CA2169374A1 - Method and apparatus for measuring continuous blood flow at low power - Google Patents
Method and apparatus for measuring continuous blood flow at low powerInfo
- Publication number
- CA2169374A1 CA2169374A1 CA002169374A CA2169374A CA2169374A1 CA 2169374 A1 CA2169374 A1 CA 2169374A1 CA 002169374 A CA002169374 A CA 002169374A CA 2169374 A CA2169374 A CA 2169374A CA 2169374 A1 CA2169374 A1 CA 2169374A1
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- level
- location
- blood flow
- determining
- temperature
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Classifications
-
- A—HUMAN NECESSITIES
- A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
- A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
- A61B5/00—Measuring for diagnostic purposes; Identification of persons
- A61B5/02—Detecting, measuring or recording pulse, heart rate, blood pressure or blood flow; Combined pulse/heart-rate/blood pressure determination; Evaluating a cardiovascular condition not otherwise provided for, e.g. using combinations of techniques provided for in this group with electrocardiography or electroauscultation; Heart catheters for measuring blood pressure
- A61B5/026—Measuring blood flow
- A61B5/0275—Measuring blood flow using tracers, e.g. dye dilution
- A61B5/028—Measuring blood flow using tracers, e.g. dye dilution by thermo-dilution
-
- A—HUMAN NECESSITIES
- A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
- A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
- A61B5/00—Measuring for diagnostic purposes; Identification of persons
- A61B5/02—Detecting, measuring or recording pulse, heart rate, blood pressure or blood flow; Combined pulse/heart-rate/blood pressure determination; Evaluating a cardiovascular condition not otherwise provided for, e.g. using combinations of techniques provided for in this group with electrocardiography or electroauscultation; Heart catheters for measuring blood pressure
- A61B5/026—Measuring blood flow
- A61B5/029—Measuring or recording blood output from the heart, e.g. minute volume
Abstract
A technique for determining blood flow in a living body by changing the thermal energy level by a predetermined amount at a site in a blood flow path and detecting temperatures at locations upstream and downstream of the site. The temperature difference at such locations is determined and the blood flow is calculated as a function of the change in energy level and of the temperature difference measured prior to and following the change in energy level.
Description
~ ~ wO 95/OSllS ~ r ~ 2 ~ 6 9 3 7 4 PCT~S~/03~1 METHOD AND APPARATUS FOR MEASURING CONTINUOUS
BLOOD FLOW AT LOW POWER
Introduction This invention relates generally to techniques for measuring blood flow in a body and, more particularly, to the use preferably of one or more temperature sensors for measuring thermal energy changes in the blood flowing through the heart and to the use o~ unique data processing techniques in response thereto for determining cardiac output.
Backqround of the Invention While the invention can be used generally to measure blood flow at various locations in a body, it is particularly useful in measuring blood flow in the heart so as to permit the measurement of cardiac output.
Many techniques for measuring cardiac output have been suggested in the art. Exemplary thermodilution techniques described in the technical and patent literature include:
"A Continuous Cardiac Output Computer Based On Thermodilution Principles", Normann et al., Annals of Biomedical Engineering, Vol. 17, 1989; "Thermodilution Cardiac Output Determination With A single Flow-Directed Catheter", Forrester, et al., American Heart Journal, Vol. 83, No. 3, 1972; "Understanding Techniques for Measuring Cardiac Output", Taylor, et al., Biomedical Instrumentation & Technology, May/June 1990; United States Patent No. 4,507,974 of M.L. Yelderman, issued April 2, 1985; United States Patent No. 4,785,823, of Eggers et al., issued on November 22, 1988; and United States Patent No. 5,000,190, of John H. Petre, issued on March 19, 1991.
WO95/05115 ~ ~ 2 1 6 9 3 7 4 PCT~S94/09064 ~
A principal limitation in the quanification of cardiac output is the existence of thermal fluctuations inherent in the bloodstream. Previous methods work with those fluctuations while observing the effects of an input signal to calculate cardiac output. The invention described herein uses a differential measurement technique to substantially eliminate the effect of the thermal fluctuations, permitting the use of a minimal thermal input signal, which allows frequent or continuous measurements.
It is desirable to obtain accurate cardiac output measurements in an effectively continuous manner, i.e., several times a minute, so that a diagnosis can be achieved more rapidly and so that rapid changes in a patient's condition can be monitored on a more continuous basis than is possible using current techniques.
Moreover, it is desirable to obtain instantaneous measurements of the cardiac output on a beat-to-beat basis to evaluate the relative changes which occur from beat to beat, as well as to determine the presence of regurgitation.
Brief Summary of the Invention In accordance with general principal of the invention, blood flow and/or cardiac output is determined rapidly, using a technique by which an indicator substance, or agent, is introduced into the bloodstream between a pair of detectors. The detectors are sensitive to a parameter functionally related to the concentration or magnitude in the bloodstream of the selected indicator agent. The detectors are positioned apart by a distance functionally sufficient to allow a measurement to be made of the differential value of the selected parameter as it WO95105115 . 1~ ~ ` 2 ~ 69374 PCT~S94/09064 exists from time-to-time between the two detectors. The indicator agent, for example, may be a substance to change the pH of the blood, a fluid bolus carrying thermal energy, or a substance to change a selected characteristic of the blood, or the direct introduction of thermal energy, or the like.
A determination is made of the difference in the values of the selected blood parameter as it exists at the two detectors, prior to the introduction of indicator agent (i.e., the first differential value). The selected indicator agent is then introduced in a predetermined magnitude. Then again a determination is made of the difference in values of the selected parameter as it exists at the locations of the two detectors (i.e., the second differential value).
Blood flow or cardiac output, depending on the specific location of the detectors, can then be determined as a function of the difference between the first differential value and the second differential value. Because the ultimate measurement of blood flow or cardiac output is based on the difference of the differences, the system operates effectively with the introduction of the indicator agent in a very low magnitude. In turn, this allows measurements to be made rapidly 80 that effectively continuous measurements are obtained.
In accordance with a preferred embodiment of the invention, for example, cardiac output can be determined rapidly and with low levels of thermal energy input. To achieve such operation, in a preferred embodiment, the technique of the invention uses a pair of temperature sensors positioned at two selected locations within a catheter which has been inserted into the path of the 2 1 6 q 3 7 4 PCT~S~ 3~61 ~
blood flowing through the heart of a living body. The sensors detect the temperature difference between the two locations. Depending on the location of the temperature sensors in the circulatory system, the measured temperature difference varies over time. It has been observed that when the temperature sensors are placed within the heart, e.g., so that one sensor lies in the vena cava, for example, and the second in the right ventricle or pulmonary artery, the temperature difference varies in a synchronous manner with the respiratory cycle.
Thus, in the preferred embodiment of the invention the temperature difference over at least one respiratory cycle is measured and averaged to provide an average temperature difference. The averaging, or integrating, action effectively eliminates, as a confounding factor in the determination of cardiac output, the effect of instantaneous blood temperature fluctuations, such as cyclical, respiratory-induced fluctuations.
To make such determinations, an average temperature difference is first calculated over a time period of at least one respiratory cycle in which no thermal energy is introduced into the blood flow path. Thermal energy of a predetermined and relatively low magnitude is then introduced into the blood flow path to produce a heating action therein at a location between the two temperature sensors. Once the temperature rise induced by the heating stabilizes, the average temperature difference between the two locations is again calculated from temperature difference measurements over a time period of at least one respiratory cycle at the higher temperature level. The difference between the average temperature differences which occurs when the thermal energy is turned on, referred to as the rising temperature change, ~ W095/05115 - ,~ s( - 2 1 6 9 3 7~ PCT~S94/ogO64 is determined. The difference between the average temperature differences which occurs when the thermal energy is turned off, referred to as the falling temperature change, is similarly determined. The cardiac output is calculated as a function of the thermal energy input and the rising and falling temperature changes.
Because a relatively low level of thermal energy is used in making measurements, the overall sequence of determinations can be safely repeated multiple times per minute, for example, so that an effectively continuous, or quasi-continuous, determination of cardiac output is obtained.
In accordance with a further embodiment of the invention, a temperature sensor that also acts as a source of thermal energy, e.g., a thermistor, is positioned at a third location in the cardiac blood flow path. Power is supplied to the sensor sufficient to elevate the temperature of the sensor from a first temperature level to a second temperature level. In one embodiment of the invention, the temperature of the sensor is changed from the first to the second level and i9 maintained constant at said second level by varying the power that i8 supplied thereto. Such varying power is proportional to the instantaneous flow velocity and, hence, assuming a constant flow area, is proportional to the instantaneous cardiac output. Measurement of the sensor heating power and the temperature increment at the sensor can thus be used to continuously effect a determination of the instantaneous cardiac output.
30 Further, for example, when the sensor is placed downstream at the outlet of one of the heart chambers, the variation in flow output over the cardiac cycle can be analyzed to provide an indication of the regurgitation characteristics of the heart outlet valve over the cardiac cycle. Moreover, such instantaneous cardiac WO9S/05115 2 1 6 9 3 7 4 PCT~S~09^C4 ~
output determination can be further refined to compensate for fluctuations in the temperature of the blood flowing through the heart by measuring the instantaneous temperature of the blood with another temperature sensor at a nearby location and appropriately taking into account such temperature variations when determining the cardiac output.
In another application, both the continuous cardiac output determinations and the instantaneous cardiac output determinations, as described above, can be combined. Thus, three temperature sensors and a source of thermal energy can all be used in combination to simultaneously provide an accurate and effectively continuous determination of time-averaged cardiac output, and a determination of instantaneous cardiac output at .
each instant of the cardiac cycle. In still another application, two temperature sensors and a source of thermal energy can be used in an appropriate sequence to provide the averaged cardiac output determination and the instantaneous cardiac output determination.
Description of the Invention The invention can be described in more detail with the help of the accompanying drawings wherein FIG. 1 shows a simplified diagrammatic view of a human heart;
FIG. 2 shows a simplified diagrammatic view of.a catheter useful in the invention;
FIG. 3 shows a flow chart depicting steps in a process used in the invention;
FIG. 3A shows a smooth temperature difference curve obtained in the process depicted in FIG. 3;
FIG. 4 shows a flow chart depicting steps in another process used in the invention;
FIG. 5 shows a flow chart depicting steps in still another process of the invention;
~ W095/05115 ~ 2 ~ 6q37~ PCT~S9410906~
FIG. 6 shows a graph depicting a temperature/time relation used in the invention;
FIG. 7A shows a simplified diagrammatic view of another catheter used in the invention;
FIG. 7B shows a ~low chart depicting steps in still another process of the invention;
FIGS. 8A, 8B, and 8C show flow charts depicting still other processes of the invention;
FIGS. 9A, 9B, and 9C show graphs of parameter relationships used in the invention; and FIG. 10 shows a graph useful ~or calibrating the flow values for a catheter used in the invention.
As can be seen in FIG. l, which represents a human heart 10 i~n a much simplified diagrammatic form, a flexible catheter 11 is inserted through the veins into the right atrium, or auricle, 12 of the heart and, thence, through the right ventricle 13 until the end of the catheter resides in or near the exit, or pulmonary, artery 14 which leads to the lungs. As is well known, blood flows (as represented by the arrows) from the input vein 15, i.e., the vena cava, into the right atrium and right ventricle and thence outwardly to the lungs and subsequently returns from the lungs into the left atrium 16, through the left ventricle 17 and thence outwardly into the aorta 18.
In accordance with the embodiment of the invention, shown with reference to FIG. 1, temperature sensors, e.g., thermistors, are carried by the catheter so that, 30 when inserted as shown in FIG. l, a first sensor 19 is positioned at a location within the vena cava 15 or right atrium 12 and a second sensor 20 is positioned at a location in or near the pulmonary artery 14.
For simplicity, the flexible catheter 11 is depicted in FIG. 2 in an extended condition with temperature WO95/05115 2 ~ 6 93 74 PCT~S94/0906~ ~
sensors 19 and 20 at two different locations for measuring temperatures T1 and T2, respectively. A power source 21 of thermal energy which is borne, or carried, by the catheter 11 is positioned in the right atrium at a location between sensors 19 and 20. In a particular embodiment, the catheter-borne source is, for example, a coil of resistive wire placed on or embedded in the surface of catheter 11, to which an AC or a DC voltage (not shown) at a controllable level is supplied so as to generate thermal energy, i.e. heat. The magnitude of the thermal energy can be suitably controlled to insert a predetermined amount of thermal energy at a selected time, which thermal energy is transferred to the blood flowing through the heart so as to raise its temperature.
The energy source is positioned at a sufficient distance from the sensor 19 that the latter is effectively thermally isolated from the site of the thermal energy source.
While the locations of the sensors 19 and 20 and the energy source 21 can be as shown in FIG. 1, alternative locations can also be used. Thus, the sensor 19 can be positioned in the vena cava 15, while the energy source 21 is located in the right atrium 12 and the sensor 20 in either the right atrium or the right ventricle.
Moreover, if sensor 19 is positioned in the vena cava 15, the entire energy source 21, which is normally elongated, need not be located in the right atrium and can have a portion thereof in the vena cava and a portion thereof in the right atrium. Such source should preferably be at least partially located in the right atrium. Further, sensor 20 may be positioned in the right ventricle near the pulmonary artery 14 or may be located in the pulmonary artery itself at or near the right ventricle.
The temperatures T1 and T2 at locations 19 and 20 upstream and downstream, respectively, from the thermal ~ WO95/05115 ~ 2 1 6937~ PCT~S94/09064 energy source 21 are monitored and processed appropriately by a digital microprocessor. In accordance with the invention, the instantaneous temperatures are obtained as the outputs T1(t) and T2(t) of the temperature sensors l9 and 20, respectively. The outputs are connected to a differential amplifier to generate an analog signal which is proportional to the temperature difference ~T(t)=T1(t)-T2(t) between them.
The temperature difference signal aT(t) is digitized and sampled at selected time intervals by an analog-to-digital/sampling circuit. The digitized sampled temperature difference values and the known thermal energy values are supplied to a digital microprocessor which then suitably proce8ses the data to provide the desired cardiac output information. The processing stages used in the host microprocessor are implemented by suitable programming of the microprocessor and are discussed below with the help of FIGS. 3-6.
The source 21 of thermal energy is alternately turned on and off. If it is assumed that thermal stability is reached after each change and that there is a substantially constant rate of blood flow, a stable temperature difference can be measured in each case. The quantity of blood flowing past the thermal energy source, i.e., the cardiac output, can be derived from such temperature difference measurements. However, such derivation is complicated by two factors which may affect the measurement of blood flow. First, the rate of blood flow through the heart is not substantially constant but surges with each heart contraction. Second, the temperature of the blood flowing through the heart is not constant but varies with each respiratory (breathing) cycle. In a preferred embodiment, the processing of the data ta~es such factors into account, as discussed below.
WO95/05115 ~ 2 1 6 9 3 7 4 PCT~S94/09064 ~
The process for determining cardiac output is performed in a microprocessor 2 which in a first embodiment is programmed to respond to the temperatures sensed at Tl and T2 and to perform the steps depicted in accordance with the flow charts shown in FIGS. 3-5. From a knowledge of such flow charts, it would be well within the skill of those in the art to appropriately program any suitable and known digital microprocessor, such as a personal computer, to perform the steps shown.
FIG. 3 depicts a basic process, identified as Process I, which is used in the overall processing of temperature data for determining cardiac output, as subsequently depicted in FIGS. 4 and 5. In the basic process shown in FIG. 3, a temperature difference as a function of time ~T(t) is determined by a differential amplifier which responds to T1(t) and T2(t). Such differences may be effectively smoothed, or filtered, to produce a smooth temperature difference curve, as shown in FIG. 3A, which varies as a function of time in a cyclic manner which depends principally on the respiratory cycle of the person whose cardiac output is being determined.
The periods ~ 2. . . ~n for each respiratory cycle are determined over n cycles. A characteristic of the temperature difference at each cycle is determined. For example, such characteristic preferably is the averaged temperature difference during each cycle (~T7l, ~ T72 ~ T72 Q T7n ) (Alternatively, for example, the peak temperature differences may be the determined characteristic.) These averaged temperature differences (~T7n) are added for the n cycles involved and are divided by n to determine an averaged temperature difference per cycle (~T7n)~ The use of Process I is depicted in the process steps shown -~ W095/OS115 ~ 2 i 6 ~ 3 7 4 PCT~S94/09064 in FIG. 4, identified as Process II.
o As seen therein, the steps of Process I are first performed when the source 21 of thermal energy (i.e., a heater) is turned off and the average AToff value per cycle is determined and suitably stored. The sampling time at which such determination is made is depicted in FIG. 6 as the sample time period Sl.
The heater 21 is then turned on for a specific time period to supply a known amount of power P to the blood flowing through the heart and, accordingly, the temperature of the blood flowing past the heater rises and the temperature difference ~T(t) rises over a transition, or delay, rise time period, t~1, shown in FIG. 6 and designated as D1, after which the temperature difference generally stabilizes over a second sample time period S2. As seen in FIG. 4, after the heater 21 is turned on and the temperature has stabilized, Process I
is performed, again over n cycles, e.g., over the time period S2, and the averaged temperature difference ~Ton is determined with the heater turned on and is suitably stored. The heater is then turned off and the temperature falls over a transition, or delay, fall time period tFl shown in FIG. 6 and designated as D2, generally to its former value.
Cardiac output is calculated using the averaged temperature differences when the energy is off and the averaged temperature differences when the energy is on, by the relationship:
F = P
Cp ( ~ Ton ~ ~ Tof f ) where: F = Flow P = Power Cp = heat capacitance ~Ton = average temperature for power on off = average temperature for power off WO95/05115 ~ .' =' 2 1 6 9 3 7 4 PCT~S94/09064 ~
As seen in FIG . 5, the steps of Process II are repeated indefinitely for N data collection cycles, a data collection cycle being designated as including the time periods Sl, Dl, S2, and D2, as shown in FIG. 6. For each data collection cycle the rise tir.le temperature difference ATR between the averaged temperature difference ~Ton at S2 and the averaged temperature difference QToff at Sl and the fall time temperature difference ATp between the averaged temperature difference ATOffat Sl and the averaged temperature difference QTon at S2 are determined.
The flow, FR is calculated for each data collection cycle from the known amount of power P introduced into the blood flow stream by the energy source, or heater 21, from the known heat capacitance of blood, Cp and from the difference in the averaged temperature differences ~Ton and QToff/ which occurs over the data collection cycle Sl + Dl + S2 in accordance with the following relationship:
FR = P
Cp (~Ton ~ A Toff~
In a similar manner, the flow FP is calculated from P, Cp and the difference in the averaged temperature differences ATon and ~Toff which occurs over the later portion of the data collection cycle Sl + Dl + S2 in ~ WO95/05115 . ~- Z1~ ~3~4 PcT~s~4,a~64 accordance with the following relationship:
.
FF = P
Cp ( 1~ Ton ~ ~ Tof f ) FR and FF can be averaged to obtain the averaged flow (F) over one data collection cycle as shown in FIG. 6.
F = F~ + FF
A suitable calibration constant can be used to adjust the values Of FR~ FF and F.
Accordingly, by using two temperature sensors l9 and 20, cardiac output can be determined several times a minute in accordance with FIGS . 3-6, yielding an effectively continuous cardiac output value. Because such measurements can be made using relatively low power levels, the danger that the heart may be damaged by the introduction of thermal energy is substantially eliminated.
It will be apparent that the foregoing technique, which has been described in connection with the direct introduction of heat as an indicator agent and the measurement of temperature, can readily be performed by those skilled in the art by using indicator agents which affect the pH of the blood or change other blood parameters.
In some situations it may be desirable to provide WO95/05115 ~ 2 1 6 9 3 7 4 PCT~S~/0~
more frequent indications of cardiac output, such as, for example, the instantaneous cardiac output or the cardiac output averaged over each individual cardiac cycle (i.e.
each heart beat). Such information can be provided using the further embodiments of the invention discussed below with reference to FIGS. 7-8. A single temperature sensor 30 at a location near the distal end of the catheter 31 (as shown in FIG . 7A) can be used to determine the instantaneous or beat-to-beat blood velocity V(t). The blood velocity can be combined with the cardiac output averaged over one or more data collection cycles to calculate instantaneous cardiac output. The process used is shown in the process depicted in FIG. 7B, identified as Process IV.
As seen therein, the initial temperature T31(t-) sensed at temperature sensor 30 as a function of time is smoothed, or filtered, in the manner as previously discussed above, and suitably measured and stored at an initial time to. A predetermined rise in temperature ~T3 of the temperature sensor itself is selected. Power is then supplied at time to to the temperature sensor 30 from a power source 30A connected thereto to cause its temperature T3(t) to rise by a predetermined amount.
Power may be supplied to the sensor in different ways according to the needs o~ the particular measurement and the relative simplicity or complexity of the required WO95/05115 ;i -; ; 2 ~ 6 93 74 PCT~S94/09064 circuitry, three such ways being depicted in FIGS. 8A, 8B, and 8C.
For example, in a first mode of operation (FIG. 8A), heating power may be supplied to the sensor in such a manner as to keep the final sensor temperature T3f (t) constant at an initial level ~T3 above the initial temperature T3i(to), i.e. T3f = T3i (to) + ~T3 even when the local blood temperature varies with time, as illustrated in FIG. 9A. Under such conditions, the sensor is maintained at a time-varying temperature increment ~T3 (t) above the instantaneous local blood temperature, Tb(t).
Alternatively, in a second mode of operation, power can be supplied to the sensor so as to continuously maintain the sensor at a fixed temperature increment above the time varying local blood temperature, as illustrated in FIG. 9B. Under such conditions, ~T3(t) =
~T3, a constant, and the sensor temperature varies according to T3f (t) = ~T3 + T3i(t).
A third mode of heating may also be convenient when the temperature sensors are temperature~sensitive resistors, or thermistors. Thus, when a thermistor is used, it may be more convenient to design an electrical heating circuit that maintains the sensor at a constant resistance increment above the resistance of the sensor that corresponds to the local blood temperature. If R is the corresponding resistance for a sensor temperature T, WO95/05115 ; ~ 2 1 6 93 7 4 PCT~S94/09064 ~
then these conditions are represented by ~R3(t) = ~R3, a constant, and the sensor resistance varies according to R3f(t) = ~R3 + R3i(t), as illustrated in FIG. 9C. The change in temperature ~T3(t) is then replaced by the change in resistance R3(t) in the ratio which is integrated over a cardiac cycle. Further details and exemplary apparatus for such modes of operation are presented and described in U.S. patent No. 4,059,982, issued to H.F. Bowman on November 29, 1977. With all three of the above approaches, power (P) is supplied to produce a temperature rise (~T) both of which are then related to-the instantaneous blood velocity and, hence, blood flow.
Techniques in which sensor heating power and temperature can be measured and used to provide more detailed information on cardiac output are described below. The technique involved can be applied to measure both instantaneous cardiac output, and the cardiac output for an individual cardiac cycle. Such detailed measurement information greatly enhances the diagnostic capability of a physician.
First, a method is described to measure instantaneous volumetric flow (which flow if measured at the location described above is the cardiac output). For each of the particular implementations described above, the power P3(t) applied to the temperature sensor 30 i9 - ~ :
~ WO95/05115 ;1~ r 2 1 69374 PCT~S94109064 controlled so as to maintain the final temperature of the sensor at a desired value T3f. The power applied to the temperature sensor 30 or, more generally, the ratio of the power applied to the sensor to the temperature increment, P3(t)/~T3(t), is directly correlated with the fluid and flow properties of the flowing liquid about the sensor.
For example, the relationship between required sensor power and local fluid velocity, V(t), is given by a correlation of the form:
P(t) = 4~kaAT(t) [1 + Cl Prn (2ap V(t) /~)m]
Where P(t) = instantaneous power to sensor k = thermal conductivity of fluid a = sensor radius ~T(t) = instantaneous temperature difference between heated sensor and unheated fluid temperature.
Cl = constant of calibration Pr = a non-dimensional "Prandtl" number which relates to the viscosity ~, heat capacity Cp and thermal conductivity k of a fluid.
n,m = power factors which are determined from ~ 35 experimental data p = fluid density ~ = viscosity V(t) = instantaneous fluid velocity WO95/05115 ~ 2 1 69374 PCT~S94/09064 ~
The fluid flow velocity in the vicinity of the sensor can be determined from the required sensor heating power. Volumetric flow in the vessel can then be determined with one further assumption for the distribution of the fluid flow within the vessel. For example, assuming a uniform velocity profile within the vessel, volumetric flow F3 is given by F = VA
where V is the fluid velocity in the vessel and A is the flow area. If the fluid flow area A is not previously known, it may be inferred from the measurement of average volumetric flow in the vessel. Such average volumetric flow can be determined, for example, by using the techniques of the invention already described above herein or by using other techniques for yielding comparable information. For example, if F is the average cardiac output, typically measured over several cardiac cycles, as described above, and V is the average fluid velocity, determined by calculating an average value for the instantaneous flow velocity over at least one cardiac cycle, then one such estimate for the average flow area A
is given by A = F/V
~ WO9S/05115 , 2 1 6 9 3 7 4 PCT~S94/09064 Therefore, given the sensor measured heating power, first the fluid velocity and then volumetric flow can be calculated at any desired instant in time, i.e., F3(t) V(E)A, yielding an instantaneous measure of volumetric flow, i.e., cardiac output.
In another embodiment, a method to measure cardiac output over a single cardiac cycle is described. As described above in different implementations, the power P3(t) applied to the temperature sensor 30 is controlled .so as to maintain the temperature of the sensor at a desired signal value T3f. The power applied to the temperature sensor 30, or more generally, as discussed above, the ratio of the power applied to the sensor to the temperature increment, i.e., P3(t)/~T3(t), is directly correlated with the properties of the fluid flow in the vicinity of the sensor.
Thus, the integrated value of the power to temperature ratio over a single cardiac cycle is directly correlated with, i.e., is proportional to the average cardiac output over the cardiac cycle, di~ 1 ~T3(t) or, alternatively expressed l r P3(t) dt F
cardia!c cycle 3~ ) WO95/05115 . ~ ~ 2 1 69374 PCT~S~1,0~4 ~
where T represents the period of the cardiac cycle and F
correspondingly represents cardiac output averaged over the cardiac cycle. Thus, the average cardiac output F
over an individual cardiac cycle then can be determined from the measured and integrated power and temperature signals from the sensor.
Furthermore, an explicit correlation for integrated power and average cardiac output over the cardiac cycle may be dispensed with if a simple qualitative indication of the change in cardiac output on a cardiac cycle-to-cardiac cycle basis is desired. To obtain such information, a given measurement of cardiac output is taken as associated with a corresponding measured value of the integrated sensor signal over a cardiac cycle.
The measurement of cardiac output could be obtained intermittently by the techniques described in this invention or other similar techniques. Since cardiac output is known to be correlated with the value of the integrated sensor signal over the cardiac cycle, any changes in the sensor signal indicate a corresponding change in cardiac output.
In certain situations, it may be desirable to compensate for temperature variations in the blood which is flowing past the sensor, as this may affect the value of F(t). A process for such compensation is depicted as Process IV in FIG. 7B wherein a temperature T2(t) is ~ WogS/OSllS . ~, ~ 2 ~ 6 93 7 4 PCT~S94/09064 sensed by a second sensor (which may be, for example, sensor 19 or sensor 20) at a location remote from sensor - 30 ~see FIG. 7A). For example, knowledge of the instantaneous blood temperature is required ~or the process in which the heated sensor is maintained at a constant increment above the local blood temperature. In this case, the temperature T2(t) is used as a proxy for the temperature T3(t) which would be measured in the absence of sensor heating.
In a further alternative embodiment, where only two sensors 19 and 20 are utilized (as shown in FIG. 2), sensor 20 can be used as the primary sensor when calculating instantaneous cardiac output (equivalent to sensor 30 in FIG. 7A) and sensor 19 can be used as the secondary temperature compensation sensor. In such an embodiment, the averaged cardiac output can be determined using sensors 19 and 20, as set forth in FIGS. 3-6 and the instantaneous cardiac output can then subsequently be determined using sensors 19 and 20, as set ~orth in FIGS.
7B and either FIGS. 8A, 8B or 8C, such average and instantaneous cardiac output determinations being made in sequence by the microprocessor to provide the cardiac information in both forms, as desired.
As mentioned above, when using the above described catheter, the various flow values which are determined in WO95/05115 ~ 2 1 6 ~ 3 7 4 PCT~S94/09064 ~
accordance with the processes as discussed above are proportional to flow but may not be equal to the actual flow values unless they are suitably calibrated since the correspondence between the calculated and actual values depends on the manner in which a particular catheter is constructed and used. A calibration constant for a particular catheter can be represented by the slope and intercept of a curve which relates the calculated flow and the actual flow, in accordance with the following relationship:
FaCtual = aFca1c. +b where, as illustrated in FIG. lO, "a" is the slope of a straight line 35 and "b" is the intercept thereof along the vertical axis. Curve 35 can be obtained by using a known catheter and known flow values therein to construct a curve 36. The best straight line fit is determined as line 35. The slope "a" and intercept "b" are thereby determined. Such determined values for "a" and "b" can be used with the calculated flow values in each case to determine the actual flow from the calculated flow.
While the above description discusses preferred embodiments of the invention, modifications thereof may occur to those in the art within the spirit and scope of the invention. Hence, the invention is not to be construed as limited to particular embodiments described, except as defined by the appended claims.
BLOOD FLOW AT LOW POWER
Introduction This invention relates generally to techniques for measuring blood flow in a body and, more particularly, to the use preferably of one or more temperature sensors for measuring thermal energy changes in the blood flowing through the heart and to the use o~ unique data processing techniques in response thereto for determining cardiac output.
Backqround of the Invention While the invention can be used generally to measure blood flow at various locations in a body, it is particularly useful in measuring blood flow in the heart so as to permit the measurement of cardiac output.
Many techniques for measuring cardiac output have been suggested in the art. Exemplary thermodilution techniques described in the technical and patent literature include:
"A Continuous Cardiac Output Computer Based On Thermodilution Principles", Normann et al., Annals of Biomedical Engineering, Vol. 17, 1989; "Thermodilution Cardiac Output Determination With A single Flow-Directed Catheter", Forrester, et al., American Heart Journal, Vol. 83, No. 3, 1972; "Understanding Techniques for Measuring Cardiac Output", Taylor, et al., Biomedical Instrumentation & Technology, May/June 1990; United States Patent No. 4,507,974 of M.L. Yelderman, issued April 2, 1985; United States Patent No. 4,785,823, of Eggers et al., issued on November 22, 1988; and United States Patent No. 5,000,190, of John H. Petre, issued on March 19, 1991.
WO95/05115 ~ ~ 2 1 6 9 3 7 4 PCT~S94/09064 ~
A principal limitation in the quanification of cardiac output is the existence of thermal fluctuations inherent in the bloodstream. Previous methods work with those fluctuations while observing the effects of an input signal to calculate cardiac output. The invention described herein uses a differential measurement technique to substantially eliminate the effect of the thermal fluctuations, permitting the use of a minimal thermal input signal, which allows frequent or continuous measurements.
It is desirable to obtain accurate cardiac output measurements in an effectively continuous manner, i.e., several times a minute, so that a diagnosis can be achieved more rapidly and so that rapid changes in a patient's condition can be monitored on a more continuous basis than is possible using current techniques.
Moreover, it is desirable to obtain instantaneous measurements of the cardiac output on a beat-to-beat basis to evaluate the relative changes which occur from beat to beat, as well as to determine the presence of regurgitation.
Brief Summary of the Invention In accordance with general principal of the invention, blood flow and/or cardiac output is determined rapidly, using a technique by which an indicator substance, or agent, is introduced into the bloodstream between a pair of detectors. The detectors are sensitive to a parameter functionally related to the concentration or magnitude in the bloodstream of the selected indicator agent. The detectors are positioned apart by a distance functionally sufficient to allow a measurement to be made of the differential value of the selected parameter as it WO95105115 . 1~ ~ ` 2 ~ 69374 PCT~S94/09064 exists from time-to-time between the two detectors. The indicator agent, for example, may be a substance to change the pH of the blood, a fluid bolus carrying thermal energy, or a substance to change a selected characteristic of the blood, or the direct introduction of thermal energy, or the like.
A determination is made of the difference in the values of the selected blood parameter as it exists at the two detectors, prior to the introduction of indicator agent (i.e., the first differential value). The selected indicator agent is then introduced in a predetermined magnitude. Then again a determination is made of the difference in values of the selected parameter as it exists at the locations of the two detectors (i.e., the second differential value).
Blood flow or cardiac output, depending on the specific location of the detectors, can then be determined as a function of the difference between the first differential value and the second differential value. Because the ultimate measurement of blood flow or cardiac output is based on the difference of the differences, the system operates effectively with the introduction of the indicator agent in a very low magnitude. In turn, this allows measurements to be made rapidly 80 that effectively continuous measurements are obtained.
In accordance with a preferred embodiment of the invention, for example, cardiac output can be determined rapidly and with low levels of thermal energy input. To achieve such operation, in a preferred embodiment, the technique of the invention uses a pair of temperature sensors positioned at two selected locations within a catheter which has been inserted into the path of the 2 1 6 q 3 7 4 PCT~S~ 3~61 ~
blood flowing through the heart of a living body. The sensors detect the temperature difference between the two locations. Depending on the location of the temperature sensors in the circulatory system, the measured temperature difference varies over time. It has been observed that when the temperature sensors are placed within the heart, e.g., so that one sensor lies in the vena cava, for example, and the second in the right ventricle or pulmonary artery, the temperature difference varies in a synchronous manner with the respiratory cycle.
Thus, in the preferred embodiment of the invention the temperature difference over at least one respiratory cycle is measured and averaged to provide an average temperature difference. The averaging, or integrating, action effectively eliminates, as a confounding factor in the determination of cardiac output, the effect of instantaneous blood temperature fluctuations, such as cyclical, respiratory-induced fluctuations.
To make such determinations, an average temperature difference is first calculated over a time period of at least one respiratory cycle in which no thermal energy is introduced into the blood flow path. Thermal energy of a predetermined and relatively low magnitude is then introduced into the blood flow path to produce a heating action therein at a location between the two temperature sensors. Once the temperature rise induced by the heating stabilizes, the average temperature difference between the two locations is again calculated from temperature difference measurements over a time period of at least one respiratory cycle at the higher temperature level. The difference between the average temperature differences which occurs when the thermal energy is turned on, referred to as the rising temperature change, ~ W095/05115 - ,~ s( - 2 1 6 9 3 7~ PCT~S94/ogO64 is determined. The difference between the average temperature differences which occurs when the thermal energy is turned off, referred to as the falling temperature change, is similarly determined. The cardiac output is calculated as a function of the thermal energy input and the rising and falling temperature changes.
Because a relatively low level of thermal energy is used in making measurements, the overall sequence of determinations can be safely repeated multiple times per minute, for example, so that an effectively continuous, or quasi-continuous, determination of cardiac output is obtained.
In accordance with a further embodiment of the invention, a temperature sensor that also acts as a source of thermal energy, e.g., a thermistor, is positioned at a third location in the cardiac blood flow path. Power is supplied to the sensor sufficient to elevate the temperature of the sensor from a first temperature level to a second temperature level. In one embodiment of the invention, the temperature of the sensor is changed from the first to the second level and i9 maintained constant at said second level by varying the power that i8 supplied thereto. Such varying power is proportional to the instantaneous flow velocity and, hence, assuming a constant flow area, is proportional to the instantaneous cardiac output. Measurement of the sensor heating power and the temperature increment at the sensor can thus be used to continuously effect a determination of the instantaneous cardiac output.
30 Further, for example, when the sensor is placed downstream at the outlet of one of the heart chambers, the variation in flow output over the cardiac cycle can be analyzed to provide an indication of the regurgitation characteristics of the heart outlet valve over the cardiac cycle. Moreover, such instantaneous cardiac WO9S/05115 2 1 6 9 3 7 4 PCT~S~09^C4 ~
output determination can be further refined to compensate for fluctuations in the temperature of the blood flowing through the heart by measuring the instantaneous temperature of the blood with another temperature sensor at a nearby location and appropriately taking into account such temperature variations when determining the cardiac output.
In another application, both the continuous cardiac output determinations and the instantaneous cardiac output determinations, as described above, can be combined. Thus, three temperature sensors and a source of thermal energy can all be used in combination to simultaneously provide an accurate and effectively continuous determination of time-averaged cardiac output, and a determination of instantaneous cardiac output at .
each instant of the cardiac cycle. In still another application, two temperature sensors and a source of thermal energy can be used in an appropriate sequence to provide the averaged cardiac output determination and the instantaneous cardiac output determination.
Description of the Invention The invention can be described in more detail with the help of the accompanying drawings wherein FIG. 1 shows a simplified diagrammatic view of a human heart;
FIG. 2 shows a simplified diagrammatic view of.a catheter useful in the invention;
FIG. 3 shows a flow chart depicting steps in a process used in the invention;
FIG. 3A shows a smooth temperature difference curve obtained in the process depicted in FIG. 3;
FIG. 4 shows a flow chart depicting steps in another process used in the invention;
FIG. 5 shows a flow chart depicting steps in still another process of the invention;
~ W095/05115 ~ 2 ~ 6q37~ PCT~S9410906~
FIG. 6 shows a graph depicting a temperature/time relation used in the invention;
FIG. 7A shows a simplified diagrammatic view of another catheter used in the invention;
FIG. 7B shows a ~low chart depicting steps in still another process of the invention;
FIGS. 8A, 8B, and 8C show flow charts depicting still other processes of the invention;
FIGS. 9A, 9B, and 9C show graphs of parameter relationships used in the invention; and FIG. 10 shows a graph useful ~or calibrating the flow values for a catheter used in the invention.
As can be seen in FIG. l, which represents a human heart 10 i~n a much simplified diagrammatic form, a flexible catheter 11 is inserted through the veins into the right atrium, or auricle, 12 of the heart and, thence, through the right ventricle 13 until the end of the catheter resides in or near the exit, or pulmonary, artery 14 which leads to the lungs. As is well known, blood flows (as represented by the arrows) from the input vein 15, i.e., the vena cava, into the right atrium and right ventricle and thence outwardly to the lungs and subsequently returns from the lungs into the left atrium 16, through the left ventricle 17 and thence outwardly into the aorta 18.
In accordance with the embodiment of the invention, shown with reference to FIG. 1, temperature sensors, e.g., thermistors, are carried by the catheter so that, 30 when inserted as shown in FIG. l, a first sensor 19 is positioned at a location within the vena cava 15 or right atrium 12 and a second sensor 20 is positioned at a location in or near the pulmonary artery 14.
For simplicity, the flexible catheter 11 is depicted in FIG. 2 in an extended condition with temperature WO95/05115 2 ~ 6 93 74 PCT~S94/0906~ ~
sensors 19 and 20 at two different locations for measuring temperatures T1 and T2, respectively. A power source 21 of thermal energy which is borne, or carried, by the catheter 11 is positioned in the right atrium at a location between sensors 19 and 20. In a particular embodiment, the catheter-borne source is, for example, a coil of resistive wire placed on or embedded in the surface of catheter 11, to which an AC or a DC voltage (not shown) at a controllable level is supplied so as to generate thermal energy, i.e. heat. The magnitude of the thermal energy can be suitably controlled to insert a predetermined amount of thermal energy at a selected time, which thermal energy is transferred to the blood flowing through the heart so as to raise its temperature.
The energy source is positioned at a sufficient distance from the sensor 19 that the latter is effectively thermally isolated from the site of the thermal energy source.
While the locations of the sensors 19 and 20 and the energy source 21 can be as shown in FIG. 1, alternative locations can also be used. Thus, the sensor 19 can be positioned in the vena cava 15, while the energy source 21 is located in the right atrium 12 and the sensor 20 in either the right atrium or the right ventricle.
Moreover, if sensor 19 is positioned in the vena cava 15, the entire energy source 21, which is normally elongated, need not be located in the right atrium and can have a portion thereof in the vena cava and a portion thereof in the right atrium. Such source should preferably be at least partially located in the right atrium. Further, sensor 20 may be positioned in the right ventricle near the pulmonary artery 14 or may be located in the pulmonary artery itself at or near the right ventricle.
The temperatures T1 and T2 at locations 19 and 20 upstream and downstream, respectively, from the thermal ~ WO95/05115 ~ 2 1 6937~ PCT~S94/09064 energy source 21 are monitored and processed appropriately by a digital microprocessor. In accordance with the invention, the instantaneous temperatures are obtained as the outputs T1(t) and T2(t) of the temperature sensors l9 and 20, respectively. The outputs are connected to a differential amplifier to generate an analog signal which is proportional to the temperature difference ~T(t)=T1(t)-T2(t) between them.
The temperature difference signal aT(t) is digitized and sampled at selected time intervals by an analog-to-digital/sampling circuit. The digitized sampled temperature difference values and the known thermal energy values are supplied to a digital microprocessor which then suitably proce8ses the data to provide the desired cardiac output information. The processing stages used in the host microprocessor are implemented by suitable programming of the microprocessor and are discussed below with the help of FIGS. 3-6.
The source 21 of thermal energy is alternately turned on and off. If it is assumed that thermal stability is reached after each change and that there is a substantially constant rate of blood flow, a stable temperature difference can be measured in each case. The quantity of blood flowing past the thermal energy source, i.e., the cardiac output, can be derived from such temperature difference measurements. However, such derivation is complicated by two factors which may affect the measurement of blood flow. First, the rate of blood flow through the heart is not substantially constant but surges with each heart contraction. Second, the temperature of the blood flowing through the heart is not constant but varies with each respiratory (breathing) cycle. In a preferred embodiment, the processing of the data ta~es such factors into account, as discussed below.
WO95/05115 ~ 2 1 6 9 3 7 4 PCT~S94/09064 ~
The process for determining cardiac output is performed in a microprocessor 2 which in a first embodiment is programmed to respond to the temperatures sensed at Tl and T2 and to perform the steps depicted in accordance with the flow charts shown in FIGS. 3-5. From a knowledge of such flow charts, it would be well within the skill of those in the art to appropriately program any suitable and known digital microprocessor, such as a personal computer, to perform the steps shown.
FIG. 3 depicts a basic process, identified as Process I, which is used in the overall processing of temperature data for determining cardiac output, as subsequently depicted in FIGS. 4 and 5. In the basic process shown in FIG. 3, a temperature difference as a function of time ~T(t) is determined by a differential amplifier which responds to T1(t) and T2(t). Such differences may be effectively smoothed, or filtered, to produce a smooth temperature difference curve, as shown in FIG. 3A, which varies as a function of time in a cyclic manner which depends principally on the respiratory cycle of the person whose cardiac output is being determined.
The periods ~ 2. . . ~n for each respiratory cycle are determined over n cycles. A characteristic of the temperature difference at each cycle is determined. For example, such characteristic preferably is the averaged temperature difference during each cycle (~T7l, ~ T72 ~ T72 Q T7n ) (Alternatively, for example, the peak temperature differences may be the determined characteristic.) These averaged temperature differences (~T7n) are added for the n cycles involved and are divided by n to determine an averaged temperature difference per cycle (~T7n)~ The use of Process I is depicted in the process steps shown -~ W095/OS115 ~ 2 i 6 ~ 3 7 4 PCT~S94/09064 in FIG. 4, identified as Process II.
o As seen therein, the steps of Process I are first performed when the source 21 of thermal energy (i.e., a heater) is turned off and the average AToff value per cycle is determined and suitably stored. The sampling time at which such determination is made is depicted in FIG. 6 as the sample time period Sl.
The heater 21 is then turned on for a specific time period to supply a known amount of power P to the blood flowing through the heart and, accordingly, the temperature of the blood flowing past the heater rises and the temperature difference ~T(t) rises over a transition, or delay, rise time period, t~1, shown in FIG. 6 and designated as D1, after which the temperature difference generally stabilizes over a second sample time period S2. As seen in FIG. 4, after the heater 21 is turned on and the temperature has stabilized, Process I
is performed, again over n cycles, e.g., over the time period S2, and the averaged temperature difference ~Ton is determined with the heater turned on and is suitably stored. The heater is then turned off and the temperature falls over a transition, or delay, fall time period tFl shown in FIG. 6 and designated as D2, generally to its former value.
Cardiac output is calculated using the averaged temperature differences when the energy is off and the averaged temperature differences when the energy is on, by the relationship:
F = P
Cp ( ~ Ton ~ ~ Tof f ) where: F = Flow P = Power Cp = heat capacitance ~Ton = average temperature for power on off = average temperature for power off WO95/05115 ~ .' =' 2 1 6 9 3 7 4 PCT~S94/09064 ~
As seen in FIG . 5, the steps of Process II are repeated indefinitely for N data collection cycles, a data collection cycle being designated as including the time periods Sl, Dl, S2, and D2, as shown in FIG. 6. For each data collection cycle the rise tir.le temperature difference ATR between the averaged temperature difference ~Ton at S2 and the averaged temperature difference QToff at Sl and the fall time temperature difference ATp between the averaged temperature difference ATOffat Sl and the averaged temperature difference QTon at S2 are determined.
The flow, FR is calculated for each data collection cycle from the known amount of power P introduced into the blood flow stream by the energy source, or heater 21, from the known heat capacitance of blood, Cp and from the difference in the averaged temperature differences ~Ton and QToff/ which occurs over the data collection cycle Sl + Dl + S2 in accordance with the following relationship:
FR = P
Cp (~Ton ~ A Toff~
In a similar manner, the flow FP is calculated from P, Cp and the difference in the averaged temperature differences ATon and ~Toff which occurs over the later portion of the data collection cycle Sl + Dl + S2 in ~ WO95/05115 . ~- Z1~ ~3~4 PcT~s~4,a~64 accordance with the following relationship:
.
FF = P
Cp ( 1~ Ton ~ ~ Tof f ) FR and FF can be averaged to obtain the averaged flow (F) over one data collection cycle as shown in FIG. 6.
F = F~ + FF
A suitable calibration constant can be used to adjust the values Of FR~ FF and F.
Accordingly, by using two temperature sensors l9 and 20, cardiac output can be determined several times a minute in accordance with FIGS . 3-6, yielding an effectively continuous cardiac output value. Because such measurements can be made using relatively low power levels, the danger that the heart may be damaged by the introduction of thermal energy is substantially eliminated.
It will be apparent that the foregoing technique, which has been described in connection with the direct introduction of heat as an indicator agent and the measurement of temperature, can readily be performed by those skilled in the art by using indicator agents which affect the pH of the blood or change other blood parameters.
In some situations it may be desirable to provide WO95/05115 ~ 2 1 6 9 3 7 4 PCT~S~/0~
more frequent indications of cardiac output, such as, for example, the instantaneous cardiac output or the cardiac output averaged over each individual cardiac cycle (i.e.
each heart beat). Such information can be provided using the further embodiments of the invention discussed below with reference to FIGS. 7-8. A single temperature sensor 30 at a location near the distal end of the catheter 31 (as shown in FIG . 7A) can be used to determine the instantaneous or beat-to-beat blood velocity V(t). The blood velocity can be combined with the cardiac output averaged over one or more data collection cycles to calculate instantaneous cardiac output. The process used is shown in the process depicted in FIG. 7B, identified as Process IV.
As seen therein, the initial temperature T31(t-) sensed at temperature sensor 30 as a function of time is smoothed, or filtered, in the manner as previously discussed above, and suitably measured and stored at an initial time to. A predetermined rise in temperature ~T3 of the temperature sensor itself is selected. Power is then supplied at time to to the temperature sensor 30 from a power source 30A connected thereto to cause its temperature T3(t) to rise by a predetermined amount.
Power may be supplied to the sensor in different ways according to the needs o~ the particular measurement and the relative simplicity or complexity of the required WO95/05115 ;i -; ; 2 ~ 6 93 74 PCT~S94/09064 circuitry, three such ways being depicted in FIGS. 8A, 8B, and 8C.
For example, in a first mode of operation (FIG. 8A), heating power may be supplied to the sensor in such a manner as to keep the final sensor temperature T3f (t) constant at an initial level ~T3 above the initial temperature T3i(to), i.e. T3f = T3i (to) + ~T3 even when the local blood temperature varies with time, as illustrated in FIG. 9A. Under such conditions, the sensor is maintained at a time-varying temperature increment ~T3 (t) above the instantaneous local blood temperature, Tb(t).
Alternatively, in a second mode of operation, power can be supplied to the sensor so as to continuously maintain the sensor at a fixed temperature increment above the time varying local blood temperature, as illustrated in FIG. 9B. Under such conditions, ~T3(t) =
~T3, a constant, and the sensor temperature varies according to T3f (t) = ~T3 + T3i(t).
A third mode of heating may also be convenient when the temperature sensors are temperature~sensitive resistors, or thermistors. Thus, when a thermistor is used, it may be more convenient to design an electrical heating circuit that maintains the sensor at a constant resistance increment above the resistance of the sensor that corresponds to the local blood temperature. If R is the corresponding resistance for a sensor temperature T, WO95/05115 ; ~ 2 1 6 93 7 4 PCT~S94/09064 ~
then these conditions are represented by ~R3(t) = ~R3, a constant, and the sensor resistance varies according to R3f(t) = ~R3 + R3i(t), as illustrated in FIG. 9C. The change in temperature ~T3(t) is then replaced by the change in resistance R3(t) in the ratio which is integrated over a cardiac cycle. Further details and exemplary apparatus for such modes of operation are presented and described in U.S. patent No. 4,059,982, issued to H.F. Bowman on November 29, 1977. With all three of the above approaches, power (P) is supplied to produce a temperature rise (~T) both of which are then related to-the instantaneous blood velocity and, hence, blood flow.
Techniques in which sensor heating power and temperature can be measured and used to provide more detailed information on cardiac output are described below. The technique involved can be applied to measure both instantaneous cardiac output, and the cardiac output for an individual cardiac cycle. Such detailed measurement information greatly enhances the diagnostic capability of a physician.
First, a method is described to measure instantaneous volumetric flow (which flow if measured at the location described above is the cardiac output). For each of the particular implementations described above, the power P3(t) applied to the temperature sensor 30 i9 - ~ :
~ WO95/05115 ;1~ r 2 1 69374 PCT~S94109064 controlled so as to maintain the final temperature of the sensor at a desired value T3f. The power applied to the temperature sensor 30 or, more generally, the ratio of the power applied to the sensor to the temperature increment, P3(t)/~T3(t), is directly correlated with the fluid and flow properties of the flowing liquid about the sensor.
For example, the relationship between required sensor power and local fluid velocity, V(t), is given by a correlation of the form:
P(t) = 4~kaAT(t) [1 + Cl Prn (2ap V(t) /~)m]
Where P(t) = instantaneous power to sensor k = thermal conductivity of fluid a = sensor radius ~T(t) = instantaneous temperature difference between heated sensor and unheated fluid temperature.
Cl = constant of calibration Pr = a non-dimensional "Prandtl" number which relates to the viscosity ~, heat capacity Cp and thermal conductivity k of a fluid.
n,m = power factors which are determined from ~ 35 experimental data p = fluid density ~ = viscosity V(t) = instantaneous fluid velocity WO95/05115 ~ 2 1 69374 PCT~S94/09064 ~
The fluid flow velocity in the vicinity of the sensor can be determined from the required sensor heating power. Volumetric flow in the vessel can then be determined with one further assumption for the distribution of the fluid flow within the vessel. For example, assuming a uniform velocity profile within the vessel, volumetric flow F3 is given by F = VA
where V is the fluid velocity in the vessel and A is the flow area. If the fluid flow area A is not previously known, it may be inferred from the measurement of average volumetric flow in the vessel. Such average volumetric flow can be determined, for example, by using the techniques of the invention already described above herein or by using other techniques for yielding comparable information. For example, if F is the average cardiac output, typically measured over several cardiac cycles, as described above, and V is the average fluid velocity, determined by calculating an average value for the instantaneous flow velocity over at least one cardiac cycle, then one such estimate for the average flow area A
is given by A = F/V
~ WO9S/05115 , 2 1 6 9 3 7 4 PCT~S94/09064 Therefore, given the sensor measured heating power, first the fluid velocity and then volumetric flow can be calculated at any desired instant in time, i.e., F3(t) V(E)A, yielding an instantaneous measure of volumetric flow, i.e., cardiac output.
In another embodiment, a method to measure cardiac output over a single cardiac cycle is described. As described above in different implementations, the power P3(t) applied to the temperature sensor 30 is controlled .so as to maintain the temperature of the sensor at a desired signal value T3f. The power applied to the temperature sensor 30, or more generally, as discussed above, the ratio of the power applied to the sensor to the temperature increment, i.e., P3(t)/~T3(t), is directly correlated with the properties of the fluid flow in the vicinity of the sensor.
Thus, the integrated value of the power to temperature ratio over a single cardiac cycle is directly correlated with, i.e., is proportional to the average cardiac output over the cardiac cycle, di~ 1 ~T3(t) or, alternatively expressed l r P3(t) dt F
cardia!c cycle 3~ ) WO95/05115 . ~ ~ 2 1 69374 PCT~S~1,0~4 ~
where T represents the period of the cardiac cycle and F
correspondingly represents cardiac output averaged over the cardiac cycle. Thus, the average cardiac output F
over an individual cardiac cycle then can be determined from the measured and integrated power and temperature signals from the sensor.
Furthermore, an explicit correlation for integrated power and average cardiac output over the cardiac cycle may be dispensed with if a simple qualitative indication of the change in cardiac output on a cardiac cycle-to-cardiac cycle basis is desired. To obtain such information, a given measurement of cardiac output is taken as associated with a corresponding measured value of the integrated sensor signal over a cardiac cycle.
The measurement of cardiac output could be obtained intermittently by the techniques described in this invention or other similar techniques. Since cardiac output is known to be correlated with the value of the integrated sensor signal over the cardiac cycle, any changes in the sensor signal indicate a corresponding change in cardiac output.
In certain situations, it may be desirable to compensate for temperature variations in the blood which is flowing past the sensor, as this may affect the value of F(t). A process for such compensation is depicted as Process IV in FIG. 7B wherein a temperature T2(t) is ~ WogS/OSllS . ~, ~ 2 ~ 6 93 7 4 PCT~S94/09064 sensed by a second sensor (which may be, for example, sensor 19 or sensor 20) at a location remote from sensor - 30 ~see FIG. 7A). For example, knowledge of the instantaneous blood temperature is required ~or the process in which the heated sensor is maintained at a constant increment above the local blood temperature. In this case, the temperature T2(t) is used as a proxy for the temperature T3(t) which would be measured in the absence of sensor heating.
In a further alternative embodiment, where only two sensors 19 and 20 are utilized (as shown in FIG. 2), sensor 20 can be used as the primary sensor when calculating instantaneous cardiac output (equivalent to sensor 30 in FIG. 7A) and sensor 19 can be used as the secondary temperature compensation sensor. In such an embodiment, the averaged cardiac output can be determined using sensors 19 and 20, as set forth in FIGS. 3-6 and the instantaneous cardiac output can then subsequently be determined using sensors 19 and 20, as set ~orth in FIGS.
7B and either FIGS. 8A, 8B or 8C, such average and instantaneous cardiac output determinations being made in sequence by the microprocessor to provide the cardiac information in both forms, as desired.
As mentioned above, when using the above described catheter, the various flow values which are determined in WO95/05115 ~ 2 1 6 ~ 3 7 4 PCT~S94/09064 ~
accordance with the processes as discussed above are proportional to flow but may not be equal to the actual flow values unless they are suitably calibrated since the correspondence between the calculated and actual values depends on the manner in which a particular catheter is constructed and used. A calibration constant for a particular catheter can be represented by the slope and intercept of a curve which relates the calculated flow and the actual flow, in accordance with the following relationship:
FaCtual = aFca1c. +b where, as illustrated in FIG. lO, "a" is the slope of a straight line 35 and "b" is the intercept thereof along the vertical axis. Curve 35 can be obtained by using a known catheter and known flow values therein to construct a curve 36. The best straight line fit is determined as line 35. The slope "a" and intercept "b" are thereby determined. Such determined values for "a" and "b" can be used with the calculated flow values in each case to determine the actual flow from the calculated flow.
While the above description discusses preferred embodiments of the invention, modifications thereof may occur to those in the art within the spirit and scope of the invention. Hence, the invention is not to be construed as limited to particular embodiments described, except as defined by the appended claims.
Claims (46)
1. A method for determining blood flow in a living body comprising the steps of:
changing the thermal energy level by a predetermined amount at a site in a blood flow path of said living body;
detecting the temperatures at a first location upstream of said site and a second location downstream of said site;
determining the temperature difference between said first and second locations at one energy level;
determining the temperature difference between said first and second locations at a changed energy level; and calculating blood flow as a function of the change in energy level and of the temperature differences measured prior to and following the change in energy level.
changing the thermal energy level by a predetermined amount at a site in a blood flow path of said living body;
detecting the temperatures at a first location upstream of said site and a second location downstream of said site;
determining the temperature difference between said first and second locations at one energy level;
determining the temperature difference between said first and second locations at a changed energy level; and calculating blood flow as a function of the change in energy level and of the temperature differences measured prior to and following the change in energy level.
2. A method in accordance with claim 1 wherein said first location is substantially thermally isolated from thermal energy changes occurring at said site.
3. A method in accordance with claim 2 wherein said blood flow path includes at least a portion of the heart of a living body and the blood flow represents cardiac output.
4. A method in accordance with claim 3 wherein said site is at least partially in the right atrium of the heart, said first location is in the vena cava of the heart, and the second location is in or near the pulmonary artery of the heart.
5. A method in accordance with claim 3 wherein said site is at least partially the right atrium of the heart, said first location is in the vena cava of the heart, and the second location is in or near the right ventricle of the heart.
6. A method in accordance with claims 1, 2, 3,-4, or 5 wherein said thermal energy is changed by applying thermal energy from a catheter borne thermal energy source which is positioned in the blood flow path and wherein temperatures are detected by temperature sensors in said catheter at said first and second locations, respectively.
7. A method for determining blood flow in a living body comprising the steps of:
changing the thermal energy at a site in a blood flow path of said living body, from a first level to a second level;
determining a characteristic of the temperature difference between the temperatures in said blood flow path at a first location upstream of said site and at a second location downstream of said site at said first level;
determining a characteristic of the temperature difference between the temperatures in said blood flow path at the first location and at the second location at said second level;
determining the difference between the characteristics of said temperature differences;
calculating blood flow as a function of the difference between the characteristics of said temperature differences and the change in thermal energy from said first level to said second level.
changing the thermal energy at a site in a blood flow path of said living body, from a first level to a second level;
determining a characteristic of the temperature difference between the temperatures in said blood flow path at a first location upstream of said site and at a second location downstream of said site at said first level;
determining a characteristic of the temperature difference between the temperatures in said blood flow path at the first location and at the second location at said second level;
determining the difference between the characteristics of said temperature differences;
calculating blood flow as a function of the difference between the characteristics of said temperature differences and the change in thermal energy from said first level to said second level.
8. A method in accordance with claim 7 wherein each of said characteristics is determined over a specified time period.
9. A method in accordance with claim 8 wherein each of said characteristics is the average of the temperature differences.
10. A method in accordance with claim 8 wherein the specified time period is at least one respiratory cycle.
11. A method in accordance with claim 7, 8, 9, or 10 wherein said blood flow path includes at least a portion of the heart of a living body where the blood flow represents cardiac output.
12. A method in accordance with claim 11 wherein said site is at least partially in the right atrium of the heart, said first location is in the vena cava of the heart, and the second location is in or near the pulmonary artery of the heart.
13. A method in accordance with claim 11 wherein said site is at least partially in the right atrium of the heart, said first location is in the vena cava of the heart, and the second location is in or near the right ventricle of the heart.
14. A method in accordance with claim 7 wherein each of said characteristic temperature difference determining steps includes the steps of:
detecting temperatures at said first location upstream of said site and at said second location downstream of said site during each cycle of one or more respiratory cycles, when said thermal energy is at said first level and when said thermal energy is at said second level;
determining temperature differences between said temperatures for each respiratory cycle, when said thermal energy is at said first level and when said thermal energy is at said second level; and determining the average of said temperature differences over multiple respiratory cycles.
detecting temperatures at said first location upstream of said site and at said second location downstream of said site during each cycle of one or more respiratory cycles, when said thermal energy is at said first level and when said thermal energy is at said second level;
determining temperature differences between said temperatures for each respiratory cycle, when said thermal energy is at said first level and when said thermal energy is at said second level; and determining the average of said temperature differences over multiple respiratory cycles.
15. A method in accordance with claim 7 wherein:
said thermal energy level changing step includes applying a predetermined amount of power P at said site to change said thermal energy level from said first level to said second level and said cardiac output calculating step includes the steps of calculating a rise-time cardiac output component as a function of the power P, the calculated average rise-time temperature difference, and the heat capacity of blood;
calculating a fall-time cardiac output component as a function of the power P, the calculated fall-time difference, and the heat capacity of blood; and calculating the cardiac output as a function of the rise-time and fall-time cardiac components.
said thermal energy level changing step includes applying a predetermined amount of power P at said site to change said thermal energy level from said first level to said second level and said cardiac output calculating step includes the steps of calculating a rise-time cardiac output component as a function of the power P, the calculated average rise-time temperature difference, and the heat capacity of blood;
calculating a fall-time cardiac output component as a function of the power P, the calculated fall-time difference, and the heat capacity of blood; and calculating the cardiac output as a function of the rise-time and fall-time cardiac components.
16. A method in accordance with claims 7, 8, 9, 10 and 15 wherein the steps therein can be successively repeated a plurality of times to provide repeated calculations of the cardiac output to provide an effectively continuous calculation thereof.
17. A method in accordance with claim 16 wherein the repeated calculations of the cardiac output can be averaged over a selected number of said plurality of times to provide an averaged cardiac output over said selected number of times.
18. A system for determining blood flow in a living body comprising:
means for changing the thermal energy level by a predetermined amount at a site in a blood flow path of said living body;
means for detecting the temperatures at a first location upstream of said site and a second location downstream of said site;
means for determining the temperature difference between said first and second locations at one energy level;
means for determining the temperature difference between said first and second locations at a changed energy level; and means for calculating blood flow as a function of the change in energy level and of the temperature differences measured prior to and following the change in energy level.
means for changing the thermal energy level by a predetermined amount at a site in a blood flow path of said living body;
means for detecting the temperatures at a first location upstream of said site and a second location downstream of said site;
means for determining the temperature difference between said first and second locations at one energy level;
means for determining the temperature difference between said first and second locations at a changed energy level; and means for calculating blood flow as a function of the change in energy level and of the temperature differences measured prior to and following the change in energy level.
19. A system in accordance with claim 18 wherein said first location is substantially thermally isolated from thermal energy changes occurring at said site.
20. A system in accordance with claim 19 wherein said blood flow path includes at least a portion of the heart of a living body and the blood flow represents cardiac output.
21. A system in accordance with claim 20 wherein said site is at least partially in the right atrium of the heart, said first location is in the vena cava of the heart, and the second location is in or near the pulmonary artery of the heart.
22. A system in accordance with claim 20 wherein said site is at least partially the right atrium of the heart, said first location is in the vena cava of the heart, and the second location is in or near the right ventricle of the heart.
23. A system in accordance with claims 18, 19, 20, 21 or 22 wherein said system includes a catheter, said thermal energy changing means is a thermal energy source carried by said catheter and positioned in the blood flow path, and said temperature detecting means are temperature sensors positioned in said catheter so that when said catheter is positioned in said blood flow path said thermal energy source is at said site and said temperature sensors are at said first and second locations, respectively.
24. A system for determining blood flow in a living body comprising:
means for changing the thermal energy at a site in a blood flow path of said living body, from a first level to a second level;
means for determining a characteristic of the temperature difference between the temperatures in said blood flow path at a first location upstream of said site and at a second location downstream of said site at said first level;
means for determining a characteristic of the temperature difference between the temperatures in said blood flow path at the first location and at the second location at said second level;
means for determining the difference between the characteristics of said temperature differences;
means for calculating blood flow as a function of the difference between the characteristics of said temperature differences and the change in thermal energy from said first level to said second level.
means for changing the thermal energy at a site in a blood flow path of said living body, from a first level to a second level;
means for determining a characteristic of the temperature difference between the temperatures in said blood flow path at a first location upstream of said site and at a second location downstream of said site at said first level;
means for determining a characteristic of the temperature difference between the temperatures in said blood flow path at the first location and at the second location at said second level;
means for determining the difference between the characteristics of said temperature differences;
means for calculating blood flow as a function of the difference between the characteristics of said temperature differences and the change in thermal energy from said first level to said second level.
25. A system in accordance with claim 24 wherein said characteristic determining means determines said characteristics over a specified time period.
26. A system in accordance with claim 25 wherein each of said characteristics is the average of the temperature differences.
27. A system in accordance with claim 25 wherein the specified time period is at least one respiratory cycle.
28. A system in accordance with claims 24, 25, 26 or 27 wherein said blood flow path includes at least a portion of the heart of a living body where the blood flow represents cardiac output.
29. A system in accordance with claim 20 wherein said site is at least partially in the right atrium of the heart, said first location is in the vena cava of the heart, and the second location is in or near the pulmonary artery of the heart.
30. A system in accordance with claim 28 wherein said site is at least partially in the right atrium of the heart, said first location is in the vena cava of the heart, and the second location is in or near the right ventricle of the heart.
31. A system in accordance with claim 24 wherein each of said characteristic temperature difference determining means includes:
means for detecting temperatures at said first location upstream of said site and at said second location downstream of said site during each cycle of one or more respiratory cycles, when said thermal energy is at said first level and when said thermal energy is at said second level;
means for determining temperature differences between said temperatures for each respiratory cycle, when said thermal energy is at said first level and when said thermal energy is at said second level; and means for determining the average of said temperature differences over multiple respiratory cycles.
means for detecting temperatures at said first location upstream of said site and at said second location downstream of said site during each cycle of one or more respiratory cycles, when said thermal energy is at said first level and when said thermal energy is at said second level;
means for determining temperature differences between said temperatures for each respiratory cycle, when said thermal energy is at said first level and when said thermal energy is at said second level; and means for determining the average of said temperature differences over multiple respiratory cycles.
32. A method in accordance with claim 24 wherein:
said thermal energy level changing means includes means for applying a predetermined amount of power P
at said site to change said thermal energy level from.
said first level to said second level; and said cardiac output calculating means includes means for calculating a rise-time cardiac output component as a function of the power P, the calculated average rise-time temperature difference, and the heat capacity of blood;
means for calculating a fall-time cardiac output component as a function of the power P, the calculated fall-time difference, and the heat capacity of blood; and means for calculating the cardiac output as a function of the rise-time and fall-time cardiac components.
said thermal energy level changing means includes means for applying a predetermined amount of power P
at said site to change said thermal energy level from.
said first level to said second level; and said cardiac output calculating means includes means for calculating a rise-time cardiac output component as a function of the power P, the calculated average rise-time temperature difference, and the heat capacity of blood;
means for calculating a fall-time cardiac output component as a function of the power P, the calculated fall-time difference, and the heat capacity of blood; and means for calculating the cardiac output as a function of the rise-time and fall-time cardiac components.
33. A system in accordance with claims 24, 25, 26 or 27 wherein said system calculates the cardiac output repeatedly to provide an effectively continuous calculation thereof.
34. A system in accordance with claim 33 and further including means responsive to the repeated calculations of the cardiac output can for averaging said calculation over a selected number of repeated calculations to provide an averaged cardiac output over said selected number of calculations.
35. A method for determining blood flow in a living body comprising the steps of:
changing the value of the selected parameter of blood at a site in a blood flow path of said living body, from a first level to a second level;
determining the difference between the value of the selected parameter in said blood flow path at a first location upstream of said site and at a second location downstream of said site at said first level;
determining the difference between the value of the selected parameter in said blood flow path at the first location and at the second location at said second level;
determining the difference between the value obtained in the first said difference determining steps and the value obtained in the second said difference determining steps; and calculating blood flow as a function of the difference between the values obtained in said first and second difference determining steps and the change in the value of the selected parameter from said first level to said second level.
changing the value of the selected parameter of blood at a site in a blood flow path of said living body, from a first level to a second level;
determining the difference between the value of the selected parameter in said blood flow path at a first location upstream of said site and at a second location downstream of said site at said first level;
determining the difference between the value of the selected parameter in said blood flow path at the first location and at the second location at said second level;
determining the difference between the value obtained in the first said difference determining steps and the value obtained in the second said difference determining steps; and calculating blood flow as a function of the difference between the values obtained in said first and second difference determining steps and the change in the value of the selected parameter from said first level to said second level.
36. A method in accordance with claim 35 wherein said blood flow path includes at least a portion of the heart of a living body where the blood flow represents cardiac output.
37. A system for determining blood flow in a living body comprising:
means for changing the value of the selected parameter of blood at a site in a blood flow path of said living body, from a first level to a second level;
means for determining the difference between the value of the selected parameter in said blood flow path at a first location upstream of said site and at a second location downstream of said site at said first level;
means for determining the difference between the value of the selected parameter in said blood flow path at a first location upstream of said site and at a second location downstream of said site at said first level;
means for determining the difference between the value of the selected parameter in said blood flow path at the first location and at the second location at said second level;
means for determining the difference between the value obtained in the first said difference determining steps and the value obtained in the second said difference determining steps; and means for calculating blood flow as a function of the difference between the values obtained in said first and second difference determining steps and the change in the value of the selected parameter from said first level to said second level.
means for changing the value of the selected parameter of blood at a site in a blood flow path of said living body, from a first level to a second level;
means for determining the difference between the value of the selected parameter in said blood flow path at a first location upstream of said site and at a second location downstream of said site at said first level;
means for determining the difference between the value of the selected parameter in said blood flow path at a first location upstream of said site and at a second location downstream of said site at said first level;
means for determining the difference between the value of the selected parameter in said blood flow path at the first location and at the second location at said second level;
means for determining the difference between the value obtained in the first said difference determining steps and the value obtained in the second said difference determining steps; and means for calculating blood flow as a function of the difference between the values obtained in said first and second difference determining steps and the change in the value of the selected parameter from said first level to said second level.
38. A system in accordance with claim 37 wherein said blood flow path includes at least a portion of the heart of a living body where the blood flow represents cardiac output.
39. A method for determining instantaneous blood flow in a living body comprising the steps of:
determining the temperature level at a selected location in a blood flow path of said living body;
heating said selected location in said blood flow path to a final temperature level;
determining the final temperature level;
determining the ratio of the heating power required to heat said selected location to said final temperature level to the change in temperature level at said selected location.
determining the temperature level at a selected location in a blood flow path of said living body;
heating said selected location in said blood flow path to a final temperature level;
determining the final temperature level;
determining the ratio of the heating power required to heat said selected location to said final temperature level to the change in temperature level at said selected location.
40. A method in accordance with claim 39 wherein said indicating step includes integrating said ratio over a single cardiac cycle and dividing the integrated value by the period of said cardiac cycle to obtain the average instantaneous blood flow in said living body over said cardiac cycle.
41. A method in accordance with claims 39 or 40 wherein said heating and said final temperature level determining steps include determining the initial temperature level at said selected location before heating said selected location;
heating said selected location to said final temperature level;
maintaining said final temperature level at a constant value while heating said selected location over time; and determining the time varying change in the temperature level at said selected location which is required to maintain said final temperature level at said constant value.
heating said selected location to said final temperature level;
maintaining said final temperature level at a constant value while heating said selected location over time; and determining the time varying change in the temperature level at said selected location which is required to maintain said final temperature level at said constant value.
42. A method in accordance with claims 39 or 40 wherein said heating and said final temperature level determining steps include heating said selected location from an initial temperature level to a final temperature level by changing the temperature level by a fixed amount; and maintaining the changing temperature level at said fixed amount over time and determining the time varying temperature level over said time.
43. A method in accordance with claims 39 or 40 wherein said heating step includes heating a sensor positioned at said selected location, said sensor having a resistance R which varies as a function of the temperature thereof and further including the steps of:
heating said sensor to a final resistance value thereby changing the resistance value of said sensor;
maintaining the changing resistance value of said sensor during heating at a fixed amount over time; and determining the time varying final resistance value of said sensor during the heating over said time;
and further wherein said ratio determining step includes determining the ratio of the heating power required to heat said sensor to said final resistance value to the fixed changing resistance value thereof.
heating said sensor to a final resistance value thereby changing the resistance value of said sensor;
maintaining the changing resistance value of said sensor during heating at a fixed amount over time; and determining the time varying final resistance value of said sensor during the heating over said time;
and further wherein said ratio determining step includes determining the ratio of the heating power required to heat said sensor to said final resistance value to the fixed changing resistance value thereof.
44. A device for insertion into the bloodstream of a living organism to obtain physiological data comprising:
an elongated flexible catheter adapted to be inserted into the venous or arterial system of the living organism;
first temperature sensing means carried by said catheter;
second temperature sensing means carried by said catheter positioned to be upstream of said first temperature sensing means when in use; and heating means carried by said catheter between said first and second temperature sensing means and positioned to be substantially thermally isolated from said second temperature sensing means when in use;
said first and second temperature sensing means and said heating means being adapted to be in communication with a data processing system.
an elongated flexible catheter adapted to be inserted into the venous or arterial system of the living organism;
first temperature sensing means carried by said catheter;
second temperature sensing means carried by said catheter positioned to be upstream of said first temperature sensing means when in use; and heating means carried by said catheter between said first and second temperature sensing means and positioned to be substantially thermally isolated from said second temperature sensing means when in use;
said first and second temperature sensing means and said heating means being adapted to be in communication with a data processing system.
45. A device in accordance with claim 44 and further comprising thermistor means carried by said catheter and positioned to be downstream of said heating means when in use, said thermistor means being adapted to be in communication with said data processing system.
46. A device for insertion into the bloodstream of a living organism to obtain physiological data comprising:
an elongated flexible catheter adapted to be inserted into the venous or arterial system of the living organism;
temperature sensing means carried by said catheter;
and thermistor means carried by said catheter and positioned to be downstream of said temperature sensing means when in use;
said temperature sensing means and said thermistor means being adapted to be in communication with a data processing system.
an elongated flexible catheter adapted to be inserted into the venous or arterial system of the living organism;
temperature sensing means carried by said catheter;
and thermistor means carried by said catheter and positioned to be downstream of said temperature sensing means when in use;
said temperature sensing means and said thermistor means being adapted to be in communication with a data processing system.
Applications Claiming Priority (2)
| Application Number | Priority Date | Filing Date | Title |
|---|---|---|---|
| US08/106,068 US5797398A (en) | 1993-08-13 | 1993-08-13 | Method and apparatus for measuring continuous blood flow at low power |
| US08/106,068 | 1993-08-13 |
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| Publication Number | Publication Date |
|---|---|
| CA2169374A1 true CA2169374A1 (en) | 1995-02-23 |
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| Application Number | Title | Priority Date | Filing Date |
|---|---|---|---|
| CA002169374A Abandoned CA2169374A1 (en) | 1993-08-13 | 1994-08-11 | Method and apparatus for measuring continuous blood flow at low power |
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|---|---|
| US (6) | US5797398A (en) |
| EP (1) | EP0712290A1 (en) |
| JP (2) | JP3557211B2 (en) |
| CA (1) | CA2169374A1 (en) |
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1993
- 1993-08-13 US US08/106,068 patent/US5797398A/en not_active Expired - Lifetime
-
1994
- 1994-08-11 WO PCT/US1994/009064 patent/WO1995005115A2/en not_active Application Discontinuation
- 1994-08-11 CA CA002169374A patent/CA2169374A1/en not_active Abandoned
- 1994-08-11 EP EP94925821A patent/EP0712290A1/en not_active Ceased
- 1994-08-11 JP JP50706695A patent/JP3557211B2/en not_active Expired - Fee Related
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1995
- 1995-05-24 US US08/449,205 patent/US5692514A/en not_active Expired - Lifetime
-
1997
- 1997-10-07 US US08/946,366 patent/US6203501B1/en not_active Expired - Fee Related
- 1997-10-07 US US08/946,367 patent/US6165132A/en not_active Expired - Fee Related
-
2000
- 2000-12-08 US US09/733,595 patent/US6565516B2/en not_active Expired - Fee Related
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2003
- 2003-02-11 US US10/364,773 patent/US6913576B2/en not_active Expired - Fee Related
-
2004
- 2004-03-09 JP JP2004065117A patent/JP2004195251A/en active Pending
Also Published As
| Publication number | Publication date |
|---|---|
| US20010000792A1 (en) | 2001-05-03 |
| EP0712290A4 (en) | 1996-07-03 |
| WO1995005115A2 (en) | 1995-02-23 |
| US20030120162A1 (en) | 2003-06-26 |
| US6203501B1 (en) | 2001-03-20 |
| JPH09501591A (en) | 1997-02-18 |
| JP2004195251A (en) | 2004-07-15 |
| EP0712290A1 (en) | 1996-05-22 |
| US5692514A (en) | 1997-12-02 |
| US6565516B2 (en) | 2003-05-20 |
| US5797398A (en) | 1998-08-25 |
| WO1995005115A3 (en) | 1995-04-13 |
| JP3557211B2 (en) | 2004-08-25 |
| US6165132A (en) | 2000-12-26 |
| US6913576B2 (en) | 2005-07-05 |
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Legal Events
| Date | Code | Title | Description |
|---|---|---|---|
| EEER | Examination request | ||
| FZDE | Discontinued |