WO2001088525A1 - Structurally programmable microfluidic systems - Google Patents

Abstract

The present invention provides microfluidic systems, which have structurally programmable (PFD), reconfigurable, and multi-sample analysis capabilities. In one embodiment, the device includes structurally programmable fluidic paths, passive microvalves (30), fluidic components based on hydrophobic microfluidic systems (PFD), and pneumatic actuators using an air-bursting actuation concept. By controlling both the length and surface properties (e.g., hydrophilic or hydrophobic) of the channels (14), the pressure drops through the designed microfluidic systems will be controlled and thus programmable.

Description

Structurally Programmable Microfluidic Systems

[0001] This invention claims priority of U.S. Provisional Patent Appl. Ser.

Nos. 60/204,214, filed May 12, 2000 and 60/209,051, filed June 2, 2000, incorporated herein by reference.

[0002] This invention was made in part with Government support under Grant

No. AF 30602-00-1-0569, awarded by the Defense Advanced Research Projects Agency. The Government may have certain rights in this invention.

FIELD OF THE INVENTION

[0003] The present invention relates to a microfluidic device and more particularly, to a structurally programmable microfluidic system based on passive fluidic restriction valves and pneumatic air-bursting actuators.

BACKGROUND OF THE INVENTION

[0004] Various microfluidic systems are known for liquid phase biochemical analysis and manipulating fluids. In order to control a small amount of fluid in the order of nanoliters per minute, microfluidic regulation systems are necessary for the fast and precise control of small fluid volumes. To achieve this goal, microvalves, pumps, flow sensors and microchannels are considered essential components using micromachining techniques.

[0005] Microfluidic devices have been used for analyzing a variety of analytes such as those in blood or serum. Silicon provides the practical benefit of enabling mass production of such systems. A number of established techniques developed by the microelectronics industry using micromachining exist and provide accepted approaches to miniaturization. Examples of the use of such micromachining techniques are found in U.S. Patent Numbers 5,194,133, 5,132,012, 4,908,112, and 4,891,120 incorporated herein by reference in their entirety. Micromechanical devices and arrays of such devices may be mechanical, electromagnetic, electrostatic fluid or pneumatic in nature. Uses for such devices are readily apparent in the field. Such microdevices have been used for application in medicine, optics, microassembly, industrial process automation, analytical instruments, photonics and aerospace. In the field of micromechanical devices, miniaturization of analyzers provide an integrated system of pumps, flow valves, physical and chemical sensors, detectors, etc., produced on microscaie structures or composites consisting of several microcomponents made from different materials.

The microscaie devices such as in the present invention typically have a plurality of grooves or microchannels and chambers etched or molded into a substrate that can be silicon, plastic, quartz, glass or plastic. The size, shape and complexity of these microchannels and chambers and their innerconnections influence the limits of a microsystem's functionality and capabilities. The size, shape and complexity of microchannels and structures depend on the materials used and the fabrication processes available for those materials. Typically system fabrication includes making trenches in a conducting material such as silicon or in a nonconducting substrate such as glass or plastic and converting them to channels by bonding a cover plate to the substrate. Alternatively, microchannels and chambers may be made into a microdevice using multiple layers of similar or dissimilar materials. The typical overall channel sizes range from about 5 to about 100 microns wide and from about 5 to about 100 microns deep. Generally, microfluidic devices provide precise control of fluids by forming various grooves or channels and chambers in a substrate. The process of forming channels can include wet chemical etching, photo lithographic techniques, controlled vapor deposition and laser drilling into a substrate.

[0007] Microfluidic biochemical analysis systems or lab-on-a-chip systems are of great interest in the area of biotechnology in terms of blood analysis, biochemical detection, drug discovery, and so forth. Prior to the present invention, many microfluidic systems have been explored and realized on glass or plastic substrates. Generally, in order to achieve a microfluidic analysis system or biochip, active microfluidic components are essential to deliver small volume of fluid samples to a reservoir and/or reaction chamber for desired analysis, reaction and detection.

[0008] In the microfluidic biochips involving active microactuators, most power will be consumed for driving active microfluidic actuators such as microvalves and pumps, whereas the power consumption for the signal conditioning circuits is negligibly small compared to the power required for driving microactuators. The present invention provides an innovative method to address the relevant difficult issues related to power consumption, which uses pressurized air or inert gas as an alternative power to drive microatuators. The pressurized air or inert gas will be injected into on-chip reservoirs, which are located on the disposable plastic biochip. So, each disposable plastic chip holds several microreserviors, which contain different air or gas pressures as an alternative power to drive liquid.

[0009] The present invention contemplates a structurally programmable microfluidic system concept for application to a disposable biochip at low cost by means of structurally programmable passive fluidic valves and pneumatic air-bursting actuators that overcome all of the above- referred problems. [0010] A great amount of effort is presently being directed toward developing practical fluidic interconnection techniques so that they can be economically produced and further offer the facility of being easy to integrate with the microfluidic system. Broadly described, microfluidic interconnections are a means of realizing a microfluidic coupling between the macro realm and the microfluidic analysis unit. Unfortunately, the microscopic nature of these interconnects makes implementation complex, tedious and time consuming, often resulting in non-optimal solutions. One of the key requirements for a reliable microfluidic interconnect technique is low dead volume. Micro- biochemical analysis units typically utilize sample, reagent and analyte volumes of the order of μL (microliter) to nL (nanoliter) range. It is essential that little or no part of this minute volume be lost in the interconnection.

[0011] Various solutions have been proposed for microfluidic interconnection, including the use of adhesives, fluidic couplers, and frictional force fitting. The proposed techniques suffer from shortcomings such as, complex assembly, non-repeatable performance, cumbersome methodology, high dead volumes, and lack of alignment.

[0012] Microfluidic interconnects are necessary for transferring controlled amount of fluids to and from a microfluidic biochemical analysis platform. It is necessary that this operation be performed without any loss of fluids in the interconnect and the pressure required to transfer fluids should be low. These requirements can only be satisfied if the microfluidic interconnect is accurately aligned with the opening on the microfluidic platform. However, it is very difficult to realize good alignment during the assembly stage, as the tolerance for misalignment is very low (typically tens ofμm). Furthermore, the realized interconnects should be robust to allow manual handling when connecting / disconnecting tubing to the microfluidic interconnect. Recently, a novel self-aligning microfluidic interconnection technique has been proposed and demonstrated by Aniruddha Puntambekar et al. in "Self- Aligning Microfluidic Interconnects with Low Dead Volume", Proc. of the μTAS conference, Enschede, The Netherlands, pp. 323- 326, May 2000. This technique allows the fabrication of self-aligning microfluidic interconnects that are fabricated in-situ after microfluidic assembly. It is possible to realize very low a dead volume in these interconnects because of the high alignment accuracy between the tubing and the opening on the microfluidic platform.

[0013] In further regard to the interconnection techniques disclosed in

Puntambekar et al. describes the structural configuration and fabrication technique used therein. Two techniques are disclosed namely, a serial assembly technique and a parallel assembly technique. Both techniques rely on the use of a thermoplastic tubing material that is deformed under controlled conditions of heat and pressure to conform to the shape of the opening in the microfluidic platform. Since both the techniques use the structure of the opening in the microfluidic platform to form the microfluidic interconnect, hence self- aligning, high accuracy in alignment is achieved.

[0014] To date, no known technique has been able to achieve the self- alignment feature and its consequent benefit of low dead volume and low-pressure drop. In this invention, we describe a novel self-aligning microfluidic interconnection technique with low dead volume and pressure drop for generic microfluidic systems and CE chips.

SUMMARY OF THE INVENTION

[0015] In this invention, we disclose an innovative concept of microfluidic systems, which have structurally programmable, reconfigurable, and multi-sample analysis capabilities. An innovative, fully integrated, plastic microfluidic chip will be developed for the dual applications of a fully stand-alone biochip as well as a wrist watch-type analyzer. The disclosed device includes structurally programmable fluidic paths, passive microvalves, fluidic components based on hydrophobic microfluidic systems, and pneumatic actuator using an air-bursting actuation concept.

[0016] Specifically, we disclose an innovative fluid control technique using structurally programmable fluidic paths and passive microvalves and fluidic components, based on hydrophobic microfluidic systems. The structurally programmable paths consist of simple passive valves and flow conduits, which have different pressure drops depending on both the structure and the surface properties of the fluidic paths. By controlling both the length and surface properties (e.g., hydrophilic or hydrophobic) of the channels, the pressure drops through the designed microfluidic systems will be controlled and thus programmable. Since most biochemical assays using a microfluidic chip require a simple fluidic control sequence, the innovative structurally programmable techniques, proposed here, can immediately address several relevant difficult issues to the fluid control in the microfluidic chips, without using a flow sensor.

[0017] The structurally programmable microfluidic system consists of simple passive valves and flow conduits, which have a series of different pressure drops depending on both the geometrical configuration and surface properties of the microfluidic channels. The microfluidic system components are serially positioned in the system as per their sensing/actuation objective. By tailoring both the length and surface properties (e.g. degree of hydrophobicity) of the channels, the pressure drops through the designed microfluidic system will be controlled and thus programmable. By using the discrete air-bursting actuation scheme proposed we could position and maneuver the fluid in the microfluidic system with high accuracy. [0018] The production of a structurally programmable fluidic system will be partly based on localized control over hydrophihcity for flow channels on a micron scale. Development of a graded or variable hydrophihcity along a flow path on a micron to millimeter scale is also of interest with the aim of completely tunable flow properties. Polymer surfaces are typically hydrophobic and there are many techniques available for the modification of such surfaces to induce hydrophihcity ranging form simple flame or thermal treatment of polyolefins, in the presence of excess oxygen, yielding carboxyl modified surfaces; targeted plasma treatment; chemical etching using acids; and more controlled chemical modification where surface active agents are introduced and bonded to the polymer surface.

[0019] The present invention provides a structurally programmable microfluidic system for a low cost biochemical analysis system using programmable passive valves and air-bursting pneumatic actuators. Air-bursting pneumatic actuators can replace active microfluidic components including power sources. In order to realize structurally programmable microfluidic systems, a state-of-the-art pneumatic actuation technique using an air-bursting actuator has to be designed, which is a very attractive approach when strong forces are desired in a microfluidic chip. In the microfluidic bio-chips involving active microactuators, most power will be consumed in driving active microfluidic actuators such as microvalves and pumps, whereas the power consumption for the signal conditioning circuits is negligibly small compared to the power required for driving microactuators. In this invention, we disclose an innovative method to address the difficult issues relevant to power, which uses a pressurized air or inert gas as an alternative power to drive the microatuators. The pressurized air or inert gas will be injected into the on-chip reservoirs, which are located on the disposable plastic biochips. So, each disposable plastic chip holds several microreserviors, which contain different air or gas pressures as an alternative power to drive liquid. Specifically, if microreservoirs in the bio-chips are discretely pressurized with air or gases, which are tuned with the already structurally programmed microfluidic systems, the fluid control of microfluidic system can be achieved by detonating the microreservoirs using electronically programmed containers.

[0020] The disclosed pneumatic actuator consists of a pressurized reservoir with a passive inlet port that provides the injection and sealing of fluid and an outlet port for the air-burst from the pressurized reservoirs. Specifically, the outlet port is comprised of a small heater to melt a thermodestructing seal at the outlet port. Since the heater is designed to produce only enough heat to melt the seal in an extremely small area, very little power will be consumed for only a short time. This will be sufficient to initiate a burst of air from the pressure reservoir.

[0021] Application of the present invention is generally a plastic-based, low cost biochemical analysis system using structurally programmable microfluidic system concepts including passive valves and pneumatic air-bursting actuators.

[0022] It is an object of the present invention to overcome the difficulties and complexities in integration of active microfluidic components into a microfluidic system.

[0023] It is also an object of the present invention to simplify controlling nano-liter range of fluidic samples using programmed microfluidic paths based on already structured fluidic restrictive valves.

[0024] It is another object of the present invention to provide a disposable biochip when a host analyzer system is provided. [0025] It is another object of the present invention to involve no moving parts, which requires a minimal control signal input and thus provides unique fluidic control for disposable biochips.

BRIEF DESCRIPTION OF THE DRAWINGS

[0026] This invention, as defined in the claims, can be better understood with reference to the following drawings. The drawings are not necessarily to scale, emphasis instead being placed upon clearly illustrating principles of the present invention.

[0027] FIG. 1 is a conceptual illustration of structurally programmable microfluidic delivery system for the dual uses of stand-alone biochip and wrist watch-type analyzer.

[0028] FIG. 2 is a schematic illustration of a passive fluidic restriction or microvalve.

[0029] FIG. 3 is a schematic illustration of a structurally programmable microfluidic delivery system. Fluidic channels are programmed to have a delivery sequence from 1 to 14.

[0030] FIG. 4A-4D depict schematic illustrations of structurally programmable fluid delivery sequences.

[0031] FIG. 5A-5B are graphical representations of the pressure drops in the system as a function of fluid position and the effect of flow rate on pressure distribution at: FIG. 5A. 1 μL/min and FIG. 5B.0.1 μLV min.

[0032] FIG. 6 is a conceptual illustration of a pneumatic air-bursting microactuator.

[0033] FIG. 7 is a schematic illustration of a pneumatic air-bursting microactuator as an alternative power source. [0034] FIG. 8 is a schematic cross-sectional view of the microfluidic interconnects assembled using the serial interconnection technique.

[0035] FIG. 9 depicts a schematic cross-sectional view of the microfluidic interconnects assembled using the parallel interconnection technique.

[0036] FIG. 10. Schematic diagram of serial interconnects: (a) bonded wafer assembly; (b) flanging operation; and (c) assembled view.

[0037] FIG. 11. A schematic diagram of parallel interconnects: (a) insertion of flanged tubing in plastic holder wafer; (b) after heat/pressure treatment; and (c) assembled view.

[0038] FIG. 12. A schematic diagram of parallel interconnects for plastic microfluidic devices: (a) insertion of flanged tubing in substrate; (b) after heat/pressure treatment; (c) insertion of flanged tubing in substrate with embossing master; (d) after heat/pressure and simultaneous embossing treatment; and (e) assembled view.

DETAILED DESCRIPTION OF THE INVENTION

Definitions

[0039] The present invention provides a structurally programmable microfluidic system based on passive fluidic restriction valves and pneumatic air-bursting actuators. The present invention also provides for the production of biochips containing such structurally programmable microfluidic system and actuators.

[0040] The term "chip", "microchip" or "bio-chip" as used herein means a microfluidic system containing microdevice components on a substrate. The chip generally includes fluidic pass-active and/or passive microvalves, fluidic components, electrical magnetic and/or pneumatic actuators, chambers, receptacles, diaphragms, detectors, sensors, ports, pumps, switches, conduits, filters, and related support systems. The term "microfluidic" refers to a system or device having a network of chambers connected by channels, tubes or other interconnects in which the channels may act as conduits for fluids or gasses. "Microfluidic analytical systems" refer to systems for forming chemical, clinical, or environmental analysis of chemical and/or biological specimens. Such microfluidic systems are generally based on a chip. These chips are preferably based on a substrate for micromechanical systems. Substrates are generally fabricated using photolithography, wet chemical etching, laser ablation, injection molding and other techniques similar to those employed in the semiconductor industry. Microfluidic systems generally provide for flow control and physical interactions between the samples and the supporting analytical structure. The microfluidic device generally provides conduits and chambers arranged to perform numerous specific analytical operations including mixing, dispensing, valving, reactions, detections, electrophoresis and the like. Increasing the number and complexity of networked or arrayed channels, reaction chambers and the like may enhance the analytical capabilities of such systems.

Flow control management may be use of a variety of mechanisms including the patterned application of voltage, current or electrical power to the substrate, for example, to induce and/or control electrokinetic flow. Alternatively, fluid flows may be induced mechanically through the application of differential pressure, acoustic energy and the like. Selective heating, cooling and exposure to light or other radiation or other inputs may be provided at selected locations, distributed throughout the substrate to promote the desired chemical and/or biological interactions. Measurements of lights or other emissions, electrochemical signals, pH, etc. may be taken from the substrate to provide analytical results. [0042] The term "substrate" is used herein to refer to any material suitable for forming a microfluidic device, such as silicon, silicon dioxide material such as quartz, fused silica, glass (borosilicates), laser ablatable polymers (including polyimides and the like), and ceramics (including aluminum oxides and the like). Further, substrate may comprise composite substrates such as laminates. A "laminate" refers to a composite material formed from several different bonded layers of same or different materials.

[0043] It should be noted that throughout the description the terms "channel" and "micro-channel" refer to structures for guiding and constraining gasses or fluids and gas or fluid flow and also include reservoir structures associates with micro-channels and will be used synonymously and interchangeably unless the text declares otherwise. Channel intersections may exist in a number of formats, including cross intersections, "T" intersections, or any number of other structures whereby two channels are in fluid communication.

[0044] The term "microfluidic" generally refers to structures or features of a device for transporting gasses or fluids that have at least one dimension or structural element in the range of from about 0.1 microns to about 500 microns. Microfluidic systems are particularly well adapted for analyzing small sample sizes. Sample sizes are typically are on the order of nanoliters and even picoliters.

[0045] As used herein, the term "detection means" refers to any means, structure or configuration that allows one to interrogate a sample within the microfluidic device using analytical detection techniques well known in the art. Thus, a detection means includes one or more apertures, elongated apertures or grooves that communicate with the microfluidic device and allow an external detection apparatus or device to be interfaced with the microfluidic device to detect an analyte. [0046] One or more electrical "signal paths" capable of carrying and/or transmitting electric current can be arranged such that the signal paths, in combination, form a circuit. As used herein, a signal path includes any conductive material that is capable of transmitting or receiving an electrical signal. In an exemplary embodiment, the conductive material is gold or copper. In another embodiment, the electrical signal is provided by one or more pulses of electrical current modulated by a controller, e.g., a CPU.

[0047] The term "motive force" is used to refer to any means for inducing movement of a sample along a channel or reservoir, and includes application of an electric potential, application of a pressure differential or any combination thereof.

[0048] The term "surface treatment" is used to refer to preparation or modification of the surface of a microchannel that will be in contact with a sample during separation, whereby the separation characteristics of the device are altered or otherwise enhanced. Accordingly, "surface treatment" as used herein includes: physical surface adsorptions; covalent bonding of selected moieties to functional groups on the surface of microchannel substrates (such as to amine, hydroxyl or carboxylic acid groups on condensation polymers); methods of coating surfaces, including dynamic deactivation of channel surfaces (such as by adding surfactants to media), methods of plasma treatment to activate (or inactivate) surface such as plasma treatment in oxygen plasma, coating of bioaffinity reagents to the surfaces of channels, polymer grafting to the surface of channel substrates (such as polystyrene or divinyl-benzene) and thin-film deposition of materials such as diamond or sapphire to microchannel substrates.

[0049] The term "laser ablation" is used to refer to a machining process using a high-energy photon laser such as an excimer laser to ablate features in a suitable substrate. The excimer laser can be, for example, of the F₂, ArF, KrCl, KrF, or XeCl type. In laser ablation, short pulses of intense ultraviolet light are absorbed in a thin surface layer of material within about 1. micron of less of the surface. Preferred pulse energies are greater than about 100 millijoules per square centimeter and pulse durations are shorter than about 1 microsecond. Under these conditions, the intense ultraviolet light photo-dissociates the chemical bonds in the material. Furthermore, the absorbed ultraviolet energy is concentrated in such a small volume of material that it rapidly heats the dissociated fragments and ejects them away from the surface of the material. Because these processes occur so quickly, there is o time for •heat to propagate to the surrounding material. As a result, the surrounding region is not melted or otherwise damaged, and the perimeter of ablated features can replicate the shape of the incident optical beam with precision on the scale of about one micrometer.

[0050] Although laser ablation has been described herein using an excimer laser, it is to be understood that other ultraviolet light sources with substantially the same optical wavelength and energy density may be used to accomplish the ablation process. Preferably, the wavelength of such an ultraviolet light source will lie in the 150 nm to 400 nm range to allow high absorption in the substrate to be ablated. Furthermore, the energy density should be greater than about 100 millijoules per square centimeter with a pulse length shorter than about 1 microsecond to achieve rapid ejection of ablated material with essentially no heating of the surrounding remaining material. Laser ablation techniques are well known in the art.

[0051] The term "injection molding" is used to refer to a process for molding plastic or nonplastic ceramic shapes by injecting a measured quantity of a molten plastic or ceramic substrate into dies (or molds). In one embodiment of the present invention, devices can be produced using injection molding. More particularly, it is contemplated to form a mold or die of a miniaturized column device wherein excimer laser- ablation is used to define an original microstructure pattern in a suitable polymer substrate. The microstructure thus formed may then be coated by a very thin metal layer and electroplated (such as by galvano forming) with a metal such as nickel to provide a carrier. When the metal carrier is separated from the original polymer, an mold insert (or tooling) is provided having the negative structure of the polymer. Accordingly, multiple replicas of the ablated microstructure pattern may be made in suitable polymer or ceramic substrates using injection-molding techniques well known in the art.

[0052] The term "LIGA process" is used to refer to a process for fabricating microstructures having high aspect ratios and increased structural precision using synchrotron radiation lithography, galvanoforming, and plastic molding. In a LIGA process, radiation sensitive plastics are lithographically irradiated at high-energy radiation using a synchrotron source to create desired microstructures (such as channels, ports, apertures and micro-alignment means), thereby forming a primary template.

[0053] One or more layers of material formed from a dimensionally stable support can form the substrate. Suitable substrate materials may be semiconductor materials, dielectric materials, plastics and polymeric resins, and alloys. Dielectric materials may be selected from the group consisting of paper, glass, quartz, silicon, polyethylene, ceramics, and mixtures thereof. In the case of polymeric substrates, the substrate materials may be rigid, semi-rigid, or non-rigid, opaque, semi-opaque or transparent, depending upon the use for which they are intended. For example, devices that include an optical or visual detection element will generally be fabricated, at least in part, from transparent materials to allow, or at least facilitate that detection. Alternatively, transparent windows of, e.g., glass or quartz, may be incorporated into the device for these types of detection elements. Additionally, the polymeric materials may have linear or branched backbones, and may be crosslinked or non-crosslinked. Examples of particularly preferred polymeric materials include, e.g., polymethylmethacrylate (PMMA), polydimethylsiloxanes (PDMS), polyurethane, polyimide, polyvinylchloride (PVC), polystyrene, polysulfone, polycarbonate, and the like. Preferably, these materials will be phenolic resins, epoxies, polyesters, thermoplastic materials, polysulfones, or polyimides and/or mixtures thereof.

[0054] In addition to constructing the substrate using conventional printed circuit board composites, alternative structures can be used. For example, for certain applications the use of plastic films, metals, glasses, ceramics, injection molded plastics, polyastomeric layers, ferromagnetic layers, sacrificial photo resist layers, shaped memory metal layers, or other suitable materials may be used. These may be bound with other substrate to form the system with or without an adhesive bonding layer.

[0055] In general, microfluidic devices can be fabricated out of any material that has the necessary characteristics of chemical compatibility and mechanical strength. One exemplary material is silicon since a wide range of advanced microfabrication and micromachining techniques have been developed for it and are well known in the art. Additionally, microfluidic devices can be produced directly in electrically insulating materials. The most widely used processes include isotropic wet chemical etching of glass or silica and molding of plastics. In another embodiment, the microfluidic devices can be produced as a hybrid assembly consisting of three layers — (1) a substrate, (2) a middle layer that forms the channel and/or chamber walls and whose height defines the wall height generally joined or bonded to the substrate and (3) a top layer generally joined or bonded to the top of the channels that forms a cover for the channels. I-n one exemplary method, the channels are defined by photolithographic techniques and etching away the material from around the channel walls produces a freestanding thin walled channel structure. Freestanding structures can be made to have very thin or very thick walls in relation to the channel width and height. The walls, as well as the top and bottom of a channel can all be of different thickness and can be made of the same material or of different materials or a combination of materials such as a combination of glass and silicon. Sealed channels or chambers can be made entirely from silicon glass and/or plastic substrates.

[0056] In one exemplary method, the microfluidic structure is formed from silicon nitride on a non-conducting substrate such as glass. Coating etched features in a silicon wafer with deposited silicon nitride forms the silicon nitride channels. The silicon nitride channels are bonded to the substrate by an intermediate thermo-oxide layer grown on the surface of the silicon nitride. The silicon wafer is etched away leaving a silicon nitride channel on the surface of the substrate. An electrically insulating material can then be applied to the substrate to support the silicon nitride structures. Sealed channels or reservoirs can be made entirely out of silicon with a substrate present only for mechanical support.

[0057] In another embodiment, rather than using conventional practice of fabricating fluid channels as trenches or grooves in a substrate. The fluid channels or reservoirs are fabricated as raised structures on a substrate. Such a substrate can be any material that will bond to silicon and is generally glass such as PYREX. In selecting the correct substrate material, important parameters are the thermal expansion coefficient being compatible with the silicon and flatness and ion content. Bonding of a silicon wafer to a substrate may be by methods well known in the art such as anodic bonding. The thickness of the silicon can be reduced through chemical or mechanical polishing. An etch resistant material such as polymeric photo resist material or hard mask such as silicon oxide, silicon nitride or some comparable material that provides high etch selectivity for subsequent processing may then be applied to the surface of the polished silicon layer. Photolithographic patterning methods are used to form the desired channel pattern as well as define thickness of the channel walls in the photo resist or hard mass material.

[0058] In additional preferred aspects, the substrate materials will comprise polymeric materials, e.g., plastics, such as polymethylmethacrylate (PMMA), polycarbonate, polytetrafluoroethylene (TEFLON.TM.), polyvinylchloride (PVC), polydimethylsiloxane (PDMS), polysulfone, and the like. Such polymeric substrates are readily manufactured using available microfabrication techniques, as described above, or from microfabricated masters, using well known molding techniques, such as injection molding, embossing or stamping, or by polymerizing the polymeric precursor material within the mold (See U.S. Pat. No. 5,512,131). Such polymeric substrate materials are preferred for their ease of manufacture, low cost and disposability, as well as their general inertness to most extreme reaction conditions. Again, these polymeric materials may include treated surfaces, e.g., derivatized or coated surfaces, to enhance their utility in the microfluidic system, e.g., provide enhanced fluid direction, e.g., as described in U.S. Pat. 5,885,470, and which is incorporated herein by reference in its entirety for all purposes.

Structurally Programmable Fluidic Paths

[0059] The Hagen-Poiseuille equation for laminar flow governs the pressure drops in microfluidic systems with laminar flow. For a rectangular channel with a high width to height ratio, the pressure drop is governed by equation (1): VλLμJQ

ΔR = - (1) wh³ where L is the length of the microchannel, μ is the dynamic viscosity of the fluid, Q is the flow rate, and w and h are the width and height of the microchannel respectively. By varying L or Q, we can directly control the pressure drop for a given set of w and h.

[0060] For the passive fluidic restriction that is illustrated in FIG. 2, an abrupt change in the width of the channel causes a pressure drop at the point of the restriction. For a hydrophobic channel material, an abrupt decrease in channel width causes a positive pressure drop as

AP₂ = 2σ, cos(θ_c) , (2)

where σ is the surface tension of the liquid, θ_c is the contact angle, and wi, h_\ and w₂, hi are the width and height of the two sections before and in the restriction, respectively. Setting h as constant through the system we can vary ΔP₂ by adjusting the ratio of

and w₂. Thus, by adjusting the both the physical structure of the system and the ratio of the restriction, we can control the pressure drops throughout the system.

[0061] In one embodiment, each channel dimension is adjusted to accommodate the exact same amount of fluid as the succeeding channel pair. At the split-off point, a passive microvalve is placed in one of the channels, to ensure that the fluid will first fill up the other channel. Fluid will first flow into channel 1 then in channel 2 because of restriction (or passive valve) Rl . The passive microvalve is designed so that when fluid fills up the channel, the pressure drop in the channel exceeds the pressure drop across the first valve. At the end of each channel pair, there are passive valves (or flow restrictions) of unequal dimensions that further regulate flow, as shown in FIG. 3 (R2 and R3). By extending this arrangement, fluid can be manipulated to an exact location as desired. In this case, the delivery sequence is 1, 2, 3, 4, 5, 6, and so on, as illustrated in FIG. 3. By changing the locations of the passive valves and/or their relative values with respect to channel size, we can program the fluid delivery sequence of a fluidic system as shown in FIG. 4 A through FIG. 4D.

[0062] Without wishing to be bound by theory, the present invention works on the principle that the pressure drop of the channels is small compared to the pressure drop across the passive microvalves. This allows the microvalves to mainly control the flow sequence. As described in Equation (2), ΔP₂ is independent of flow rate, whereas ΔP from Equation (1) is directly proportional to the flow rate. Thus, the flow rate has to be very low for the microvalves to control flow. FIG. 5A and 5B, which depict the numerical calculations using Equations (1) and (2), show the effect of flow rate on system performance. The system will be fabricated using several different methods, e.g., plastic injection molding, plastic embossing technique, polymer cast molding, laser ablation technique, and others well known in the art.

[0063] A typical programmable fluidic device (PFD) is schematically illustrated in FIG. 2. Fluid 12 enters the main microfluidic channel 14 that in fluid communication with a succeeding pair of microfluidic channels 16 and 18. As shown in the diagram, a passive microvalve 30 is placed in one branch channel 18 to ensure that the fluid will first fill up the other channel 16. Generally each channel dimension is adjusted to accommodate the exact same amount of fluid as the succeeding channel pair. Fluid will first flow into channel 16 and then channel 18 because of restriction or passive valve 30. After filling the first branch 16 and restricted branch 18, a second restriction or passive valve 22 placed at the end of microchannel 16 so that fluid will first flow into branch 18. After branch 18 fills, restriction 31 prevents further flow until the pressure fills the succeeding branches 20 after branch 16 and passed restriction or passive valve 22. The passive microvalves are designed so that when fluid fills up a channel the pressure drop in the channel exceeds the pressure drop across the previous valve. At the end of each channel pair, there are passive valves or flow restrictions of unequal dimensions that further regulate flow such as the paired restrictions 22 and 31.

[0064] In general, the passive microvalve is designed so that when fluid fills up the preceding channel, the pressure drop in the channel exceeds the pressure drop across the first valve. At the end of each channel pair, there are passive microvalves or flow restrictions of unequal dimensions that further regulate flow as shown in FIG. 4. By extending this arrangement, fluid can be manipulated to an exact location within the PFD in a delivery sequence of any order desired. For example, the fluid can be manipulated to an exact location in a delivery sequence of filling channel 24a first then channel 24b, then channel 24c, 24d, so forth, until channels 24f and 24h are filled as illustrated in FIG. 4b and c. By changing the locations of the passive valves and/or their relative sizes with respect to channel size, the PFD can be programmed for a fluid delivery sequence of any order as shown in FIG. 4.

[0065] For the passive fluidic restriction that is illustrated in FIG. 3, fluid enters through channel 44 of a certain width 46, length 50 and depth 47. There is then placed a passive fluidic restriction 42 with an appropriate width 48 and depth 49 that causes a decrease in pressure. In order for the fluid of channel 41 to pass through restriction 42, there needs to be an increase in the pressure of channel 41 sufficient to cause flow through restriction 42 and into channel 43 therein exiting through port 45.

[0066] In a typical microfluidic system, the substrate for the microfluidic channels will include a port in fluid communication with channel 14. A fluid sample 12 such as blood or other fluid for analysis is deposited with the substrate port typically relying on pneumatic or hydraulic pressure or capillary forces to deposit the fluid in the microchannel 14. Generally, sample liquid deposited in hole ports extending into or through the substrate will wick into channels or chambers within the device by capillary forces between the fluid and the surrounding surfaces. As only very small amounts of fluid are needed for the microfluidic system of the present invention, the size of the drops of fluid can be quite small. Drops generally enter through a port or hole in substrate and are drawn into a chamber or channel by both gravity and capillary forces. Generally, capillary forces draw fluids from larger channels to smaller channels. The minimum cross-sectional dimension of a channel largely controls capillary forces. For example, capillary forces will wick a fluid from a channel 10 microns by 20 microns into a contiguous channel having a dimension of 10 microns by 10 microns cross-sectional. Hence, simple capillary forces may be relied on to draw fluid from or into microfluidic channels so long as the microfluidic channels have a smaller cross-sectional dimension than the smallest cross-sectional dimension of the previous port. The production of PFDs will be based partly on the localized control over hydrophihcity for flow channels on a micron scale. Development of a graded or variable hydrophihcity along a flow path on a micron to millimeter scale is also possible. Polymer surfaces are typically hydrophobic and there are many techniques well known in the art for modification of such surfaces to induce hydrophihcity ranging from simple thermal treatment of polyolefins in the presence of excess oxygen yielding carboxymodified surfaces.

The cross-sectional dimensions of channels and restrictions will typically be selected to provide sufficient pressure drop and will depend upon the wettability of the material bordering the channel, the fluid to be retained therein, the length of the channel, the depth of the channel and the like. The wettability of the channel will depend upon the substrate materials used.

[0068] The channels of the PFD such as the main channel 14 will typically have a minimum cross-sectional dimension of between from about 0.1 to about 1,000 microns. Typically the minimum cross-sectional dimension will be from about 0.5 to about 100 microns. The passive microvalves or restrictions will generally be smaller than the main fluid channels ideally having a minimum cross-sectional dimension of between from about 5 and about 95% of the minimum cross-sectional dimension of the main channel. More typically, the restriction will generally have a cross-sectional dimension of between from about 10 and about 50% of the minimum cross-sectional dimension of the main channel. Typically the channels are trough-like recesses adapted for microfluid flow therein and may have widths of about 5 to about 1,000 microns preferably from about 10 to about 100 microns. The channels generally have a channel depth being from about 5 microns to 1,000 microns preferably from about 10 microns to about 500 microns.

[0069] In one embodiment, novel miniaturized devices of the present invention can be laser ablated into a substrate other than silicon or silicon dioxide materials. The use of laser ablation techniques in the practice of the invention enables highly symmetrical and accurately defined devices to be fabricated in a wide class of polymeric and ceramic substrates to provide a variety of miniaturized systems. In this regard, miniaturized channels may be provided which have widths ranging from about 0.5 to about 1000 microns, typically ranging from about 1 to about 500 microns, and channel lengths ranging from about 5 microns to about 100 millimeters, typically ranging from about 25 microns to about 10 millimeters. [0070] In the practice of the invention, laser ablating a set of desired features in a selected substrate using a step-and-repeat process to form discrete units may form PFD's. In this regard, it is particularly contemplated to laser ablate the subject devices in condensation polymer substrates including polyimides, polyamides, polyesters and polycarbonates. Further, the instant invention may be practiced using either a laser ablation process or a LIGA process to form templates encompassing a set of desired features, whereby multiple copies of miniaturized devices may be mass-produced using injection molding techniques well known in the art. More particularly, it is contemplated herein to form miniaturized devices by injection molding in substrates comprised of materials such as the following: polycarbonates; polyesters, including poIy(ethylene terephthalate) and poly(butylene terephthalate); polyamides, (such as nylons); polyethers, including polyformaldehyde and poly(phenylene sulfide); polyimides, such as KAPTON and UPILEX; polyolefin compounds, including ABS polymers, Kcl-F copolymers, poly(methyl methacrylate), polystyrene- butadiene) copolymers, poly(tetrafluoroethylene), poly(ethylenevinyl acetate) copolymers, poly(N-vinylcarbazole) and polystyrene.

[0071] Laser ablation of microchannels in the surfaces of the above-described substrates has the added feature of enabling a wide variety of surface treatments to be applied to the microchannels before formation of the sample processing compartment. That is, the open configuration of laser-ablated microchannels produced using the method of the invention enables a number of surface treatments or modifications to be performed which are not possible in closed format constructions, such as in prior micro-capillaries. More specifically, laser ablation in condensation polymer substrates provides microchannels with surfaces featuring functional groups, such as carboxyl groups, hydroxyl groups and amine groups, thereby enabling chemical bonding of selected species to the surface of the subject microchannels using techniques well known in the art. Other surface treatments enabled by the open configuration of the instant devices include surface adsorptions, polymer graftings and thin film deposition of materials such as diamond or sapphire to microchannel surfaces using masking and deposition techniques and dynamic deactivation techniques well known in the art of liquid separations.

Pneumatic Microactuator

[0072] The microfluidic systems of the present invention can be used in a variety of chemical analysis systems using active pneumatic pumps to overcome the change in pressure required to fill various segments of the microfluidic system. Such an analysis system will require pneumatic actuators and/or power sources for pumping of fluids within the device. In a microfluidic device using active microactuators, most power would be consumed in driving active elements such as microvalves and pumps whereas the power consumption for the signal conditioning circuits of such devices is negligible compared to the power required for the driving elements. Therefore, the present invention provides an alternative power source to drive the microactuators. The alternative power source in derived from pneumatic bursting microactuators derived from pressurized air or inert gas or other fluids injected into on-chip reservoirs located throughout the chip or device. The microdevices of the present invention can hold one or more micro-reservoirs that contain pressurized gas or fluid for use as an alternative power source.

[0073] In one embodiment, the pneumatic actuator consists of a pressurized reservoir with an outlet port for release of the pressurized gas or fluid within the reservoir.

[0074] In one example depicted in FIG. 6a, the pneumatic actuator 60 is shown schematically. A reservoir microstructure 62 having an outlet port 64 is laser ablated or otherwise formed into the substrate 70. The reservoir microstructure may be formed in any geometry and with any aspect ratio to provide a reservoir compartment having a desired volume. The reservoir microstructure 62 has an outlet port 65. In one embodiment, a rupture disk 64 closes off the outlet 65.

[0075] As used herein, a "rupture disk" is a frangible segment of the reservoir capable of breaking, melting or otherwise rupturing upon actuation. Actuation is by direct application of force on the rupture disk or by an actuator capable of initiating the rupturing of the rupture disk upon communication of a signal along one or more signal paths by a suitable controller. An actuator may be by any electrical or chemical ignition, heat or reaction, pressure, mechanical force or other means for rupturing such a disk. The rupture disk may be comprised of the same or different material as the reservoir substrate wall 70. .

[0076] In another embodiment, the rupture disk is comprised of a burstable diaphragm 66 in communication with a burst activator 68. The ruptured disk 64 seals off the reservoir 62 until the burst activator 68 is activated by a signal such as an electrical current applied from a source such as a controller.

[0077] Typically, outlet port 65 is in fluid communication with a sample processing compartment or other reservoir by way of an interconnecting microchannel 74. In another example depicted in FIG. 6b, the reservoir microstructure 60 has an inlet port 84 fluidly connected to an interconnecting microchannel 82. In one embodiment, the interconnecting channel 82 comprises a seal 83 so that the reservoir 62 can be charged with a gas or fluid and then sealed so that there is substantially no gas or fluid flow in or out of the reservoir 62. The seal 83 may be composed of any sealable material such as plastics or metals. Such an inlet port allows for the reservoir 62 to be charged with a gas or fluid after manufacturing of the device. In an alternative embodiment, the device is manufactured under pressure and is charged with a gas from the atmosphere under which it is manufactured without the need for a separate inlet port. The inlet port is optionally divertably connected to an external source of fluid or gas from which the reservoir compartment may be charged.

[0078] In one exemplary embodiment, the rupture disk is comprised of a thermal plastic diaphragm 66 in communication with an electric heater actuator 68 fabricated on the polymeric diaphragm 66. Upon the application of electric current to the electric heater, the polymeric diaphragm melts and allows the pressurized gas or fluid in the reservoir 62 to discharge through the microfluidic interconnection 74. The delivered pressurized gas or fluid works then as an alternative pressure source to provide motive force for pumping fluids or activation of microvalves or other power requiring microactuators.

[0079] In another embodiment, the rupture disk 64 is comprised of a diaphragm made from a low-melting temperature metallic material comprising one or more metals selected from the group consisting of tin, copper, silver, bismuth, indium, lead, magnesium, zinc, gallium, aluminum, antimony and alloys thereof. In one exemplary embodiment, the low-melting diaphragm is composed of a low-melting point metallic material with a melting point less than about 300°C. In another embodiment, the low-melting point diaphragm is composed of an electrically conductive alloy having a melting point less than about 300°C. In another embodiment, the low-melting diaphragm is composed of an electrically conductive alloy having a melting temperature from about 150 to about 250°C. In another embodiment, the low-melting diaphragm is composed of an electrically conductive alloy having a melting temperature from about 175 to about 215°C. See U.S. Pat. Nos. 5,328,660, 5,538,686, and 5,580,520, incorporated herein by reference. [0080] In one exemplary form, the structurally programmable microfluidic systems of the present invention utilize the pneumatic air bursting actuators to pump fluids such as sample fluid, buffer and analytic compounds through the fluidic device upon activation of the burst activator. In another embodiment, a series of pneumatic air bursting actuators can be used increasing stepwise the pressure applied to a sample or other fluid to the device under controlled conditions. For example, upon activation of a first air bursting actuator, a pressure can be applied to a fluid to fill the main branch 14 and first secondary branch 16 of the fluidic device. Upon appropriate signaling, a second pneumatic air bursting actuator can be activated to apply an appropriate additional amount of pressure to overcome the pressure drop of the restrictions 18 and 22 thereby filling the next level of programmed branches.

[0081] In another embodiment, the activator 68 is placed within the reservoir

62 wherein the actuator is a heating means by which the temperature inside the reservoir 62 can be regulated. The heating means may be a resistance type heating means or any type of suitable heating means known in the art. Upon activation, the temperature of the heater increases thereby heating the contents of reservoir 62 producing an increase of pressure of the contents thereby causing a downward deflection of rupture disk 64. In another embodiment, such rupture disk may comprise a frangible wall portion. Such frangible wall portion may be scored to facilitate bursting. Upon actuation, the temperature of the heater increases thereby heating the gas or fluid contents of reservoir 62 increasing the pressure thereof and producing a downward deflection of the frangible rupture disk 64 and displacing fluid from the reservoir compartment 62 into interconnection microchannel 74 and thereafter into an appropriate processing compartment. The pressure inside reservoir 62 is generally chosen to provide an appropriate pulse of pressure upon actuation. [0082] The contents of reservoir 62 are generally any gas or fluid compatible with the microfluidic device of the present invention. I-n an exemplary embodiment, a gas is used. Typical suitable gasses include N₂, CO₂, NO₂, NH₃, 1₂, O₂, or other inert gas such as the noble gasses.

[0083] As will be appreciated, various layer fabrication techniques can be employed for definition of the laminate or microfluidic device structures in accordance with the present invention. For example, microdevice sensors or actuators can be created through surface etching of microdevice structural components on two or more laminate strips, followed by sandwiching the strips to form a working microdevice embedded in the laminate layers. Alternatively, in certain embodiments the use of sacrificial layers during laminate construction is contemplated. After layering additional laminate layers on top of the sacrificial layer (typically photoresist, an easily erodable plastic, or a chemically etchable material), the sacrificial layer can be partially or completely removed to leave freestanding, suspended or movable extended microactuation structures such cantilever beams, movable flaps, louvers, and diaphragms (which can be considered equivalent to a cantilever beam pinned at all edges). In still other layering techniques, selective or patterned deposition can be used to partially or completely construct laminate layers. Alternatively, chambers or receptacles for holding microdevices or microdevice components can be created by the sandwiching or sacrifice techniques as previously described, drilling, punching, mold forming, or any other conventional technique known to those skilled in the art. Microdevices can then be formed, deposited, emplaced or otherwise embedded in the laminate defined chamber. Similar apparatus and methods of fabricating microfluidic devices are also taught and disclosed in U.S. Pat. Nos. 5,858,195, 5,126,022, 4,891,120, 4,908,112, 5,750,015, 5,580,523, 5,571,410, 5,885,470. BioChips

[0084] A typically biochip including microfluidic structurally programmable fluidic path system with passive microvalves is schematically illustrated in FIG. 1. The devices of the present invention include structurally programmable fluidic paths, passive microvalves, fluidic components based on hydrophobic microfluidic systems, and pneumatic actuator using an air-bursting actuation concept as illustrated in FIG. 1.

[0085] Generally, the top planar surface of the substrate is then mated, e.g., placed into contact with, and bonded to the planar surface of the bottom substrate, covering and sealing the grooves and/or indentations 16 in the surface of the bottom substrate, to form the channels and/or chambers (i.e., the interior portion) of the device at the interface of these two components. The substrate is provided with holes in the top portion of the device oriented such that they are in communication with at least one of the channels and/or chambers formed in the interior portion of the device from the grooves or indentations in the bottom substrate. In the completed device, these holes function as reservoirs for facilitating fluid or material introduction into the channels or chambers of the interior portion of the device, as well as providing ports at which electrodes may be placed into contact with fluids within the device, allowing application of electric fields along the channels of the device to control and direct fluid transport within the device.

[0086] These devices may be used in a variety of applications, including, e.g., the performance of high throughput screening assays in drug discovery, immunoassays, diagnostics, genetic analysis, and the like. As such, the devices described herein, will often include multiple sample introduction ports or reservoirs, for the parallel or serial introduction and analysis of multiple samples. Alternatively, these devices may be coupled to a sample introduction port, e.g., a pipettor, which serially introduces multiple samples into the device for analysis. Examples of such sample introduction systems are described in e.g., U.S. Pat. No. 6,045,056 and U.S. Pat. No. 5,880,071, and is hereby incorporated by reference in its entirety for all purposes.

[0087] Microfluidic systems have been employed in the separation of biological macromolecules, in the performance of assays, e.g., enzyme assays, imriiunoassays, receptor binding assays, and other assays in screening for affectors of biochemical systems. Generally, such systems employ microscaie channels and/or chambers through which various reactants are transported, where they may be mixed with additional reactants, subjected to changes in temperature, pH, ionic concentration, etc., separated into constituent elements and or detected.

[0088] Although the devices and systems specifically illustrated herein are generally described in terms of the performance of a few or one particular operation, it will be readily appreciated from this disclosure that the flexibility of these systems permits easy integration of additional operations into these devices. For example, the devices and systems described will optionally include structures, reagents and systems for performing virtually any number of operations both upstream and downstream from the operations specifically described herein. Such upstream operations include sample handling and preparation operations, e.g., cell separation, extraction, purification, amplification, cellular activation, labeling reactions, dilution, aliquoting, and the like. Similarly, downstream operations may include similar operations, including, e.g., separation of sample components, labeling of components, assays and detection operations. Assay and detection operations include without limitation, probe interrogation assays, e.g., nucleic acid hybridization assays utilizing individual probes, free or tethered within the channels or chambers of the device and/or probe arrays having large numbers of different, discretely positioned probes, receptor/ligand assays, immunoassays, and the like.

[0089] The systems described herein generally include microfluidic devices, as described above, in conjunction with additional instrumentation for controlling fluid transport and direction within the devices, detection instrumentation for detecting or sensing results of the operations performed by the system, processors, e.g., computers, for instructing the controlling instrumentation in accordance with preprogrammed instructions, receiving data from the detection instrumentation, and for analyzing, storing and inteφreting the data, and providing the data and inteφretations in a readily accessible reporting format. A variety of controlling instrumentation may be utilized in conjunction with the microfluidic devices described above, for controlling the transport and direction of fluids and/or materials within the devices of the present invention. As noted above, the systems described herein preferably include structurally programmable fluidic paths, passive microvalves, fluidic components based on hydrophobic microfluidic systems, and pneumatic actuators using an air-bursting actuation concept as illustrated in FIG. 1.

[0090] As such, the controller systems for use in conjunction with the microfluidic devices typically include an electrical or mechanical power supply and circuitry for concurrently delivering appropriate signals to a plurality of electrodes that are placed in electrical contact with the actuators contained within the microfluidic devices.

[0091] In the microfluidic systems described herein, a variety of detection methods and systems may be employed, depending upon the specific operation that is being performed by the system. Often, a microfluidic system will employ multiple different detection systems for monitoring the output of the system. Examples of detection systems include optical sensors, temperature sensors, pressure sensors, pH sensors, conductivity sensors, and the like.

[0092] Integration of these functions into a single unit facilitates connection of these instruments with the computer by permitting the use of few or a single communication port(s) for transmitting information between the controller, the detector and the computer. As noted above, either or both of the controller system and/or the detection system are coupled to an appropriately programmed processor or computer which functions to instruct the operation of these instruments in accordance with preprogrammed or user input instructions, receive data and information from these instruments, and inteφret, manipulate and report this information to the user.

[0093] The computer typically includes appropriate software for receiving user instructions, either in the form of user input into a set parameter fields or in the form of preprogrammed instructions. The software then converts these instructions to appropriate language for instructing the operation of the fluid direction and transport controller to carry out the desired operation.

[0094] The computer portion of the system is capable of performing a number of functions in the context of the overall microfluidic system, generally, and specifically with respect to the monitoring and control methods described herein. Specifically, the computer typically includes appropriate programming for instructing the application of voltages to the device in order to carry out a desired fluid transport profile, which is either input by the user, or is contained in a separate program.

[0095] FIG. 6 illustrates a pneumatic air-bursting microactuator as an alternative power source in a microfluidic system. The pneumatic air- bursting actuator consists of a pressurized reservoir with a passive inlet port that provides the injection and sealing of pressurized air, and a plastic microvalve collapsible by heat, releasing an air burst from the pressurized reservoir. The bursting of pressurized air by electrically heating a small thermoplastic valve is analogous to the explosion of dynamite by igniting a fuse, so, very strong forces can be achieved from these pressurized reservoirs. FIG. 7 illustrates a schematic structure of a pneumatic air-bursting microactuator for the case of silicon bulk processing and patterned thermoplastic with "detonating" heater. A pneumatic air-bursting microactuator can be fabricated either using silicon bulk processing or plastic micromachining including plastic injection techniques and embossing techniques. A pneumatic air-bursting microactuator can be fabricated using through- etched silicon wafer and SU-8 thermoplastic material. Once electric currents applied, then the "detonating" heater will increase heat and so polymer diaphragm will be melt down injecting pressurized air (or inert gas) into microfluidic channels working as an alternative pressure source to pump and deliver.

In realizing a stand-alone microfluidic biochip, the development of a power supply for the smart control of chips has been considered one of the most difficult and challenging tasks. Although an appropriate chip power supply like a chip-battery or a commercially available battery (e.g., cadmium, alkaline, NiMH, lithium ion) can be mounted on standalone biochips with active microfluidic components, the advantage of low-cost plastic chips will vanish due to the high cost of batteries or the extra packaging costs. Furthermore, the chip-battery usually requires an additional voltage regulation circuit to maintain a constant out voltage, and the totally available power from chip-batteries is usually not sufficient for driving the active microfluidic components. So, an on-chip battery in a disposable plastic microfluidic biochip is not an optimized solution for cost reduction, although there have been numerous efforts to reduce the cost of IC circuits and power supplies mounted over plastic chips. Flip-chip or embedding techniques will be also an alternative approach to reduce the cost, but both techniques are still not very good at address the low cost issues. We need an innovative method to address this painful bottleneck towards the realization of a low-cost plastic disposable biochip.

[0097] Specifically, if microreservoirs in the biochips are discretely pressurized with air or gases, which are tuned with the already structurally programmed microfluidic systems, the fluid control of the microfluidic system can be achieved by detonating the microreservoirs as electronically programmed.

Microfluidic Interconnections

[0098] The present invention also provides for novel self-aligning fluidic interconnections and methods of making such interconnections, namely the serial assembly method and the parallel assembly method, both resulting in low dead volume and low pressure drop across the interconnects. The serial assembly technique is especially useful for microfluidic systems that require low number of interconnects, such as CE (capillary electrophoresis) chips and the parallel technique is especially useful for generic biochemical microfluidic analysis systems that require high-density microfluidic interconnects.

[0099] The present invention provides a fully integrated, self-aligning microfluidic interconnection technique with low dead volume for applications in microfluidic systems. The present invention discloses the use of thermoplastic or other suitable polymer tubing in conjunction with a concentric inlet structure on a microfluidic platform. Generally, any thermoplastic polymer or thermoset polymer can be used if it has the required properties to be deformed within the forming process and remain intact during use and can be formed without undue harm to the substrate. Such polymers include, but are not limited to: acetate rayon, acrylic resins, acrylonitrile-butadiene- styrene (ABS) resins and acrylic resins, aliphatic and aromatic polyamides, aliphatic and aromatic polyesters, allyl resin, (Allyl), AS resins, butadiene resins, chlorinated polyethylene, conductive resins, copolymerised polyamides, copolymers of ethylene and vinyl acetate, cuprammonium rayons and natural and synthetic rubbers, EEA resins, epoxy resins (e.g. , bisphenol, dihydroxyphenol, and novolak), ether ketone resins, Ethylene vinyl alcohol, (E/NAL), fluorine resins, fluorocarbon polymers, fluoroplastics, (PTFE), (FEP, PFA, CTFE, ECTFE, ETFE), high density polyethylenes, ionomer resins, Liquid Crystal Polymer, (LCP), low density polyethylenes, Melamine formaldehyde, (melamine resins), natural polymers such as cellulosics, nylons, Phenol-formaldehyde Plastic, (PF) phenolic resins, Polyacetal, (Acetal), Polyacrylates, (Acrylic), Polyacrylonitrile, (PAN), (Acrylonitrile), Polyamide, (PA), (Nylon), Polyamide-imide, (PAI), Polyaryletherketone, (PAEK), (Ketone), Polybutadiene, (PBD), polybutylene terephthalate, Polybutylene, (PB), Polycarbonate, (PC), polycarbonates, Polydicyclopentadiene, (PDCP), Polyketones, (PK), polyester block copolymers, polyesters, polyesterurethane, polyesterurethaneurea, polyether and polyester block polymers, polyether ketoneketone (PEKK), polyetherether ketone (PEEK), Polyetherimide, (PEI), polyethers, Polyethersulfone, (PES), polyetherurethane, polyetherurethaneurea, polyethylene isophthalate, polyethylene terephthalate, Polyethylene, (PE), Polyethylenechlorinates, (PEC), polyglycohc acid, polyhexamethylene terephthalate, Polyimide, (PI), polylactic acid, Polymethylpentene, (PMP), poly-m-phenylene isophthalamide, polyolefins, Polyphenylene Oxide, (PPO), Polyphenylene Sulfide, (PPS), Polyphthalamide, (PTA), poly-p-phenylene terephthalamide, Polypropylene, (PP), polysiloxanes such as polydimethyl siloxane, Polystyrene, (PS), polysulfides, Polysulfone, (PSU), polytetrafluoroethylene, Polyurethane, (PU), polyvinyl acetate, polyvinyl alcohols, Polyvinylchloride, (PVC), Polyvinylidene Chloride, (PVDC), polyvinylidene fluoride and polyvinyl fluoride, rayon, reconstituted silk and polysaccharides, reinforced polyethylene terephthalate resins, segmented polyurethane elastomers, silicone resins, spandex or elastane elastomers, styrene type specific resins, thermoplastic polyurethane elastomers, thermosetting synthetic polymers such as phenol-formaldehyde copolymer, triacetate rayon, unsaturated polyester resins, urea resins, urethane resins, vinyl chloride resins, vinyl polymers, and vinylidene chloride resins. This group includes copolymers, teφolymers and mixtures of the species listed.

[00100] Thermoplastic resins suitable for use in as the tubings of this invention must have a glass transition temperature ("Tg") less than about 550°C, preferably less than about 500°C, and more preferably less than about 400°C, and most preferably less than about 350°C. Examples of suitable thermoplastic resins for use in the practice of the invention include polyamides, polyesters, cellulose esters, polyethylene, polypropylene, poly (vinyl chloride) or PVC, poly (vinylidene fluoride) or PVF2, polyphenylsulfones and polytetrafluoroethylene or PTFE.

1

[00101] Microfabrication of polymeric substrates for use in the devices of the invention maybe carried out by a -variety of well-known methods. In particular, polymeric substrates may be prepared using manufacturing methods that are common in the microfabrication industry, such as injection molding or stamp molding/embossing methods where a polymeric substrate is pressed against an appropriate mold to emboss the surface of the substrate with the appropriate channel structures. Other suitable microfabrication techniques are also suitable for preparation of polymeric substrates, including, e.g., laser drilling, etching techniques, and photolithographic techniques. Such photolithographic methods generally involve exposing the polymeric substrate through an appropriate photolithographic mask, i.e., representing the desired pattern of channels and chambers, to a degradative level of radiation, e.g., UV light for set periods of time. The exposure then results in degradation of portions of the surface of the substrate resulting in the formation of indentations that correspond to the channels and/or chambers of the device.

[00102] Polymeric substrate materials may be rigid, semi-rigid, or non-rigid, opaque, semi-opaque or transparent, depending upon the use for which they are intended. For example, devices which include an optical or visual detection element, e.g., for use in fluorescence based or colorimetric assays, will generally be fabricated, at least in part, from a transparent polymeric material to facilitate that detection. Alternatively, transparent windows of, e.g., glass or quartz, maybe incoφorated into the device to allow for these detection elements. Additionally, the polymeric materials may have linear or branched backbones, and may be cross-linked or non-cross-linked. Examples of preferred polymeric materials include, e.g., polyamide, polyester, cellulose esters, polyethylene, polypropylene, poly(vinyl chloride), poly(vinylidene fluoride), polyphenylsulfones, polytetrafluoroethylene. polymethylmethacrylate (PMMA), polycarbonate, polydimethylsiloxane (PDMS), polystyrene, polysulfone, polyurethane and the like.

[00103] While the polymeric substrates used in the devices of the present invention can be fabricated as one piece, they are generally fabricated in two or more parts. Specifically, a first planar substrate element is provided having one or more grooves and/or wells, corresponding to the fluid channels and/or chambers, manufactured, e.g., molded or machined, into one of its planar surfaces. These grooves provide the bottom and side walls of the channels and chambers of the devices. A second planar substrate element is then mated with the first to define the top wall of the channels and chambers. The two members are bonded together in order to ensure that the channels and chambers in the substrate are fluid tight. Bonding of the two members may be accomplished by a number of methods that are known in the art, such as through the use of adhesives, e.g., UV curable adhesives, or by sonically welding one member to the other. Alternatively, the two planar elements maybe bonded by applying pressure to the joined pair under elevated temperatures, sufficient to bond the two planar elements together.

[00104] This invention also discloses the uses of a polymeric (or plastic)

"holding layer" for interconnections. The microfluidic platforms are formed using conventional photolithography, etching and bonding techniques. The inlets and outlets of the microfluidic channels (microchannels) have a concentric-cavity structure wherein, the cavity adjacent to the microchannels is concentric and has a larger size than the cavity adjacent to the surface of the holding layer. This larger size allows for a flange area of the cavity to be produced beyond the edges of the smaller cavity. In one embodiment, the concentric-cavity structure is realized by interchanging different sized drill bits at the same drill position to achieve perfect alignment. Optionally, one or more air- vents can be drilled in the flange area of the larger cavity to allow trapped air to escape.

[00105] In one embodiment of this invention, namely the serial assembly technique, deformable flanging polymer tubing is placed into the concentric cavities at its glass transition temperature ("Tg"), under appropriate controlled temperature and pressure conditions within this concentric-cavity structure so that the flanging polymer tubing is deformedly expanded into the flange areas of the larger concentric cavity. Generally, a substantial portion of the flange area is filled with the flanging polymer tubing material creating a flanged, polymer bushing within the concentric cavities and leaving an interior channel adapted for receiving a second polymer tubing. Preferably, the tubing is thermoplastic. Optionally, the flanging polymer tubing has a rigid insert on the inside of tubing that has an outside diameter that substantially fills the inside diameter of the flanging polymer tubing to prevent inward deformation of the flanging polymer tubing.

[00106] Under these conditions, the tubing material deformedly and substantially fills the concentric-cavity structure including the flanged areas. This realizes a self-aligned microfluidic interconnection. A second polymer tubing, namely the interconnection polymer tubing, is then inserted within the inner cavity defined by the flanging polymer tubing bushing. Preferably, the outside diameter of the interconnection polymer tubing is substantially the same as the inside diameter of the flanging polymer tubing bushing thereby substantially mating the surfaces of these two polymer tubes. In one embodiment, the interconnection tubing is substantially matched to the bushing so that the two polymers remain in a substantially leak-free joining due to pressure and friction. In another embodiment, the two polymers are of appropriate composition so that under appropriate temperature and pressure conditions, the polymers become substantially thermally bonded. In yet another embodiment, an appropriate adhesive, bonding or sealing agent is used between the surfaces of the polymers.

[00107] In another embodiment, namely the parallel assembly technique, the flanging polymer tubing is pre-formed into an appropriate bushing shape substantially conforming to a mating concentric-cavity structure. The tubing is then inserted within the concentric-cavity structure in a thermoplastic substrate. The substrate is then bonded with the microfluidic platform thus generating multiple interconnects simultaneously. [00108] In one embodiment, the tubing-substrate arrangement is then sandwiched between rigid supports and pressurized in a heated state. The thermoplastic substrate is chosen to have a glass transition temperature below that of the flanging polymer tubing material. Under these conditions the flanging polymer tubing fuses with the substrate to form a substantially leak-free bond.

[00109] FIG. 8 is a schematic cross-sectional view of the microfluidic interconnects assembled using the serial interconnection technique. In one embodiment, the microfluidic device is created on a two-wafer substrate structure. Microfluidic channels are formed on wafer 110, generally using conventional photolithography and etching techniques as well known in the art. At least one inlet 116 is defined on wafer 112, generally using mechanical boring (e.g., a diamond coated drill bit), photolithography, etching, or other appropriate hole forming method. Optionally, at least at least one outlet 118 is also defined on wafer 112. Wafers 110 and 112 are then bonded together using adhesives, glass-to-glass fusion, or other appropriate bonding technique as known in the art to form the microfluidic device. In another embodiment (not shown) the microfluidic device is formed within a single substrate.

[00110] The position of the inlets and outlets is also defined photolithographically on an additional wafer, the holding layer 120, in the serial assembly technique. In one embodiment, the concentric cavities can be formed by drilling two sets of holes for each at least one inlet and at least one outlet on wafer 120. First hole 130 is drilled such that it passes through the entire thickness of the wafer 120. Next the drill bit for hole 130 is removed and replaced by drill bit having a larger diameter for hole 132. The position of the wafer relative to the drilling machine chuck is maintained substantially the same between the changes of drill bits. Next a second hole 132 is drilled substantially concentric with hole 130 such that it extends partly through the thickness of the wafer 120. This creates the concentric-cavity structure shown in FIG. 10a to create a flange 128 of the tubing 124. The concentric-cavity structure can be created by various means including, but not limited to, diamond coated drill bits, etching, chemical treatment, laser machining, and lamination. Optionally, a smaller hole 134 is drilled into the flange area of the cavity 132 that passes through the entire thickness of the wafer 120. This hole forms the air vent to release air as the flanging polymer tubing 124 extends into space of hole 132 to create the tubing flange 128 of the flanging polymer tubing bushing.

[00111] Wafer 120 is then flipped over such that the now bushing filled larger cavity 132 and the flanging polymer tubing bushing 124 are in contact with the top surface of the microfluidic device, i.e., wafer 112. The two wafers are aligned and bonded using glass-to-glass fusion bonding or other suitable bonding method as well known in the art. The alignment accuracy in this step determines the alignment accuracy of the interconnects.

[00112] After assembling the wafer stack by fusion bonding, the entire assembly can then be heated to a temperature corresponding to the glass transition temperature of the thermoplastic tubing 124. Optionally, tubing 124 has a rigid insert 140 as shown in FIG. 10b during interconnect forming stage to prevent inward deformation. When thermoplastic tubing (e.g. PEEK) is heated at about its transition temperature, it will soften and deform outwards, under applied pressure, filling up the larger cavity 132 in the so-called "flanging" operation. One or more optional air- vents 134 help at this stage as they allow any trapped gas to escape from the flanging cavity 132 and softened flanging polymer tubing bushing material can now occupy substantially the entire cavity. As is evident from the sequence of operation, this technique is self-aligning. After the flanging operation, the rigid insert 140 is withdrawn (if used) and a second interconnection polymer tubing 126 is force-fitted within tubing 124. Interconnection polymer tubing 126 is chosen such that it has an internal diameter 114 corresponding generally to the microchannel 116 dimension. This minimizes dead volume as well as pressure drop across interconnect.

[00113] FIG. 9 is a schematic cross-sectional view of the microfluidic interconnects assembled using the parallel interconnection technique. The microfluidic device is created within wafers 110 and 112 using a technique similar to those listed above. In one embodiment, a TEFLON® like polymer is spin-coated on either surface of the microfluidic platform. In this application at least one inlet hole 116 is formed in wafer 112 and the at least one outlet hole 18 is formed in wafer 110 as shown in FIG. lie.

[00114] In another embodiment, the position of the cavities is defined photolithographically on two holder layers. Generally, the holder layers comprise a plastic or other polymer material but may be made of any suitable material compatible with the techniques employed herein. One substrate 120 has cavities corresponding to wafer 112 and the second substrate 122 has cavities corresponding to wafer 110. The concentric-cavity structure is made at the appropriate location, for example by using a flat-end drill bit. However, the concentric-cavity structures may be created by various methods as known in the art.

[00115] Next, thermoplastic flanging polymer tubing 124 is flanged externally, e.g., by using a flanging machine, wherein the flanging polymer tubing 124 is deformedly made under applied heat and/or pressure to create flange 128. Next, the flanging polymer tubing bushing 124 is inserted through the concentric-cavity structure as shown in FIG. 11a. This assembly is then heated to the glass transition temperature of the substrate. Generally, the support layer 120 has a significantly lower glass transition temperature than the flanging polymer tubing bushing 124. Preferably, the support layer 120 has a glass transition temperature that is at least from about 20°C lower than the flanging polymer tubing bushing 124, preferably at least from about 25°C lower than the flanging polymer tubing bushing 124, more preferably at least from about 30°C lower than the flanging polymer tubing bushing 124, most preferably at least from about 50°C lower than the flanging polymer tubing bushing 124. Under these conditions when pressure is applied to the assembly, as shown in FIG. 11a, the flanging polymer tubing bushing 124 fuses with the substrate 120 to form a substantially leak-free joint. Furthermore, this treatment also renders the surface of the plastic substrate substantially smooth and flat.

[00116] Next, both tubing/substrate assemblies are positioned and bonded to the mating surface of the microfluidic platform. This creates aligned microfluidic interconnects on both sides of the microfluidic platform. This arrangement can then be used to realize stacked microfluidic assemblies wherein, the tubing from one microfluidic assembly substantially aligns with the tubing on another microfluidic platform with a clamping tubing in between.

[00117] FIG. 12e is a schematic cross-sectional view of the microfluidic interconnects assembled using the parallel interconnection technique for a microfluidic platform. Generally, the microfluidic channels are created on the substrates by hot embossing. Typically, the substrates are made from plastic or other suitable polymer. The position of the holes is defined photolithographically on the two substrates. In this application the at least one inlet hole is drilled in substrate 120 and the at least one outlet hole is formed in substrate 122 as shown in FIG. 12e. In this case too, the concentric-cavity structure may be created by various methods. [00118] Next, flanging polymer tubing 124 is flanged externally, e.g., by using a flanging machine, wherein the flanging polymer tubing 124 is deformed to form a flanging polymer tubing bushing having a flange 128 under applied heat and/or pressure. Next, the flanging polymer tubing bushing 124 is inserted through the concentric-cavity structure as shown in FIG. 12a. Preferably, this assembly is then heated to about the glass transition temperature of the substrate 120. The layer 120 has a significantly lower glass transition temperature than the flanging polymer tubing bushing 124. Under these conditions when a high pressure is applied to the assembly, as shown in FIG. 12b, the flanging polymer tubing bushing 124 fuses with the substrate 120 for a substantially leak-free joint. Furthermore, this treatment also renders the surface of the substrate substantially smooth and flat. Also, this step is used to generate at least one microfluidic channel 118 in at least one of the substrates 120 and 122 wherein, an embossing master 144, as shown in FIG. 12c, replaces the support layer 140.

[00119] Next, the tubing/substrate assemblies are positioned and bonded, typically using thermoplastic bonding. This creates aligned microfluidic interconnects on both sides of the microfluidic platform. This arrangement too, can then be used to realize stacked microfluidic assemblies wherein, the tubing from one microfluidic assembly substantially aligns with the tubing on another microfluidic platform with a clamping tubing in between.

[00120] A partial list of exemplary thermoplastic polymers that can be used for the serial and/or parallel interconnects disclosed in this application include: Table 1.

[00121] The value of the glass transition temperature shown in Table 1 is highly dependent on exact formulation, added materials, processing conditions during manufacture. The stated range reasonably covers the bulk material with a broad tolerance level. The biocompatibility levels vary significantly based on whether only the stated material is used or additives are used to modify the chemical properties of the material.

[00122] The processing conditions for serial interconnects generally comprise

(a) a glass-glass fusion bonding at a temperature of from about 400 to about 680 C, preferably at a temperature of from about 550 to about

650 C, more preferably at a temperature of from about 600 to about 640 ^°C; (b) a thermoplastic tubing glass transition for flanging at a temperature of from about 30 to about 500 ^°C, preferably at a temperature of from about 60 to about 400 C, more preferably at a temperature of from about 75 to about 350 ^°C, and is dependent upon material; and (c) the pressure during flanging is typically from about 0.1 MPa to about 5 MPa, preferably from about 0.25 MPa to about 2.5 MPa, more preferably from about 0.5 MPa to about 1.5 MPa.

[00123] The processing conditions for parallel interconnects generally comprise (a) a glass-glass fusion bonding (which forms the microfluidic platform) and can be by (i) fusion bonding (direct bonding) at a temperature of from about 400 to about 680 C, preferably at a temperature of from about 550 to about 650 C, more preferably at a temperature of from about 600 to about 640 C and at a pressure of from about 0.1 MPa to about 5 MPa, preferably from about 0.25 MPa to about 2.5 MPa, more preferably from about 0.5 MPa to about 1.5 MPa; or by (ii) interfacial layer bonding (indirect bonding), e.g., by use of a thermoplastic interfacial layer that is spin-coated on the wafers, or by UV curable adhesive, etc, at a temperature of from about 60 to about 400 ^°C, preferably at a temperature of from about 80 to about 360 C, more preferably at a temperature of from about 100 to about 340 ^°C and at a pressure of from about 0.1 MPa to about 5 MPa, preferably from about 0.25 MPa to about 2.5 MPa, more preferably from about 0.5 MPa to about 1.5 MPa; (b) a thermoplastic tubing glass transition for flanging at a temperature of from about 30 to about 500 ^°C, preferably at a temperature of from about 60 to about 400 ^°C, more preferably at a temperature of from about 75 to about 350 ^°C, and is dependent upon material and at a pressure of from about 0.1 MPa to about 5 MPa, preferably from about 0.25 MPa to about 2.5 MPa, more preferably from about 0.5 MPa to about 1.5 MPa; (c) heat/pressure treatment to form a substantially leak proof bond and to create a substantially flat surface on the substrate where the tubing is pressed flush with the surface at a temperature of from about 30 to about 350 C, preferably at a temperature of from about 60 to about 300 C, more preferably at a temperature of from about 75 to about 250 ^°C, and is dependent upon material and at a pressure of from about 0.05 MPa to about 30 MPa, preferably from about 0.1 MPa to about 25 MPa, more preferably from about 0.2 MPa to about 20 MPa; and (d) bonding of the holder layers to the bonded microfluidic assembly using an interfacial layer at a temperature of from about 35 to about 250 ^°C, preferably at a temperature of from about 45 to about 200 ^°C, more preferably at a temperature of from about 55 to about 165 ^°C, and is dependent upon material and at a pressure of from about 0.01 MPa to about 5 MPa, preferably from about 0.05 MPa to about 2.5 MPa, more preferably from about 0.1 MPa to about 0.75 MPa;

Although the present invention has been discussed with respect to the preferred and alternative embodiments, it will be apparent to those skilled in the art that the present invention is not limited to these embodiments. Therefore, a person of ordinary skill in the art will understand that variations and modifications of the present invention are within the spirit and scope of the present invention.

Claims

What is claimed:

1 A structurally programmable microfluidic device comprising: a) a body structure comprising a substrate having at least one first fluid flow channel therein having a first and second end; b) at least two second fluid flow branch channels in fluid communication with the second end of at least one first fluid flow channel; c) a flow controller for applying a motive force operably coupled to first end of the at least one first fluid flow channel so as to controllably direct material flow through the channels d) wherein at least one said second fluid flow branch channel has at least one fluid flow passive restriction operatively associated with the channel for selectively restricting fluid flow to the branch channels; e) wherein the first fluid flow channel is capable of receiving a fluid sample and is further capable of conducting the fluid sample to the branch channels upon application of motive force; and f) wherein the fluid flow passive restriction causes the branch channels to selectively fill with fluid in a predetermined order upon application of the motive force.

2 The device of claim 1, wherein the flow controller comprises at least one actuator for applying at least one predetermined motive force to flow a detectable amount of at least one fluid from the source of fluid into the first channel during operation of the device, and for flowing fluid into at least one second channel branch.

3 The device of claim 1, wherein the first channel is one member of an array of interconnected channels disposed within the body structure.

4 The device of claim 1, wherein the first channel comprises at least one cross- sectional dimension between about 0.1 microns and about 500 microns. The device of claim 1, wherein the first channel comprises at least one cross- sectional dimension between about 0.1 microns and about 100 microns.

The device of claim 1, wherein the first channel comprises at least one cross- sectional dimension between about 0.1 microns and about 50 microns.

The device of claim 1, wherein the device further comprises one or more fluid in the channel.

The device of claim 1, wherein the fluid comprises a liquid containing a nucleic acid or a protein.

The device of claim 1, wherein the body structure comprising an aggregation of two or more layers.

The device of claim 1, wherein the body structure comprises at least one material selected from the group consisting of glass, quartz, silicon, polysilicon, a polymer, a plastic, polymethylmethacrylate, polycarbonate, polytetrafluoroethylene, polyvinylchloride, polydimethylsiloxane, and polysulfone.

The device of claim 1, wherein the surface properties of the channels are further modified by surface treatment.

The device of claim 1, wherein the surface treatment is a method of surface treatment selected from the group consisting of methods of physical surface adsoφtions, methods of covalent bonding of selected moieties to functional groups on the surface of channel substrates, methods of coating surfaces, methods of plasma treatment to activate or inactivate surface properties, methods of coating the surfaces of channels with bioaffϊnity reagents, methods of polymer grafting to the surface of channel substrates, and methods of thin-film deposition of materials to channel substrates.

The device of claim 1, wherein the substrate additionally comprises at least one sample-handling region in fluid communication with at least the first fluid flow channel.

The device of claim 1, wherein the additional branch channels are in fluid communication with the second fluid flow branch channels.

The device of claim 1, wherein the motive force is supplied by a pneumatic actuator.

The device of claim 1, wherein the pneumatic actuator is a reservoir comprising: a) at least one chamber substantially containing at least one gas or fluid under pressure; b) an outlet port for release of at least a portion of the pressurized gas or fluid within the reservoir; c) a rupture disk; d) wherein the rupture disk is ruptured upon activation.

The device of claim 1, wherein the actuation is by direct application of force on the rupture disk.

The device of claim 1, wherein the pneumatic actuator further comprises an activator operatively connected to the rupture disk.

The device of claim 1, wherein the activator is capable of initiating the rupturing of the rupture disk upon communication of a signal along one or more signal paths by a suitable controller. The device of claim 1, wherein the activator is a heater.

A device as defined in claim 1 wherein the heater includes a thin film of resistance heating material on an inner surface or embedded on rupture disk of the reservoir and includes means for connecting a pulsed electric power source to said film.

A structurally programmable microfluidic system according to claim 1 wherein said a plastic-based disposable cartridge can be inserted into a host electronic host analyzer system.

A structurally programmable microfluidic system according to claim 1 wherein said a passive valve can include controllable hydrophobicity using an electrokinetic component.

The device or system of claim 1, further comprising at least one reservoir disposed in the body structure fluidly coupled to the first channel, to flow a fluid into the first channel.

The device or system of claim 1, further comprising at least one set of additional fluid flow branch channels disposed in the body structure, the subsequent channels intersecting at least one of the previous fluid flow branches.

The device or system of claim 1, wherein the at least one set of additional fluid flow branch channels is fluidly coupled to a reservoir.

The device of claim 1, wherein the device is formed by micromachining.

The device of claim 1, wherein the micromachining is by one or more methods selected from the group consisting of etching, laser ablation, LIGA, and injection molding. A device according to claim 1, wherein the body structure is a microchip.

A device of claim 6, wherein the microchip is a biochip for DNA sample processing or sequence identification or clinical diagnostics including blood analysis.

The device of claim 1, wherein the device is plastic-based disposable cartridge type biochip comprising at least one structurally programmable fluidic path array; at least one hydrophobic passive valve; and at least one pneumatic air-bursting microactuator and is adapted for mixing and analysis of analytes.

A microfluidic interconnection system comprising: a. a body structure comprising a substrate having a top surface and a bottom surface; b. wherein the substrate comprised at least one flanging cavity fabricated therein wherein the cavity has an open upper end in communication . with the top surface of the substrate and a substantially closed lower end; c. wherein the open upper end of the flanging cavity is substantially smaller than the bottom surface of the flanging cavity, thereby defining a flange area within the lower area; d. the body structure having at least one microfluidic port fabricated therein, the at least one microchannel being in fluid communication with the lower end of the flanging cavity; e. a first polymer tubing wherein the lengths, cross-sections, or combinations thereof, of the tubing are dimensioned to adaptedly fit through the open upper end of the flanging cavity and extend substantially the depth the flanging cavity; f. wherein the first polymer tubing substantially fills the flange area of the cavity to form a polymer tubing bushing, wherein the bushing has an interior channel adapted for containing a second polymer tubing; g. a second polymer tubing having walls defining an interior microfluidic channel and having an outside diameter adapted to fit within the first polymer tubing bushing and wherein the second polymer tubing is matingly joined with the first polymer tubing; and h. wherein the microfluidic channel is substantially in alignment with and fluidly in connection with the microfluidic port of the body.

33. A microfluidic interconnection system comprising: a. a body structure having at least first and second substrate layers, the first substrate layer disposed on top of the second substrate layer wherein the first and second substrate layers each comprise a top and a bottom planar surface; b. wherein the first substrate layer comprises at least one flanging cavity wherein the cavity has an open upper end in communication with the top planar surface of the first substrate and an open lower end in communication with the bottom planar surface of the first substrate; c. wherein the open upper end of the flanging cavity is substantially smaller than the open lower end of the flanging cavity, thereby defining a flange area between the cavity walls of the first substrate layer and the top planar surface of the second substrate layer; d. the second substrate layer having at least one microfluidic port fabricated therein, the at least one microchannel being in fluid communication with the top planar surface end of the flanging cavity; e. a first polymer tubing having walls defining an inside channel wherein the lengths, cross-sections, or combinations thereof, of the tubing are dimensioned to adaptedly fit through the open upper end of the flanging cavity and extend substantially the depth the flanging cavity; f. wherein the first polymer tubing substantially fills the flange area of the cavity to form a polymer tubing bushing, wherein the bushing has an interior channel adapted for containing a second polymer tubing; g. a second polymer tubing having walls defining an inside microfluidic channel and having an outside diameter adapted to fit within the first polymer tubing bushing and wherein the second polymer tubing is matingly joined with the first polymer tubing; and h. wherein the microfluidic channel is substantially in alignment with and fluidly in connection with the microfluidic port of the body.

34. The system according to claim 31 or 32, wherein the substrate layers independently comprise two or more layers.

35. The system according to claims 32, 33, or 34 wherein the substrate further comprises at least one air vent in fluid communication with the flange area of the flanging cavity.

36. The system according to claims 32, 33, or 34 wherein the polymer tubing comprises at least one thermoplastic resin.

37. The system according to claim 36, wherein the thermoplastic resin has a glass transition temperature ("Tg") less than about 550°C.

38. The system according to claim 35, wherein the thermoplastic resin has a glass transition temperature ("Tg") less than about 500°C.

39. The system according to claim 36, wherein the thermoplastic resin has a glass transition temperature ("Tg") less than about 400°C.

40. The system according to claim 36, wherein the thermoplastic resin has a glass transition temperature ("Tg") less than about 350°C.

41. The system according to claim 36, wherein the thermoplastic resin is selected from the group consisting of wherein the polymeric material is selected from the group consisting of polyamide, polyester, cellulose esters, polyethylene, polypropylene, poly(vinyl chloride), poly(vinylidene fluoride), polyphenylsulfones, polytetrafluoroethylene. Polymethylmethacrylate, polyetheretherketone, polyamide, polypropylene, polycarbonate, polydimethylsiloxane, polystyrene, polysulfone, and polyurethane.

42. The system according to claims 32, 33, or 34 wherein the substrate comprises one or more materials selected from the group of glass, silicon, metal and polymeric substrates.

43. The system of claim 42, wherein the body structure comprising at least 50% polymeric materials.

44. The system of claim 43, wherein the polymeric material is selected from the group consisting of wherein the polymeric material is selected from the group consisting of polyamide, polyester, cellulose esters, polyethylene, polypropylene, poly(vinyl chloride), poly(vinylidene fluoride), polyphenylsulfones, polytetrafluoroethylene. Polymethylmethacrylate, polyetheretherketone, polyamide, polypropylene, polycarbonate, polydimethylsiloxane, polystyrene, polysulfone, and polyurethane.

45. The system of claim 44, wherein the body structure is formed by micromachining.

46. The system of claim 45, wherein the micromachining is by one or more methods selected from the group consisting of photolithography, etching, bonding, laser ablation, LIGA, injection molding and embossing.

47. A device according to claim 46, wherein the body structure is a microchip.

48. The system of claim 46, wherein the at least one microchannel has a dimension between about 0.1 microns and about 500 microns.

49. The system according to claims 32, 33, or 34 wherein the body structure is fabricated by bonding a first polymeric substrate to a second polymeric substrate, wherein the at least one flanging cavity is formed between the first and second substrate.

50. The system of claim 49, wherein the bonding is by thermal bonding.

51. A method of forming a microfluidic interconnection system comprising: a. forming at least one flanging cavity in a body structure comprising a substrate having a top surface and a bottom surface; b. wherein the cavity has an open upper end in communication with the top surface of the substrate and a closed lower end; c. wherein the open upper end of the flanging cavity is substantially smaller than the bottom surface of the flanging cavity, thereby defining a flange area within the lower end; d. the body structure having at least one microfluidic port fabricated therein, the at least one microchannel being in fluid communication with the lower end of the flanging cavity; e. bonding a first polymer tubing into the flanging cavity wherein the lengths, cross-sections, or combinations thereof of the tubing are dimensioned to adaptedly fit through the open upper end of the flanging cavity and extend substantially the depth the flanging cavity, wherein the first polymer tubing substantially fills the cavity defined within the holding substrate upon acceptable pressure, chemical and/or thermal treatment to form a polymer tubing bushing upon treatment, wherein the bonded polymer tubing thereby creates a bushing with an interior channel adapted for containing a second polymer tubing; f. bonding a second polymer tubing into the bushing wherein the second polymer tubing has walls defining an inside microfluidic channel and has an outside diameter adapted to fit within the first polymer tubing bushing; and g. wherein the microfluidic channel is substantially in alignment and fluidly in connection with the microfluidic port of the body.

52. A method of forming a microfluidic interconnection system comprising: a. forming at least one flanging cavity in a body structure having at least first and second substrate layers, the first substrate layer disposed on top of the second substrate layer wherein the first and second substrate layers each comprise a top and a bottom planar surface; b. wherein the at least one flanging cavity is formed in the first substrate layer, wherein the cavity has an open upper end in communication with the top planar surface of the first substrate and an open lower end in communication with the bottom planar surface of the first substrate; wherein the open upper end of the flanging cavity is substantially smaller than the open lower end of the flanging cavity, thereby defining a flange area between the cavity walls of the first substrate layer and the top planar surface of the second substrate layer; the second substrate layer having at least one microfluidic port fabricated therein, the at least one microchannel being in fluid communication with the top planar surface end of the flanging cavity; c. bonding a first polymer tubing into the flanging cavity wherein the lengths, cross-sections, or combinations thereof of the tubing are dimensioned to adaptedly fit through the open upper end of the flanging cavity and extend substantially the depth the flanging cavity, wherein the first polymer tubing substantially fills the cavity defined within the holding substrate upon acceptable pressure, chemical and/or thermal treatment to form a polymer tubing bushing upon treatment, wherein the bonded polymer tubing thereby creates a bushing with an interior channel adapted for containing a second polymer tubing; d. bonding a second polymer tubing into the bushing wherein the second polymer tubing has walls defining an inside microfluidic channel and has an outside diameter adapted to fit within the first polymer tubing bushing; and e. wherein the microfluidic channel is substantially in alignment and fluidly in connection with the microfluidic port of the body.

53. The system according to claim 51 or 52, wherein the substrate layers independently comprise two or more layers.

54. The system according to claims 53 wherein the substrate further comprises at least one air vent in fluid communication with the flange area of the flanging cavity.

55. The system according to claims 53 wherein the polymer tubing comprises at least one thermoplastic resin.

56. The system according to claim 55, wherein the thermoplastic resin has a glass transition temperature ("Tg") less than about 550°C.

57. The system according to claim 55, wherein the thermoplastic resin has a glass transition temperature ("Tg") less than about 500°C.

58. The system according to claim 55, wherein the thermoplastic resin has a glass transition temperature ("Tg") less than about 400°C.

59. The system according to claim 55, wherein the thermoplastic resin has a glass transition temperature ("Tg") less than about 350°C.

60. The system according to claim 55, wherein the thermoplastic resin is selected from the group consisting of wherein the polymeric material is selected from the group consisting of polyamide, polyester, cellulose esters, polyethylene, polypropylene, poly(vinyl chloride), poly(vinylidene fluoride), polyphenylsulfones, polytetrafluoroethylene. Polymethylmethacrylate, polyetheretherketone, polyamide, polypropylene, polycarbonate, polydimethylsiloxane, polystyrene, polysulfone, and polyurethane.

61. The system according to claims 53 wherein the substrate comprises one or more materials selected from the group of glass, silicon, metal and polymeric substrates.

62. The system of claim 61, wherein the body structure comprising at least 50% polymeric materials.

63. The system of claim 62, wherein the polymeric material is selected from the group consisting of wherein the polymeric material is selected from the group consisting of polyamide, polyester, cellulose esters, polyethylene, polypropylene, poly(vinyI chloride), poly(vinylidene fluoride), polyphenylsulfones, polytetrafluoroethylene. Polymethylmethacrylate, polyetheretherketone, polyamide, polypropylene, polycarbonate, polydimethylsiloxane, polystyrene, polysulfone, and polyurethane.

64. The system of claim 63, wherein the body structure is formed by micromachining.

65. The system of claim 64, wherein the micromachining is by one or more methods selected from the group consisting of photolithography, etching, bonding, laser ablation, LIGA, injection molding and embossing.

66. A device according to claim 65, wherein the body structure is a microchip.

67. The system of claim 66, wherein the at least one microchannel has a dimension between about 0.1 microns and about 500 microns.

68. The system according to claims 53 wherein the body structure is fabricated by bonding a first polymeric substrate to a second polymeric substrate,

69. The system of claim 68, wherein the at least one flanging cavity is formed between the first and second substrate.

70. The system of claim 68, wherein the bonding is by fusion bonding.

71. The system of claim 69, wherein the bonding is by interfacial layer bonding.

72. The system of claim 70, wherein the bonding is at a temperature of from about 400 to about 680 ^°C.

73. The system of claim 70, wherein the bonding is at a temperature of from about 550 to about 650 ^°C.

74. The system of claim 70, wherein the bonding is at a temperature of from about 600 to about 640 ^°C.

75. The system of claim 71, wherein the bonding is at a temperature of from about 60 to about 400 ^°C.

76. The system of claim 71, wherein the bonding is at a temperature of from about 80 to about 360 ^°C.

77. The system of claim 71, wherein the bonding is at a temperature of from about 100 to about 340 ^°C.

78. The system of claim 60, wherein the flanging and/or bonding of the thermoplastic tubing flanging is at a temperature of from about 30 to about 500 ^°C.

79. The system of claim 60, wherein the flanging and/or bonding is at a temperature of from about 60 to about 400 ^°C.

80. The system of claim 60, wherein the flanging and/or bonding is at a temperature of from about 75 to about 350 C.

81. The system of claim 60, wherein the flanging and/or bonding of the thermoplastic tubing flanging is at a pressure of from about 0.1 MPa to about 5 Mpa.

82. The system of claim 60, wherein the flanging and/or bonding of the thermoplastic tubing flanging is at a pressure of from about 0.25 MPa to about 2.5 Mpa.

83. The system of claim 60, wherein the flanging and/or bonding of the thermoplastic tubing flanging is at a pressure of is from about 0.5 MPa to about 1.5 MPa.