JP2000037371A - Blood light absorptive material concentration measuring instrument - Google Patents

Abstract

PROBLEM TO BE SOLVED: To lessen the delay of measurement and to facilitate calculation by changing the weighting subtraction waveform between both waveforms of prescribed two wavelengths of a beam attenuation degree change component, determining the power of the weighting subtraction waveform, determining the change component ratio of the beam attenuation degree of the tissue in accordance with this power and calculating the concn. ratio of blood light absorptive materials in accordance with this change component ratio. SOLUTION: A change component ratio Φ calculation section 11 of the beam attenuation degree of the tissue including blood consists of a beam attenuation degree change component detecting means A which determines the change component ΔA2, ΔA3 of the beam attenuation degree of the tissue transmitted light of the two waveforms, a weighting subtraction waveform calculating section B which determines the respective Ks of the weighting subtraction waveform a ΔA32=ΔA3(1-K)-ΔA2K between both waveforms determined by this detecting means A by changing the Ks, a weighting value detecting means C which determines the power of the weighting subtraction waveform determined by this calculating means B with respect to the respective Ks and determines the value of the K to minimize (maximize) the power and a Φ calculating means D which determines Φ=K/(1-K) determined by this detecting means C. When the power attains the minimal value (maximal value), Φa(Φv)=K/(1-K) is determined and the concn. ratio of the plural light absorptive materials in the arterial blood is calculated. The concn. ratio of the plural light absorptive materials in the venous blood is calculated from Φv.

Description

DETAILED DESCRIPTION OF THE INVENTION

[0001]

BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention relates to a blood oximeter for continuously measuring the concentration ratio of a plurality of light-absorbing substances in blood based on the pulsation of transmitted light or reflected light of a living tissue, such as a pulse oximeter. The present invention relates to a device for measuring the concentration of a medium light absorbing substance. The present invention is also generally used for measuring a dye dilution curve by measuring a blood dye concentration, and for measuring other light-absorbing substances in blood.

[0002]

2. Description of the Related Art A conventional apparatus of this kind has a problem that a pulse wave is disturbed by a body motion and a measurement error occurs. In addition, the pulse wave is small, and there is a problem that a measurement error occurs due to external noise and noise due to the device itself. This has become a major problem as the applications and applications of this type of device have expanded.

As a countermeasure for these problems, there is (1) a method of correcting a Φ column, but its effect is limited. (2) E
Although there is a method of averaging pulse waves in synchronization with CG, there is a drawback that simultaneous measurement of ECG is required, and the measurement result is delayed. (3) There is a so-called adaptive-filter system, but the operation is complicated and it is impossible to respond immediately.

[0004]

SUMMARY OF THE INVENTION The present invention has been made in view of the above-mentioned drawbacks of the conventional body movement countermeasures. The object of the present invention is to provide a measurement apparatus of this kind which has a small measurement delay and is easy to calculate. An object of the present invention is to provide a device having a body movement countermeasure that can be easily applied to multiple wavelengths and multiple components.

[0005]

Here, two wavelengths λ2 and λ
Let us consider the case of irradiating the living tissue with light 3. The logarithm of the tissue transmitted light L2, L3 is determined. ΔA2≡ΔLogL2 (1) ΔA3≡ΔLogL3 (2) is a change in the dimming degree of the tissue containing blood. In the absence of motion, these are all components of arterial blood. If the component of arterial blood is indicated by suffix a, then ΔA2 = ΔAa2 (3) ΔA3 = ΔAa3 (4) The ratio between the two is defined as follows. Φa≡ΔAa3 / ΔAa2 (5)

The measurement of the concentration ratio of the light-absorbing substance in arterial blood by this type of measuring apparatus, for example, the measurement of the oxygen saturation of arterial blood and the measurement of the dye dilution curve are performed based on this Φ. Therefore, if an artifact or noise is added to ΔA2 or ΔA3, information on the target arterial blood is disturbed, and an error occurs in the measurement.

The causes of the artifacts are as follows. (1) Optical system distortion: This is due to a change in the positional relationship between light emission and light reception in the probe. The variation ratio of the transmitted light due to this is common to all wavelengths. This can be greatly reduced if the contact between the probe and the living body is improved by, for example, a double-sided adhesive tape.

(2) Variation in venous blood thickness: The cause of body motion artifacts is acceleration applied to the body. Since the venous blood vessels are soft and the venous blood is less filled, the venous blood in the blood vessels is easily moved by the acceleration. This causes artefacts. This is because, in the measurement of oxygen saturation, the arterial blood and the venous blood have different oxygen saturations, and in the measurement of the pigment dilution curve, the blood pigment concentration differs between the arterial blood and the venous blood. . Here, the component of the venous blood is indicated by a suffix v, the change in the degree of tissue dimming due to the change in the amount of venous blood in the tissue is defined as ΔAv2, ΔAv3, and the ratio between the two is defined as Φv≡ΔAv3 / ΔAv2 (6).

(3) Variation in thickness of arterial blood: Since arterial blood has a high degree of filling in blood vessels, it does not move much due to acceleration. Moreover, even if the pulse wave waveform is distorted due to body movement, it is easy to prevent the Φ calculation from being affected. For example, if a regression line between pulse waves of two wavelengths is obtained, the slope Φa = ΔAa3 / ΔAa2 (7) is obtained.

(4) Fluctuation in thickness of tissue other than blood: In the movement of blood in the tissue, fluctuation in the thickness of tissue other than blood naturally occurs, which distorts the pulse wave. However, the effect is small since the light absorption is small compared to blood. In a pulse oximeter, the influence of tissues other than blood can be reduced by using light of three wavelengths. At this time, the influence of variations in the thickness of tissue other than blood as an artifact is also reduced.

(5) Noise: Noise due to the inside of the apparatus and the surrounding environment. The present invention has been made in order not to be affected by such artifacts or noise, and its principle will be described below.

When the fluctuation of the venous blood is superimposed on the pulsation of the arterial blood, the following can be written from the equations (5) and (6). ΔA2 = ΔAa2 + ΔAv2 (8) ΔA3 = ΔAa3 + ΔAv3 = Φa ΔAa2 + Φv ΔAv2 (9) It is defined as follows. ΔA32 ≡ΔA3 (1-K) -ΔA2K = [Φa ΔAa2 + Φv ΔAv2] (1-K)-[ΔAa2 + ΔAv2] K = ΔAa2 [Φa (1-K) -K] + ΔAv2 [Φv (1-K ) -K] (10) Let K = 0 to 1. Thus, ΔA32 = ΔA3 to −ΔA2. Defining M ≡K / (1-K) (11), M = 0 to at K = 0 to 1. (1
The expression (0) is as follows. ΔA32 = [ΔAa2 (Φa-M) + ΔAv2 (Φv-M)] (1-K) (12)

If M = Φa, ΔA32 = ΔAv2 (Φv−Φa) (1-K) (13) This is the waveform of venous blood. If M = Φv, ΔA32 = ΔAa2 (Φa−Φv) (1-K) (14) This is the waveform of arterial blood.

Assuming that λ2 = 900 nm and λ3 = 660 nm, Φ
a <Φv, so if M = Φa, ΔA32>0; if there are no artifacts, this value is 0 if M = Φv, ΔA32 <0

ΔA32 decreases as K increases from zero. If K continues to increase and M = Φa, ΔA32 = ΔAv2 (Φv-Φa) (1-K)> 0 (15) This is the venous blood waveform, and if there is no fluctuation in venous blood, This is general noise.

If K continues to increase and M = Φv, ΔA32 = ΔAa2 (Φa−Φv) (1-K) <0 (16) This is the waveform of arterial blood.

The magnitude relationship between ΔAv2 and ΔAa2 in the above equation is that if the magnitude of the artifact is small, ΔAv2 <ΔAa2,
If the size of the artifact is large, then ΔAv2> ΔAa2. The polarity of ΔA32 is reversed on the way from M = Φa to M = Φv.

Therefore, the change in the power of ΔA32 with an increase in K is (1) minimum at M = Φa, and (2) M = Φa
In Φv, if the artifact is small, it is maximal, and if the artifact is large, it is minimal. In any case, the first minimum is when M = Φa. However, this is based on the above-mentioned wavelength selection.

Based on the above considerations, the following two methods can be considered for correctly obtaining Φa of the arterial blood component when there is body movement. (1) Find the minimum point where ΔA32 has no arterial blood component, and obtain M = Φa in that case. (2) Find the minimum point where ΔA32 has no arterial blood component, and analyze the frequency of ΔA32 in that case, and analyze the frequency of the original waveform ΔA2 or ΔA3. Based on the comparison between the two, noise and artifacts are determined. , The original waveform is passed through a filter whose pass band is determined based on the frequency component, thereby improving the SN of the pulse wave. However, in order to compare the results of the frequency analysis of different ΔA32, the magnitude of the power needs to be a standardized value. The standardization method is, for example, that the highest frequency band in the frequency band is a noise component and does not depend on the optical wavelength of a signal. I do.

Regardless of the number of optical wavelengths, if the pass band of the filter is determined using appropriate two wavelengths, the change in the dimming degree of the other wavelengths can also be passed through this filter to cause artifacts and other noises. Is removed.

When there is a body movement, the power of ΔA32 becomes M =
Since K takes a local maximum or a local minimum at Φv, M = Φv
Can be detected. From this Φv, information on venous blood can be obtained. For example, in measuring oxygen saturation,
The oxygen saturation of venous blood can be determined. This is useful as a parameter indicating the supply and demand of oxygen in the tissue. In addition, for example, in the measurement of a dye dilution curve, a pigment movement curve in a vein can be simultaneously measured by intentionally adding a body motion. Based on the above principle, the present invention has the following configuration.

According to a first aspect of the present invention, a concentration ratio of a plurality of light-absorbing substances in blood is measured based on a pulsation of a tissue transmitted light or a reflected light obtained by irradiating a living tissue with light of a plurality of wavelengths. In the blood-absorbing substance concentration measuring device, a dimming degree change detecting means for obtaining a dimming degree change of the tissue transmitted light or the reflected light, and a predetermined two wavelengths of the dimming degree change obtained by the dimming degree change detecting means. Weighted subtraction waveform Δ between both ΔA2 and ΔA3 waveforms
A32 = ΔA3 (1-K) -ΔA2K
Weighted subtraction waveform calculation means for calculating the weighted subtraction waveform power obtained by the weighted subtraction waveform calculation means for each K, and a weight value detection means for obtaining the value of K at which the power takes the minimum or maximum, and the weight value Φ calculating means for obtaining Φ = K / (1-K) from the value of K obtained by the detecting means, and concentration ratio calculating means for calculating the concentration ratio of the light absorbing substance in blood based on Φ obtained by the Φ calculating means. And characterized in that:

According to a second aspect of the present invention, the concentration ratio of a plurality of light-absorbing substances in blood is measured based on the pulsation of tissue transmitted light or reflected light obtained by irradiating a living tissue with light of a plurality of wavelengths. In the blood-absorbing substance concentration measuring device, a dimming degree change detecting means for obtaining a dimming degree change of the tissue transmitted light or the reflected light, and a predetermined two wavelengths of the dimming degree change obtained by the dimming degree change detecting means. Weighted subtraction waveform Δ between both ΔA2 and ΔA3 waveforms
A32 = ΔA3 (1-K) -ΔA2K
Weighted subtraction waveform calculation means for determining the power of each weighted subtraction waveform determined by the weighted subtraction waveform calculation means for each K, and a weight value detection means for determining the value of K at which the power is minimum or maximum, A frequency analysis is performed on a filter that passes the change in the dimming degree obtained by the change detection means and a weighted subtraction waveform at the value of K obtained by the weight detection means, and the pass band of the filter is determined based on the analysis result. A pass band determining means for determining, a Φ calculating means for obtaining a ratio Φ of a change in the dimming degree of each of the two outputs of the filter, and a concentration ratio of a light absorbing substance in blood based on Φ obtained by the Φ calculating means. And a concentration ratio calculating means.

According to a third aspect of the present invention, in the device according to the second aspect, the passband determining means performs frequency analysis of the original waveforms of ΔA2 and ΔA3 obtained by the dimming degree change detecting means,
In addition, the frequency analysis of the weighted subtraction waveform at K at which the power becomes minimum or maximum is performed, the two are compared, and the pass band is determined based on this.

According to a fourth aspect of the present invention, in the apparatus according to the second or third aspect, in the frequency analysis, a moving average is taken for each fixed number of samples in a fixed section, and a difference waveform between two moving average waveforms having different numbers of samples is obtained. , And the power of the band is calculated based on this.

According to a fifth aspect of the present invention, in the apparatus according to the second aspect, the filter takes a moving average for each of a fixed number of samples in a fixed interval, and calculates a difference between two moving average waveforms having no difference or different numbers of samples. It is characterized by obtaining a waveform.

According to a sixth aspect of the present invention, in the device according to the second aspect, the wavelength used is three or more wavelengths, and the pass band determining means determines the pass band of the filter based on the two wavelengths. The filter of this pass band is characterized by passing a change in dimming degree for wavelengths other than ΔA2 and ΔA3.

According to a seventh aspect of the present invention, in the apparatus according to the first or second aspect, the weight value detecting means sets the value of K obtained for a section when the present apparatus performs continuous measurement to Ka as follows. In the section, the power of the weighted subtraction waveform near Ka is obtained, the value of K at the minimum or maximum power is obtained, and this is set as a new Ka.

According to an eighth aspect of the present invention, in the device according to the first or second aspect, the weight value detecting means obtains power for a plurality of weighted subtraction waveforms for each of the plurality of K values, and based on this, Power is minimal or maximal
It is characterized in that the value of K is estimated.

The invention of claim 9 is the invention of claim 1, 2 or 7
Wherein the concentration ratio of the blood light-absorbing substance determined by the concentration ratio calculating means is an arterial blood oxygen saturation or another blood light-absorbing substance concentration ratio.

According to a tenth aspect of the present invention, in the apparatus according to the first, second or seventh aspect, the concentration ratio of the blood light-absorbing substance determined by the concentration ratio calculating means is the venous blood oxygen saturation or other blood absorption ratio. It is characterized by a substance concentration ratio.

[0032]

FIG. 1 (a) shows an example of the overall configuration of a pulse oximeter according to the present invention. This device includes a probe 1 and a device main body 2.

The probe 1 has a light irradiating section 3 for generating light.
And a light receiving unit 4. The light irradiator 3 includes a light source 5 composed of an LED. The two LEDs emit light of different wavelengths λ2 and λ3, respectively. The light irradiating unit 3 and the light receiving unit 4 are arranged so as to face each other via the living tissue.
The light receiving unit 4 includes a photodiode, receives light from the light irradiation unit 3, and converts the light into an electric signal.

The apparatus main body 2 includes an analog processing unit 7 and an A / D
It includes a conversion unit 8, a light source driving unit 9, and a digital processing unit 10. The analog processing unit 7 is a circuit that removes noise and amplifies the output signal of the light receiving unit 4, and the A / D conversion unit 8
Is a circuit for converting an output signal of the analog processing unit 7 into a digital signal. The light source driving section 9 is a circuit for driving the LED of the light source 5. The digital processing unit 10 is constituted by a computer and exchanges signals with a CPU (Central Processing Unit) that performs arithmetic control, a memory in which a processing program is written, and data in which data necessary for processing is written, and an external signal. It has an input / output interface for performing. This digital processing unit 10 is represented by functional blocks as follows:
Φ calculation unit 11 for calculating Φ from the output of A / D conversion unit 8,
It comprises a converter 12 for obtaining the oxygen saturation SaO2 of arterial blood based on the Φ calculated by the Φ calculator 11, and a controller 13 for controlling the entire apparatus. The value of SaO2 thus obtained is usually described as SpO2.

The Φ calculation unit 11 is described in more detail in FIG.
The result is as shown in FIG. That is, the dimming degree change detecting means A for obtaining the dimming degree changes ΔA2 and ΔA3 of the tissue transmitted light of two wavelengths, and the ΔA obtained by the dimming degree change detecting means A
2, Weighted subtraction waveform between both waveforms of ΔA3 ΔA32 = ΔA3 (1-K) -ΔA
The weighted subtraction waveform calculating means B for obtaining K for each K by changing K, and the power of the weighted subtraction waveform obtained by the weighted subtraction waveform calculating means B for each K to obtain 2K Φ = K / (1-K) from the weight value detecting means C for obtaining the value and the value of K obtained by the weight value detecting means C.
And Φ calculating means D for determining

Next, the operation of the thus constructed apparatus will be described. FIG. 2 is a flowchart of a process performed by the digital processing unit 10. Description will be made with reference to this figure.

When the operation of the apparatus starts, the light irradiating section 3 irradiates the living tissue with light. And the digital processing unit 10
Starts fetching data from the A / D converter 8 (step 101). Here we have two wavelengths λ2, λ
3 is transmitted light data of a living tissue by light.

Next, the logarithm of the transmitted light L2, L3 at each of the two wavelengths is taken, and the change ΔA in the degree of extinction of the tissue is obtained (step 102). That is, ΔA3 = ΔlogL3 (17) ΔA2 = ΔlogL2 (18) is obtained.

Next, K is set to the minimum value (step 1).
03). For example, if K is changed in the range of 0 to 1, K = 0.

Next, ΔA32 = Δ at the set value of K
A3 (1-K) -ΔA2K is obtained (step 104).

Next, the obtained ΔA32 is passed through a predetermined filter to remove its low-frequency component and high-frequency component (step 105). This is to eliminate obvious noise. For example, the filter may take the difference between the moving average waveform at 10 points and the moving average waveform at 60 points.

Next, the power of ΔA32 is obtained, and the result is recorded (step 106). The power is determined as follows. ΔA32 power = ΣΔA32 ² (20) Next, K is added to the current K by a predetermined change ΔK, and K = K + ΔK
(Step 107).

Next, it is determined whether the value of K has reached the maximum value (step 108). That is, in the above example, it is determined whether or not K = 1. If the value of K has not reached the maximum value yet, the process returns to step 104. If the value of K is the maximum value, the relationship between K and power as shown in FIG. 3 is obtained. Thereafter, the processing of steps 109 to 111 and the processing of steps 112 to 114 are performed.

In step 109, the value of K when the power of ΔA32 first takes the minimum value (when K is increased from 0)
Is determined (see FIG. 3). In step 110, Φa = K / (1-K) is obtained from the K, and in step 111, the obtained Φa
The oxygen saturation SaO2 of the arterial blood is determined from this.

When ignoring individual differences in tissue thickness variation, the oxygen saturation SaO2 of arterial blood can be obtained from two wavelengths. Φa is expressed by the following theoretical formula, and Eo, Er,
This is because F is known (see Japanese Patent Application No. 10-23048). Therefore, Sa can be obtained by substituting Φa obtained in the following equation. Φa = ΔAa3 / ΔAa2 = [{Ea3 (Ea3 + F)} ^1/2 -Ex3] / [{Ea2 (Ea2 + F)} ^1/2 -Ex2] (21) where Ea3 = SaEo3 + (1-Sa ) Er3 (22) Ea2 = SaEo2 + (1-Sa) Er2 (23) Eo2, Eo3; extinction coefficient of oxyhemoglobin at wavelengths λ2, λ3. Er2, Er3; extinction coefficient of reduced hemoglobin at wavelengths λ2, λ3. F; scattering coefficient. Sa; SaO2 (oxygen saturation of arterial blood) Ex2, Ex3; tissue terms at wavelengths λ2, λ3 (terms representing the effect of tissue) If individual differences in pulsation of tissue thickness are ignored, these should be real numbers.

On the other hand, at step 112, K at which the power of ΔA32 takes the maximum value is determined (see FIG. 3), and step 11 is performed.
In step 3, Φv = K / (1-K) is obtained from K. In this example, K where the power takes the maximum value is determined because the artifact is small, but K where the power takes the minimum value is determined when the artifact is large. In any case, Φa = K / (1
-K) K which is larger than K at the time of and takes an extreme value is obtained. In step 114, the oxygen saturation SvO2 of the venous blood is obtained from the obtained Φv. To find SvO2 from Φv:

As in the case of the equation (21), Φv is represented by the following theoretical equation. Since Eo, Er, F are known, Sv is obtained by substituting Φv into this equation. Φv = ΔAv3 / ΔAv2 = [{Ev3 (Ev3 + F)} ^1/2 -Ex3] / [{Ev2 (Ev2 + F)} ^1/2 -Ex2] (24) where Ev3 = SvEo3 + (1-Sv ) Er3 (25) Ea2 = SvEo2 + (1-Sv) Er2 (26) Sv; SvO2 (oxygen saturation of venous blood) Other symbols are as described above.

Next, an actual measurement example according to the above embodiment will be described. FIG. 4 shows an original waveform obtained by logarithmically converting two wavelengths of tissue transmitted light, which is data in the section 117-126 sec. This waveform shows changes ΔA2 and ΔA3 in the degree of extinction of the tissue (step 10).
1-102). The probe was for a fingertip, and the probe was attached to the fingertip of the subject with a double-sided adhesive tape. The light wavelength is λ2 =
900 nm and λ3 = 660 nm. During the measurement, the oxygen saturation of the arterial blood was kept almost constant by air respiration. At 117.5-119sec, a periodic shock was applied to the back of the hand.

Next, when ΔA32 = ΔA3 (1-K) -ΔA2K is obtained, the actual example is as shown in FIGS. 5 and 6.

FIG. 5 is for the section 123-126 sec.
This is when there are no artifacts. The two numbers at the top center of each subgraph indicate K and 1-K. From K = 0
Draw a ΔA32 waveform in 0.1 steps until K = 1, and subgraph 1-11
Shown in Looking at this, ΔA32 between K = 0.3 and K = 0.4
Is reversed. Small graph 12 shows the details in between
-14. As can be seen, at K = 0.35, the pulse wave component has just disappeared, leaving only noise.

As described above, when the weight K in the weighted subtraction is sequentially increased, there is a point where the power of ΔA32 is minimized. In this case, when K is Φ = K / (1-K), ,
This Φ is equivalent to ΔA3 / ΔA2 of arterial blood, and is set as Φa (steps 103 to 110).

In the measurement of arterial blood oxygen saturation, this Φa is converted to arterial blood oxygen saturation SaO2, which is converted to SpO2
(Step 111). In this way,
The oxygen saturation of arterial blood can be determined. Although the section to be calculated is 3 seconds, the section may be larger or smaller than 3 seconds.

FIG. 6 is for section 117-120 sec.
This is when there are artifacts. The two numbers at the top center of each subgraph indicate K and 1-K. From K = 0
Draw a ΔA32 waveform in 0.1 steps until K = 1, and subgraph 1-11
Shown in Looking at this, ΔA3 between K = 0.3 and K = 0.4
2 is reversed. A small graph shows the details in between
12-14. Looking at this, at K = 0.35, the pulse wave component just disappeared, and only the artifact was found.

As described above, when the weight K in the weighted subtraction is sequentially increased, there is a point where the power of ΔA32 is minimized. In this case, when Φ = K / (1-K), ,
This Φ is equivalent to ΔA3 / ΔA2 of arterial blood, and is set as Φa (steps 103 to 110).

In the measurement of arterial blood oxygen saturation, this Φa is converted to arterial blood oxygen saturation SaO2, which is converted to SpO2
(Step 111). In this way,
The oxygen saturation of arterial blood can be determined.

The waveform of the small graph 15 in FIG. 6 is for the case where the power of ΔA32 is maximized. At this point, the case where the venous blood component is erased and only the arterial blood component is left. Therefore, Φ at this point is Φv corresponding to the oxygen saturation of venous blood (steps 112 to 113).

In the measurement of the venous blood oxygen saturation, this Φv is converted into the venous blood oxygen saturation SvO2, which is output (step 114). Thus, the oxygen saturation of the venous blood can be obtained.

K at which the power of ΔA32 is minimized or maximized
In order to obtain with high accuracy, the increase ΔK of K must be made sufficiently small. But it requires many computation steps. There are two methods for devising K with high accuracy and reducing the number of steps.

One is a method of obtaining Ka by increasing ΔK and then obtaining a more accurate Ka from the vicinity of Ka by reducing ΔK. The other is a relatively large ΔK as ΔK
In this method, the power of A32 is obtained, and the local maximum value of the power is estimated from the relationship between K + ΔK and the power of ΔA32.

In the above embodiment, when performing continuous measurement, steps 103 to 109 may be performed as follows. First, K corresponding to the previous Φa is stored as Ka, and in the next new section, K is sequentially changed by a predetermined width before and after Ka to obtain the power of ΔA32 for each K. The value of K having the minimum power among them is defined as Ka of the new section. By doing this,
The calculation time can be shortened, and an abnormal state when the artifact is large (the point of Φa and the point of Φv are both minimal and confused because the power takes a minimal value) can be avoided.

Also, when the object to be measured is not the oxygen saturation, but also, for example, the measurement of the blood pigment concentration in the measurement of the dye dilution curve, the point where the power becomes minimum corresponds to the arterial value Φ, that is, Φa.

Also, when multiple wavelengths are used instead of two wavelengths, Φa is determined as described above for an appropriate combination of two wavelengths, thereby obtaining a plurality of light-absorbing substances in arterial blood without the influence of artefacts. The concentration ratio can be measured (for examples using multiple wavelengths, refer to Japanese Patent Application No. 7-48
20, see Japanese Patent Application No. 10-23048).

Next, a second embodiment will be described.
The overall configuration of the device according to the present embodiment is different from the device shown in FIG. 1A in that the Φ calculation unit is different in the digital processing unit 10. FIG. 7 shows the configuration of the Φ calculation unit. As shown in this figure, the Φ calculation unit includes a dimming degree change detecting means E for obtaining the dimming degree changes ΔA2 and ΔA3 of the tissue transmitted light of two wavelengths, and a ΔA2 obtained by the dimming degree change detecting means E. , Δ
Weighted subtraction waveform between both A3 waveforms ΔA32 = ΔA3 (1-K) -ΔA2K
Is calculated by the weighted subtraction waveform calculation means F obtained by changing K 1, the filter G that passes the ΔA2 and ΔA3 obtained by the dimming degree change detection means E, and the weighted subtraction waveform calculation means F The power of the weighted subtraction waveform is obtained for each K, and a weight value detection means H for obtaining a value of K at which the power becomes a minimum, and a frequency analysis is performed on the weighted subtraction waveform at the value of K obtained by the weight value detection means H. A pass band determining means I for determining a pass band of the filter G based on the analysis result, and a ratio Φ of two outputs ΔA 2 and ΔA 3 of the filter G
And Φ calculating means J for determining

Next, the operation of the present apparatus configured as described above will be described. FIG. 8 is a flowchart of a process performed by the digital processing unit 10. Description will be made with reference to this figure.

The output of the A / D converter 8 is logarithmically converted to ΔA
2.Calculate ΔA3, weighted subtraction waveform ΔA32 = ΔA3 (1-K) -ΔA2K
Is obtained for each K by changing K, the power of the weighted subtraction waveform is obtained for each K, and the value of K at which the power becomes minimum or maximum is shown in FIG. 2 of the first embodiment. This is the same as steps 101 to 109.

Next, the frequency spectrum of ΔA32 at the value of K at which the power becomes minimum is obtained (step 12).
1). FIG. 9 shows a process for obtaining this frequency spectrum. The waveform of ΔA32 at the value of K at which the power becomes minimum is extracted (step 150). Next, the average number is set to 1 (step 151). A moving average of the set average number is obtained (step 152). Next, a difference between the moving average waveform of the previous moving average number and the current moving average waveform is obtained (step 153). Power of difference waveform found
= ΣΔA32 ² and this correction value is recorded (step 154). Next, the average number is increased by one (step 155). Next, it is determined whether the average number has reached a predetermined value Kmax (step 156). If the result of this determination is No, step 15
Returning to step 2, if Yes, the frequency spectrum analysis of ΔA32 ends (step 157).

On the other hand, the frequency spectrum of the original waveform ΔA3 is also obtained by the same processing as the processing shown in FIG. 9 (step 1).
25). Next, the spectrum of the difference between the frequency spectrum of ΔA32 and the frequency spectrum of ΔA3 is obtained (step 12).
2). From the spectrum of the difference, a wavelength region having the maximum value is determined (step 123). A filter pass band is determined from this wavelength range (step 124).

On the other hand, ΔA2 and ΔA3 are obtained from the logarithm obtained in the logarithmic conversion step 102 (steps 126 and 12).
7) Pass these through the filter of the pass band determined in step 124 (steps 128 and 129). From the filtered ΔA2 and ΔA3, Φ = ΔA3 / ΔA2 is determined (step 130), and SpO2 is calculated based on the Φ (step 131).

In the steps described above, step 1
21 will be specifically described based on actual measurements. In this example, the two wavelengths are the same as the wavelengths used in the first embodiment, and the original waveforms ΔA2 and ΔA3 are the same as those shown in FIG.

Assuming that K determined in step 109 is Ka, step 121 (steps 150 to 15 in FIG. 9) is performed.
In 7), the frequency spectrum of ΔA32 at K = Ka is obtained as follows. First, a moving average of the waveform ΔA32 is obtained.
The average number ave is as follows. ave = 1-2-3-5-9-13-17-25-33-49-65-97-129 Here, the data used is 60 samples per second. Next, a difference waveform of each moving average waveform of the adjacent average number is obtained. The name of each difference waveform is, for example, 1ave-2ave is 2av
e, 25ave-33ave should be like 33ave. Next, the power of each difference waveform is obtained. This also takes the integral of the square as described above. As a result, a relationship between the average number and the power is obtained, which is used as a frequency spectrum. Similarly, the original waveforms ΔA2, ΔA3
, A frequency spectrum is obtained. Examples are shown in FIGS. The former is a section 123-12 of the waveform shown in FIG.
The latter is the frequency spectrum of ΔA32 in the section 117-120 sec. The original waveforms ΔA2 and ΔA3 are given by ΔA3 = ΔA32 (K =
0), ΔA2 = ΔA32 (K = 1). For reference, the frequency spectrum for other K is also shown. For reference, FIGS. 12 and 13 show the power of each frequency component, where K is the horizontal axis. Thus, the change in power when K is changed can be clearly understood.

Next, the difference between the spectrum of the original waveform and the spectrum of ΔA32 (K = Ka) is determined (step 122). To find the difference, it is necessary to match the two levels somewhere. Here, the highest frequency band, 1a
The power of the frequency band of ve-2ave was matched. FIG.
FIG. 11 shows such a case. The difference spectrum is shown in FIGS.

Next, the maximum band of the difference spectrum is obtained (step 123). This is both sections 123-126sec, 117-12
It is the same at 0 sec, and the frequency band is 33ave-49ave.

Next, the original waveforms ΔA2 and ΔA3 are passed through filters (steps 128 and 129). Specifically, 33 and 49 moving averages of the original waveform are obtained, and the difference is obtained. 117-12
The result of applying this for 6 seconds is shown in FIG. 16 (step 130).

Φ is calculated from the waveform of FIG. Specifically, the slope of the regression line between ΔA2 and ΔA3 was determined every 1 second. The result is shown in FIG. Also shown for comparison are cases where Φ is obtained without passing through a filter. As shown in this figure, according to the present apparatus, disturbance of Φ due to body movement is greatly reduced.

In the present embodiment, the difference between the spectrum of the original waveform and the spectrum of ΔA32 (K = Ka) is obtained. The difference may be obtained by the ratio of the two spectra, or the band of the filter may be determined by another method. can do.

In the present embodiment, two wavelengths are used. However, if the original waveform of another wavelength is passed through a filter determined by two wavelengths even if the wavelength is multi-wavelength, artifacts and noise are similarly reduced. Removed. For this reason, devices that use multiple wavelengths, such as pulse oximeters that take into account the pulsation of tissue thickness excluding blood, dye dilution curve measurement devices, and other devices that measure the concentration ratio of light-absorbing substances in blood, are used. Extremely useful.

In the above-described apparatus, the light receiving section is arranged opposite to the light irradiating section with the living tissue interposed therebetween to receive transmitted light. However, the light irradiating section and the light receiving section are arranged on the same side of the living tissue. The same operation and effect can be obtained by a configuration that receives reflected light.

[0078]

According to the first to fifth aspects of the present invention, the calculation is easy, the measurement result can be obtained quickly, and the case where multiple wavelengths are required and the measurement of multiple components are simple. With the configuration, measurement can be performed without being affected by body movement.

According to the present invention of claims 6 to 10, in addition to the above effects, when multiple wavelengths are required, measurement can be performed with a simpler configuration.

[Brief description of the drawings]

FIG. 1 is a diagram showing a configuration of an apparatus according to a first embodiment.

FIG. 2 is a flowchart for explaining the operation of the apparatus according to the first embodiment.

FIG. 3 is a diagram showing a relationship between the power of a weighted subtraction waveform ΔA32 and a weight K;

FIG. 4 is a diagram showing an original waveform of a change in the dimming degree of light transmitted through a biological tissue.

FIG. 5 is a diagram showing a waveform of ΔA32 when the value of K is sequentially changed from 1 to 0.

FIG. 6 is a diagram showing a waveform of ΔA32 when the value of K is sequentially changed from 1 to 0.

FIG. 7 is a diagram illustrating a configuration of a main part of an apparatus according to a second embodiment.

FIG. 8 is a flowchart for explaining the operation of the device according to the second embodiment.

FIG. 9 is a detailed diagram of step 121 shown in FIG. 8;

FIG. 10 is a diagram showing a frequency spectrum of ΔA32.

FIG. 11 is a diagram showing a frequency spectrum of ΔA32.

FIG. 12 is a diagram corresponding to FIG. 10 and showing a change in a frequency spectrum due to K;

FIG. 13 is a diagram corresponding to FIG. 11 and showing a change in frequency spectrum due to K;

FIG. 14 is a view showing a difference spectrum when K = 0 and K = 0.35.

FIG. 15 is a diagram showing a difference spectrum when K = 0 and K = 0.35.

FIG. 16 is a diagram showing a filter output waveform.

FIG. 17 is a diagram showing a time series of Φ obtained by the present invention and a conventional Φ.

[Explanation of symbols]

 Reference Signs List 3 light irradiation unit 4 light receiving unit 10 digital processing unit 11 Φ calculation unit 12 conversion unit 13 control unit

 ────────────────────────────────────────────────── ─── Continuing on the front page (72) Inventor Masayoshi Fuse 1-31-4 Nishiochiai, Shinjuku-ku, Tokyo Japan F-term (reference) 2G059 AA06 BB13 EE01 EE02 EE11 GG02 GG03 HH01 HH02 HH06 MM01 MM09 MM20 NN10 4C038 KK01 KL07 KM01 KX02

Claims (10)

[Claims] 1. A blood light-absorbing substance concentration for measuring a concentration ratio of a plurality of light-absorbing substances in blood based on a pulsation of tissue transmitted light or reflected light obtained by irradiating a living tissue with light of a plurality of wavelengths. In the measuring apparatus, a dimming degree change detecting means for obtaining a dimming degree change of the tissue transmitted light or the reflected light; and two predetermined wavelengths ΔA2 and ΔA3 of the dimming degree change obtained by the dimming degree change detecting means. Weighted subtraction waveform between waveforms ΔA32 =
A weighted subtraction waveform calculation means for calculating ΔA3 (1-K) -ΔA2K for each K by changing K, and the power of the weighted subtraction waveform determined by the weighted subtraction waveform calculation means for each K is obtained. Or take the maximum
Weight value detection means for calculating the value of K; Φ calculation means for obtaining Φ = K / (1-K) from the value of K obtained by the weight value detection means; and Φ calculated by the Φ calculation means A blood-light-absorbing substance concentration measuring device, comprising: a concentration ratio calculating means for calculating a blood light-absorbing substance concentration ratio. 2. A light-absorbing substance concentration in blood for measuring a concentration ratio of a plurality of light-absorbing substances in blood based on a pulsation of a transmitted light or a reflected light obtained by irradiating a living tissue with light of a plurality of wavelengths. In the measuring apparatus, a dimming degree change detecting means for obtaining a dimming degree change of the tissue transmitted light or the reflected light; and two predetermined wavelengths ΔA2 and ΔA3 of the dimming degree change obtained by the dimming degree change detecting means. Weighted subtraction waveform between waveforms ΔA32 =
A weighted subtraction waveform calculation means for calculating ΔA3 (1-K) -ΔA2K for each K by changing K, and the power of the weighted subtraction waveform determined by the weighted subtraction waveform calculation means for each K is obtained. Or maximal
Weight value detecting means for calculating the value of K; a filter for passing the dimming degree change calculated by the dimming degree change detecting means; and a frequency analysis for a weighted subtraction waveform at the value of K calculated by the weighting value detecting means. And a pass band determining means for determining a pass band of the filter based on the analysis result; a Φ calculating means for determining a ratio Φ of a change in dimming degree of each of two outputs of the filter; A concentration ratio calculating means for calculating a concentration ratio of the blood light-absorbing substance based on the Φ. 3. The pass band determining means performs a frequency analysis of the original waveforms of ΔA2 and ΔA3 obtained by the dimming degree change detecting means, and performs a frequency analysis of a weighted subtraction waveform at K at which the power becomes minimum or maximum. The apparatus for measuring a concentration of a light-absorbing substance in blood according to claim 2, wherein the pass band is determined based on the comparison. 4. In the frequency analysis, a moving average is taken for each fixed number of samples in a fixed section, a difference waveform between two moving average waveforms having different numbers of samples is obtained, and the power of the band is calculated based on the difference waveform. 3. The method according to claim 2, wherein
Or the blood light absorbing substance concentration measuring device according to 3 above. 5. The filter according to claim 2, wherein the filter calculates a moving average for each of a fixed number of samples in a fixed section and obtains a difference waveform between two moving average waveforms having different numbers of samples as they are. Blood light absorbing substance concentration measuring device. 6. The wavelength used is three or more wavelengths, and the pass band determining means determines the pass band of the filter based on two wavelengths among them, and applies ΔA 2, ΔA 3
3. The blood light-absorbing substance concentration measuring apparatus according to claim 2, wherein the amount of change in the dimming degree for wavelengths other than the above is passed. 7. The weighted value detecting means calculates the power of a weighted subtracted waveform near Ka for the next section, assuming that the value of K obtained for a certain section when the apparatus performs continuous measurement is Ka. Find the value of K at the minimum or maximum of,
3. The apparatus for measuring the concentration of a light-absorbing substance in blood according to claim 1, wherein the apparatus is used as a new Ka. 8. The weight value detecting means obtains power for a plurality of weighted subtraction waveforms for each of a plurality of K values, and estimates a K value at which the power becomes minimum or maximum based on the obtained power. The blood light-absorbing substance concentration measuring device according to claim 1 or 2. 9. The method according to claim 1, wherein the concentration ratio of the light-absorbing substance in the blood determined by the concentration ratio calculating means is the oxygen saturation of arterial blood or another concentration ratio of the light-absorbing substance in the blood. Blood light absorbing substance concentration measuring device. 10. The method according to claim 1, wherein the concentration ratio of the blood light-absorbing substance determined by the concentration ratio calculating means is the oxygen saturation of venous blood or another blood light-absorbing substance concentration ratio. The light-absorbing substance concentration measuring device according to the above.