CROSS-REFERENCE TO RELATED APPLICATION
The present application claims the benefit under 35 U.S.C. § 119(e) of U.S. Provisional Patent Application No. 60/615,525, filed Sep. 30, 2004, the disclosure of which is incorporated herein by reference.
BACKGROUND OF THE INVENTION
1. Field of the Invention
Embodiments of the present invention relate generally to microfluidic operations and deformable polymer membranes.
2. Background Information
Microfluidic components for performing a variety of operations are integral parts of micro-total analysis system applications. For example, cell sorters have become a vital component in micro total analysis systems aiming to investigate biological events at the single cell level. However it has not been easy to integrate different microfluidic components together into a single chip. This has been due to the different and sometimes difficult fabrication requirements for each of the microfluidic components. For example, pumping in micro total analysis is generally achieved using external devices such as syringes or peristaltic pumps or using voltages across the channels generating electrokinetic or electroosmotic flow.
Essential processes such as bonding, aligning, clamping and interconnections for realizing a micro total analysis system generally cause significant device failure rates. Making components from the same basic unit and material facilitates the integration of operations and components. For example, polymers such as poly(dimethyl siloxane) (PDMS) can be used to fabricate various components in microfluidic devices. In addition, easy fabrication processes and simplicity of the device greatly help in integration of these components into a single device.
BRIEF DESCRIPTION OF THE FIGURES
FIGS. 1A and 1B provide schematics (a top view and a side view, respectively) of a deformable membrane operably connected to a microfluidic channel.
FIGS. 2A and 2B show exemplary designs for peristaltic pumps.
FIGS. 3A, 3B, and 3C graph measurements of fluid flow rate versus the frequency of pressure applied to the operating channels for several peristaltic pump configurations.
FIGS. 4A and 4B show exemplary designs for microfluidic sorting devices.
FIG. 5 demonstrates sorting action in an exemplary microfluidic sorting device.
FIG. 6 shows results of measurements to determine the fidelity of an exemplary sorter device as a function of the actuation times of the deformable membranes.
FIGS. 7A and 7B show an exemplary design for a microfluidic mixer. FIG. 7C shows an idealized depiction of the distribution of the liquids in the microfluidic channels before and after mixing in the device of FIGS. 7A and 7B.
FIGS. 8A, 8B, and 8C show data from video microscopy of microfluidic channels containing fluorescent microbeads passing through an exemplary active mixer unit.
FIGS. 9A, 9B, and 9C show an exemplary design for a device that can be used to mix fluids within a microfluidic channel.
FIGS. 10A and 10B show exemplary designs for devices that can be used to break and coalesce droplets of fluid.
FIG. 11 shows an exemplary design for a device that can be used to introduce precise amounts of sample into a carrier fluid.
FIG. 12 shows the operation of an exemplary device for dispensing a substance into a microfluidic channel.
FIGS. 13A, 13B, and 13C show exemplary devices for piezoelectric membrane actuation.
FIG. 14 shows a diagram of a system used for deflecting and pulsing membranes.
FIG. 15 shows a method for fabricating a microfluidic device using single layer soft lithography.
DETAILED DESCRIPTION OF THE INVENTION
Embodiments of the present invention provide deformable polymer membranes as active components of a microfluidic system. The deformable membranes perform functions associated with the manipulation of liquids in a microfluidic channel. Because the polymer membranes are disposed in the same polymer layer as the active microfluidic channel, manufacture of the microfluidic device is simplified. Although deformable membranes have been exemplified using PDMS, the present invention is not so limited as other elasomeric polymers can be used to fabricate membranes. Using a deformable membrane unit such as that shown schematically in FIGS. 1A and 1, consisting of an active microfluidic channel, an operating channel, and a membrane separating the channels, microfluidic components functioning, for example, as pumps, sorters, and mixers can be designed and fabricated.
Referring to FIG. 1, a basic deformable membrane as part of a microfluidic device is illustrated. FIG. 1A provides a top-down view and FIG. 1B provides a side-view of the same section of a microfluidic device. A microfluidic channel 10 is formed in a solid polymer block 20. An operating channel 30 is operably connected to a membrane 40 that is formed by the intersection of the microfluidic channel 10 and the operating channel 30 within the polymer block 20. As used herein, the term operating channels refers to channels that are operably connected to the deformable membrane to allow for deformation, actuation, and or pulsing of the membrane. The polymer block housing 20 is attached to a substrate 50. Actuation, deflection, or pulsing of the membrane 40 causes a change in the flow characteristics of the fluid contained within the microchannel 10. By placing the membrane in the same polymer layer as the microfluidic channel, device fabrication is facilitated.
In embodiments of the present invention, peristaltic pumping of fluids within a microchannel is effectuated using deformable membranes and operating channels that are disposed in the same polymer layer as the active microfluidic channel. The deformable membrane unit can be actuated, for example, by pressurizing the operating channels with a gas or liquid. Peristaltic pumps were realized by placing multiple deformable membrane units (a membrane unit is a pair of membranes disposed on opposite sides of a microfluidic channel) in series along a microfluidic channel. Referring now to FIG. 2, two different designs were built to compare pumping efficiency as a function of the placement of deformable membranes along a microfluidic flow channel 60. In one example, a symmetric parallel design had membranes 70 and operating channels 80 placed symmetrically on opposite sides of the microfluidic channel 60, as shown schematically in FIG. 2A. In another example, an asymmetric alternating design had membranes 70 and operating channels 80 placed asymmetrically on each side of the fluid channel 60, as shown schematically in FIG. 2B. In the example shown in FIG. 2B, deformable membrane units 70 were staggered by 50 μm on opposite sides of the active microfluidic channel 60. Alternative dimensions are possible. The membranes on a side of the channel can be separated from each other, for example, by distances of about 200 μm to about 50 μm. Pumping was visualized using a diluted solution of 1 μm fluorescent poly(styrene) beads in water. Three different phase angles of actuation for the membranes, 60°, 90°, and 120° (corresponding to actuation patterns of (100, 110, 111, 011, 001, 000), (100, 110, 011, 001), and (101, 100, 110, 010, 011, 001), where 1 indicates the membrane is actuated (distended into the microfluidic channel) and 0 indicates the membrane is not actuated), were tested and it was found that for these designs, the 60° phase angle of actuation provided the fastest flow rate for both exemplary designs.
Several different parameters, including the external regulated pressure, frequency of actuation, microfluidic channel width, membrane thickness, channel height, and gap between air channels, were tested. Typical operating channel width was 100 μm. Flow rates were calculated by measuring the time taken for fluorescent beads to traverse through a 2.7 mm long serpentine channel. FIG. 3A shows the frequency dependence of flow rate for an exemplary parallel membrane device design (as diagrammed in FIG. 2A) for different external applied pressures. The microfluidic device in this example had a membrane thickness of 20 μm, a microfluidic channel width of 20 μm, a channel height of 100 μm, and an operating channel gap of 50 μm. As can be seen from the graph in FIG. 3A, the flow rate increases to a maximum at about 30 Hz and then drops down rapidly as frequency of actuation increases for all external pressures applied. It is believed that these results can be attributed to the spring force effect of the membrane in which, after a certain frequency, the membrane does not revert back to its original position thereby reducing the volume displacement of the fluid achieved. Also, as the external pressure applied is increased, the maximum flow rate obtained increases. It is believed that the increased external pressure applied to the membrane increases the deflection of the PDMS membrane thereby increasing the volume of the fluid displaced.
FIG. 3B shows the flow rate dependence at different frequencies of actuation for two exemplary parallel design devices (as diagrammed in FIG. 2A) that had microfluidic channel widths of 20 μm and 30 μm. The devices had membrane thicknesses of 20 μm, channel heights of 100 μm, and operating channel gaps of 50 μm. The pressure applied to the operating channels was 50 psi. As seen from the graph in FIG. 3B, the trend of the flow rate dependence on the frequency of actuation is the same while the maximum flow rate obtained using the 30 μm width channel is higher than that of the 20 μm.
The results acquired from two exemplary designs for deformable membrane unit placement in a peristaltic pump (as shown in FIGS. 2A and B) are shown in FIG. 3C. In both cases, the dimensions of both the microfluidic and the operating channels were the same and the pressure applied was 30 psi. The membrane thickness was 20 μm, the microfluidic channel width was 30 μm, the microfluidic channel height was 100 μm, and the operating channel gap was 50 μm. As seen from the graph in FIG. 3C, the alternating design provided about twice the maximum flow rate of the parallel design. It was also found that the alternating design example prototype was only better at the higher pressure of 30 psi while the parallel design performed slightly better than the alternating design at pressures of 10 and 20 psi. It is believed that this observed enhancement can be attributed to the fact that in the alternating design, the deflection of the membrane is higher because it is not as constrained by the membrane on the other side of the microfluidic channel.
By controlling the various parameters of actuation and dimensions of the components of the basic deformable membrane unit, it is possible to control the flow velocities and rates. In general, channel aspect ratios of about 1:2 to about 1:10 (width to height) and widths of about 10 to about 100 μm have been used in embodiments of the present invention. Additionally, in general, average membrane thicknesses of about 5 to about 50 μm and distances between membranes located on a side of a channel of about 50 to about 200 μm can be used in embodiments of the present invention. The height and the width of the membranes are typically determined by the dimensions of the intersection of the microchannels that form the membranes which in turn are user-defined variables.
Referring now to FIG. 4, the placement of operating channels in several exemplary microfluidic sorting devices is diagrammed. In this example, deformable membranes 90 were placed along the main inlet microfluidic channel 100 or along each of the branch outlet microfluidic channels 110. The sample flow was hydrodynamically focused in the main microfluidic channel by using sheath flows from intersecting sheath microfluidic channels 120 on either side of the sample solution inlet channel 100. Two exemplary designs are shown: in FIG. 4A deformable membranes 90 are placed alongside the main inlet microfluidic channel 100, and in FIG. 4B deformable membranes 90 are placed alongside each of the branch microfluidic channels 110 (the branch channels are labeled “outlet to bin 1” and “outlet to bin 2” in FIG. 4B). In one embodiment, the membrane units 90 were activated by increasing air pressure in the operating channel 130 and causing the membrane 90 to deflect into the microfluidic channel 100 or 110. A diluted sample solution of 6 μm fluorescent poly(styrene) beads was hydrodynamically focused using branch sheath flows of DI water from either side of the main channel. To sort particles contained in a flow in the main microfluidic channel in an exemplary device having deformable membrane units placed alongside the main microfluidic channel, either the left or the right membrane is deflected into the channel to guide the microfluidic stream into the left or the right outlet channel, respectively. It was found that placement of the deformable membrane unit in the main microfluidic channel far from the Y-branch results in poorer sorting fidelity because of recovery of the laminar streams before reaching the Y-branch. Additionally, a sorting device may also optionally comprise a device for interrogating the sample stream and providing input to the switcher that activates the pressure in the operating channels. For example, the device for interrogating may be a UV-vis, fluorescence, or Raman detector that detects the presence of a cell, a virus, a bacterium, a label molecule, or a nanoparticle. When the detector detects a species of interest, it communicates to the switcher to direct the species into a selected outlet channel. For example, when the light intensity is above a certain threshold from a CCD camera used to detect a nanoparticle, the above-threshold signal can be converted using an algorithm to provide a voltage to the solenoid valves and cause a switcher to activate pressure in the operating channels.
In an exemplary design, the deformable membrane units were placed in the branch channels that when actuated would increase the resistance to flow in the respective branch thereby diverting the direction of flow of the sample to the other branch. FIG. 5 shows 6 μm beads sorted by using the deformable membrane units placed in a branch channel. The exemplary device pictured in FIG. 5 had a microfluidic channel width of 100 μm, a channel height of 100 μm, a membrane thickness of 20 μm. In the left-hand picture (FIG. 5A), actuation of the deformable membranes in the right branch outlet directs the 6 μm bead to the left branch outlet. In the right-hand picture (FIG. 5B), actuation of the deformable membranes in the left branch outlet directs the 6 μm bead to the right branch outlet. This exemplary device design worked with approximately 100% fidelity for hydrodynamically focused beads, that is to say, approximately 100% of the beads went to the right branch when the deformable membrane on the left branch was actuated and vice versa.
Referring now to FIG. 6, the time of actuation of the deformable membrane units along each of the branch channels for the device shown in FIG. 5 was varied and the number of beads that flowed into each of the branches in a minute of hydrodynamically focused sample flow was counted. FIG. 6 shows that the sorting ratios at different ratios of times of actuation for the deformable membrane units. The actuation times of the deformable membrane units in the right and left branch channels were varied and the number of beads flowing through each branch was measured using video microscopy. Data was also collected when the deformable membrane units were both open and both actuated. Even when both the deformable membrane units were actuated, there was flow of bead solution into both branches as the deformable membranes in this exemplary device did not act as a binary valve.
FIGS. 7A and 7B show an exemplary design for a single mixer segment of an active microfluidic mixer having deformable membrane units. FIG. 7B is a close-up of the deformable membranes 140 of the active mixer shown in FIG. 7A. In this example, two microfluidic inlet channels 150 carrying solutions for mixing intersect and form a single inlet channel 160 which then has two microfluidic channels 170 branching from the inlet channel 160 each having a deformable membrane unit (a pair of deformable membranes 140). Liquids from microfluidic inlet channels 1 and 2 flow into the main channel 160 and the flow is then branched again into two microfluidic channels 170 each having a deformable membrane unit with operating channels 180 that can be, for example, pneumatically controlled. The branches are then combined again to complete the first mixer segment and the outlet channel 190 can be connected to more mixer segments. These units, when actuated, increase the resistance to flow in the branches thereby introducing plugs as shown in FIG. 7C. FIG. 7C shows the concept of an active mixer using the deformable membrane unit with an idealized depiction of the distribution of the liquids along the microfluidic channels. The labels, A, B, C, and D above the boxes in FIG. 7C correspond to the labels in FIG. 7A indicating different segments of the microfluidic mixing device. Consider two liquids 1 (grey) and 2 (white) flowing from microfluidic input channels 1 and 2. The flow will be laminar in the main channel and liquid 1 will flow to the upper branch and liquid 2 through the lower branch. Mixing is achieved by actuating the deformable membrane unit in each branch in an alternating fashion at different frequencies. Due to the increase in resistance to flow when the deformable membrane unit is actuated in the upper branch, the liquids 1 and 2 flow predominantly through the lower branch, and similarly through the upper branch when the lower branch deformable membrane unit is actuated. Thus alternating plugs of liquid 1 and liquids 1 and 2 as shown in FIG. 7C flow through the upper branch while alternating plugs of liquid 2 and liquids 1 and 2 flow through the lower branch. When the flow from the two branches combine (as shown in FIG. 7C, box D), the diffusion length required for mixing of the two liquids is greatly reduced.
FIG. 8 shows qualitative results of mixing that were obtained in an exemplary mixer of FIG. 7A in which liquid 1 is a diluted solution of 1 μm fluorescent poly(styrene) beads in DI water and liquid 2 is DI water. In FIG. 8A, video microscopy of the main channel shows the distribution of the fluorescent beads after the first mixer segment before and after actuation of the deformable membrane units. The bead solution is limited to the left side of the main channel due to laminar flow before actuation, but after actuation, the bead solution is more uniformly distributed in the main channel. FIG. 8B shows the temporal average of the fluorescent intensity of each point within the channel before and after actuation of the deformable membrane units. In this Figure, 180 and 280 frames were averaged respectively from the video sequences before and after. FIG. 8C is a graphical representation of the data shown in FIGS. 8A and B, in which the y-axis is intensity in arbitrary units plotted across the channel before and after mixing. The deformable membrane units were actuated at a frequency of 1.66 Hz in an alternating fashion and video microscopy of the main channel after the first and the second segments was performed. As can be seen from FIG. 8, the bead solution showed mixing with the DI water.
Referring now to FIG. 9, a device for inducing and enhancing mixing between solutions is diagrammed. FIG. 9A provides a top view of the device, FIG. 9B provides a side view of a slice through the indicated region, and FIG. 9C provides a side view showing the microfluidic channel 230 and membranes 250. In this embodiment, two or more inlet microfluidic channels 200 formed in a unitary section of polymer 210 form a junction 220 at which they merge into a single inlet microfluidic channel 230. The angle at which the two inlet channels 200 meet in the diagram is arbitrary and any angle up to and including about 180° can be used. Operating channels 240 abut deformable membranes 250 formed from the polymer block 210. The polymer block 210 is attached to a substrate 260. Actuation of the membranes 250 causes the solutions to mix. A device such as this is useful, for example, for breaking up laminar flows of solutions. Although four deflectable membranes are pictured, a device could contain more or less than four membranes. Additionally, the membranes may also be staggered relative to a membrane on the opposite side of a channel. The frequencies of actuation of the membranes is dictated in part by the flow rates for the solutions and the desired state of the resulting solution.
Referring now to FIG. 10A, an exemplary device is pictured that can be used to break and coalesce droplets formed from the mixing of immiscible fluids. In this embodiment, two inlet microfluidic channels 270 form a junction 280 at which they merge into a single microfluidic channel 290. The angle at which the two inlet channels 270 meet in the diagram is arbitrary and any angle up to and including about 180° can be used. Operating channels 300 abut membranes 310 formed in a unitary section of polymer. Actuation of the membranes 310 causes the droplets that are formed from the meeting of the immiscible fluids in the microfluidic channel 290 to be mixed through the breaking and coalescing of the droplets. Although four deflectable membranes are pictured, such a device could contain more or less than four membranes. Additionally, the membranes may also be staggered relative to the membrane on the opposite side of a channel. Such a device is useful, for example, for facilitating the reaction between species contained in immiscible liquids. In an additional embodiment shown in FIG. 10B, two or more junctions 280 of two microfluidic inlet channels 270 are merged to form an additional junction 320 leading into a microfluidic channel 330 containing membranes 340 operably connected to operating channels 350 and capable of being deformed into the microfluidic channel 330. Although four deflectable membranes are pictured, such a device could contain more or less than four membranes. Additionally, the membranes may also be staggered relative to the membrane on the opposite side of a channel. Such a device is useful, for example, for mixing droplets formed in the two or more junctions of the microfluidic channels 280. The droplets may contain, for example, two different reactive species and the mixing of the droplets facilitates the reaction between the two different species.
Referring now to FIG. 11, a schematic shows a device that can be used for the introduction of precise amounts of a first fluid into a second fluid. Additionally, this device can be used for precise droplet formation via emulsion by using immiscible liquids as the sample and carrier fluids. In FIG. 11, a microfluidic channel 360 is provided for a first fluid, such as, for example, a sample stream, and a membrane 370 is operably connected to an operating channel 380. A second microfluidic channel 390 is located between the first microfluidic channel 360 and a third microfluidic channel 400 in which a second fluid can flow. In operation, deflection of the membrane 370 delivers a plug of the sample stream through the connecting channel 390 and into a carrier fluid contained in microfluidic channel 400. In general, the dimension labeled TIC can range from about 10 to about 200 μm and the dimension labeled LIC can range from about 5 to about 100 μm. The dimension TIC, for example, can be used to determine the drop size.
FIG. 12 shows an exemplary device for controllable release of a fluid into a microchannel. In FIG. 12, a microfluidic channel 410 is flanked by deformable membranes 420 that separate operating channels 430 from microfluidic channel 410. The number of operating channels chosen depends on the number of reagents or fluids that are desired to be added to the microfluidic channel and one, two, three, four, and more are possible. There can be only one operating channel 430 and membrane 420, or as many as desired. In operation, a first liquid is pumped into an operating channel with increasing pressure until the membrane 420 separating the operating channel 430 from the microfluidic channel 410 is ruptured and releases the first liquid into the microfluidic channel 410. Introduction of reagents can be controlled and sequenced by controlling the membrane thickness and the pumping of the various fluids in the operating channels. Higher pressures and thinner membranes will allow the membranes to rupture more quickly. The microfluidic channel can be a reaction vessel for performing a variety of chemical and biochemical reactions. In the case of a PDMS housing bonded reversibly to a substrate, the opening of the membrane occurs at lower pressures, since delamination or peeling of the membrane can occur at the substrate-PDMS interface. Additionally, in operation the membrane separating the operating channel enclosing the first fluid can be ruptured through increasing the flow rate of a second fluid in the microfluidic channel. The increased flow rate increases pressure within the microfluidic channel and results in membrane rupture and introduction of the first fluid into the second fluid.
The micro-fluidic channels represent micro-sized fluid passages that may have a cross-sectional dimensions, channel width, channel height, channel diameter, etc. that may be not greater than approximately one millimeter (mm, one-thousandth of a meter, also 1000 μm). In various embodiments the cross-sectional dimension may be not greater than approximately 500 micrometers (μm, one millionth of a meter), 200 μm, 100 μm, 50 μm, or 10 μm. The invention is not limited to any known minimum cross-sectional dimension for the channels. In various embodiments the cross-sectional dimension may be greater than approximately 0.001 μm (1 nm), greater than approximately 0.01 μm (10 nm), or greater than approximately 0.1 μm (100 nm). The optimal dimension of the channel may depend upon the characteristics of the fluids and/or particles to be conveyed therein. An exemplary micro-fluidic channel which may be used for one or more of an inlet, outlet, or focusing channel, may comprise a rectangular channel having a channel width of approximately 100 μm and a channel height of approximately 50 μm. The rectangular shape and specific dimensions are not required. These miniaturized channels are often useful for handling small sized samples and allow many channels to be constructed in a small substrate, although this is not a requirement. There is no known minimum or maximum length for the channels. Commonly the channel lengths are at least several times their width and not more than several centimeters.
PDMS may offer certain advantages such as compatibility with biological materials and chemicals and transparency to facilitate alignment, although the use of PDMS is not required and other materials may optionally be employed for forming the housing containing the membranes and microchannels. Any machinable, etchable, reformable, moldable, stampable, embossable, or castable elastomeric material (a material that is capable of deforming when pressure is applied and returning to its original shape when pressure is removed) may potentially be used. In general, there are a wide variety of formulations for elastomeric polymers, and a choice of materials may be based upon considerations such as elasticity, gas and/or liquid permeability, cost of fabrication, and/or temperature stability. Suitable polymers include among others, polyurethanes, silicones, polybutadiene, polyisobutylene, polyisoprene, elastomeric formulations of polyvinylchloride, polycarbonate, polymethylmethacrylate, polytetrafluoroethylene (Teflon®), and combinations of these materials. It may be appropriate to form focusing devices of polymers because these materials are inexpensive and may be injection molded, hot embossed, and cast.
In general, almost any non-absorbent material capable of presenting a smooth surface can be used to form the substrate. Possible substrates that could be used include glass; silicon; polymers, such as for example, PDMS, polystyrene, and polyethylene; silicon nitride; silicon dioxide; and metals, such as for example, gold, aluminum, and the like. The housing in which the channels and the membranes are formed may be reversibly or irreversibly attached to the substrate. For example, a PDMS housing can be reversibly attached to, for example, a PDMS or a glass surface through van der Waals forces. Additionally, adhesives such as silicone adhesives and epoxies can be used to bond the housing to the substrate. Choice of method of bonding is dependent in part on the materials chosen for the housing and the substrate, the desired user-chosen operating pressure ranges, and functional compatibility with operating fluids chosen for a particular application and can be effectuated according to well-known methods in the art. Additionally, PDMS, for example, can be oxidatively sealed to, for example, PDMS, silicon, polystyrene, polyethylene, silicon nitride, or glass by exposing the surfaces to be bonded to an air plasma and bringing the surfaces into contact within about a minute after oxidation.
The invention is generally not limited to any known process flow. Suitable process flows may comprise an aqueous, organic, or biological solution. The process flow may contain a species of interest. The species of interest may comprise a biological material, such as a cell, organelle, liposome, biological molecule or macromolecule, enzyme, protein, protein derivative, protein fragment, polypeptide, nucleic acid, DNA, RNA, nucleic acid derivative, biological molecule tagged with a particle, fluorescently labeled biological molecule, charged species, or charged protein. Additionally, a process flow may contain reagents for chemical reactions and the products of chemical reactions.
In general, the deformable membranes can be actuated (deflected) pneumatically, hydraulically, piezoelectrically, thermopneumatically, and magnetically. Pneumatic and hydraulic actuation can be accomplished by pumping a gas or liquid, respectively, into an operating channel. Typically, the gas or liquid can be supplied and vented through a valve that is controlled by a valve drive and a computer generating a programmed actuation pattern that is converted into a control signal. Piezoelectric actuation can be accomplished using, for example, the devices shown in FIG. 13. In FIG. 13, piezoelectric disk 431 is mounted with either a support plate 432 or a support structure 433 above an actuation reservoir 439. An operating channel 434 is separated from microfluidic channel 435 by deformable membrane 436. Polymer housing 437 is attached to substrate 438. FIG. 13C provides a top down view of the device shown in FIG. 13A. Deformation of the piezoelectric disk 431 into the operating channel 434 causes the membrane 436 to actuate. Piezoelectric disks are commercially available from, for example, Piezo Systems, Inc (Cambridge, Mass.).
Precursors for poly(dimethyl siloxane), Sylgard A and B were obtained from Dow Corning Inc. 1 and 6 μm YG fluorescent poly(styrene) beads used to visualize flow were obtained from Polysciences Inc. SU-2035 Photoresist was obtained from Microchem Corp.
An actuation system consisting of hardware and software components was constructed for pneumatically controlling the operating channels. Referring to FIG. 14, the actuation system consisted of a control computer 440 generating a programmed actuating pattern that is converted into a control signal through a digital output board (NI MIO-16XE-10, National Instruments) 450. The control signal operates the valve drive (NI SCCDO01, National Instruments) 460 that converts the control signals into the appropriate power leveled operating power patterns for switching the solenoid valves (LHDA1223111H, Lee company) 470. Regulated external gas pressures (10-30 psi) were provided to the normally closed port of the manifold on which the solenoid valves were mounted allowing the operating channels to be pressurized or vented.
The valve drives are enclosed in the signal conditioning box (NI SCC2345, National Instruments) having two RJ45 connectors, two sets of banana connectors and four LEDs. Two sets of banana connectors are to provide the external power which then is converted into the pulsing power by valve drives. There are eight valve drives and each set of banana connector is connected to four valve drives so that enough external power is supplied. Two 12 V power supplies are connected to the banana connectors. The role of valve drive is to turn on and off the external power for solenoid valves so that it generates the patterned pulsing power with particular frequencies.
The application for the actuation system was written in C language. In order to increase the response time to maximum, Graphic User Interface (GUI) was not implemented. Actuation patterns for performing synchronized actuation of the different deformable membrane units were implemented in the software depending on the microfluidic operations. Video microscopy was done using a Canon Digital camera ZV20 that captures the video via an S-video port from the Hamamatsu color CCD camera mounted onto an inverted fluorescence microscope.
FIG. 15 provides a general outline for microfluidic chip fabrication using standard single layer soft lithography. In FIG. 8, a photoresist on silicon master is prepared using standard photolithography using a thick SU-8 photoresist spun at thickness of 100 μm. This is followed by micromolding with PDMS after which the PDMS mold is peeled off the master and bonded to, for example, a glass or PDMS substrate. Other methods for forming microfludic structures are known and the invention is not limited to a particular method of forming the structures.
Designs of the micro fluidic channels to be fabricated were drawn to scale using L-Edit (Tanner Research) and chrome masks were printed using a Micronics laser writer at Stanford nanofabrication facility.
SU-8 2035 photoresist was spun onto 4″ silicon wafers at 2000 rpm for 30 sec. The wafers were then baked at 65° C. for 6 min. and at 95° C. for 20 min. The wafers are then exposed using UV light (365 nm) at a dose of ˜400 mJ/cm2. The exposed wafers were then baked at 65° C. for 1 min and at 95° C. for 5 min. After post-exposure bake, the wafers were immersed in SU-8 developer for ˜10 min. to develop the unexposed regions. The SU-8 photoresist on the wafer was then silanized for 1 hr by placing the wafers in close proximity with a few drops of trimethylchlorosilane in a vacuum desiccator. The silanized photoresist on the wafer was used as the master for subsequent micromolding experiments.
Ten parts by weight of Sylgard A were added to 1 part by weight of Sylgard B, mixed thoroughly and degassed to remove any air bubbles to form the PDMS precursor. PDMS precursor was poured onto the silanized master and then cured at 65° C. for 1 hr. The cured PDMS was peeled off the master and holes were punched for reservoirs. In order to irreversibly seal the PDMS to a glass cover, the PDMS and the glass cover were placed in a plasma cleaner and treated with plasma (100 W) generated from ambient air for 1 min. and brought into conformal contact within 30 sec.
In the examples shown liquids were flowed through microfluidic channels using gravity. Other methods are also possible, including, for example, pumps and syringes.