BACKGROUND

Nonlinear dynamics, in conjunction with microfluidics, play a central role in the design of the devices and the methods according to the invention. Microfluidics deals with the transport of fluids through networks of channels, typically having micrometer dimensions. Microfluidic systems (sometimes called labs-on-a-chip) find applications in microscale chemical and biological analysis (micro-total-analysis systems). The main advantages of microfluidic systems are high speed and low consumption of reagents. They are thus very promising for medical diagnostics and high-throughput screening. Highly parallel arrays of microfluidic systems are used for the synthesis of macroscopic quantities of chemical and biological compounds, e.g., the destruction of chemical warfare agents and pharmaceuticals synthesis. Their advantage is improved control over mass and heat transport.

Microfluidic systems generally require means of pumping fluids through the channels. In the two most common methods, the fluids are either driven by pressure or driven by electroosmotic flow (EOF). Flows driven by EOF are attractive because they can be easily controlled even in complicated networks. EOF-driven flows have flat, plug-like velocity profile, that is, the velocity of the fluid is the same near the walls and in the middle of the channel. Thus, if small volumes of multiple analytes are injected sequentially into a channel, these plugs are transported as non-overlapping plugs (low dispersion), in which case the dispersion comes mostly from the diffusion between plugs. A main disadvantage of EOF is that it is generated by the motion of the double layer at the charged surfaces of the channel walls. EOF can therefore be highly sensitive to surface contamination by charged impurities. This may not be an issue when using channels with negative surface charges in DNA analysis and manipulation because DNA is uniformly negatively charged and does not adsorb to the walls. However, this can be a serious limitation in applications that involve proteins that are often charged and tend to adsorb on charged surfaces. In addition, high voltages are often undesirable, or sources of high voltages such as portable analyzers may not be available.

Flows driven by pressure are typically significantly less sensitive to surface chemistry than EOF. The main disadvantage of pressure-driven flows is that they normally have a parabolic flow profile instead of the flat profile of EOF. Solutes in the middle of the channel move much faster (about twice the average velocity of the flow) than solutes near the walls of the channels. A parabolic velocity profile normally leads to high dispersion in pressure-driven flows; a plug of solute injected into a channel is immediately distorted and stretched along the channel. This distortion is somewhat reduced by solute transport via diffusion from the middle of the channel towards the walls and back. But the distortion is made worse by diffusion along the channel (the overall dispersion is known as Taylor dispersion).

Taylor dispersion broadens and dilutes sample plugs. Some of the sample is frequently left behind the plug as a tail. Overlap of these tails usually leads to cross-contamination of samples in different plugs. Thus, samples are often introduced into the channels individually, separated by buffer washes. On the other hand, interleaving samples with long buffer plugs, or washing the system with buffer between samples, reduces the throughput of the system.

In EOF, flow transport is essentially linear, that is, if two reactants are introduced into a plug and transported by EOF, their residence time (and reaction time) can be calculated simply by dividing the distance traveled in the channel by the velocity. This linear transport allows precise control of residence times through a proper adjustment of the channel lengths and flow rates. In contrast, dispersion in pressure-driven flow typically creates a broad range of residence times for a plug traveling in such flows, and this diminishes time control.

The issue of time control is important. Many chemical and biochemical processes occur on particular time scales, and measurement of reaction times can be indicative of concentrations of reagents or their reactivity. Stopped-flow type instruments are typically used to perform these measurements. These instruments rely on turbulent flow to mix the reagents and transport them with minimal dispersion. Turbulent flow normally occurs in tubes with large diameter and at high flow rates. Thus stopped-flow instruments tend to use large volumes of reagents (on the order of ml/s). A microfluidic analog of stopped-flow, which consumes smaller volumes of reagents (typically μL/min), could be useful as a scientific instrument, e.g., as a diagnostic instrument. So far, microfluidic devices have not be able to compete with stopped-flow type instruments because EOF is usually very slow (although with less dispersion) while pressure-driven flows suffer from dispersion.

In addition, mixing in microfluidic systems is often slow regardless of the method used to drive the fluid because flow is laminar in these systems (as opposed to turbulent in larger systems). Mixing in laminar flows relies on diffusion and is especially slow for larger molecules such as DNA and proteins.

In addition, particulates present handling difficulty in microfluidic systems. While suspensions of cells in aqueous buffers can be relatively easy to handle because cells are isodense with these buffers, particulates that are not isodense with the fluid tend to settle at the bottom of the channel, thus eventually blocking the channel. Therefore, samples for analysis often require filtration to remove particulates.

SUMMARY ACCORDING TO THE INVENTION

In one aspect, a microfluidic system includes a substrate with one or more inlet ports, one or more outlet ports, and an array of traps in fluidic communication with the inlet and outlet ports, and with one or more means for applying force to introduce a fluid via the one or more inlet ports so that a plurality of fluid portions are trapped in the array of traps. Each fluid portion includes reagents sufficient for an autocatalytic reaction including a first species of molecule in a concentration such that each fluid portion contains no more than a single molecule of the first species, and wherein the fluid portions are trapped at least for a period of time sufficient for the autocatalytic reaction to occur such that the single molecule is amplified.

In another aspect the fluid portions are moved to the one or more outlet ports after the period of time has passed. The one or more of the outlet ports may also be an inlet port.

In another aspect, the single molecule is a single biological molecule. The single biological molecule may be DNA or RNA and the autocatalytic reaction may be a polymerase-chain reaction.

In yet another aspect, the system further includes a detector to detect, monitor, or analyze one or more properties of the autocatalytic reaction during and/or after it has occurred. The properties may include light emission. The reagents may include one or more fluorescent labels and the detector may detect emissions from the one or more fluorescent labels.

In a further aspect, the fluid portions are trapped for a period of time that may range from tens of microseconds to hours to weeks.

In another aspect, the array of traps includes an array of connected microchannels. The system may further include one or more valves that are used to control the flow of fluid portions into and/or out of the traps.

In another aspect, each fluid portion is a fluid plug. Each fluid plug may be substantially spherical in shape. In addition, each fluid plug may be substantially surrounded by and immiscible in a carrier fluid.

In yet another aspect, the means for applying force may be a means for applying pressure.

