second microfluidic channel 106 increases along the longitudinal direction of the channel (e.g., along the direction of fluid propagation or the positive Z direction in the example shown in FIG1 ), thereby resulting in a decrease in fluid resistance. Any increase in the cross-sectional area of the channel 106 along the longitudinal direction of the first microfluidic channel, rather than just an increase in width, can also be used to reduce fluid resistance. Alternatively or additionally, the fluid in the channel 108 can be subjected to an increase in fluid resistance relative to the fluid resistance of the channel 106 (e.g., by reducing the cross-sectional area of the channel 108 along the longitudinal direction). Thus, it can be said that fluid is drawn in response to a change in the relative fluid resistance between the second and first microfluidic channels, which change causes fluid to be drawn from the first channel 108 into the second channel 106. The change in relative fluid resistance can occur over the entire particle sorting region or a portion of the sorting region that is less than the entire particle sorting region. The change in relative fluid resistance can occur along the direction of fluid flowing through the particle sorting region (e.g., along the longitudinal direction of the particle sorting region as shown in FIG1 ).

As the fluid resistance at the indentation 114 and/or in the channel 106 gradually decreases, a larger amount of fluid flows into the second microfluidic channel 106. In addition, the repeated movement of fluid into the second channel 106 reduces the amount of fluid in the first channel 108. For the configuration shown in FIG1 , the repeated fluid withdrawal increases the concentration of particles in the first channel 108 while decreasing the concentration of particles in the second microfluidic channel 106, thereby "filtering" or "purifying" the fluid in the second microfluidic channel 106.

The resulting concentrated particle streamlines can be incorporated into a separate processing area of the microfluidic device 100 or removed from the device 100 for additional processing and/or analysis. Similarly, the "filtered" fluid in the second channel 106 can be incorporated into a separate area of the microfluidic device 100 or removed from the device 100 for additional processing and/or analysis. In certain embodiments, particles 102 entering the device 100 are "pre-aggregated" into a desired fluid streamline position aligned with the first microfluidic channel 108. By pre-aggregating the particles 102 into the desired position, the probability of the particles inadvertently entering the second microfluidic channel 106 can be reduced.

Another advantage of the particle movement technology described herein is that it can be used for gathering particles along one or more streamlines. For example, as mentioned above, partial fluid can be extracted into one or more parallel microfluidic channels from initial microfluidic channel. In some cases, the parallel microfluidic channels comprising the fluid extracted can be recombined with initial microfluidic channel in downstream subsequently, so that particle is confined to the specified streamline in single channel. The advantage of this technology combining fluid movement with inertial lift is that how much fluid can be removed from each side of initial channel by control, particles can be gathered into the desired position (for example, near channel wall, in channel center, or midway and other position between channel wall and channel center) in downstream channel, thereby for the design and use increase flexibility of microfluidic device. By contrast, for the microfluidic system mainly based on inertial gathering, the position of gathering flow can only be limitedly selected in channel.

Having reviewed fluid extraction and inertial forces, a microfluidic device for moving/sorting particles between fluids can now be described. In particular, the fluid and particle movement techniques described herein can be used to move fluids across channels without associated particle movement, so that particles are indirectly transported to another fluid. FIG2 is a schematic diagram illustrating an example of a device 200 for moving particles between fluids. The device 200 includes a first inlet microfluidic channel 204 and a second inlet microfluidic channel, both of which are separated by a partitioning structure 205, such as a wall or other object, that prevents mixing between the first inlet 204 and the second inlet 206. At the ends of the partitioning structure 205, the first and second inlet microfluidic channels (204, 206) are fluidically coupled to a particle movement region having three different fluid flow regions (a second microfluidic channel 208, a first microfluidic channel 210, and a third microfluidic channel 212).

The second microfluidic channel 208 is separated from the first microfluidic channel 210 by island structures 218 in the first array 214. The third microfluidic channel 212 is separated from the second microfluidic channel 210 by islands 218 in the second array 216. Each adjacent island structure in the first array 214 and each adjacent island structure in the second array 216 are separated by a gap for fluid movement. The boundaries of the microfluidic channels are defined by device walls 220 and the walls of the islands. The microfluidic channel walls 220 of the second channel 208 are tilted at an angle away from the islands 218 (oriented obliquely relative to the islands), thereby increasing the width of the second channel along the direction of fluid propagation, thereby reducing fluid resistance and causing fluid to be drawn from the first channel 210 into the second channel 208. In contrast, the walls 220 of the third channel 212 are tilted at an angle toward the islands 218, thereby decreasing the width of the third channel 212 along the direction of fluid propagation, thereby increasing fluid resistance and causing fluid to be drawn from the third channel 212 into the first channel 210.

During operation of the device 200, particles 202 flowing within the first fluid in the first inlet channel 204 interact with the channel walls, causing them to experience inertial lift forces, which cause the particles 202 to move or aggregate toward the central streamline of the liquid flow. The fluid path of the central microfluidic channel 210 is substantially aligned with the fluid path of the first inlet channel 204, so the aggregated particles and a portion of the first fluid from the channel 204 flow into the first channel 210. Once the particles 202 enter the first microfluidic channel 210, they experience inertial lift forces from the island structures 218, which continue to cause the particles 202 to aggregate along one or more central streamlines extending through the channel 210. Simultaneously, due to reduced fluid resistance, some of the first fluid is drawn into the second microfluidic channel 208 in the "fluid movement" region. Because the particles 202 experience inertial lift forces in the first channel 210, most of the particles 202 remain in the first channel 210 rather than being carried into the second channel 208 by the first fluid.

A second fluid is provided in the second inlet channel 206. The second fluid can be the same fluid as the carrier fluid used to introduce particles from the inlet channel 204, or a different fluid. The second fluid primarily flows from the second inlet 206 into the third microfluidic channel 212. A portion of the second fluid is drawn from the inlet 206 and/or the third channel 212 into the first microfluidic channel 210 during the "fluid movement" region. As the second fluid propagates downward along the channel 212, the withdrawal of the second fluid occurs due to an increase in fluid resistance experienced by the second fluid (e.g., a decrease in channel width). Consequently, a larger amount of the second fluid begins to flow within the first channel 210. In certain embodiments, such as those with sufficiently long microfluidic channels, the second fluid can even cross from the third microfluidic channel 212 into the first microfluidic channel 210 and ultimately enter the second microfluidic channel 208 to merge with the first fluid. However, because the particles 202 experience inertial lift in the first channel 210, the majority of the particles 202 remain in the first channel 210 and are not carried into the second channel 208 by the first or second fluids. Thus, the combination of the fluid movement region and the particle movement region allows particles 202 to be separated from the first fluid. That is, when the second fluid is introduced into the channel 210, the particles 202 are moved from the first fluid to the second fluid. Because inertial forces hold the particles 202 within the channel 210, in certain embodiments, the particles can also be separated from the combined first and second fluids passing through the channel 208.

If the amount of first fluid drawn from first channel 210 into second channel 208 remains equal or substantially equal to the amount of second fluid introduced into the first channel from the third channel over the length of the device, the amount of fluid propagating through channel 210 can be kept substantially constant. Similarly, because inertial lift forces keep the number of particles propagating through channel 210 substantially constant, the overall concentration of particles does not change significantly even if the fluid within channel 210 changes. That is, the concentration of particles 202 in the first fluid at the beginning of channel 210 is substantially equal to the concentration of particles 202 in the second fluid near the end of channel 210 after the fluid movement has been completed.

In some embodiments, the first fluid is not completely drawn from channel 210 into channel 208. Instead, the microfluidic device can be configured such that, after operating the device, both the first fluid and the second fluid merge within channel 210. In this case, the first and second fluids can propagate side by side according to laminar flow, or can mix due to, for example, diffusion. In either case, if the amount of the first fluid drawn into channel 208 is substantially equal to the amount of the second fluid introduced into channel 210 from channel 212 over the length of the device, the concentration of particles 202 relative to any fluid within channel 210 can remain substantially constant.

Any of the fluid flows from the second, first, or third channels can be coupled to a separate region of the microfluidic device or removed from the device for additional processing or analysis. In certain embodiments, the variation in size/fluidic resistance of the second and third channels can be configured to ensure that an equal amount of fluid flows into the third channel and out of the second channel at each cell.

Another example of a device capable of transferring particles between fluids is shown in FIG3 , which is a schematic diagram showing an example of a device 300 comprising two inlet microfluidic channels (304, 306) that are coupled to a single microfluidic channel 305 to merge the fluids. The confluent channel 305 is further coupled to a particle movement region that includes two different flow regions (a second microfluidic channel 308 and a first microfluidic channel 310). The second microfluidic channel 308 is separated from the first microfluidic channel 310 by an array of island structures 312, wherein each island 312 is separated from an adjacent island 312 by a notch 314 for fluid movement. In addition, the top wall 316 of the second microfluidic channel 308 is tilted away from the island 312 at a certain angle to reduce fluid resistance in the Z direction.

During operation of the device 300, a first fluid ("fluid 1") containing particles 302 is introduced into the first inlet channel 304, while a second fluid ("fluid 2") without particles is introduced into the second inlet channel 306. Assuming that the fluids are introduced at a flow rate corresponding to a low Reynolds number (and therefore laminar flow), the two different fluids hardly mix in the confluence region 305, i.e., the two fluids continue to flow essentially in layers close to each other. The fluid path in the confluence region 305 is aligned with the fluid path of the first microfluidic channel 310, so that the merged fluids flow primarily into the first channel 310. When the two fluids enter the first microfluidic channel 310, the particles 302 in the first fluid are subjected to an inertial lift force from the island structure 312, which is transverse to the flow direction and holds the particles 302 within the first microfluidic channel.

At the same time, the increased width of second microfluidic channel 308 (due to the sloped channel walls 316) reduces fluidic resistance, so that a portion of the first fluid (the first fluid closest to the island structure) is drawn into second channel 308 at each gap between islands 312. Because the first fluid flows as a layer above the second fluid, little to no second fluid is drawn into second channel 308. After propagating a sufficient distance past islands 312, most of the first fluid is drawn into second channel 308, while particles 302 and most or all of the second fluid remain in first channel 310. Therefore, the microfluidic device configuration shown in FIG3 can also be used to transfer particles from one fluid to a different second fluid. If the amount of the first fluid flowing through inlet 304 is substantially equal to the amount of the second fluid flowing through inlet 306, then the concentration of particles 302 in the second fluid within channel 310 (and after the first fluid is drawn) can remain substantially equal to the concentration of particles 302 in the first fluid within inlet 304. In some embodiments, the propagation distance is sufficiently long that the second fluid is also drawn into second microfluidic channel 308. In this case, the concentration of particles 302 in the second fluid within the first microfluidic channel 310 may increase to a level higher than the concentration of particles within the channel 304 .

In certain embodiments, repeated particle and fluid movement can be used to perform size-based intra-fluid particle separation. FIG4 is a schematic diagram showing an example of a device 400 for size-based particle sorting. The device is constructed identically to the device 300 shown in FIG3 . During operation of the device 400, a first fluid (“fluid 1”) containing particles of different sizes (large particles 402 and small particles 403) is introduced into the first inlet channel 304, while a second fluid (“fluid 2”) without particles is introduced into the second inlet channel 406. The first and second fluids can be fluids of the same or different species. Again, assuming that the fluids are introduced at a flow rate corresponding to a low Reynolds number (and therefore laminar flow), the two different fluids hardly mix in the confluence region 405, i.e., the two fluids continue to flow substantially in layers close to each other. When the two fluids enter the first microfluidic channel 410, the force on the large particle 402 is large enough to keep the particle 402 within the first microfluidic channel 410. In contrast, the force on the small particle 403 is not strong enough to prevent the small particle 403 from being drawn into the second microfluidic channel 408 along with the first fluid. After repeated particle movement and fluid extraction over a sufficiently long distance, most of the first fluid and small particles 403 are drawn into the second channel 408, while large particles 402 and most of the second fluid remain in the first channel 410. This process is also called fractionation and can be used to separate particles and fluids based on size.

There are a variety of reasons why large particles 402 are preferentially left relative to small particles 403. The first, inertial lift is highly nonlinear on particle diameter. For example, it is believed that the scale of the inertial lift near the channel wall is in the range of a ³ to a ⁶ , where a is the particle diameter, so large particles are subjected to much larger force than small particles. Larger inertial lift can be used to remove particles from the fluid flow that moves upward from one or a row of island-like structure arrays to the next one near the island. Other information about the relationship between particle size and inertial lift can be found in " Particle Segregation and Dynamics in Confined Flows " (Physical Review, 2009) by Di Carlo et al., which is incorporated herein by reference in its entirety. The second, the equilibrium position of large particles is usually further away from the wall than small particles, so further away from the fluid extraction channel, and is more likely to fall on streamlines, not moving to the extraction channel. Therefore, large particles can be kept in a specified row, and the small particles flowing near the island move upward from a row of the array to the next row.

Thus, fractionation is accomplished by repeatedly (1) using inertial lift to move large particles away from the channel walls, and then (2) moving fluid without large particles to an adjacent channel. In certain embodiments, fractionation can also be used to sort particles from a source fluid (e.g., blood) into an adjacent target fluid (e.g., buffer) across a fluid flow line.

For example, FIG5 is a schematic diagram illustrating an example of a device 500, which can also be used to separate particles based on size. The construction of device 500 is identical to device 200 shown in FIG2 . The fluid resistance in third microfluidic channel 512 gradually increases as the channel width decreases, while the fluid resistance in second microfluidic channel 508 gradually decreases as the channel width increases. Therefore, during operation of device 500, repeated fluid movement of a first fluid ("fluid 1") from first microfluidic channel 510 to second microfluidic channel 508 occurs at the gaps between islands 518 in first array 514. Similarly, repeated fluid movement of a second fluid ("fluid 2") from third microfluidic channel 512 to first microfluidic channel 510 occurs at the gaps between islands 518 in second array 516. The fluid extraction force is sufficient to draw small particles 503 away with the first fluid, but insufficient to counteract the inertial lift force on large particles 502. Consequently, the large particles remain flowing along streamlines within first microfluidic channel 510. After repeated particle and fluid movement, large particles 502 begin to flow along streamlines in the second fluid that has moved into the first channel 510. If the amount of fluid flowing from channel 510 to channel 508 is substantially equal to the amount of fluid flowing from channel 512 into channel 510 over the length of the particle sorting/or movement region, then the amount of fluid flowing in channel 510 can be kept substantially constant.

The microfluidic device shown in Figures 1-5 uses inertial lift from the microfluidic channel walls and from a periodic matrix of island structures to implement particle movement across fluid streamlines. Techniques other than inertial lift can be used to help particles move across fluid streamlines. For example, internal forces generated by Gordian and/or High Stokes flows, such as inertial focusing, can be used to move particles across fluid streamlines and/or retain particles within the microfluidic channel. Alternatively or additionally, external forces such as magnetic forces, acoustic forces, gravity/centrifugal forces, optical forces, and/or electric forces can be used to move particles across fluid streamlines. In addition, the shape of the rigid island structures separating different flow regions is not limited to that shown in Figures 1-5. For example, the rigid island structures can have a shape similar to a pillar, a cube, or other polyhedron, where the top or bottom surfaces are or can be congruent polygons. In some cases, such as at high flow rates, it is advantageous to use islands with streamlined, tapered ends because this helps minimize the formation of flow recirculation (eddies), which can interfere with flow in unpredictable and undesirable ways. Other shapes of rigid island structures are also possible. The long axis of the rigid island structure can be oriented at a certain angle relative to the average flow direction of the fluid, the average flow direction of the particles, or the long axis of the sorting area. The shape of the channel section is not limited to the approximate rectangle shown in Figures 1-5. The channel section can include a bend or a considerable width change. Regarding the cross section, the channel described in Figures 1-5 can be square, rectangular, trapezoidal, or circular. Other shapes of channel cross sections are also possible. The channel depth can be uniform on the particle sorting area, or the channel depth can vary laterally or longitudinally. In addition, although the microfluidic channel shown in Figures 1-5 is a path that is approximately straight, the channel can be arranged into other different layouts. For example, in some embodiments, the microfluidic channel can be formed to have a spiral configuration. For example, the first microfluidic channel and the second microfluidic channel can be arranged into a spiral configuration, wherein the first and second microfluidic channels are still separated by the island structure array, but the longitudinal direction of the fluid flowing through the channel will follow a roughly spiral path. In some embodiments, the size or shape of the island structures can vary along the length of the sorting region (e.g., along the direction of fluid flow) and/or along the width of the sorting region (e.g., transverse to the direction of fluid flow). In some embodiments, the proportion of fluid passing between the island structures varies at different locations within the channel. For example, the percentage of fluid passing through the first gap between two island structures can be higher or lower than the percentage of fluid passing through the next adjacent gap between the two island structures.

While some embodiments shown in Figures 1-5 include two inlet channels, additional inlet channels can be incorporated into the microfluidic channels. In some embodiments, three, four, or more inlet channels can introduce fluid into the region of the device where particles are moved through fluid exchange and inertial lift. For example, in some embodiments, there can be three inlet channels, one for blood, one for dye, and one for buffer flow. Using the combination of fluid movement and inertial lift technology disclosed herein, white blood cells from the bloodstream can be moved into a reagent stream and then into a buffer stream.

In some embodiments, the devices described herein can be used with other microfluidic modules (for example, including filters for filtering out particle subpopulations of certain sizes) to manipulate fluids and/or particles. Additionally, the devices described herein can be used in series and/or in parallel within a microfluidic system.

Design parameters of microfluidic devices

The effects of various design parameters on the operation of the microfluidic device will now be described. For reference only, FIG7 is a schematic diagram showing a top view of a typical particle sorting region 700 comprising several rows of island structures 710, each row of islands being separated from adjacent rows of islands by corresponding internal microfluidic channels 703. In addition, there are external microfluidic channels 705a extending above the island array and external microfluidic channels 705b extending below the island array. The primary direction of fluid flow is indicated by arrow 701. The width of the external channel 705a (defined along the Y direction) expands along the length of the channel, while the width of the external channel 705b (defined along the Y direction) contracts along the length of the channel. For ease of discussion below, the channels and islands can be understood as being arranged into separate "units" (see unit 1, unit 2, and unit 3 in FIG7 ). Specifically, FIG7 shows three units of an array having two internal channels and two external channels.

The relevant design parameters for the particle sorting region 700 are cell length, width, and fluid movement. Here, width w refers to the size of the internal microfluidic channel 703, while length l refers to the length of the island structure 710 within the cell. The internal channel thus has a fixed width w and fluidic conductance g. The expansion channel has a width w _e,i and a fluidic conductance g _e,i , where i refers to the cell number. Similarly, the contraction channel has a width w _c,i and a fluidic conductance g _c,i , where i refers to the cell number. The fluid movement f is the fraction of the flow q in the internal channel that moves between the rows (channels) at each cell. The net flow in the internal channel remains unchanged because at each cell, flow fq (the product of f and q) moves out of these channels (at the openings 707 between the island structures) and flow fq moves into these channels (at the openings 707 between the island structures). In contrast, the net flow in the external channel 705 varies. The net flow in the converging channel 705b decreases by fq at each cell, while the net flow in the diverging channel 705a increases by fq at each cell. Thus, the width of the external channel of each cell can also be adjusted to provide the desired fluid movement across the array.

For each successive cell, the conductance of the expanding channel increases by fg (the product of f and g), while the conductance of the converging channel decreases by fg. The changing conductance translates into a proportional change in volumetric flow rate because the pressure drop p per cell is approximately equal across all channels in the cell, and the relationship between flow rate and conductance is q = pg. Therefore, when the conductance of an expanding channel (e.g., channel 705a) increases by fg, the flow rate in the expanding channel increases by fq. Similarly, when the conductance of a converging channel (e.g., channel 705b) decreases by fg, the flow rate in the converging channel increases by fq.

The conductance of each channel in the array is a function of its size and the viscosity of the fluid. In the array shown in Figure 7, each channel is assumed to have a rectangular cross-section and therefore has a conductance described by the following equation:

Here, η is the fluid viscosity, l is the channel length, w is the channel width, h is the channel height, and α = h/w. More precise formulas based on infinite series are also available (Tanyeri et al., "A microfluidic-based hydrodynamic trap: Design and implementation (Supplementary Materials)", Lab on a Chip (2011)). Computational models or empirical methods can be used to determine the conductance for more complex channel geometries. (Note: For simplicity, in this specification, we focus on the conductance g rather than the fluid resistance R. The two quantities are simply related by g = 1/R.)

After the conductance of the internal channel is determined, the conductance g _e,i of the i-th unit of the expansion channel can be expressed as

g _e，i ＝if g

That is, the conductance of the first unit is fg, and the conductance of each successive unit increases by fg. Note that the flow rate qe _,i in the i-th unit of the expansion channel is related to the conductance by qe _,i = pge _,i , and q = pg, then the flow rate in the i-th unit of the expansion channel can be expressed as

q _e,i =if q

Thus, the flow rate in the first unit is fq, and the flow rate in each successive unit increases by fq.

The conductance g _c,i of the i-th unit of the contraction channel can be expressed as

g _c，i ＝2g-if g

That is, the conductance of the first unit is 2g-fg, and the conductance of each successive unit increases by fg. Note that the flow rate qc _,i of the i-th unit of the contracting channel is related to the conductance by qc _,i = pgc _,i , and q = pg, then the flow rate in the i-th unit of the expanding channel can be expressed as

q _c,i = 2q-if q

Thus, the flow rate in the first unit is 2q-fq, and the flow rate in each successive unit increases by fq.

The width of the dilated channel _{, we,i} , is chosen to give the desired g _,i , and the width of the contracted channel, wc _,i, is chosen to give the desired gc _,i . In practice, these widths can be found by evaluating the conductance for a wide range of channel widths (using the above formula) and then interpolating to find the channel width that gives the desired channel conductance.

The number of cells required to increase the flow in the diverging channel to q and to reduce the flow in the converging channel to q is n = 1/f. Thus, in the nth cell, the flow rates in all channels are equal: _qe,n = qc _,n = q.

After n units, the array can be "reset." For example, FIG8 is a schematic diagram illustrating a top view of an example particle sorting region 800, similar to region 700 shown in FIG7, wherein a "reset" region is introduced after n island structures 710 of the aforementioned units. During the reset, contracting channel 705b and the adjacent inner channel 703 combine to form a new contracting channel 709b, expanding channel 705a becomes an inner channel, and a new expanding channel 709a is introduced.

Unit length, width, movement, flow rate and particle size are the factors that most significantly affect device performance. Briefly, the impact of each factor is as follows:

The cell length determines how far (and how long) the inertial lift force acts on the particle, and thus how far laterally the particle migrates in each cell. To retain the particle, the cell must be long enough to allow the particle to escape the fluid it will travel through at the next opening between islands.

Flow rate also affects the magnitude of the inertial lift force and the lateral distance a particle travels in each cell. The (lateral) distance traveled per longitudinal distance is approximately proportional to the flow rate. To retain the particle, the flow rate must be fast enough to allow the particle to escape the fluid to be transported at the next opening between the islands.

• The motion does not directly affect particle migration, but rather determines how far (ie, how much fluid fraction) a particle must migrate to escape the fluid to move at the next opening between islands. The greater the motion, the farther the particle must migrate.

Cell width affects performance in two ways. First, cell width (and height) influences the magnitude of the inertial lift force acting on a particle, with the force decreasing as cell width increases. Second, width relates movement to the distance a particle must travel. In other words, for a given movement, the larger the cell width, the farther the particle must travel to escape the fluid to move through the next opening between islands.

The magnitude of the inertial lift force is strongly dependent on particle size, with the force increasing dramatically as particle size increases (D. Di Carlo., "Inertial Microfluidics," Lab on a Chip (2009)). Consequently, large particles migrate farther laterally per unit cell than small particles. It is this difference in migration speed that allows for size-based sorting of particles.

For exchanging fluids between channels, the above factors can be selected to ensure that particles of interest are retained in the rows of the array. For size-based particle sorting applications, the above factors can be selected so that a subpopulation of large particles is retained in the rows, while a subpopulation of small particles is not.

For example, the following set of parameters can be used for blood debulking (i.e., separating white blood cells (WBCs) from red blood cells (RBCs) and platelets): a cell length of approximately 200 μm, a cell width of approximately 50 μm, a cell depth of approximately 52 μm, a migration of approximately 3.0%, and a flow rate of approximately 80 μL/min/row (average velocity of 0.51 m/s). This set of parameters is very effective for separating white blood cells with very few red blood cells and platelets left over. In this case, white blood cells are spherical, typically >8 μm in diameter. Red blood cells are disc-shaped, approximately 7 μm in diameter, and approximately 1.5 μm thick (thus, they are expected to behave similarly to medium-sized spheres). Platelets are disc-shaped, 3 μm in diameter.

The islands have a length of 200 μm (i.e., the same as the cell length) and a width of 50 μm. The islands' purpose is simply to isolate the channels to create suitable flow conditions within the device. Therefore, the island width has no particular functional significance. The islands can be made slightly narrower or wider without significantly affecting device performance.

However, island width does affect fabrication difficulty. Fabrication difficulty primarily depends on the aspect ratio (height divided by width) of the structures within the microfluidic device. Devices with smaller aspect ratios are easier to manufacture at low cost and high yield. The aspect ratio can be determined in two ways. The minimum aspect ratio is the structure height h divided by the minimum structure width w _min . The overall aspect ratio is the structure height h divided by the diameter D of a circle with the same area as the structure. Here, D can be expressed as where A is the area of the structure.

Because the islands in the example above have a width of 50 μm and a height of 52 μm, they have a minimum aspect ratio of 1.04 and an overall aspect ratio of 0.46. This allows for simple fabrication of molded polydimethylsiloxane and epoxy devices, as well as injection-molded plastic devices. Therefore, the device is not only extremely useful from a functional perspective, but also fundamentally scalable and economical from a commercial perspective. Furthermore, the set of device parameters listed above can be varied to sort particles of other sizes.

A microfluidic device configured to move particles of a given size can, in certain embodiments, be scaled to effectively move particles of different sizes. For example, for a device that uses inertial lift to move particles across fluid streamlines, the size of the particle movement region can be scaled according to particle size and the flow conditions can be varied, as long as the particle Reynolds number, R _p , is maintained. The particle Reynolds number can be expressed as:

Wherein _U is the maximum channel velocity, a is the particle diameter, ν is the kinetic viscosity of the fluid and _D is the hydraulic diameter of the channel.For the channel with a rectangular cross section of width w and height h, _D can be expressed as (2wh)/(w+h), where h is the channel height and w is the channel width. For example, suppose there is a mobile region 1 that can effectively move particles of size a. A method for designing a mobile region 2 that can effectively move particles of size 2a is to amplify all dimensions of the mobile region 1 by 1 times (i.e., to double the length, width and height of all features). In order to maintain the same R _p in the mobile region 2, the maximum channel velocity u m must be reduced to 1/2.

Other methods of scaling the size of the particle movement region and flow conditions according to particle size are also possible.

For separation devices with straight channels that rely on inertial lift to move particles across streamlines, the following guidelines are provided for the design and operation of the device:

First, as described in "Inertial Microfluidics" (Di Carlo, Lab on a Chip (9), 3038-3046, 2009) (incorporated herein by reference in its entirety), the ratio of the transverse (across the channel) particle velocity _Uy to the longitudinal (in the direction of fluid flow) velocity _Uz is proportional to the particle Reynolds number _Rp and can be expressed as:

Here, _Um is the maximum channel velocity, a is the particle diameter, ν is the kinematic viscosity of the fluid, and _Dh is the hydraulic diameter of the channel. (For a channel with a rectangular cross-section of width w and height h, _Dh = (2wh)/(w+h).) The goal of the particle sorting apparatus described herein, in certain embodiments, is to use inertial lift to efficiently move particles across streamlines (i.e., to maximize _Uy / _Uz ). To this end, it is recommended that the channel dimensions and flow conditions be selected so as to maximize the particle Reynolds number, _Rp, in the particle channel to the extent permitted by other practical constraints such as operating pressure. Throughout the apparatus, the particle Reynolds number, Rp, in the particle channel is preferably greater than about 0.01, but it may be much greater than this, perhaps greater than 100. _{An Rp} of approximately 1 is a good intermediate goal.

For a given particle diameter a and kinematic viscosity ν, the target particle Reynolds number _Rp can be achieved through multiple different combinations of channel size and channel velocity. One possible strategy to increase _Rp is to select a very small hydraulic diameter _Dh (relative to a). However, the channel resistance is a quartic function of _Dh , and the cost of choosing an unnecessarily small _Dh is a significant increase in operating pressure. In contrast, the operating pressure scales linearly with the channel velocity _Um , so a good alternative strategy is to design the device with a moderate hydraulic diameter _Dh and then increase the channel velocity _Um (and thus _Rp ) as needed during operation to achieve high particle yield. For channels with a square cross-section, i.e., _Dh = w = h, a value of _Dh of approximately five times the particle diameter a is a reasonable choice: _Dh = 5a.

Second, the length (in the longitudinal direction) of the openings between islands is preferably, but not necessarily, greater than about a and less than or equal to about w. If the length of the openings is less than a, the openings may become blocked by particles, thereby interrupting flow through the openings. Openings with a length approximately equal to w are less likely to be blocked by particles and provide sufficient space between the two islands for fluid to cross to adjacent channels. Openings with a length greater than w are effective, but offer no particular benefit and come at the expense of wasted space.

Third, the length of the islands, l, is preferably greater than or equal to the length of the openings between the islands. Because particles experience inertial lift when they travel close to the islands rather than through the openings, particles should travel most of their longitudinal distance against the islands rather than across the openings between them. Stated another way, if the length of the islands and the openings between them are equal, then the particles experience inertial lift for exactly 50% of their distance. On the other hand, if the length of the islands is four times the length of the openings, then the particles experience inertial lift for 80% of their distance.

A loose upper limit on the length of the islands, l, is the length required for the particles to migrate to their equilibrium aggregation position. Any additional channel length beyond the length required for the particles to reach equilibrium does not help the particles to move across the streamlines. The formula for the channel length _Lf required for the particles to reach equilibrium is found in "Inertial Microfluidics" (Di Carlo, Lab on a Chip (9), 3038-3046, 2009) and can be expressed as:

Here μ is the dynamic viscosity, w is the channel width, ρ is the fluid density, U _m is the maximum channel velocity, a is the particle diameter, and f _L is a dimensionless constant ranging from about 0.02 to 0.05 for channels with aspect ratios (h/w) from about 2 to 0.5. Although L _f provides an upper bound, it is a loose upper bound and exceeds the optimal length l of the island. This is because the lift force on the particle is very strong near the channel wall (proportional to a ⁶ ), but decreases significantly with distance from the wall (proportional to a ³ near the center of the channel). Therefore, if the particles are kept near the channel wall by using islands with a length l significantly less than L _f , then the sorting device will cause the particles to move across the streamline more efficiently.

Given these conditions, a reasonable intermediate value for the island length is about l = 4w. This is an approximation and will necessarily depend on the values chosen for other parameters, such as the fluid movement _fs .

Fourth, the fluid movement _fs should be greater than 0.2%, and preferably greater than 1.0%. If the fluid movement is small, for example 0.1%, then the total number of movements (cells) required to move the particles across the width of the sorting array is very large, and the device itself must be very long. Assuming that the maximum channel velocity _Um is high enough to keep the particle Reynolds number _Rp within the specified range, then extremely small movements, for example 0.1%, should not be required. Depending on the maximum channel velocity _Um , a fluid movement _fs in the range of about 1% to 5% should work well for a device designed and operated as outlined herein.

For any given device design and particle size a, the final parameter selection is the device operating flow rate, which directly determines the particle Reynolds number _Rp and the maximum channel velocity _Um in the particle channel. For devices designed as summarized, there will be a lower end flow rate that provides good performance. Below this flow rate threshold, inertial lift will not be enough to move particles far enough away from the island walls to avoid moving through these islands with the fluid, thereby resulting in low particle production. Although the formula provided here allows for a rough estimate of the flow rate threshold, the most accurate and relevant method for determining the flow rate threshold is based on experience.

If the sorting device is used to classify particles according to size (ie, to have two or more different sized populations of particles), the operating flow rate should be selected so that the inertial lift force is sufficient to move the large particles without moving the small particles.

The design and operating parameter guidelines described here have been found to be suitable for cell sorting devices. However, other design and optimization strategies may also lead to effective, high-performance particle sorting devices.

Microfluidic device dimensions

For approximately spherical particles transported through a microfluidic device having at least two channels separated by an array of island-like structures with gaps between adjacent islands (e.g., see FIG1 ), the depth (e.g., measured along the X direction in FIG1 ) and width (e.g., measured along the Y direction in FIG1 ) of each microfluidic channel are preferably in the range of about 2 to about 50 times the diameter of a single particle. Relative to rigid structures formed with gaps through which fluid is extracted, the width of the structure can be up to about 10 times the width of a single microfluidic channel, while the length of the structure can be between about 0.25 times the channel width and about 50 times the channel width.

For example, for a roughly spherical particle having a diameter of about 8 μm, a microfluidic device having a configuration similar to that shown in FIG1 , having two microfluidic channels separated by an array of rigid structures, can have the following parameters: each microfluidic channel and island structure can have a depth of about 50 μm, each microfluidic channel can have a width of about 50 μm, each island structure can have a width of about 50 μm, and each island structure can have a length of about 200 μm.

Examples of other sizes are set forth below.

For example, the distance between the outer walls of the regions containing different fluid flow ranges, i.e., measured transversely to the direction of fluid flow, can be set to between about 1 μm and about 100 mm (e.g., about 10 μm, about 50 μm, about 100 μm, about 500 μm, about 1 mm, about 50 mm, about 10 mm, or about 50 mm). Other sizes are also possible. The width of each fluid flow range measured transversely to the direction of fluid flow can be set to between about 1 μm and about 10 mm (e.g., about 50 μm, about 100 μm, about 250 μm, about 500 μm, about 750 μm, about 1 mm, or about 5 mm). Other distances are also possible.

The length of the gap/opening between the island structures measured along the fluid flow direction (e.g., along the Z direction in Figure 1) can be set to between about 500nm and about 1000μm (e.g., about 1μm, about 2μm, about 54μm, about 106μm, about 50μm, about 100μm, about 200μm, about 500μm, or about 750μm). In some embodiments, the length of each continuous opening is greater than or less than the length of the previous opening. For example, in a channel configured to have a reduced fluid resistance along a fluid path, each continuous opening can be larger so that more fluid is extracted through the opening. The island structures separating different fluid flow areas can be configured to have a length between about 10nm and about 1000μm and a width between about 10nm and about 1000μm. Other gaps and island structure sizes are also possible.

The height of the fluid flow region and the island structure in the particle movement region (e.g., measured in the X direction along FIG. 1 ) is in the range of about 100 nm to about 10 mm. For example, the height of the channel can be about 500 nm, about 1 μm, about 5 μm, about 10 μm, about 50 μm, about 100 μm, about 500 μm, about 750 μm, about 1 mm, or about 5 mm. Other heights are also possible. The cross-sectional area that the microfluidic flow region has can fall in the range of, for example, about 1 μm ² to about 100 mm ² .

Microfluidic systems

In certain embodiments, the particle movement region of the microfluidic device described herein is part of a larger optional microfluidic system having a network of microfluidic channels. Such microfluidic systems can be used to facilitate control, manipulation (e.g., sorting, separation, segregation, mixing, aggregation, concentration), and separation of liquids and/or particles from complex parent samples. During separation, microfluidic elements provide important functions, such as biofluid processing or reproducible mixing of particles and samples.

For example, a microfluidic system can include additional areas that sort particles based on size and/or shape using techniques other than inertial lift. These other techniques can include, for example, deterministic lateral displacement. The additional area can utilize an array of notched networks, where fluid flowing through a notch is unevenly distributed into subsequent notches. The array comprises a network of notches arranged so that fluid flowing through the notches is unevenly divided, even though the notches may be of the same size. In contrast to the techniques described herein for separating particles based on a combination of inertial lift and fluid extraction, deterministic lateral displacement relies on collisions that occur when particles come into direct contact with the pillars that form the notches. The fluid flow is aligned at a small angle (flow angle) relative to the array's line of sight. Particles within the fluid flow that are larger than a critical size migrate along the array's line of sight, while particles that are smaller than the critical size follow the flow in a different direction. Flow within the device generally occurs under laminar flow conditions. Within the device, particles of different shapes can behave as if they were of different sizes. For example, lymphocytes are spheres with a diameter of approximately 5 μm, while red blood cells are biconcave discs with a diameter of approximately 7 μm and a thickness of approximately 1.5 μm. The long axis (diameter) of an erythrocyte is larger than the long axis of a lymphocyte, but the short axis (thickness) is smaller. If the long axis of the erythrocytes is aligned with the flow as they are driven through the array of pillars by the flow, then their fluid size is actually their thickness (about 1.5 μm), which is smaller than that of a lymphocyte. When an erythrocyte is driven through the array of pillars by the fluid flow, it tends to align its long axis with the flow and behave like a particle about 1.5 μm wide, which is actually "smaller" than a lymphocyte. Therefore, the area of decisive lateral displacement can separate cells based on their shape, even though the cells may have the same volume. In addition, particles with different deformabilities behave as if they have different sizes. For example, two particles with undeformed shapes can be separated by decisive lateral displacement because when the more deformable particle contacts an obstacle in the array, it can deform and change shape. Therefore, separation in the device can be achieved based on any parameter that affects the hydraulic size, including the particle's external dimensions, shape and deformability.

Additional information regarding microfluidic channel networks and their fabrication can be found in, for example, US Patent Application Publication No. 2011/0091987, US Patent Nos. 8,021,614, and 8,186,913, each of which is herein incorporated by reference in its entirety.

In some embodiments, the microfluidic system is included in a part for preparing a fluid sample carrying particles before the fluid sample is introduced into the particle movement zone. For example, Fig. 6 A is a schematic diagram showing a top view of an example of a microfluidic system 600, which includes a particle movement zone 601 (labeled as "sorting unit") similar to the particle movement zone shown in Figure 1. Other structures can also be used as the particle movement zone, such as any structure shown in Figures 2-5. For a structure comprising two or more inlet channels, system 600 can include one or more auxiliary fluid sources for these inlets.

6B is an enlarged view schematically illustrating the particle sorting region 601. As shown in the enlarged view, the region 601 includes a first microfluidic channel 650 and a second microfluidic channel 652 extending along the first microfluidic channel 650. The second and first microfluidic channels are separated from each other by an array of island structures 654, wherein each island in the array is separated from an adjacent island in the array by a notch or opening 656, which fluidically couples the first microfluidic channel 650 with the second microfluidic channel 652. The fluid resistance of the region 601 varies along the longitudinal direction of the region 601 (e.g., along the direction of fluid propagation) so that the fluid flowing in the first microfluidic channel 650 passes through the opening 656. Inertial lift causes particles flowing near the island structures in the first microfluidic channel 650 to cross streamlines, so that they do not enter the second microfluidic channel 652 with the fluid.

System 600 also includes a filtration section 603 (labeled "filter") and a particle aggregation section 605 (labeled "aggregation channel") located upstream of particle movement region 601. Filter section 603 includes a layout of multiple pillar structures of different sizes. According to this structural layout, filter section 603 is configured to filter particles contained in the incoming fluid based on particle size (e.g., average diameter) so that only particles of a predetermined size or smaller can be passed to the next stage of system 600. For example, for complex matrices, such as bone marrow aspirates, filter section 603 can be configured to remove bone fragments and fibrin clots to improve the efficiency of downstream enhanced concentration. In a typical layout, filter section 603 can include a pillar array with pillar sizes and array offsets that are designed to deflect particles larger than a certain size, thereby separating them from the main suspension. Typically, the size limit is determined based on the maximum particle size that can pass through the subsequent stage of system 600. For example, the filter 603 may be configured to filter/block the passage of particles having an average diameter greater than 50%, 60%, 70%, 80% or 90% of the minimum width of the passage in the particle movement region 601 .

Filter section 603 is fluid-bound to particle gathering section 605. Particle gathering section 605 is configured to, before particles are provided to particle movement region 601, pre-aggregate the particles exiting filter section 603 to the desired fluid flow line position. The advantage of making particles pre-aggregate is that it reduces the distribution of particles across channel width to a narrow lateral range. The aggregation line of particles can then be reset so that the probability of the wrong channel in particle sorting region 601 is reduced (for example, avoiding particles from entering the second microfluidic channel 652 but entering the first microfluidic channel 650). Pre-aggregation can be achieved using inertial aggregation technology. More details of inertial aggregation can be found in, for example, U.S. Patent No. 8,186,913, which is incorporated herein by reference in its entirety.

Once the particles have been sorted in the particle movement area 601, the sorted particles can be coupled to a separate processing area of the microfluidic system 600, or removed from the system 600 for additional processing and/or analysis. For example, the second channel of the particle movement area 601 is coupled to the first outlet 607, and the second channel of the particle movement area 601 is coupled to the second outlet 609.

external force

Other functionalities can be added to microfluidic systems to enhance the aggregation, concentration, separation, and/or mixing of particles. For example, in certain embodiments, additional forces can be introduced that result in targeted improvements in particle flow. Additional forces can include, for example, magnetic forces, acoustic forces, gravitational/centrifugal forces, electric forces, and/or inertial forces.

Fabrication of microfluidic devices

The process of manufacturing a microfluidic device according to the present disclosure is described as follows. First, a substrate layer is provided. The substrate layer can include, for example, glass, plastic, or a silicon wafer. An optional thin film layer (e.g., silicon dioxide) can be formed on the surface of the substrate layer using, for example, thermal or electron beam deposition. The substrate and optional thin film layer provide a foundation on which a microfluidic region can be formed. The thickness of the substrate can fall within the range of about 500 μm to about 10 mm. For example, the thickness of substrate 210 can be about 600 μm, 750 μm, 900 μm, 1 mm, 2 mm, 3 mm, 4 mm, 5 mm, 6 mm, 7 mm, 8 mm, or 9 mm. Other thicknesses are also possible.

After providing a substrate layer, microfluidic channels are formed on the substrate layer. The microfluidic channels include the various fluid flow paths for the particle movement region, as well as other microfluidic components of the system, including any filtration segments, inertial focusing segments, and magnetophoresis segments. Microfluidic channels for other processing and analysis components of the microfluidic device can also be used. The microfluidic channels and lid are formed by depositing a polymer (e.g., polydimethylsiloxane (PDMS), polymethyl methacrylate (PMMA), polycarbonate (PC), or cycloolefin polymer (COP)) in a mold that defines the fluid channel region. Once the polymer is cured, it is transferred and bonded to the surface of the substrate layer. For example, PDMS can first be injected into a mold (e.g., an SU-8 mold made using a two-step photolithography process (MicroChem)) that defines the microfluidic network of channels. The PDMS is then cured (e.g., by heating at 65°C for approximately 3 hours). Before transferring the solid PDMS structure to the device, the surface of the substrate layer is treated with oxygen plasma to enhance bonding. Alternatively, the microfluidic channels and lids can be made of other materials such as glass or silicone.

application

The new microfluidic techniques and devices described here can be used in a variety of different applications.

Centrifugal displacement

The particle movement techniques and devices disclosed herein can be used for centrifugal displacement. Generally, centrifugation is understood to include concentrating subpopulations within a fluid by applying centrifugal force to the fluid. Typically, this process requires a device with moving parts, which are prone to wear and breakage. In addition, the moving parts require complex and expensive manufacturing processes. Another problem with centrifugation is that it is a process that is typically used in closed systems, that is, centrifugation requires manual transfer of samples to and from the centrifuge.

In contrast, the currently disclosed technology can significantly increase the concentration of fluid components using relatively simple microstructures without the need for moving parts. This type of technology can be implemented as part of a single-port microfluidic system, so that fluid samples can be transferred to and from the particle movement area without manual intervention. In addition, particle movement can be extended to devices that require large processing capacity (i.e., the volume rate of fluid that can be processed) without significantly degrading particle separation efficiency. For example, the device disclosed herein can be configured to allow fluid flow of up to 10, 25, 50, 75, 100, 250, 500, 1000, 5000 or 10,000 μl/min. Other flow rates are also possible. For example, using the device 100 in Figure 1 as an example, if the second and first microfluidic channels 106, 108 have a depth of about 50 μm and a width of about 50 μm, then the device 100 can achieve a combined sample flow rate of up to about 5 mL/min. Changing the channel dimensions can change the maximum volume flow rate that the device can achieve. Additionally or alternatively, the volumetric flow rate can be tuned by varying the length of the island structures (see the above section "Design parameters of microfluidic devices"). Furthermore, multiplexing multiple channels (e.g., operating multiple particle sorting regions in parallel) can allow even higher flow rates.

In some embodiments, the particle mobilization technology allows for separation of particles based on size. For example, in a fluid sample containing particles of two different sizes, a particle sorting region according to the present disclosure can be used to separate large particles from small particles (e.g., by siphoning small particles from the fluid sample into an adjacent microfluidic channel while using inertial lift to retain large particles in the original microfluidic channel). In other examples, the fluid sample can include particles of three or more different sizes, wherein the particle sorting region is designed to sort the particles into different regions based on their different sizes.

Thus, in certain embodiments, particle mobilization techniques can provide significant cost and time savings over conventional centrifugation processes.Examples of applications where microfluidics can be used instead of centrifugation devices include bone marrow and urine analysis.

Detection of infectious agents

Additionally, the particle mobilization technology disclosed herein can be used as part of a research platform for studying analytes of interest (e.g., proteins, cells, bacteria, pathogens, and DNA), or as part of a diagnostic assay for diagnosing a patient's underlying disease state or infectious agent. By separating and aggregating particles in a fluid sample, the microfluidic devices described herein can be used to measure a variety of biological targets, including small molecules, proteins, nucleic acids, pathogens, and cancer cells. Further examples are described below.

Rare cell detection

The microfluidic devices and methods described herein can be used to detect rare cells, such as circulating tumor cells (CTCs) in blood samples or fetal cells in blood samples from pregnant females. For example, the enrichment of primary tumor cells or CTCs in blood samples can be enhanced to provide a rapid and comprehensive understanding of cancer. By combining the particle deflection technology described herein with magnetophoresis, different types of cells (such as circulating endothelial cells in heart disease) can be detected. Therefore, microfluidic devices can be used as powerful diagnostic and prognostic tools. The target and detection cells can be cancer cells, stem cells, immune cells, white blood cells, or other cells, including, for example, circulating endothelial cells (using antibodies to epithelial cell surface markers such as epithelial cell attachment molecule (EpCAM)) or circulating tumor cells (using antibodies to cancer cell surface markers such as melanoma cell adhesion molecule (CD146)). The systems and methods can also be used to detect CTC populations, small molecules, proteins, nucleic acids, or pathogens.

Fluid exchange

The microfluidic devices and methods described herein can be used to move cells from one carrier fluid to another. For example, the disclosed particle movement technology can be used to move cells into and out of a fluid stream containing reagents such as drugs, antibodies, cell stains, magnetic beads, cryoprotectants, lysine reagents, and/or other analytes.

The single particle transport zone can contain multiple parallel fluid streams (from multiple inlets) through which the moving cells pass. For example, white blood cells can move from the blood stream to a stream containing a dye and then to a buffer stream.

In bioprocessing and related fields, the described devices and techniques can be used to allow the aseptic, continuous transfer of cells from old media (containing waste) to fresh growth media. Similarly, extracellular fluids and cell products (such as antibodies, proteins, sugars, lipids, biopharmaceuticals, alcohols, and various chemicals) can be withdrawn from a bioreactor in a sterile, continuous manner while the cells remain within the bioreactor.

Isolation and analysis of cells

The microfluidic devices and methods described herein can be used to fractionate cells based on biophysical properties such as size. For example, the devices and methods can be used to fractionate blood into separate streams of platelets, red blood cells, and white blood cells. In another example, the devices and methods can be used to fractionate white blood cells into separate streams of lymphocytes, monocytes, and granulocytes.

They can be separated by being transported to the separate fluid outlet by the graded cell stream.As an alternative, the cell stream can be detected and analyzed (for example, using optical techniques) in real time to determine the number of cells in each liquid stream or the properties of cells in each liquid stream, for example size or granularity.

These techniques can be used to change cells or their carrier fluids before or during sorting to promote their classification and/or analysis. For example, large magnetic beads can be attached to special cell types to increase the effective size of such cells. Controlled cell collections can also be used to increase the effective size of cells. The temperature, density, viscosity, elasticity, pH, permeability and other properties of the fluid can change, thereby or directly affect the sorting process (for example, inertial action is viscosity-dependent), or indirectly affect the sorting process by changing the properties of the cells (for example, osmotic expansion or contraction).

Fluid sterilization and purification

The microfluidic devices and methods described herein can be used to remove pathogens, pollutants, and other specific contaminants from fluids. By allowing the contaminants to migrate across fluid flow lines, the contaminants can be removed from a fluid sample and collected as a separate waste stream.

Harvesting algae for biofuel

Harvesting algae from the growth medium is a major expense in biofuel production because algae grows in a very dilute suspension with neutral buoyancy, which makes it difficult to efficiently extract and enrich the algal biomass. The microfluidic devices and methods described herein can provide an efficient way to collect algae that does not rely on density or filtration. The described devices and techniques allow algae in a growth chamber to be extracted from the growth medium and enriched to a high capacity density. This can be done as a single step or as part of a continuous process. In addition, because the devices described herein can sort cells in a size-dependent manner, they can be designed to sort and enrich only larger algae that have reached maturity, while returning the immature smaller algae to the chamber.

Micro heat exchanger

The apparatus and methods described herein can be used to process not only liquid flows, but also gaseous and multiphase flows. A typical application of interest is in high-efficiency heat exchangers for integrated circuits. The high power density in microchips requires efficient removal of waste heat. This cooling becomes increasingly difficult as microchips are stacked together, reducing the overall surface-to-volume ratio. Liquid cooling, in which heat is transferred from a heat source to a flowing liquid, is one method for increasing the cooling rate of microchips. This cooling can be particularly effective near the boiling point of the liquid, as considerable energy is absorbed in the liquid-to-gas transition. However, the accumulation of vapor (bubbles) on the heat transfer surface can drastically reduce the heat flux. The apparatus and techniques described herein can be used to remove bubbles from the heat transfer surface, maximizing the liquid's heat absorption rate (via phase change) while minimizing the surface area in contact with the vapor. For this application, one side of the sorting module contacts the microchip heat source. As the liquid flowing through the module absorbs heat, bubbles form on the heat transfer surface and are then swept into the liquid stream (via fluid drag) once they reach a critical size. The sorting array then directs these bubbles across the sorting module and away from the heat transfer surface, maximizing heat flux from the microchip to the coolant.

Example

The present invention will be further described in the following examples, which do not limit the scope of the invention described in the claims.

Example 1: Volume reduction

Debulking is the removal of plasma, red blood cells (RBCs), platelets, and other small components (e.g., magnetic beads) from blood and other mixed fluids, such as bone marrow aspirates (BMAs), relative to nucleated cells (e.g., white blood cells (WBCs)). This is typically achieved by density centrifugation, which separates the blood into separate fractions based on density. However, in the following example, we describe the use of a separation device that relies on inertial lift to debulk the blood.

Device fabrication

To fabricate the microfluidic device, standard SU8 photolithography and soft etching techniques were used to fabricate the master mold and PDMS microchannels, respectively. Briefly, negative photoresist SU8-50 (Microchem Corp, MA) was spun at 2850 rpm to a thickness of approximately 50 μm, exposed to UV light through a polyester emulsion-printed photomask (Fineline Imaging, CO) defining the microfluidic network of channels, and then developed in BTS-220 SU8-developer (J.T. Baker, NJ) to form a relief mold. A 10:1 mixture of Sylgard 184 elastomer matrix and curing agent (Dow Corning, MI) was then cast onto the relief mold, allowed to cure in an oven at 65°C for 8 hours, and then removed from the SU8 master mold to form a microfluidic device cover with patterned channels. The inlet and outlet holes of the channels were punched out using a conventional sharp needle. The devices were then cleaned of particulates using low-residue tape and oxygen plasma bonded to pre-cleaned 1 mm thick microscope slides.

The device has the following parameters: cell length of 200 μm, cell width of 50 μm, cell depth of 52 μm (designed based on the cell length discussed above for Figure 7). The island structure has a length of 200 μm (ie, the same as the cell length) and a width of 50 μm.

Sample preparation

The performance of the debulking device was evaluated in a large number (n=63) of independent experiments. In each experiment, ≥2 mL of fresh whole blood with EDTA or ACD anticoagulant was diluted 1:1 with phosphate-buffered saline (PBS) (1X) with 1% F68 Pluronic.

Experimental steps and results

For each run, blood samples were introduced into the device using a syringe pump operating at 60 μL/min. A buffer co-flow (phosphate-buffered saline (1X) and 1% F68 Pluronic) was introduced into the device using a syringe pump operating at 272 μL/min. The composition of the input, product, and waste products was analyzed using a hematology analyzer (Sysmex KX21N). Manual counts were also performed on the product and waste using Neubaur and Nageotte chambers to ensure accurate data for cell types present at low concentrations (specifically white blood cells in the waste product and red blood cells and platelets in the product).

The data from the experiments are summarized in Figures 9A-9D. The median white blood cell yield was 85.9%, while the median neutrophil yield was 93.4%. The neutrophil yield was slightly higher than the total white blood cell yield, which is consistent with the fact that neutrophils are the largest white blood cell subset in terms of actual size. The purity of the product was very good. The median carryover rate for red blood cells (i.e., the percentage of input red blood cells that ended up in the product) was only 0.0054%, and the median carryover rate for platelets was only 0.027%, indicating that very few red blood cells and platelets remained in the same fluid stream as the white blood cells and neutrophils. Compared to other methods of microfluidic fractionation, the method proposed here is sensitive to the flow rate in the device. This is because inertial lift depends strongly on the flow rate. In an array, the relevant flow rate is the flow rate per row.

Additional experiments were also performed in which the flow rate was varied to evaluate its effect on debulking. Figures 10A-10B show a plot of white blood cell yield versus flow rate per row. At low flow rates (<10 μL/min), the inertial lift force was too weak to displace the white blood cells from the moving fluid stream, resulting in a yield of approximately 0%. As the flow rate increased toward 60 μL/min, the inertial lift force increased, and a greater proportion of the white blood cells escaped the moving fluid stream and reached the product, increasing the yield. At the highest flow rate (>60 μL/min), the inertial lift force was large enough to allow most of the white blood cells to cross the array and enter the product, and the white blood cell yield stabilized at approximately 85%. For reference only, the flow rate per row in the debulking experiment in Figure 9 was 80 μL/min.

The high dependence of leukocyte yield on flow rate suggests that the fractionation size threshold can be controlled by adjusting the flow rate per row. For any given flow rate, the inertial lift force on the particles depends on their size. Therefore, one would expect the curve shown in Figures 10A-10B to shift to the left for larger cells (or large cell subpopulations within the leukocyte population), and to shift to the right for smaller cells (or small cell subpopulations within the leukocyte population). Operating at a per-row flow rate that is large enough to allow large cells (e.g., neutrophils) to move across the array, but not large enough to allow small cells (e.g., lymphocytes) to move across the array, may be a way to isolate specific cell subpopulations with high purity.

Example 2: Evaluating the Effects of Particle Size, Fluid Flow Rate, and Fluid Movement

The effects of design and process factors on device performance are demonstrated in two experiments. The first experiment used fluorescent beads to demonstrate the effects of particle size and per-row flow rate on yield, while the second experiment used white blood cells (WBCs) to show the effect of migration on yield. The devices used in these experiments were fabricated using the same method described above in Example 1.

For the first experiment, several fluorescent beads of different sizes were used across a range of per-row flow rates. Each bead size was tested independently. In each case, a sample containing beads suspended in a buffer (phosphate-buffered saline (1X) with 1% F68 Pluronic) was introduced into the device along with the buffer flow.

Figure 11 shows the yield of fluorescent beads of different sizes at different flow rates. The total flow rate (sample + buffer) was selected to give the flow rate per row shown, and then the relative input flow rates of sample and buffer were selected so that 18% of the total input flow was sample. At the end of the device, particles that have migrated downward exit the device through the product channel and are collected in a vial. Particles remaining at the top of the array exit the device through the waste channel and are collected in a separate vial. The volumes of the product and waste vials were measured by mass, and the enrichment of the particles was determined using a standard Neubauer and Nageotte counting chamber. The relative yield was calculated as the fraction of the output beads in the product.

Several trends stand out in the resulting data. First, for any given bead size, the yield increases with flow rate. This is because inertial lift forces increase with flow rate. Second, for any given flow rate, the yield increases with bead size. This is because inertial lift forces increase dramatically with particle size. Taking 80 μL/min/row as an example, for 20 μm and 10 μm beads, the yield is 100%. This then drops to 68% for 8 μm beads, 1% for 7 μm beads, and 0% for 6 μm beads. Third, for any given device, the per-row flow rate provides a way to fine-tune the critical particle size. For example, to separate 10 μm particles from particles ≤7 μm, 80 μL/min is an ideal per-row flow rate. To separate 8 μm particles from particles ≤6 μm, 150 μL/min is an ideal per-row flow rate.

In a second experiment, the effect of fluid movement on leukocyte yield was evaluated. Leukocytes were isolated using the hydroxyethyl starch sedimentation method. Specifically, 1 mL of 6% hydroxyethyl starch (Stemcell Technologies HetaSep) was added to 10 mL of fresh whole blood, mixed, and then left to settle for 30 minutes. The top layer, enriched in leukocytes (and depleted in erythrocytes), was then aspirated using a pipette. The sample was introduced into one of six different devices, each with a different migration (2.5%, 3.0%, 3.2%, 3.4%, 3.6%, or 4.0%) and size as described above. In each case, the sample entered the device along with the buffer flow. The total flow rate (sample + buffer) was selected to give a per-row flow rate of 80 μL/min, and the relative input flow rates of sample and buffer were then selected so that 18% of the total input flow was sample. At the end of the device, leukocytes that had migrated downward (across the array) exited the device through the product channel and were collected in a vial. Leukocytes remaining on the top of the array exit the device through a waste channel and are collected in a separate vial. The volumes of the product and waste vials are measured by mass, and the enrichment of leukocytes is determined using a standard Neubauer and Nageotte counting chamber. Relative yield is calculated as the fraction of output leukocytes in the product.

FIG12 is a graph showing the yield of white blood cells as a function of fluid displacement. From a 2.5% displacement to a 3.2% displacement, the yield drops from 96% to 93%. Beyond a 3.2% displacement, the yield drops off steeply, reaching 71% at a 4.0% displacement. This indicates that for the smaller displacements tested, the migration of white blood cells due to inertial lift was large enough to allow substantially all white blood cells to escape the fluid moving between islands. However, for larger displacements, some white blood cells, presumably smaller white blood cells, were unable to escape the fluid moving between islands and thus remained in the waste product.

Multiplexing device

In some embodiments, island arrays, such as those described herein, can be multiplexed to create devices with very high throughput because the footprint of each array is very small. Figure 13 is an image of a standard microscope slide (25 mm x 75 mm) that accommodates 46 arrays operating in parallel and arranged in 23 duplexes. Multiplexing the arrays allows for a combined blood sample throughput of up to approximately 1.4 mL/min.

Other implementations

It will be understood that while the invention has been described in conjunction with the detailed description thereof, the foregoing description is intended to illustrate and not limit the scope of the invention, which is defined by the scope of the appended claims.

Claims (10)

1. A microfluidic device comprising: a particle sorting region comprising a first microfluidic channel, a second microfluidic channel extending along a first side of the first microfluidic channel, a first array of islands separating the first microfluidic channel from the second microfluidic channel, wherein each island in the first array is separated from an adjacent island in the first array by a gap that fluidically connects the first microfluidic channel to the second microfluidic channel, wherein a boundary of the second microfluidic channel is defined by walls of the second microfluidic channel and by walls of the islands of the first array, and wherein the walls of the second microfluidic channel are inclined away from the islands of the first array at an angle such that a width of the second microfluidic channel is greater than a width of the second microfluidic channel along a direction of fluid propagation. a third microfluidic channel extending along a second side of the first microfluidic channel, the second side opposite the first side of the first microfluidic channel, a second array of islands separating the first microfluidic channel from the third microfluidic channel, wherein each island in the second array is separated from an adjacent island in the second array by a gap fluidically connecting the first microfluidic channel to the third microfluidic channel, wherein a boundary of the third microfluidic channel is defined by walls of the third microfluidic channel and by walls of the islands of the second array, wherein the walls of the third microfluidic channel are inclined at an angle toward the islands of the second array such that a width of the third microfluidic channel decreases along the direction of fluid propagation, and wherein , the cross-section of the gaps between the islands in the second array increases along the longitudinal direction, the cross-section of each gap between the islands in the second array is defined along a plane transverse to the fluid flow through the gap, wherein the first microfluidic channel, the second microfluidic channel and the first array of islands are arranged such that the fluid resistance of the second microfluidic channel decreases relative to the fluid resistance of the first microfluidic channel along the longitudinal direction of the particle sorting region, so that during operation of the microfluidic device, a portion of fluid from the fluid sample in the first microfluidic channel passes through the first array into the second microfluidic channel, and wherein the first microfluidic channel, the third microfluidic channel and the second array of islands are arranged such that the fluid resistance of the third microfluidic channel increases relative to the fluid resistance of the first microfluidic channel along the longitudinal direction of the particle sorting region, so that during operation of the microfluidic device, a portion of fluid from the fluid sample in the third microfluidic channel passes through the second array into the first microfluidic channel, wherein the reduction in fluid resistance of the second microfluidic channel relative to the fluid resistance of the first microfluidic channel is a function of the increasing cross-sectional area of the second microfluidic channel along the longitudinal direction of the particle sorting region, wherein boundaries of the microfluidic channels are defined by device walls and walls of the islands, and wherein the width of the first microfluidic channel is substantially constant along the longitudinal direction. 2. The microfluidic device of claim 1 , wherein the cross-section of the gaps between the islands in the first array increases longitudinally, the cross-section of each gap between the islands in the first array being defined along a plane transverse to the flow of the fluid passing through the gap. 3. The microfluidic device of claim 1 , wherein the increase in fluidic resistance of the third microfluidic channel relative to the fluidic resistance of the first microfluidic channel is a function of a decreasing cross-sectional area of the third microfluidic channel along the longitudinal direction of the particle sorting region. 4. The microfluidic device according to claim 1 further comprises: a first inlet channel; and a second inlet channel, wherein each of the first inlet channel and the second inlet channel is fluidically connected to the particle sorting region. 5. A method for moving particles between fluid samples in a microfluidic device, the method comprising: flowing a first fluid sample comprising a plurality of first particles into a particle sorting region of the microfluidic device, wherein the particle sorting region comprises a first microfluidic channel, a second microfluidic channel extending along a first side of the first microfluidic channel, a first array of islands separating the first microfluidic channel from the second microfluidic channel, a third microfluidic channel extending along a second side of the first microfluidic channel, and a second array of islands separating the first microfluidic channel from the third microfluidic channel, the second side being opposite the first side of the first microfluidic channel, each island in the first array being separated from an adjacent island in the first array by a gap fluidically connecting the first microfluidic channel to the second microfluidic channel, and each island in the second array being separated from an adjacent island in the second array by a gap fluidically connecting the first microfluidic channel to the third microfluidic channel, wherein the cross-section of the gaps between the islands in the second array increases along a longitudinal direction, the cross-section of each gap between the islands in the second array being defined along a plane transverse to a flow of fluid through the gap, and wherein the width of the first microfluidic channel is substantially constant along the longitudinal direction; and flowing a second fluid sample into the particle sorting region, wherein the fluidic resistance of the second microfluidic channel is substantially constant. The fluid resistance of the first microfluidic channel varies along the longitudinal direction of the particle sorting region, so that a portion of the first fluid sample enters the second microfluidic channel from the first microfluidic channel through the gaps between the islands in the first array, and the variation of the fluid resistance of the second microfluidic channel relative to the fluid resistance of the first microfluidic channel includes an increase in the cross-sectional area of the second microfluidic channel along the longitudinal direction, wherein the fluid resistance of the third microfluidic channel varies along the longitudinal direction of the particle sorting region relative to the fluid resistance of the first microfluidic channel, so that a portion of the second fluid sample enters the first microfluidic channel from the third microfluidic channel through the gaps between the islands in the second array, and the variation of the fluid resistance of the third microfluidic channel relative to the fluid resistance of the microfluidic channel includes a decrease in the cross-sectional area of the third microfluidic channel along the longitudinal direction, wherein the first microfluidic channel, the second microfluidic channel and the first array of islands are further arranged to generate an inertial lift force, which inertial lift force substantially prevents the plurality of first particles from propagating with the portion of the first fluid sample passing through the gaps between the islands in the first array, and wherein the inertial lift force causes the plurality of first particles to move across fluid streamlines, so that the plurality of first particles are transported from the first fluid sample to the second fluid sample in the first microfluidic channel. 6. The method of claim 5, wherein the first fluid sample is delivered to a first microfluidic channel and the second fluid sample is delivered to a third microfluidic channel. 7. The method of claim 5, wherein the first fluid sample comprises a plurality of second particles, and wherein the plurality of second particles propagate with the fluid portion of the first fluid sample that enters the second microfluidic channel. 8. The method of claim 7, wherein the first particles are larger than the second particles. 9. The method of claim 5, wherein the first fluid sample comprises a different fluid than the second fluid sample. 10. The method of claim 5, wherein the amount of the first fluid sample entering the second microfluidic channel from the first microfluidic channel is substantially equal to the amount of the second fluid sample entering the first microfluidic channel from the third microfluidic channel, thereby maintaining a substantially constant concentration of the first particles in the first microfluidic channel.