Patent ID: 12220704

DETAILED DESCRIPTION

Interactions among particles within a fluid (e.g., cells, e.g., 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, virus, algae, any prokaryotic or eukaryotic cells, cell clusters, organelles, exosomes, droplets, bubbles, pollutants, precipitates, organic and inorganic particles, beads, bead labeled analytes, magnetic beads, and/or magnetically labeled analytes), the fluids in which the particles travel (e.g., blood, aqueous solutions, oils, or gases), and rigid structures can be controlled to perform various microfluidic operations on both the particles and fluid. In particular, such interactions may entail shifting the particles across fluid streamlines, through either the displacement of the fluid or the particles themselves. Examples of microfluidic operations that can be performed by controlling these interactions include, but are not limited to, increasing the concentration of particles in a carrier fluid, reducing the volume of a fluid sample, reducing the concentration of particles within a fluid, shifting particles from one carrier fluid to another fluid, separating particles within a fluid based on particle size (e.g., average diameter), focusing particles within a carrier fluid to a single-streamline (or to multiple different streamlines), precise positioning of particles at any position within a micro-channel, and mixing (defocusing) particles. Moreover, any of the above operations can be executed simultaneously with other techniques (e.g., magnetic sorting) to enhance the operation's effectiveness.

Several different mechanisms can be employed to create the forces capable of shifting particles across fluid streamlines. Any of the following techniques may be used individually or in combination to induce particle shifting within a fluid. A first type of force is referred to as “bumping” (also called deterministic lateral displacement (DLD)). Bumping is direct interaction between a rigid wall of a structure and a particle that arises due to the size of the particle relative to the wall. Since the center of a particle having radius rpcannot pass closer to an adjacent structure than rp, if the particle center lies on a streamline that is less than rpfrom the structure, the particle will be bumped out by the structure to a distance that is at least rpaway. This bumping may move the particle across fluid streamlines.

Another type of force is called inertial lift force (also known as wall force or wall induced inertia). The inertial lift force is a fluidic force on a particle that arise when then the particle and fluid flow near a wall. Though not well understood, the inertial lift force is a repulsive force arising due to a flow disturbance generated by the particle when the particle nears the wall. In contrast to bumping, the inertial lift force is a fluidic force on a particle, not a force due to contact with a rigid structure. A particle flowing near a micro-channel wall experiences an inertial lift force normal to the wall. At high flow rates, the inertial lift force is very strong and can shift the particle across streamlines.

Another type of force is a result of pressure drag from Dean flow. Microfluidic channels having curvature can create additional drag forces on particles. When introducing the curvature into rectangular channels, secondary flows (i.e., Dean flow) may develop perpendicular to the direction of a flowing stream due to the non-uniform inertia of the fluid. As a result, faster moving fluid elements within the center of a curving channel can develop a larger inertia than elements near the channel edges. With high Dean flow, drag on suspended particles within the fluid can become significant.

Another type of particle shifting occurs with high Stokes number flow. The Stokes number (Stk) describes how quickly a particle trajectory changes in response to a change in fluid trajectory. For Stk greater than 1, a lag exists between the change in fluid trajectory and the change in particle trajectory. Under high Stokes flow conditions (e.g., a Stokes number greater than about 0.01), changing the fluid flow direction can be used to force particles across streamlines. Further details on Dean flow and high Stokes number can be found, for example, in U.S. Pat. No. 8,186,913, which is incorporated herein by reference in its entirety. In both high Stokes flow applications and Dean flow applications, the fluid displacement causes the particles to cross fluid streamlines.

Other techniques for shifting particles include viscoelastic and inertio-elastic focusing. Details on those methods can be found in “Sheathless elasto-inertial particle focusing and continuous separation in a straight 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 foregoing techniques for shifting particles are “internal,” in that they use fluid flow and/or structures of the microfluidic channel itself to generate the forces necessary to shift particles across streamlines. In some cases, other external mechanisms can also be used in conjunction with one or more of the internal forces to alter the course of particles traveling within a fluid. For example, in some cases, externally applied magnetic forces, gravitational/centrifugal forces, electric forces, or acoustic forces may be used to cause a shift in particle position across fluid streamlines. Further information on how to apply such forces can be found, e.g., in WO 2014/004577 titled “Sorting particles using high gradient magnetic fields,”, U.S. Pat. No. 7,837,040 titled “Acoustic focusing,” WO 2004/074814 titled “Dielectrophoretic focusing,” and “Microfluidic, Label-Free Enrichment of Prostate Cancer Cells in Blood Based on Acoustophoresis,” Augustsson et al., Anal. Chem. 84(18), Sep. 18, 2012,

The present disclosure focuses primarily on combining inertial lift forces with periodic fluid extraction to shift particles across fluid streamlines to modify the concentration of and/or to filter particles in a fluid, though it should be understood that inertial lift forces may be replaced with or used in addition to other forces, such as those described above. As an example of combined inertial, particle containing fluids may be introduced into a microfluidic channel having an array of rigid island structures separating the channel from an adjacent microfluidic channel. As fluid is extracted from the first microfluidic channel into the second microfluidic channel through gaps between the island structures, the particles are drawn nearer to the island structures. As the particles reach nearer to the island structures, the particles experience a repulsive force (e.g., an inertial lift force) away from the direction of fluid extraction such that the particles cross fluid streamlines. The combination of fluid extraction and the repulsive forces may be used to perform positioning of particles, increasing the concentration of particles within a fluid, decreasing the concentration of particles within a fluid, particle mixing, fluid mixing, and/or shifting of fluids across particle streams, among other operations.

The mechanisms for shifting particles may be size-based and therefore can be used to perform size-based manipulation of particles (e.g., based on the average diameter of the particles). Through the repeated shifting of particles and/or displacement of fluid using any of the above-mentioned techniques, various different microfluidic operations may be performed, such as focusing particles to one or more fluid streamlines, increasing the concentration of particles within a fluid, performing volume reduction of a fluid, filtering particles from a fluid, and/or mixing different particles from different fluid streams. In general, “focusing” particles refers to re-positioning the particles across a lateral extent of the channel and within a width that is less than the channel width. For example, the techniques disclosed herein can localize particles suspended in a fluid within a length of the channel having a width of 1.05, 2, 4, 6, 8, 10, 20, 30, 40, 50, 60, 70, 80, 90, or 100 times the average diameter of the particles. In some implementations, the particles are focused to a streamline of a fluid. In some implementations, a streamline defines a width that is substantially equal to or slightly greater than a hydraulic diameter of the particle. Particles may have various sizes including, but not limited to, between about 1 μm and about 100 μm in average diameter.

Altering Particle Concentration Using Inertial Lift Forces

FIG.1is a schematic that illustrates a top view of an example of a microfluidic device100capable of shifting the position of particles102across fluid streamlines while the fluid propagates through the microfluidic device100. As will be explained, the particle shifting across fluid streamlines relies on the inertial lift forces experienced by particles as fluid is periodically extracted from a microfluidic channel, though other repulsive forces may be used in place of or in addition to inertial lift forces. For reference, a Cartesian coordinate system is shown, in which the x-direction extends into and out of the page.

During operation of the device100, a fluid carrying the particles102is introduced through an inlet microfluidic channel104. In this and other implementations of the particle shifting devices, the fluid can be introduced through the use of a pump or other fluid actuation mechanism. The inlet channel104splits into two different fluid flow channels (second microfluidic channel106and first microfluidic channel108substantially parallel to the second microfluidic channel106) that are separated by a 1-dimensional array of rigid island structures110. The 1-dimensional array of island structures110extends substantially in the same direction as the flow of the fluid through the second and first microfluidic channels. Each island structure110in the array is separated from an adjacent island110by an opening or gap114through which fluid can flow. Each gap114in the example ofFIG.1has the same distance between adjacent islands110. In other implementations, different gaps can have different distances between adjacent islands110. For example, in some implementations, a length of each subsequent opening (e.g., as measured along the fluid propagation direction—the z-direction inFIG.1) in the first array is greater than a size of a previous opening in the array. Furthermore, although a 1-dimensional array is shown inFIG.1, the islands110may be arranged in different configurations including, for example, a two-dimensional array of islands. The boundaries of the fluid flow regions within the microfluidic channels are defined by the device walls112and the walls of the islands110.

As the fluid propagates substantially along the z-direction (i.e., the longitudinal direction) from the inlet channel104to the channels (106,108), particles102experience a force (in this example, an inertial lift force) that causes the particles102to shift across fluid streamlines and travel along the first microfluidic channel108. These inertial lift forces are in the negative y-direction (see short arrows adjacent to each particle102inFIG.1).

For instance, when a particle102is located in the inlet channel104and approaches the top wall112, the particle experiences an inertial lift force that pushes the particle down toward the first microfluidic channel108. Once in the first microfluidic channel108, the particle102may approach a wall of the first island110, such that it again experiences an inertial lift force pushing the particle102down, maintaining the particle within the first microfluidic channel108. The repeated application of the inertial lift force to the particle102in each of the “particle shift” regions shown inFIG.1thus serves to separate/filter the particle from the fluid propagating through the second microfluidic channel106.

At the same time, portions of the fluid traveling in the first microfluidic channel108are extracted (e.g., siphoned)/pass into the second microfluidic channel at one or more “fluid shift” regions (seeFIG.1) in the device100. In the example ofFIG.1, each fluid shift region corresponds to an opening or gap that extends between the first microfluidic channel108and the second microfluidic channel106. Each “fluid shift” region primarily allows fluid to be extracted from the first microfluidic channel108into the second microfluidic channel106. The movement of fluid into the gaps tends to pull the particles102toward the gaps as well, since the particles follow the fluid streamlines. However, as the particles move closer to the gaps114, they approach the island structures112, which impart an inertial lift force causing the incident particles to cross fluid streamlines in a direction away from the gaps114. That is, the particles102shift from a fluid streamline passing into the second microfluidic channel106to a fluid streamline that continues to flow in the first microfluidic channel108. As a result, the particles102continue to propagate in the first microfluidic channel108and are not shifted into the second microfluidic channel106with the fluid. If there were no fluid shifting from the first microfluidic channel108to the second microfluidic channel106, the particles would migrate as a result of inertial focusing toward equilibrium focusing positions where the inertial lift force and shear gradient force are balanced. However, by shifting the fluid across the channels, the particles102tend to follow the fluid toward areas where the inertial lift force is much stronger than the shear gradient force, thus causing the particles to shift across streamlines in a very efficient and controlled manner.

In the present example, the fluid is extracted through the fluid shift regions as a result of decrease in fluidic resistance along a longitudinal section of the fluid shift region. That is, for a fluid of constant viscosity, the gaps114between adjacent islands110increase the channel area through which the fluid can flow, resulting in a reduced fluidic resistance. As fluid propagates through the device100and arrives at a gap114, a portion of the fluid will flow into the gap114and subsequently into the second microfluidic channel106(i.e., the fluid portion is extracted into channel106). The decrease in fluidic resistance also can occur as a result of the increasing channel width in the second microfluidic channel106. In particular, the second microfluidic channel wall112is slanted at an angle away from the islands so that the width of the second microfluidic channel106increases along the channel's longitudinal direction (i.e., in the direction of fluid propagation or the positive z-direction), thus causing a decrease in fluidic resistance. Any increase in the cross-sectional area of the channel106along the longitudinal direction of the first microfluidic channel, not just an increase in width, also can be employed to reduce the fluidic resistance. Alternatively, or in addition, the fluid may experience an increase in fluidic resistance in channel108relative to the fluidic resistance of channel106(e.g., through a decrease in the cross-sectional area of the channel108along the longitudinal direction). Thus, it may be said that the fluid is extracted in response to a change in the relative fluidic resistance between the second and first microfluidic channels. The change in the relative fluidic resistance may occur over the entire particle sorting region or over a portion of the sorting region that is less than the entire particle sorting region. The change in the relative fluidic resistance may occur over along the direction of the fluid flow through the particle sorting region (e.g., along a longitudinal direction of the particle sorting region as shown inFIG.1).

With progressively lower fluidic resistance at the gaps114and/or in channel106, greater amounts of fluid flow into the second microfluidic channel106. Furthermore, the repeated shifting of fluid into the second channel106reduces the amount of fluid in the first channel108. This constant fluid extraction thus increases the particle-to-fluid concentration in the first channel108, while decreasing the concentration of particles in the second microfluidic channel106, such that the fluid in the second microfluidic channel106is “filtered” or “purified.” In some implementations, the particle shifting techniques disclosed herein may be capable of increasing the particle concentration from an initial fluid sample by up to 10, 25, 50, 75, 100, 200, 300, 400, or 500 times the initial particle to fluid concentration. Such concentration increases can result in particle yields from fluid samples of up to 90%, up to 95%, up to 99% or even 100%.

In some implementations, the increases in particle concentrations may be achieved using multiple microfluidic devices configured to employ the particle shifting techniques disclosed herein. For example, the output of a first microfluidic device configured to increase the particle concentration of an incoming fluid sample by 10× may be coupled to an input of a second microfluidic device configured to increase the particle concentration of an incoming fluid sample by50X, for an overall increase in particles concentration from the initial fluid sample of 500×.

In addition to increasing particle concentration, the repeated particle shifting may also be used to focus the particles along one or more desired positions/streamlines within the fluid propagating through the lower channel108. For instance, as previously explained, portions of fluid may be extracted from an initial microfluidic channel into one or more parallel microfluidic channels. In some instances, the parallel microfluidic channels containing the extracted fluid then may be re-combined with the initial microfluidic channel downstream so that the particles are confined to designated streamlines in a single channel. An advantage of this technique of combining fluid shifting with inertial lift force is that particles may be focused to desired positions within the downstream channel (e.g., near the channel wall, at the middle of the channel, or halfway between the channel wall and the middle of the channel, among other positions) by controlling how much fluid is removed from each side of the initial channel, providing increased flexibility to the design and use of microfluidic devices. In contrast, for microfluidic systems based primarily on inertial focusing, one cannot choose the position of the focused stream within the channel.

The resulting concentrated and focused particle streamline may be coupled to a separate processing region of the microfluidic device100or removed from the device100for additional processing and/or analysis. Likewise, the “filtered” fluid in the second channel106may be coupled to a separate region of the microfluidic device100or removed from the device100for additional processing and/or analysis. In some implementations, the particles102entering the device100are “pre-focused” to a desired fluid streamline position that is aligned with the first microfluidic channel108. By pre-focusing the particles102to a desired position, the probability that particles inadvertently enter into the second microfluidic channel106can be reduced.

Other microfluidic device configurations different from the implementation shown inFIG.1also may be used to concentrate particles based on repeated particle and fluid shifting. For example,FIG.2is a schematic that illustrates an example of a device200for particle and fluid shifting, in which the particle shifting area includes two different microfluidic channels for extracting fluid, rather than one microfluidic channel. The device200includes an inlet microfluidic channel204that is fluidly coupled to a particle shifting region that has three different fluid flow regions (an second microfluidic channel206, a first microfluidic channel208, and a third microfluidic channel210). The second microfluidic channel206is separated from the first microfluidic channel208by a first array212of islands216. The third microfluidic channel210is separated from the first microfluidic channel208by a second array214of islands216. Each adjacent island in the first array212and each adjacent island in the second array214is separated by a gap for fluid shifting. The boundaries of the microfluidic channels are defined by the device walls218and the walls of the islands. The microfluidic channel walls218are slanted at angles away from the islands so that the widths of the second and third microfluidic channels (206,210) increase along the fluid propagation direction (i.e., the positive z-direction), thus causing a decrease in fluidic resistance in each channel.

The device200operates in a similar manner to the device100. In particular, as fluid propagates substantially along the z-direction from the inlet channel204to the channels (206,208,210), particles202within the fluid experience inertial lift forces in the “particle shift” regions upon approaching the walls of the inlet channel204and the walls of the island structures216. The inertial lift forces in the inlet channel204push the particles202toward the center of the fluid flow (i.e., the inertial lift forces “focus” the particles toward central fluid streamlines), such that they primarily flow into the first microfluidic channel208. Once the particles202enter the first microfluidic channel208, they experience inertial lift forces from the island structures216that continue to focus the particles202along one or more central streamlines extending through the channel208. At the same time, fluid is extracted into the second and third microfluidic channels (206,210) in the “fluid shift” regions due to the reduced fluidic resistance. The combination of the fluid shift regions and the particle shift regions serve to focus particles from the incoming fluid into the first channel208, while increasing the concentration of the particles at the same time. Any of the resulting fluid streams (from the second, first, or third channels) may be coupled to a separate region of the microfluidic device200or removed from the device200for additional processing or analysis. In some implementations, the variation in size/fluidic resistance of the second and third channels can be set so as to ensure that equal amounts of fluid flow in from the third channel and out the second channel at each unit.

In some cases, particle and fluid shifting can be used to create multiple different streams of focused/concentrated particles. For instance,FIG.3is a schematic of a device300in which particle shifting concentrates particles from one stream along two different microfluidic channels. The device300includes an inlet microfluidic channel304that is fluidly coupled to two different fluid flow regions (a second microfluidic channel306and a third microfluidic channel310). A single island structure312positioned at the coupling point between the inlet channel404and the second and third channels (306,310) splits fluid propagating from the inlet channel304into two streams: one propagating along the second channel306and one propagating along the third channel310. Downstream from the first island structure312, the second microfluidic channel306is separated from the third microfluidic channel310by both a first array314of islands318and a second array316of islands318. Each adjacent island in the first array314and each adjacent island in the second array316is separated by a gap for fluid shifting.

During operation of the device300, a fluid containing particles302enters from the inlet channel304. The fluid is separated by island312causing the fluid and the particles within the fluid to flow into either the second microfluidic channel306or the third microfluidic channel310. Once the particles302have entered the second and third channels (306,310), the particles remain concentrated within those channels due to repeated particle shifting (e.g., as a result of inertial lift forces) that occurs when the particles302approach the islands318. A first microfluidic channel308is used to repeatedly extract fluid from the second and third channels (306,310). In particular, the first channel308progressively increases in width, resulting in a lower fluidic resistance. Fluid is extracted from the second and third channels (306,310) at the gaps between the islands318and follows this path of lower resistance. The device300thus takes a fluid containing randomly distributed particles and focuses/concentrates those particles into two separate streamlines in the second and third microfluidic channels306,310. The resulting particle streamlines and may be coupled to separate outputs for additional processing or analysis.

The particle and shifting techniques described herein also may be used to shift particles from a first fluid to a second different fluid, where the concentration of the particles in the second fluid can be increased.FIG.4is a schematic that illustrates an example of a device400capable of shifting particles from one carrier fluid to another. The device400that includes two inlet microfluidic channels (404,406) coupled to a single microfluidic channel405for merging the fluids. The merging channel405is, in turn, coupled to a particle shifting area that includes two different flow regions (second microfluidic channel408and first microfluidic channel410). The second microfluidic channel408is separated from the first microfluidic channel410by an array of island structures412, in which each island412is separated from an adjacent island412by a gap414for fluid shifting. In addition, the top wall416of the second microfluidic channel408is slanted at an angle away from the islands412in order to decrease the fluidic resistance between the second and first microfluidic channels along the downstream fluid direction.

During operation of the device400, a first fluid (“Fluid1”) containing particles402is introduced in the first inlet channel404and a second fluid (“Fluid2”) having no particles is introduced into the second inlet channel406. Assuming the fluids are introduced at flow rates corresponding to low Reynolds numbers (and thus laminar flow), there is little mixing between the two different fluids in the merge region405, i.e., the two fluids essentially continue flowing as layers adjacent to one another. The fluid pathway within the merge region405is aligned with the fluid pathway of the first microfluidic channel410such that the merged fluids primarily flow into the first channel410. As the two fluids enter the first microfluidic channel410, the particles402within the first fluid experience inertial lift forces from the island structures412that are transverse to the direction of flow and that keep the particles402within the first microfluidic channel.

At the same time, the increasing width of the second microfluidic channel408(due to the slanted channel wall416) decreases the fluidic resistance in the openings414between the channels, such that portions of the first fluid are extracted into the second channel408at each gap between the islands412. Because the first fluid flows as a layer above the second fluid, it is primarily the first fluid that is extracted into the second channel408from the first channel410. After propagating for a sufficient distance past the islands412, most of the first fluid is extracted into the second channel408, whereas the particles402and most of the second fluid remain in the first channel410. Accordingly, the microfluidic device configuration shown inFIG.4is useful for transferring particles from one fluid to a second different fluid. In some implementations, the propagation distance is long enough so that the second fluid also is extracted into the second microfluidic channel408. In that case, the concentration of the particles402in the first microfluidic channel410can be increased. Although the implementation shown inFIG.4includes two inlet channels, additional inlet channels may be coupled to the microfluidic channels used for altering the particle concentration.

The microfluidic devices shown inFIGS.1-4implement particle shifting across fluid streamlines using inertial lift forces from the microfluidic channel walls and from the periodic arrays of island structures. Techniques other than inertial lift force may be used to shift particles across fluid streamlines. For example, internal repulsive forces arising due to bumping against the island structures, high Dean flow and/or high Stokes flow, such as inertial focusing, can be used to shift particles across fluid streamlines. Alternatively, or in addition, external forces such as magnetic forces, acoustic forces, gravitational/centrifugal forces, and/or electrical forces may be used to shift particles across fluid streamlines.

Additionally, the shape of the rigid island structures that separate different flow regions is not limited to the shapes shown inFIGS.1-4. For example, the rigid island structures may have shapes similar to posts, cuboids, or other polyhedrons in which the top and bottom faces are, or can be, congruent polygons. In some circumstances, such as at high flow rates, it is advantageous to use islands with streamlined, tapered ends (such as the shape of the island structures inFIGS.1-4), as the taper helps minimize the formation of flow re-circulations (eddies) that disrupt flow in unpredictable and undesirable ways. Other shapes for the rigid island structures are also possible. The long axis of the rigid island structures may be oriented at an angle with respect to the average flow direction of the fluid, the average flow direction of the particles, or the long axis of the region for altering the particle concentration. The shapes of the channel segments are not limited to the approximately rectangular shapes shown inFIGS.1-4. The channel segments may include curves or substantial changes in width. In cross-section, the channels described inFIGS.1-4may be square, rectangular, trapezoidal, or rounded. Other shapes for the channel cross-sections are also possible. The channel depth may be uniform across the region for altering the particle concentration, or the channel depth may vary laterally or longitudinally. Additionally, thoughFIGS.1-4show the microfluidic channels as approximately rectilinear pathways, the channels may be configured in other different arrangements. For example, in some implementations, the microfluidic channels may be formed to have a spiral configuration. For instance, the first microfluidic channel and the second microfluidic channel may be arranged in a spiral configuration, in which the first and second microfluidic channel are still be separated by the array of island structures, but where the longitudinal direction of fluid flow through the channels would follow a generally spiral pathway.

In some implementations, the microfluidic devices can be designed to incorporate redundancy so as to prevent particles that unintentionally pass with fluid through openings in a first array of island structures from ultimately being collected with the filtered fluid. For example, in some cases, the devices may be designed to include two or more “confinement channels” operating in parallel, i.e., two or more channels, such as channel108inFIG.1, that are designed to impart repulsive forces to substantially prevent particles from passing through openings in the island array. Since particles would need to overcome the repulsive forces associated with each additional channel, the probability of a particle escaping with fluid that passes through openings between islands decreases as more confinement channels are added.

In some implementations, the devices described herein may be used in conjunction with other microfluidic modules for manipulating fluids and/or particles including, for example, filters for filtering sub-populations of particles of certain sizes. In addition, the devices described herein may be used in series and/or in parallel within a microfluidic system.

Altering Particle Concentration/Reducing Fluid Volume Using Inertial Focusing and Fluid Shifting

Altering the concentration of particles within microfluidic samples is not limited to techniques that rely on a combination of fluid shifting with inertial lift forces and/or bumping forces to direct particles across fluid streamlines. Other internal forces, such as inertial focusing or viscoelastic focusing may be used in combination with fluid shifting as well.

With respect to inertial focusing, an inherent advantage is that the fluid forces depend on higher speed flows rather than low Reynolds number operation, thus leading to higher throughput, which is otherwise a common limitation of microfluidic devices.

Inertial focusing uses inertial forces to enable the precise lateral positioning of particles within a microfluidic channel, e.g., along a common streamline. Inertial focusing is based upon the notion that laminar flow of a fluid through microfluidic channels can result in the continuous and accurate self-ordering of particles suspended within the fluid from a randomly distributed state. In general, sorting, ordering, and focusing of particles in an inertial focusing system depends, inter alia, on the geometry of the microfluidic channel, the ratio of particle size to hydrodynamic cross-sectional size of the channel, and the speed of the fluid flow. Various channel geometries may require a predetermined particle-to-volume ratio of the particle to be focused to achieve a desired inter-particle spacing and thereby maintain ordering and focusing of those particles.

In general, a maximum particle-to-volume ratio for a specified particle size and channel geometry for inertial focusing alone can be determined using the formula:

MaxVolumeFraction=2⁢N⁢π⁢a23⁢h⁢w
where N is the number of focusing positions in a channel, a is the average focused particle diameter, h is the microfluidic channel height, and w is the channel width. Higher ratios may be achieved when additional forces are applied to the particles.

Different microfluidic channel geometries can be used to achieve inertial focusing of particles. For example, the microfluidic channel can be a symmetrically curved channel, such as S-shaped, sinusoidal, or sigmoidal. The channel can have various cross-sections, such as a rectangular, elliptical, or circular cross-section. Alternatively, the channel can be an asymmetrically curved channel having various shapes, cross-sections, and configurations as needed for a particular application (e.g., each curve in the channel can be a different size, or, for example, the odd-numbered curves in a channel may be a first size and shape and the even-numbered curves may be a second size and shape, or vice versa). For example, the channel can generally have the shape of a wave having large and small turns, where a radius of curvature can change after each inflection point of the wave. The maximum particle-to-volume ratio can be adjusted as necessary for the particular geometry.

The channel can be configured to focus particles within a fluid sample into one or more discrete streamlines at one or more equilibrium positions within the channel. In general, separation, ordering, and focusing are primarily controlled by a ratio of particle size to channel size and the flow characteristics of the system, but is independent of particle density. For example, analytes can have a hydrodynamic size that is in the range of about 1000 microns to about 0.01 microns. More particularly, analytes can have a hydrodynamic size that is in the range of about 500 microns to about 0.1 micron, such as between about 100 microns and about 1 micron. In general, the analyte size is limited by channel geometry. Analytes that are both larger and smaller than the above-described ranges can be ordered and focused within inertial focusing regions having laminar flow conditions.

Lateral migration of particles in shear flow arises from the presence of inertial lift, attributed mainly to the shear-gradient-induced inertia (lift in an unbounded parabolic flow) that is directed down the shear gradient toward the wall, and the wall induced inertia which pushes particles away from the wall. Particles suspended in fluids are subjected to drag and lift forces that scale independently with the fluid dynamic parameters of the system. Two dimensionless Reynolds numbers can be defined to describe the flow of particles in closed channel systems: the channel Reynolds number (Rc), which describes the unperturbed channel flow, and the particle Reynolds number (Rp), which includes parameters describing both the particle and the channel through which it is translating:

Rc=Um⁢DhvandRp=Rc⁢a2Dh2=Um⁢a2v⁢Dh.

Both dimensionless groups depend on the maximum channel velocity, Um, the kinematic viscosity of the fluid, and ν=μ/ρ (μ and ρ being the dynamic viscosity and density of the fluid, respectively), and Dh, the hydraulic diameter, defined as 2wh|(w+h)(w and h being the width and height of the channel, respectively, for a channel having a rectangular or square cross-section). The particle Reynolds number has an additional dependence on the particle diameter α. The definition of Reynolds number based on the mean channel velocity can be related to Rcby Re=2/3Rc. Inertial lift forces dominate particle behavior when the particle Reynolds number, Rp, is of order 1. Typically, particle flow in microscale channels is dominated by viscous interactions with Rp«1. In these systems, particles are accelerated to the local fluid velocity because of viscous drag of the fluid over the particle surface. Dilute suspensions of neutrally buoyant particles are not observed to migrate across streamlines, resulting in the same distribution seen at the inlet, along the length, and at the outlet of a channel. As Rp, increases, migration across streamlines occurs in macro scale systems. An example of Rp, that allows localization of a flux of cells from a blood sample within a rectangular or square channel is about 2.9, but this can range from about 0.02 to 2.9 or higher. Again, different microfluidic channel geometries can be used to achieve inertial focusing of particles, resulting in corresponding Reynolds numbers suitable for those channel geometries. Examples and further discussion of inertial focusing can be found, for example, in U.S. Pat. No. 8,186,913, which is incorporated herein by reference in its entirety.

Generally, inertial focusing is used to focus particles to one or more equilibrium positions and then flow the different focused streams of particles to distinct outputs, where the particles are then collected. However, by adding the repetitive removal of fluid from the focused stream, the ability of inertial focusing to substantially increase particle concentration within a fluid (and/or reduce the concentration of particles in a fluid sample) may be greatly improved. In particular, the technique relies on two different behaviors that enable a substantial and rapid reduction in fluid volume: 1) a fast depletion of the near wall regions and 2) a reduced shear gradient lift driven migration of particles to their equilibrium positions.

FIG.5is a schematic illustrating a top view of an example of a particle shifting area500of a microfluidic device, in which the particle shifting area500relies on inertial focusing in combination with repeated fluid extraction to enhance volume reduction from a particle-rich fluid sample. Fluid samples may be provided to particle shifting area500using, e.g., pumps, in a manner similar to that described with respect to other embodiments disclosed herein. The particle shifting area500includes an array of island structures504separating an elongated second fluid flow region506from an elongated first fluid flow region508. The first fluid flow region508may also be called the “focusing channel” and the second fluid flow region506may be called the “particle-free channel.” In the present example a particle containing fluid sample is introduced into flow region508, whereas, a particle-free fluid sample, which may be the same or different fluid as that propagating in region508, is introduced into flow region506.

Each island504is separated from an adjacent island504in the array by a corresponding gap510that allows fluid to cross between the second and first flow regions. In contrast to the devices shown inFIGS.1-4, the first flow region508has an undulating channel wall512(e.g., approximately sinusoidal in shape) in which the channel width (along the y-direction inFIG.5) alternates between being narrow and enlarged along the longitudinal direction (along the z-direction inFIG.5). Additionally, each island structure504has a curved contour that follows the curvature of portions of the peaks and troughs in the channel wall512. That is, a side of each island and an opposing side of the second channel have substantially matching contours. In the present example, this leads to flow region508having an undulating longitudinal pathway through which the particle-carrying fluid sample propagates.

More specifically, a first turn through flow region508is narrow and the matching contours of the wall512and island504have small radii of curvature, whereas a second adjacent turn through flow region508is wider and the matching contours of the wall512and island have larger radii of curvature. This pattern of a relatively small radius of curvature followed by a relatively larger radius of curvature is repeated over the length of the flow region508. Thus, the microfluidic channel is asymmetrically curved to create higher fluid speeds closer to the wall512than away from the wall512. Depending on the flow rate of a particle carrying fluid, the fluid pathway curvature of the first flow region508may generate inertial forces that focus and retain particles502along one or more fluid streamlines within the first flow region508.

Additionally, the fluidic resistance near the gaps510between islands504decreases so that a portion of fluid tends to follow the low resistance path and shift/flow into the second flow region506. This fluid flow also tends to pull particles502traveling with fluid in the direction of the gaps510. However, in certain implementations, the inertial forces generated by the undulating fluid pathway of this region are great enough to shift the particle502across fluid streamlines and away from the gaps510so that the particle502remains suspended in the portion of fluid traveling through the first flow region508. The second fluid flow region506can be configured to have a width that progressively increases so the fluidic resistance in that region decreases over the channel length. As a result, greater amounts of particle-free fluid will shift into the second fluid flow region farther downstream along the channel, and lead to an increase in particle concentration in the first fluid flow region508.

FIG.6Ais a schematic depicting how fluid streamlines may behave within a microfluidic device600that combines inertial focusing with repeated fluid extraction. The structure of the device shown inFIG.6Ais similar to the device500and includes an input region601, where a fluid suspension containing a dilute concentration of particles (e.g., cells) is introduced. As the dilute sample of particles enters the device, the fluid sample is accelerated when the microfluidic channel converges toward a first narrow neck region603. A particle-free layer (labeled “cell free layer” inFIG.6)605forms after the fluid sample passes through the neck region603as a result of the cells moving away from the wall by Dean flow. A portion of this particle-free layer605then passes/is siphoned off toward the second fluid flow region606at the first island structure612, whereas the particles remain in the first fluid flow region608. The amount of the fluid sample that passes into the second fluid flow region606depends on the hydraulic resistance of the openings and the second fluid flow region606relative to the hydraulic resistance of the first fluid flow region608. The process of accelerating the particle-rich fluid to create a particle-free layer, and passing the particle-free layer into the second fluid flow region606is repeated multiple times at each island612until the end of the device where the separate flows may be captured for further processing or removal from the device. For instance, the device600may be understood as having a repeating array of focusing units and siphoning units arranged in parallel (i.e., a “focusing-siphoning unit pair”). An example of the regions corresponding to a single focusing unit607and a single siphoning unit609are depicted inFIG.6A. The focusing unit607includes the area adjacent to an island structure612where the walls of the microfluidic channel have relatively high curvature to induce inertial focusing. The siphoning unit609includes the area adjacent to the same island structure, but opposite to that of the focusing unit607, that has relatively less curvature and which provides a wider pathway for fluid to travel, resulting in a lower hydraulic resistance. In the example shown inFIG.6, the width of each siphoning unit609(as determined along a direction transverse to fluid flow) increases along the direction of fluid flow, leading to lower fluidic resistance and therefore an increase in the amount of fluid passing from the first fluid flow region608.

FIG.6Bincludes plots of simulated fluid flow for different cross-sections of the device600shown inFIG.6A. The plots inFIG.6Bdepict the Dean flow vectors and velocity profile which causes the formation of the cell free layer. As can be seen from these plots, the overall flow speed, and thus the inertial force, of the fluid sample decreases along the length of the microfluidic channel as fluid passes into the second fluid flow region606. In other words, to achieve a given level of volume reduction, the flow speed must be reduced to a fixed degree, independent of the number of units used.

An important design consideration for a device that combines inertial focusing with repeated fluid extraction is the percentage of the fluid that is siphoned at each siphoning unit. Ideally, the greater the amount of particle-free fluid that is removed at each siphoning unit, the quicker one will be able to obtain a desired particle concentration in the particle-rich fluid. However, it is also the case that the higher the percentage of fluid that is siphoned, the greater is the risk that particles will be carried away with the siphoned fluid if the inertial forces do not shift the cells out of the larger siphoned fluid fraction.

FIG.7is a plot that depicts the results of a calculation based on the device structure shown inFIG.6B. The calculation was performed to determine the Cell Free Flow Fraction as a function of the number of siphon-focusing unit pairs and the percentage of fluid that passes into the particle-free layer at each opening between the island structures of the device. “Cell Free Flow Fraction” refers to the fraction of all fluid that has been siphoned out. For example, if the siphon percentage is 10%, then after one unit the cell free flow fraction is 10%. The other 90% remains in focusing units. Then, in the second unit remove 10% of the remaining 90% is removed (i.e., 9% of the overall fluid). Thus, after two units the Cell Free Flow Fraction is 19%. This continues on. The plot also includes two horizontal dashed lines, with the top line representing a factor of 50 times reduction in fluid volume of the particle-rich fluid, and the bottom dashed line representing a factor of 10 times reduction in fluid volume of the particle-rich fluid. The four different curves inFIG.7represent siphoning at four different percentages, with the smallest siphon percentage corresponding to the bottom curve and the highest siphon percentage corresponding to the top curve in the plot. As shown inFIG.7, higher siphon percentages (i.e., the percentage of fluid siphoned at each siphon unit) decrease the overall number of units required to reach an equivalent volume reduction factor seen at the intersections of the 10× and 50× dashed lines.

A microfluidic device that combines inertial focusing and siphoning is not limited to the configuration shown inFIG.5. For example, in some implementations, a combined inertial focusing and siphoning device may have a configuration that includes an second fluid flow channel, a first (center) fluid flow channel and a third fluid flow channel similar to the device shown inFIG.2, with the exception that the device would be constructed to induce inertial focusing in the center channel. For example, the center channel may be configured to have an undulating pathway/shape in which the channel width (as determined transverse to the direction of fluid flow) alternates between narrow and enlarged. This may be achieved by constructing each of the first and second array of island structures to have matching contours that alternate between regions of high and low curvature. As in the example ofFIG.2, fluid passes into the second and third channels at the openings/gaps between the island structures. Alternatively, in some implementations, the device can be constructed to induce inertial focusing in the second and third fluid flow channels. For example, each of the second and third fluid flow channels may be configured to have an undulating pathway/shape in which their widths alternate between narrow and enlarged. This may be achieved by constructing the walls of the second channel and the opposing array of island structures to have matching contours that alternate between regions of high and low curvature, whereas the walls of the third channel and an opposing array of island structures may also have matching contours that alternate between regions of high and low curvature. At the gaps/openings between the island structures in each array, fluid may pass from the second channel into the center channel and from the third channel into the center channel.

In some implementations, the combined inertial focusing and siphoning device may have two fluid inputs, similar to the device400shown inFIG.4, so that the device acts as a fluid exchanger, where particles are transferred from a first fluid to a second fluid. That is, a first fluid sample may be introduced through input406, whereas a second different fluid sample containing particles402may be introduced into input404. Initially, a portion of the second fluid sample containing the particles402and first fluid sample propagate side by side through channel410. The walls of the first channel410and the island structures412may be configured so that the first channel410has an undulating pathway/shape in which the width of the channel alternates between narrow and enlarged (similar to the configuration shown inFIG.5). The undulating channel410leads to focusing of the particles along streamlines within the first fluid sample in channel410. Simultaneously, portions of the second fluid sample that are free of particles402are extracted from channel410at the gaps414between islands412into the second channel408. After repeated extraction of the second fluid sample, the particles402eventually are entirely transferred to the first fluid sample within channel410, and the second fluid sample is particle free.

In some implementations, a microfluidic device includes a particle shifting area having multiple channels that rely on inertial focusing in combination with repeated fluid extraction. Using multiple channels allows, in some implementations, a substantial increase in the throughput of a microfluidic device. For example, multiple copies of the particle shifting area500shown inFIG.5may be arranged in parallel. The output of each of the channels containing the particles may be delivered to a common repository. Similarly, the output of each of the channels containing the particle-free fluid also may be delivered to a different common repository.

In contrast to conventional centrifugation, an advantage of devices that use the combined inertial focusing and siphoning techniques is that particles are exposed to heightened forces for a shorter duration (e.g., fractions of seconds) than during centrifugation (e.g., several minutes). Additionally, compaction of particles does not occur in the microfluidic volume reduction process. Cell compaction, which may occur in centrifugation processes, is known to mechanically damage certain cells as well as alter gene expression (see, e.g., Peterson, B. W., Sharma, P. K., Van Der Mei, H. C. & Busscher, H. J. “Bacterial Cell Surface Damage Due to Centrifugal Compaction,” Applied and Environmental Microbiology 78, 120-125 (2012), incorporated herein by reference in its entirety). Additionally, the short duration over which cells may be exposed to heightened forces in a combined siphoning and inertial focusing device results in little or no restructuring of cells' interiors. In contrast, centrifugation techniques are susceptible to causing the dislocation of organelles. Moreover, there is no need for sterile breaks between steps in the combined siphoning and inertial focusing devices, unlike when transferring samples from a centrifuge. Thus, compared to centrifugation, the combined siphoning and inertial focusing devices offer a more efficient closed system for performing common biomedical tasks.

Increasing Particle Concentration/Reducing Fluid Volume Using Viscoelastic Focusing

As explained above, viscoelastic focusing also may be used in combination with fluid shifting to alter the concentration of particles within a fluid sample. In some implementations, viscoelastic focusing includes the addition of specified concentrations (e.g., micromolar concentrations or other concentrations) of one or more drag-reducing polymers (e.g., hyaluronic acid (HA)) to a fluid that results in a fluid viscoelasticity that can be used to control the focal position of the particles within the moving fluid at different Reynolds numbers (Re).

With viscoelastic focusing, the volumetric flow rate at which a particle-carrying fluid is driven results in the formation of a localized streamline in the fluid at or near a center of the channel. The localized streamline defines a width that is substantially equal to or slightly greater than a hydraulic diameter of a particle within the fluid. By adding the drag-reducing polymer to a Newtonian fluid (e.g., water or a physiological saline solution), the particle in the fluid is focused into the localized streamline, creating particle-free regions at the edges of the channel (e.g., the regions closest to the channel boundaries or walls).

Thus, similar to inertial focusing, viscoelastic focusing enables the precise positioning of particles within a fluid along a common streamline. In contrast to inertial focusing, viscoelastic focusing has an equilibrium position at the center of the channel cross-section, i.e., along a longitudinal path extending in a direction of fluid flow and centered between walls of the channel. Viscoelastic focusing also works across large ranges of flow rates and Reynolds numbers. The technique of viscoelastic focusing thus can be coupled with fluid extraction as described herein (e.g., repetitive removal/siphoning of fluid from the focused stream) to substantially alter particle concentration within a fluid.

Any of the devices described herein may be used with viscoelastic focusing to focus particles to a streamline within a fluid and alter the particles' concentration within the fluid. For example, viscoelastic focusing may be used with the device200shown inFIG.2. A pump (not shown) connected to the inlet of channels206and210may be operated to drive a fluid that carries suspended particles202. In some implementations, the pump is operated to drive the fluid through the channels at volumetric flow rates that result in the formation of a localized streamline in the fluid at or near a center of the center channel208, e.g., defined by the axis220. The localized streamline220defines a width that is substantially equal to or greater than a hydraulic diameter of the particle202. The particles in the fluid are focused into the localized streamline220. The localized streamline220represents a portion of the fluid into which the suspended particles202are focused. That is, the suspended particles are focused into a streamline formed by the fluid flow at or near a center of the channel208. At the same time, fluid may be extracted at the gaps/openings between the islands212,214that separate the second and third channels206,210from the center channel208. Because the particles are focused to a center streamline, the particles202are located further away from the gaps between islands and are less likely to be carried out of the center channel208with the portions of the fluid sample being extracted into the second and third channels206,210. That is, at each gap a portion of particle-free fluid is extracted from the center channel208into either channel206or channel210, resulting in an increase in the concentration of particles within the center channel208. After repeated siphoning of fluid at the gaps, the concentration of the particles may be increased, e.g., from 10 to 100 times or more.

The fluid in which the particles202are suspended and which is flowed through the channels206,208,210can include a Newtonian fluid, e.g., water or other Newtonian fluid, or a drag-reducing polymer mixed with a Newtonian fluid. In general, any polymer (or material) that can decrease a drag on particles, e.g., by exerting viscoelastic normal stresses on the particles, at the volumetric flow rates described herein can be implemented as an alternative or in addition to HA. In other words, any material (e.g., polymer, or other material) which, when mixed with a Newtonian fluid, alters a drag on a particle suspended in the fluid-material mixture, relative to a drag on the particle suspended in the Newtonian fluid without the material can be implemented as an alternative or in addition to HA. Such materials can include, e.g., polyethylene oxide (PEO), polyacrylamide, gelatin, to name a few. The particles can include rigid particles, e.g., beads, or deformable particles. In some implementations, the particles can include biological particles, e.g., cells. The drag-reducing polymer can include hyaluronic acid (HA). The molecular weight of HA can be between 350 kDa and 1650 kDa. The Reynolds number of the fluid flow can be between 0.001 and 4500, e.g., between 0.01 and 20, between 0.01 and 15, between 0.01 and 10, between 0.01 and 1, between 0.1 and 1000, between 0.1 and 100, between 0.1 and 20, between 0.1 and 10, between 0.1 and 1, between 1 and 1000, between 1 and 100, or between 1 and 20. The concentration of the drag-reducing polymer can be between about 0.001-1% g/mL (0.00001-0.01 g/mL) such as between about 0.01-0.1% g/mL (0.0001-0.001 g/mL). Further discussion of viscoelastic focusing can be found, e.g., in WO 2015/116990, which is incorporated herein by reference in its entirety.

Microfluidic Device Design Parameters

The effect of various design parameters on the operation of the microfluidic device will now be described. For reference,FIG.16is a schematic illustrating a top view of an example particle and fluid shifting region1600containing a row of island structures1610. The row of island structures1610separates an “extraction” microfluidic channel1605from a “particle” microfluidic channel1607. The primary direction of fluid flow is indicated by the arrow1601. The width of the extraction channel1605(defined along the y-direction) expands along the length of the channel, whereas the width of the particle channel1607(defined along the y-direction) remains essentially constant along the length of the channel. During operation of the device, fluid is extracted into the extraction channel1605through the openings between the islands1610, while particles traveling within the particle channel1607are retained in the particle channel1607by repulsive forces, e.g., inertial lift forces. For the purposes of the following discussion, the channels and islands may be understood as being arranged into separate “units” (see Unit1, Unit2and Unit3inFIG.16). Specifically,FIG.16illustrates three units of an array with each unit including a portion of the exterior microfluidic channel1605, an island1610, and a portion of the particle channel1607.

The relevant design parameters for the particle and fluid shifting region1600include the length of each unit, the width of each channel, and the fluid shift for each unit. The fluid shift, fs, is the fraction of the fluid flow, q, that shifts between channels at each unit (i.e., at the openings between the island structures). Together these parameters determine the fluid conductance of the channels in each unit of the device. Thus, each unit has a particle channel with length li, a particle channel width wp,i, and a particle channel fluidic conductance gp,i, where i refers to the unit number. Each unit also has an extraction fluid channel with length li, an extraction channel width we,i, and an extraction channel fluidic conductance ge,i, where i refers to the unit number. In the example described here, all channels are rectangular in shape and the fluid shift is the same for each unit. The basic method presented here can be easily modified for non-rectangular (e.g., curving) channels and varying shift.

At each unit, the total flow divides between the particle and extraction fluid channels in proportion to their relative fluidic conductances. Thus, the fraction of the total flow that flows through the particle channel1607in the ithunit is

fp,i=qp,iqp,i+qe,i=gp,igp,i+ge,i
where qp,iand qe,iare the flow rates of the particle and extraction fluid channels, respectively. Similarly, the fraction of the total flow that flows through the extraction fluid channel1605in the ithunit is

fe,i=qe,iqp,i+qe,i=ge,igp,i+ge,i

The dimensions of the particle channel1607are chosen to optimally shift particles across streamlines (e.g., away from the extraction fluid channel1605). Because the flow rate qp,ichanges along the length of the device, the particle channel dimensions may be altered to maintain optimal particle shifting. For example, as qp,idecreases, the unit length limay be increased to compensate for the weakening inertial lift force operating on particles.

The dimensions of the extraction fluid channel1605are chosen to provide a conductance ge,isuch that a precise fraction of the fluid in the particle channel1607shifts to the extracted fluid channel at each unit. This fractional amount is called the fluid shift, fs. The result of this shifting is that the fraction of flow in the particle channel decreases by a fixed factor at each unit:

fp,i=(1-fs)⁢fp,i-1

For example, if fs=0.1, then fraction of flow in the particle channel will be 90% of the fraction of flow in the particle channel of the previous unit. More generally, because fp,0=1,

fp,i=(1-fs)i

Thus, for the example case shown inFIG.16in which the particle and fluid shifting region is divided into three units, fs=0.1, fp,1=0.9, fp,2=0.81, and fp,3=0.729.

Recall that the fraction of flow in the particle channel is also described by

fp,i=gp,igp,i+ge,i

Substituting for fp,iand solving for ge,i:

ge,i=((1-fs)-i-1)⁢gp,i

Thus, for each unit the conductance of the extracted fluid channel can be written in terms of the conductance of the particle channel and the fluid shift. The fluidic conductance, g, of each channel is a function of its dimensions and the fluid viscosity. In the device described here, each channel is rectangular and therefore has conductance that can be expressed as

g≈(h41⁢2⁢η⁢l⁢α)⁢(1-0.6⁢3⁢α)

Here, η fluid viscosity, l is channel length, w is channel width, h is channel height, and α=h/w. A more accurate infinite series-based formula is also available (Tanyeri et al., “A microfluidic-based hydrodynamic trap: Design and implementation (Supplementary Material).”Lab on a Chip(2011).) Computational modeling or empirical methods can be used to determine the conductance of more complex channel geometries. (Note that in this description it is simpler to focus on fluidic conductance, g, rather than fluidic resistance, R. The two quantities are simply related by g=1/R.)

Using the above formulas, a microfluidic device for increasing the concentration of particles within a fluid sample may be implemented as follows:1. The dimensions of the particle channel are chosen for each unit in the device. As mentioned, the dimensions are chosen to optimally shift particles away from the extracted fluid channel.2. Using these dimensions and the fluid viscosity, the particle channel conductance gp,iis determined for each unit using the rectangular channel conductance formula (or an equivalent method).3. The extraction fluid channel conductance ge,iis then evaluated for each unit using the previously determined gp,iand fs. The width of the extraction fluid channel, we,i, is then chosen to give the desired ge,ifor each unit. In practice, the width may be determined by evaluating fluidic conductance (using the above formula) across a wide range of channel widths and then interpolating to find the channel width that gives the desired channel conductance.

For concentrators with straight channels that rely on inertial lift forces to shift particles across streamlines, the following are device design and operation guidelines:

First, as described in “Inertial Microfluidics,” Di Carlo, Lab Chip (9), 3038-3046, 2009 (incorporated herein by reference in its entirety), the ratio of the lateral (across channel) particle velocity Uyto the longitudinal (in direction of fluid flow) velocity Uzis proportional to the particle Reynolds number Rp:

UyUz∝Rp=Um⁢a2v⁢Dh
Here Umis the maximum channel velocity, a is the particle diameter, ν is the kinematic viscosity of the fluid, and Dhis the hydraulic diameter of the channel. (For channels of rectangular cross-section with width w and height h, Dh=(2wh)/(w+h).) Because it is the aim of the particle concentrator device described here to use inertial lift forces to efficiently move particles across streamlines (e.g., maximize Uy/Uz), it is recommended that the channel dimensions and flow conditions be selected so as to maximize particle Reynolds number Rpin the particle channel to the extent permitted by other practical constraints, such as operating pressure. Throughout the device, the particle Reynolds number Rpin the particle channel should ideally be greater than about 0.01, though it may be much larger than this, possibly greater than 100.

For a given particle diameter α and kinematic viscosity ν, a target particle Reynolds number Rpcan be achieved through many different combinations of channel dimensions and channel velocities. One strategy for increasing Rpwould be to select a very small (relative to a) hydraulic diameter Dh. However, channel resistance has a quartic dependence on Dh, and choosing an unnecessarily small Dhcomes at the cost of highly increased operating pressure. On the contrary, the operating pressure scales linearly with channel velocity Um, so a good alternative strategy is to design a device with a modest hydraulic diameter Dhand then increase channel velocity Um(and therefore Rp) at the time of operation as needed to achieve high yield of particles. For a channel with square cross-section, such that Dh=w=h, a value of Dhapproximately five times the particle diameter a is a reasonable choice: Dh=5α.

Second, the length of the openings (in the longitudinal direction) between islands should be greater than about α and less than or equal to about w. If the length of the opening is less than α, the opening may clog with particles, thereby disrupting flow through the opening. An opening with length approximately equal to w is unlikely to clog with particles and provides adequate room for fluid to cross between islands to the adjacent channel. An opening with a length greater than w will work but provides no particular benefit and comes at the cost of wasted space.

Third, the length of the islands l should be greater than or equal to the length of the openings between islands. As aforementioned, it is the aim of the particle concentrator device to use inertial lift forces to efficiently move particles across streamlines. Because particles only experience inertial lift forces as they travel alongside islands, particles should travel most of their longitudinal distance alongside islands, rather than across openings between islands. Put another way, if the length of islands and the length of the openings between islands are equal, then particles experience inertial lift forces along just 50% of the distance they travel. On the other hand, if the length of the islands is four times the length of the openings, then particles experience inertial lift forces along 80% of the distance they travel.

A loose upper limit on the length of islands l is the length required for particles to migrate to equilibrium focusing positions. Any additional channel length beyond what is required for particles to reach equilibrium does not contribute to shifting particles across streamlines. A formula for the channel length Lfrequired for particles to reach equilibrium is given in “Inertial Microfluidics,” Di Carlo, Lab Chip (9), 3038-3046, 2009:

Lf=π⁢μ⁢w2ρ⁢Um⁢a2⁢fL

Here μ is dynamic viscosity, w is channel width, ρ is fluid density, Umis the maximum channel velocity, α is the particle diameter, and fLis a dimensionless constant ranging from about 0.02 to 0.05 for channels with aspect ratios (h/w) ranging from about 2 to 0.5. While Lfprovides an upper bound, it is a loose upper bound and exceeds the optimal length of islands l. This is because the lift force on particles is very strong near the channel wall (proportional to α6), but falls off sharply with distance from the wall (proportional to α3near the center of the channel). Thus, a concentrator device will more efficiently shift particles across streamlines if the particles are kept near the channel wall by using an island length l that is significantly less than Lf.

Given these considerations, a reasonable intermediate value for the island length is about l=4w. This is an approximate value and necessarily depends on the values selected for other parameters, such as the fluid shift fs. It is also important to note that the length of the islands l need not be constant along the length of the device. Rather, as the maximum channel velocity Umand particle Reynolds number Rpin the particle channel decrease, the lengths of the islands can be increased to compensate. For example, a factor of two decrease in Rpcan be compensated by a factor of two increase in island length l. Up to a point, the lateral deflection distance of particles per unit is expected to be roughly proportional to the island length l.

Fourth, the fluid shift fsshould be greater than 0.2% and ideally greater than 1.0%. If the fluid shift is small, e.g., 0.1%, then the total number of shifts (units) needed to achieve a significant volume reduction, e.g., 10×, is very large and the device itself must therefore be very long. Provided the maximum channel velocity Umis sufficiently high to place the particle Reynolds number Rpin the prescribed range, an extremely small shift, e.g., 0.1%, should not be necessary. Depending on the maximum channel velocity Um, a fluid shift fsin the range of about 1% to 5% should perform well for a device designed and operated as outlined here.

It is important to note that the fluid shift fs, like the length of the islands l, need not be constant along the length of the device. Rather, as the maximum channel velocity Umand particle Reynolds number Rpin the particle channel decrease, the fluid shift fscan be reduced to compensate. For example, a factor of two decrease in Rpcan be compensated by a factor of two decrease in fluid shift fs. Either or both of these compensation strategies can be implemented to optimize device efficiency and performance.

For any given device design and particle size α, the final parameter choice is the device operating flow rate, which directly determines the maximum channel velocity Umand the particle Reynolds number Rpin the particle channel. For a device designed as outlined, there will be a minimum flow rate required for good performance. Below this threshold flow rate, the inertial lift forces will be insufficient to shift particles far enough from the island wall to avoid being shifted as fluid is extracted (siphoned), thus resulting in low yield of particles. While the formulas provided here enable one to make rough estimates of the threshold flow rate, the most accurate and relevant method of determining the threshold flow rate is empirically.

Other design and optimization strategies may also result in effective, high performance concentrator devices.

A microfluidic device that is configured to shift particles of a given size can, in some implementations, be scaled to effectively shift particles of a different size. For instance, for a device that employs inertial lift forces to shift particles across fluid streamlines, one can scale the dimensions of the particle shifting area with particle size and alter the flow conditions, so long as the value of the particle Reynolds number, Rp, is preserved. The particle Reynolds number can be expressed as:

Rp=Um⁢a2v⁢Dh
where Umis the maximum channel velocity, α is the particle diameter, ν is the kinematic viscosity of the fluid, and Dhis the hydraulic diameter of the channel. (For channels of rectangular cross-section with width w and height h, Dh=(2wh)/(w+h).) For example, consider a Shifting Area 1 that effectively shifts particles of size α. One method of designing a Shifting Area 2 that effectively shifts particles of size 2α is scale all dimensions of Shifting Area 1 by a factor of 2 (i.e., double the length, width, and height of all features). To maintain the same Rpin Shifting Area 2, the maximum channel velocity Ummust be decreased by a factor of 2.

Other methods of scaling the dimensions of particle shifting areas and flow conditions with particle size are also possible.

Ease of microfluidic device manufacturing is largely determined by the aspect ratio (height divided by width) of the device structures, with smaller aspect ratio devices being easier to manufacture at low cost and with high manufacturing yield. We can define the aspect ratio in two ways. The minimum aspect ratio is the structure height, h, divided by the minimum structure width, wmin. 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=√(4A/π), where A is the area of the structure.

As an example, for a microfluidic device having substantially straight channels, the island structures may have a length of about 50-1000 μm, a width of about 50 μm, and a height of about 52 μm. With these dimensions, the minimum aspect ratio of the islands is 1.04, and the overall aspect ratio is in the range 0.92-0.21. The aspect ratio could be further reduced by increasing the width of the islands. In another example, for a microfluidic device having curved channels, the island structures may have an irregular shape with a wminin the range of about 42-80 μm, A in the range of about 18,000-61,000 μm2, and a height of 52 μm. With these dimensions, the minimum aspect ratio of the islands is in the range 1.24-0.65, and the overall aspect ratio is in the range 0.34-0.19.

In both cases, the low aspect ratio of the structures enables straightforward fabrication of molded PDMS and epoxy devices, as well as injection molded plastic devices. This is a major advantage of this class of devices: they are not only extremely useful from a functional perspective, but they also are fundamentally scalable and economical from a commercial perspective.

Microfluidic Device Dimensions

For generally spherical particles being transported through a microfluidic device having at least two channels separated by an array of island structures, with gaps between adjacent islands (see, e.g.,FIG.1), the depth (e.g., as measured along the x-direction inFIG.1) and width (e.g., as measure along the y-direction inFIG.1) of each microfluidic channel is preferably in the range of about 2 times to about 50 times the diameter of a single particle. With respect to the rigid structures that form the gaps through which fluid is extracted, the width of the structures may be up to about 10 times the width of the a single microfluidic channel, whereas the length of the structures may be between about 0.25 times the channel width up to about 50 times the channel width.

As an example, for a generally spherical particle having a diameter of about 8 microns, a microfluidic device having two microfluidic channels separated by an array of rigid structures similar to the configuration shown inFIG.1may have the following parameters: each microfluidic channel may have a depth about 52 μm, each microfluidic channel may have a range of widths between about 10 μm to about 5000 μm, each island structure may have a width of about 50 μm, each island structure may have a length of about 200 μm.

Other examples of dimensions are set forth as follows.

For instance, the distance between the outer walls of the area containing the different fluid flow regions, i.e., as measured transverse to the fluid flow direction, can be configured to be between about 1 μm to about 100 mm (e.g., about 10 μm, about 50 μm, about 100 μm, about 500 μm, about 1 mm, about 5 mm, about 10 mm, or about 50 mm). Other sizes are possible as well. The width of each fluid flow region/channel (e.g., the width of second and first microfluidic channels106and108inFIG.1), measured transverse to the fluid flow direction, can be configured to be between about 1 am to 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 possible as well.

The length of the gaps/openings between the island structures, as measured along the fluid flow direction (e.g., along the z-direction inFIG.1), can be configured to be between about 500 nm to about 1000 μm (e.g., about 1 μm, about 2 μm, about 5 μm, about 10 μm, about 50 μm, about 100 μm, about 200 μm, about 500 μm, or about 750 μm). In some implementations, the length of each successive opening is greater than or less than the length of the last opening. For example, in a channel configured to have a decreasing fluidic resistance along the fluid pathway, each successive opening may be larger so that a greater amount of fluid is extracted through the opening. The island structures that separate different fluid flow regions can be configured to have a maximum length between about 10 nm to about 10 μm, and a maximum width between about 10 nm to about 10 μm. Other dimensions for the gaps and island structures are possible as well.

The height of the fluid flow regions and the island structures within the particle shifting area (e.g., as measured along the x-direction inFIG.1) are within the range of approximately 100 nm to approximately 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 possible as well. The microfluidic flow regions can have a cross-sectional area that falls, e.g., within the range of about 1 μm2to about 100 mm2.

Microfluidic Systems

In some implementations, the particle shifting areas of the microfluidic devices described herein are 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., separation, segregation, mixing, focusing, concentration), and isolation of liquids and/or particles from a complex parent specimen. During the isolation process, microfluidic elements provide vital functions, for example, handling of biological fluids or reproducible mixing of particles with samples.

For example, the microfluidic system may include additional areas for separating particles according to size and/or shape using other techniques different from inertial lift forces. These other techniques include, for example, deterministic lateral displacement. These additional areas may employ an array of a network of gaps, in which a fluid passing through a gap is divided unequally into subsequent gaps. The array includes a network of gaps arranged such that fluid passing through a gap is divided unequally, even though the gaps may be identical in dimensions. In contrast to the techniques described herein for separating particles based on a combination of inertial lift forces and fluid extraction, deterministic lateral displacement relies on bumping that occurs when the particle comes into direct contact with posts forming the gaps. The flow of the fluid is aligned at a small angle (flow angle) with respect to a line-of-sight of the array. Particles within the fluid having a hydrodynamic size larger than a critical size migrate along the line-of-sight in the array, whereas those having a hydrodynamic size smaller than the critical size follow the flow in a different direction. Flow in the device generally occurs under laminar flow conditions. In the device, particles of different shapes may behave as if they have different sizes. For example, lymphocytes are spheres of ˜5 μm diameter, and erythrocytes are biconcave disks of ˜7 μm diameter, and ˜1.5 μm thick. The long axis of erythrocytes (diameter) is larger than that of the lymphocytes, but the short axis (thickness) is smaller. If erythrocytes align their long axes to a flow when driven through an array of posts by the flow, their hydrodynamic size is effectively their thickness (˜1.5 μm), which is smaller than lymphocytes. When an erythrocyte is driven through an array of posts by a hydrodynamic flow, it tends to align its long axis to the flow and behave like a ˜1.5 μm-wide particle, which is effectively “smaller” than lymphocytes. The area for deterministic lateral displacement may therefore separate cells according to their shapes, although the volumes of the cells could be the same. In addition, particles having different deformability behave as if they have different sizes. For example, two particles having the undeformed shape may be separated by deterministic lateral displacement, as the particle with the greater deformability may deform when it comes into contact with an obstacle in the array and change shape. Thus, separation in the device may be achieved based on any parameter that affects hydrodynamic size including the physical dimensions, the shape, and the deformability of the particle.

Additional information about microfluidic channel networks and their fabrication can be found, for example, in U.S. Patent App. Publication No. 2011/0091987, U.S. Pat. Nos. 8,021,614, and 8,186,913, each of which is disclosed herein by reference in its entirety.

In some implementations, a microfluidic system includes components for preparing a particle carrying fluid sample prior to introducing the fluid into a particle shifting area. For instance,FIG.8is a schematic that illustrates an example of a microfluidic system800that includes a particle focusing area801(labeled “Concentrating units”), similar to the particle focusing area shown inFIG.5that relies on inertial focusing and siphoning/fluid extraction for increasing particle to fluid concentration and/or for obtaining a low particle concentration fluid. The system800additionally includes a filter section803(labeled “Filter”) and a particle focusing section805(labeled “Focusing Units”) upstream from the particle shifting area801. The filter section803includes an arrangement of multiple different-sized post structures.

Based on the arrangement of the structures, the filter section803is configured to filter particles contained in an incoming fluid according to the particle size (e.g., average diameter), such that only particles of a pre-defined size or less are able to pass to the next stage of the system800. For instance, for complex matrices, such as bone marrow aspirate, the filter section803may be configured to remove bone chips and fibrin clots to improve the efficiency of enhancing concentration downstream. In an example arrangement, the filter section803may include an array of posts having a pillar size and array offset designed to deflect particles above 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 later stages of the system800. For example, the filter803may be configured to filter/block passage of particles that have an average diameter greater than 50%, greater than 60%, greater than 70%, greater than 80% or greater than 90% of the minimum width of a channel in the particle shifting area801.

The filter section803is fluidly coupled to the particle focusing section805. The particle focusing section805is configured to pre-focus particles exiting the filter section803to a desired fluid streamline position, before the particles are provided to the particle shifting area801. An advantage of pre-focusing the particles is that it reduces the distribution of particles across the channel width to a narrow lateral extent. The focused line of particles then can be repositioned so that the probability of the particles inadvertently entering the wrong channel (e.g., the channel for obtaining “filtered” fluid in the particle shifting area801) is reduced. Pre-focusing can be achieved using inertial focusing techniques. Further details of inertial focusing are described above in the section entitled “Particle Shifting Using Inertial Focusing.”

Once the particle to fluid concentration has been increased in the particle shifting area801, the “filtered” fluid and/or the particles may be coupled to a separate processing region of the microfluidic system800or removed from the system800for additional processing and/or analysis. For example, the second channel of the particle shifting area801is coupled to a first outlet807, whereas the first channel of the particle shifting area801is coupled to a second outlet809.

External Forces

Other functionality may be added to the microfluidic system to enhance the focusing, concentrating, separating, and/or mixing of particles. For instance, in some implementations, additional forces may be introduced which result in target specific modification of particle flow. The additional force may include, for example, magnetic forces, acoustic forces, gravitational/centrifugal forces, electrical forces, and/or inertial forces.

FIGS.9A-9Care schematics illustrating three different examples of microfluidic devices that rely on magnetophoresis used together with the particle shifting techniques described herein to focus different types of particles along different corresponding streamlines within a microfluidic device. In general, magnetophoresis employs high magnetic field gradients for sorting magnetically labeled particles flowing within a microfluidic channel of a device. The magnetic field gradients are produced by placing one or more magnets adjacent to the microfluidic channel, in which the configuration of the magnets gives rise to a magnetic flux gradient profile that extends across the microfluidic channel. The magnetically labeled particles are subsequently “pulled” by the gradient. Depending on the positioning of the gradient profile, the magnetically labeled particles can be focused to one or more desired positions within the microfluidic channels. Further details on the application of magnetophoresis to microfluidic devices can be found, for example, in WO 2014/004577, incorporated herein by reference in its entirety.

In the first example shown inFIG.9A, a microfluidic device900aincludes a particle shifting area901fluidly coupled to magnetophoresis area703. The particle shifting area (labeled “Focusing” inFIG.9A)901is constructed in a similar manner as the device100shown inFIG.1. Briefly, the focusing area901includes two separate fluid flow regions: a second fluid flow region and a first fluid flow region separated by a ID array of island structures, each of which is separated from an adjacent island structure by a gap. As fluid propagates through the first flow region, a portion of the fluid is extracted into the second flow region, while an inertial lift force is exerted on the particles, which keeps the particles traveling within the first flow region. Of course, other forces (such as inertial focusing) may be used in addition or as an alternative to keep particles within the first fluid flow region. Both the second and first fluid flow regions of the particle shifting area are fluidly coupled into the magnetophoresis area903, which is void of island structures.

The magnetophoresis area903is configured to include a magnetic field gradient that extends across the microfluidic channel. For example, the microfluidic device900amay include one or more magnets907adjacent to the magnetophoresis area903, in which the magnets907create the magnetic field gradient. For ease of illustration, the magnets907are shown at the bottom of the page to indicate their position relative to the microfluidic devices (900a,900b, and900c) along the longitudinal direction of fluid flow. However, it should be understood that in operation, the magnets907are more likely to be positioned above and/or below the fluidic channel in the magnetophoresis area903(i.e., along the x-axis inFIGS.9A-9C) of each of the devices900a,900band900c.

Referring again toFIG.9A, two different types of particles are included in the fluid introduced into the focusing area901. A first type of particle may include a desired analyte (e.g., a cell, platelet, or bacteria) that is bound to a magnetic marker such as a magnetic bead. The second type of particle may include a second analyte that has no substantial magnetic component. As the two different types of particles pass through the focusing area901, the particles are concentrated in the first fluid flow region and are focused along a fluid streamline. The focused particles then pass into the magnetophoresis area903, where the magnetic field gradient exerts a force on the particles bound to the magnetic beads. The force generated by the interaction of the field gradient with the magnetic beads causes the magnetically labeled particles to deviate from the propagation direction of the original fluid streamline. In particular, the magnetically labeled particles follow the magnetic gradient and form a new stream of particles. The direction of the magnetic gradient, and thus the path that the magnetically labeled particles follow may depend on the orientation and arrangement of the magnets907near the magnetophoresis area903. The two different streams of particles, i.e., a stream containing magnetically labeled particles and a stream of non-magnetically labeled particles, then may be separately collected at an output of the magnetophoresis area903(referred to as “labeled particles” and “unlabeled particles” inFIG.9A).

In the second example shown inFIG.9B, the particle shifting area is constructed in a similar manner as the device200inFIG.2. Again, a fluid containing a first type of particle that is bound to a magnetic marker and a second type of particle that has no substantial magnetic component is introduced into the focusing area901. The fluid shifting and inertial lift forces (or, e.g., inertial focusing forces) focus both types of particles within a first fluid flow region between two arrays of island structures. The focused particles then exit the particle shifting area and are fluidly coupled into the microfluidic channel of the magnetophoresis area903. Once the particles enter the magnetophoresis area903, the magnetic field gradient generated by the magnets907exerts a force on the magnetically labeled particles, causing them to diverge from the propagation direction of the original focused stream. In the example ofFIG.9B, the stream of particles flowing from the focusing area901include a first set of magnetically labeled particles, a second set of magnetically labeled particles, and a third set of non-labeled particles. As shown inFIG.9B, the gradient is arranged such that the magnetically labeled particles are deflected either to the top or bottom of the channel, whereas the non-labeled particles continue to follow their original focused trajectory through the magnetophoresis area703. Again, the labeled and unlabeled particles, once separated, may be collected at an output of the magnetophoresis area903for extraction or further analysis.

The third example shown inFIG.9Cdemonstrates sorting of particles in a manner opposite to that ofFIG.9B. The focusing area901inFIG.9Cis constructed in a similar manner to the device300shown inFIG.3. In particular, the focusing area901includes an initial island structure configured to separate an incoming fluid containing magnetically labeled and non-labeled particles into two separate channels (i.e., a second fluid channel (upper channel inFIG.9C) and a third fluid channel (lower channel inFIG.9C), where the particles are focused into streamlines. Once the focused streams of particles pass into the microfluidic channel of the magnetophoresis area903, the magnetic field gradient generated by the magnets907causes the magnetically labeled particles to diverge towards the center of the first channel (center channel inFIG.9C) and form a third focused stream. After deflection by the magnetic gradient, the second and third streams are left with unlabeled particles. Again, both the unlabeled and labeled particles, once separated, may be collected at an output of the magnetophoresis area903for extraction or further analysis.

While the examples shown inFIGS.9A-9Cperform the focusing and magnetic separation of particles in separate stages, such functions can be performed in a single stage.FIGS.9D-9Fare schematics illustrating three different examples of microfluidic devices (900d,900e,900f) that rely on the use of magnetophoresis with the particle shifting techniques described herein to focus different types of particles along different corresponding streamlines in a single stage. Again, the microfluidic devices900include one or more magnets907to create the magnetic field gradient. The magnets907inFIGS.9D-9Fare shown at the bottom of the page to indicate their position relative to the microfluidic devices (900d,900e, and900f) along the longitudinal direction of fluid flow. However, it should be understood that in operation, the magnets907are more likely to be positioned above and/or below the fluidic channel in the magnetophoresis area903(i.e., along the x-axis inFIGS.9D-9F) of each of the devices900d,900e, and900f.

Referring toFIG.9D, the focusing area is constructed in a similar manner as the device100shown inFIG.1. That is, the focusing area includes a second microfluidic channel separated from a first microfluidic channel by an array of island structures. In contrast toFIGS.9A-9C, the magnetic field gradient from the magnets907extends across both the second and first fluid flow regions of the focusing area. When a fluid containing both magnetically labeled particles and unlabeled particles is introduced into the particle shifting area, the particles are initially constrained within the first microfluidic channel due to inertial lift forces. However, the magnetically labeled particles may experience a force (depending on the arrangement of the magnetic field gradient) from the magnetic field that overcomes the inertial lift force. In certain implementations, the magnetically generated force may cause the labeled particles to diverge from the stream of unlabeled particles and pass through openings between the island structures.

FIGS.9E-9Fare schematics illustrating alternative configurations of microfluidic devices that combine particle shifting areas with magnetophoresis. Similar to the example ofFIG.9D, the examples shown inFIGS.9E-9Fillustrate how a magnetic field gradient can cause magnetically labeled particles to diverge from an initially focused stream of particles and form new focused particles streams. InFIG.9E, magnetically labeled particles are deflected through openings between island structures to a second (upper channel inFIG.9E) and third (lower channel inFIG.9E) microfluidic channel, whereas a focused stream of non-labeled particles remain within a first (center channel inFIG.9E) microfluidic channel that is located between the two arrays of island structures. InFIG.9F, the inertial lift forces near the island structures maintain the non-labeled particles along focused streams within a second (upper channel inFIG.9F) and third (lower channel inFIG.9F) microfluidic channel. In contrast, a magnetic field gradient generated by the magnets907causes magnetically labeled particles to pass through openings in the island structures into a center microfluidic channel that is located between the second and third microfluidic channels.

The magnetic markers used for labeling particles can include spherical bead-like materials having one or more inner magnetic cores and an outer coating, e.g., a capping polymer. The magnetic cores can be monometallic (e.g., Fe, Ni, Co), bimetallic (e.g., FePt, SmCo, FePd, and FeAu) or can be made of ferrites (e.g., Fe2O3, Fe3O4, MnFe2O4, NiFe2O4, CoFe2O4). The magnetic markers can be nanometers or micrometers in size, and can be diamagnetic, ferromagnetic, paramagnetic, or superparamagnetic, in which size corresponds to an average diameter or average length. For example, the magnetic markers can have a size of approximately 1 μm, approximately 600 nm, approximately 500 nm, approximately 300 nm, approximately 280 nm, approximately 160 nm, or approximately 100 nm. Other marker sizes are possible as well. The outer coating of a marker can increase its water-solubility and stability and also can provide sites for further surface treatment with binding moieties. The magnetic markers each have a magnetic moment in the range of about 1 KA/m to about 100 kA/m. For example, in some implementations, the magnetic markers have a magnetic moment of about 35 kA/m

In general, the magnetic markers may be bound to target analytes in a fluid using binding moieties. A binding moiety is a molecule, synthetic or natural, that specifically binds or otherwise links to, e.g., covalently or non-covalently binds to or hybridizes with, a target molecule, or with another binding moiety (or, in certain embodiments, with an aggregation inducing molecule). For example, the binding moiety can be a synthetic oligonucleotide that hybridizes to a specific complementary nucleic acid target. The binding moiety can also be an antibody directed toward an antigen or any protein-protein interaction. Also, the binding moiety can be a polysaccharide that binds to a corresponding target. In certain embodiments, the binding moieties can be designed or selected to serve, when bound to another binding moiety, as substrates for a target molecule such as enzyme in solution. Binding moieties include, for example, oligonucleotides, polypeptides, antibodies, and polysaccharides. As an example, streptavidin has four sites (binding moieties) per molecule that will be recognized by biotin. For any given analyte, e.g., a specific type of cell having a specific surface marker, there are typically many binding moieties that are known to those of skill in the relevant fields.

For example, certain labeling methods and binding moiety techniques are discussed in detail in U.S. Pat. No. 6,540,896 entitled, “Microfabricated Cell Sorter for Chemical and Biological Materials” filed on May 21, 1999; U.S. Pat. No. 5,968,820 entitled, “Method for Magnetically Separating Cells into Fractionated Flow Streams” filed on Feb. 26, 1997; and U.S. Pat. No. 6,767,706 entitled, “Integrated Active Flux Microfluidic Devices and Methods” filed on Jun. 5, 2001.

The surface of the magnetic markers can be treated to present functional groups (e.g., —NH2, —COOH, —HS, —CnH2n-2) that can be used as linkers to subsequently attach the magnetic markers to the target analytes (e.g., antibodies, drugs). In some cases, the surface treatment makes the magnetic markers essentially hydrophilic or hydrophobic. The surface treatment can include the use of polymers including, but not limited to, synthetic polymers such as polyethylene glycol or silane, natural polymers, derivatives of either synthetic or natural polymers, and combinations thereof.

In some implementations, the surface treatment does not result in a continuous film around the magnetic marker, but results in a “mesh” or “cloud” of extended polymer chains attached to and surrounding the magnetic marker. Exemplary polymers include, but are not limited to, polysaccharides and derivatives, such as dextran, pullanan, carboxydextran, carboxmethyl dextran, and/or reduced carboxymethyl dextran, PMMA polymers and polyvinyl alcohol polymers. In some implementations, these polymer coatings provide a surface to which targeting moieties and/or binding groups can bind much easier than to the marker. For example, in some embodiments magnetic markers (e.g., iron oxide nanoparticles) are covered with a layer of 10 kDa dextran and then cross-linked with epichlorohydrin to stabilize the coating and form cross-linked magnetic markers.

Additional information on the fabrication, modification, and use of magnetic markers can be found, for example, in PCT Pub. No. WO/2000/061191, U.S. Patent App. Pub. No. 20030124194, U.S. Patent App. Pub. No. 20030092029, and U.S. Patent App. Pub. No. 20060269965, each of which is incorporated herein by reference in its entirety.

Fabrication of Microfluidic Devices

A process for fabricating a microfluidic device according to the present disclosure is set forth as follows. A substrate layer is first provided. The substrate layer can include, e.g., glass, plastic or silicon wafer. An optional thin film layer (e.g., SiO2) can be formed on a surface of the substrate layer using, for example, thermal or electron beam deposition. The substrate and optional thin film layer provide a base on which microfluidic regions may be formed. The thickness of the substrate can fall within the range of approximately 500 μm to approximately 10 mm. For example, the thickness of the substrate210can be 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 possible as well.

After providing the substrate layer, the microfluidic channels formed above the substrate layer. The microfluidic channels include the different fluid flow pathways of the particle shifting area, as well as the other microfluidic components of the system, including any filtering sections, inertial focusing sections, and magnetophoresis sections. Microfluidic channels for other processing and analysis components of a microfluidic device also may be used. The microfluidic channels and cover are formed by depositing a polymer (e.g., polydimethylsiloxane (PDMS), polymethylmethacrylate (PMMA), polycarbonate (PC), or cyclo olefin polymer (COP)) in a mold that defines the fluidic channel regions. The polymer, once cured, then is transferred and bonded to a surface of the substrate layer. For example, PDMS can be first poured into a mold (e.g., an SU-8 mold fabricated with two step photolithography (MicroChem)) that defines the microfluidic network of channels. The PDMS then is cured (e.g., heating at 65° C. for about 3 hours). Prior to transferring the solid PDMS structure to the device, the surface of the substrate layer is treated with O2plasma to enhance bonding. Alternatively, the microfluidic channels and cover can be fabricated in other materials such as glass or silicon.

APPLICATIONS

The new microfluidic techniques and devices described herein can be used in various different applications.

Centrifugation Replacement

The particle shifting techniques and devices disclosed herein can be used as replacements for centrifugation. In general, centrifugation is understood to include the concentrating of sub-components within a fluid through the application of centrifugal forces to the fluid. Typically, this process requires devices that have moving parts, which are prone to wear and breakage. Moreover, the moving parts require complex and costly fabrication processes. Another problem with centrifugation is that it is a process typically applied in a closed system, i.e., centrifugation requires manually transferring samples to and from a centrifuge.

In contrast, the presently disclosed techniques are capable of substantially increasing the concentration of fluid components using relatively simple micro-structures without the need for moving parts. The techniques can be implemented as part of a single open microfluidic system, such that fluid samples may be transferred to or from the particle shifting area without manual interference. Additionally, particle shifting can be extended to devices requiring large throughput (i.e., volume rate of fluid that can be processed). For example, the devices disclosed herein may be configured to enable up to 10, 25, 50, 75, 100, 250, 500, 1000, 5000, or 10000 μl/min of fluid flow. Other flow rates are also possible. For instance, using device100inFIG.1as an example, if the second and first microfluidic channels106,108have depths of approximately 50 μm and widths of approximately 50 μm, the device100may be capable of achieving a combined sample flow rate of up to about 5 mL/min. Varying the channel sizes may alter the maximum volumetric flow rate of which the device is capable. Furthermore, multiplexing multiple channels may enable even higher rates of flow. Thus, in certain implementations, the particle shifting techniques may provide substantial cost and time saving advantages over traditional centrifugation processes. Examples of applications where a microfluidic replacement for a centrifuge device may be useful include bone marrow and urine analysis.

Detecting Infectious Agents

In addition, the particle shifting techniques disclosed herein can be used as part of a research platform to study analytes of interest (e.g., proteins, cells, bacteria, pathogens, and DNA) or as part of a diagnostic assay for diagnosing potential disease states or infectious agents in a patient. By separating and focusing particles within a fluid sample, the microfluidic device described herein may be used to measure many different biological targets, including small molecules, proteins, nucleic acids, pathogens, and cancer cells. Further examples are described below.

Rare Cell Detection

The microfluidic device and methods described herein may be used to detect rare cells, such as circulating tumor cells (CTC) in a blood sample or fetal cells in blood samples of pregnant females. For example, the concentration of primary tumor cells or CTCs can be enhanced in a blood sample for rapid and comprehensive profiling of cancers. By combining the particle deflection techniques described herein with magnetophoresis (seeFIG.7), different types of cells can be detected (e.g., circulating endothelial cells for heart disease). Thus, the microfluidic device may be used as a powerful diagnostic and prognostic tool. The targeted and detected cells could be cancer cells, stem cells, immune cells, white blood cells or other cells including, for example, circulating endothelial cells (using an antibody to an epithelial cell surface marker, e.g., the Epithelial Cell Adhesion Molecule (EpCAM)), or circulating tumor cells (using an antibody to a cancer cell surface marker, e.g., the Melanoma Cell Adhesion molecule (CD146)). The systems and methods also can be used to detect small molecules, proteins, nucleic acids, or pathogens.

Fluid Exchange

The microfluidic device and methods described herein may be used to shift cells from one carrier fluid to another carrier fluid. For example, the particle shifting techniques disclosed could be used to shift cells into or out of a fluid stream containing reagents, such as drugs, antibodies, cellular stains, magnetic beads, cryoprotectants, lysing reagents, and/or other other analytes.

A single particle shifting region could contain many parallel fluid streams (from many inlets) through which a shifted cell would pass. For example, white blood cells could be shifted from a blood stream into a stream containing staining reagents and then into a buffer stream.

In bioprocessing and related fields, the devices and techniques described may be used to enable sterile, continuous transfer of cells from old media (containing waste products) into fresh growth media. Similarly, extracellular fluids and cellular products (e.g., antibodies, proteins, sugars, lipids, biopharmaceuticals, alcohols, and various chemicals) may be extracted from a bioreactor in a sterile, continuous manner while cells are retained within the bioreactor.

Fluid Sterilization and Cleansing

The microfluidic device microfluidic device and methods described herein may be used to remove pathogens, pollutants, and other particular contaminants from fluids. By shifting contaminants across fluid streamlines, contaminants may be removed from a fluid sample and collected as a separate waste stream.

Harvesting Algae for Biofuels

Harvesting algae from growth media is a major expense in the production of biofuels because algae grow in very dilute suspensions at near neutral buoyancy, making efficient extraction and concentration of algal biomass difficult. The microfluidic device and methods described herein can provide an efficient means of harvesting algae that does not depend on either density or filtration. The devices and techniques described enable the algae in a growth tank to be extracted from the growth media and concentrated to a high volume density. This can be done either as a single step or as part of a continuous process. Additionally, because the devices described herein can sort cells in a size-dependent manner, they can be designed to sort and concentrate only the larger algae that have reached maturity, returning smaller, immature algae to the growth tank.

EXAMPLES

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

Device Fabrication

Various experiments were performed to analyze the behavior of microfluidic devices having asymmetrically curved channels (see, e.g., the section above entitled “Increasing Particle Concentration/Reducing Fluid Volume” and the device shown inFIG.5) that combine inertial focusing with fluid extraction to achieve volume reduction of particle-rich fluid samples. That is, the devices included a focusing channel (see, e.g., channel508inFIG.5) in which particles were focused using inertial focusing techniques and a particle-free channel/second fluid flow channel (see, e.g., channel506inFIG.5) into which fluid from the focusing channel was extracted. The experiments are described in Examples 1 to 5 below. The devices used in those examples were designed and fabricated as follows.

For each microfluidic device, standard SU8 photolithography and soft lithography techniques were used to fabricate the master mold and the PDMS microchannels, respectively. Briefly, negative photoresist SU8-50 (Microchem Corp, Massachusetts) was spun at 2850 RPM to a thickness of approximately 50 μm, exposed to ultraviolet light through a mylar emulsion printed photomask (Fineline Imaging, Colorado) that defines the microfluidic network of channels, and developed in BTS-220 SU8-Developer (J. T. Baker, New Jersey) to form a raised mold. A 10:1 ratio mixture of Sylgard 184 Elastomer base and curing agent (Dow Corning, Michigan) was then poured over the raised mold, allowed to cure in an oven at 65° C. for 8 hours and then removed from the SU8 master mold to form the microfluidic device cover having the patterned channels. Inlet and outlet holes to the channels were punched using custom sharpened needle tips. The devices were then cleaned of particulate using low-residue tape and oxygen plasma bonded to pre-cleaned 1 mm thick glass microscope slides.

For experiments where high pressure deformation of PDMS was a concern, epoxy devices were used instead. Epoxy devices were constructed using PDMS molds created by treating PDMS channels with tridecafluoro-1,1-2,2-tetrahydrooctyl-trichlorosilane (Gelest) and then pouring PDMS over the silanized channels. After 24 hours of curing at 65° C., the molds were carefully separated from the silanized channels. Holes were punched into PDMS molds at the inlets and outlets using a 0.75 mm diameter Harris Uni-Core biopsy punch. Teflon coated wire (0.028 inch diameter, McMaster-Carr) was inserted gently into these holes as to not deform the surface of the PDMS mold. Tygon tubing (0.02″ I.D., 0.06″ O.D.) was then guided onto teflon coated wire and suspended ˜1 mm from the mold surface. Epoxacast 690 (Smooth-On) was mixed and degassed for 30 minutes prior to pouring into the mold. At the same time as molds were filled, slides were coated with epoxy by laying a glass slide on a drop of epoxy atop a flat PDMS surface. After ˜28 hours, the devices were cooled temporarily to −22° C. to prevent deformation, the Teflon wire was removed and devices removed from the molds. Then the glass slides were removed from the PDMS slabs and heated to 55° C. and devices were pressed against slides ensuring bonding.

Particle and Cell Suspensions

The devices used in the Examples described below were tested over a wide range of flow conditions using fluorescent polystyrene beads and white blood cells as exemplar particles. Polystyrene particle suspensions were created using 4.4 μm diameter blue-fluorescent beads (Polysciences), 9.9 μm diameter green-fluorescent beads (ThermoFisher Scientific) and 15 μm diameter red-fluorescent beads (Invitrogen). Each was suspended to a final length fraction of 0.1 in an equivalent density solution (1.05 g/mL) of 1×PBS, 0.1% Tween20, and iodixanol. White blood cells (buffy coat) were isolated using deterministic lateral displacement with a co-flow of buffer solution.

Fluorescent Counting and Cell Counting

Fluorescent and high resolution imaging of fluid samples were accomplished using an automated Nikon TiE inverted microscope with a Retiga 2000R monochromatic camera as well as a Vision Research Phantom v4.2 high speed monochromatic camera.

Hemocytometers and Nageotte chambers were utilized for measuring particle concentrations in white blood cell yield experiments at dilutions dependent upon the output cell concentrations.

Example 1: Cell Free Layer Growth and Siphon Percentage

The combined siphoning and inertial focusing design takes advantages of fast-acting inertial forces, which generate a particle-free layer near the walls of the microfluidic channel. This particle-free fluid layer then is controllably siphoned off leaving the particles once again closer to the walls where the inertial forces are strongest. The process of focusing and siphoning may be repeated until a desired volume reduction is achieved. When using a microfluidic device to enhance the concentration of particles within a fluid or to extract a particle-free fluid, an important design consideration may include controlling the percentage of fluid that is siphoned relative to the dynamics of the formation of the particle-free layer. In inertial focusing systems, the focusing behavior is a cumulative result of numerous parameters including the channel geometry as well as flow speed (See, e.g., Di Carlo, D. “Inertial microfluidics,” Lab Chip 9, 3038 (2009) and Martel, J. & Toner, M. “Inertial Focusing in Microfluidics,” Annual Review of Biomedical Engineering 16, 371-396 (2014), incorporated herein by reference in their entirety). For instance, curved structures are generally more efficient than planar structures at achieving focusing over a given channel length while in some implementations are also more sensitive to changes in flow speed.

Using asymmetrically curved structures similar to the structures described with respect toFIGS.5-6, we characterized the formation of a particle-free layers for a range of focusing channel widths (between 50 μm to 200 μm) and over a range of flow rates (between 10 μL/min and 3000 μL/min) depending on the channel width. Each of the devices tested included a series of five focusing-siphoning unit pairs (see, e.g.,FIG.6) followed by an expansion into a 500 μm wide straight section. The particle-free layer width of the resulting output fluid was measured downstream of the focusing units after the channel had fully expanded based on a 10% relative intensity threshold across the channel width (i.e., the intensity is normalized to between 0% to 100%, after which the position at which the intensity reaches 10% is identified. See, e.g., Martel, J. M. & Toner, M. “Particle Focusing in Curved Microfluidic Channels,” Sci. Rep. 3, 1-8 (2013), incorporated herein by reference in its entirety).

The width of the particle-free layer at the optimal flow rate for each channel width was compared to one another as shown inFIG.10. Specifically,FIG.10shows the “cell-free fraction” versus flow rate, in which each of the data points represents the maximum fraction of the fluid (as measured across the channel width) that is free of particles for each different sized channel. The legend beneath the plot indicates the channel widths. As is evident from the graph shown inFIG.10, the narrower channels achieve significantly higher maximum particle-free layer width than the wider channels (50 μm wide-38%, 75 μm-46%, 100 μm-42%, 125 μm-30%, 150 μm-15%, 200 μm-13%). The variation in particle-free layer width over a range+/−50% of the optimal flow rate (flow rate which achieves the maximum particle-free layer width) was lower for the wider channels (50 μm wide-12%, 75 μm-23%, 100 μm-16%, 125 μm-15%, 150 μm-4.6%, 200 μm-5.5%).

Using the reference data we determined that there was a nearly linear relationship between the optimal flow rate, QOptimal(i.e., the flow rate that resulted in the greatest width for particle-free layer formation), and the focusing unit width, wfocus=1.0911e−07*QOptimal(μL/min)+4.4789E−05m. Based on the foregoing relationship, it is possible to create a device that maintains a high level of particle-free layer formation efficiency as fluid is siphoned from the focusing channel and as the flow rate through the focusing channels decreases.

The relationship between the formation of the particle-free layer and a maximum siphon percentage was also studied. The siphon percentage is the percentage of flow in the focusing channel that is siphoned out at the next opening between islands. The amount siphoned is determined by the relative fluidic resistances of the focusing and siphon channels. In particular, a set of devices was designed using a range of siphon percentages (7%, 10%, 12% and 15%) for a fixed input flow rate of 500 μL/min. The flow rate of 500 μL/min was chosen to be within the optimal flow rate range of the narrower more efficient focusing unit widths. A comparison of the focusing performance of these devices indicates that, depending on the volume reduction factor desired, the siphon percentage must be below 10% for a factor of 10 volume reduction and 7% for a factor of 50 volume reduction. The volume reduction factor is equivalent to the concentration factor and may be expressed as one divided by the fraction of flow in the focusing channel. For example, if 5% of the total flow is in the focusing channel, the volume reduction factor is 20.FIG.11includes images of fluorescently tagged white blood cells flowing through the focusing-siphoning units of the microfluidic device, in which each image corresponds to a different siphon percentage for a factor of 10 volume reduction. As is evident from the images, the loss of particles from the focusing channel into the second fluid flow channel in the 15% siphon percentage device is quite noticeable.

As the foregoing results demonstrate, the combined siphoning and inertial focusing techniques enable the control of the volume reduction factor in a well-regulated manner. In some implementations, it may be possible to obtain a specific volume reduction factor thereby tailoring a specific sample volume for downstream molecular assays independent of the input sample volume.

For the experiments described below, we have selected two specific designs for detailed characterization. The two selected designs are a factor of 10 (“10×”) concentrator (this device included 26 focusing-siphoning unit pairs and had a 10% siphon percentage) and a factor of 50 (“50×”) concentrator (this device included 152 focusing-siphoning unit pairs and had a 7% siphon percentage).

Example 2: Flow Rate Dependence

Another factor that may be considered in a microfluidic system for performing volume reduction and/or increasing the particle concentration within a fluid is the flow speed of the fluid sample through the microfluidic device. Accordingly, the sensitivity to flow rate was also investigated. Using isolated white blood cells (buffy coat), the yields of both the 10× and 50× devices were analyzed between input flows rates of 100 μL/min and 1000 μL/min. Yield is calculated on a relative basis between the number of cells in the stream flowing in the focusing channel and the number of cells in the second fluid flow region or, alternatively, as the total number of cells in the stream flowing in the focusing channel divided by total cells in the focusing channel and the second fluid flow channel combined. A high yield of greater than 95% for the devices was maintained between 400 and 600 μL/min but beyond that the drop off in yield began to be significant. For instance, multiple separate streams containing the white blood cells began to form at 1000 μL/min.

FIG.12is a plot of relative white blood cell yield versus flow rate for both the 10× and 50× devices. In general, the system loss (e.g., due to cells lost in transfers between various containers, in tubing, etc.) comparing the input number of cells to total cells coming out of the focusing channel and the second flow channel combined was typically low, around 10%. For flow rates lower than 400 μL/min, the drop off in yield was consistent with an overall lack of focusing. For example, in the case of negligible inertial effects, one would expect a yield equivalent to the flow split, such as 10% and 2% for the 10× and 50× devices, respectively. The increase in yield by increasing the flow rate from 100 to 400 μL/min was indicative of the improvement of focusing with Reynolds number as inertial effects increase. The decrease in yield after 600 μL/min was a likely a consequence of PDMS deformation at the higher driving pressures leading to significantly different focusing patterns.

The exact range of input and output flow rates depend on the particle size and channel dimensions used. To efficiently achieve higher throughput for a given design, multiplexing of channels may be needed.

Example 3: Size Dependence

Inertial forces are strongly dependent upon the size of the particles being focused. Accordingly, the performance of the combined inertial focusing and siphoning devices were evaluated to understand the sensitivity to particle size. In particular, a variety of polystyrene particle sizes (4 μm-10 μm) were run simultaneously through the 10× and 50× devices in order to determine the size range of particles that are deflected from the focusing channel to the second fluid flow region where the “particle-free” layer was desired.FIG.13is a plot of the foregoing experiment and suggests a trend where smaller particle sizes have lower relative yields (i.e., (total cells in product)/(total cells in product+total cells in waste)) compared to larger particle sizes, i.e., the smaller a particle is, the greater the probability that the particle will escape the focusing channel through a gap between island structures. If relative yields above 90% are desired, a cutoff particle size for this threshold can be interpolated as approximately 8.5 μm for the 10× device and approximately 8 μm for the 50× device. This slight difference may be attributed to the significantly lower velocities at the end of the 50× concentrator where the focusing becomes more sensitive to particle size.

The foregoing results showing the sensitivity of the combined siphoning and inertial focusing devices to particle size may lead to several possible advantageous applications. For instance, the size dependence can be beneficial for cleanup of biological samples (e.g., removing bacteria) as particles smaller than a cutoff size will be siphoned off from the focusing channel into the second fluid flow channel, thus improving the final sample purity or decreasing undesired biological sample contamination.

Example 4: Volume Fraction Dependence

Another factor that was analyzed was the effect of inter-particle interactions on the focusing behavior. Generally, conventional inertial focusing devices have a strict requirement that the input fluid sample concentrations be low in order to achieve high quality focusing (see, e.g., Lee, W., Amini, H., Stone, H. A. & Di Carlo, D. “Dynamic self-assembly and control of microfluidic particle crystals,” Proceedings of the National Academy of Sciences 107, 22413 (2010), incorporated herein by reference in its entirety). A theoretical concentration limitation is given by the limit of a continuous line of adjacently touching particles at the equilibrium positions along the entire channel length or a length fraction of 1 (see, e.g., Di Carlo, D. “Inertial microfluidics,” Lab Chip 9, 3038 (2009), incorporated herein by reference in its entirety). We investigated the operational cutoff of the particle concentration for the 10× and 50× devices by varying the input concentration of white blood cells processed at 500 μL/min.

FIG.14is a plot illustrating the relative yield of the white blood cells at this flow rate for different input concentrations. As the plot indicates, there is a sharp maximum limit at an input concentration of approximately 1 million cells per milliliter. The particle concentration at which the particle interactions will start affecting the performance of the device threshold was reached in the devices of approximately at approximately 80M cells per milliliter. This high particle concentration may be attributable to the fact that the operational success or yield of the devices does not require that all of the particles fall on a single streamline. Instead, the cell free layer formation near the walls leads to a much higher concentration at which the yield decreases (i.e., rather than requiring all particles to pack into the limited space of a single narrow stream, we only required that particles be packed into the region of fluid that is not siphoned, which can accommodate far more particles). The foregoing experimental results indicate that the particle-free layer formation is not as sensitive to particle volume fraction as the single stream or high quality inertial focusing as previously understood (see, e.g., Di Carlo, D. “Inertial microfluidics,” Lab Chip 9, 3038 (2009), incorporated herein by reference in its entirety).

Example 5: Achieving Greater Than 50× Volume Reduction

We also analyzed the ability of the microfluidic volume reduction devices to obtain substantially high throughputs and volume reduction. For example, in some cases, large numbers of the devices shown inFIGS.1-5may be operated in parallel to increase the overall system throughput (i.e., the overall volume of fluid processed). For instance, in one possible design, multiple volume reduction devices (e.g., device100) may each have a separate fluid input to receive a fluid sample, where the output of each device is coupled to a common output channel for collecting either concentrated particles or the filtered fluid sample.

Alternatively, or in addition, two or more devices may be constructed in series so that particle concentration/volume reduction is modified at each stage (i.e., device) of the overall system. To demonstrate the application of serial volume reduction, we constructed a microfluidic system containing serially integrated devices: in particular, we used ten parallel 10× devices that feed into a single 50× device for a theoretical overall volume reduction of 500×.FIG.15is a schematic that illustrates a top view of the design of the system1500used to study volume reduction, which includes ten parallel 10× concentrator devices1502and a single 50× concentrator device1504. The operation of the system1500proceeds as follows: (i) dilute particles enter the system1500and are focused in the separate 10× concentrators1502into ten parallel focused streams; (ii) the ten parallel focused streams then are sent through a series of converging channels1506; (iii) the converged streams then are refocused as they enter the 50× device1504; and (iv) finally, all the particles exit through the bottom product outlet of the 50× device.

Due to the pressure requirements and PDMS deformation, the systems used for the experiments were fabricated in rigid epoxy in place of PDMS [Eugene J. Lim et al. “Inertio-elastic focusing of bioparticles in microchannels at high throughput,” Nature Communications.2014] (see, e.g., Martel, J. M. & Toner, M. “Particle Focusing in Curved Microfluidic Channels,” Sci. Rep. 3, 1-8 (2013), incorporated herein by reference in its entirety). To test the yield, white blood cells at an input concentration of 100,000 per mL were introduced into the system. The yield of the integrated system was consistently above 95% and exhibited a volume reduction factor of ˜411. Thus, for a 30 mL input sample containing 100,000 white blood cells per mL, the sample will be reduced by the microfluidic system into 73 μL+/−1.2 μL (n=5) with greater than 95% of the original cells (95.7%+/−3.6%, n=5). The discrepancy between the 411 volume reduction factor and 500 designed volume reduction factor is a difference of only a few microliters of product which was difficult to control as the input flow rate of 4 mL/min (pump driving force limitation) and the product flow rate of <10 μL/min. That is to say, that while the device was designed to perform 500× volume reduction, it actually performed 400× volume reduction. It is believed that the relative resistances of the product and waste channels were slightly off, such that slightly more volume went to the product than desired. Additionally, the tiny product volume may have caused some measurement error. Tiny fabrication imperfections in the microfluidic system can alter this balance as well.

Centrifugation used for washing cells, exchanging media and/or concentrating a sample for subsequent assays is one of the most widely utilized processes in the biomedical sciences. The system1500and the foregoing experimental results demonstrate that the microfluidic siphoning and inertial focusing devices are capable of accomplishing the foregoing common biomedical tasks typically performed with centrifugation in a continuous flow and sterile manner at throughputs of up to 4 mL/min (240 mL/hour) and at volume reduction factors of 20-fold or higher. Furthermore, the typical limitation on throughput of microfluidic devices is also mitigated using the combined siphoning and inertial focusing techniques. While we have presented a non-integrated single device which achieves a throughput of 500 μL/min at a volume reduction factor of 50×, the devices can be further arranged in parallel to obtain a set of greater than 40 channels (20 mL/min or 1200 mL/hr), diminishing the run time for the larger volume samples.

While much of the advancement presented is in terms of improving experimental methods there has also been a key finding about the nature of inertial focusing. The realization that the particle-free layer formation is not as sensitive to particle volume fraction as the single stream or high quality inertial focusing previously predicted may be intuitive, but also brings to light a new means of comparing inertial focusing device performance. There are typically five different geometries utilized in inertial focusing and typically are each compared by the length required to achieve a minimum streak width. By changing the definition of optimal focusing from minimizing streak width to the dynamic formation of the particle-free layer, new insights into the dynamics of focusing for different microfluidic structures can be investigated and directly compared. This new means of comparison could standardize how the effectiveness of this class of microfluidic devices is measured.

OTHER EMBODIMENTS

It is to 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.