WO2006030405A1

WO2006030405A1 - A transducer apparatus for measuring biomedical pressures

Info

Publication number: WO2006030405A1
Application number: PCT/IE2005/000100
Authority: WO
Inventors: Vincent Casey
Original assignee: University Of Limerick
Priority date: 2004-09-14
Filing date: 2005-09-14
Publication date: 2006-03-23

Abstract

A transducer apparatus has a transducer (1) which lies between a bandage (8) and a patient's leg (9), for example. An active part (2) has a sensor element (15) comprising an array of sub-elements (46) of trapezoidal shape in cross-section. The sensor element (15) is a dielectric between electrodes (16, 18) on substrates (6, 17). The assembly is enveloped by encapsulant polymer films 5(a) and 5(b). The arrangement of the sensor element (15) allows a linear response and a very thin non-intrusive shape which allows reliable calibration.

Description

"A transducer apparatus for measuring biomedical pressures"

INTRODUCTION

This invention pertains to transducers for estimating the pressure applied to body-tissue by an object such as a medical device, bandage or dressing.

The pressure developed, P, beneath a membrane is governed by the tension, T, in the membrane and the curvature, K, of the membrane according to the law of Laplace, P=TK. In the case of bandages and wound dressings applied to cylindrical bodies, i.e. a compression bandage on a limb, the form P=NTZr is frequently used by clinicians to estimate the bandage applied pressure, where N is the number of complete wraps, r is the radius of curvature of the limb and the wrapping tension T is assumed to be constant for each wrap. However, real limbs do not have regular geometric shapes and so the pressure will vary locally with curvature. In addition the actual bandage tension is not measured or known. However, some bandages have visual aids to help the clinician establish a constant extension in the bandage as it is applied, typically 50% extension. However, bandage extension only provides a crude estimate of the actual tension in the bandage and so pressures calculated using the law of Laplace cannot be expected to reflect the actual sub-bandage pressure at a given location on a limb or support tissue with a great degree of accuracy.

In compression therapy and many other medical dressing/bandage applications it is desirable to have a transducer which establishes the actual pressure applied by a device such as a bandage to the underlying tissue, in a continuous way, during the time while the bandage is being applied but also over prolonged periods of time while the patient is undergoing treatment. In view of the importance of the calf muscle pump on venous return, information on the actual sub-bandage pressure during movement would be particularly useful to the clinician.

While a number of biomedical pressure sensor devices have been developed, significant errors arise with any such intrusive sensor, no matter how low its profile is, due to the so called 'hammocking-effecf whereby a membrane such as bandage, cuff liner or indeed skin, stretches and curves in order to adjust to the shape of the sensor. In the case of a bandage on a limb, the bandage lifts away from the supporting tissue in the vicinity of the sensor forming a profile somewhat like that of a hammock. The sensor area is a critical parameter in the calibration of pressure sensors. It is normal to assume a constant area based upon the geometric area of the sensor. However, the 'hammocking- effect' changes the effective area of the sensor which in turn varies with applied pressure/tension in a complex manner. Consequently, it is difficult to calibrate such sensors reliably for the wide range of tissue properties, curvatures and bandage properties encountered in actual biomedical application environments.

Force and pressure are often sensed using the 'spring balance principle', where, in a primary transduction step, an elastic element transforms the measurand into a deflection or deformation, which, in a secondary transduction step, is converted into an electrically useful signal. Common pressure transducers tend to depend on the deformation of a spring, cantilever or diaphragm in the primary transduction step while electrical and optical techniques are used frequently in the secondary transduction step.

The compliance of polymer materials and rubbers, in addition to their elasticity, makes them attractive candidate materials for the primary transduction step. Weighing mats by Miller et al. in U.S. Pat. No. 3875481 and load cell type transducers have been developed using elastomer elements where change in capacitance is used to infer the applied pressure. Improvements in relation to hysteresis and linearity of response have resulted from the use of voids and structured elastomer elements by Haberl et al. in U.S. Pat. No. 4266263 and by Seimiya et al. in U.S. Pat. No. 5693886. However, it appears that such devices would have high profiles and would be complex to manufacture in volume.

US4584625 (Kellogg) describes a capacitive tactile sensor having gas-filled compression cells.

US6526043 (Boie et al) describes a tactile transducer having a planar deformable layer and multiple conductive elements on each side. This arrangement appears to be complex. GB2199953 (Medimatch) describes a biomedical pressure transducer having fluid-filled cells.

US4869265, describes a biomedical pressure transducer having a pressurisable chamber containing integral membrane switch type electrical contacts interposed between the tissue and an apparatus such as a tourniquet cuff. The normally closed contacts are opened when the pressure within the chamber is equal to the pressure applied by the cuff. This approach does not appear to provide a continuous dynamic estimate of the pressure applied and suffers from signal damping introduced by the fluid line. In addition this device appears to be complex.

In a planar transducer disclosed in US6636760 a structured elastomer element is used in a fiber optic based semi-rigid sensor suitable for use in the measurement of pressures under surgical tourniquet cuffs. While this approach is effective, there is still a need for less intrusiveness and simpler manufacturing.

The invention is directed towards providing an improved transducer for measuring biomedical pressures.

SUMMARY OF THE INVENTION

According to the invention there is provided a pressure transducer apparatus for measuring biomedical pressures, the transducer comprising:

a primary transduction part having a deformable sensor element, and

a secondary transduction part for measuring deformation of the sensor element, to generate an indication of pressure on the element, wherein:

the sensor element comprises a plurality of co-planar deformable sub- elements, and - A -

the secondary transduction part comprises an electrode on each side of the sensor element so that the sensor element is a dielectric, and a circuit for monitoring change in capacitance between the electrodes.

In one embodiment, the sub-elements are arranged in an array.

In one embodiment, the sub-elements have a trapezoidal shape in cross-section in at least one dimension.

In one embodiment, the span to base ratio of the sub-elements is in the range of 0.1 to 1.

In one embodiment, the ratio of the sub-element base width to height is in the range of 1 to 5.

In another embodiment, the sensor element material is silicone.

In one embodiment, the apparatus further comprises an encapsulant comprises opposed encapsulant layers, and preferably the encapsulant is flexible.

In one embodiment, the apparatus further comprises a substrate for each electrode, each substrate having a thickness of less than 150 microns, wherein the sensor element has a thickness of less than 30 microns, and wherein the thickness of each encapsulant layer is less than 150 microns.

In one embodiment, the overall thickness of the apparatus in a normal plane through the sensor element is less than 750 microns.

In a further embodiment, the sub-elements each have a trapezoidal cross section in at least one dimension and comprise a pair of parallel opposing surfaces in contact with the electrodes, and each sub-element has two or four free surfaces inclined at an angle of between 0 and 60 degrees to the normal to the plane of the electrodes. In one embodiment, the apparatus further comprises a screening electrode adhesively joined to the encapsulant at an external surface, and which is itself enclosed in an encapsulation layer.

In one embodiment, the apparatus further comprises an adhesive member for attaching to medical dressings, devices or to living tissue.

In one embodiment, the apparatus further comprises means for calibrating in terms of known^' values of pressure, wherein the calibrating means comprises inflatable bladders which contact the apparatus above and below the plane of the sensor element in order to impose predetermined pressure to the sensor element.

In another aspect, the invention provides a method of producing any pressure transducer apparatus as defined above, the method comprising the steps of:

providing a first electrode;

depositing and bonding the sensor sub-elements onto the first electrode; and

providing a second electrode in contact with the sub-elements on the opposed side so that the sub-elements are a dielectric.

In one embodiment, the sub-elements are moulded onto the first electrode.

In another embodiment, the sub-elements are moulded in a mould which is formed using a lithography stamp, in turn moulded from a micro-machined silicon master.

In one embodiment, an adhesion promotion agent is applied to the first electrode before depositing the sub-elements.

In one embodiment, the electrodes and electrical leads are on substrates. In one embodiment, the method comprises the further step of enveloping the sensor sub- elements and the electrodes in opposed encapsulant layers.

In one embodiment, the encapsulant layers comprise polymer film.

In one embodiment, the encapsulant layers are coated with adhesive on their inner surfaces.

In one embodiment, part of an electrode protrudes from the encapsulation, and an edge connector is connected to the protruding part of the electrode.

In another embodiment, the encapsulant layers are applied under controlled tension.

In one embodiment, the tension is controlled to produce a pressure acting normal to the plane of the sensor element.

In one embodiment, the tension is controlled to adjust the zero offset capacitance of the transducer apparatus

DETAILED DESCRIPTION OF THE INVENTION

Brief Description of the Drawings

The invention will be more clearly understood from the following description of some embodiments thereof, given by way of example only with reference to the accompanying drawings in which: -

Fig. 1 is a schematic diagram depicting a flexible pressure transducer disposed between a bandage and a human limb as well as electronic processing circuitry;

Fig. 2 is a cross-sectional view in the direction of the arrows II - II of Fig. 1 showing details of the sensor construction of the invention; Figs. 3 and 4 respectively illustrate a plan view and an exploded perspective view of the flexible pressure transducer;

Fig. 5 is a plan view and Fig. 6 a cross-sectional view in the direction of the arrows VI - VI, of a deformable layer of the transducer;

Fig. 7 is a plan view of an alternative deformable layer;

Fig. 8 is a block diagram showing the electronic processing circuitry in more detail;

Fig. 9 is a detailed circuit diagram illustrating an arrangement for connecting the transducer to the circuitry;

Fig. 10 is a transducer calibration curve;

Fig. 11 is an output display illustrating the course of the sub-bandage pressure variation as a bandage is applied to a human limb;

Fig. 12 is an output display illustrating the course of the sub-bandage pressure variation due to limb movement;

Figs. 13 and 14 respectively illustrate a plan view and an exploded perspective view of an alternative embodiment of flexible pressure transducer;

Figs. 15 and 16 respectively illustrate a plan view and an exploded perspective view of a further embodiment of flexible pressure transducer of the invention; and

Fig. 17 is a schematic diagram depicting a flexible pressure transducer array disposed between a bandage and a human limb as well as electronic processing circuitry. Description of the Embodiments

Referring to Figs. 1 and 2, a transducer 1 having an active part 2 is disposed between a limb 9 and a compression bandage 8. The transducer 1 has a plurality of layers of materials including a deformable polymer sensor element 15, and the layers are stacked in a thickness direction. The transducer 1 is positioned at a desired location on the limb 9 using an adhesive retainer 7 such as a medical adhesive tape or adhesive pads.

In overview, the primary transduction is compression of the sensor element 15, and the secondary transduction is monitoring change in capacitance between electrodes (16, 18) on either side of the sensor element 15, which is a dielectric. Each of the layers of the transducer 1 is flexible and bendable out of a regular plane of the flexible transducer about both a lengthwise axis and a width- wise axis. This is achieved because of the configuration of the sensor element and the thin layers surrounding the sensor element. For example electrode layers are of Cu but are nevertheless flexible because of the very small thickness. This flexibility allows excellent accuracy and non-intrusiveness for a patient.

A transducer apparatus comprises the transducer 1 and processing electronics including a Universal Transducer Interface (UTI) module 11 connected to the transducer 1 via a screened cable 10, for monitoring changes in capacitance between the electrodes to determine pressure. The screened cable 10 is terminated with standard electronic connectors 3 and 41. The UTI module 11 in turn is connected via a serial port cable 12 to a personal computer 13. The active part 2 is electrically connected to the terminal end 43 using planar conductor tracks 4. The transducer 1 and these electronic items together form a complete transducer apparatus.

Referring to Figs. 2, 3, and 4 polyimide (such as Kapton, ™ DuPont) tape 125μm thick is used as a base substrate 6 and a top substrate 17. A top electrode 18, and a base electrode 16 as well as the conductor tracks 4 comprising a top electrode interconnect track 38, a base electrode interconnect track 39, a top electrode flying track 19 and a screen track 37 are formed from 35μm thick copper foil which is adhesively joined to the polyimide substrates 6 and 17. An insulating polyester film is used as an encapsulation 5 (a) and 5(b) to enclose the entire assembly and seal it from the environment. The base electrode 16 and the top electrode 18 have a circular geometry. The diameter of the base electrode 16 is 8mm. The top electrode 18 has a diameter lmm smaller than that of the base electrode 16. This mismatch in the diameter of the electrodes reduces the criticality of alignment of the electrodes in order to obtain sensors with closely matched characteristics. The diameter of the electrodes is not restricted to the value used here but may be chosen to suit the particular sensor application.

Referring to Figs. 5 and 6 the deforniable polymer sensor element 15 comprises silicone rubber (polydimethylsiloxone, PDMS). The sensor element 15 may be formed directly onto the copper base electrode 16 using a variety of techniques such as photolithographic printing, moulding and soft lithography. A rigid polymer mould formed using a soft lithography stamp that in turn was moulded from a micro-machined silicon master is used to pattern the sensor element 15 onto either the top 18 or base copper electrode 16. Strong adhesion between the sensor element 15 and the copper electrode is achieved by treating the copper surface with an adhesion promotion agent such as DC1200 (™ Dow Corning).

In this embodiment the thickness dimensions are:

• substrate layers (6, 17) 125 microns,

• sensor element (15) 25 microns,

• encapsulant layers (5(a) and 5(b)), each 10-150 microns, and

• electrode copper plus adhesive (16, 18), each 45 microns.

The total thickness can be from less than 500 microns up to a maximum of lmm but is typically 700-800 microns (0.7-0.8mm), and the overall sensor width can be 10-20mm. The reproducibility of the sensor element 15 thickness allows effective use of a capacitive transduction measurement technique.

As shown most clearly in Figs. 5 and 6 the sensor element 15 comprises multiple sub- elements 46. The pattern of the sub-elements 46 is chosen so as to achieve the optimum deformation characteristics for the application. The shape in plan for the sub-elements 46 can be chosen from a range including square, (this embodiment), striped (cfr. strips

47 in Fig. 7), rectangular, circular, or crossed for example. Important sensor element parameters to control relate to the cross-sections of the sub-elements, the spacing between neighbouring sub-elements, and the height of the sub-elements.

During manufacture, the shape of the sub-elements 46 is determined by the shape of the master mould used. Micro-machined silicon moulds which have been formed through anisotropic etching of silicon produce pits and trenches with side walls tapered at an angle of 54.7° to the normal to the plane of the wafer. Etching of square planar patterns will produce pits of trapezoidal cross section in silicon. Etching of lines will produce trenches of trapezoidal cross section. The etch depth will control the eventual height of the moulded sub-elements 46. However, the etch depth also determines the ratio of the contact surface area to that of the non-contact surface area for each sub-element.

Referring particularly to Fig. 6, the optimum span to base width ratio, (G/F), for sub- element 46 cross-sections is within a range from 0.1 to a maximum 1. At ratios greater than 1, excessive deformations will be encountered in the structure which will cause increased non-linearity in the relationship between the compressive stress in the body and the deformation of the body, and, a reduced working range for the device. A G/F ratio of less than 0.1 on the other hand, may cause contact between adjacent sub- elements 46 under load conditions. The ratio of the base width, F, to the height, D, of the individual sub-elements 46, F/D, can vary from a value of 1 to 5. A value of less than 1 would lead to a structure that would have relatively low resistance to shear loading. For values of F/D greater than 5 the sensitivity of the transducer would be so low as to make the device impractical as a simple pressure transducer. In addition, this would lead to increased hysteresis in the relationship between applied pressure and the deformation.

An important index in determining the level of hysteresis and non-linearity in structured elastomer layers is the shape factor, S, which for a given deformable sub-element 46 is defined as the ratio of one loaded surface area, i.e. top surface, to the load-free surface area. For stripe shaped sub-elements (Fig. 7) the shape factor is given by S = W/2R where R is given by R = D/cos(A) and W is the side width of the stripe. For a pyramid shaped element (Figs. 5 and 6) with top flat surface side length W the shape factor is given by S = W² /(4R(W + \(F - W))) . Shape factors somewhere in the range 0.2 to

0.7 but preferably towards the lower end of the scale are recommended in order to optimise elastic properties.

The deformable sensor element of the embodiment of Fig. 7 has sub-elements in the form of strips 47 with a similar cross-section along the lines VI - VI as the sensor element 15 of Figs. 5 and 6

The invention may be implemented using other suitable regular geometric shapes for the electrodes such as square or rectangle. The dimensional extent of the sensor element 15 is chosen such that it extends at least 0.5mm beyond the perimeter of the base electrode 16 in order to electrically isolate the base electrode 16 from the top electrode 18 and also to ensure electrical isolation between a given electrode and the conductor track of the opposing electrode .

In addition to the base electrode 16 the base substrate 6 also supports a series of conductor tracks, 4. The screen track 37 is used to isolate the electrode conductor tracks 38, 39 from each other and is positioned centrally on the base substrate 6. The base electrode conductor track 39 electrically connects the base electrode 16 to the terminal end 43 of the transducer. It is positioned to one side of the screen track 37. A top electrode conductor track 38 is positioned on the other side of the screen track 37 and is of similar dimensional extent to it. As already described, the sensor element 15 is affixed to the base electrode 16 and extends over the edge of the base electrode 16.

The top substrate 17 supports the top electrode 18 and the top electrode flying track 19. The top electrode flying track 19 is connected to the top electrode 18 at one end and extends over the end of the top substrate 17 at the other end by about 5mm. The top substrate 17 has a minimum width equal to the diameter of the top electrode 18 plus lmm and may have a maximum width equal to the width of the base substrate 6. Its length is chosen such that it accommodates the top electrode 18 and provides for an overlap of the top electrode flying track 19 with the top electrode interconnect track 38 of at least 2mm. The top electrode 18 and the top electrode flying track 19 are positioned on the top substrate 17 such that when the top substrate 17 is fixed in place the top electrode 18 faces the sensor element 15 with its centre directly over the centre of the base electrode 16 and the top electrode flying track 19 is directly above the top electrode interconnect track 38 of the base substrate 6. Electrical connectivity between the top electrode interconnect track 38 and the top electrode 18 is completed by soldering the top electrode flying track 19 to the top electrode interconnect track 38. The widths and spacing of the conductor tracks is chosen to correspond with one of the standard PCB values so as to facilitate the use of standard cable interconnectors and PCB connectors. In the embodiment described here, the track width and spacing is lmm.

The encapsulation polymer film 5 is used to package the transducer assembly. The encapsulation polymer film is coated on one side with thermally activated adhesive. Two layers of the film are used to achieve optimum encapsulation, a base encapsulation layer 5(a) and a top encapsulation layer 5(b).

The base substrate 6 is placed onto the adhesive coated side of the base encapsulation 5(a) and is positioned so that the terminal end 43 extends a distance of at least lcm beyond the end of the base encapsulation layer 5(a). This is sufficient to allow the use of pcb type edge connectors or other standard electrical connectors to be used to connect the transducer via standard cable assemblies to the UTI module 11. The base encapsulation layer 5 (a) extends a distance of at least 2mm beyond the remaining sides and end of the transducer. The top encapsulation layer 5(b) is then placed onto the base encapsulation layer 5 (a), adhesive coated side facing the adhesive coated side of the base encapsulation layer with all edges corresponding to the edges of the base encapsulation layer.

Encapsulation of the transducer and zero adjustment of the sensor zero pressure capacitance are achieved in the one process as follows. The lateral edges of the encapsulation layers are clamped between metal straight edge clamps of a tensioning device. The tension of the device is adjusted so as to produce a pressure acting normal to the plane of the transducer which presses the two substrates into intimate contact. By measuring the sensor capacitance during this process the zero offset capacitance of the sensor may be adjusted to a value corresponding to the lower limit of the sensor working range. For 8mm diameter electrodes and a deformable polymer layer thickness of 25μm a typical value would be of the order 15-25pF. The encapsulation film is then heated to 110°C to activate the adhesive. The assembly is allowed cool to room temperature before being removed from the tensioning device. Alternatively, a very thin insulating polymer film such as Teflon tape 42 (Fig. 14) may be wrapped under tension around the sensor end of the transducer in order to bring the sensor electrodes into intimate contact with the sensor element and adjust the zero offset capacitance of the sensor to the lower limit of its working range. In this case encapsulation may be achieved by simple lamination with polyester laminate film or by moulding encapsulant around the transducer.

In use, the transducer 1 may be placed between a compression bandage 8 and a limb 9 or tissue (Fig. 1). The transducer 1 is sufficiently thin that it does not displace substantially the tissue from its normal location in relation to the compression bandage 8. The pressure applied by the bandage 8 to the tissue may be determined using the transducer 1 in conjunction with the electronics circuitry. When pressure is applied to the transducer 1, the deformation of the sensor element 15 will result in a change in the degree of separation of the sensor top electrode 18 and the sensor base electrode 16. This will change the capacitance of the sensor 2 which is measurable using the electronic measurement system, illustrated schematically in Fig. 6 and Fig. 7.

With reference to Figs. 8 and 9, the transducer 1 may be connected directly to the universal transducer interface module 11 or the cable assembly 10 may be used for this purpose. The cable assembly 10 comprises at least one coaxially screened cable 48 and a second cable which may or may not be screened 49. One electrode is connected via the screened cable 48 to the A pin 26 of the UTI chip 21. The other electrode is connected to pin D 27 of the UTI chip 21 using the second cable 49. A reference capacitor 24 is also connected via coaxially screened cable 50 to the A pin 26 of the UTI chip 21. The other lead of reference capacitor 24 is connected to the C pin 28 of the UTI chip 21. Pin F 30 of the UTI chip 21 is connected to ground. The supply voltage is connected to the UTI chip 21 via pins Sel2 34 and PD 35 directly and via the potential divider resistors 32 (25kΩ) and 33 (lkΩ) to pin E 31. The UTI chip 21 is designed for use in smart microcontroller-based systems. One output-data wire reduces the number of interconnect lines and reduces the number of opto-couplers required in isolated systems. Continuous auto-calibration of offset and gain of the complete system is performed by using a three-signal technique. Low frequency interference is removed by using an advanced chopping technique. The function selection can be configured in both software and hardware. The mode of the UTI chip 21 can be used to measure a relatively large value of capacitance up to 30OpF. The output signal of the UTI 21 is period modulated. A microcontroller 23 is used to measure the period modulated signals from the UTI chip 21, to process the measured data and to output digital data to a personal computer 13 via a serial RS232 communication interface 22. The RS232 interfacing chip 22 offers serial communication between the microcontroller 23 and, for instance a personal computer 13. A 12 MHz crystal is used for the on-chip oscillator of the microcontroller 23. The power supply 25 of 5 V is used for the UTI 21, microcontroller chip 21 and the RS232 interface 22. The counting function of the microcontroller 23 is used to measure the periods of the output signal from the UTI chip 21. The periods T_AB, T_AC and T_AD, are quantized in numbers, NA_B, N_AC and NAD, respectively by the microcontroller 23 and these numbers in turn relate to the offset capacitance C₀ the reference capacitor C_ref 28 and the sensor 2 capacitor Cs . When no capacitor is connected between pin A 26 and

pin B 29, the sensor capacitance may be found using C₅ = — — —C _f and if C_ref is

^ AC ~ N AB accurately known, C_x may be accurately measured. When pressure is applied to the sensor, the capacitance changes. The quantized number NAD is proportional to this

capacitance. Consequently, the ratio — — — is a direct indicator of the sensor

N AC - N AB capacitance which in turn is representative of the pressure applied. This ratio will hereafter be referred to as the sensor output represented by Xs.

Known pressures may be used to calibrate the sensor output, Xs, in any desired pressure units. The method of transducer encapsulation used here facilitates absolute pressure calibration using a simple air compression chamber since the encapsulation provides an air-tight seal for the transducer. The transducer is placed in the compression chamber and connected to an air-tight lead-through which facilitates connection of the sensor to the control electronics. The air pressure within the chamber is adjusted in a controlled way to provide known pressures thereby allowing transducer calibration.

Alternatively, application of such known pressures is conveniently achieved using a pair of soft rubber bladders confined within a rigid box which defines a fixed volume. The sensor is placed between the bladders and an air pump is used to inflate the bladders which are connected together using a pressure equalisation tube. A mercury sphygometer, pressure dial gauge or solid state pressure sensor or other such accurate pressure measurement device is connected to the bladder pump line. The bladders press against the flexible pressure transducer as they are inflated thereby applying a pressure to the transducer which corresponds to the hydrostatic pressure within the bladders. Pressure may be increased or reduced by inflating or deflating the bladders respectively. In a clinical setting a particularly simple calibration is possible by applying two reference weights to the sensor corresponding to two datum points within the desired pressure measurement range, e.g. 20mmHg and 60mmHg for sub-bandage pressure measurement. The weights are shaped such that they have a bottom smooth surface which has a geometry which corresponds exactly to the planar geometry of the sensor. Reliable and reproducible applied pressures are obtained once a soft elastomer layer 2mm thick is affixed to the base of the weights.

The transducer output is correlated with pressure by applying known pressures to the sensor and recording a calibration curve such as that shown in Fig. 10. The linear response of the sensor makes this invention particularly easy to calibrate over the range of interest for sub-bandage pressure measurement. The calibration curve yields a sensor offset value (intercept of best fit line and the abscissa), X₀, and a calibration constant (slope of best fit line) S_M to calculate the pressure, P, according to the formula P = S_MX_S - X₀ where P is in pressure units corresponding to those used in the generation of the calibration curve. A GUI displays the value of P as a function of time as depicted in Fig. 11. The update rate is programmable but is typically in the range 10 to 50 readings per second. Another embodiment is shown in Figs. 13 and 14 where a pre-tensioning film 42 which is selected from a range of insulating polymers including Teflon (™Du Pont) and Cling Film (Saran, ™Dow Chemical Company) is wrapped under tension around the electroded end of substrates 6 and 17 in order to adjust the sensor zero pressure capacitance to a desired value within the sensor working range. This embodiment is particularly advantageous where it is desirable to use a moulded encapsulation layer such as silicone rubber rather than a polyester or other such preformed film as the transducer packaging.

A third embodiment is shown in Figs. 15 and 16 where the base electrode 16 and top electrode 18 oppose one another and are separated by the sensor element 15. There are no substrate layers 6, 17, and necessity for these is avoided because the base electrode 16 and top electrode 18 are formed directly onto the encapsulation layers 5. This embodiment is particularly advantageous where highly compliant flexible pressure transducers are required.

It will be appreciated that the compliance of the flexible pressure transducer may be adjusted by the choice of the encapsulation polymer layer 5 which may be selected from a wide range of polymers including polyester, rubber and other biomedically approved polymers. In addition, further control over transducer compliance is afforded through the choice of encapsulation polymer layer 5 thickness. A practical range of thicknesses for polyester encapsulation polymer layers 5 is from 12 to 150 microns. The lower limit is determined largely as a compromise between the degree of compliance desired and the level of durability desired for the flexible pressure transducer 1. Generally, compliance will increase with reduced thickness. However, durability and ease of handling will decrease with reduced thickness. The upper limit is largely determined by the durability and ease of handling desired. Thicker encapsulation will generally confer increased durability and increased ease of handling possibly at the expense of an increase in intrusiveness and/or reduced the accuracy.

EMI screening consistent with best practice in sensors technology may be provided by adhesively joining a continuous layer of copper foil, or by spray coating a proprietary screening film, onto the base encapsulation layer and connecting this electrically to the screen track 37 which in turn is connected to the UTI ground. This screening film may be extended over the encapsulated interconnect track region of the transducer but should not cover the top electrode sensor region nor should it contact electrically the terminal area 43 of the transducer. The entire screening layer may be conveniently electrically isolated by encapsulating it in 12μm thick insulating adhesive tape.

The invention may also be used to form one-dimensional and two-dimensional flexible pressure transducer arrays. Fig. 17 demonstrates a one-dimensional transducer array comprising a linear series of three sensors which is particularly advantageous for measuring pressure gradients along a limb during compression therapy (such as used in the treatment of venous leg ulcers).

The transducer may be sterilized by heating in an autoclave or by exposure to UV radiation.

While preferred embodiments of the invention have been described, changes in the construction and the arrangement of parts and the performance of steps can be made by those skilled in the art, which changes are encompassed within this invention as defined by the appended claims

It will be appreciated that the transducer combines a pressure sensor, sensor package (encapsulant) (which in addition to normal packaging functions provides a means of zero offsetting the sensor) and interconnects in a single integrated solution which has an ultra low profile, is flexible, so that it can be incorporated into arrays which conform to body parts/tissue. Because of the extremely low profile of the sensor (minimally intrusive) and its structure it is relatively immune to 'hammocking" errors.

The invention is not limited to the embodiments described but may be varied in construction and detail.

Claims

1. A pressure transducer apparatus for measuring biomedical pressures, the transducer comprising:

a primary transduction part having a deformable sensor element, and

the sensor element comprises a plurality of co-planar deformable sub- elements, and

2. A pressure transducer apparatus as claimed in claim 1 , wherein the sub-elements are arranged in an array.

3. A pressure transducer apparatus as claimed in claims 1 or 2, wherein the sub- elements each have a trapezoidal shape in cross-section in at least one dimension.

4. A pressure transducer apparatus as claimed in any preceding claim, wherein the span to base ratio of the sub-elements is in the range of 0.1 to 1.

5. A pressure transducer apparatus as claimed in any preceding claim, wherein the ratio of the sub-element base width to height is in the range of 1 to 5.

6. A pressure transducer apparatus as claimed in any preceding claim, wherein the sensor element material is silicone.

7. A pressure transducer apparatus as claimed in any preceding claim, further comprising an encapsulant comprising opposed encapsulant layers.

8. A pressure transducer apparatus as claimed in claim 7, wherein the encapsulant is flexible.

9. A pressure transducer apparatus as claimed in claims 7 or 8, further comprising a substrate for each electrode, each substrate having a thickness of less than 150 microns, wherein the sensor element has a thickness of less than 30 microns, and wherein the thickness of each encapsulant layer is less than 150 microns.

10. A pressure transducer apparatus as claimed in claim 11, wherein the overall thickness of the apparatus in a normal plane through the sensor element is less than 750 microns.

11. A pressure transducer apparatus as claimed in any preceding claim, wherein the sub-elements have a trapezoidal cross section in at least one dimension and comprise a pair of parallel opposing surfaces in contact with the electrodes, and each sub-element has two or four free surfaces inclined at an angle of between 0 and 60 degrees to the normal to the plane of the electrodes.

12. A pressure transducer apparatus as claimed in any of claims 7 to 11, further comprising a screening electrode adhesively joined to the encapsulant at an external surface, and which is itself enclosed in an encapsulation layer.

13. A pressure transducer apparatus as claimed in any preceding claim, further comprising an adhesive member for attaching to medical dressings, devices or to living tissue.

14. A pressure transducer apparatus as claimed in any preceding claim, further comprising means for calibrating in terms of known values of pressure, wherein the calibrating means comprises inflatable bladders which contact the apparatus above and below the plane of the sensor element in order to impose predetermined pressure to the sensor element.

15. A method of producing a pressure transducer apparatus of any preceding claim, the method comprising the steps of:

providing a first electrode;

depositing and bonding the sensor sub-elements onto the first electrode; and

16. A method of producing a transducer apparatus as claimed in claim 15, wherein the sub-elements are moulded onto the first electrode.

17. A method of producing a transducer apparatus as claimed in claim 16, wherein the sub-elements are moulded in a mould which is formed using a lithography stamp, in turn moulded from a micro-machined silicon master.

18. A method of producing a transducer apparatus as claimed in claim 17, wherein an adhesion promotion agent is applied to the first electrode before depositing the sub-elements.

19. A method of producing a transducer apparatus as claimed in claim 18, wherein the electrodes and electrical leads are on substrates.

20. A method of producing a transducer apparatus as claimed in any of claims 15 to 19, comprising the further step of enveloping the sensor sub-elements and the electrodes in opposed encapsulant layers.

21. A method of producing a transducer apparatus as claimed in claim 20, wherein the encapsulant layers comprise polymer film.

22. A method of producing a transducer apparatus as claimed in claims 20 or 21, wherein the encapsulant layers are coated with adhesive on their inner surfaces.

23. A method of producing a transducer apparatus as claimed in any of claims 20 to 22, wherein part of an electrode protrudes from the encapsulation, and an edge connector is connected to the protruding part of the electrode.

24. A method of producing a transducer apparatus as claimed in any of claims 20 to 23, wherein the encapsulant layers are applied under controlled tension.

25. A method of producing a transducer apparatus as claimed in claim 24, wherein the tension is controlled to produce a pressure acting normal to the plane of the sensor element.

26. A method of producing a transducer apparatus as claimed in claim 25, wherein the tension is controlled to adjust the zero offset capacitance of the transducer apparatus