A PIEZOELECTRIC SENSOR IN A LIVING ORGANISM FOR FLUID PRESSURE MEASUREMENT.

Abstract: The invention relates to a fluid pressure sensor for a lead (18) or catheter preferably intended to be placed in a living organism, such as the heart of a human being, said sensor comprising a piezoelectric element (4, 10, 11, 14) delivering an electric signal when subjected to a pressure variation, said piezoelectric element (4, 10, 11, 14) being designed to exhibit circumferential sensitivity and being disposed on at least parts of the outer surface of a rigid annular or tubular supporting structure (2). The rigidity of said sensor is such that the signal from the sensor relating to the pressure transferred to the sensor through an ongrowth of for instance tissue on the sensor would be at least 90 % of the signal from the sensor without the ongrowth.

Description: A piezoelectric sensor in a living organism for fluid pressure measurement Technical field The present invention relates to a pressure sensor for mounting in the distal end portion of a lead or catheter in accordance with the preamble of the attached independent claim. The sensor thus preferably is intended to be placed in a living organism, such as the heart of a human being, and comprises a piezoelectric element delivering an electric signal when subjected to a pressure variation. Background Art With the aid of a pressure sensor in the heart tachycardias, such as fibrillation, can be detected and also distinguished. Atrial fibrillation can e.g. be separated from ventricular fibrillation and electrical activity can be separated from mechanical activity. This is of importance for controlling the type of stimulation therapy to be delivered by a heart stimulator. From the time derivative of the blood pressure, dP/dt, an indication is obtained of the contractility of the heart, which in turn is a measure of the degree of work of the body. In H. Heynen et al. "Klinische Erfahrungen mit eine auf dem rechtsventrikularen dP/dt basierenden VVIR-Schrittmacher- syste ", Herzchrittmacher, Ausgabe 13, Nr. 1, pp.34-44, 1993 rate adaptive pacing based on right ventricular dP/dt maκ is described. A general problem with this kind of measurement is to avoid flow generated pressure components, and for the kind of pressure measurement described it is desirable that the sensor is sensitive to pressure variations all around the periphery of the sensor. US, A, 5, 271, 408 discloses a pacing lead with a flow meter for measuring intravascular blood flow comprising piezoelectric sensors that are sensitive to pressure variations all around the periphery of the sensor. In order to function correctly, the flow meter should be centred in the flow to be measured. None of the devices described in the above documents relate to a device for determining capture by means of pressure and none of the above documents address the problem of overgrowth that might cover the pressure sensors and that thus might affect the measurements. The purpose of the present invention therefore is to provide a pressure sensor, which is sensitive to pressure variations all around its periphery and which will not be affected by overgrowths. Disclosure of the Invention This purpose is achieved by means of a pressure sensor for a lead or catheter preferably intended to be placed in a living organism, for instance in the heart of a human being, in accordance with the characterizing portion of the independent claim. The invention is based on the insight that the signal from a sensor body having a specified rigidity essentially will not be affected by overgrowth since the overgrowth simply will transmit the pressure changes in a manner similar to the way pressure is transmitted by a liguid. The reason for this is the large difference in rigidity or compliance between the sensor and the overgrowth. The overgrowth will have properties similar to the properties of ordinary cardiac tissue, i. e. have a large compliance. The pressure sensor thus comprises a piezoelectric element disposed on at least parts of the outer surface of a rigid annular or tubular supporting structure. Alternatively, the piezoelectric element itself is formed as an annular or tubular rigid body. By disposing the piezoelectric element on the outer surface of a rigid supporting structure the piezoelectric element can be made of comparatively hard material necessitating only small movements or deformations of the piezoelectric material to obtain reliable measurement signals. This results in that the pressure will be sensed substantially without distortion . The sensor according to the invention can be used to detect evoked response, that is the response by the heart to stimulation pulses from a heart stimulator. After the stimulation of a heart, verification is needed that the heart actually did contract. This contraction is normally verified by study of the electrical activity of the heart, by a so-called IEGM. With a pressure sensor according to the invention this detection of the heart response to a stimulation pulse can be detected in a reliable way by measuring the blood pressure. Further, the pressure sensor according to the invention will function correctly even if the sensor should happen to be positioned against a heart wall since the sensor is sensitive in all radial directions. By forming the supporting structure of the sensor annular or tubular the advantage is obtained that electric conductors and guide wires can easily be passed through the sensor to the electrode tip. According to advantageous embodiments of the sensor according the invention, the supporting structure is formed of e.g. titanium, titanium alloy, titanium nitride, platinum, platinum alloy, niobium, niobium alloy, tantalum, tantalum alloy or carbon. The layer of conductive material applied to the piezoelectric element may be formed of e.g. titanium, titanium alloy, titanium nitride, platinum, platinum alloy, carbon, niobium, niobium alloy, tantalum, tantalum alloy or gold as these materials are biocompatible . According to another advantageous embodiment of the sensor according to the invention a low impedance charge amplifier is connected to the electrodes of the piezoelectric element to measure charge signals generated by the piezoelectric element when subjected to pressure variations. The piezoelectric element has a capacitance of the order of 1 nF and the lead has a leakage resistance that can be as low as about 50 kohm. These properties of the sensor result in that the sensor forms a high- pass filter having a cut-off freguency of about 3 kHz. With the sensor according to the invention it is desirable to measure frequencies down to e.g. 0.2 Hz. With the above mentioned embodiment of the sensor according to the invention this problem is solved by short-circuiting the piezoelectric element and measuring the charge generated in the piezoelectric material instead of measuring the voltage from the piezoelectric element. Expressed in other words, the signal from the piezoelectric element is consequently amplified by a charge amplifier with low input impedance instead of by a voltage amplifier in order to avoid the necessity of a high insulation resistance between the conductors of the piezoelectric element. Brief Description of the Drawings To explain the invention more in detail by way of examples, chosen embodiments of the sensor according to the invention will now be described with reference to the accompanying drawings in which Figures 1 through 3 show three embodiments of the sensor according to the invention in longitudinal section, connected to a charge amplifier in figures 1 and 2, Figure 4 shows an alternative embodiment of the sensor configuration, Figure 5 illustrates the mounting of a sensor according to the invention in a pacing lead, and Figure 6 shows the equivalent electrical circuit of the sensor in figures 1 and 2 connected to a charge amplifier. Description of Preferred Embodiments In figure 1 a longitudinal section of a first embodiment of the sensor according to the invention is shown. This embodiment comprises a sensing element in the form of a tube 2 which on its outer surface is covered with a layer 4 of a piezoelectric material, preferably a ceramic piezoelectric material. Preferred materials are PZT and potassium-sodium- niobate. The layer may be applied on the tube 2 by sputtering, laser-ablation or any other suitable method. On the outer surface of the piezoelectric layer 4 a layer 6 of a conductive material, e.g. titanium, titanium alloy, titanium nitride, platinum, platinum alloy, carbon, niobium, niobium alloy, tantalum or tantalum alloy is applied in a similar way. In this way the tube 2 and the layer 6 form electrodes which are connected to an amplifier 8 for measuring the charge produced in the piezoelectric layer 4 when it is subject to pressure variations . The tube 2 is a rigid tube, that may be placed inside the outer silicon rubber insulation of a pacing lead 18 with the electrodes connected through two conductors to a charge amplifier mounted in the pacemaker, see figure 5. Alternatively, the human tissue can replace one of the conductors, i. e. the piezoelectric material is in direct contact with blood or tissue. In this case the piezoelectric material should have such properties that the material could be considered biocompatible. In the case of a bipolar lead the sensor can be positioned outside the inner coiled conductor and the inner silicon insulation but inside the outer silicon insulation. As materials for the tube 2 and the conducting layer 6 are chosen for instance titanium, titanium alloy, titanium nitride, platinum, platinum alloy, niobium, niobium alloy, tantalum, tantalum alloy or carbon and titanium, titanium alloy, titanium nitride, platinum, platinum alloy, carbon, niobium, niobium alloy, tantalum, tantalum alloy or gold respectively. In figure 2 a second embodiment of the sensor according to the invention is shown comprising a tube of a piezoelectric material 10 provided around the inner tube 2 of platinum-iridium alloy, titanium or carbon. Electric contact between these two tubes 2, 10 may for instance be formed with an electrically conducting glue. On the outer surface of the piezoelectric tube 10 a conducting layer 6 is applied as in the embodiment in figure 1. Also, like in the previously described embodiment, the tube 2 and the layer 6 form electrodes that are connected to a charge amplifier 8 mounted in a pacemaker for measuring electrical charges generated in the piezoelectric tube, when the tube is subjected to radial pressure variations. Figure 3 shows a third embodiment of the sensor element of the sensor according to the invention comprising a rigid tube 11 of a piezoelectric material with inner and outer electrodes 13, i.e. the supporting structure and the piezoelectric layer are both made of piezo-electric material and may also be made of the same material. In this case the tube is provided with an inner conducting layer as well. Figure 4 shows an embodiment in which the piezoelectric element is formed as longitudinal strips 14 of piezoelectric material interspacedly distributed around the outer surface of the supporting tube 2. As mentioned above an important feature of the sensor according to the invention resides in the fact that it is sensitive to pressure variations all around its periphery. Thus it will operate reliable also if it should happen to be positioned against a heart wall. Since the sensor has an annular or tubular design electric conductors and guide wires can easily be passed through the sensor to the electrode tip 16. This can for instance be seen in figure 5 illustrating the positioning of a pressure sensor according to the invention in a lead 18 at 20 in the distal end portion. Since the pressure sensor according to the invention comprises a rigid supporting structure for the piezoelectric element, this element can be chosen such that only small deformations are needed to get reliable measurement signals. The operation of the sensor thus will not be affected by overgrowths . The rigidity of the sensor depends on the material in the support and/or the piezoelectric material, on the amount of material and on the shape of the sensor and the rigidity required by the invention thus can be obtained in a number of ways . The rigidity of the sensor thus preferably is defined with the aid of a silicon elastomer having a modulus of elasticity of at least 1.49 MPa measured at a deformation of 6 %. One example of a suitable commercially available elastomer is "Sylgard ® 567 Primerless Silicone Encapsulant" manufactured by DOW CORNING CORPORATION, Midland, Michigan (this standard silicon elastomer further has a durometer hardness of 38 points Shore A according to CTM 0099, a tensile strength of 200 psi(l,38 MPa) according to CTM 0137A, an elongation of 100% according to CTM 0137A, a specific gravity of 1.24xl0 3 kg/m 3 at 25°C according to CTM 0022, a volume coefficient of thermal expansion of 8.49xl0 ~4 CC/CC°C according to CTM 0653 and a lap shear of 200 psi (1,38 MPa) according to CTM 0243, all based on a sample thickness of 125 mils (3.2 mm)). For the sake of comparison it should however be noted that a solid sensor element made of potassium-sodium-niobate (which is the preferred material) would have a modulus of elasticity of about 80 GPa. When the rigidity of the sensor is to be determined, the sensor is first immersed in distilled water at 20°C to a depth of 50 mm in a pressure chamber and the pressure is increased with 2 kPa. The charge generated in the sensor by the increase of pressure is then measured by means of an electrometer. The sensor is removed from the pressure chamber and dried. The active parts of the sensor are then covered with a 1 mm thick layer of the silicon elastomer defined above and and the sensor is again immersed in distilled water in the pressure chamber to a depth of 50 mm. The pressure in the pressure chamber is again raised to 2kPa and the charge generated in the sensor is again measured by means of an electrometer. If the charge is 90% of the previous signal or more, then a first requirement for the rigidity of the sensor has been met . The above test is then repeated with the difference that the pressure is varied sinusoidally up to said 2 kPa, starting with a frequency of 1 Hz. The frequency is increased from 1 to 30 Hz and the amplitude of the signal from the sensor is measured. The amplitude of the signal from the sensor, when covered by said layer of silicon elastomer, should not deviate more than 3 dB from the amplitude of the signal from a sensor not covered with silicon elastomer over the entire frequency range between 1 and 30 Hz. This is the second requirement for the rigidity of the sensor. The pressure for instance could be varied by means of pulsed pressurised air or by means of a movable plunger forming one wall, or a part thereof, of the pressure chamber. The piezoelectric element can electrically be represented as a voltage source V in series with a capacitance C p i ezo - The lead can electrically be represented as a resistance Ri ead/ see figure 6. The piezoelectric element forming the sensing element is connected through the lead to a charge amplifier, located inside the pacemaker 22. The capacitance Ci ezo of the piezo-element is normally of the order of 1 nF and the insulation resistance Rι eac ι between the lead wires can be as low as 50 kohm. These two quantities form a high-pass filter having a cut-off frequency of the order of kHz, in the worst situation of approximately 3 kHz, when a voltage amplifier is used. With the sensor according to the invention it is desirable to measure pressure variations of frequencies down to 0.2 Hz. This problem is solved in the sensor according to the invention by connecting to the piezoelectric element a charge amplifier 8 having a low input resistance Ri for measuring the charge produced in the piezoelectric element, when it is subjected to pressure variations, instead of measuring the voltage from the piezoelectric element. The amplification in this case is given by the ratio C p i ezo /Cι and the cut-off frequency is equal to l/(2πR] . Cι) As appears from figure 5 the piezoelectric element in practice will be short-circuited by the input resistance Ri of the charge amplifier 8 and the requirement for a high insulation resistance between the conductors of the piezoelectric element is eliminated. In the case the sensor is to be covered with a layer of silicone, polyurethane or similar in use, then the test silicon layer is applied on top of this layer. The above definitions of the rigidity is based on the rationale that the compliance of the silicon elastomer is close to but less than the compliance of tissue grown on the sensor and thus can be used to define the rigidity of the sensor. Although normally a rigidity giving a charge from a covered sensor that is 90% of the charge from from the not covered the sensor normally is sufficient, it of course could be preferable if the rigidity is such that charges amounting to more that 95% were obtained, for instance 97%, 99% or even approaching 100%. The silicon elastomer in the embodiment illustrated above preferably is in the form of a commercially available silicon tube having suitable dimensions. A length of the tube is simply pulled over the sensor, care being taken that the tube extends well past the ends of the sensor. The layer of silicon elastomer however also could be applied by dipping the sensor in liquid, uncured silicon elastomer, if necessary several times until the required thickness has been obtained, and then curing the elastomer.

Claims: Claims 1. A fluid pressure sensor for a lead (18) or catheter preferably intended to be placed in a living organism, such as the heart of a human being, said sensor comprising a piezoelectric element (4,10,11,14) delivering an electric signal when subjected to a pressure variation, said piezoelectric element (4,10,11,14) being designed to exhibit circumferential sensitivity and being disposed on at least parts of the outer surface of a rigid annular or tubular supporting structure (2) , characterized in that the rigidity of said sensor is such that a) the charge generated by said sensor when said sensor is covered by a 1 mm thick layer of a silicon elastomer having a modulus of elasticity of at least 1.49 MPa (as measured with 6% deformation of the material) and is immersed in distilled water at 20°C to a depth of 50 mm in a pressure chamber and subjected to an increase of pressure amounting to 2 kPa, is at least 90% of the charge generated by said sensor when immersed into distilled water at 20°C to a depth of 50 mm in a pressure chamber and subjected to an increase of pressure of 2 kPa but not covered with said silicon elastomer layer, b) the amplitude of the signal from said sensor when said sensor is covered by a 1 mm thick layer of said specific, commercially available silicon elastomer and is immersed in distilled water at 20°C to a depth of 50 mm in a pressure chamber and subjected to a pressure increase varying sinusoidally between 0 and 2 kPa at a frequency increasing from 1 to 30 Hz over the entire frequency range between 1 and 30Hz does not deviate more than 3 dB from the amplitude of the signal obtained from said sensor when said sensor is immersed in distilled water at 20°C to a depth of 50 mm in a pressure chamber and subjected to a pressure increase varying sinusoidally between 0 and 2 kPa at a frequency increasing from 1 to 30 Hz but not covered by said silicon elastomer. 2. The sensor according to claim 1, characterized in that the piezoelectric element is formed as a layer (4) of piezoelectric material on the outer surface of said supporting structure (2) . 3. The sensor according to claim 1, characterized in that the piezoelectric element is formed as strips (14) of piezoelectric material distributed at interspaces around the outer surface of the supporting structure (2). 4. The sensor according to claim 1, characterized in that the piezoelectric element is formed of a tube (11) or ring of piezoelectric material, positioned around said rigid supporting structure (2 ) . 5. The sensor according to any of claims 1 through 4, characterized in that said supporting structure (2) also is formed of a piezoelectric material. 6. The sensor according to claim 5, characterized in that said supporting structure (2) is formed of a piezoelectric material different from the material in the piezoelectric layer (4, 10, 11, 14) 7. The sensor according to any of the claims 1 through 4, characterized in that said supporting structure (2) is formed of titanium, titanium alloy, titanium nitride, platinum, platinum alloy, niobium, niobium alloy, tantalum, tantalum alloy or carbo . 8. The sensor according to any of the claims 1 through 7, characterized in that said supporting structure (2) comprises a rigid tube or ring. 9. The sensor according to any of the claims 1 through 8, characterized in that layers of conductive material (6,13) are arranged in contact with the piezoelectric element (4,10,14) to form its electrodes, said layers comprising titanium, titanium alloy, titanium nitride, platinum, platinum alloy, carbon, niobium, niobium alloy, tantalum or tantalum alloy. 10. The sensor according to claim 8, characterized in that a low input impedance charge amplifier (8) is connected to the electrodes (6,13) of the piezoelectric element (4,10,11,14) to measure charge signals generated by the piezoelectric element when subjected to pressure variations. 11. A lead comprising a fluid pressure sensor according to any one of claims 1 - 10 characterized in that said piezoelectric element is formed as an annular or tubular rigid body (11) . 12. The lead according to claim 11, characterized in that said lead is a bipolar lead and in that the sensor (20) is mounted coaxially with the lead, between the inner and outer insulating layers of the lead (18).

Inventors: KJELLMAN CHARLOTTE, LIDMAN JOHAN, LJUNGSTROEM KARIN, NILSSON KENTH, KJELLMAN, CHARLOTTE, LIDMAN, JOHAN, LJUNGSTROEM, KARIN, NILSSON, KENTH
Applicants: KJELLMAN CHARLOTTE, LIDMAN JOHAN, LJUNGSTROEM KARIN, NILSSON KENTH, ST JUDE MEDICAL, KJELLMAN, CHARLOTTE, LIDMAN, JOHAN, LJUNGSTROEM, KARIN, NILSSON, KENTH, ST. JUDE MEDICAL AB
Application year: 2001
Main classification: A61B5
IPC codes: A61B5, A61B19, G01L9