JPH0467857B2 - - Google Patents

Description

DETAILED DESCRIPTION OF THE INVENTION [Industrial Application Field] The present invention relates to a cardiac output measuring device that calculates cardiac output by integrating blood temperature data diluted with an indicator, for example. The present invention relates to a cardiac output measuring device having an automatic measurement start function and a function of determining the start point of integration in real time.

[Prior Art] Conventionally, an indicator dilution method has been used to measure cardiac output by right heart catheterization for cardiac function testing. This indicator dilution method includes a thermodilution method that calculates cardiac output from heat diffusion, a dye dilution method that calculates cardiac output from changes in illuminance due to dye diffusion, and a method that calculates cardiac output from changes in resistance due to electrolyte diffusion. There are electrolyte dilution methods to determine the amount. This thermodilution method will be explained.

In right heart catheterization, as shown in Figure 16, a catheter 4 is introduced from the jugular vein, femoral vein, cubital vein, etc., passes through the superior vena cava or inferior vena cava, the right atrium, and the right ventricle. is placed in the pulmonary artery. Catheter 25
In this case, the discharge port 26 is located in the right atrium, and the thermistor 1 is located in the pulmonary artery.
are placed respectively. Now, when a liquid with a temperature higher or lower than the blood temperature is injected into the right atrium from the discharge port 26, the liquid is diffused and diluted in the right atrium and right ventricle. The temperature of this diluted liquid is detected by the thermistor 27 located in the pulmonary artery, and a dilution curve (diagram of temperature change with respect to time) (Figure 17) of the temperature is obtained, and based on the area of the curve, etc.・Calculate cardiac output using the following formula (1) using the Hamilton method.

CO＝S _i・C _i・(T _b −T _i )・V _i ／S _b・C _b・∫ ^∞ / ₀ ΔT _b d
t……(1) where, CO: cardiac output, S _i : specific gravity of injected liquid C _i : specific heat of injected liquid, V _i : amount of injected liquid T _i : temperature of injected liquid, T _b : blood Temperature S _b : Specific gravity of blood, C _b : Specific heat of blood ∫ ^∞ ₀ ΔT _b dt: Area of thermodilution curve. Calculating cardiac output using the Stewart-Hamilton method is well known as a method that yields relatively accurate results, and cardiac output measuring devices based on this method are also available on the market.

However, such conventional cardiac output measurement devices require a relatively high level of skill and complicated operation, and therefore, it is difficult to determine how easy they are to use, how easy they are to operate, and how accurate they are. Being able to measure things is becoming more important. In order to meet such needs, the inventors of the present application have proposed the following method when calculating cardiac output using a dilution method, particularly a thermodilution method.
found the problem.

: In the thermodilution method described above, as can be seen from the formula for calculating cardiac output, the injection amount of the indicator liquid and its temperature greatly affect the accuracy of the calculated value. Incidentally, normally, the liquid temperature T _i is accurately measured with a thermistor or the like, and then the catheter 26 is primed and the indicator liquid is injected into the blood vessel.
This indicator liquid is cooled or heated by body temperature until it is injected, causing an error in the measured value. In other words, even if an indicator liquid set at a normal temperature is injected, the liquid that is first injected into the blood vessel is the indicator liquid that has been cooled or heated to the above body temperature, and then the liquid that is set at the normal temperature is the indicator liquid that is injected into the blood vessel. An indicator fluid is injected into the blood vessel. The cause of measurement error is that the total heat capacity of the indicator liquid injected into the blood vessel is unknown. The quantity V _i to be determined is determined by the temperature of the indicator liquid T _i . The capacity inside the catheter is determined by the above Stewart Wart.
This is a factor that is abstracted from Hamilton's equation, and is generally fixed depending on the outer diameter (French size) of the catheter. Conventionally,
Estimates are made from the above-mentioned injected liquid volume, catheter size, etc., and a correction constant is obtained by comparing with the experimentally obtained cardiac output value, and this correction constant is applied to the calculation of cardiac output. In this way, the influence of the residual fluid mentioned above is reduced to obtain a more accurate cardiac output.

However, the above correction constant obtained from the amount of liquid remaining in the catheter in the conventional example is just a list of numerical values, and has no meaning to the measurer. Calculation errors occurred when determining the correction constants, and input errors occurred when inputting the determined correction constants into the measuring device.

Note that this error is a problem that occurs not only in the thermodilution method but also in all cardiac output measurements using the dilution method, in view of the fact that a catheter is widely used in the indicator dilution method.

:Also, as can be seen from Figure 17, the operation of injecting the indicator liquid and the start of sampling of blood data necessary for calculating the area (integral) of the dilution curve must be performed at a predetermined timing. This can cause errors in calculations, such as starting the calculation of the integral a long time after the drop. However, in the past, the location where the indicating liquid is injected and the location where the measuring device is installed are often forced to be separated from each other. Since the start, that is, the start of integration was instructed, there was always the possibility of a measurement error occurring. Furthermore, when one person performs the operation, it is complicated because two operations must be performed quickly within a short period of time.

These error factors and complications are not limited to the thermodilution method, but are also problems of indicator dilution in general.

: In the measurement of cardiac output based on the thermodilution method shown in equation (1) above, the temperature of the injected indicator liquid T _i
It is essential to measure accurately. This is because not only the term (T _b −T _i ) in the above equation,
This is because T _i is also related to the heat capacity of the residual fluid, which is the error factor explained in .

Typically, this indicator liquid is immersed in an ice or hot solution that is maintained at a predetermined temperature. Therefore, if the measurement is to be carried out by an experienced measurer, or if absolutely accurate measurement of cardiac output is not required, an ice or hot agent to which the above-mentioned indicator liquid can be applied is recommended. In some cases, the temperature of the solution may be regarded as the indicated liquid temperature T _i . Nevertheless, in the past, the indicator fluid temperature was measured each time, making cardiac output measurement inefficient.

That is, in a cardiac output measurement device using the dilution method, -1: In cases where it is acceptable to consider the temperature of the above-mentioned ice agent or hot agent solution as the indicated liquid temperature T _i , the temperature required for calculation is -2: If the data is necessary for calculation but can be obtained without taking the trouble of measuring it, use the manually preset data instead. The problem is that the device itself detects what the measurer intends, or detects such a designation from the operator, and if this is detected, it uses data preset separately manually in place of it. Appears as.

On the other hand, in the method of measuring cardiac output using the thermodilution method or indicator dilution method described above, cardiac output is measured intermittently every time the indicator liquid is injected, so cardiac output cannot be measured continuously. It cannot be used to measure stroke volume. Furthermore, if measurements are to be performed frequently, the total amount of liquid to be injected increases, which increases the burden on the subject, and at the same time increases the risk of infection associated with the liquid injection operation, which is undesirable.

In order to eliminate the disadvantages associated with cardiac output measurement according to the conventional indicator dilution method, in particular the thermodilution method, which allows only one-shot measurement, the applicant of the present invention filed a patent application filed in Japanese Patent Application No. No. 244586 (Unexamined Japanese Patent Publication No. 1983-
125329), we proposed a completely new cardiac output measuring device that enables continuous cardiac output measurement. With this continuous cardiac output measurement device, long-term cardiac output measurement and continuous recording became a challenge.

Furthermore, the above-mentioned patent application No. 59-244586 calculates the relationship between cardiac output and blood flow velocity by simultaneously measuring cardiac output and blood flow velocity using the thermodilution method, and from now on, this relationship is The gist is to calculate the cardiac output from the and blood flow velocity. In other words, since there is a fixed relationship between cardiac output and blood flow velocity, unless you are looking for the absolute cardiac output value, you can only find the blood flow velocity value and calculate the relative cardiac output value. This makes it possible to notice changes.

[Problems to be Solved by the Invention] As a technique for disclosing the start point of the above-mentioned integration, there is, for example, Japanese Patent Publication No. 58649/1983 (``Liver Function Testing Apparatus''). This technique samples data once every second, stores 1200 of these data in memory, finds the rising point of the signal from among the stored data, and uses that rising point as the starting point of integration. It is. However, this technology
Although it is practical for liver function tests, it is not suitable for measuring cardiac output using the indicator dilution method. This is because, in this measurement, a catheter is inserted into a patient's blood vessel and an indicator is injected into the blood vessel. That is, if cardiac output measurement, that is, integration, is started before this preliminary preparation is completed,
For example, in the device disclosed in Japanese Patent Publication No. 60-58649, meaningless data for a long period of time before injection is simply stored in the memory. There is no problem if a large capacity memory is provided, but a large capacity memory increases the cost. Furthermore, with a small-capacity memory, there is a problem in that storing useless data takes up memory space, and the memory capacity becomes insufficient when storing the important data after the indicator is injected. Therefore, this conventional technology still does not solve the problem that the operator has to pay attention to synchronizing the time point at which the measurement of the device starts and the time point at which the indicator injection starts. There isn't.

Therefore, the present invention solves the problem of being easy to use, easy to operate, and capable of relatively accurate measurement. This is achieved by proposing a cardiac output measuring device that is equipped with an automatic measurement start function that does not cause problems and further reduces the operational burden on the measurer.

[Means for Solving the Problems and Their Effects] The configuration of the present invention for achieving the above-mentioned problems includes: a cardiac output measuring device that measures cardiac output based on a dilution method; blood data detection means for outputting data related to blood diluted with the indicated indicator liquid; blood data detected by the blood data detection means at a first time; and blood data detected by the blood data detection means at a second time after the first time.
a change amount calculation means for calculating the amount of change from the blood data at the time of , a comparison means for comparing whether the calculated time change amount is equal to or greater than a predetermined value, and a time change amount calculated by the comparison means It is characterized by comprising an integrating means that starts integrating the blood data when it is determined that the blood data is equal to or greater than a predetermined value, and a cardiac output calculating means that calculates the cardiac output based on the integrated value.

In the dilution method using an indicator, it is important to integrate the blood data values from the time when the blood data values change due to the injection of the indicator. The comparison means monitors the change in real time, and if the change satisfies a predetermined condition, integration is started, so the operator does not have to worry about synchronizing the start of indicator injection and the start of measurement. Furthermore, the real-time nature of the change point detection operation eliminates the waste of storing meaningless measurement data before the start of injection.

[Examples] Examples according to the present invention will be described in detail below with reference to the accompanying drawings.

<Appearance of Example Device> FIGS. 1A to 1C are plan views of a continuous measurement and recording device for cardiac output to which the present invention is applied, respectively.
A front view and a back view are shown. Figure 2A shows the external appearance of the catheter connected to this measuring device.
Figure B shows a cross-sectional view along the longitudinal direction of the distal end portion of this catheter, and Figure 2C shows a cross-sectional view of the opening of the catheter.

The principle of this measuring device is to measure the initial cardiac output CO ₀ based on indicator dilution methods such as thermodilution method, electrolyte dilution method, dye dilution method, etc., and the blood flow rate when this cardiac output is measured. V ₀ is measured, and then a parameter S is determined that links this blood flow velocity V ₀ to the initial cardiac output CO ₀ (here, CO ₀ =v ₀ ·S). After finding this parameter, all you need to do is measure the blood flow velocity v _x at any time, and from this blood flow velocity v _x and the parameter S, the cardiac output CO _x (=v _x・S) can be found.
The measuring method for determining the initial cardiac output CO ₀ may be any of the indicator dilution methods described above, but in the embodiments shown in FIG. 1A and subsequent figures, it is particularly based on the thermodilution method.

The external appearance of the measuring device 100 shown in FIGS. 1A to 1C will be explained. FIG. 1A is a top view of apparatus 100. 50 is a recording device such as a well-known plotter. This plotter 50 outputs measurement results and the like. Reference numerals 52 and 53 are switches for specifying the output mode on the plotter 50 when measuring the initial cardiac output CO ₀ by the thermodilution method. The switch 52 displayed as “thermodilution value” indicates the cardiac output.
Specifies to output CO as a numerical value (as shown in Figure 5). The switch 53 labeled ``thermodilution curve'' specifies that changes in blood temperature <sub>Tb</sub> , etc. be output as a curve (FIG. 4). As mentioned above, the switches 52 and 53 are used only to measure the initial cardiac output value CO <sub>0</sub> ("intermittent measurement"), whereas the switches described next are mainly used to measure the cardiac output value. Continuous measurement (used for ``continuous measurement'')

Switches 57 and 58 specify the paper feed speed of plotter 50 (20 mm/hour and 10 mm/minute, respectively).
The switch 54 displayed as “progress chart” shows the cardiac output.
This is a switch that instructs to output the 12-hour progress of CO _x , etc. on the plotter 50 (Fig. 8).
The switch 55 that is displayed outputs the continuously obtained data such as cardiac output CO _x stored in the memory on the plotter 50 for the past 30 minutes (6th
7). The switch 80 labeled "Memo" is a memo switch, and when this switch is pressed, the recording paper
It is fed about cm. The switch 56 labeled "continuous recording" is a switch that instructs to output changes in CO _i , T _b , v _{x ,} etc. on recording paper in real time (FIGS. 6 and 7). The above-mentioned "memory playback" mode records the past 30 minutes of memory data, but even during this "playback" the measuring device 1
00 stores data such as cardiac output measured in real time. Therefore, before long, the past data recorded on the recording paper and the currently stored real-time data become temporally close and coincide with each other. In that case, instead of stopping the "Playback" mode,
It automatically switches to the ``continuous recording'' mode mentioned above.

Switches 59, 60 (Fig. 1A), 61, 62
(Figure 1B of the front view) etc. are switches for manually setting the desired data when inputting the desired data to the device 100, and the setting values input by these switches are as follows: In order,
Catheter diameter (unit: French), injection volume of indicated liquid (unit: ml), subject's body surface area (unit:
m ² ), the initial CAL value (unit: L/min). initial
The CAL value will be described later. On these data input switches, press ▲ if the value to be set is greater than the displayed value (displayed from among the □), and press ▼ if it is smaller.

Switch 6 displayed as “CONTINUOUS”
Reference numeral 4 denotes a switch for setting the mode of the measuring device to continuous mode, and when the mode is set to continuous mode, this switch/indicator 64 lights up.
The switch 65 labeled "SINGLE" is a single (intermittent) mode switch/indicator, and is mainly used when returning from continuous mode to single mode when it is desired to reset the above parameter S again. This switch 65 lights up when pressed to enter the intermittent mode (SINGLE mode). Display 68 is a four-digit LED that digitally displays cardiac output. The switches 66 and 67, which are displayed as ``BLOOD'' or ``INJECTATE,'' determine whether the temperature displayed on the 3-digit LED display 69 is the blood temperature T _b or the temperature of the indicator liquid T _i . It is to be specified.

The switch/indicator 70 labeled "ENTRY" indicates that the measurement of the initial cardiac output value CO ₀ using the thermodilution method has been completed, and that the measuring device 100
Displays that it is now possible to register (ENTRY) CO ₀ as data. When this switch 70 is pressed while this ENTRY indicator 70 is lit, the initial cardiac output value CO ₀ is registered as data. This registration of CO ₀ is performed to find the parameter S, but the reason we are planning to register multiple CO ₀ is to obtain a more accurate
The purpose is to determine CO ₀ , that is, the initial cardiac output value CO ₀ is measured several times using the thermodilution method, and the measurer judges and registers the data that seems reliable. This is to set the average value of the registered data as the initial value CO ₀ .

The switch/indicator 72 labeled "START" indicates the initial cardiac output value CO ₀ due to thermodilution.
When the measurement preparation is complete, press this switch to start the integration based on the Stuart-Hamilton equation.

The display 71 is a bar graph, and mainly displays the value of blood temperature T _b in real time.
73 and 74 are the connector 15 of the catheter 4 (FIG. 2A) which has a built-in blood temperature sensor, and the temperature probe 12 (FIG. 2A) for measuring the temperature of the injected fluid, respectively.
This is a connector for connecting the connector 16 shown in the figure. 63 is a power switch, and 78 is an indicator lamp indicating that the built-in battery is being used.

Switch 75 in FIG. 1C is a switch for manually setting the temperature of the injection liquid. 76 is an RS232C interface connector for communicating with other external measurement devices, and 78 is a connector for transmitting analog signals such as blood temperature T _b , blood flow velocity v _x , cardiac output CO _x, etc. to other external measurement devices. The output terminal 77 is a connector for connecting a power cable.

<Functions of the measuring device 100> The features of the embodiment device 100 shown in FIGS. 1A to 1C and 2A to 2C are as follows: In order to measure the initial value CO ₀ , the thermodilution method is used. To compensate for the influence of indicator liquid remaining in the catheter when injecting indicator liquid,
Instead of what the measurer was looking for from a table of values determined in advance based on experimental results, the measurer can create a cable with correction constants in advance and use the table as preliminary data to access. This measuring device is provided with a switch 59 for setting and inputting the outer diameter size (unit: French) of a catheter for injecting liquid, and a switch 60 for setting and inputting the amount of liquid injection (unit: ml). Note that the influence of the indicator liquid remaining in the catheter described above also depends on the indicator liquid temperature T _i . In this embodiment, T _i for this purpose is either actually measured by the thermistor 12a of the temperature measurement probe 12 (FIG. 2A), as described later, or
Or, instead of that, manual switch 69
(Figure 1B) is used to input settings.

:Blood temperature after injecting the indicated liquid into the blood vessel.
Automatically recognizes changes in T _b . That is, the integral starting point (point of change in blood temperature T _b ), which is essential for accurately performing the integral calculation based on the above-mentioned Stewart-Hamilton equation (1), can be recognized without bothering the measurer.

: The temperature probe 12 is used to measure the temperature of the indicator liquid, but if the connector 16 of this probe 12 is not connected to the measuring device 100, this disconnection is detected and the switch 75 ( Input the value previously set in FIG. 1C) as the indicated liquid temperature T _i .

:Once, measure the initial value CO ₀ by thermodilution method,
Once the parameter S is determined, the continuous cardiac output CO _i , blood temperature T _b , blood flow velocity v _x, etc. are continuously stored in the memory and output on recording paper as necessary. This output mode is as follows: (): Measured CO _x , etc. are output on the recording paper of the plotter 50 in approximately real time (Fig.
Figure 7). This is "continuous record" switch 56
This is done by pressing .

(): The past 30 minutes stored in memory.
Data such as CO _x is reproduced and output (Figures 6 and 7). This is done by pressing the "memory record" switch 55. As mentioned above, this "memory recording" mode automatically shifts to the "continuous recording" mode.

(): Data such as CO _x for the past 12 hours is compressed and reproduced and output on the plotter 50 (Fig. 8).
This is done by pressing the "progress chart" switch 54. When measuring data on a patient who requires rest and noiselessness, the collected data is not recorded on the plotter 50 and the printer prints noisy, but instead is stored in memory and used when necessary. By pressing the switch 54, the data is output on paper.

:The parameter S can be set manually via the switch 62. Therefore, this S
CO _x can be continuously determined from the and blood flow velocity v _x . This is useful when you want to know the relative changes in CO _x .

: With the bar graph display 71, changes in blood temperature T _b can be visually recognized when determining the initial cardiac output value CO ₀ by the thermodilution method. In particular, this bar graph displays the blood temperature baseline (reference temperature value) in addition to displaying temperature changes moment by moment.
T _b0 ) and maximum temperature T _bnax can be fixed and displayed.

:This device is equipped with a battery (14 in Figure 9) to protect the stored data from failures such as power outages.
8) is built-in. When the output voltage of the battery 148 drops, the LED 78 is blinked to alert the user.

: To an external measuring device other than this measuring device 100,
Terminal 78 outputs cardiac output CO _x , etc. as an analog signal.
an analog output circuit 142 for outputting from
It is equipped with 151.

: To an external measuring device other than this measuring device 100,
It is equipped with a circuit 144 that outputs CO _x etc. as a digital signal using an RS232 interface.

<Structure of Catheter> A catheter incorporating a thermistor for measuring blood temperature T _b and a thermistor for measuring blood flow velocity v according to the thermodilution method is shown in FIG. 2A. In the figure, catheter 4 is configured to have four lumens. This catheter 4
is a pressure detection port 18 provided at the tip, a balloon 17 made of a flexible elastic body attached several mm back from the tip so as to cover the entire tip of the catheter tube, and the balloon 17 is inflated and deflated. For this purpose, there is a balloon side hole 25 provided on the side of the tube inside the balloon for injecting or extracting air (preferably carbon dioxide gas), and a hole 10 from the tip.
Thermistor 1 is placed ~20mm away, thermistor 2 is placed 10~15mm closer to the base, and thermistor 2 is further spaced 8.5~38cm from thermistors 1 and 2, and 12~40cm from the tip. and discharge ports 3 provided at positions spaced apart within a range of. Thermistor 1 is used to measure the thermal equilibrium temperature for measuring blood flow velocity v ₀ (or v _x ). Thermistor 2 measures the initial cardiac output using the thermodilution method.
Used to measure diluted blood temperature T _b where required to measure CO ₀ . The discharge port 3 is for discharging an indicator liquid.

The catheter 4 shown in Fig. 2A is inserted into the jugular vein, femoral vein, or cubital vein, and is used in the pulmonary artery via the superior vena cava or inferior aorta, the right atrium, and the right ventricle. Considering that the flow direction is from the proximal end of the catheter to the distal end, the thermistor 2 is provided closer to the distal end (ie, downstream of the blood flow) when viewed from the discharge port 3. On the other hand, in the case of a catheter that is conduited from a peripheral artery and used in the aorta, the thermistor 2 is placed on the proximal end side of the catheter when viewed from the discharge port 3 because the blood flow direction is opposite. .

FIG. 2B shows a sectional view of the main part of the catheter 4.
In the figure, pressure detection port 18, balloon side hole 2
5 (Fig. 2A), thermistors 1 and 2, and discharge port 3 are as follows:
Each communicates with the four lumens. These four lumens are a pulmonary artery pressure lumen 19 that transmits the pressure of the pulmonary artery, a balloon lumen 20 that is an air passage that inflates and deflates the balloon 17, a thermistor lumen 21 that stores thermistors 1 and 2 and their lead wires, and a thermistor lumen 21 for dilution. an injection lumen 23 through which an indicator liquid passes. Furthermore, these four lumens are independent, and are further connected to a pulmonary artery pressure measuring tube 8, a balloon tube 6, a thermistor tube 13, and an indicator fluid injection tube 10 at the rear end of the catheter, as shown in FIG. 2A. ing. Each tube 8, 6, 10 is provided with a connector 7, 9, 11 at its rear end. The thermistor tube 13 is connected to the connector 15, and thermistors 1 and 2 are installed in this tube.
Lead wires 22 and 24 respectively connected to are passed through. Connector 15 is connector 73 of device 100
connected to.

Furthermore, the catheter of the embodiment shown in FIGS. 2A to 2C will be described in detail. FIG. 2B is an enlarged sectional view of the thermistor 1, thermistor 2, and balloon 17, and FIG. 2C is a sectional view of the catheter tube 4 taken along the line -'. As mentioned above,
The four-lumen structure of this catheter consists of a balloon lumen 20, a pulmonary artery pressure lumen 19, an injection lumen 23 (FIG. 2C), and a thermistor lumen 21, as shown in FIG. 2C. The balloon lumen 20 has a balloon side hole 25,
It communicates with the balloon tube 6. The pulmonary artery pressure lumen 19 has a pressure detection port 18 and communicates with the pulmonary artery pressure measurement tube 8 . The injection lumen 23 has a discharge port 3 located 12 to 40 cm from the tip of the catheter, and communicates with the indicator liquid injection tube 10 on the proximal side. Further, the thermistor lumen 21 has side holes 26 and 27 in which the thermistors 1 and 2 are attached, respectively, at a position 1 to 2 cm away from the tip and at a position further 1 to 1.5 cm away from there toward the base side. . Also, the lumen 21 connects the thermistor leads 22 and 24 from the thermistors 1 and 2.
It also has a built-in thermistor tube 13 and communicates with the thermistor tube 13 on the base side.

It is more preferable that the thermistor 1 be a self-heating type thermistor and be located downstream of the thermistor 2 in the blood flow direction. That is, when measuring cardiac output with the catheter 4, it is necessary to accurately measure the temperature of the blood flow diluted by the thermistor 2, and the thermistor 2 is made less susceptible to the influence of the thermistor 1, which is a self-heating type. It's for a reason.
Even so, in order to position the thermistor 1 correctly in the pulmonary artery, this thermistor 1 must be placed in the second A.
Differently from the figure, it may be located further upstream from the thermistor 2 by 10 to 15 mm in the blood flow direction.

The characteristics of the thermistor 1 used in the catheter of the example are B _25-45 = 3500K, R(37) = 1000Ω, and its size is 1.18 ¹ ×0.4 ^w ×0.15 ^t (unit: mm). The characteristics of thermistor 2 are B _25-50 = 3500K, R
(37) = 14Ω, its size is 0.75 ¹ × 0.1 ^w × 0.15 ^t . The thermistor 1 preferably generates a heat value of 0.01 to 50 joules; a higher heat value may raise the blood temperature or even damage the blood vessel wall if it comes into contact with it, and a lower heat value is preferred. Either amount is not preferable because the detection sensitivity becomes low. In addition, thermistor 2 is B _25-45 =
3980K, R(37) = 40KΩ, its size is 0.50 ¹ ×
It may be 0.16 ^w × 0.15 ^t .

Incidentally, the probe 12 shown in FIG. 2A is a probe for measuring the temperature of the indicator liquid, and near its tip there are:
A thermistor 12a is provided, and a connector 16 for connecting to the connector 74 of the measuring device 100 is provided on the terminal end side.

<Operating procedure of measuring device> In order to better understand this measuring device 100,
The operating procedure seen from the operator's side will be explained using FIG. 1A, FIG. 1B, FIG. 3A, and FIG. 3B.

First, in step S1, the injected liquid temperature probe 12 is
is connected to the measuring device 100. At step S2,
The temperature probe 12 is immersed in the injection instruction liquid in an ice-cooled container (in this example, provided within the measuring device 100). This liquid may be physiological saline or the like. After priming the catheter 4 in step S3, it is inserted up to the pulmonary artery in step S4.
The catheter 4 is inserted through a vein in the upper or lower limb, and placed in the pulmonary artery. Next, step S5
Then, connect the connectors of the catheter shown in FIG. 2A to the main body of the device. When these connectors are connected to the device, first, by sequentially moving the intravascular indwelling position of the catheter 4, the pressure detection port 1 is
8, central venous pressure via tube 8, connector 9;
Confirm based on blood pressure values and pressure waveforms such as right atrial pressure, right ventricular pressure, and pulmonary artery pressure. After the catheter is indwelled, the pulmonary artery pressure is measured, and the balloon 17 is inflated to occlude the pulmonary artery to determine the pulmonary artery wedge pressure. In this way, the catheter 4 is left in place.

Next, in step S6, predetermined values are set using the switches 59-61. The switch 62 is set when measuring the relative CO _x (described later), and when measuring the initial cardiac output value CO ₀ by the thermodilution method described now is not necessary.

In step S7, the measuring device is powered on. While confirming that there is no abnormal display (indicating a device failure) in step S8, the process waits for the SINGLE indicator 65 to light up in step S9. indicator 6
When 5 lights up, BLOOD indicator 6
6 lights up, and the display 69 shows the blood temperature in the pulmonary artery.
T _b is displayed. If you want to know the temperature of the injectate at this point, press the INJECTATE switch 67, and the display on the display 69 will change to the display of the injectate temperature T _i . At this point, the bar graph display 71 shows:
It should display the baseline of blood temperature T _b .

Note that the display of T _i automatically returns to the BLOOD display after 30 seconds have elapsed, and the display 69 displays T _b again. The reason why the display automatically returns to T _b is also convenient _for the operator.
This is because it is more meaningful as information.

Furthermore, in step S10, the program waits for the START indicator 72 to light up. When this indicator lights up, the thermodilution initial cardiac output value CO ₀
This means that you are ready to start measuring. In step S11, the mode of information to be output to the plotter 50 is selected by the switch 52 or 53, if necessary. In step S12, the pot 11 is opened and liquid is injected into the blood vessel.

After liquid injection, the measuring device automatically reads the change in T _b and determines the optimal time to start integration. If the operator presses the START switch 72 before detecting this optimum point, the integration of T _b starts from the point at which the operator presses the START switch 72. If switch 72 is not pressed,
Integrate T _b data from the time determined by the device itself. The cardiac output measurement calculation using this thermodilution method usually completes in ten seconds or more, but changes during that time are stored in a predetermined memory (RAM 132 in FIG. 9) and displayed on the display 69 and bar graph display. 71. Furthermore, if the switch 53 has been pressed, the change in blood temperature _Tb as shown in FIG. 4 is output to the plotter 50. These displays allow the operator to confirm that the device is operating normally. In FIG. 4, the blood temperature T _b is moving upward toward the negative side. Note that the date, time, etc. will also be output on this graph for the convenience of the measurer.
The output example shown in FIG. 4 has an output speed of 300 mm/min.

When the measurement of the cardiac output CO ₀ by thermodilution is completed, the value is displayed digitally on the display 68, and if the switch 52 has been pressed, the plotter 50 displays the value as shown in FIG. Output the measurement results. In this Fig. 5, in addition to the date and time, cardiac output CO, body surface area BSA input from switch 61, blood temperature BT (=T _b ), catheter outer diameter CAT (=Fr), and indicated liquid injection are shown. volume IV (=V _i ), cardiac index CI (=CO/BSA), infusion fluid temperature IT (=T _i ), and "**ENTRY*" indicating that it has been registered.
*” indications, etc. are also recorded. Completion of measurement of the initial cardiac output CO ₀ can be confirmed by lighting the ENTRY indicator 70 (step S15).

If the measurer determines that the obtained data such as CO is reliable, he or she presses the ENTRY switch 70 to register this data (step S17). After a while, the START indicator 72 lights up again. Now, by thermodilution method,
The measurement of initial cardiac output CO ₀ is completed.

If you want to obtain more accurate CO, repeat the steps from step S11 described above to obtain a plurality of measurement data. Press the ENTRY switch 70 for each measurement, register the
Obtain the average value from the CO ₀ values to determine the initial cardiac output value CO ₀ value.

Once sufficiently reliable initial cardiac output CO ₀ data has been obtained in the manner described above, it is time to start continuous CO _x measurements. This is simply
This is done by simply pressing the CONTINUOUS switch 64. When you press this switch 64,
CONTINUOUS indicator 64 lights up,
The entire measuring device is in continuous measurement mode. still,
When the SINGLE switch 65 is pressed, the measuring device goes into intermittent mode again.

When the device enters the continuous device mode and the continuous recording switch 56 is pressed, as shown in FIGS. 6 and 7,
Blood temperature BT (=T _b ), CO (=CO _x ), blood flow velocity v (=
v _x ) is output on the plotter 50. Figure 6 is 10
Figure 7 shows the output at a paper feed speed of 20 mm/hour. Furthermore, Figures 6 and 7
Although not shown in the figure, it goes without saying that various data such as the date shown in FIG. 5 are output together with the graph. In the example shown in FIG. 7, at around 4:30 pm, the thermodilution method is used again to measure the initial cardiac output value and CO ₀ value.
In order to clearly indicate on the graph that this initial cardiac output CO ₀ has been remeasured, the device 10 of this embodiment
0, a vertical line is inserted as shown in FIG.

<Configuration of Measuring Device> The entire circuit connection in the measuring device 100 with the catheter 4 and the injectate temperature probe 12 connected is as shown in FIG. This measuring device 100 has two electrically isolated measuring circuits 120 (FIG. 9A) and 130 (FIG. 9B).
Figure). The circuit 120 mainly converts electrical signals (voltage) from the thermistors 1 and 2 in the catheter 4 and the thermistor 12a in the probe 12 into temperature data, and sends the data to the measurement recording circuit 130 via the optical communication circuit 108. send. Measurement recording circuit 1
30 is the temperature data from the measurement circuit 120,
The initial cardiac output value CO ₀ , blood flow velocity v ₀ (or v _x ), parameter S, continuous cardiac output CO _x, etc. are calculated and displayed on the plotter 50 or on various displays. Execute. The local CPU 105 controls the measurement circuit 120, and the main CPU 133 controls the measurement recording circuit 130.

<Measurement circuit 120> Now, the catheter type sensor 150 shown in FIG. 9 uses the catheter 4 shown in FIG. It has a built-in thermistor 2 that detects the blood temperature inside the body. This catheter-type sensor 150 is introduced to the pulmonary artery by right heart catheterization in the same manner as described above. Connector 1 of sensor 150
5 is coupled to the connector 73 of the main body 100. The thermistor 2 is connected to the thermistor 2 via the lead wire 24.
It is connected to a constant voltage circuit 112 that drives the blood and a blood temperature detection circuit 113 that measures the temperature of the blood.

The pulmonary artery blood temperature signal detected by the thermistor 2 is detected by the blood temperature detection circuit 113 as a voltage signal E _b . On the other hand, thermistor 1 is
It is connected to a thermistor temperature detection circuit 115 and a constant current circuit 111 via a lead wire 22 .
Then, a predetermined current I _c is supplied to the thermistor 1 from the constant current circuit 111, and the thermistor 1 is heated. Further, the temperature signal detected by the thermistor 1 is sent to the thermistor temperature detection circuit 115, and this detection circuit 1
15, it is detected as a voltage _Et . The thermistor 12a in the injectate temperature probe 12 is driven by a low voltage circuit 101, and its temperature changes are detected by the circuit 10.
2, it is detected as a voltage value E _i . thus,
The local CPU 105 has three thermistors 12
The output voltages E _i , E _t , and E _b from a, 1, and 2 are converted into the temperature T _i of the infusion liquid, the resistance R _t of thermistor 1, the temperature T _t of thermistor 1, and the blood temperature T _b .

This operation will be explained in more detail. local
The CPU 105 drives an analog switch 103 having a multiplexer function to perform time division processing.
E _i , E _t , and E _b are input to the 14-bit A/D converter 104, and the above voltage values are sequentially converted into digital values in the RAM.
107. These voltage values are ROM10
From the voltage-temperature conversion table stored in 6,
Converted to temperature data. Further, the resistance value R _t of the thermistor 1 is calculated by R _t =E _t /I _c . Here, I _c is the current flowing through the thermistor. The local CPU 105 sends data such as T _i , R _t , T _t , and T _b to the measurement recording circuit 130 via an optical communication line. The communication control (for example, performed by the well-known polling/selecting method) is carried out by the local CPU 10.
5 and main 133.

By using a relatively high-resolution 14-bit A/D converter 104, it has become possible to measure up to a high cardiac output region, as will be described later.

<Measurement recording circuit 130> In the measurement recording circuit 130, the temperature T _i of the injection liquid and the thermistor 1 are determined by this simple communication control procedure.
resistance R _t , temperature of thermistor 1 T _t , blood temperature T _b
etc., and store them in the RAM 132 in chronological order. The CPU 133 calculates the initial cardiac output value CO ₀ , the blood flow velocity v ₀ (or v _x , and the continuous cardiac output CO _{x )} based on the above data.

An RTC (real-time lock circuit) 147 counts real time and is further used for time monitoring such as when the LED display 69 changes the display from T _i to T _b after 30 seconds.

FIG. 10A shows the structure of data stored in RAM 132. The capacity of this RMA is an area (132a) that stores data such as T _b obtained by A/D conversion every 40 ms for 30 minutes, and a region (132a) that stores data such as continuous cardiac output CO _x at any time from these data. 30 measurements
12 minutes of storage area (132b) and 1.5 minute average value of data such as CO _x obtained by continuous calculation.
It is only possible to store at least the area (132c) for storing time.

FIG. 10B shows three address counters 160, 16 for storing data in RAM 132.
1,162 and the three areas (132a, 132b,
132c. A data storage address counter (SC pointer) 160 points to an address at which data from the measurement circuit 120 is stored in the RAM 132. The data read address counter (RC pointer) 161 is connected to the RAM 132 for calculating the initial cardiac output value CO ₀ .
Reads the data stored in the memory and points to the address where the calculated value will be stored. Data print address counter (PC pointer) 1
62 points to data required when outputting data on the plotter 50. The reason why three pointers are used in this way is that data storage, data reading, data output, etc. are performed independently and in parallel as subroutines.

<Calculation of initial cardiac output CO ₀ > In Figure 11, the main CPU 133 is the local CPU.
Receives measurement data from 105 and stores it in RAM 132.
This is a routine that stores in . As mentioned above, the SC pointer is used as the pointer at this time.

13A and 13B show a control routine of the CPU 133 for measuring the initial cardiac output value CO ₀ value.
The control shown in FIGS. 13A and 13B also includes an automatic measurement start function.

Manual start First, in step S40, check the READY state of the measuring device. This state is based on the premise that at least no hardware failure has occurred in the entire device, and is a state in which a stable blood temperature T _b is detected. The stable state of T _b is T _b within a certain time width
The variance of the data is calculated and the determination is made based on whether the variance is below a predetermined value. Furthermore, this
When READY state is detected, as mentioned above,
START indicator 72 lights up. In step S41, it is determined whether the START switch 72 has been pressed.

If it has been pressed, this is an instruction by the operator to start manual measurement. step
S50 uses the real-time clock RTC to check the time t.
147 and sets the pointer value according to the time t to the RC pointer 161 in step S51. In step S52, the injection liquid temperature T _i is set to RAM1.
I know from 32. In step S53, one T _b is read from the RAM 132 according to the RC pointer 161. In step S54, a baseline temperature T _b0 is detected from data before time t.
The detection of T _b0 is obtained from, for example, the average temperature of the low variance state used to determine the READY state. ΔT _b at step S55 (Fig. 17)
is calculated according to the following formula.

ΔT _b =t _b0 −T _b Here, the sign of blood temperature T _b is positive in the direction of temperature decrease. In step S56, this ΔT _b
, and calculate Σ△T _b (integral). This ΣΔT _b corresponds to ∫ ^∞ ₀ ΔT _b dt in the Stewart-Hamilton equation. In step S57, it is determined whether the blood temperature has not decreased even after a certain period of time has elapsed. That is, in step S57, when the temperature drop ΔT _b is smaller than the predetermined threshold value TH _a , it is determined that the indicator liquid has not yet reached the thermistor 2, and the counter ERTIM is incremented in step S58. If the state of ΔT _b ≦TH _a continues until ERTIM overflows, it is assumed that some condition has occurred, and the occurrence of an error state is displayed in step S60. This display is displayed on the display 68.
"Er" is displayed along with the error code number.

If the device operates normally and the indicator is injected, ΔT _b >
TH _a . Step S53 until ΔT _b > TH _a
~Step S57 Step S58 Step S59 The integration (ΣΔT _b ) performed in step S56 in the loop of step S53 is expected to be almost “0” because the blood temperature T _b has no difference from T _b0 . Ru. Furthermore, even if there is a noise-like change, the integral is "0" because the change is in the positive direction and in the negative direction.

When ΔT _b >TH _a is detected, the process advances to step S61. This will be explained with reference to FIG. Step S61
detects the peak of the dilution curve. Step 61
T _bnax stores the maximum temperature drop from T _b0 during temperature drop, and is initially initialized to "0". If there is a temperature drop (until a peak is detected), in step S61, since ΔT _b >ΔT _bnax , ΔT _bnax is updated in step S62.
In step S63, the temperature value ΔT _bc (used in step S68), which is a guideline for terminating the integral calculation, is updated. Theoretically, the longer the integration is performed, the higher the accuracy will be. Even in Stewart-Hamilton, the interval of integration extends to infinity. However, increasing the integration interval does not improve measurement efficiency;
In addition, it is easily affected by noise near T _b0 .
ΔT _bc is set as a lower limit temperature that increases measurement efficiency (shortens the integration interval) without significantly reducing accuracy. Since the peak temperature ΔT _bnax is large or small and ΔT _bc is also affected, ΔT _bc is updated in step S63 every time ΔT _bnax is updated. In step S64, the RC pointer is incremented to point to the new temperature data, and the step
Returning to S53, the above-described control is repeated. thus,
Integration is performed until a peak is detected.

When the blood temperature stops decreasing, the determination of ΔT _b > ΔT _bnax in step S61 becomes NO, so step
Proceed to S66. In this step S66, ΔT _bnax −ΔT _b >TH _b is checked. Here, TH _b is a predetermined threshold value.
As shown in No. 18, when the above relationship is satisfied, it can be considered that the peak has been exceeded.
In step S67, a peak flag is set. This peak flag indicates that a peak has been detected at least once after the start of measurement. In step S68, the blood temperature increases (the curve decreases),
Check whether ΔT _b reaches the critical temperature ΔT _bc at which integration stops.

When the temperature rise ΔT _b has not yet reached ΔT _bc ,
The process advances to step S64, and then returns to step S53 to continue the integration.

When the temperature rise ΔT _b becomes ΔT _bc , the process advances to step S69 and the peak flag is checked. Since the peak flag has been set, the process proceeds to step S70 and subsequent steps to calculate the cardiac output _CO0 . Here, the significance of the peak flag becomes clear. Due to the influence of noise etc., in step S68, the peak is not detected even once.
It is possible that ΔT _b ≦ΔT _bc . Therefore,
A peak flag is set only when a peak such that ΔT _bnax −ΔT _b >TH _b is detected, and until this flag is set, integration is prevented from being stopped by mistake once started. Furthermore, if the above-mentioned TH _a and TH _b are set to the same threshold value, the peak flag and related control (step
S66, S67, S69, S70) are no longer needed.

Note that when emphasis is placed on increasing the accuracy of integration, it is desirable that ΔT _bc be approximately equal to ΔT _b0 .
Further, the explanation of step S70 and subsequent steps will be given after explaining the automatic start control.

In this way, the calculation control of the initial cardiac output CO ₀ at the time of manual start is performed.

Automatic Start Next, the case where the START switch 72 is not pressed in step S41 will be explained. There are two cases in which the START switch 72 is not pressed: the operator intends to perform a manual start until the START switch 72 is pressed, and the operator intends to perform an automatic start. There is a street.
Furthermore, in this embodiment, even if the START switch 72 is not pressed, if there is a predetermined change in temperature T _b ,
It is determined that the operator intends automatic start. That is, until a predetermined change in temperature T _b is detected while the START switch 72 is not pressed,
While monitoring this temperature change, it is also necessary to monitor the depression of the switch 72 at the same time. This temperature monitoring is performed after step S43, and whether the switch 72 has been pressed is detected by interrupt between steps S42 and S47. If the above-mentioned interrupt occurs between steps S42 and S47, control is forcibly transferred to the above-mentioned step S50.

Now, if the switch 72 is not pressed, T _b is set according to the RC pointer 161 in step S42.
Read from RAM132. The loop of step S42 → step S43 → step S44 → step S42 is
In this embodiment, since the change in T _b is determined from the change in the moving average value of T _b (16 samples), this is to obtain the required number of samples from the memory 132 . Note that the moving amount of the moving average is one sample data. Once the required number of samples is collected, the total number of samples for averaging will be collected every time one piece of sample data is read from the memory 132. In step S45, this moving average value T _b (n) is determined. In step S46, the difference from the previous moving average value T _b (n-1) is determined. If this difference is greater than or equal to a predetermined threshold value _THd , it is determined that the blood temperature has changed due to the injection of the indicator liquid, and in step S49, the time t at the time of this change is calculated. This is because, due to the time lag caused by taking the moving average, the time when the actual temperature change occurred will be in the past than the time when it was determined that the change occurred. The subsequent control is the same as the case using the START switch 72 described above.

If it is determined in step S47 that the change is less than the threshold TH _d , then in step S48 the change is determined to be the current average value.
Move T _b (n) to the previous average value T _b (n-1) storage area.

Here, the moving average interval (in the above example, 40ms x
16 pieces = 640 ms), the time difference between T _b (n) and T _b (n-1), and the relationship between the threshold value TH _d will be explained.
In the control shown in FIG. 13A, depending on whether the temperature change with respect to time change (=temperature gradient) and the time difference between T _b (n) - T _b (n-1)/n and n-1 exceeds TH _d , Changes in the dilution curve due to heartbeat are detected. The time difference between n and n-1 is 640 ms. By the way, in order to detect changes in the dilution curve without being affected by noise, etc., TH _d is preferably about 0.1°C. For example, in a human dilution curve, it takes about 2 seconds for a 0.1°C change to appear. Therefore, the first
It is necessary to slightly change the control shown in Figure 3A. The changes are as follows. The moving average calculation of the dilution curve temperature T _b is performed every 16 samples (=640 ms) and stored in the memory. The moving average value calculated and memorized in this way is expressed as T _b (k), and the difference between it and the moving average value T _b (l) calculated and memorized approximately 2 seconds ago is 2
Divide by seconds to find the slope of the curve, and combine this with TH _d =
It is compared with 0.1°C. That is, if T _b (k) - T _b (l)/2 seconds ≧0.1, it is assumed that a temperature change due to heartbeat has been recognized. In this way, when finding the slope of the curve and determining the integration interval from the past data of about 2 seconds, t in step S49 in Fig. 13A means the past data of 2 seconds ago. .

In addition, since the moving average calculation cycle and the time interval of 2 seconds are not synchronized, in the above example, each T _b
I end up having to memorize all of (k).
Therefore, instead of the average value T _b (l), the real time value T _b two seconds ago may be used instead.

Thus, the control of integral calculations during manual start and automatic start has been explained.

Creating a table of correction coefficients Returning to the explanation of step S71. Now, as mentioned above, according to the Stewart-Hamilton equation, CO=S _i・C _i・(T _b −T _i )・V _i /S _b・C _b・∫ ^∞ / ₀
ΔT _b dt where, CO: cardiac output, S _i : specific gravity of injected liquid C _i : specific heat of injected liquid, V _i : amount of injected liquid T _i : temperature of injected liquid, T _b : temperature of blood S _b : Specific gravity of blood, C _b : Specific heat of blood. The above formula becomes CO=A・(T _b −T _i )/∫ ^∞ / ₀ ΔT _b dt. Here, A=S _i ·C _i ·V _i /S _b ·C _b . In this constant A, S _i , S _b , C _b , C _i etc. are fixed, but the amount of injected liquid V _i is that even if the indicator liquid set at the normal temperature is injected into the blood vessel, there is no It is affected by the fact that it remains and is not flushed out into the bloodstream. That is, the above V _i is only a nominal value. Furthermore, the amount of liquid remaining in the catheter becomes a problem because the amount of heat of the indicator liquid injected into the blood vessel becomes unknown.
Since the absolute value of this amount of heat also depends on the temperature of the indicator liquid, conventionally, as mentioned above, a table determined experimentally in advance is indexed based on the catheter size, the indicator liquid injection volume V _i , etc. , a correction constant A' was calculated, but in this embodiment, this correction constant is
A' is converted into a table (FIG. 12) and stored in the ROM 131. When subtracting one correction constant A' from this table, as shown in FIG. 12, the outer diameter Fr of the catheter 4 and the nominal liquid injection volume ml1 inputted from the switches 59 and 60 are used as address data. That is, the CPU 133 operates the switch 59,
60 set values are read through the I/O port 134, and these and the indicated liquid temperature T _i are stored in the ROM 131.
is converted into the above address data for addressing. The correction constant data to be stored in the above table is obtained in advance from actual measurement results obtained by varying the amount of injected liquid for catheters of various sizes. If the catheter size and injection fluid volume used in actual measurement using the thermodilution method include those that do not correspond to those in the table above, linear interpolation will be performed from the closest ones in the table.
Find the corresponding correction constant A′.

Return to the explanation of the flowchart. When the calculation of the integral of ΣΔT _b is completed up to step S68, as described above, in step S71, data is stored in the ROM 131 based on the catheter outer diameter Fr, indicated liquid volume ml, etc. inputted from the switches 59 and 60. The correction constant A' is read from the table shown in FIG. 12 (FIG. 12). In step S71, the above-mentioned Stewart
Based on Hamilton's equation, initial cardiac output value CO ₀
Calculate. In step S72, a parameter S necessary for calculating the continuous cardiac output CO _x is calculated. In step S73, these values are changed to, for example,
The data is output on the plotter 50 as shown in FIGS. 4 and 5.

<Continuous CO _x measurement> First, we will explain the principle by which CO _x can be calculated continuously. If the resistance value of thermistor 1 is _Rt , and the current value given to thermistor 1 by the constant current circuit 111 is _Ic , the amount of heat generated per unit time when thermistor 1 is heated is: _Ic2・^R It becomes _t . If the heated thermistor 1 is placed in blood with a blood flow velocity v, the heated thermistor 1 will be cooled depending on the blood flow velocity v. The amount of heat exemplified by blood flow is K・v・( _Tt - _Tb , where Tb is the temperature of the blood, _Tt is the temperature of the heated thermistor, and _K is the proportionality constant. Now, the temperature of thermistor 1 is is maintained at a temperature such that the amount of heat generated by heating and the amount of heat cooled are equal.This temperature is the thermal equilibrium temperature.

The above can be expressed as the following equation (2).

I _c ² ·R _t =K·v·(T _t −T _b )...(2) From equation (2), equation (3) for determining the blood flow velocity v is derived.

v=(1/K)( _Ic2・_Rt )/( _Tt − _Tb ⁾ ...(3) That is, the data _Rt obtained from thermistor 1,
T _t and the blood temperature obtained from thermistor 2
The blood flow velocity v can be determined from T _b . Incidentally, since the heating thermistor 1 is driven by the constant current circuit 111, the potential difference _E0 between both ends of the lead wire of the heating thermistor 1 may be detected instead of detecting the resistance value. Details of this case are described in the above-mentioned Japanese Patent Laid-Open No. 125329/1983. Furthermore, although the constant current value I _c can be handled by detecting the current of the constant current circuit 111, it is also possible to provide it in the ROM 131 as a constant term like the proportionality constant K.

Now, if the cross-sectional area of the pulmonary artery is S, then the difference between the initial cardiac output value CO ₀ and the initial blood flow velocity value v ₀ is (4)
There is a relationship as shown by the formula.

CO ₀ = S・v ₀ ...(4) Therefore, the initial cardiac output value CO ₀ and the initial blood flow velocity value v
From this, calculate the blood vessel cross-sectional area S according to equation (4), and use this value S as a calibration value (parameter) in the measurement recording circuit 1.
Hold within 30. In this way, once the parameter S is determined, the CPU 133 calculates the parameter S for the continuously measured blood flow velocity v _x .
By multiplying by , it is possible to obtain a continuous cardiac output value CO _x . That is, CO _x =S·v _x =(S/K)·(I _c ² ·R _t )/(T _t −T _b ) (5).

FIG. 14A details step S72 of FIG. 13 for calculating the parameter S. First, in step S110 of FIG. 14A, the local CPU 10
5 to heat the thermistor 1 via an optical communication path. In step S112, thermistor 1
Wait for the heating and cooling to reach thermal equilibrium. Eventually, T _t , R _t , T _{b ,} etc. are sent from the local CPU, and according to the flowchart in Figure 11,
Stored in RAM 132 in chronological order. Therefore, in step S113, data in the RAM 132 is accessed using the RC pointer 161 according to time t when a certain period of time has elapsed. In step S114 (3)
Calculate the initial blood flow velocity v ₀ based on the formula and step
In S115, the parameter S is determined from the relationship between the initial cardiac output value CO ₀ and the blood flow velocity v ₀ according to equation (4). The preparation for continuous CO _x measurements was thus completed.

Next, according to the flowchart in Figure 14B,
Explain the procedure for calculating CO _x . It is assumed that the initial cardiac output CO ₀ and the parameter S have already been measured before starting this control.

In step S80, it is checked whether the measuring device 100 is in continuous mode. CONTINUOUS to this mode
The transition is made by pressing the switch 64. If it is continuous mode, in step S81, local CPU1
05 to heat the thermistor 1 via an optical communication path. In step S82, thermistor 1
Wait for the heating and cooling to reach thermal equilibrium. Eventually, T _t , R _t , T _{b ,} etc. are sent from the local CPU, and according to the flowchart in Figure 11,
Stored in RAM 132 in chronological order. Therefore, data in the RAM 132 is accessed using the RC pointer 161 according to time t when a certain period of time has elapsed. In step S84, the blood flow velocity is determined based on equation (3).
v Calculate _x .

In step S86, it is confirmed whether continuous mode has been canceled. This is because the parameter S representing the blood vessel cross-sectional area usually changes over time. Therefore, even if the blood vessel cross-sectional area S is once held as a parameter, changes in the blood vessel cross-sectional area S may make it impossible to obtain an accurate cardiac output. Therefore, the initial cardiac output CO is appropriately measured using the thermodilution method in preparation for the next continuous measurement of CO _x .

Step S87 unless continuous mode is canceled.
Proceed to and calculate CO _x =v _x・S. In step S88, v _x and _CO
Calculate the average value for 1.5 minutes. These values are
It is stored in the area 132c in the RAM 132. and,
In step S90, in order to calculate the next CO _x ,
Increment the RC pointer by 1.

In this way, each time T _t , R _t , T _b , etc. are sent,
CO _x , etc. can be calculated continuously. In this way, the cardiac output CO _x could be measured by measuring only the blood flow velocity v _x .

<Data Recording> In addition to recording thermodilution curves and _{thermodilution} values as shown in Figs. Continuous recording function to output (by switch 56), past 30
A memory playback function (by switch 55) that outputs the measured values for each minute to the plotter 50, a progress chart output function (by the switch 54) that outputs the measured values for the past 12 hours (average value every 3 minutes) to the plotter 50, etc. It has an output function.

Output formats for continuous recording and recording playback are specified in Section 6.
Figure (Paper feed speed 10ms/min), Figure 7 (20ms/min)
As shown in (time).

Figure 8 shows the output format after the elapse of time. In this progress chart, CO (=CO _x ) and blood temperature BT (=
Record the value of T _b ) along with the graph.

Now, since such recording is performed in continuous mode (when the CONTINUOUS indicator 64 is lit), in parallel with output to the plotter 50, the local CPU 105
It is necessary to import data from the computer and perform calculations/memories such as CO _x . For this control,
To access the RAM 132, as shown in FIG. 10B, the PC pointer 162 is sent to the plotter output.
For data acquisition, SC pointer 160 is used for calculation/
An RC pointer 160 is used for storage.

<Substitute for Indicator Liquid Temperature> Since the indicator liquid injected into the blood vessel from the catheter 4 is placed in an ice-cooled container, its temperature can be said to be almost determined by the ice-cooling temperature. Therefore, when this ice-cooling temperature is known (for example, zero degrees), it can be regarded as the indicated liquid temperature.

For this purpose, in this measuring device, the temperature probe 12
is not connected, it detects this and changes the temperature set by switch 75 to the liquid temperature.
It is regarded as T _i and used to calculate the initial cardiac output CO ₀ . Therefore, since the thermistor 12a of the temperature probe 12 is connected to the constant voltage circuit 101, the injected liquid temperature detection circuit 102 detects the current flowing through the thermistor 12a to be zero. Corresponding to this zero current
When the local CPU 105 receives E _i , the local CPU sends a signal to that effect to the main CPU 133 . Alternatively, when converting E _i for zero current in the CPU 105 to temperature T _i , it is converted to an impossible value and sent to the main CPU 133. These allow the CPU 133 to detect that the temperature probe 12 is not connected.

FIG. 15 shows this control procedure. In Figure 13,
Instead of step S52 in which the main CPU 133 reads T _i from the RAM 132, in step S100 of FIG. 15, it is determined whether the T _i sent from the local CPU 105 is data indicating that the probe 12 is not connected. to judge. If the data indicates the connection status, in step S101, this RAM 132
T _i within is used for measurement. If not connected, in step S102 the data set in switch 75 is transferred.
Import as T _i .

<Measurement of relative change in cardiac output> The embodiment of the measuring device described above actually determines the initial cardiac output CO ₀ by the thermodilution method, and measures this CO ₀
The parameter S is determined from the relationship between the initial blood flow velocity v ₀ and the cardiac output CO _x can be determined continuously by simply measuring the blood flow velocity v _x at any time. . That is, CO _x ∝v _x , and changes in v _x reflect relative changes in CO _x . Therefore, if we record the change in v _x , we are directly recording the relative change in CO _x . On the other hand, if the above-mentioned cardiac output CO is the same for each subject and the measurer is familiar with the measurement, its approximate value is known. Therefore, in the measuring device of this embodiment, this approximate cardiac output CO is measured by the switch 62.
input as the "initial CAL value" (=CO ₀ ), and internally set the parameter S based on this CO ₀ .
By calculating this, it is also possible to omit the measurement of the initial cardiac output value CO ₀ by thermodilution.

<Experimental results> The measurement device shown in Fig. 9A and Fig. 9B was
Connecting the catheter 4 shown in the figure, this measurement system was used for one experiment by applying it to a circulation model circuit, and one for an animal experiment using an anesthetized dog (an adult mongrel dog weighing 15 kg). The former model circuit is composed of a constant temperature bath, a pump, etc. In the latter animal experiment, catheter 4 was inserted into the right internal jugular vein of the dog.
was inserted and advanced into the pulmonary artery. An electromagnetic blood flow meter was used to compare the results of this measurement system. FIG. 19 shows the measurement results obtained with the circulation model circuit. In Figure 19, the horizontal axis (X-axis) is the cardiac output (unit: L/min) obtained using the electromagnetic blood flow meter, and the vertical axis (Y-axis) is the cardiac output obtained using this measurement system. For cardiac output, the correlation is Y = 0.89X + 0.82, and the correlation coefficient is r = 0.966 (number of samples = 18
). On the other hand, in the results of animal experiments, the correlation is Y = 1.48X - 0.78, and the correlation coefficient is r = 0.944 (number of samples = 29
). In either result, 0 to 4 L/min
Good results were obtained in this area. Also, 0~
Good results were confirmed even in the high cardiac output region of 10 L/min, as shown in Figure 19. 0~10L/
Good results were obtained in the high central output region of min.
This is probably due to the use of a 14-bit A/D converter 104.

<Application to other dilution methods> As is clear from the above explanation, the continuous cardiac output measurement in the above-mentioned embodiment is performed once the initial cardiac output CO is measured by some method. If so, then the parameter S is measured and held, and thereafter the cardiac output CO _x is continuously determined by simply measuring the blood flow velocity v _x . Therefore, the present invention is not limited to obtaining the initial cardiac output CO by thermodilution measurement as in the above embodiment.
For example, the initial cardiac output CO ₀ can also be obtained by a dye dilution method, an electrolyte dilution method, or the like. In this case, in the dye dilution method, the amount of pigment in the blood is measured by measuring changes in illuminance, such as at the earlobe, and in the electrolyte dilution method, changes in blood resistance are measured using two electrodes installed in a catheter. Obtain the cardiac output CO.

That is, the function of inputting the amount of data (for example, Fr, ml) that is the basis of the correction constant A' can also be applied to general indicator dilution methods using a catheter. Also,
The automatic measurement start function can be directly applied to detecting the change points of the above-mentioned changes in illuminance and changes in resistance value. In addition, the continuous cardiac output measurement function is based on the initial cardiac output obtained by the general indicator dilution method.
It is also applicable to obtain the parameter S based on CO ₀ and then calculate CO _x from the blood flow velocity v _x at any time.

Furthermore, as other variations and modifications, the temperature probe 12
We propose the following mechanism to detect whether or not the That is, a biased protrusion is provided at a portion of the main body of the measuring device to which the connector 16 of the probe is connected. When the connector 16 is connected, this protrusion is pushed and retracted. When the protrusion is pressed, the end of the protrusion is arranged to provide a ground signal to the measurement circuit 120, for example. The measurement circuit 120 can determine whether the probe 12 is connected or disconnected based on the presence or absence of this ground signal. Note that the protrusion is electrically insulated from the circuit 120 except for its ends. This is to keep the entire circuit 120 electrically connected.

As another modification, blood flow rate measurement using thermistor 1 not only detects thermal equilibrium under constant current, that is, changes in thermistor resistance using voltage, etc., but also detects the temperature difference necessary to keep the temperature difference from blood temperature constant. It is also possible to extend the current and measure the current. In other words, the temperature of the thermistor is controlled at the upper limit of 42°C, which does not affect living organisms, and the current is measured. Further, the blood flow rate sensor is not limited to a thermistor, and other means may be used.

If the cardiac output measuring device according to the embodiment described above is viewed from the viewpoint of a measuring device having a function of automatically detecting the starting point of integration, the device according to this embodiment has the following features:
Thermistors 1 and 2 and detection circuits 112 and 113 as blood data detection means, and changes in blood data detected by the blood data detection means at a first time and blood data at a second time after the first time. Calculate the amount, compare whether the calculated amount of change over time is greater than or equal to a predetermined value, and if it is determined that the amount of change over time is greater than or equal to the predetermined value, calculate to start integrating the blood data. It includes a CPU 133 as means, a comparing means, and an integrating means, and a CPU 133 as a cardiac output calculating means for calculating the cardiac output based on the integral value.

In particular, in step S46, the CPU 133 determines whether the difference between the moving average value Tb(n) at a certain time and the moving average value Tb(n-1) a predetermined time ago has become larger than the threshold value THd. Tb(n-1) can be considered to be "blood data at the first time" and Tb(n) can be considered to be "blood data at the second time after the first time."

Furthermore, step S54 in FIG.
When it is determined that {Tb(n-1)-Tb(n)}) is equal to or greater than a predetermined value, means for determining reference blood data Tb ₀ based on blood data values before the first time is configured. do.

Furthermore, updating ΔTbc in step S63 corresponds to "means for determining an integration interval based on the peak value to end the integration."

[Effects of the Invention] As explained above, according to the cardiac output measuring device equipped with the automatic measurement start function of the present invention, the cardiac output measuring device measures cardiac output based on the dilution method. A blood data detection means for outputting data related to the blood diluted by the indicator liquid injected into the blood vessel, and a first blood data detection means detected by the blood data detection means.
and a change amount calculating means for calculating the amount of change between the blood data at the time and the blood data at the second time after the first time, and comparing whether or not the calculated time change amount is equal to or greater than a predetermined value. a comparing means; an integrating means that starts integrating the blood data when it is determined that the temporal change amount calculated by the comparing means is equal to or greater than a predetermined value; and an integrating means that calculates cardiac output based on the integrated value. It is characterized by comprising a cardiac output calculation means.

In the dilution method using an indicator, it is important to integrate the blood data value from the time when the blood data value changes due to the injection of the indicator. is monitored in real time by a comparison means, and if this change satisfies a predetermined condition, integration is started, so the operator does not have to worry about synchronizing the start of indicator injection and the start of measurement. Furthermore, because of the real-time nature of the change point detection operation, the waste of storing meaningless measurement data before the start of injection is also eliminated.

According to one aspect of the invention, said blood data sensing means comprises a temperature sensor, said blood data being data related to the temperature of blood diluted by an indicator liquid injected into a blood vessel.

According to one aspect of the invention, the integrating means not only has the function of detecting the starting point of integration;
Furthermore, it is also possible to detect the end point.

According to one aspect of the present invention, the end point of the integration by the integrating means is predicted from the peak value, so that the calculation of the integration, that is, the cardiac output can be efficiently measured without reducing accuracy.

According to one aspect of the present invention, the change calculating means calculates a difference between moving average values of blood data. According to this configuration, the integration start point can be detected with high accuracy without being affected by noise, and therefore the measurement accuracy of cardiac output is improved.

[Brief explanation of the drawing]

1A, 1B, and 1C are a plan view, a front view, a back view, and a second view of the measuring device according to the embodiment, respectively.
Figures A, 2B, and 2C are an overall perspective view, a sectional view of a main part, and a sectional view of an opening, respectively, of a catheter used in the measuring device of the embodiment, and Figures 3A and 3B are a diagram of the measuring device of the embodiment. Procedure diagram explaining the operating procedure, Part 4
Figure 5 is a diagram showing an example of the result output when cardiac output is measured based on the thermodilution method using the embodiment device, and Figures 6 to 8 are output results of continuous measurement. Figures 9A and 9B are circuit diagrams of the main parts of the measuring device, Figures 10A and 10B are diagrams explaining the RAM management within the measuring device, and Figure 11 , FIG. 13A, FIG. 13B,
14A, 14B, and 15 are flowcharts showing the control procedure according to the embodiment, and FIG. 12 is a diagram showing a table storing the correction constant A.
Figures 6 and 17 are diagrams explaining the principle of cardiac output measurement using the conventional thermodilution method, and Figure 18 is a diagram explaining the operation for automatically starting cardiac output measurement.
FIG. 19 is a diagram showing the results of an experiment conducted using the apparatus of the embodiment. In the figure, 1, 2, 12a... thermistor, 4...
Catheter, 6, 8, 10, 13, 14... tube, 7, 9, 11, 15, 16... connector,
12...Injection liquid temperature probe, 50...Protuter, 52, 53, 54, 55, 56, 57, 58
...Switch, 59, 60, 61, 62, 75...
...Setting input switch, 64, 65, 66, 67,
70, 72...Switch/indicator, 68
...4-digit LED display, 69...3-digit LED display,
71... LED bar graph, 73... Catheter connection connector, 74... Temperature probe connection connector, 100... Measuring device, 101, 112... Constant voltage circuit, 111... Constant current circuit, 102... Infusate temperature Detection circuit, 103...Analog switch, 104...A/D converter, 105, 133...
...CPU, 106,131...ROM, 107,1
32...RAM.

Claims (1)

[Scope of Claims] 1. In a cardiac output measuring device that measures cardiac output based on a dilution method, blood data outputs data related to blood diluted by an indicator liquid injected into a blood vessel. a detection means; a change amount calculation means for calculating the amount of change between the blood data detected by the blood data detection means at a first time and the blood data at a second time after the first time; a comparison means for comparing whether the amount of change over time is greater than or equal to a predetermined value; and an integration means for starting integration of the blood data when it is determined that the amount of change over time calculated by the comparison means is greater than or equal to the predetermined value; A cardiac output measuring device comprising an automatic measurement start function and a cardiac output calculating means for calculating the cardiac output based on the integral value. 2. The first aspect of claim 1, wherein the blood data detection means includes a temperature sensor, and the blood data is data related to the temperature of blood diluted by an indicator liquid injected into a blood vessel. A cardiac output measuring device equipped with an automatic measurement start function described in . 3. The integrating means is configured to calculate the first and second values calculated by the comparing means.
means for determining reference blood data based on blood data values before the first time when it is determined that the difference in blood data at the time is equal to or greater than a predetermined value; calculation means that starts integrating a difference value between the blood data after time 1 and the determined reference blood data, and integrates the difference value every time, and the blood data is the reference blood data. 2. The cardiac output measuring device with an automatic measurement start function according to claim 1, further comprising a stop means for terminating the integration when substantially equal to . 4. The integrating means is configured to calculate the first and second values calculated by the comparing means.
means for determining reference blood data based on blood data values before the first time when it is determined that the difference in blood data at the time is equal to or greater than a predetermined value; a calculation means that starts integrating a difference value between the blood data after time 1 and the determined reference blood data, and integrates the difference value every time, and detects a peak value of the difference value. The device for measuring cardiac output with an automatic measurement start function according to claim 1, characterized in that the device includes means for determining an integral interval for terminating the integration based on the peak value. . 5. The change amount calculating means includes a plurality of registers for holding blood data at a plurality of times, and two average value registers, a first and a second average value register for holding a moving average value of the blood data, Blood data sampled at a plurality of times in the vicinity of the first time is held in the plurality of registers, and a first moving average value at the first time is calculated from the blood data held in these registers. , this first moving average value is held in the first average value register, blood data sampled at a plurality of times in the vicinity of the second time is held in the plurality of registers, and blood data sampled at a plurality of times in the vicinity of the second time is held in these registers. calculates a second moving average value at the second time from the blood data obtained, and stores this second moving average value in the second average value register; 2. The cardiac output measuring device having an automatic measurement start function according to claim 1, wherein the difference between the first and second moving average values stored in the average value register of 2 is calculated. measuring device.