NON-CONTACT THERMAL CONTROL OF SMALL VOLUME AND RELATED APPARATUS THEREOF

FIELD OF THE INVENTION The present invention relates to methods and apparatus for rapidly and accurately measuring and controlling the temperature of a small volume sample. More specifically, the present invention relates to methods and apparatus for measuring the temperature of the sample in performing non-contact thermocycling on small, micro to nanoliter, volume samples.

BACKGROUND OF THE INVENTION

Numerous analytical methods require that a sample be heated to a particular temperature and then cooled to a particular temperature. Often, sequential heating and cooling steps, known as thermocycling, are required. Various methods involve cycling through two or more stages all with different temperatures, and/or involve maintaining the sample at a particular temperature stage for a given period of time before moving to the next stage. Accordingly, thermocycling of samples can become a time consuming process, hi addition, these methods often require the precise control of temperature at each stage of the cycle; exceeding a desired temperature can lead to inaccurate results.

Two factors that are typically important, therefore, in the performance of effective thermocycling on a sample are the speed and homogeneity of the apparatus and the methods used. Cycle times are largely defined by how quickly the temperature of the sample can be changed, and relate to the heat source itself and the rate of heat transfer to the sample. Uniformity of sample temperature is important to ensure that

reproducible and reliable results are obtained. Typically, increasing cycle speeds makes it harder to maintain homogenous sample temperatures.

The concept of using elevated temperatures to effect chemical, biological and biochemical reactions is commonly known and expressed as the law of Arrhenius. Generally, an increase in temperature of a reaction translates into an increase in the rate of the reaction. Reaction parameters, such as the activation of the reaction, the increase in dissolution of the reaction components, the desolvation of the substrate and the specificity of the catalysis are temperature dependent. Exact or nearly exact maintenance of a reaction temperature is often critical in most biochemical/biological processes to guarantee their successful completion. Therefore, great efforts are made in the daily routine of a chemical/biochemical laboratory to control the temperature conditions during a reaction. It is expected that better temperature control increases the performance of most reactions, for example, increasing the specificity of proteolytic reactions. There is particular interest in rapid and homogenous thermocycling when performing DNA amplification via polymerase chain reaction (PCR). PCR is a process by which a single molecule of DNA (or RNA) from an organism can be amplified by a factor of 10 ⁶ to 10 ⁹ . This procedure requires the repetition of heating and cooling cycles in the presence of an original DNA target molecule, specific DNA primers, deoxynucleotide triphosphates, and DNA polymerase enzymes and cofactors. Heating accounts for a denaturing of the sample while cooling results in annealing of the sample. At a temperature typically between the denaturing and annealing temperatures, extension of the annealed primers using an enzyme occurs to replicate the DNA strand or portion of the strand. Extension of the primer can also occur at the same temperature as annealing, depending on the specifics of the reaction. Each

heating/cooling cycle produces a doubling of the target DNA sequence, leading to an exponential accumulation of the target sequence. PCR based technology has been applied to a variety of analyses, including environmental and industrial contaminant identification, medical and forensic diagnostics, and biological research. There are a number of biochemical reactions that require accurate and rapid thermocycling. Additionally, there are reactions whose specificity can be enhanced when conducted in a rapid and accurate thermocycling environment. The PCR reaction places very high demands on the accuracy of the thermocycling parameters and is, therefore, an ideal assay to test the accuracy of the thermocycling method and apparatus.

U.S. Pat. No. 4,683,202 generally describes the PCR concept, in which a stretch of DNA is copied using a polymerase. Generally, the procedure involves annealing a piece of primer DNA at a first temperature to any stretch of single- stranded DNA template with a complementary sequence. The DNA polymerase copies the primed piece of DNA at a second given temperature. At a third given temperature, the newly copied DNA and the primer dissociate from the template DNA, thereby regenerating single-stranded DNA. The temperature of the sample is returned to the first temperature to allow the primer to attach itself to any strand of single-stranded DNA with a complementary sequence, including the DNA strands that were synthesized in the immediately preceding cycle. In that manner, the template DNA is amplified or reproduced any number of times, depending on how many times the template DNA occurs in the sample, and the number of cycles completed. The procedure can also be performed using RNA.

Most existing methods and techniques of thermocycling in benchtop instrumentation are indirect with respect to the effect of the heating source on the

sample. Most thermocycling approaches heat and/or cool a circulating medium, such as water or air, that affects the container which holds the sample and, subsequently, subjects the sample itself to the desired thermocycling process. The rate of the cycling process depends on the effectiveness of the heat transfer between the circulating medium and the sample.

For example, U.S. Pat. No. 5,504,007 discloses a thermocycle apparatus having a body containing a thermally conductive liquid. The liquid is contained within the body of the apparatus, and the temperature of the liquid alternated between lower and higher temperatures in repeating cycles. A well or container for holding a sample of material is held in contact with the liquid and conducts the cyclic temperature changes of the liquid to the sample.

U.S. Pat. No. 5,576,218 discloses a method for the thermocycling of nucleic acid assays using a blended fluid stream produced from constant velocity, constant volume, and constant temperature fluid streams. Using these streams, a variable temperature, constant velocity, constant volume fluid stream is introduced into a sample chamber for heating and cooling the samples contained therein. The temperature of the blended fluid stream is varied by diverting and altering the ratio of the constant temperature fluid streams relative to one another.

U.S. Pat. No. 5,508,197 discloses a thermocycling system based on the circulation of temperature controlled water directly to the underside of a thin-walled polycarbonate microtiter plate. The water flow is selected from a manifold fed by pumps from heated reservoirs.

Other methods are reported for heating a sample through the use of heated air. U.S. Pat. No. 5,187,084 discloses an apparatus and method for performing thermocycling on a sample using an array of sample containing vessels supported in a

reaction chamber, through which air at controlled temperatures is forcibly circulated as a heat-transfer medium in heat exchange relationship with the vessels. The temperature of the air is controlled as a function of time to provide a preselectable sequence defining a temperature profile. The profile is a repetitive cycle that is reproduced to effect replication of and amplification of the desired sequence of the DNA.

U.S. Pat. No. 5,460,780 discloses a device for rapidly heating and cooling a reaction vessel through various temperatures in PCR amplification utilizing a device for heating at least one side wall of a reaction vessel, device for cooling the heating device at repeated intervals and device for moving the reaction vessel and/or heating and cooling relative to each other. In one embodiment, heated air is used to heat the reaction vessel.

Similarly, U.S. Pat. No. 5,455,175 demonstrates that rapid, non-contact PCR can be accomplished in glass capillaries using air heated by foam lining the chamber in which the capillaries are placed; the foam is heated first by a halogen lamp.

Another common approach for thermocycling is through intimate contact between a reaction vessel holding the reaction medium and a heating block that is rapidly heated and cooled (for example, by using a Peltier element that can both heat and cool). That is the basis of most commercially available PCR instrumentation. For example, U.S. Pat. No. 5,525,300 discloses an apparatus for generating a temperature gradient across a heat conducting block.

U.S. Pat. No. 5,498,392 discloses chip-like devices for amplifying a preselected polynucleotide in a sample by conducting a polynucleotide polymerization reaction. The devices comprise a substrate microfabricated to defme a sample inlet port and a mesoscale flow system, which extends from the inlet port. A

polynucleotide polymerization reaction chamber containing reagents for polymerization and amplification of a polynucleotide is in fluid communication with the inlet port. A heat source and, optionally, a cooling source are used to heat and/or cool the chip. Wilding and co-workers, Nucleic Acids Res., 24:380-385 (1996), demonstrated that PCR could be carried out in a microfabricated silicon glass chip- like chamber. By contacting enclosed 12 microliter reaction chambers microfabricated in glass to a block heater which cycled between two temperatures, they were able to obtain effective and reproducible PCR amplification, as confirmed by removing the PCR product and evaluating it using capillary electrophoresis. Similarly, Northrup and co-workers, Anal. Chem., 68:4081-4086 (1996), accomplished PCR amplification of DNA in a microfabricated silicon PCR device that could be directly interfaced with an electrophoretic chip for PCR product analysis. The device contained disposable polypropylene liners to retain the PCR mixture which could be cycled between two temperatures using polysilicon heaters in direct contact with the PCR chamber and cooled either passively or by air drawn along the heater surfaces of the reaction chamber. The device was interfaced with the electrophoretic chip by forcing it into the 1 mm drilled holes in the electrophoretic chip.

All of the above references, however, describe PCR amplification methods wherein the vessel containing the sample is contacted directly by a heater or another heat source, which transfers heat to the vessel in which the sample is contained. The vessel, in turn, heats the sample. Since these techniques rely on the intimate contact between the circulating medium and the reaction vessel, the surface-to-volume ratio of the reaction vessel is of utmost importance to the effectiveness of the heating step; the higher that ratio the better the PCR reaction.

PCT publication WO 96/41864 discloses a diode laser heated microreaction chamber with a sample detection device. A heat source, such as an IR or UV source, is used to heat the reagents to a thermally induced chemical reaction. Such heating device can be used, for example, in conjunction with the microfabricated reactor described in U.S. Pat. No. 5,639,423.

U.S. Patent Nos. 6,413,766 and 6,210,882, which are incorporated herein by reference, disclose thermocycling using both a non-contact heating source and a non- contact cooling source. The heating source is provided by optical energy from an IR source. The cooling source is provided by forcing air across the reaction vessel. The temperature sensor in the system, however, is a thermocouple that requires direct contact with the sample fluid.

None of the above references teach methods and apparatus for performing ultrafast and reliable thermocycling using all non-contact heating source, cooling source, and temperature sensor for providing sharp and rapid transitions from one temperature to another.

The possibilities of thermocycling on a device in which temperature control is achieved using a temperature sensor that is predetermined by the initial design of the chip are limited, as the location of the temperature sensor is typically part of the chip itself. Thus, those microdevices used in thermocycling are spatially constrained; and the devices are not flexible with respect to temperature sensing on different locations within or at the microdevice structure.

In addition, the design of single-use modules for various diagnostic and monitoring purposes with integrated temperature sensor is very complex, and becomes cumbersome and difficult to use especially when numerous samples are to

be tested. Therefore, the inexpensive production of such devices, normally a major advantage of microfabrication technology, is compromised.

One method for fast thermal cycling is non-contact heating via infrared (IR) excitation of the vibrational bands of water. IR-mediated PCR in microdevices, has been shown to be capable of fast thermocycling. The use of a broadband tungsten lamp/convective fan allows for remote heating and cooling, and facilitates building the temperature control hardware into the instrumentation and not into the device, allowing for simple and cost-effective microdevice fabrication.

For a non-contact (remote) method to be effective for selective heating of microdomains on a fluidic chip, it must be coupled with a comparably simple but sensitive method for temperature sensing. While we have frequently utilized manually-inserted thermocouples for direct sensing of solution temperature, these require surface passivation to avoid inhibition of PCR (as would a thermocouple fabricated into the chamber). A temporary but effective solution to this problem has been the creation of a thermocouple reference chamber adjacent to the PCR chamber. However, a more viable solution involves a method that, like the heating, senses temperature through a non-contact process, again minimizing the cost of microchip fabrication. Additionally, remote (non-contact) temperature sensing is most applicable on microfluidic chips where the channels or chambers are so small (less than about 100 μm) that a thermocouple will not fit therein.

There is a need, therefore, for improved methods and apparatus for remote temperature sensing of analytical samples in a fast and reliable manner. There is a further need for such methods and apparatus for use with miniaturized thermocylcing, such as that for the polymerase chain reaction (PCR) amplification. Remote temperature sensing is used herein to describe temperature measuring without directly

S

contacting the solution of interest. Remote temperature sensing offers two very important benefits for small volume microdevice thermocycling applications. First, there is no added thermal mass from the detection system. That is important in the rapid heating and cooling needed for decreasing the reaction times. Second, no additional components need be fabricated into the microchip itself in order to accurately measure the temperature of a solution within a microfabricated chamber. That decreases the cost, and may allow a single system to monitor temperatures in a number of positions on a single device with no additional connections.

SUMMARY OF THE INVENTION

Remote sensing of the temperature of a solution within a small volume chamber can be accomplished by using a pyrometer, which is normally used for remotely measuring the temperature of a surface by measuring the radiation from that surface. However, because the small volume of the present invention is typically a closed reservoir or container, the present inventor has developed methods and apparatuses to use the pyrometer to interrogate the temperature inside the closed reservoir or vessel, rather than the temperature of the surface of the material enclosing the reservoir or container. The "closed" vessel, as used herein, does not necessarily imply that the vessel is closed to the outside, but to convey that the fluid in the vessel is not directly exposed to interrogation by the remote temperature sensor. Generally, the vessel is contained inside a microfluidic device. As such, a layer of material, that forms the microfluidic device, e.g. glass, encloses the vessel. However, the channel is open to allowing fluids to flow to and from the vessel via microchannels of the microfluidic device.

A pyrometer generally contains an optical system and a detector. The optical system focuses the energy emitted by the surface of an object onto the detector, which is sensitive to the radiation. The output of the detector is proportional to the amount of energy radiated by the target object, and the response of the detector to the specific radiation wavelengths. This output can be used to infer the objects temperature by the Stefan Boltzmann equation:

J^εσT ⁴ (eq. 1) where J is the energy detected by the detector, ε is the emissivity of the object, and σ is the Stefan-Boltzmann constant. Pyro ^meter are ^we U known in the art and have been used in various capacities to measure surface temperatures.

Therefore, it is an object of the present invention to provide a method and apparatus for temperature sensing of a small volume of fluid inside a closed reservoir with a pyrometer.

It is further object of the present invention to provide a method and apparatus for temperature control, particularly for thermal cycling, on a microchip having all remote heating, cooling, and remote temperature sensing.

π

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 shows a preferred thermocycling system of the present invention; FIG. 2 shows the instrumental layout of pyrometer-controlled PCR of Example 1. The pyrometer was placed off axis from the tungsten lamp to ensure the pyrometer was not heated during experiments. During and after calibration of the pyrometer, the microdevice was secured to the setup to ensure the same location above the PCR chamber was measured for accurate temperature sensing. The digital control lines from the data acquisition (DAQ) card to the relays are shown as dotted lines. FIG, 3 shows a calibration curve of the pyrometer to a thermocouple. A. Five mock PCR cycles of a three temperature thermal cycling program controlled by a thermocouple (black line, left Y-axis) were used to calibrate the pyrometer output voltage (red line, right Y-axis). Data values from each hold step were averaged and used in attaining the calibration constants. It is apparent that the time lag between the two temperatures was larger in the initial cycles compared to the later cycles and this effect was likely due to the entire device heating as the cycles progressed (explained in more detail at the end of the text). B. The lag time between the maximum heating rates measured by the thermocouple and pyrometer are plotted for the various transition periods of the cycling shown in Figure 2A. Heating to denature entailed the temperature change from 72 95 ⁰ C, heating to extension was 60 - 72 ⁰ C, and cooling to anneal was 95 - 60 ⁰ C. While it appears that the lag time was largest during the heating to extension, the large error, due to the changing overall device temperature, was large enough that the results were not statistically different.

FIG. 4 shows non-contact amplification of λ-phage DNA. A. Thirty cycles of a two-temperature heating protocol using the pyrometer calibrated with the trace

shown in Figure 2 A. B. After thermal cycling was complete, analysis of the product within the PCR chamber by capillary gel electrophoresis demonstrated that an amplified product was present which migrated at the expected size (shown by *). The other peaks present at 3.5 and 4.0 minutes correspond to primer and dimer peaks, respectively. Amplification with no template DNA was subsequently performed and is shown as the bottom electropherogram.

FIG. 5 shows calibration of the pyrometer using boiling points. Water (black line) or an azeotrope (grey line) was heated until the solution boiled (n = 3). As more easily observed with water, the surface temperature of the glass still increased during the phase change, but the change in the heating rate could be quantitatively determined by taking the derivative of these traces with respect to time. The derivative of the azeotrope heating trace is shown in the inset with the unbroken arrow indicating the time point that boiling occurred. This time point was then used to determine at which pyrometer output voltage (broken line) corresponded to the boiling temperature.

FIG. 6. shows control of PCR using boiling point calibration. The pyrometer was calibrated using the data in FIG. 3 and used to control amplification of a gene fragment from B. anthracis in a 250 nL PCR chamber. The product peak (shown by *) migrated at the expected size relative to a DNA ladder with size of fragments given above ladder peaks in bp. The inset details the sizing information for the 211 bp product (open square). A control amplification showed no detectable amplified product at the expected size (bottom electropherogram).

FIG. 7 shows increased heating rates using a Au-coated mirror. A. A gold- coated mirror was placed 5 cm above the PCR chamber and used to focus stray IR radiation back onto the device. With this mirror, the cycling time, relative to Figure

3A and 5, was reduced 44% to 18.8 minutes. B. Amplification of a 211 bp gene fragment from B. anthracis was performed using this faster cycling method and a 214 + 2 bp peak was observed (peak shown by *, sizing data as the open square in the inset) when compared to the same DNA sizing ladder as shown in FIG. 6. Amplification with no template DNA produced no detectable product (bottom electropherogram).

FIG. 8 shows increased overall temperature of the microfluidic device. The average normalized outputs recorded by both the thermocouple and pyrometer at each denature, anneal, and extend steps were plotted for each cycle in a 30 cycle, thermocouple-controlled mock PCR. As can be seen, the thermocouple traces are relatively constant since the temperature is being controlled with this sensor; however, the pyrometer signal increases over the first 15 cycles indicating the temperature of the microdevice surface is increasing during this time. Plots are named by the method used to sense and the hold step, for example, "TC Denature" are values recorded from the thermocouple during the denature holds and "Pyro Anneal" are values recorded from the pyrometer during the anneal holds. One standard deviation is shown for each data point for clarity.

FIG. 9 shows thermal cycling in glass devices of 2.2-mm thickness was found to be much slower (~0.5 h) than the polymer counterparts (<10 min). FIG. 10 shows (a) a typical microchamber; (b) a microchamber with contact heating, an indirect heating mechanism; and (c) non-contact heating, a direct heating mechanism requiring less overall energy.

FIG. 11 shows that the PID feedback control algorithm, optimized for the 72° C hold, was effective for thermal cycling.

FIG. 12 shows that the use of pulse-width modulation (PWM) allowed precise and accurate non-contact temperature control on microchips, even with cycling rates exceeding 20 ⁰ C s ^'1 .

FIG. 13 shows a standard method for thermocouple calibration developed using a conventional PCR instrument, (a) An example trace and (b) calibration curve are shown (76.6 mV ⁰ C ¹ ).

FIG. 14 shows water boiling with the thermocouple position (a) at the center of the chamber; and (b) at the edge of the chamber. The inaccurate boiling point seen in (a) was indicative of the thermocouple absorbing IR radiation.

DETAILED DESCRIPTION OF THE INVENTION

The present invention is generally directed to an apparatus and method for performing remote, rapid, accurate temperature measurement on small volume samples. Remote temperature measurement, in the context of this application, is used to describe temperature measuring without directly contacting the solution of interest. The term "small volume" as used herein refers to volumes in the picoliters (pL) to microliters (μL) range, preferably about 100 pL to about 100 μL, most preferably about 1 nL to about 10 μL.

The present invention uses a pyrometer, preferably an infrared (IR) pyrometer, to remotely measure temperature of a small volume sample. A pyrometer generally contains an optical system and a detector. The optical system focuses the energy emitted by the surface of an object onto the detector, which is sensitive to the radiation. The output of the detector is proportional to the amount of energy radiated by the target object, and the response of the detector to the specific radiation wavelengths. This output can be used to infer the objects temperature by the Stefan- Boltzmann equation:

J=SaT ⁴ (eq. 1) where J is the energy detected by the detector, ε is the emissivity of the object, and σ is the Stefan-Boltzmann constant. Pyrometer are well known in the art and have been used in various capacities to measure surface temperatures. The present invention, however, uses the pyrometer to measure the temperature of the small volume enclosed inside a microfluidic chip, rather than the temperature of the surface of the material enclosing the small volume.

In an embodiment of the present invention, several temperature sensors can be used to simultaneously detect temperatures at a plurality of locations on a microfluidic chip.

The remote temperature sensor of the present invention is most appropriate for use with miniaturized analytical processes, particularly those on chips or microflluidic devices. Preferably, those PCR processes, along with the use of the present remote temperature sensor, also use remote heating and cooling. In doing so, the chip is manufactured without having to fabricate integral heating, cooling and temperature measuring elements for each individual chip. That results in significant simplification of the chip and a reduced cost of manufacturing. Further, the remote heating, cooling, and temperature sensing apparatus can be repeatedly reused with other chips and does not have to be discarded with the chip after use. Because those elements are remote, physical connection between the chip and the heating means, cooling means, and temperature sensing means do not have to be engineered into the chip itself. Applications of the thermocycling method of the present invention are numerous and generally encompass any analytical system in which the temperature of a sample is regulated and/or changed. The present invention is particularly applicable to analytical systems wherein fast or ultrafast transition from one temperature to the next is needed, and in which it is important that exact or nearly exact temperatures be achieved.

For example, the present apparatus and methods are suitable for testing and incubation and treatment of biological samples typically analyzed in a molecular biology laboratory or a clinical diagnostic setting. The accuracy of the thermocycling method of the present invention makes it particularly suitable for use in nucleic acid replication by the polymerase chain reaction (PCR). Any reaction that benefits from

precise temperature control, rapid heating and cooling, continuous thermal ramping or other temperature parameters or variations can be accomplished using this method discussed herein. Other applications include, but are not limited to, the activation and acceleration of enzymatic reactions, the deactivation of enzymes, the treatment/incubation of protein-protein complexes, DNA-protein complexes, DNA- DNA complexes and complexes of any of these biomolecules with drugs and/or other organic or inorganic compounds to induce folding/unfolding and the association/dissociation of such complexes. The following applications illustrate the usefulness of the present thermocycling apparatus and methods, representing only some of the possible applications.

A common procedure in the protocols of molecular biology is the deactivation of proteins through heat. One of the most basic procedures in molecular biology is the cleavage of proteins and peptides into discrete fragments by proteases/digestion enzymes, such as trypsin. A thermocycling procedure is typically used to activate the enzyme at an elevated temperature followed by: the incubation of the enzyme during the reaction to sustain the enzymatic catalysis; the heat inactivation of the enzyme; and the final treatment/analysis at ambient temperature. Typically, the reaction components are incubated at 4O ⁰ C for 60 minutes until the reaction is completed, after which the enzyme activity has to be stopped to avoid unspecific cleavage under uncontrolled conditions. Many enzymes, such as trypsin, can be irreversibly inactivated by incubation for 10 minutes at higher temperature, such as 95 ⁰ C. The sample is then cooled back to ambient temperature and ready for downstream analysis. Such deactivation of enzymes is taught, for example, in Sequencing of proteins and peptides: Laboratory Techniques in Biochemistry and Molecular Biology, ed G. Allen, pages 73-105.

The same principle of heat inactivation can be used to inactivate restriction endonucleases that recognize short DNA sequences and cleave double stranded DNA at specific sites within or adjacent to the recognition sequence. Using the appropriate assay conditions (for example, 4O ⁰ C for 60 min), the digestion reaction can be completed in the recommended time. The reaction is stopped by incubation of the sample at 65 ⁰ C for 10 minutes. Some enzymes may be partially or completely resistant to heat inactivation at 65 ⁰ C, but they may be inactivated by incubation for 15 minutes at 75 ⁰ C. Such methods are taught, for example, by Ausubel et al. Short Protocols in Molecular Biology, 3rd Ed., John Wiley & Sons, Inc. (1995) and Molecular Cloning: A Laboratory Manual, J. Sambrook, Eds. E. F. Fritsch, T. Maniatis, 2nd Ed.

Similar to the heat inactivation of proteins for the control of enzymatic activity, the sample processing of proteins for electrophoretic analysis often requires the denaturation of the protein/peptide analyte before the separation by electrophoretic means, such as gel electrophoresis and capillary electrophoresis, takes place. For example, a 5 minute heat denaturation (which provides for the destruction of the tertiary and secondary structure of the protein/peptide) at 95 ⁰ C. in an aqueous buffer in the presence or absence of denaturing reagents, such as SDS detergent, allows the size dependent separation of proteins and peptides by electrophoretic means. That is taught, for example, in Gel Electrophoresis of Proteins: A Practical Approach, Eds. B, D. Hames and D. Rickwood, page 47, Oxford University Press (1990).

Thermocycling of samples is also used in a number of nonenzymatic processes, such as protein/peptide sequencing by hydrolysis in the presence of acids or bases (for example, 6M HCl at 11O ⁰ C. for 24 hours) into amino acids. Studies involving the investigation of the interaction of biomolecules with drugs and/or drug

candidates are frequently conducted under conditions requiring precise temperature control to obtain binding characteristics, such as kinetic association/dissociation constants.

Those applications for the thermocycling taught by the present invention will find use, for example, as a diagnostic tool in hospitals and laboratories such as for identifying specific genetic characteristics in a sample from a patient, in biotechnology research such as for the development of new drugs, identification of desirable genetic characteristics, etc., in biotechnology industry-wide applications, and in scientific research and development efforts. Thus, the samples subjected to the thermocycling methods of the present invention will vary depending on the particular application for which the methods are being used. Samples will typically be biological samples, although accurate heating and cooling of non-biological samples is equally within the scope of this invention. A suitable vessel or reservoir according to the methods of the present invention is one in which extremely low volumes of sample can be effectively tested, including sample volumes in the nanoliter range. The sample vessel must be made of a material that allows the penetration of IR light wavelengths, such as quartz glass, glass, silicon, transparent plastics, and the like. Preferably, the reaction vessel or container will have a high surface-to-volume ratio. A high surface-to-volume ratio leads to a decrease in the thermal time constant, which can lead to an increase in the efficiency of the thermocycling. A high surface-to-volume ratio, while not as important for the heating step, is related to the effectiveness of the cooling step. Various examples of suitable reaction vessels can be given, including but not limited to, microchambers, capillary tubes, microchips and microtiter plates. A preferred example of a suitable reaction vessel or reservoir is a

microchamber made from thin-walled glass. Another preferred embodiment is a glass capillary tube. Such capillaries are typically used in capillary electrophoresis ("CE"). Suitable inner diameters of the capillaries having an outer diameter of about 370 μm typically vary between about 15 μm and 150 μm. Thermal gradients that lead to convection are substantially reduced in capillary tubes which are available commercially. Alternatively, the vessel vessel or reservoir may be any component of a microfluidic system where it is desired to measure the temperature of the fluid contained therein.

Another preferred example of a suitable reaction vessel is the channel structure incorporated into a microfabricated device, such as the micro fabricated substrate described by Wilding et al. in Nucleic Acids Res., 24:380-385 (1996), and U.S. Patent Nos. 5,726,026 and 6,184,029, which are encorporated herein by reference. Other reaction vessels with characteristics suitable for rapid thermocycling are shown in U.S. Patent Nos. 6,413,766 and 6,210,882, which are incorporated herein by reference.

Any other reaction vessel, such as a microtiter plate (96, 384 or 1636 wells), can be used according to the methods of the present invention, provided that the vessel is made of a material which allows FR radiation to directly heat the sample and has a surface-to-volume ratio sufficient to allow for cooling within the time parameters discussed below. A method for preparing a suitable microfabricated device is discussed in the example section. Further guidance in preparing such microfabricated device is provided, for example, in U.S. Pat. Nos. 5,250,263; 5,296,114; Harrison et al., Science 261 :895-897 (1993); and McCormick et al., Anal. Chem., 69:2626-2630 (1997), which are incorporated herein by reference.

A preferred system for thermocycling of the present invention is depicted in FIG. 1, which shows the system of the present invention in which microchip 600 contains at least one closed reservoir 602. The temperature of the small volume of fluid inside the reservoir can be measured using a remote temperature sensor (pyrometer) 606. The microchip 600 is preferably placed on a movable stage 608, which may be generally ring-like to leave the underside of the microchip 600 exposed, and which can be motorized or moved manually. A cooling source 614 is directed underneath the reservoir 602. Although only one cooling source 614 is shown, multiple cooling jets may be used to direct air above and/or below the reservoir 602. A heating source 610 may be a lamp having its emitted light whose frequency can be controlled, for example, by a filter. The light to which the reservoir 602 is exposed may be further limited by an aperture and/or a light restricting device to control the amount of heating to the reservoir 602. The microchip 600 may sit on stage that is a frame that supports the microchip 600 on its periphery so that the microchip 600 is exposed to the light from the lamp 610. The stage 608 may be moveable by a motor or manually so that different reservoirs 602 on the same chip may be interrogated by moving the reservoir 602 in proximity of the remote temperature sensor 606. Alternatively, multiple remote temperature sensors 604 can be used to monitor multiple reservoirs 602 on the microchip 600. It is preferable that the remote temperature sensors 606, the cooling source

614 and the heating source 610 as shown in FIG. 1 are operatively associated with a microprocessor 622. The remote temperature sensor 606 may electrically connected to the microprocessor 622 via a data acquisition card. The cooling source 614 may be a cooling fan or associated with a compressed gas source for cooling the small volume reservoir 602 in the microchip 600. The heating source 610 and the cooling source

614 may be electronically connected directly to and controlled by the microprocessor 622. The microprocessor 622 preferably contains systems for controlling and receiving data from each of the heating source 610, the cooling source 614, and the remote temperature sensor 606. Although FIG. 1 shows only one remote temperature sensor 604 associated with one reservoir 602, a plurality of temperature sensors/reservoirs are appropriate for the present invention. For example, if a microchip contains multiple reservoirs, the number of remote temperature sensors used can equal the number of reservoirs being used on the microchip. As can be seen from FIG. 1 , according to the present invention, a remote heat source, a remote cooling source, and remote temperature sensors are used. That allows for the repeated introduction of any number of reaction vessels in and out of the apparatus. Thus, the present invention provides an economic advantage over other thermocycling apparatus, in that it is only a relatively inexpensive microchip, capillary tube, or other reaction vessel that must be changed for every sample. Some methods provided in the art require the physical attachment of the heating, cooling, and/or temperature sensing systems to the reaction vessel itself. Therefore, unless the reaction vessel could be completely cleaned to ensure that contamination from one sample to another did not occur, a new chip attached to a new heating, cooling, and/or temperature sensing device would have to be provided for every sample. While for ease of reference, only two sample-containing vessel were shown and/or described in FIG. 1, it is equally within the scope of the present invention to thermocycle more than two samples at the same time. In addition, because the heating and cooling sources are relatively stationary in the apparatus of the present invention, the reaction vessel can be moved in any direction relative to the heating and/or cooling sources.

Heating of the sample is preferrably accomplished through the use of optical energy from a remote heat source. Preferably, this optical energy is derived from an IR light source which emits light in the wavelengths known to heat water, which is typically in the wavelength range from about 0.775 μm to 7000 μm. For example, the infrared activity absorption bands of sea water are 1.6, 2.1 , 3.0, 4.7 and 6.9 μm with an absolute maximum for the absorption coefficient for water at around 3 μm. The IR wavelengths are directed to the vessel containing the sample, and because the vessel is made of a clear or translucent material, the IR waves act directly upon the sample to cause heating of the sample. Although some heating of the sample might be the result of the reaction vessel itself absorbing the irradiation of the IR light, heating of the sample is primarily caused by the direct action of the IR wavelengths on the sample itself.

Typically, the heating source will be an IR source, such as an IR lamp, an IR diode laser or an IR laser. An IR lamp is preferred, as it is inexpensive and easy to use. Preferred IR lamps are halogen lamps and tungsten filament lamps. Halogen and tungsten filament lamps are powerful, and can feed several reactions running in parallel. A tungsten lamp has the advantages of being simple to use and inexpensive, and can almost instantaneously (90% lumen efficiency in 100 msec) reach very high temperatures. A particularly preferred lamp is the CXR, 8 V, 50 W tungsten lamp available from General Electric, That lamp is inexpensive and convenient to use, because it typically has all the optics necessary to focus the IR radiation onto the sample; no expensive lens system/optics will typically be required.

In a preferred embodiment, the optical energy is focused on the sample by means of IR transmissible lenses so that the sample is homogeneously irradiated. That technique avoids "hotspots" that could otherwise result in the creation of undesirable

temperature differences and/or gradients, or the partial boiling of the sample. The homogeneous treatment of the sample vessel with optical energy therefore contributes to a sharper temperature profile. The homogenous sample irradiation can further be enhanced through the use of a mirror placed on the opposite site of the IR source, such that the reaction vessel is placed between the IR source and the mirror. That arrangement reflects the radiation back onto the sample and substantially reduces thermal gradients in the sample. Alternatively, the radiation can be delivered by optical IR-transparent fiberglass, for example, optical fiberglass made from waterfree quartz glass that is positioned around the reaction vessel and that provides optimal irradiation of the sample.

With optical energy heating, various optical instruments, such as lenses, mirrors, filters, apertures, etc., may be used to efficiently deliver the energy to the reservoir. For example, mirrors and lens may be used to focus the radiation on the reservoir to improve heating efficiency. Heating can be effected in either one step, or numerous steps, depending on the desired application. For example, a particular methodology might require that the sample be heated to a first temperature, maintained at that temperature for a given dwell time, then heated to a higher temperature, and so on. As many heating steps as necessary can be included. Similarly, cooling to a desired temperature can be effected in one step, or in stepwise reductions with a suitable dwell time at each temperature step. Positive cooling is preferably effected by use of a non-contact air source that forces air at or across the vessel. Preferably, that air source is a compressed air source, although other sources could also be used. It will be understood by those skilled in the art that positive cooling results in a more rapid cooling than simply allowing the vessel to

cool to the desired temperature by heat dissipation. Cooling can be accelerated by contacting the reaction vessel with a heat sink comprising a larger surface than the reaction vessel itself; the heat sink is cooled through the non-contact cooling source. The cooling effect can also be more rapid if the air from the non-contact cooling source is at a lower temperature than ambient temperature.

Accordingly, the non-contact cooling source should also be positioned remotely to the sample or reaction vessel, while being close enough to effect the desired level of heat dissipation. Both the heating and cooling sources should be positioned so as to cover the largest possible surface area on the sample vessel. The heating and cooling sources can be alternatively activated to control the temperature of the sample. It will be understood that more than one cooling source can be used. Positive cooling of the reaction vessel dissipates heat more rapidly than the use of ambient air. The cooling means can be used alone or in conjunction with a heat sink. A particularly preferred cooling source is a compressed air source. Compressed air is directed at the reaction vessel when cooling of the sample is desired through use, for example, of a solenoid valve which regulates the flow of compressed air at or across the sample. The pressure of the air leaving the compressed air source can have a pressure of anywhere between 10 and 60 psi, for example. Higher or lower pressures could also be used. The temperature of the air can be adjusted to achieve the optimum performance in the thermocycling process. Although in most cases compressed air at ambient temperature can create enough of a cooling effect, the use of cooled, compressed air to more quickly cool the sample, or to cool the sample below ambient temperature might be desired in some applications.

Monitoring and controlling is accomplished by use of a microprocessor or computer programmed to monitor temperature and regulate or change temperature.

An example of such a program is the Lab VIEW program, available from National Instruments, Austin, TX. Feedback from the temperature sensor is sent to the computer. In one embodiment, the temperature sensor provides an electrical input signal to the computer and/or other controller, which signal corresponds to the temperature of the sample.

Signals from the computer, in turn, control and regulate the heating and cooling means, such as through one or more switches and/or valves. The desired temperature profile, including dwell times, is programmed into the computer, which is operatively associated with heating and cooling means so as to control heating and cooling of the sample based upon feedback from the temperature sensor and the predetermined temperature profile.

Accordingly, the methods of the present invention provide for the use of virtually any temperature profile/dwell time necessary. For example, cleavage of proteins through use of proteases or digestion enzymes might require use of different temperatures, each of which must be precisely maintained for various amounts of time. Activation of restriction endonucleases might similarly require achieving and maintaining two or three different temperatures. Protein or peptide sequencing can require the steady maintenance of a high temperature for an extended period of time. The above apparatus provide for rapid heating and cooling of a sample in a precise and easy to replicate manner. Heating can be effected for example as quickly as 1O ⁰ C per second when using approximately 15 to 50 μL volumes of sample in a microchamber and as rapidly as 100°C. per second when using nL volume samples in a capillary. Cooling can be effected quickly, typically in the range of between about 5 and 5O ⁰ C per second. The increased effectiveness of heating and cooling improves the cycling process and sharpens the temperature profile. This means that the desired

reaction can be conducted under more optimal thermal conditions than in conventional instruments. Thermal gradients in the reaction medium frequently observed in instrumentation using a contact heat source are detrimental to the specificity of the reaction. Those thermal gradients are substantially reduced in the IR mediated heating, particularly when the heat source is strong enough to penetrate the aqueous mixture and provide sufficient irradiation to the opposite side of the reaction vessel. Non-contact, remote rapid cooling, heating, and temperature sensing, such as that provided in the present invention, also contributes to the ability to obtain sharp transition temperatures in minimum time and to achieve fast and accurate temperature profiles.

In operation, the pyrometer must be properly calibrated to measure the temperature of the small volume inside the closed reservoir of the microfluidic device rather than just the external wall temperature. The calibration account for the lag time associated with temperature equilibrium between the surface of the microfluidic device and the small volume inside the closed reservoir of the micofluidic device. At the very least, the method involves the use of the transition temperatures of two different calibration substances. In certain embodiments of the present invention, the calibration substances are reference solutions. In these embodiments, the calibration method involves the use of the boiling points of at least two different reference solutions. The criteria for choosing the reference solutions included 1) boiling points that were within the range of temperatures normally used for the small volume (in the case of PCR, approximately 60-95 ⁰ C); and 2) the solutions needed to approximate IR absorption in a manner similar to absorption by the solution used inside the small volume (in the case of PCR, at least 10% water). Preferably, the reference solutions include either water and an azeotrope or two different azeotropes. The azeotrope or

azeotropes may be selected from any suitable azetrope known in the art, for example, from those described in the CRC Handbook of Chemistry and Physics. During calibration, the reference solutions, one at a time, are flowed into the reservoir and heated. The rising temperature of each solution is measured through surface sensing (above the reservoir) with the pyrometer which generated the pyrometer a voltage versus time plot. When the solution reaches its boiling point, the temperature remained constant during the phase change, which can be detected with the pyrometer as a change in slope. Using the derivative of these plots, the pyrometer voltage minima can be identified and used to calibrate the pyrometer against the known boiling point for the reference solution. This method yields at least two reference points with which to calibrate the pyrometer. More reference points may be obtained by using more than two reference solutions.

Alternatively, the microfluidic device can be prefabricated with two calibration chambers (reservoirs), each chamber being filled with a different reference solution. The calibration process would be the same where each of the chamber is heated and the boiling point is determined. This prefabricated device saves the user of having to prepare and flow the reference fluids into the chamber for calibration.

In other embodiments of the present invention, the calibration substance is a solid. For example a solid with a flash point at a certain temperature may be used as one or more of the calibration substances, with the flash point acting as a reference temperature. In certain embodiments of the invention, the solids are selected from the alkali metals and alkaline earth metals.

In still other embodiments of the present invention, the calibration substance is a gas that undergoes a pseudostate change at a specific temperature, with the pseudostate change acting as a reference temperature. Any gas with a detectable

pseudostate transition at a suitable temperature may be used as a calibration substance of the present invention. In certain embodiments, the gas may undergoes a pseudostate transition from clear to cloudy or from cloudy to clear at the reference temperature. Without further description, it is believed that one of ordinary skill in the art can, using the preceding description and the following illustrative examples, make and utilize the compounds of the present invention and practice the claimed methods. The following example is given to illustrate the present invention. It should be understood that the invention is not to be limited to the specific conditions or details described in the examples.

Example 1 - Non-cootact temperature control of PCR on microfluidic chip

In this example, a simple, effective, and robust temperature sensing method is described which, together with IR-mediated heating, enables completely non-contact temperature control for performing PCR in microfluidic devices. This involves an IR pyrometer sensing the surface temperature above a PCR chamber with a design that assures the rapid equilibration between the PCR solution and the chamber surface. Thermal modeling is used to define physical properties of the device and environment needed to achieve this rapid equilibration, and the model verifies an experimentally- observed thermal equilibration of the device during the initial PCR thermal cycles. In line with a completely non-contact method of temperature control, calibrating the surface temperature relative to the PCR solution temperature is accomplished using the boiling point of water and an azeotrope within the chip. The effectiveness of the method is illustrated with the successful PCR of a fragment of B. anthracis gene in a fluidic channel too small for conventional thermocouple-based sensing. Using a gold-

coated parabolic mirror to focus stray IR radiation back onto the device allows for 30 cycles of PCR amplification to be complete in 19 minutes.

Experimental Method

Microfluidic devices were made as previously described (Roper et al., Anal.

Chim. Acta 2006, 569, 195-202, which is incorporated herein by reference). Glass wafers (Telic Company, Valencia, CA) pre-coated with a layer of chrome and positive photoresist were exposed to UV radiation through a film mask. Irradiated photoresist and the underlying chrome were then removed and the exposed glass was etched using hydrofluoric acid. Diamond-tipped drill bits (Tripple Ripple, 1.1 mm diameter, Crystalite Corp., Lewis Center, OH, USA) were used to drill fluidic access holes and after drilling, microdevices were cleaned with a 3:1 (v:v) solution Of H ₂ SO ₄ )H ₂ O ₂ followed by a 5: 1 : 1 (v:v:v) solution of H ₂ OiNH ₄ OH^ ₂ O ₂ . Clean bottom plates, 200 μm thick, were then placed into contact with the etched devices and bonded together at 640 ⁰ C for 8 hours.

The photomask design consisted of two ellipses, each 0.75 mm wide and 3 mm long. The two ellipses were separated by 1.5 mm (from one center of the ellipse to the other center). Two channels, each impinging on the end of each ellipse, were used for filling the ellipses with PCR solution and insertion of the thermocouple. Etch depths were either 200 μm deep or 100 μm deep producing PCR reaction volumes of 550 nL and 230 nL, respectively. Thermocouple and Pyrometer PCR Setups

The PCR setup was similar to systems described by Easley et al. {Lab Chip 2006, 6, 601-610, which is incorporated herein by reference) and Legendre et al. {Anal Chem. 2006, 78, 1444-1451 , which is incorporated herein by reference).

Briefly, the microdevice was mounted on a Plexiglas stage that allowed access to the infrared radiation and convective cooling from the bottom of the stage. Two solid state relays (CMX60D10, Crydom Corp., San Diego, CA, USA) were placed in series with a 12 V power supply to power both the fan and the tungsten lamp (CXR, 8 V, 50 W, General Electric, Cleveland, OH, USA). The pyrometer (MI-N5, Mikron Infrared, Inc., Oakland, NJ, USA) was oriented 45° from vertical above the microdevice and a red diode laser within the pyrometer was used to align the sensing area of the pyrometer onto the surface of the device above the PCR chamber (Figure 1). When the microdevice was placed on the stage and aligned with both the pyrometer sensing area and the focal spot of the tungsten lamp, the physical location of the microfluidic chip was recorded using double-sided tape. In some experiments stated in the text, a gold-coated parabolic mirror (Edmund Optics, Barrington, NJ, USA) was placed 50 mm above the microdevice to focus the stray radiation from the lamp onto the top of the microfluidic chip. A thermocouple (T-240C, Physitemp Instruments, Inc., Clifton, NJ, USA) was used to initially calibrate the pyrometer. The thermocouple voltage was digitized and amplified using a commercially-available circuit (TAC-386-T, Omega Engineering, Stamford, CT, USA) and a homemade circuit for a total amplification factor of 1875 (75 mV/°C). This voltage was recorded using a data acquisition (DAQ) card (PCI- 6014, National Instruments, Austin, TX, USA). The pyrometer voltage output was fed into the same DAQ card using a 1 kω resistor across the current output of the pyrometer.

To control the output of the IR lamp, a proportional integral derivative feedback algorithm (written in LabView, National Instruments) using the thermocouple or pyrometer input was used to control the duty cycle of a 1000 Hz

TTL output to the solid state relay in series with the lamp. The control to the relay in series with the fan was either on (TTL high) or off (TTL low). The thermocouple was calibrated by filling the 550 nL microfluidic device with PCR buffer, inserting the thermocouple, and placing mineral oil above the fluidic access holes. The microdevice and thermocouple were then placed on a conventional PCR thermal cycler (GeneAmp 2400 PCR System, Perkin-Elmer, Wellesley, MA, USA) and the temperature of the conventional cycler was increased in increments of 5 ⁰ C and held for 30 s at each increment. The average thermocouple voltage versus temperature of the conventional cycler was used to calibrate the thermocouple. The pyrometer was calibrated by placing the thermocouple within the microdevice and the IR-lamp used to perform 5 mock PCR cycles with 15 s hold times at 95, 60, and 72 ⁰ C. The average pyrometer output voltage at the average temperature recorded by the thermocouple was used for calibration. When reporting the data for the pyrometer or thermocouple, all values are average + one standard deviation.

Calibration of the pyrometer by measurement of boiling points

A 230 nL microfluidic device was filled with water or a 13.1:38.2:48.7 (v:v:v) mixture of water:2-propanol:toluene and positioned in the focal spot of the tungsten lamp. The lamp was subsequently turned on at full power and the pyrometer output voltage recorded as a function of time. This procedure was repeated three times with each solution and the derivative of these traces were used to determine the pyrometer output voltage that correlated with the known boiling points of the solutions (see Results and Discussion for more details).

PCR Procedure

Microdevices were rinsed with acetone, dried, and passivated with 10 μL SigmaCote (Sigma-Aldrich, St. Louis, MO) prior to making PCR mastermix solutions. All PCR reagents were from Sigma-Aldrich and primers were from MWG Biotech, Inc. (High Point, NC, USA). For PCR amplification, a 10 μL PCR solution was made which was composed of 10 mM Tris-HCl, 50 mM KCl, pH 8.3, 3 mM MgCl ₂ , 0.2 mM each dNTP, 0.4 μM forward and reverse primer, 0.1 units/μL Taq polymerase, and either 7 ng λ-phage DNA or purified DNA from 3000 colony forming units of anthrax spores (spores and anthrax primers kindly donated by Dr. Sanford H. Feldman, Department of Comparative Medicine, University of Virginia Health System). PCR protocol for a 500 bp fragment of λ-phage DNA consisted of an initial denaturation at 95 °C for 60 s, followed by 30 cycles of 95 ⁰ C at 15 s, 68 ⁰ C for 15 s, and a final extension at 72 ⁰ C for 120 s (Huhmer et al., Anal. Chem. 2000, 72, 5507-5512, which is incorporated herein by reference). Protocol for amplification of a 211 bp fragment of the virulence B gene on pXOl of the anthrax genome consisted of a 60 s initial denaturation at 95 ⁰ C, followed by 30 cycles of 95 ⁰ C for 5 s, 62 ⁰ C for 5 s, and 72 ⁰ C for 5 s. A final extension at 72 ⁰ C was performed for 30 s. Control amplifications for each of the thermal cycling protocols were performed without template DNA and are shown in the appropriate figures.

After PCR was complete, amplified product was removed from the device, diluted with 24 μL of water, and 0.8 μL of a PCR marker (N3234S, New England BioLabs, Inc., Ipswich, MA, USA). This mixture was then separated on a commercial capillary electrophoresis instrument using laser induced fluorescence detection as described by Legendre et al. Sizing results of amplified PCR products are given as average + standard deviation as determined by comparison of the migration time of the amplified product to the migration time of the sizing ladder.

Results and Discussion

The contents of an optical pyrometer are relatively simple with an optical filter and a detector, typically a thermopile (a collection of thermocouples in series). The basis for using a pyrometer as a temperature sensor is that as an object is heated, the blackbody radiation from the object shifts to shorter wavelengths which fall into the range that can be detected by the pyrometer. In this report, the pyrometer is composed of a thermopile situated behind a Ge lens which is transparent to wavelengths between 8-14 μm. As the solution in the PCR chamber is heated by the IR radiation from the tungsten lamp, the glass above the chamber is also heated producing blackbody radiation that is detected by the pyrometer and recorded as a change in the output voltage of the thermopile. Output voltages of the thermopile were calibrated to temperatures using either a thermocouple, or the boiling points of reference solutions and used to successfully control PCR amplification in microdevices.

Pyrometer Calibration and Temperature Sensing Time Lag

The pyrometer was placed off the vertical axis of the lamp to ensure that the pyrometer was not heated during experiments (FIG. 2). Since the emissivity of glass is known, the surface temperature of the top glass plate can be deduced directly with the pyrometer; however, due to difficulties in maintaining the exact optical properties of glass pre- and post-thermal bonding, a calibration that defined the pyrometer output voltage (from the surface) relative to the PCR solution temperature was performed initially with a thermocouple. One consequence of correlating the surface temperature with solution temperature is the lag time associated with temperature

equilibration between these two regions. The optimal design of a microdevice would have a negligible time lag, allowing the surface temperature to accurately reflect the solution temperature; alternatively, a suboptimal design would render the surface temperature completely insensitive to changes in the corresponding solution temperature. The pyrometer was calibrated by performing five mock PCR cycles between 60 ⁰ C, 95 ⁰ C, and 72 ⁰ C with a dwell time of 15 s at each, and comparing the temperature in the PCR chamber (inserted thermocouple) with that on the surface (pyrometer output voltage). As shown in Figure 3 A, the pyrometer signal was increasing and decreasing in proportion to the solution temperature and was linear. However, further analysis of the traces indicated that the temperature measured by the thermocouple increased prior to an increase in the pyrometer output, indicative of a lag associated with the equilibration of the two regions. Quantitative information on this time lag was made by calculating the difference between the heating rate maxima sensed by the thermocouple and that of the pyrometer (obtained by taking the derivative of the traces with respect to time), the time lag for cooling was measured in a similar manner with the minima of the cooling rates (FIG. 3B). Heating occurred in two steps, the first from 60 - 72 ⁰ C and the second from 72 - 95 ⁰ C , leading to time lags of 1.2 + 0.8 s and 0.3 ± 0.2 s, respectively, whereas cooling involved transitioning from 95 - 60 ⁰ C and corresponded to a time lag of 0.4 + 0.2 s. These positive values indicated that the surface temperature always lagged the solution temperature; however, since the absolute numbers were small, the surface temperature provided an acceptable method for indirect control of the solution temperature. The large error in the measurements likely resulted from the incidental heating of the entire device (not just the PCR domain) over the course of the experiment (discussed in detail below). Ideally, the device could be designed and fabricated such that this

time lag would be zero, which would allow the surface temperature to correspond exactly with the solution temperature. To this end, a one-dimensional heat transfer model was used to investigate the properties of the device and environment that contributed to this time lag. We hypothesized that the underlying reason for the time lag was that the solution in the PCR chamber was heating faster than the glass above the PCR chamber; however, a more qualitative method of defining the reason for this time lag was obtained with a one-dimensional heat transfer model. Using device and environment specifications, the model predicted a time lag of 0.3 s between the solution and surface temperature, correlating, within error, with the data obtained.

The model also defined several parameters that could minimize this time lag - among these were fabrication of a device with a thin cover plate (the thinner the top plate, the more accurately the two temperatures would correlate), a material with a large thermal diffusivity, and performing the experiments in an environment with a small convective heat transfer coefficient (e.g., in a pre-heated chamber). The use of the pyrometer for these cases is currently under investigation.

PCR Amplification ofλ-Phage DNA

Having defined the parameters needed for accurate pyrometer-controlled temperature cycling, amplification of a 500 bp fragment of bacterial λ-phage DNA in a 550 nL volume was attempted. After calibration of the pyrometer, the thermocouple was removed and thermocycling carried out using 15 s dwell times at 95 and 68 ⁰ C for 30 cycles (FIG. 4A). Following the final extension for 120 s at 72 ⁰ C, the PCR solution was removed from the microchip and analyzed by capillary electrophoresis. As shown in FIG. 4B, amplified product was observed indicating that the pyrometer

provided an effective approach for controlling the temperature in low volume chambers on microdevices for enzyme-mediated DNA amplification. While the cycling was relatively slow (~30 minutes) this was due to not having a reflective surface above the device, as was reported by Easley et al. and Legendre et al. for other IR-mediated PCR amplifications.

Boiling Point Calibration

The pyrometer calibration remained valid only when the same microdevice was placed in the same location above the tungsten lamp. This result was due to different devices having slightly different glass thickness and resulting in different rates of heat transfer. Consequently, it became clear that a more simple method for calibrating the pyrometer without the need for a thermocouple or other external hardware was needed. Due to the complexity of using emissivity as a means for obtaining a direct temperature measurement, a method was used to calibrate the pyrometer by the boiling points of reference solutions. The criteria for choosing the solutions included boiling points that were within the range of temperatures normally used in PCR (-60-95 ⁰ C) and the solutions needed to be composed of at least 10% water to approximate IR absorption in a manner similar to absorption by the PCR solution. Two solutions were defined as ideal for this purpose: pure water and an azeotrope consisting of water:2-propanol:toluene at a ratio of 13.1:38.2:48.7 (v:v:v). The rising temperature of the solution was measured through surface sensing {above the microchannel) with the pyrometer which generated the pyrometer voltage versus time plot given in FlG. 5. When the solution reached its boiling point, the temperature remained constant during the phase change, which was detected with the pyrometer as a change in slope (n = 3, FIG. 5). Using the derivative of these plots, the

pyrometer voltage minima were identified and used to calibrate the pyrometer against the known boiling point values for these reference solutions, 76.3 ⁰ C for the azeotrope and 99.6 ⁰ C for water (CRC Handbook of Chemistry and Physics, Student Ed, 69 ^th ed.\ Weast, R. C, Ed.; CRC Press: Boca Raton, FL, 1988). This analysis yielded two reference points with which to calibrate the pyrometer. While one could be skeptical about creating a calibration curve with only two points, it was known from the previous calibration that the pyrometer signal was linear in this temperature range (see FIG. 3A), and it was difficult to find azeotrope solutions that met the aforementioned criteria. It is noteworthy that the microdevice used for all experiments in this section contained channels that were smaller (100 μm in depth) than the thermocouple (-160 μm) to demonstrate the practicality of controlling temperature in low volume conditions (230 nL ellipse).

Following this calibration, the pyrometer was tested for its effectiveness in controlling temperature cycling for the PCR amplification of a 211 bp gene fragment from B. anthracis. After completion of 30 cycles in 33.7 minutes, electrophoretic analysis of the solution in the PCR chamber indicated the presence of a 210 + 3 bp DNA amplicon (FIG. 6). This positive result indicated that with this type of calibration method the need for IR-PCR chambers large enough for insertion of a thermocouple are no longer needed, and lowers the volume limit of these chambers to sub-nanoliter. The use of the boiling point calibration takes advantage of the well- known physical properties of reference solutions, and provides a simple means for calibration and will allow for more portable and self-contained chip-based PCR devices in the future.

Pyrometer Temperature Control with Increased Heating Rates

The cycling times for the amplifications shown in FIGS. 4 and 6 were slow compared to previously-reported IR-mediated amplifications by Easley et al. and Legendre et al. Two factors contributed to the longer cycle times: the substantial mass of glass surrounding the PCR chamber and the lack of a reflective surface above the device to redirect stray radiation back onto the PCR chamber. A simple method for increasing the heating and cooling rates is by etching away a large portion of the mass around the PCR chamber (Easley et al.). However, it is more difficult to increase the heating rates with a reflective surface while still allowing the pyrometer an unobstructed view of the PCR chamber (see FIG. 2). While it has been shown that reflection of IR can be accomplished using a gold-coated glass slide placed 1 mm above the PCR chamber (Easley et al. and Legendre et al.), this was not compatible with the clearance needed for the pyrometer to probe the chamber surface. Alternatively, a 25 mm diameter parabolic gold mirror with a 50 mm focal distance was positioned above the device, allowing ample room for the pyrometer to measure the surface. With the aim of improving heating rates, the pyrometer was calibrated by the boiling point method, and amplification of a fragment of the B. anthracis genome was attempted.

As seen in Figure 7A, with the same thermal cycling parameters used in both FIGS. 4 and 6 to amplify the B. anthracis fragment, the heating rates were increased 44%, with completion of the PCR in 18.8 minutes. The reduced cycling time did not appear to adversely influence the amplification of the 211 bp amplicon, which was sized by capillary gel electrophoresis as a 214 + 2 bp peak relative to DNA standards (FIG 7B), close to the expected 211 bp target. While an overall cycling time of 18.8 minutes was still slower than the IR-mediated PCR reactions of Easley et al. and Legendre et al., the goal of demonstrating remote pyrometer temperature sensing was

accomplished with reasonably fast PCR. Cycling speed could be further enhanced by decreasing the hold time at each temperature and, as mentioned previously, etching away the glass around the PCR chambers for increased heating and cooling rates (Easley et al.).

Device Heating During Thermal Cycling

In addition to the ability to sense temperature in a non-contact fashion, the pyrometer also allowed for observation of a phenomenon that had not been observed with thermocouple-controlled IR-PCR. A qualitative evaluation of temperature cycles indicated that, as thermocycling ensued, successive cycle times decreased, leading to the theory that as the heat dissipated from the PCR chamber to the microchip, the ambient temperature of the entire device increased until the device came to equilibration. If the average temperature of the microdevice itself increased significantly over the course of the temperature cycling, this could be detrimental to the PCR amplification since recalibration of the pyrometer at the elevated temperatures would be required (i.e., the temperature gradient between the solution in the chamber and the surface above the chamber would be different than what was calibrated). While the reactions appeared to be successful, as shown in FIGS. 4B, 6, and 7B, it was possible that amplification efficiency was not optimal due to inaccurate temperature sensing because of the increased ambient temperature at cycles later in the sequence. With the goal of determining the relative temperature change of the entire device during thermal cycling and what design/fabrication parameters could be altered to alleviate this effect, a more quantitative investigation of this phenomenon was attempted.

Evidence to support the hypothesis that the entire microfluidic chip was warming over the course of the thermal cycling sequence was provided by the model discussed above. The times for each successive temperature cycle decreased until the fifth cycle indicating that the entire device had come to an equilibrium temperature by this time. Due to thermal contact between the PCR chambers and the "frame" (the glass around the PCR chambers not involved in storing fluids), this excess glass was also heated at the same time as the PCR chambers, serving as a heat "capacitor" in that heat could be stored during the initial cycles to increase the heating rates of later cycles. The model delineated two methods to reduce this effect: first, thermally isolate the PCR chambers from the frame, and second, implement flow conditions that produce low heat transfer coefficients. While the former method is simply a matter of etching the glass to isolate the PCR chambers, the latter would be difficult to implement as a material of this type would take an excessively long time to cool from a heated state. Practical methods that circumvent (but do not alleviate) this phenomenon are given below.

More quantitative evidence supporting the heating of the entire device was obtained when 30 cycles of mock PCR was controlled by an inserted thermocouple and the pyrometer used to probe the surface temperature (n = 3). As shown in FIG. 8, during each hold step, the normalized pyrometer output increased during the initial cycles until leveling off during the middle and final cycles - this indicated that, while the solution temperature was constant during each hold step (as would be expected with temperature control by the thermocouple), the device temperature did increase until an equilibration was reached. The percentage that the signal changed over the 30 cycles shown in FIG. 8 were 2.7 ± 1.0 %, 4.2 + 1.3 %, and 3.6 + 1.3 % (for denature, anneal, and extension temperatures). While these numbers are not large,

again, this effect may have lowered the amplification efficiency in this report, hi the future, the calibration constants will be updated in a time-dependent or cycle number- dependent manner through the control program so that more accurate temperatures are measured with each cycle. Alternatively, a pre-heating step could be used to equilibrate the device during the first cycle, for example by using a "hot-start" Taq polymerase.

Conclusion

The first completely non-contact PCR device was described using a tungsten lamp and convective air source for heating and cooling, respectively, and a pyrometer for temperature sensing. In the future, we predict that the heat transfer modeling will allow for optimization of microdevice design prior to experimentation instead of just providing a qualitative means by which to explain the phenomenon observed. Also, modeling will help dictate other materials in which to fabricate the microdevices so that the optimal heat transfer properties are utilized which may allow for direct temperature sensing of the PCR solution.

Calibration by boiling solutions of known compositions is ideal not only since it may allow for a completely enclosed system, but also allows for temperature control of solutions in smaller volumes than what can be currently performed using a thermocouple. While DNA amplifications are only shown in this report, this method of temperature control could easily be applied to other temperature-sensitive processes. No matter the application, use of a completely non-contact system will allow for simplification and cost reduction of microdevice fabrication as it alleviates the need for fabricating metal-based temperature sensors into the device.

Example 2 - Development of instrumentation and fabrication methods for rapid non-contact DNA amplification is glass microchip

The temperature of solution in microfluidic reaction chambers can be precisely controlled using a simple tungsten filament lamp for heating and forced air convection for cooling. Non-contact system was shown effective for capillary (Oda et al., Anal Chem 1998, 70, 4361; and Huhmer et al., Anal Chem 2000, 72, 5507, which are incorporated herein by reference) and microchip (Giordano et al., Anal Biochem 2001, 291, 124, which is incorporated herein by reference) platforms. However, the microchip work was carried out on polyimide devices, which are not amenable to downstream integration with electrophoresis due to the incompatibility of this polymer with optical detection. Considering the fact that the research in our lab on solid phase extraction (SPE) and microchip electrophoresis (ME) has been done almost exclusively on glass microdevices, it was imperative for the PCR system to be adapted to this material if a true micro-total analysis system (μTAS) was to be realized. Unfortunately, as shown in FIG. 9, this substrate was found to be detrimental to the heating rates (at typical chip thicknesses of 2.2 mm), resulting in PCR cycling times (0.5 h) that were hardly an improvement over those of conventional heating block instruments. When the large surface area-to- volume ratio in microfluidic chambers is combined with the glass thermal conductivity (1.4 Wm ^"1 °C ^"! ) that is 10-fold larger than polymer devices, heat is dissipated from the solution much to quickly to permit effective thermal cycling. Furthermore, our group has shown that glass surfaces may inhibit the PCR reaction (Giordano et al., Electrophoresis 2001, 22, 334, which is incorporated herein by reference), especially at such large surface area-to-volume ratios (~0.1 μm ^"1 ). In order to integrate PCR with other sample processing and analysis techniques on glass devices, it was

necessary to overcome the problems encountered using this substrate. The work in this example outlines the steps taken toward achieving this goal, in which novel device fabrication, system design techniques, and surface passivation chemistry allowed the cycling rates in glass devices to exceed those observed with previous polymer systems. As a result, 25 cycles of successful PCR is accomplished in only 5 minutes on glass devices, a dramatic improvement over the 30-90 min required in conventional systems.

Background Polymerase Chain Reaction on Microfluidic Chips

Since the technique was first described two decades ago, the polymerase chain reaction (PCR) has become an essential tool in the field of genetic analysis, providing an in vitro method to amplify DNA sequences of interest (described further in Chapter 1). However, while conventional techniques are improving in speed, they are still time-consuming (1-3 hours per amplification), and the reagents are expensive at the volumes needed for manual transfer of samples between pre-treatment, amplification, and post-amplification analysis steps. Wittwer et al. (Biotechniques 1991, 10, 76, which is incorporated herein by reference) emphasized that precise thermal control during temperature holds and rapid transition times between temperature set points should result in optimal PCR efficiency. Although transition times in conventional cyclers are much too slow to realize this, miniaturization of the PCR has been shown to do just that, significantly reducing amplification time as well as reagent consumption. Decreasing volume and increasing the surface area-to-volume ratio using microscale reaction chambers allows homogeneous solution temperatures to be achieved much more rapidly than in conventional heating blocks. Microfluidic

systems provide the possiblity to cycle rapidly as well as to integrate methods into single μ-TAS.

Mechanisms of Thermal Cycling As an alternative method of heating, our group has pioneered the use of infrared (IR) radiation to selectively heat a static reaction solution through excitation of vibrational modes of water molecules. This direct heating method, as shown in FIG. 1OC, can additionally reduce the time and energy needed for thermal cycling in microfluidic chambers; the reaction container (substrate) is not heated and can even serve as a heat sink for enhancement of the cooling steps to follow. The greatest potential for maximizing amplification speed and efficiency lies in the use of this type of direct heating mechanism, one in which the majority of the input energy is transferred actively and selectively into the reaction solution — namely through radiative or electrical means — to limit the thermal inertia to the solution of interest, typically <1 μL of aqueous solution. This method drastically decreased amplification times as well as increased efficiencies when coupled with convective air cooling of nL or low μL reaction volumes. Non-contact, infrared-mediated PCR carried out in glass capillaries (Oda et al. and Hulmer et al.) resulted in heating and cooling rates as high as 65 and -20 ⁰ C s ^"1 , respectively, and was shown to be more efficient than conventional methods. Non-contact PCR in polyimide microchips (Giordano et al.) was accomplished in less than 4 min by reducing the cycle number to 15 and adding 6.9 x 10 ⁴ starting copies. However, the polyimide substrate is less conducive to integrating optical components for the detection of products because it is not transparent throughout the entire visible spectrum. Using the same infrared-mediated system with a more characterized substrate (glass), it was possible to integrate non-

contact PCR with microchip electrophoresis (ME) in 12.2 min, representing the fastest achieved on glass devices to date {19). However, because the surrounding glass acted as a heat sink to the PCR chamber, heating and cooling rates of 13.4 and - 6.4 ⁰ C respectively, fell dramatically short (by ~5-fold) of the rates achieved in the capillary system (Oda et al. and Hulmer et al). We hypothesized that it should be possible to approach. — or even surpass — the cycling rates achieved by the capillary system with further optimization of the thermal characteristics of our non-contact system. This enhancement was, indeed, possible (as discussed below), resulting in the fastest PCR reaction achieved to-date in glass devices at 25 cycles in only 5 min. Furthermore, this non-contact method provides the means to reduce the cost of individual devices, especially when single-use, disposable devices are desired. We have also developed an interferometric approach for non-contact means of temperature sensing (U.S. Patent Application Publication No. 2004/0131504, which is incorporated herein by reference), allowing complexities and cost to be confined to the instrumentation, not the microdevice. This approach allowed for the development of non-contact temperature control system for microfluidics.

Temperature Control Programming The previous microchip PCR control program (written in LabVIEW code) utilized proportional feedback control, in which the power delivered to the heating lamp was proportional to the temperature difference between the signal and set point. While this system was functional, it should be possible to enhance the temperature control using a proportional-integral-derivative (PID) algorithm. For this reason, an improved PCR control program was written in LabVIEW using PID feedback control. Briefly, the power delivered to the digitally-controlled lamp was programmed as a

flicker rate (FR), which essentially defined the duty cycle of the lamp. The FR was defined as

FR = P {T _s - T) + D — (T _S - T)+ lj(T _s -T)dt (eq. 2) dt

where P, D, and I are the proportional gain, derivative gain (or damping constant), and integral gain (or controller reset level), respectively. T ₅ is the set temperature, and T is the temperature signal being measured. The control system was then tested for its ability to reach a set temperature of 72 ⁰ C and maintain that temperature. As can be seen in the figure, the set temperature was reached rapidly without overshooting and was maintained without significant ringing at 72.18 ± 0.05 ⁰ C. The system was subsequently tested for its ability to carry out a mock PCR cycling (94 ⁰ C, 15 s; 3 cycles of 60 ⁰ C for 2 s, 72 ⁰ C for 3 s, and 94 ⁰ C for 3 s; 72 ⁰ C). As shown in FIG. 11 , the PID control was sufficient for the mock cycling. Note, however, that the most optimal control was attained at the 72 ⁰ C set point, which is the temperature that was used to optimize the system, and the control at the other temperatures was less precise. This suggested that the PID constants should be a function of the set temperature. However, because the PID constants will need to be optimized for each different microchip configuration, the extent of control shown in FIG. 11 was deemed sufficient at this point, with all temperatures held at <0.5 ⁰ C deviation. More recently, the system was further optimized by making use of the pulse- width modulation (PWM) functions available with LabVIEW programming. The counter pin from the DAQ card (National Instruments) was used as the digital control of the tungsten filament lamp, which allowed a pulse frequency of 1 kHz at any specified duty cycle, effectively giving an analog voltage control that was defined by the flicker rate shown above. Using this method, the temperature control

programming could be greatly simplified. The cycling was no longer split into discrete steps, but the set temperature was simply changed at the user-defined times, and the feedback control was allowed to adjust accordingly. With the PWM programming, temperature control was much more refined, even using device configurations with large cycling rates. FIG. 12 shows the utility of this method, in which heating and cooling rates exceeding 20 ⁰ C s ^"1 were still well-controlled at the set temperatures.

Calibration Methods A standard method was developed for calibration of the miniature thermocouples for microchip PCR. The sensor is first immersed in a polypropylene PCR tube filled with 50 μL of distilled water, then the tube is placed in a conventional heating block-based thermal cycler which is set to transition through various temperatures relevant to PCR (50, 61, 72, 83, and 94 ⁰ C). The average signal (in volts) is then taken for each set point during the last 10 s of the 30-s hold periods. FIG. 13 illustrates the quality of calibration that is possible using this method. The thermocouple signal (in volts) during two separate runs of cycling is shown in FIG. 13 a, with two runs carried out to obtain voltage values in triplicate at each set temperature. An example calibration curve is shown in FIG. 13b, with excellent linearity (R ² = 0.9998), as expected. Error bars are included (orange), with all RSD values at 0.1% or below (n=3). This curve resulted in a calibration equation of T = (13.062 ⁰ C V ^"1 x S) + 0.3078 ^C C, where T is the temperature in ⁰ C and S is the signal in V, The reciprocal of this slope gave the calibration sensitivity to be 76.6 mV ⁰ C ' . Since the amplification circuit was designed for a final sensitivity of -75 mV ⁰ C ^" , this data matched very closely with the expected value.

While the heating block method proved to be very useful and robust, it was discovered that the miniature thermocouple may be absorbing radiation from the tungsten filament lamp used for heating. This effect seemed to be more pronounced as the microchip substrate was thinned around the reaction chambers. In order to test the extent of this absorption, the thermocouple was placed in the center of a microfluidic chamber, which corresponded to the center of the lamp's focal region, and the lamp was turned on full power. The idea was to heat the solution to boiling and to detect this boiling point as a discontinuity in the temperature versus time trace. Since the chamber was filled with distilled water, it was expected that boiling would occur at 99.5 ⁰ C (boiling point of water adjusted for the elevation and barometric pressure). However, as shown in FIG. 14a, the boiling discontinuity was seen at 102.5 ± 0.5 ⁰ C (three separate runs are shown for each). This incorrect boiling point determination provided proof that the thermocouple was biased by infrared absorption. In fact, when the thermocouple sensor was placed in the focus of the lamp, excluding the microchip and solution, the signal increased very rapidly. A second test was then performed, in which the thermocouple was placed at the edge of the reaction chamber, rather than the center. In this configuration, the boiling point was correctly determined to be 99.6 ± 0.5 ⁰ C (FIG. 14b). Presumably, with the thermocouple slightly out of the central focus of the lamp, the radiation absorption bias becomes negligible at this time-scale. From these observations, it was concluded that the thermocouple sensor should always be placed at the edge of the reaction chambers, not in the center where radiation absorption bias can occur.

Following the boiling point determination experiments above, it was hypothesized that this method could be used to calibrate the system without the necessity of a conventional PCR instrument. Furthermore, with ongoing research in

our laboratory concerning non-contact temperature detection, this method could provide a means for non-invasive calibration. Even if a radiation absorption bias was present, this should provide a means to compensate for that bias. Of course, a calibration could not be carried out with only one boiling point; at least two points were needed. Ideally, several solutions with boiling points spanning the range of typical PCR temperatures (50-100 ⁰ C) could be used for this purpose, but the solutions must not change composition upon boiling, or the boiling points would constantly change value. It was, therefore, proposed that azeotropic solutions would be ideal for this purpose, because — by definition — their composition would not change upon boiling.

Accordingly, a search was conducted for azeotropic solutions that met three criteria : 1) having boiling points between 50-100 ⁰ C; 2) containing at least 10% water to provide similar IR absorption; and 3) being readily available. The CRC Handbook of Chemistry and Physics had listed hundreds of azeotropic solutions (binary and ternary) along with their compositions and boiling points. Based on the above three criteria, the boiling point calibration list was narrowed down to the solutions shown in Table 1.

solution X %X Y %y Z %Z bp ('C)

1 Water 1Q0 9956

2 Water 42.5 butyl alcohol 57.5 92.7

3 Water 60 π-butyl alcohol 14.6 n-octanβ 25 86.1

4 Water 13 1 isopropyl alcohol 38 2 toluene 48.7 76 3

Table 1 Clearly, water (solution 1) would serve as one of the solutions. However, it was difficult to find solutions that included at least 10% water yet had a boiling point lower than 95 ⁰ C to provide a wider range of points in the calibration. Based on this,

as well as immediate availability of the solvents, solution 4 from Table 1 was chosen as the other calibration standard. A two-point, noninvasive calibration could then be carried out using pure water (solution 1 ; bp = 99.6 ⁰ C) and a ternary azeotrope (solution 4; water:2-propanol:toluene, 13.1:38.2:48.7 (by volume); bp = 76.3 ⁰ C). Ideally, all four of the solutions in Table 1 could be used to provide a 4-point calibration in the PCR temperature range. This method has been proven useful for calibration of a completely non-contact temperature sensing and control system for microchip PCR, in which an infrared pyrometer is used for temperature detection.

Thermal Management in Microfluidic Devices

Not only is thermal control important for microchip PCR, it becomes a vital consideration in microscale chemistry such as synthetic or biological reactions. In the pursuit of μ-TAS, in which a variety of processes of diverse temperature may be desired, control of temperature in discrete zones becomes an import parameter. While some have investigated thermal management in microscale devices, few have emphasized the importance of thermal isolation with emerging lab-on-a-chip systems. Although glass has proven an ideal substrate for microfluidic applications (Easly et al.; and Lagally et al., Journal of Physics D: Applied Physics 2004, 37, R245) due to chemical and electrical inertness, and its optical clarity, it is naturally incompatible with heat isolation due to the relatively large thermal conductivity ( 1.4 W m ^{"1 0} C ^'1 ) compared to plastic devices such as polycarbonate (0.16 W m ^{"1 0} C ^"1 ). Initial work was carried out by Poser et al. (Sensors and Actuators, A: Physical 1997, A62, 672) in silicon devices, where the reaction chambers were bordered with "thermal gaps" to insulate them from the bulk silicon, providing fast cycling with low power consumption. Krishnan et al. (Sensors and Actuators, A: Physical 2002, A95, 250)

showed that heating a non-isolated reaction chamber to 65 ⁰ C in a glass device caused a chamber 5 mm away to heat to 47 ⁰ C, whereas the polycarbonate-glass device heated the adjacent chamber to only 37 ⁰ C. Using a silicon-glass device, thermal crosstalk between regions up to 10 mm apart was shown to be severe (Yang et al., Journal of Micromechanics and Microengineeήng 2005, 15, 221), with equilibrium temperature of the entire device reaching at least 80 ⁰ C if one section was held at 95 ⁰ C. Three thermal isolation methods were then devised to solve this problem; thermal conduits, silicon back dicing, and silicon back etching were shown to be effective techniques for isolating reaction chambers (Yang et al.). However, the thermal conduit process was largely manual and not conducive to automation, the silicon back dicing technique could only create trenches as deep as 450 μm depth, and the silicon back etching method required complex fabrication steps. Sadler et al. {IEEE Transactions on Components and Packaging Technologies 2003, 26, 309) reported on thermal management of continuous-flow PCR devices fabricated from ceramic, utilizing air gaps for thermal isolation of heated zones. Ceramic substrates, however, are incompatible with optical interrogation and standard semiconductor fabrication technologies, and the PCR times on these devices were unimpressive at 1.5 hours. Furthermore, the air gaps were milled, which is less compatible with batch fabrication than an etching process. A more ideal solution was devised by Kajiyama et al. {Genome Research 2003, 13, 467), in which microfabrication was used to create discrete, independently temperature-controlled islands for DNA microarray hybridization studies. Even in a device with high thermal conductivity (148 W m ^" ' ⁰ C ^"1 ), with islands only 500 μm apart, nearly complete thermal isolation was successful using automatable fabrication steps. The aforementioned studies outline the importance of thermal control in microfluidics. Prior to this work, a method was

still needed for thermal isolation of discrete zones in glass microdevices that is conducive to high density architectures and automated fabrication. The work that follows will outline the development of such a method.

Initial descriptions from our lab on non-contact nanoliter- scale PCR in glass capillaries (Oda et al.; and Hulmer et al.) suggested that limiting the amount of glass (thermal mass) surrounding the PCR chamber could allow for enhanced thermal cycling. In this section, a method is described for controlled thermal isolation of microfluidic reaction chambers for PCR, extending the concepts proven by silicon back etching (Yang et al.) and thermal island fabrication (Kajiyama et al.) to create a mimick of the capillary structures shown previously to be so effective for thermal transfer during temperature cycling (Oda et al.; and Hulmer et al.). The glass is removed using the same wet etching protocol used to create microchannels and is, therefore, conducive to batch fabrication. Theoretical modeling was shown to fit well with experimental heating and cooling curves, and the method was used to fabricate thermally-isolated PCR microchips capable of DNA amplification in only 5 min. This work represents the fastest PCR carried out in glass devices to date. Furthermore, the fabrication protocol is compatible with standard photolithography and wet etching, making it useful for creating discrete, thermally-isolated zones for μ- TAS.

Concluding Remarks

A non-contact, infrared-mediated thermal cycling system was further optimized and adjusted for DNA amplification in glass microfluidic devices. Instrumentation developments provided increased thermal control, and surface passivation of the glass allowed the sensitive enzymatic reaction to be carried out

successfully and reproducibly. Additionally, the fabrication method outlined in this work was shown provide heat transfer characteristics that matched well with theoretical predictions, producing thermally-isolated microchip reaction chambers in glass that allowed non-contact PCR in only 5 minutes in a 270-nL volume. Since this thermal isolation method is compatible with standard photolithography and wet etching, it is feasible to extrapolate to more complex devices. In situations where differential thermal control is necessary on a single device, heat transfer characteristics can be precisely manipulated using microfabrication. Formation of discrete, thermally-isolated zones could be envisioned to allow for separate 'hot' and 'cold' regions on a single μ-TAS. Stable thermal gradients could also be designed into the device architecture. Regardless of the particular application, the thermal isolation methods presented in this work should prove to be a useful fabrication tool for a broad range of tnicrofluidic devices.

With the rapid cycling achieved in this work, it should be possible to study the base-incorporation rate of Taq polymerase. With a typical rate of 60- 150 bp s ¹

(Hashimoto et al., Lab on a Chip 2004, 4, 638), typical PCR amplicons require at least a few seconds at the extension step of PCR (72 ⁰ C). The rates achieved in this work essentially provide a method to cycling faster than the enzyme can incorporate bases. It should, therefore, be possible to study the kinetics of the enzyme using this platform.

Although certain presently preferred embodiments of the invention have been specifically described herein, it will be apparent to those skilled in the art to which the invention pertains that variations and modifications of the various embodiments shown and described herein may be made without departing from the spirit and scope

of the invention. Accordingly, it is intended that the invention be limited only to the extent required by the appended claims and the applicable rules of law.