GB2569417A

GB2569417A - Microfluidic drive system

Info

Publication number: GB2569417A
Application number: GB1812482.6A
Authority: GB
Inventors: Harrison Tom; Hatfield Stuart; Crossley Peter
Original assignee: TTP Ventus Ltd
Current assignee: TTP Ventus Ltd
Priority date: 2018-07-31
Filing date: 2018-07-31
Publication date: 2019-06-19
Anticipated expiration: 2038-07-31
Also published as: GB201812482D0; GB2569417B

Abstract

A microfluidic drive system includes a resonant gas pump 1 comprising a substantially cylindrical cavity 110, defined by cavity walls 170 and having inlet and an outlet aperture 150, 160 and a piezoelectric actuator 200 arranged to generate oscillatory motion of the cavity walls to drive a gas between the inlet and outlet. The microfluidic drive system further comprises a drive circuit (14, figure 3) arranged to apply a voltage waveform across the piezoelectric actuator, such that the oscillations of the cavity have a frequency of at least 500 Hz and a microfluidic channel (9, figure 1A) arranged in fluid communication with the inlet or outlet such that in use the varying gas pressure provides a driving force to move a liquid through the microfluidic channel. A liquid reservoir may be arranged between the pump and the channel. A sensor(s) and / or transducer may be arranged to measure to measure power supplied to the pump, gas pressure, gas flow rate, liquid pressure, liquid differential pressure and liquid flow rate. A user interface element may be included which preferably comprises a light emitting element, sound emitting element and vibrating element. A further pump is also claimed.

Description

The present invention relates a microfluidic drive system, in particular a microfluidic drive system employing pressure-driven flow.

BACKGROUND

Microfluidic devices are often used in point-of-care healthcare devices to provide local testing of biological samples to provide near instant diagnostic information.

In such devices, it is desirable to avoid cross-contamination between biological samples from different patients when processed by testing equipment. Many known devices address this problem by having disposable and reusable elements, with the sample only coming into physical contact with the disposable part such as a microfluidic cartridge.

Some microfluidic systems move liquids solely under capillary action such as lateral flow systems. However this can limit the flexibility or complexity of an assay. To address these limitations, the sample often needs to be actively pumped through the disposable element and therefore a pump is required. It is therefore preferable to incorporate the pump in the reusable part of the system so that the cost of the disposable part can be minimised. Therefore, the interface between the reusable and disposable parts of the system needs to prevent the sample from contaminating the pump and reusable components of the device, thereby preventing crosscontamination of subsequent samples.

One approach to addressing these requirements is to use pressure-driven flow, alternatively referred to as air-over-liquid. This technique involves the use of an air pump, where the air pressure generated by the pump is used to drive the liquid flow of the sample. In this regime, a combination of a buffer of air and optionally a filter between the reusable part of the system and the disposable part is used to form an interface that prevents the sample from contaminating the pump or other aspects of the reusable part. Pressure-driven flow systems often employ fast-acting pressure regulation hardware. A combination of high response bandwidth and the compressibility of air combine to deliver relatively smooth liquid flow.

In point of care applications, there is a growing trend to create smaller, lighter devices. This might be a reduced-footprint benchtop instrument, or more pertinently, a handheld device which may need to be small and lightweight for ergonomic reasons. Given this trend, it is often desirable for any pump system used to be similarly small and lightweight. However, traditional pressure-driven flow systems are typically bulky and may require several pieces of hardware, including for example an external pressure source, the need for which is not ideal for portable/handheld devices, and pressure regulation hardware such as valving.

Some such lightweight devices utilise traditional pumps such as diaphragm pumps to provide the driving force to move the liquid sample. However such pumps tend to introduce oscillations in air pressure resulting in oscillatory liquid flow which is prevents the correct functioning of such devices or limits the accuracy of the results. In particular many point-of-care applications require a smooth liquid flow to achieve the accuracy of a sorting process, the precision of a dosing or metering operation, or the repeatability of microfluidic droplet generation. For such devices additional damping hardware may be necessary to reduce the oscillations. This additional hardware increases the cost, size and complexity of such devices, which precludes their inclusion in the above described, partly disposable lightweight handheld devices.

Accordingly there is a need for a microfluidic drive system which can provide a smooth fluid flow output whilst being small and readily able to be incorporated into such devices.

SUMMARY OF THE INVENTION

According to a first aspect of the invention, there is provided a microfluidic drive system comprising: a resonant gas pump comprising a substantially cylindrical cavity defined by cavity walls, the cavity having an inlet and an outlet aperture and a piezoelectric actuator arranged to generate oscillatory motion of the cavity walls to drive a gas between the inlet and outlet; a drive circuit arranged to apply a voltage waveform across the piezoelectric actuator, such that the oscillations of the cavity have a frequency of at least 500 Hz; a microfluidic channel arranged in fluid communication with the inlet or outlet of the pump, such that, in use, the varying gas pressure provides a driving force to move a liquid through the microfluidic channel.

With the microfluidic drive system according to the present invention a significantly reduced pressure pulsatility is provided due to arrangement of the high frequency resonant gas pump which provides a smooth fluid flow. With the claimed arrangement no damping hardware is required, reducing the system complexity, cost and size. Furthermore, the arrangement is capable of displacing a greatly reduced volume of gas per cycle which further reduces any oscillations in the driven liquid. The microfluidic drive system has the additional benefit over traditional pumps of being quiet and having vibration-free operation. These factors are particularly advantageous when incorporated into a point-of-care device where the reduced vibration improves the user experience for both the device user (typically a healthcare professional) and the patient.

Preferably the microfluidic drive system further comprises at least one sensor arranged to measure a property of the system and connected to the drive circuit such that, in use, the drive circuit adjusts the applied voltage waveform in response to a measured property received from the sensor. In this way, the system may compensate for the effect of factors such as gas temperature, pressure and composition, and by part to part variation to ensure accurate and precise liquid flow is maintained. In particular, the resonant mode of operation means that the absolute repeatability of the rate of fluid flow driven by the pump is limited by the above factors. By providing a sensor arranged to measure a property of the system and controlling the characteristics of the voltage waveform applied to the actuator to drive the oscillations, the pump output may be controlled to provide reliable desired flow characteristics through the microfluidic channel.

BRIEF DESCRIPTION OF THE DRAWINGS

Examples of the invention will now be described, by way of example only, with reference to the accompanying drawings, in which:

Figure 1A and 1B schematically illustrate a microfluidic drive system according to the present invention in “push” and “pull” arrangements respectively;

Figure 1C and 1D schematically illustrate a resonant gas pump used in the microfluidic drive system according to the present invention;

Figure 2A and 2B schematically illustrate a microfluidic drive system according to the present invention incorporating a reservoir;

Figure 3 schematically illustrates a drive circuit used in a microfluidic drive system according to the present invention;

Figure 4 schematically illustrates a microfluidic drive system according to the present invention which incorporates a sensor which feeds back to the drive circuit;

Figure 5 schematically illustrates a microfluidic drive system according to the present invention which incorporates a gas pressure sensor which feeds back to the drive circuit;

Figure 6 schematically illustrates a microfluidic drive system according to the present invention which incorporates a gas flow rate sensor which feeds back to the drive circuit;

Figure 7 schematically illustrates a microfluidic drive system according to the present invention which incorporates a liquid pressure sensor which feeds back to the drive circuit;

Figure 8 schematically illustrates a microfluidic drive system according to the present invention which incorporates a liquid flow rate sensor which feeds back to the drive circuit;

Figure 9 schematically illustrates a microfluidic drive system according to the present invention which incorporates multiple sensors which feed back to the drive circuit;

Figures 10A and 10B schematically illustrate a microfluidic drive system according to the present invention which incorporates a check valve;

Figure 11 schematically illustrates a microfluidic drive system according to the present invention which incorporates a filter;

Figures 12A and 12B schematically illustrate a microfluidic drive system according to the present invention which incorporates an accumulator;

Figures 13A to 13C schematically illustrate pump arrangements for a microfluidic drive system according to the present invention which incorporate multiple pumps;

Figures 14A to 141 schematically illustrate pump arrangements for a microfluidic drive system according to the present invention which incorporate a manifold device;

Figure 15 schematically illustrates a pump electrical interface for a microfluidic drive system according to the present invention.

Figure 16 schematically illustrates a microfluidic drive system according to the present invention.

DETAILED DESCRIPTION

Figures 1A and 1B schematically illustrate basic examples of the microfluidic drive system according to the present invention. The system comprises a resonant piezo-acoustic gas pump 1 having an output 11 and input 12 and a microfluidic channel 9 arranged in fluid communication with the output 11 (Figure 1A) or input 12 (Figure 1B) of the pump 1. Figures 1C and 1D illustrate the features of the resonant piezo-acoustic gas pump 1 which comprises a substantially cylindrical cavity 110 having an inlet and outlet aperture 150, 160 and a piezoelectric actuator 200 arranged to generate oscillatory motion of the cavity walls 170, as shown in Figure 1D. The oscillations of the cavity 110 have a frequency of at least 500 Hz which are provided by applying a voltage waveform across the piezoelectric actuator 200 with a drive circuit in order to provide the driving force to move a liquid through the connected microfluidic channel 9.

In certain examples the pump is arranged such that it moves only a small volume of gas each cycle that it operates, as represented by the airflow 2. The small volume of gas is generally between about 1 pL and 1 mL to provide the advantages in terms of smooth flow and low pulsatility. However improved effects can be achieved with smaller volumes, for example less than 500 pL, with particularly advantageous effects found for 1 nL to 100 pL. Once practically integrated into a pressure-driven flow system, the ratio of the volume of the gaseous headspace used to drive the fluid and the volume of air the pump moves per cycle is so substantial so as to create virtually no pressure pulsatility in the headspace once the pump is running. Pulse-free gas pressure in the headspace will result in a pulse-free pressure differential across the liquid that will ultimately translate to smooth liquid flow 6. The terms “low pulsatility” and “smooth output” are used interchangeably.

A particularly advantageous pump arrangement is illustrated in Figures 1C and 1D. In this example, the cavity 110 is defined by end walls 120 and 130, and a side wall 140. The cavity is substantially cylindrical in shape, although elliptical and other shapes could be used. The cavity 110 is provided with a nodal air inlet 150, which in this example is unvalved although could be valved and located substantially at the centre of the end wall 130. There is also a valved air outlet 160 located substantially at the centre of end wall 130. The upper end wall 120 is defined by the lower surface of a disc 170 attached to a main body 180. The inlet and outlet pass through the main body 180.

In the example of Figures 1C and 1D the actuator 200 comprises a piezoelectric disc attached to a disc 170. Upon actuation, the actuator 200 is caused to vibrate in a direction substantially perpendicular to the plane of the cavity 110 as shown in Figure 1D, thereby generating radial pressure oscillations within the fluid in the cavity 110.

In Figure 1A, the pump pneumatic output 11 is connected to a microfluidic channel 9 so that positive air pressure 3 can be supplied to a headspace 4 adjacent to a liquid 5. This positive air pressure can be used to create a pressure differential between the headspace and the gas pressure at other gas-liquid interfaces within the system such as the downstream pressure 10 represented in Figure 1A. If this pressure differential is sufficient, it can drive movement of the liquid 5 as represented by the liquid flow 6. In Figure 1B, a similar example is shown, but where the pump pneumatic input 12 is connected to the microfluidic channel 9 so that negative air pressure 13 can be supplied to the headspace 4. In this case, a pressure differential can be created in the opposite direction so as to drive the fluid in the opposite direction to that described in Figure 1A. Figure 1A can be considered a “push” example, and Figure 1B a “pull” example.

In either the push or pull cases, the pump 1 can be connected to more than one microfluidic channel to drive multiple liquid flows in parallel, for example. It is also conceivable that valves might be used in conjunction to prevent or allow the airflow 2 to a headspace 4, or the onward liquid flow driven by the pressure in the headspace, such that liquid flow in individual channels could be enabled or prevented.

One possible limitation of the examples proposed in Figures 1A and 1B is that the total liquid volume available to pump is limited by the volume offered by the microfluidic channel 9. Certain applications may require that a greater volume of liquid is used, for example in the field generally referred to as “large volume injection”, where liquid volumes in the 2 to 100 ml range are often used. Figures 2A and 2B present two simplified schematic examples where a liquid reservoir 8 has been introduced to hold a reserve volume of liquid that is independent of the volume of the microfluidic channel 9. Push and pull variants are shown in Figures 1A and 1B respectively. In these examples, liquid is driven from the reservoir into the microfluidic channel to the rest of the system such that a greater volume of liquid may be held and transported through the system. Amongst other applications, these embodiments shown in Figures 2A and 2B have potential applications in aspiration and dispensing in cell culture applications. In particular in aspiration applications, the reservoir provides a large volume in which to capture or trap a certain volume of liquid such as a waste liquid from a cell culture or preparation process. This reservoir may be necessary to prevent the liquid from reaching the pump, which might cease operating if certain internal parts were to be exposed to liquid.

In some examples of either the push or pull cases presented in Figures 2A and 2B, the pump is connected to more than one liquid reservoir, the reservoirs themselves connected to one or more microfluidic channels to drive multiple liquid flows in parallel, for example. It is also possible to provide valves in conjunction in such examples so that that liquid flow could be selectively enabled or prevented within the channels.

Variable pump drive scheme

The examples so far described use a resonant piezo-acoustic gas pump with inherently low-pulsatility output used to drive smooth liquid flow by applying a voltage to the actuator with drive electronics. Being resonant, one significant challenge is that the absolute repeatability of the pump output will vary as a result of factors such as gas temperature, pressure and composition, and by part to part variation. This variation means that in certain applications or environments it may not be possible to simply drop such a pump into a microfluidic system and drive accurate and precise liquid flows.

To achieve reliably accurate and precise liquid flow in such situations requires a versatile pump drive scheme in which the applied voltage waveform can be varied to adapt the pump output to provide certain desired flow characteristics. That is, a drive scheme in which a quantity or quality of the drive waveform supplied to the pump can be varied to adjust the output of the pump in order to compensate for the variability. Such quantities might include but are not limited to: the frequency of the waveform; the peak amplitude of the waveform or the ratio and absolute amplitudes of the peaks and troughs where the waveform is more complex than a sinusoidal or square form; the ratio of the width of phases of the waveform, such as peaks and troughs, or the mark-space ratio where a square waveform is used; and the differential of the voltage in time, such as the ramp rate on a trapezoidal waveform. Such qualities might include but are not limited to the shape of the waveform, such as whether it is square, sinusoidal, tristate, trapezoid, sawtooth or of some other form.

Figure 3 shows one example of a drive scheme incorporated in the current invention. The drive scheme comprises a supply voltage 22 provided to drive electronics 14, and specifically to a voltage regulator 23 which powers a microprocessor 24, and separately to a DC-DC voltage converter 25, such as a boost converter or buck-boost converter. The voltage converter is controlled by control signals 26 from the microprocessor that might include a digital control line to enable or shutdown the voltage converter, and some mechanism for controlling the output of the voltage converter, such as an analogue control signal from the microprocessor’s digital-to-analogue converter.

The output of the voltage converter (hereafter, the “drive voltage”) 27, whose level is controlled by the control signals 26, is provided to a current and voltage sense circuit 28 so that the voltage supplied to and the current sunk by an output stage 29 and pump 1 can be measured by the microprocessor by means of analogue signals 30 provided by the current and voltage sense circuit. The drive voltage 27 is ultimately supplied via the current and voltage sense circuitry to an output stage 29, for example an H-bridge. The output stage contains a number of transistors that are controlled by signals provided by the microprocessor 31. The control signals might themselves be controlled by timing algorithms operating within the software on the microprocessor, or perhaps by software configuration of the Pulse Width Modulation (PWM) module within a timer peripheral implemented within the hardware of the microprocessor. The control of the output stage transistors enables the creation of a time-varying voltage, the drive waveform 15, at the output 32 of the drive electronics. In the case that an H-bridge topology is used for the output stage, this is achieved by switching the transistors such that the drive voltage and a ground voltage can be selectively applied to either side of the active element of the pump 1, such as a piezo transducer, although other electroactive technologies would also benefit from the same scheme. The versatile drive scheme disclosed in Figure 3 provides a drive waveform of square nature, where the frequency, amplitude, mark-space ratio and number of voltage levels can be adjusted.

Sensor feedback to drive electronics

Figure 4 shows an example where the drive electronics 14 provide a drive waveform 15 to the pump. A sensing element 16 makes a measurement 17 of a property of the system 18. A mechanism for the drive electronics to interrogate the sensing element is provided, such as an analogue or digital signal, or a digital communication implementation that enables the drive electronics to communicate with the sensing element. The measured property (or quantity) might be related to the input to the pump, such as the power supplied to the pump, or the output of the pump, such as the gas pressure or gas flow generated by the pump.

Alternatively, the measured property might be related to the liquid 5, such as the pressure at some point in the liquid relative to ambient gas pressure, or the differential pressure across all or some section of the liquid, for example between 9 and 10, or the flow rate of the liquid.

The measured property may be processed by the drive electronics, either in hardware via analogue or digital electronics processing techniques, or by processing within the software running on the microprocessor, or both. In either case, the processing results in a transfer function which translates the measured quantity into a set of requirements on one or more quantity or quality of the drive waveform provided to the pump.

In this way, the output of the pump may be adjusted by varying the drive waveform to alter the characteristics of the fluid flow. In particular, the drive waveform may be varied such that the measured property of the system is controlled, for example to approach a predetermined desired value.

For example, it might be desirable to minimise the response time to changes in a set point, or to drive a liquid flow rate as close as possible to a target value, or to drive liquid flow rate in a varying manner according to some desired profile, such as an oscillatory motion profile or a linear or non-linear ramp. It might also be desirable to minimise perturbations in the flow of the liquid, or in the pressure of the gas.

In the example shown in Figure 5, the sensing element is a gas pressure transducer 19 and the first quantity is gas pressure 3 in the headspace of the microfluidic channel 9. For a given liquid 5 with broadly fixed fluidic properties, within a given microfluidic channel 9 a relationship may be established between the gas pressure in the headspace and the resulting liquid flow 6, if the downstream gas pressure 10 is known. In this way, the drive circuit may apply an appropriate voltage waveform to the piezoelectric actuator such that the gas pressure in the headspace approaches a desired value, thereby causing the flow to approach a corresponding desired value.

In the example shown in Figure 6, the sensing element is a gas flow transducer which measures the flow of gas from the pump into the headspace adjacent to the liquid. The gas flow into the headspace will cause the pressure in the headspace to rise, increasing the pressure differential across the liquid and the downstream pressure measurement. When this pressure differential is sufficient, the liquid will begin to flow. As the liquid flows, the volume of the headspace will increase and so the pressure within it will decrease. Assuming that the pump is controlled to deliver a constant gas flow rate via reference to the gas flow transducer, it is conceivable that eventually an equilibrium will be reached whether the gas flow rate equals the liquid flow rate, thereby providing the means to deliver controlled dosing of the liquid. One benefit of making measurements on the gas, whether of the gas pressure as in Figure 5, or the gas flow rate as in Figure 6, is that the measurement apparatus - such as a sensor or transducer - does not come into contact with the liquid and as such there is no concern of cross-contamination where the sensor is not disposable and is reused.

In the example of Figure 7, the sensing element is a liquid pressure transducer measuring the pressure of the liquid 5. The control approach with this example is similar to that explained for the arrangement in Figure 5.

In Figure 8, the sensing element is a liquid flow transducer 33 measuring the rate of flow 6 of the liquid within the microfluidic channel. The control approach with this example is similar to that explained for the arrangement in Figure 6. The benefit of this approach over measuring the gas flow rate as in Figure 6 is in having a direct measurement of liquid flow and therefore not needing to infer liquid flow rate from gas flow rate. The drawback of this approach is that the liquid flow transducer is wetted by the liquid, and therefore could lead to cross contamination between samples if reused. However, the emergence of suitably small, low cost and accurate liquid flow transducers onto the market means that the use of the same is compatible with microfluidic drive system proposed. In this case, the liquid flow transducer would be incorporated into the disposable part of the system, such as a microfluidic cartridge.

These examples are simplifications of the system and more complex arrangements are possible. A more complex practical example may include multiple sensing elements. In Figure 9, the microfluidic drive system comprises three sensing systems: voltage and current sensing 28 of the waveform 15 supplied to the pump 1 such that the power supplied to the pump can also be measured. This provides an indication of the drive level of the pump. Such an indication can then be used to limit the drive power supplied to the pump, which can help to avoid damage to the pump through overdriving it. A gas pressure transducer 19 measures the gas pressure 3 in the headspace of the microfluidic channel. This allows the drive electronics 14 to relate the drive power to the pressure generated by the pump, which can be considered as a measure of pump efficiency.

Knowledge of the pump efficiency allows the estimation of certain parameters, such as remaining runtime for a battery-powered device before it is necessary to recharge the battery, or the identification of a failure of the device. For example, if the pump actuator fails or deteriorates, then it is a regular observation that the pump efficiency falls. With knowledge of the drive efficiency one might also adopt a more insightful approach to controlling the pump. For example, above certain drive powers the pump may self-heat to an extent that it becomes inefficient. Monitoring pump efficiency as a function of drive power therefore provides a means to set an upper limit on drive power. Furthermore, the change in pump efficiency over time can be monitored, for example as the pump wears and generates less pressure for a given power input. From this information it is possible to estimate parameters such as the life expectancy of the pump.

Furthermore if a reduction in the pressure generated for a given power input is observed, this may indicate that there is an undesirable gas leak in the system; this may then be confirmed with the further addition of a gas flow sensor. Finally this example includes a liquid flow rate transducer 33 so that the drive electronics can make adjustments to the drive waveform in order to control the flow rate.

As already explained, in some examples the property of the system measured with the sensor is used to determine a fault condition. For example if the flow rate drops below a certain level or the liquid or gas pressure falls below a predetermined threshold the system may determine the presence of a fault in the system. For example the drive electronics may additionally include a user interface element which indicates the presence of a fault dependent on the quantity sensed by the one or more sensors. The user interface may be a simple component such as a light-emitting element, a sound emitting element or a vibrating element which provides a signal upon activation to alert a user to the fault. Such elements are of low cost such that they may readily be incorporated into such low cost, simple systems. Alternatively the user interface may be a more complex element such as a graphical display which can provide further details of the error identified in the system. A user interface may also more generally indicate data regarding the characteristics of the fluid flow, the system performance, the remaining battery charge, and age of the components to provide information to the user.

Additional system features

Certain examples of the invention incorporate a check valve 34 to hold system pressure at a desired value, as shown in Figure 10A. The pump 1 draws air through the check valve and out of a system volume such as a reservoir 8. When the pump is disabled, or the pressure difference across the check valve is insufficient to open the valve, then the valve may remain closed, thereby holding pressure in the reservoir. The check valve may be a passive valve such as a duck-bill valve or another kind, or an active valve such as one operated by a solenoid. Optionally a vent valve 35 or a fixed orifice can be used to set the rate at which air may bleed into the reservoir so as to equalise the pressure between the reservoir and atmosphere. Again, the vent valve might be a passive valve or orifice, or an adjustable valve such as a variable solenoid valve or manual needle valve.

Figure 10B shows a similar arrangement where the orientation of the pump and check valve with respect to the system is reversed, such that the pump is used to increase the pressure in the reservoir. The benefit of using a check valve in this way is a reduction in the duty cycle, power consumption and wear of the pump. This is because the pump may have a back leak, which if left unchecked would allow pressure to flow back through the pump and into the reservoir; the check valve opposes this back flow.

Figure 11 shows the use of a filter 36 with the pump 1 so as to prevent particles and contaminants above a certain size from entering the pump 1. Such particles can increase the risk of damage to the pump and contribute to a shorter operational life. The filter may have a pore size in the range 0.1 to 20 pm and advantageously substantially smaller than the minimum dimension within the valve structure of the pump. The filter membrane might be made from be PTFE, some other polymer, glass fibre or similar materials.

The filter 36 may be connected to the pump 1 via an intermediary channel such as a length of tube, or integrated directly into the housing of the pump, or integrated with the casework of the enclosing product or system, or situated somewhere else relative to the pump but in every case should at least act on the air being supplied to the pump. In certain applications the pump the output of the pump should also be filtered. For example where there is concern that any material that the pump may shed during operation could present a contamination risk to the rest of the related product or system.

The example of Figure 12a incorporates a reservoir, referred to as an accumulator 37, which can be used to accumulate a pressure difference relative to some other pressure such as atmospheric pressure. Here, the pump 1 removes air from the accumulator. Independently, the vent valve 35 can be operated such that the pressure outside the vent valve A can be controlled. The benefit to this approach is that the change in pressure, dP/dt, at A, can be much faster than the pump could otherwise achieve on its own without use of the accumulator and therefore this approach can reduce system response time. It may also help to reduce perturbations in the system performance, such as the pressure or flow. The vent valve in this case is an active valve and might be controlled with a control signal 38, such that with the pump running continuously, the pressure in the accumulator is as shown 39 and the pressure outside the system at point A is as shown 40. Figure 12b shows a similar arrangement where the pump increases the pressure within the accumulator during operation.

Figure 13A shows two pumps used in series so as to increase the pressure output of the combined pumping system. Figure 13B shows an arrangement with two pumps used in parallel to increase the flow output of the combined pumping system. Extending this approach further, Figure 13C shows a number of pumps used in series and parallel to increase both the pressure and flow of the combined system, relative to a single pump.

Figure 14A shows external manifold apparatus 41 to link the inlets of multiple pumps together. In Figure 14B a similar arrangement is shown where the apparatus 42 links the outlet of multiple pumps together. Pumps may themselves have more than one pumping chamber. The pumping chambers may be stacked upon one another as shown in Figures 14 G to I.

In Figure 14C, a pump with two pumping chambers 43, 44 is shown, each having an inlet and an outlet. Together these two chambers compose a single pump. The advantage of a pump with multiple chambers is to increase the total capacity of the pump, for example to generate higher pressure or high flow than a single pumping chamber is capable of. A separate advantage where the cambers are used in series is to reduce the pressure that each chamber must generate for a given application, which can reduce the pressure difference across the valve in each chamber and in doing so, reduce stress on and extend the life of the valve and therefore the pump. A manifold apparatus 45 is provided such that the flow or pressure generated by each pumping chamber can be linked somehow to the other chamber. This can be achieved with the pumping chambers in series, or with the chambers in parallel. Various configurations of manifolding arrangements are possible. Figure 14D shows a pump with two chambers, where the outlet of one is connected to the inlet of the other via a manifold apparatus, so that the chambers are pneumatically in series. Figure 14E shows a pump with two chambers, where the inlets of the two chambers are connected to one another via a manifold apparatus, so that the chambers are pneumatically in parallel, with a single pump inlet and two pump outlets, which has particularly beneficial properties for vacuum applications. Figure 14F shows a pump with two chambers, where the outlets of the two chambers are connected to one another via a manifold apparatus, so that the chambers are pneumatically in parallel, with a single pump outlet and two pump inlets. Of course, a combination of what is shown in Figure 14E and Figure 14F is possible, where the inlets of the two chambers are connected to one another via a manifold apparatus, and the outlets of the two chambers are connected to one another, either via the same manifold apparatus or a second independent one, so that the chambers are pneumatically in parallel, with a single pump inlet and a single pump outlet.

Figure 14G shows an example of manifolding apparatus where ‘L’ shaped pipe pieces 48, 49, 50, 51 are used. Each piece mates with a counter-part piece; 48 with 49 and 50 with 51. There is a raised bead 52 on the male part 49, 51 that forms an interference fit with the inner wall of the female part 48, 50, 53 so as to provide an airtight seal. Each coupled pair forms a ‘C’ shaped pipe assembly 54 that provides a flow path to link the inlet or outlet of one pumping chamber with the inlet or outlet of another. The ‘C’ shaped assembly is itself a push-fit into the pump body 55. The length of pipe that is inserted with the pump body has two raised beads 56. One is intended to form an airtight seal with the pump body; the other is intended to locate within a companion trough running around the interior circumference of the radial flow channels in the pump body so as to define the desired depth of engagement for the ‘C’ shaped assembly within the pump body. The ‘L’ shaped pipe pieces each have two rectangular wings 57 that serve as end stops to prevent the ‘C’ shaped pipe assembly from being inserted too deeply into the pump body. The male ‘L’ shaped pipe pieces 49, 51 each also have two further rectangular wings 58 that serve as end stops to prevent the male ‘L’ shaped pipe piece 49, 51 from being inserted too deeply into the female ‘L’ shaped pipe piece 48, 50.

One outcome of linking the inlets and outlets of the pumping chamber together is that it becomes possible to create an ‘axial flow’ pump; that is to say, a pump where the air flow moves through the pump in an axial direction, rather than in and out through radial tubes (as discussed later and shown in Figure 14E). For applications where the pump is integrated tightly into an enclosing product or module, having the choice between an axial flow or a radial flow configuration is advantageous in offering greater design freedom; ultimately product engineers are free to select the variant that best meets their requirements, for example minimising product size.

Figure 14G shows one possible embodiment of the axial flow pump. There is a raised boss 88 on the top surface of the pump. This defines a single, largediameter inlet for the pump. The boss is intended to sit against a filter assembly, such as a filter disc of the same diameter, so as to filter the incoming air to remove particulates before the air then passes through the pump. Such a boss arrangement is advantageous in promoting tight integration between the filter and pump, thereby reducing the overall size of the pump and filter assembly. Having flown through the filter, the airflow then splits, with part of the flow entering the upper pumping chamber through the inlet hole 86 and the other part entering the lower pumping chamber via the ‘C’ shaped pipe assembly defined by the ‘L’ shaped pipes 50 and 51. The output from each pumping chamber is combined by a second ‘C’ shaped pipe assembly defined by the ‘L’ shaped pipes 48 and 49. Having re-joined, the output flow of the pump then exits via the single exhaust hole for the pump 87 on the bottom side of the pump. In this way, the resulting airflow 91 moves axially through the pump, in entering at top side and exiting at the bottom.

Figure 14G also shows an exhaust boss 89. The exterior cylindrical face of the exhaust boss in intended to seal against a compliant part of the enclosing product or module, such as an o-ring. In that way, this embodiment can form an air-tight seal at the exhaust of the pump. There are three cuboid protrusions close to the exhaust boss, collectively labelled 90, shown on two aspects in Figure 14D. These protrusions have their longest axis aligned with the cylindrical axis of the exhaust boss, but do not extend as far along this axis as the exhaust boss. In this way, the end face of these protrusions forms a hard end stop for the mating ο-ring and defines the maximum and desired axial depth that the exhaust boss should mate with the ο-ring, and by extension in this case, the maximum depth of engagement of the pump in the enclosing product or module.

Figure 14H shows two sets of alignment features, one helpful in assemble the pump itself, and one helpful in assembling the pump into an enclosing product or module. During assembly of the pump, the separate parts of the upper pump body 96 and lower pump body 97 must be brought together axially to form the pump. The rotational alignment of these parts to one another is important. A partly-circular protrusion, the “lower alignment feature” 93, extends radially beyond the main diameter of the lower pump body 97. This protrusion is intended to engage with two companion features, the “upper alignment features” collectively shown as 92, on the upper pump body 96. The upper alignment features each have a frustoconical form. The frustoconical form provides a lead in as the upper and lower pump bodies come together axially. The lead in allows greater rotational tolerance initially as the lower alignment feature 93 and the upper alignment features 92 first come together. This is as depicted in 98, where the lower alignment feature 93 is shown not centrally located with respect to the two upper alignment features 92. As the parts engage more deeply as depicted in 99, the lead in of the upper alignment features 92 causes the lower alignment feature 93 to become better rotationally aligned, and hence the upper and lower pump bodies to be better aligned to one another.

Also shown in Figure 14H are two other partly-circular protrusions 94, 95, from the lower pump body 97, each with a hole in the centre. The hole in each is intended to sit over a companion pin in the enclosing module or product housing, so as to provide rotational alignment between the pump and the module or product. For completeness, the X-Y alignment of the pump to the module or product housing is taken care of by the part of the product housing that receives the exhaust boss 89 of the pump as shown in Figure 14D; the Z alignment of the pump to the module or product housing is taken care of by the engagement of the three cuboid protrusions close to the exhaust boss, collectively labelled 90, shown on two aspects in Figure 14D, and an o-ring as previously disclosed.

Figure 141 shows another actual approach to addressing the task of manifolding two pumping chambers together. In this case, the inlet 59, 60 and outlet 61, 62 of each pumping chamber are brought out to a pair of tubes 63, 64 with a raised bead running around the circumference of each tube 65. In total there are four tubes for a two chamber pump. These tubes are compatible with flexible tubing or alternatively allow the pump to be plugged into a manifold block of some kind. The raised bead 65 is intended to promote an airtight sealing line against tubing or a manifold block.

Figure 15 shows an electrical interface 46 that is compatible with a sprung-pin connector 47 or similar component so as to improve the speed or ease or both of which the pump can be physically installed or removed from the location of operation such as a containing product or system.

Figure 16 shows one possible embodiment of the microfluidic drive system of the present invention. The drive electronics 14 includes a microprocessor 24, which itself includes at least Pulse Width Modulation (PWM) 66, Digital to Analog Converter (DAC) 67 and Analog to Digital Converter (ADC) 68 modules. The PWM provides drive signals 71 to an H-bridge output stage 72. The H-bridge output stage is supplied a drive voltage 73 by the high voltage power supply 74. The high voltage power supply is for example a boost or buck-boost DC to DC voltage level converter. The microprocessor controls the high voltage (also referred to as drive voltage) via an analog control signal 75 produced by the DAC and controlled in software by the microprocessor 67. The drive voltage is supplied to the H-bridge via the current sense 75 in such a way that the current sunk by the H-bridge and the pump 1 can be measured. The current sense might include for example a shunt resistor and transimpedance amplifier such that the measured current can be converted into a proportional voltage signal. The drive voltage is measured by the voltage sense 76. The voltage sense might comprise a potential divider to provide a lower voltage proportional to the drive voltage. Both the voltage and current sense circuits might include a voltage clamping circuit to limit the maximum amplitude supplied to the ADC in order to prevent damage. Both the voltage sense and current sense signals 77, 78 are provided to the microprocessor via the ADC 70 and can be processed in the software.

The pump 1 is driven by a waveform supplied by the output stage 79. The pump inlet has a filter 80 to prevent particulates from entering the pump and causing blockage, wear or damage. The pneumatic output pressure of the pump is connected to a liquid reservoir 81 so as to drive liquid flow from the reservoir outlet in conceptually the same manner as outlined in Figure 2A, A bleed valve 82 connected to the pneumatic path via a T junction provides variable control of the pressure discharge from the pneumatic path to atmosphere. The bleed valve can be set to discharge pressure quickly, which has the advantage of allowing the liquid flow to be stopped more quickly when the pump is disabled as desired, and the disadvantage that the pump 1 has to be driven harder to maintain a certain pressure within the pneumatic path. Alternatively, the bleed valve can be set to discharge pressure slowly, which has the advantage that the pump can be driven more softly to maintain the same pressure, which may extend the operational life of the pump, at the expense of allowing the liquid flow to be stopped more quickly when the pump is disabled. An active valve such as a solenoid valve could be used in place of the bleed valve shown such that the benefits of stopping liquid flow quickly and yet driving the pump softly can be realised.

The pressure in the reservoir is measured by a pressure sensor 83 which provides a signal to the microprocessor via the ADC. Alternatively, a pressure sensor of a digital communications type might be connected to a digital communication interface 85 of the microprocessor. Measurement of the pressure in this way enables the microprocessor to control the output of the pump such that the desired pressure is generated. It may also serve as a fault detection approach, wherein if lower than the expected pressure is generated by the pump, the system may identify that there is a fault with the pump and take appropriate action, such as to shut off the pump and notify the system operator.

A liquid flow sensor 82 measures the flow rate of liquid from the reservoir and is connected to the microprocessor via a digital communications interface. Measurement of the liquid flow enables the microprocessor to control the output of the pump such that the desired liquid flow rate is generated. In this embodiment, this liquid flow is then provided to a microfluidic droplet generator chip. The liquid in this case is water. A second drive system (not shown) comprising an independent drive electronics, pump, filter, bleed valve, reservoir and flow rate sensor, provides a second liquid flow, in this case oil, to the droplet generator chip. Each drive system by this implementation allows for the accurate control of pulsation-free liquid flow rate of its respective liquid into the droplet generator. The result is the stable and homogeneous generation of oil and water emulsion, which is of interest in a wide range of microfluidic applications including digital PCR and cell encapsulation.

Applications for the microfluidic drive system

The microfluidic drive system of the present invention provides a reliable smooth liquid flow at low cost and complexity and as such is well suited to a wide range of applications.

Applications in which the invention may be employed include, for example: anaesthesia; apparel control including but not limited to inflation, deflation and thermal control; balloon inflation; blood filtration; blood gas analysis; blood pressure monitoring including but not limited to continuous and oscillatory methods; breast pumping; breath sampling; capnography; vehicle seat conditioning including but not limited to heating, cooling and pressure adjustment by means of the inflation and deflation of bladders within or adjacent to the seat cushion; compression therapies including but not limited to continuous and intermittent methods for the treatment of deep vein thrombosis, oedema and other circulatory conditions; domestic appliances including but not limited to vacuum and positive pressure systems; erectile dysfunction; fluidic systems including but not limited to pressure-driven liquid flows; food and beverage packaging and preservation applications; fragrance dispensing; gas detection or sensing including but not limited to diagnostics and air quality measurement; gas sampling; general healthcare monitoring; infusion; medical catheters; medical diagnostic devices including but not limited to hand-held devices or otherwise, including point-of-care devices; medical suction and smoke extraction during surgery; micro fuel cells; military applications including but not limited to therapy delivery and chemical detection and monitoring; nebulizer systems; oxygen concentration; printhead pressure control; seat control including but not limited to mechanical and thermal control and adjustment; security including but not limited to improvised explosive device detection and border security; sleep therapies including but not limited to continuous positive airway pressure; sports applications including but not limited to compression, monitoring and apparel; therapeutic surfaces including but not limited to seat and bed cushion adjustment; ventilation systems including but not limited to positive end expiratory pressure valves; wound care including but not limited to negative pressure wound therapy.

Claims

1. A microfluidic drive system comprising:

a resonant gas pump comprising a substantially cylindrical cavity defined by cavity walls, the cavity having an inlet and an outlet aperture and a piezoelectric actuator arranged to generate oscillatory motion of the cavity walls to drive a gas between the inlet and outlet;

a drive circuit arranged to apply a voltage waveform across the piezoelectric actuator, such that the oscillations of the cavity have a frequency of at least 500 Hz;

a microfluidic channel arranged in fluid communication with the inlet or outlet of the pump, such that, in use, the varying gas pressure provides a driving force to move a liquid through the microfluidic channel.

2. The microfluidic drive system of claim 1 further comprising:

a liquid reservoir arranged between the pump and the microfluidic channel such that the pump is in fluid communication with the microfluidic channel via the liquid reservoir and, in use, the varying gas pressure is applied to a liquid in the liquid reservoir to drive the movement of liquid through the microfluidic channel.

3. The microfluidic drive system of claim 1 or claim 2 further comprising:

at least one sensor arranged to measure a property of the system, the sensor connected to the drive circuit such that, in use, the drive circuit adjusts the applied voltage waveform in response to a measured property received from the sensor.

4. The microfluidic drive system of claim 3 wherein the at least one sensor comprises one or more of:

a sensor arranged to measure the power supplied to the pump;

a pressure transducer arranged to measure the gas pressure generated by the pump;

a gas flow transducer arranged to measure the gas flow rate generated by the pump;

a pressure transducer arranged to measure the pressure within the liquid at a point in the system relative to ambient pressure;

a pressure transducer arranged to measure the differential pressure across all or a portion of the liquid in the system;

a liquid flow transducer arranged to measure the flow rate of the liquid in the system.

5. The microfluidic drive system of claim 4 wherein at least one sensor is a sensor arranged to measure the power supplied to the pump, the sensor comprising a current sense subsystem.

6. The microfluidic device of claim 5 wherein the current sense subsystem comprises a resistor arranged to measure the current supplied to an electronic stage arranged to drive the pump.

7. The microfluidic drive system of claim 3 to claim 5 wherein the drive system is configured to adjust the applied voltage waveform such that the measured property converges to a predetermined set point value.

8. The microfluidic drive system of any of claims 3 to 7 wherein the drive circuit is configured to compare the measured quantity to a predetermined threshold value to determine a fault condition.

9. The microfluidic drive system of claim 8 further comprising a user interface element configured to indicate the fault condition.

10. The microfluidic drive system of claim 9 wherein the user interface element comprises one or more of a light-emitting element, a sound-emitting element and a vibrating element

11. The microfluidic drive system of any preceding claim wherein the cavity radius, a, and height, h, of the pump satisfy the following inequalities:

— is greater than 1.2; and — is greater than 4 x 10'10m.

a

12. The microfluidic drive system of any preceding claim wherein at least one of the apertures of the pump is a valved aperture.

13. The microfluidic drive system of any preceding claim wherein the volume of air displaced by the pump per cycle is less than 1 mL and preferably between about 1nL to 10OpL.

14. The microfluidic drive system of any preceding claim wherein the drive circuit comprises an electronic output stage which drives the pump.

15. The microfluidic drive system of claim 14 where the output stage is an Elbridge.

16. The microfluidic drive system of any preceding claim further comprising a valve or fixed orifice which provides a junction between the flow path between the pump and the microfluidic channel or the flow path between the liquid reservoir and atmospheric pressure such that the differential pressure between the flow path and atmosphere provides a corresponding air flow through the valve or orifice.

17. The microfluidic drive system of any preceding claim further comprising an accumulator reservoir arranged to accumulate a pressure difference relative to atmospheric pressure.

18. The microfluidic drive system further comprising at least one further pump arranged in series and/or parallel with the first pump.

19. The microfluidic drive system according to any preceding claim wherein the pump cavity comprises two pumping chambers, each chamber having an inlet and outlet aperture.

20. The microfluidic drive system according to claim 19 wherein the chambers are arranged in series or parallel.

21. The microfluidic drive system according to claim 18 to 20 further comprising a manifold apparatus arranged to connect the inlet and outlet apertures of the pumps or pumping chambers.

22. The microfluidic drive system according to claim 21 wherein the outlet of a chamber is connected to the inlet of another chamber such that the chambers are pneumatically in series.

23. The microfluidic drive system according to claim 21 or claim 22 wherein the inlets of two chambers or the outlets of the two chambers are connected to one another such that the chambers are pneumatically in parallel.

24. The microfluidic drive system according to any of claims 21 to 23 wherein the pumping chambers are stacked and the manifold apparatus comprises a pipe or pipe assembly to connect stacked pumping chambers.