US20080269581A1 - Method and apparatus for measuring blood volume - Google Patents

Abstract

In one aspect, a conductance catheter is provided for measuring the volume of a fluid. The conductance catheter comprises a series of electrodes and a circuit to compensate for variations in sensitivity of the electrodes in the catheter. In another aspect, a resistivity sensor is provided for determining the resistivity of a fluid. The sensor comprises a series of electrodes spaced such that the total distance between endmost electrodes does not exceed the diameter of the catheter deploying the sensor.

Description

• This application claims priority from U.S. Provisional Application No. 60/913,734 filed on Apr. 24, 2007, the contents of which are incorporated herein by reference.
TECHNICAL FIELD
• The present invention relates generally to data acquisition systems and particularly to measuring blood volume in a living organism.
DESCRIPTION OF THE PRIOR ART
• In the field of cardiac research the standard test for measuring cardiac efficiency is the pressure volume graph. This test correlates Left Ventricle (LV) chamber pressure and volume as the heart contracts and expands. Pressure and volume values are important for quantifying efficiency in any pump system, and can be used to calculate volumetric efficiency of such systems. Cardiac efficiency is a useful measurement for studying heart disease, by quantifying the progress of the disease and measuring the effectiveness of the treatment.
• Recently, gene altered mice have increased in popularity as a means for studying heart disease, and for modelling human heart disease. Typically, LV data is measured using a catheter that is inserted into the LV. The catheter typically has separate instrumentation for measuring blood pressure and blood volume. There are several drawbacks to using data taken from anaesthetized mice, most significantly the fact that it has been found that cardiovascular data taken from an anaesthetized specimen differs significantly from free-roaming specimens.
• In order to measure cardiovascular data from a free-roaming specimen, an inplanted device is required that can operate while the specimen is active, and transmit data to the exterior of the specimen for processing. This need presents several design problems, notably size and battery life. Particularly, a reduced size provides a less invasive device, and a longer battery life decreases the number of surgical operations required to change or recharge a device. The need to reduce repeated trauma due to surgery and the cost of the surgery are driving reasons for the need to extend battery life in biological implants. These concerns are heightened when extending the application to human specimens.
• There are numerous devices that have been developed for measuring physiological pressure in living specimens, e.g., those shown in U.S. Pat. Nos. 4,796,641; 4,846,191; and 6,033,366. These devices include a catheter having a pressure sensor that is inserted into an area in the specimen having a physiological pressure, such as an artery. The sensors include a pressure transmitting catheter filled with a pressure transmitting fluid. A pressure transducer communicates with the fluid to provide an electric pressure signal representing variations in physiological pressure that can be transmitted to the exterior of the specimen. These devices are only concerned with measuring pressure, and the use of a fluid filled catheter can lead to undesirable frequency response characteristics and may exhibit head pressure artefacts.
• Other devices, e.g., that shown in U.S. Pat. No. 6,409,674 provide an implantable sensor being anchored to the interior wall of the LV in a living specimen. The sensor acquires and transmits data from within the heart to an external data receiver. This device is concerned with only measuring a single parameter, and specifically illustrates measuring pressure.
• The volume of a liquid in a cylindrical chamber, such as the left ventricle of a heart, can be derived by measuring the conductance of the fluid. The volume (V) can be calculated according to the following equation:
• $V=\frac{1}{\alpha }\rho {\mathrm{<figure-callout\; label="L"\; filenames="US20080269581A1-20081030-D00013.png"\; state="\{\{state\}\}">L</figure-callout>}}^2G.$
• The variable α represents a non dimensional correction factor attribute to the fact that the electrical field created with catheter based volumetry is not typically uniform in its distribution throughout the blood volume. The variable ρ represents the resistivity of the blood in the LV. It is important for this value to be as accurate as possible and as representative of the actual blood inside the volume being measured. This variable has the potential to change over time or due to research intervention and thus cannot be considered always consistent. The resistivity has been shown to vary with temperature, hematocrit and blood velocity. Moreover, it is possible that changes in electrolyte concentrations also alter resistivity. The variable L represents the distance between the sensing electrodes, which is fixed by the nature of the catheter or other measurement instrument being used. Finally, G is the actual conductance value that is measured by the electrodes. G may include a correction factor for an overestimation of G caused by the electric field entering the muscle. Such a correction may then represent G as G=G_block−G_correction.
• Both α and ρ can be problematic to measure when performing conductance based cardiac volumetry, since while conductance based cardiac volume is a widely used, its accuracy is limited by the non-uniformity of the transmitted electric field and the non-fixed value of the blood resistivity.
• Studies have shown that the positioning of a catheter or other measuring device in, e.g., a ventricular chamber, is important in accurately determining volume. In particular, this can occur when the catheter is not centred within the chamber. However, knowing the position of the catheter within the ventricle can be difficult with existing technology and relies upon the user's objective skill and experience.
• It is therefore an object of the following to obviate or mitigate at least one of the above-mentioned disadvantages.
SUMMARY OF THE INVENTION
• In one aspect, there is provided a sensing tip for measuring the volume of a fluid comprising one or more electrodes for measuring the volume and a resistivity sensor disposed on the tip in communication with the fluid to incorporate a current measurement of resistivity during measurement of the volume.
• In another aspect, there is provided a system for measuring volume of a fluid comprising a sensing tip having a plurality of pairs of electrodes, each pair of electrodes being connected to a circuit to compensate for variations in sensitivity of respective pairs of electrodes according to the positioning of the respective pairs along the sensing tip.
• In yet another aspect, there is provided a method for calibrating a sensing tip used for measuring volume of a fluid comprising inserting the sensing tip into a plurality of cuvettes containing fluids having differing properties, and when in each well, the method comprises: obtaining a plurality of conductance signals using a plurality of electrodes on the sensing tip; adjusting the conductance signals to compensate for variations is sensitivity of the pairs of electrodes due to the positioning of the electrodes along the sensing tip; obtaining a measurement of resistivity; and using the conductance signals and the measurement of resistivity to calibrate the sensing tip.
• In yet another aspect, there is provided a method for positioning a sensing tip disposed in a ventricle comprising obtaining an excitation waveform generated by one pair of electrodes disposed on the sensing tip; obtaining a conductance waveform sensed by another pair of electrodes disposed on the sensing tip; comparing the waveforms to determine a phase shift between the waveforms; and adjusting the positioning until the phase shift is deemed acceptable.
BRIEF DESCRIPTION OF THE DRAWINGS
• An embodiment of the invention will now be described by way of example only with reference to the appended drawings wherein:
• FIG. 1 pictorially shows a wireless cardiovascular data acquisition system.
• FIG. 2 is a schematic representation of the system ofFIG. 1.
• FIG. 3 is a magnified view of a portion the heart shown inFIG. 1.
• FIG. 4ais a partial plan view of the pressure sensing device ofFIG. 2.
• FIG. 4bis a sectional view of the sensing device shown inFIG. 4aalong the line B-B.
• FIG. 5 is an electric schematic of the pressure sensing device.
• FIG. 6 is a schematic diagram of the transmitter processing module ofFIG. 2.
• FIG. 7 is a schematic diagram of the receiver processing module ofFIG. 2.
• FIG. 8 is a timing diagram for the timing controller ofFIG. 6.
• FIG. 9 is a flow chart showing an acquisition and transmission cycle.
• FIG. 10 shows another embodiment of the sensing tip ofFIG. 3.
• FIG. 11 shows a sensing tip having a resistivity sensor
• FIG. 12 shows a circuit for measuring fluid volume.
• FIG. 13 shows a calibration circuit.
• FIG. 14 shows another calibration circuit.
• FIG. 15 illustrates a sectional view of an off-centred catheter in a ventricle.
• FIG. 16 illustrates a sectional view of a centred catheter in the ventricle ofFIG. 15.
DETAILED DESCRIPTION OF THE DRAWINGS
• A system and method are provided for accurately handling the parameters for measuring blood volume in real time. In particular, it has been found that in order to perform accurate and repeatable conductance measurements (parameter G), it is important to eliminate variations in the measuring electrode sensitivity from the equation. The following provides a calibration circuit that takes into account electrode sensitivity at the time of calibration and thus eliminates such variability. This also compensates for the non-uniformity of the electric field generated by a conductance measuring device. In order to compensate for the electrode sensitivity, an adjustable gain and offset can thus be used in each segment of a conductance measurement.
• In the following, a resistivity sensor is also provided such that the resistivity (parameter ρ) is measured on a continuous or as-needed basis rather than at time-spaced intervals or relying on a “most recent” measurement. This compensates for variations in resistivity discussed above and thus provides more accurate volume measurements.
• The calibration and resistivity sensor can be incorporated into any deployable medical device such as an implant or catheter. The following example describes an implant transmitter, however, it will be appreciated that the principles for measuring resistivity and for calibrating volume measurements discussed herein are equally applicable to other devices such as catheters.
• It will also be appreciated that the following principles are also applicable to measuring volume in any fluid and should not be considered limited only to blood.
• Also provided is a system and method for determining the position of a catheter or other measuring device in a ventricle by comparing excitation and sensed waveforms and correcting the position until the phase angle is minimized.
• Referring now toFIG. 1, one embodiment of a wireless cardiovascular data acquisition system is generally denoted bynumeral10. Thesystem10 operates to measure physical parameters of aheart12 located within abody14. Theheart12 andbody14 form part of a living organism, such as a gene altered mouse or a human. Theheart12 includes a heart chamber, in this example a Left Ventricle (LV)16 that in part communicates with thebody14 via aheart valve18. Asensing tip22 is situated in theLV16 by insertion thereof through thevalve18, and has acommunication path24 leading to a transmittingdevice20 implanted in aportion15 of thebody14, which in this example is external to theheart12. In the example shown inFIG. 1, theportion15 is in proximity of the body's clavicle. It will be appreciated that the transmittingdevice20 may be situated anywhere as desired, e.g. within theheart12 or heart chamber (i.e. LV16).
• The transmittingdevice20 wirelessly transmits data to a receivingdevice26 that in this example is attached to abelt27 external to thebody14. The receivingdevice26 may display data on ascreen28 as shown inFIG. 1, and may comprise akeypad30 for scrolling between different views. A schematic of thesystem10 is shown inFIG. 2.
• Referring now toFIG. 2, thepath24 communicates data acquired by thesensing tip22 to atransmitter processing module32 in the transmittingdevice20. The transmittingdevice20 is powered by obtaining energy from abattery34, and has atransmitter36. It will be appreciated that the use of abattery34 is for illustrative purposes only and that any suitable means for powering the transmittingdevice20 may be used such as power scavenging (converting environmental energy into electricity) or RF power transmission (energy transmitted to thedevice20 from an external source through a radio frequency signal).
• Since theprocessing module32 is preferably implanted in thebody14, the signal sent via thetransmitter36 should pass through body tissue before reaching the air. The attenuation of an RF signal by different body materials is typically highly frequency dependent. Therefore, thetransmitter36 should be selected so as to minimize the attenuation of the signal it transmits. Typically, a lower frequency is preferred to transmit the signals since the lower the frequency, the greater the depth of penetration. However, the lower the frequency, the higher the wavelength and thus the longer the antenna required at the receiving end. Therefore, thetransmitter36 should be chosen to balance these requirements depending on the particular application. A suitable frequency to achieve such a balance is 40 MHz. The power consumed by thetransmitter36 should also be considered so that it can be faithfully detected at its receiving end whilst conserving energy.
• The transmittingdevice20 communicates wirelessly with the receivingdevice26 through areceiver40. Thedevice26 has areceiver processing module38 that is adapted for processing data received from thedevice20. Thedevice26 is powered by abattery42 or suitable AC or DC power source (not shown). Thedevice26 has a series of signals (44-50) for providing electrical representations of measurements acquired using thesensing tip22, including apressure signal44, avolume signal46, atemperature signal48, and an electrocardiogram (ECG)signal50.
• InFIG. 2 these signals are shown as being external to theprocessing module38 and communicably connected to anexternal computing device52 having an analog-to-digital (A/D)converter54 connected thereto. However, it will be appreciated that the A/D converter54 may be included in either theprocessing module38 orprocessing module32, andcomputing device52 may be replaced by any suitable alternative such as processing capabilities provided by theprocessing module38. The communicable link between the receivingdevice26 and thecomputing device52 and/or A/D converter54 may be any hardwired or wireless communication channel, e.g., using Bluetooth technology.
• Thecomputing device52, external or internal to the receivingdevice26, may be any device that is capable of acquiring data and communicating with theprocessing module38. In the example shown inFIG. 2, thedevice52 is a standard personal computer (PC) having a monitor, central processing unit (CPU), keyboard, and mouse.
• Thesensing tip22 is shown in greater detail inFIG. 3. Thesensing tip22 has arounded end70 to facilitate the deployment thereof through thevalve18. In this example, aproximal electrode62 and adistal electrode60 each following the circumference of thesensing tip22 flank a pair ofinner electrodes64,66, apressure sensing device68, and atemperature sensing device69. Theelectrodes60,62,64 and66 are used to measure the volume of blood in theLV16 and are herein collectively referred to as the volume sensing device denoted bynumeral67. Theproximal electrode62 transmits a signal, and the distal electrode receives same to create an electric field in theLV16. Theinner electrodes64,66 sense this electric field to perform a conductance measurement indicative of the volume in theLV16. Theinner electrodes64,66 can be modeled conceptually as measurement probes on either side of a “resistor”, wherein the “resistor” represents the resistivity of the blood in theLV16, theinner electrodes64, are arranged to measure the potential across the “resistor”. The volume measurement and/or volume signal may also be referred to as a conductance measurement and/or conductance signal respectively, and it will be appreciated that this terminology may herein be considered interchangeable.
• Thepressure sensing device68 is used to sense the pressure of the blood in the LV Thetemperature sensing device69 is used to sense the temperature of thebody14, since it is substantially uniform throughout. Thetemperature sensing device69 is preferably comprised of a thermistor or equivalent component. Thevolume sensing device67,pressure sensing device68, andtemperature sensing device69 communicate data to the transmittingdevice20 through thepath24, thus thepath24 typically carries a number of wires, enabling data to be transmitted from thesensing tip22 to thedevice20. The length of thepath24 is dependent upon the location of thedevice20 relative to theheart12.
• Although thetemperature sensing device69 is shown inFIG. 3 as part of thesensing tip22, it will be appreciated that thedevice69 may be situated anywhere in thebody14 enabling the internal temperature of thebody14 to be measured, and this may be inside or outside of theheart12.
• An embodiment of thesensing tip22 is shown inFIGS. 4aand4b. It will be appreciated that the relative dimensions of thesensing tip22 have been exaggerated for illustrative purposes only. Thepressure sensing device68 may be any device capable of sensing a pressure. In this example, the pressure sensing device comprises a piezoresistive deflection sensor, specifically a cantileveredsensor beam80 having abase portion82 that is attached to the housing of thesensing tip22. Abase window85 in thesensing tip22 enables the base of thebeam80 to experience external pressure, and atip window86 enables the tip of thebeam80 to experience external pressure. A layer ofsealant88 inhibits thebeam80 from direct contact with its surrounding environment. However, thelayer88 permits external pressure to effect flexure of thebeam80 due to variations in the pressure of the surrounding blood. It can be seen inFIG. 4bthat electrical wires run from thesensing devices67,68 and69 to thepath24.
• An implementation of thebeam80 is shown schematically inFIG. 5, being a strain gauge sensor, on which two resistors R_x1and R_x2are mounted. When the beam bends as a result of a pressure experienced thereby, the resistances of these resistors change in opposite directions. That is, the resistance of one of the resistors increases while that of the other one decreases. As a result, the accompanying electronic circuits may be designed in a fully differential architecture which provides a higher signal to noise ratio (SNR) compared to a single ended architecture.
• The following lists suitable specifications for thepressure sensing device68, but shall in no way be considered limited thereto: nominal resistance of each resistor R_x1, R_x2being 10,000 Ohms; gauge factor of 70-80; total resistor manufacturing tolerance of +/−10-15%; maximum resistance value mismatch between the resistors of 2.4%; temperature coefficient of resistance of +5%/100° F.; and a breakdown voltage of 20V.
• These exemplary specifications illustrate that typically there may be non-idealities for thesensing device68 that would preferably be addressed when designing the circuitry therefor. For instance, due to process variations, the resistances of R_x1and R_x2are in all likelihood not going to be equal. This may generate some offset at the output. Moreover, since the resistance of the resistors R_x1and R_x2is a temperature dependent parameter, the temperature coefficient of resistance (TCR) may cause an offset due to mismatch. Hence, even if the offset is cancelled at one temperature it may not be zero at another temperature. Finally, the temperature coefficient of the gauge factor (TCGF) makes the gain of thesensing device68, temperature dependent.
• The above parameters are typically sources for measurement inaccuracies. As a result, the output of thesensing device68 may have some offset error and be dependent on temperature. In order to compensate for the above parameters, typically a signal conditioning scheme is utilized. In the example shown inFIG. 5, a Wheatstone bridge configuration is used to measure the resistance variations with two current sources I₁and I₂.
• As indicated above, R_x1and R_x2change in opposite direction as a function of strain or equivalently blood pressure in the heart as: R_x1=R₀₁(1+GF.x) and R_x2=R₀₂(1+GF.x) where R₀₁and R₀₂are the sensor resistances at zero strain, GF is the gauge factor of thesensing device68, and x is the strain. The two current sources I₁and I₂complete the bridge, and are preferably integrated into theprocessing module32 as shown inFIG. 5. In order to cancel out the resistor mismatch, TCR, and TCGF, the following equations should be valid: R₀₁I₀₂−R₀₁I₀₁=0; and TCI=(TCR+TCGF); where TCI represents the temperature coefficient of the current sources, R₀₁and R₀₂represent the resistor values at the reference temperature, and I₀₁and I₀₂represent the current of the two current sources at the reference temperature. The technology used to implement theprocessing module32 should be capable of implementing a current source with any specific temperature coefficient, and the current sources should preferably be designed to have the lowest possible supply voltage sensitivity.
• A block diagram of thetransmitter processing module32 is shown inFIG. 6. Themodule32 comprises asensing block90 and a transmittingblock92 controlled by atiming controller94. Thebattery34 which is connected to themodule32 may be controlled by a switch Thebattery34 is preferably a miniature battery of a suitable size and having a battery life that is as long as possible. A suitable battery has a life of 180 mAh, weight of 2.3 g, 1.5 Vdc, and a volume of 0.57 cc. Theswitch96 may be, e.g., magnetic or radio controlled, i.e. any suitable device capable of controlling the main power to themodule32 from thebattery34. Between thetiming controller94 and theswitch96 is a voltage regulator that provides a regulated voltage to thetiming controller94 for controlling theblocks90 and92. With the above battery specifications, a suitable regulated voltage is a 1V output.
• Thesensing block90 includes a current source block100 for the pressure sensing device68 (described above with current sources I₁and I₂) to compensate for sensor non-idealities, and are the basis of temperature compensation for thepressure sensing device68. Theblock90 also includes a conductancecurrent source102 for generating the electric field using theelectrodes60 and62; and a thermistorcurrent supply104 for thetemperature sensing device69, that preferably comprises a high resistance thermistor for minimal current drain. The outputs from these current sources (100-104) are sent to thesensing tip22 over thepath24.
• The measurements acquired by thesensing devices67,68 and69 are sent back to thesensing block90 over thepath24. The temperature signal is fed through anamplifier106 and sampled and held for transmission by a sample and holdcomponent112. Similarly, the pressure signal is fed to anamplifier110 and sample and holdcomponent116; and the volume signal is fed to anamplifier108 and sample and holdcomponent114. Theamplifiers106,108 and110 are preferably used to encourage the fidelity of the signals. The sample and holdcomponents112,114 and116 hold the signal samples while thetiming controller94 switches power from thesensing block90 to thetransmission block92.
• Thetransmission block92 has amultiplexer118 and a voltage controlled oscillator (VCO)120. Themultiplexer118 will read the samples from the blocks112-116 and arrange the signals for transmission by theVCO120. For example, themultiplexer118 may arrange the signals in sequential order for transmission. TheVCO120 is connected to anantenna121 and together make up thetransmitter36 shown inFIG. 2. Asuitable VCO120 is a Colpitts type that consumes an average current of 32 μA. Theantenna121 is preferably connected in parallel with the frequency determining inductor of theVCO120, and preferably serves as an FM transmitter with a 42 MHz transmission frequency.
• A block diagram of thereceiver processing module38 is shown inFIG. 7. Themodule38 comprises ademultiplexer122 connected to thereceiver40 of the receivingdevice26. Thedemultiplexer122 separates the signals that have been transmitted by thetransmitter36 and received by thereceiver40. If the signals are transmitted as analog signals, thedemultiplexer122 separates the received signal into individual analog signals, and in this example would provide three individual signals, atemperature signal124, apressure signal126, and avolume signal128. Thetemperature signal124 may be immediately available asoutput48, and thepressure signal126 may be immediately available asoutput44 for further processing and/or transmission to thecomputing device52. It will be appreciated that themodule38 may also comprise a further internal component for processing and analysing thesignals124,126 and128, e.g., for display purposes. Moreover, themodule38 may comprise an alarm or other device to notify a wearer of the receivingdevice26 of abnormal heart conditions. Thedisplay28 may also be used with such additional processing to output heart parameters or a computed index that represents heart health.
• Thevolume signal128 may be sent through abuffer129 and be available asoutput46. Thevolume signal128 may also be captured atblock130 for further processing to extract the ECG signal. Thispreliminary signal130 is preferably converted using an analog-to-digital converter (A/D)132, which enables signal manipulation while preserving the integrity of the original signal. It will be appreciated that the A/D132 would not be needed if the signals received have already been converted to digital signals. The A/D132 has two identical outputs, one of which is input to a digital signal processor (DSP)134. TheDSP134 is used to clean the ECG signal from the volume signal, and allows for complex signal processing. The extraction of the ECG signal is described in greater detail later.
• The signal emerging from theDSP134 is inverted by aninverter136. Theinverter136 may also be part of theDSP134. The other output from the A/D132 is buffered by thebuffer138 and the inverted signal and the buffered signal are summed at140 to produce the ECG signal142 that may also be available asoutput45. Thebuffer138 is used to maintain the synchronicity of the raw volume signal and the digitally manipulated version (i.e. by the DSP134). The delay imposed by theDSP134 would otherwise affect the results of thesum140. Thesummer140 adds the two volume signals, and since one has been inverted, the conductance part of the volume signal will be eliminated and the remaining signal will represent theECG signal142.
• Thesensing block90 and the transmittingblock92 are selectively powered using thetiming controller94 in order to conserve power. A timing diagram is shown inFIG. 8 illustrating the operation of thetiming controller94. The period T represents an entire monitoring cycle for thesystem10 including measurement and transmission. Specifically, T₁represents the period in which thesensing block90 is powered in order to obtain the necessary measurements and sample and hold the signals; and T₂represents the period in which the transmittingblock92 is powered in order to execute transmission of data from the transmittingdevice20 to the receivingdevice26.
• For example, a 2 kHz sampling rate provides a period T of 500 μs to sample and transmit data. If the acquisition period T₂is 20 μs, and transmission period T₃is 50 μs, there exists 430 μs during each cycle, in which either theblock90 or theblock92 is waiting. Thetiming controller94 uses this timing scheme to selectively turn off either theblock90 or block92 that is not being used to conserver power, which provides an increase in battery life.
• Another benefit arises from using such an energy saving timing scheme, namely the reduction of noise. Specifically, since theblock90 is powered whilst theblock92 is not, thetransmitter36 will not be affected by the noise generated by the signal conditioning, and, conversely, the sensing circuitry (block90) will not be subject to noise from thetransmitter36. A 10 μs period, represented by T₃, is left between the end of one period and the beginning of the next, which enables any circuitry that needs stabilizing to do so.
• Therefore, since the transmittingblock92 typically cannot transmit data that has not yet been collected, it would be wasting power while thesensing block90 is performs its function. If the transmittingblock92 is turned off when it is not needed, power is not consumed, and thus conserved. Similarly, thesensing block90 typically is not adding any data while thetransmitter36 is sending the previous sample, and thus does not need to consume power during that time.
• FIG. 9 shows a flow chart illustrating an example of the steps taken by thesystem10 during one complete cycle T, and the subsequent processing by the receivingdevice26. Thesensing block90 is powered which enables the current sources to power themeasurement devices67,68 and69 and obtain the measurements. These measurements are then amplified and undergo a sample and hold. Thesensing block90 is then powered “off” and the transmittingblock92 is powered “on”, wherein the time lag between theses steps is represented by T₃as explained above. Once theblock92 has power, themultiplexer118 is then able to obtain the signals stored in the sample and hold components112-116, and combine these signals for transmission. In this example themultiplexer118 preferably operates by arranging the signals in a particular sequential order that would be known to thedemultiplexer122 in order to enable thedemultiplexer122 to separate the signals at the receiving end.
• The multiplexer18 passes this “combined” signal to theVCO120 that uses theantenna121 to transmit the “combined” signal to receivingdevice26. At this point, a complete measurement cycle has been executed, and the signal that has been transmitted continues to the receivingdevice26 for further processing and/or output. The transmittingdevice20 may then repeat this cycle as required or desired.
• The receivingdevice26 receives the “combined” signal from thereceiver40. The signal is passed to thedemultiplexer122 where it is separated into its components. The temperature and pressure signals124 and126 respectively, may be available as outputs or for further processing by themodule38. Thevolume signal128 may be buffered and output at46, and may also be obtained for extracting theECG signal142 and providingoutput45. The extraction of the ECG signal142 from theraw volume signal128 is described in greater detail below, while referring to the functional blocks shown inFIG. 7 that relate thereto.
• As indicated above, the conductance orvolume signal128 acquired using thevolume sensing device67 is used to extract theECG signal142.
• The conductance signal acquired using thevolume electrodes67 consists of the conductance value of the blood in theLV16, any noise generated by the system or in the environment, and the ECG signal142 that is picked up as a component of environmental noise. As described above, in this example, the raw signals are collected and transmitted, e.g., as a combined analog waveform, without performing any signal conditioning, to the receiving device When the combined signal is received by the receivingdevice26, the individual pressure, volume and temperature signals (124,126 and128) are separated, and a process begins to separate the various components of the volume signal128 (i.e. at130).
• Theconductance signal128 is the result of an electrical field generated, by means of theelectrodes60,62, from the apex of the heart to the carotid artery. Due to myocardial contact of the conductance rings, the resulting conductance signal will also carry the ECG signal. It is generally common practice to use signal conditioning and filtering to eliminate the environmental and ECG noise components to extract theconductance signal128. In this embodiment, signal conditioning is used to not only remove the ECG component of noise to extract the conductance signal, but also to separately condition the ECG signal142 to remove the conductance portion of the signal. The result is that anECG signal142 can be collected without introducing any additional instrumentation into theLV16. Therefore, thesensing tip22 can be used to provide a more thorough cardiac assessment, using a single device.
• Once the signal is obtained at130, an A/D converter132 in theprocessing module38 converts the raw signal to a digital signal and passes the signal to each of an ECG digital signal processor (DSP) and abuffer138. Once the respective signals are processed, they are summed and afinal ECG signal142 is produced.
• In another embodiment, thevolume sensing device67 comprises a plurality of inner electrode rings, for example four as shown inFIG. 10. Since the optimal conductance measurement is performed by transmitting along the entire length of theLV16, and different organisms have differentsized hearts12, it may be desirable to incorporate multiple sets of inner electrode ring pairs. InFIG. 10, theLV16 shown inFIG. 3 is provided, as well as anLV1016 from a smaller organism shown in dashed lines. Thepair164,166 is similar to thepair64, described above, however, thesensing tip22 now includes thepairs168,170;172,174; and176,178 arranged progressively closer together and situated between theouter electrode pair60,62.
• In such an embodiment, it may be possible to selectively operate any of the electrode rings as a transmitting ring, but typically theelectrode60 would remain as the receiving electrode. In the example shown inFIG. 10, theelectrode170 would be selected as the optimal transmitting electrode for theLV1016 and then the inner sensing electrode pair would comprise theelectrodes164 and174. Therefore, numerous configurations of receiving, and sensing electrodes can be selectively chosen in order to obtain an optimal conductance signal, depending on the size of the LV (e.g.16 or1016).
• Therefore, thesystem10 enables the monitoring of a heart in a living organism by measuring both pressure and volume in a chamber of the heart, preferably theLV16. The pressure and volume measurements are acquired using asingle sensing tip22 and are communicated to a transmittingdevice20 to be wirelessly transmitted to a receivingdevice26, wherein they are used to monitor the heart. Thesystem10 may also incorporate a temperature measurement that can be transmitted with the volume and pressure measurement to provide further data for monitoring. Thesystem10 may also extract an ECG signal from the volume measurement. This allows the monitoring of up to four signals that can be used to determine the health of a heart.
• In addition to a compact design, thesystem10 may also incorporate an energy saving timing scheme that reduces the power required per acquisition cycle and thus increases the operational lifetime of the transmittingdevice20.
• As noted above a calibration scheme and resistivity meter can also be incorporated into a system such assystem10. In yet another embodiment, shown inFIG. 11,catheter22 includes a resistivity sensor (generally numeral200). In one implementation,resistivity sensor200acomprises aseries electrodes202 arranged substantially parallel to the axis of thecatheter22. In another implementation,resistivity meter200bcomprises a set of spaced apart rings204, e.g. four. The spacing is such that D is less than the diameter of therings204. The resulting field is thus small enough that it does not experience any significant effect from changing volume. In general, theresistivity sensor200 can be configured using any four electrodes spaced such that the distance apart of the further electrodes does not exceed the diameter of the catheter.
• Theresistivity sensor200 operates on a similar principle to that of a conductance catheter. Four electrodes are deployed in series, and a constant current flows through the fluid. The current travels between the twooutermost electrodes204. The two inner electrodes are used to sense the voltage created by the resistance of the blood. By measuring voltage and by applying a known current, the resistance can be obtained according to the relationship V=IR.
• Using the configuration shown inFIG. 11, a continuous or as-needed resistivity measurement can be made, which in turn enables more accurate blood volume measurements.
• As noted above, theresistivity sensor200 can be used to measure the volume of any fluid, e.g. by considering the equation for a volume of fluid in a cylinder, namely V=ρL²G. This equation is a simplified version of the typical equation used to measure LV volume, namely where α=1 and disregarding any correction factor for G. It has been found that if ρ changes in a cylindrical volume, the volume reading changes by a directly proportional amount, even though the volume in the cylinder is actually the same. By using theresistivity sensor200, the value for ρ can be changed at any time and to any desired degree, while not changing the value for the volume of the fluid. This enables the system to be calibrated using a known volume, independent of the value for ρ.
• It can therefore be seen that by incorporating aresistivity sensor200, an accurate and current resistivity value can be obtained in real time rather than relying on the accuracy of the latest measurement. Also, where resistivity measurements are obtained by drawing blood, significant discomfort can be avoided. Although two different variations ofsensor200 are shown, it will be appreciated that typically only one variation would be used.
• A schematic diagram of avolume measurement circuit206 is shown inFIG. 12. Thecircuit206 provides a comprehensive system for calibrating conductance signals for estimating blood volume, e.g. in an LV. The electric field distribution from a dipole catheter is a series of concentric rings emanating from the distal and proximal rings, as discussed above. Due to this configuration, the voltage sensed bydifferent ring segments208 varies depending on where they are located in the field. The effect is similar to a line of listeners in front of a speaker: those is the front will hear a louder signal than those in the rear.
• In thecircuit206 shown inFIG. 12, eachsegment208 of thecatheter22 is assigned an individual gain and offsetcircuit210. This enables a custom gain to be applied to each pair of rings. The effect of adjusting the gain is beneficial for both signals that may be too “weak” as well as those that could be too “strong”. In either case, interpretation can be affected and thus compensating for such effects improves the quality of the measurement. As can be seen inFIG. 12, the segments are summed atstage212 and adisplay214, e.g. a PC monitor is used to view the volume measurement.
• In order to set the individual gains, a calibration fluid having a known conductance value can be used, which corresponds to the fluid being measured, (blood or otherwise). This allows the signals to be normalized. In general, thecatheter22 is placed in a well containing a fluid of a known conductance value. The readings can then be displayed on an electronic display. The circuit then adjusts the gains for the individual segments such that the signals are all reading the same voltage output for a given solution. In addition to adjusting voltage amplitude, thecircuit206 can also adjust for baseline voltage such that all the signals have the same span and output for a given segmental volume. The calibration fluid can also be used to calibrate thesystem10 for linear output of volume. By using a series of graduate volumes, the output of thecatheter22 can be recorded as it moves from one volume to the next.
• The calibration system can also use a series of wells into which is placed fluids of differing but know conductivities. This enables verification of the linearity of thesystem10 as it measures the value of conductance G for each solution. The voltage outputs can be displayed in a plot to verify linearity and hence accuracy of the system.
• FIGS. 13 and 14 illustrate two arrangements for calibrating thecatheter22 to account for the actual and varying sensitivity of the electrodes (60-66).
• In the arrangement shown inFIG. 13, two wells of known volume in avolume cuvette220 are used to determine the volume calibration slope. In the example shown inFIG. 13, amulti-segment control box222 connected to thecatheter22 provides a segmental output that is proportional to the volume of liquid defined by the diameter of the cylinder in thecuvette220, and the spacing of each of thesegments208. As noted above, theρ sensor200 incorporated into thecatheter22 can be used to compensate for any chance in ρ that may occur. By determining the value of ρ and incorporating such a value into the calibration measurement, it is possible to change the ρ value of the fluid being measured, without changing the output voltage. The volume measured is then independent of ρ. For this arrangement, the actual value of conductivity is not required. It can be seen that in this arrangement, a correction factor can be maintained if it should change.
• Thecontrol box222 also includes the gain and offsetcircuitry210, which is adjusted such that all thesegments208 deliver the same voltage when in a given cuvette well. The gain/offsetcircuit210 can be applied in the same way regardless of the method of calibration. For example, using 20 and 50 ml volume wells, when placed in the 20 ml volume well, the following readings may be observed, shown in Table 1 below:

• TABLE 1
  Example calibration readings

  	V @	V @	Gain/Offset	Gain/Offset
  Segment	20ml	50 ml	correction @ 20 ml	correction @ 50ml
  1	2	6	−2	0
  2	0	3	−2	0
  3	5	8	−2	0
  4	3	5	−2	0
  5	−1	−2	−2	0

• One concern is that with a +/−10 volt A/D224, the range could be exceeded if the properties of the fluid change, e.g. when a saline bolus in injected to correct for parallel conductance caused by myocardial contribution. The system described herein enables the user to dial in the span and gain desired by the user and being suitable to a wide range of subjects. Theρ sensor200 in this arrangement would not need to be calibrated as it would not see a volume change going from one cuvette well to another. The sensitivity of theρ sensor200 can be corrected however if it is not reading the correct values for the fluid. This can be done since in the arrangement shown, the ρ values are known for each fluid. Being able to account for varying electrode sensitivities can be beneficial for the accuracy of the measurements.
• The output from thecontrol box222 is fed into an A/D converter224 and the volume is calibrated according to the linear relationship y=mx+b. Two of the cuvette wells, with known volumes, can thus be used to determine the value for m. In this arrangement, the value of ρ is being compensated for in thecontrol box222. It is also possible to output the value of ρ to thePC214 and use it as a scaling factor for the volume calibration.
• In the arrangement shown inFIG. 14, anequal volume cuvette230 can be used, which has a plurality of wells with equal volumes. To calibrate, a fluid having a known and precise value of conductance can be inserted into each well. In this arrangement, acontrol box232 provides a segmental output that is proportional to the conductance of the liquid defined by the diameter of the cylinder (fluid well) and the spacing of eachsegment208. Both the conductance and resistivity voltages generated by the calibration fluids are sent from the control box to thePC214, with the segmental output again being fed through A/D converter224. ThePC214 determines the slope for both the conductance G and the resistivity ρ and the values for each segment are summed to provide G_total.
• The values of ρ and G are applied to the volume formula
• $V=\frac{1}{\alpha }\rho L^2G.$
• Since both ρ and G have been obtained in a controlled and accurate method and can be monitored on an ongoing basis, the total volume accuracy can be maintained.
• It will be appreciated that the calibration of thecircuit206 can be done prior to use of thecatheter22, but may also be configured to be performed periodically while thecatheter22 is deployed to enable real time calibration. In this way, a known voltage can be used to represent a conductance level.
• It has also been found that the position of a catheter or other measuring device can be determined by comparing the excitation and measured conductance waveforms using a phase angle detector. By comparing such waveforms, it is possible to observe a phase shift between the two waveforms, which is caused by the capacitive nature of the myocardium. This is caused by the catheter being off-centred in the ventricle. An AC waveform moving through a capacitor will incur a phase delay. Since blood is a purely resistive material, it does not add appreciably to the observed phase shift. As will be discussed below, it has been recognized that by superimposing the sinusoidal electrode stimulation waveform with the sinusoidal sensed voltage waveform, the catheter electrode position within the ventricle can be determined.
• The above can be incorporated into a real-time feedback scheme that allows the use of such a measurement of phase-angle to adjust the catheter electrode position so that they are in the optimum position within the ventricle. Optimum positioning is obtained when the phase angle is at a minimum, indicating that the electric field is in a center position that minimizes field incursion into the myocardium.
• Turning now toFIGS. 15 and 16, a sectional view of aventricle16 is shown. InFIG. 15, thesensing tip22 of a catheter is offset or off-centred in theventricle16 and thus it can be seen that theexcitation signal202 is affected by the capacitive effect of the myocardium The sensedvoltage304 is thus offset from theexcitation signal302 thus creating aphase angle306 when the two waveforms are observed. It will be appreciated that thewaveforms302,304 can be viewed using any monitoring equipment, e.g. in real-time measurements, during calibration stages etc.FIG. 16 shows that by centering thesensing tip22 in theventricle16, theexcitation signal302 travel substantially through the blood rather than thecapacitive myocardium300 and thus thephase angle306 is minimized. Based on this observation, thewaveforms302,304 can be superimposed in a display as shown inFIGS. 15 and 16 (e.g. using display214 or terminal52) to provide feedback to a user so that the positioning of thesensing tip22 can be adjusted until thephase angle306 is minimized thus indicating when thesensing tip22 is substantially centered. In another embodiment, the zero-crossings of the waveforms can be compared to provide a numerical offset factor that can be observed as thesensing tip22 is adjusted until this offset value is minimized. By providing such feedback, further accuracy of the volume measurement can be attained.
• Although the invention has been described with reference to certain specific embodiments, various modifications thereof will be apparent to those skilled in the art.

Claims (20)

1. A sensing tip for measuring the volume of a fluid comprising one or more electrodes for measuring said volume and a resistivity sensor disposed on said tip in communication with said fluid to incorporate a current measurement of resistivity during measurement of said volume.

2. The sensing tip according toclaim 1 wherein said resistivity sensor comprises a series of resistivity electrodes spaced from each other such that a total distance between endmost ones of said resistivity electrodes does not exceed the diameter of said sensing tip.

3. The sensing tip according toclaim 2 wherein said resistivity electrodes are diametrically spaced.

4. The sensing tip according toclaim 2 wherein said resistivity electrodes are axially spaced.

5. The sensing tip according toclaim 1 configured for continuously obtaining said measurement of resistivity.

6. A system for measuring volume of a fluid comprising a sensing tip having a plurality of pairs of electrodes, each pair of electrodes being connected to a circuit to compensate for variations in sensitivity of respective pairs of electrodes according to the positioning of said respective pairs along said sensing tip.

7. The system according toclaim 6 wherein said circuit comprises individual gain and offset circuits to adjust the strength of a respective signal.

8. The system according toclaim 7 comprising an output for displaying said respective signal to facilitate adjustment of said gain and offset circuits.

9. The system according toclaim 6 comprising a control box for calibrating said sensing tip, said control box including a gain and offset circuit for each pair of electrodes to adjust the strength of a respective signal; and a monitor for displaying said respective signal to facilitate adjustment of said gain and offset circuits.

10. The system according toclaim 9 wherein said control box obtains a resistivity measurement from said sensing tip and incorporates a value of resistivity into a calibration measurement.

11. The system according toclaim 9 comprising a plurality of cuvettes for containing fluids of differing properties.

12. A method for calibrating a sensing tip used for measuring volume of a fluid comprising inserting said sensing tip into a plurality of cuvettes containing fluids having differing properties, and when in each well, said method comprises: obtaining a plurality of conductance signals using a plurality of electrodes on said sensing tip; adjusting said conductance signals to compensate for variations is sensitivity of said pairs of electrodes due to the positioning of said electrodes along said sensing tip; obtaining a measurement of resistivity; and using said conductance signals and said measurement of resistivity to calibrate said sensing tip.

13. The method according toclaim 12 wherein said differing properties comprises differing volumes and said measurement of resistivity is obtained using a resistivity measurement on said sensing tip.

14. The method according toclaim 12 wherein volume measurements obtained by said sensing tip are calibrated according to a linear relationship.

15. The method according toclaim 12 wherein said differing properties comprises differing known values of conductance and said method comprises providing a segmental output that is proportional to the conductance of the fluid in the respective cuvette.

16. A method for positioning a sensing tip disposed in a ventricle comprising obtaining an excitation waveform generated by one pair of electrodes disposed on said sensing tip; obtaining a conductance waveform sensed by another pair of electrodes disposed on said sensing tip; comparing said waveforms to determine a phase shift between said waveforms; and adjusting said positioning until said phase shift is deemed acceptable.

17. The method according toclaim 16 wherein said comparing comprises superimposing said waveforms to provide a visual indication of said phase shift.

18. The method according toclaim 16 wherein said phase shift is determined by comparing zero-crossings of said waveforms.

19. The method according toclaim 16 performed in real time while said sensing tip is measuring volume of fluid in said ventricle.

20. The method according toclaim 16 performed during a calibration operation.

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