JP4945415B2 - Component concentration measuring apparatus and component concentration measuring apparatus control method - Google Patents

Description

  The present invention relates to a non-invasive component concentration measuring device and a component concentration measuring device control method, and more particularly, to a non-invasive measuring device and measuring method for non-invasive measurement of glucose concentration as a blood component, that is, a blood glucose level.

  To date, various methods based on transcutaneous electromagnetic wave irradiation or radiation observation have been tried as noninvasive component concentration measurement methods. All of these utilize the interaction with an electromagnetic wave of a specific wavelength that a blood molecule of interest has, for example, glucose molecules in the case of blood glucose level, that is, absorption or scattering.

  However, the interaction between glucose and electromagnetic waves is small, and there is a limit to the intensity of electromagnetic waves that can be safely irradiated to a living body. Furthermore, since living bodies are scatterers against electromagnetic waves, , Has not been effective enough.

  Among the conventional techniques using the interaction between glucose and electromagnetic waves, a photoacoustic method for observing sound waves generated in a living body by irradiating the living body with electromagnetic waves has attracted attention.

  The photoacoustic method is that when a living body is irradiated with a certain amount of electromagnetic waves, the electromagnetic waves are absorbed by the molecules contained in the living body, and the portion irradiated with the electromagnetic waves is locally heated to cause thermal expansion and generate sound waves. Since the pressure of this sound wave depends on the amount of molecules that absorb electromagnetic waves, it is a method of measuring the amount of molecules in the living body by measuring the pressure of sound waves. Moreover, among the photoacoustic methods, heat is generated in a local area irradiated with light, and thermal expansion is caused without diffusion of the heat. Call the law.

  A sound wave is a pressure wave propagating in a living body and has a characteristic that it is less likely to scatter than an electromagnetic wave. ). FIG. 22 is a diagram showing a configuration example of a conventional component concentration measuring apparatus using a photoacoustic method as a conventional example. Light pulses are used as electromagnetic waves. In this example, blood sugar, that is, glucose is used as a measurement target as a blood component. The driving power source 604 supplies a pulsed excitation current to the pulse light source 616, and the pulse light source 616 generates a light pulse having a sub-microsecond duration, and the light pulse is irradiated to the living body test part 610. The light pulse generates a sound wave called a pulsed photoacoustic signal in the living body test part 610, the photoacoustic signal is detected by the ultrasonic detector 613, and the photoacoustic signal is converted into an electric signal proportional to the sound pressure. Is done.

  The waveform of the electrical signal is observed by a waveform observer 620. The waveform monitor 620 is triggered by a signal synchronized with the excitation current, and an electrical signal proportional to the sound pressure is displayed at a fixed position on the screen of the waveform monitor 620, and the signals are integrated and averaged for measurement. Can do.

  By analyzing the amplitude of the electrical signal proportional to the sound pressure thus obtained, the blood glucose level in the living body test part 610, that is, the amount of glucose is measured. In the case of the example shown in FIG. 22, light pulses having a sub-microsecond pulse width are repeatedly generated at a maximum of 1 kHz, and 1024 light pulses are averaged to obtain an electric signal proportional to the sound pressure. Is not obtained.

  Therefore, a second conventional example using a light source that is continuously intensity-modulated is disclosed for the purpose of improving accuracy (see, for example, Patent Documents 1 and 2). FIG. 23 shows the configuration of the second conventional apparatus. In this example as well, blood sugar is the main measurement target, and high accuracy is attempted using a plurality of light sources having different wavelengths.

  In order to avoid complicated explanation, the operation when the number of light sources is two will be described with reference to FIG. In FIG. 23, light sources having different wavelengths, that is, a first light source 601 and a second light source 605 are driven by a driving power source 604 and a driving power source 608, respectively, and output continuous light.

  The light output from the first light source 601 and the second light source 605 is intermittently driven by a chopper plate 617 that is driven by a motor 618 and rotates at a constant rotational speed. Here, the chopper plate 617 is formed of an opaque material, and the chopper plate 617 has a relatively small number on the circumference on which the light of the first light source 601 and the second light source 605 passes on the circumference around the axis of the motor 618. An opening is formed.

With the above-described configuration, the light output from each of the first light source 601 and the second light source 605 is intensity-modulated at a relatively simple modulation frequency f ₁ and modulation frequency f ₂ , and then multiplexed by the multiplexer 609. Then, the living body test part 610 is irradiated as one light beam.

The living body test region 610 photoacoustic signal of frequency f ₁ is generated by the light of the first light source 601, the photoacoustic signal having the frequency f ₂ is generated by the light of the second light source 605, these photoacoustic The signal is detected by the acoustic sensor 619 and converted into an electric signal proportional to the sound pressure, and the frequency spectrum is observed by the frequency analyzer 621.

  In this example, the wavelengths of the plurality of light sources are all set to the absorption wavelength of glucose, and the intensity of the photoacoustic signal corresponding to each wavelength is measured as an electrical signal corresponding to the amount of glucose contained in the blood. The

Here, the relationship between the intensity of the measured value of the photoacoustic signal and the value obtained by measuring the glucose content by separately collected blood is stored in advance, and the amount of glucose is measured from the measured value of the photoacoustic signal. ing.
JP-A-10-189 WO 2005/107592 A1 University of Oulu (University of Oulu, Finland) thesis “Pulse photoacoustic techniques and glucodesistion in human blood and tissue” (IBS 951-42-6690-0, 2001/95.

  The light absorption of a solution obtained by mixing a target component with a liquid may vary depending on the temperature of the solution. For this reason, if the temperature of the measurement object changes during the measurement of the component concentration using the photoacoustic method, the blood component concentration cannot be measured accurately.

  When the intensity-modulated light is irradiated, the temperature of the light irradiation point where the intensity-modulated light is irradiated may gradually increase. Since this increase in temperature is inevitable, it is conceivable to monitor the temperature of the light irradiation point. However, it is difficult to monitor the temperature of the light irradiation point with a thermometer without blocking the light. Even if it can be achieved, a thermometer cannot detect a rapid temperature change.

  Further, when the surface of a human body or the like is a soft object to be measured, or when the object to be measured moves, the position of the measurement site that absorbs the intensity-modulated light may shift. For example, when there is a blood vessel that absorbs intensity-modulated light under the epidermis that easily stretches, the position of the blood vessel that is irradiating the intensity-modulated light through the skin changes rapidly depending on the posture of the subject.

  If it deviates from the measurement site where the intensity modulated light is absorbed, the amount of absorption of the intensity modulated light changes extremely. Further, even within the measurement site that absorbs the intensity-modulated light, the temperature rapidly decreases when it deviates from the light irradiation point. For this reason, if it deviates from the measurement site that absorbs the intensity-modulated light, the blood component concentration cannot be measured accurately.

  Accordingly, an object of the present invention is to monitor the shift of the position of a measurement site that absorbs intensity-modulated light without using a thermometer.

  The change in the absorbance of the intensity-modulated light due to the temperature change is remarkable in the liquid contained in the solution. In view of this, a calibration light generating means for detecting a change in light absorption of the liquid is provided. By detecting the change in light absorption of the liquid, the temperature change of the solution can be monitored without using a thermometer.

The component concentration measuring apparatus according to the present invention is:
Two measuring light generating means for generating and outputting light having different wavelengths with the same absorption of the liquid in a solution in which the target component is mixed with the liquid;
Calibration light generating means for generating and outputting light having a wavelength equal to that of one of the measurement light generating means, the absorption exhibited by the target component;
The light from one of the measurement light generation means is intensity-modulated at a predetermined frequency and output, and the light from one of the measurement light generation means, the other of the measurement light generation means, and the Each of the light from the calibration light generating means, the light modulating means for outputting the intensity modulated in the opposite phase with each other at the constant frequency; and
The intensity modulated light of the light from one of the measurement light generating means, the intensity modulated light of the light from one of the measurement light generating means, and the intensity modulated light of the light from the other of the measurement light generating means are synthesized. Synthetic light for measurement, and light for intensity modulation from one of the light generation means for measurement and light for intensity calibration from the light generation means for calibration are emitted toward the solution. Light emitting means for
Normalized sound wave generated from the solution by intensity-modulated light from one of the measurement light generation means, measurement sound wave generated from the solution by the measurement synthetic light emitted from the light emission means, and A sound wave detecting means for detecting a calibration sound wave generated from the solution by the calibration synthetic light emitted by the light emitting means;
A concentration calculating means for calculating the concentration of the target component based on the measurement sound wave detected by the sound wave detecting means;
Equipped with a,
It characterized that you detect the temperature change of the liquid based on the detected normalized sound wave and said calibration wave the acoustic wave detection means.

  A calibration light generating means is provided, and the target object is irradiated with calibration synthetic light to measure sound waves that are not affected by the concentration of the target component. By observing a change in the intensity of the sound wave generated by the calibration composite light, it is possible to monitor whether or not there is a sudden change in the absorbance of the liquid in the solution.

In constituent concentration measuring apparatus according to the present invention, it is determined whether within the ratio of the detected intensity with the intensity of the sound wave for normalization of the calibration sound wave of the acoustic wave detection means is a predetermined It is preferable to further include a calibration sound wave determination unit that further includes a calibration sound wave determination unit that determines whether or not the temperature change of the liquid in the solution is within an allowable range .
By further providing the calibration sound wave determination means, it can be confirmed that the temperature change of the solution is within an allowable range. Thereby, the measurement accuracy of the component concentration can be maintained.

The component concentration measuring apparatus according to the present invention further comprises output means for outputting a change in temperature when the calibration sound wave determination means is determined to be outside the range, and the output means has a temperature change. When the message is output, it is preferable that the concentration calculation unit stops calculating the concentration of the target component .
By further providing an out-of-threshold determination means, it can be determined that there has been a sudden temperature change in the solution. Since the output means notifies the subject that the temperature has changed in the solution, the component concentration can be accurately measured by re-measuring the component concentration.

In the component concentration measurement apparatus according to the present invention, the temperature at which the temperature change of the liquid is calculated based on the amount of change in the ratio between the intensity of the calibration sound wave detected by the sound wave detection means and the intensity of the normalization sound wave. It is preferable to further include a calculation unit.
By further providing the temperature calculating means, the temperature of the liquid can be calculated even if the temperature changes suddenly. By calculating the temperature of the liquid, the contribution of the liquid absorption coefficient due to the temperature can be corrected.

In the component concentration measuring apparatus according to the present invention, the concentration of the target component calculated by the concentration calculating unit is calculated based on the temperature change calculated by the temperature calculating unit and the change rate of the absorbance of the solution and the target component with respect to the temperature. It is preferable to further include density correction means for correcting.
By further providing the density correction means, the density correction means corrects the density of the target component with the calculated temperature calculated by the temperature calculation means. Thereby, the contribution of the temperature to the concentration of the target component can be automatically corrected.

In the component concentration measuring apparatus according to the present invention,
The liquid is water;
The target component is glucose or cholesterol,
It is preferable that the solution is blood.
According to the present invention, it is possible to non-invasively measure the concentration of glucose or cholesterol in blood that is less affected by the temperature of the living body test part.

The component concentration measuring device control method according to the present invention is:
Two measuring light generating means for generating and outputting light having different wavelengths with the same absorption of the liquid in a solution in which the target component is mixed with the liquid, a calibration light generating means, a light modulating means, A control method for a component concentration measuring apparatus comprising a control circuit for controlling light emitting means, sound wave detecting means, and concentration calculating means,
The control circuit causes one of the measurement light generation means to output light, and causes the light modulation means to output the light from one of the measurement light generation means after intensity modulation at a predetermined constant frequency. The light emitting means emits intensity-modulated light from one of the measurement light generation means toward the solution, and the sound wave detection means emits light from one of the measurement light generation means. A normalization sound wave detection procedure for detecting a normalization sound wave generated from the solution by intensity-modulated light;
Before, after or simultaneously with the normalization sound wave detection procedure, the control circuit causes the two measurement light generation means to output light, and causes the light modulation means to emit light from one of the measurement light generation means. And the light from the other of the measurement light generating means are modulated in intensity with the constant frequency in opposite phases to each other and output to the light emitting means, and the intensity modulated light of the light from one of the measurement light generating means and The measurement combined light combined with the intensity-modulated light from the other of the measurement light generation means is emitted toward the solution, and the sound wave detection means is caused by the measurement combined light emitted from the light emission means. A measurement sound wave detection procedure for detecting a measurement sound wave generated from the solution;
Before, after or simultaneously with the measurement sound wave detection procedure, the control circuit generates light having the same wavelength as the one of the measurement light generation means in the calibration light generation means. And output the light from one of the measurement light generating means by intensity modulation at a predetermined constant frequency and output the light from one of the measurement light generating means and the calibration light. Light from the light generating means is intensity-modulated in phase opposite to each other at the constant frequency and output, and the light emitting means directs intensity-modulated light from one of the measurement light generating means to the solution. And radiating the combined light for calibration obtained by synthesizing the intensity-modulated light of the light from one of the measurement light generation means and the intensity-modulated light of the light from the calibration light generation means toward the solution, The measuring light generating means is connected to the sound wave detecting means. Sonic for normalization generated from the solution by the intensity-modulated light of the light from one of, and, by calibration synthetic light emission of the light emitting means to detect a calibration waves generated from the solution, the acoustic wave detection means A temperature change detection procedure for detecting a temperature change of the liquid based on the normalized sound wave and the calibration sound wave detected by
After the measurement sound wave detection procedure, the control circuit causes the concentration calculation unit to calculate the concentration of the target component based on the measurement sound wave detected by the sound wave detection unit;
It is characterized by having.

  A calibration light generating means is provided, and the target object is irradiated with calibration synthetic light to measure sound waves that are not affected by the concentration of the target component. By observing a change in the intensity of the sound wave generated by the calibration composite light, it is possible to monitor whether or not there is a sudden change in the absorbance of the liquid in the solution.

In the component concentration measuring device control method according to the present invention,
The component concentration measuring device further includes a sonic wave determination means for calibration,
The control circuit further determines whether or not the calibration sound wave determination means has a ratio between the intensity of the calibration sound wave detected by the sound wave detection means and the intensity of the normalization sound wave within a predetermined range. It is preferable that a calibration sound wave determination procedure for determining whether or not the temperature change of the liquid in the solution is within an allowable range is included after the temperature change detection procedure.
By further providing the calibration sound wave determination means, it can be confirmed that the temperature change of the solution is within an allowable range. Thereby, the measurement accuracy of the component concentration can be maintained.

In the component concentration measuring device control method according to the present invention,
The component concentration measuring device further comprises output means,
When it is determined that the calibration sound wave determination means is out of the range, the control circuit further outputs an output procedure to the effect that the output means has a temperature change after the calibration sound wave determination procedure. Yes, and
In the concentration calculation procedure, it is preferable that when the output unit outputs that the temperature has changed, the concentration calculation unit stops calculating the concentration of the target component .
By further providing an out-of-threshold determination means, it can be determined that there has been a sudden temperature change in the solution. Since the output means notifies the subject that the temperature has changed in the solution, the component concentration can be accurately measured by re-measuring the component concentration.

In the component concentration measuring device control method according to the present invention,
The component concentration measuring device further includes a temperature calculating means,
The control circuit further causes the temperature calculation means to change the temperature of the liquid based on the amount of change in the ratio between the intensity of the calibration sound wave detected by the sound wave detection means and the intensity of the normalization sound wave. It is preferable to have a temperature calculation procedure for calculation after the temperature change detection procedure.
By further providing the temperature calculating means, the temperature of the liquid can be calculated even if the temperature changes suddenly. By calculating the temperature of the liquid, the contribution of the liquid absorption coefficient due to the temperature can be corrected.

In the component concentration measuring device control method according to the present invention,
The component concentration measurement apparatus further includes a concentration correction unit,
The control circuit further causes the concentration correction unit to calculate the target component calculated by the concentration calculation unit based on a temperature change calculated by the temperature calculation unit and a change rate of the absorbance of the solution and the target component with respect to temperature. It is preferable to have a density correction procedure for correcting the density after the density calculation procedure and the temperature change detection procedure .
By further providing the density correction means, the density correction means corrects the density of the target component with the calculated temperature calculated by the temperature calculation means. Thereby, the contribution of the temperature to the concentration of the target component can be automatically corrected.

In the component concentration measuring device control method according to the present invention,
The liquid is water;
The target component is glucose or cholesterol,
It is preferable that the solution is blood.
According to the present invention, it is possible to non-invasively measure the concentration of glucose or cholesterol in blood that is less affected by the temperature of the living body test part.

  According to the present invention, it is possible to monitor a temperature change of a solution without using a thermometer by detecting a change in light absorption of a liquid.

  Embodiments of the present invention will be described with reference to the accompanying drawings. The embodiment described below is an example of the configuration of the present invention, and the present invention is not limited to the following embodiment.

  In the following embodiments, a component concentration measuring device and a component concentration measuring device control method will be described as a blood component concentration measuring device or a blood component concentration measuring device control method. If the living body test part described in the present embodiment is replaced with a measurement object, blood is replaced with a solution, water is replaced with a liquid, and glucose or cholesterol is replaced with a target component, a liquid component concentration measuring device or a liquid component concentration measuring device control is provided. It can be implemented as a method. For example, the solution includes not only blood but also a solution constituting a living body such as lymph and tears. The target component is not limited to glucose or cholesterol, but also includes components in the solution constituting the living body, such as “lymph component” and “tear component”. Thus, various components can be measured according to the measurement object.

  Moreover, in the structure of the following embodiment and an Example, if it replaces with a biological test part and a fruit is put, it will function as a fruit sugar content meter. This is because sucrose and fructose, which are the sweetness components of fruits, have absorption at a wavelength similar to that of glucose, which is a blood sugar component. Thus, in the range not departing from the spirit of the present embodiment, It goes without saying that the measuring apparatus and the measuring apparatus control method can be applied to various objects.

(First embodiment)
The blood component concentration measurement apparatus according to the present embodiment includes a light generating unit that generates two light beams having different wavelengths, and an electrical intensity of each of the two light beams having different wavelengths by signals having the same frequency and opposite phases. Light modulating means for modulating, light emitting means for combining two light beams of different wavelengths, which have been intensity-modulated, into one light beam and emitting the light toward a living body; and a sound wave generated in the living body by the emitted light. A blood component concentration measuring device comprising: a sound wave detecting means for detecting; and a blood component concentration calculating means for calculating a blood component concentration in the living body from the pressure of the detected sound wave. The blood component concentration calculating means according to the present embodiment can be applied not only to the present embodiment but also to embodiments described later.

  Furthermore, in the blood component concentration measuring apparatus according to the present embodiment, the light generating means sets the wavelength of one wave of light to a wavelength at which the blood component exhibits characteristic absorption, and the wavelength of the other one wave of light. Can also be set to a wavelength at which water exhibits absorption equal to that of the one-wave light wavelength.

  With reference to FIG. 1, the configuration according to the present embodiment will be described. FIG. 1 shows a basic configuration of a blood component concentration measuring apparatus according to this embodiment. In FIG. 1, a first light source 101 as a part of light generation means is intensity-modulated by a drive circuit 104 as a part of light modulation means in synchronization with an oscillator 103 as a part of light modulation means. Yes.

  On the other hand, the second light source 105 as a part of the light generating means is intensity-modulated in synchronism with the oscillator 103 by a drive circuit 108 as a part of the light modulating means. However, the output of the oscillator 103 is fed to the drive circuit 108 via the 180 ° phase shift circuit 107 as a part of the light modulation means. As a result, the second light source 105 is supplied to the first light source 101. On the other hand, the intensity is modulated by a signal whose phase is changed by 180 °.

  Here, the wavelength of each of the first light source 101 and the second light source 105 shown in FIG. 1 is set such that the wavelength of one wave of light has a characteristic absorption of blood components, and the other one wave The wavelength of light is set to a wavelength at which water exhibits absorption equivalent to that at the wavelength of the one wave of light.

  The first light source 101 and the second light source 105 each output light having different wavelengths, and each output light is combined by a combiner 109 serving as a light emitting unit, and as a light beam as a subject. The living body test part 110 is irradiated. The sound wave generated in the living body test part 110 by the light output from each of the irradiated first light source 101 and the second light source 105, that is, the photoacoustic signal, is detected by ultrasonic waves as a part of the sound wave detecting means. And is converted into an electric signal proportional to the sound pressure of the photoacoustic signal. The electrical signal is synchronously detected by a phase detection amplifier 114 as a part of sound wave detection means synchronized with the oscillator 103, and an electrical signal proportional to the sound pressure is output to the output terminal 115.

  Here, the intensity of the signal output to the output terminal 115 is proportional to the amount of light output from each of the first light source 101 and the second light source 105 absorbed by the components in the living body test portion 110. The intensity of the signal is proportional to the amount of components in the living body test part 110. Therefore, a blood component concentration calculating means (not shown) calculates the amount of the component to be measured in the blood in the living body test part 110 from the measured value of the intensity of the signal output to the output terminal 115.

  Since the blood component concentration measuring apparatus according to the present embodiment modulates the intensity of two light beams of different wavelengths output from the first light source 101 and the second light source 105 with signals having the same period, that is, the same frequency, There is a characteristic that the frequency characteristic of the sound wave detector 113 is not affected by the nonuniformity, and this point is superior to the existing technology.

  As described above, the blood component concentration measuring apparatus according to the present embodiment can measure blood components with high accuracy.

  The blood component concentration measurement apparatus control method according to the present embodiment includes a light generation procedure in which the light generation unit generates two light beams having different wavelengths, and a light modulation unit that has different wavelengths generated in the light generation procedure. An optical modulation procedure for electrically modulating the intensity of each of the two light waves with a signal having the same frequency and an opposite phase, and the light emitting means includes two light beams of different wavelengths that are intensity-modulated in the optical modulation procedure. A light emitting procedure for combining the luminous flux and emitting the light toward the living body; a sound wave detecting means for detecting a sound wave generated in the living body by the light emitted in the light emitting procedure; and A blood component concentration measuring device control method that sequentially includes a blood component concentration calculating procedure for calculating a blood component concentration in a living body from pressure.

  Furthermore, in the blood component concentration measuring device control method according to the present embodiment, the light generation procedure sets the wavelength of one wave of light to a wavelength at which the blood component exhibits characteristic absorption, and another one wave of light. Is set to a wavelength at which water exhibits absorption equal to that at the wavelength of the one wave of light, and a blood component concentration measuring device control method for generating two waves of different wavelengths can be obtained.

  Here, as a method for electrically intensity-modulating each of the two light beams having different wavelengths, two light beams having different wavelengths are generated, and then each of the two light beams having different wavelengths is phased at the same frequency. Alternatively, a method may be used in which the intensity is electrically modulated using a modulator using signals that differ by 180 °, and the drive circuit 104 and the drive circuit 108 are the first light source 101 and the second light source, respectively, as in the example shown in FIG. A direct modulation method in which the intensity is modulated at the same time as causing the light to emit light 105 may be used.

  Next, each of the two light beams having different wavelengths whose intensity is modulated by the above procedure is combined into one light beam by, for example, a combiner 109 shown in FIG. A sound wave generated in the living body by each of the two light beams having different wavelengths, that is, a photoacoustic signal, is detected by, for example, an ultrasonic detector 113 shown in FIG. The phase detection amplifier 114 shown in FIG. 1 performs synchronous detection, and an electric signal proportional to the photoacoustic signal is output to the output terminal 115. Next, in the blood component concentration calculation procedure, the blood component concentration in the living body is calculated from the pressure of the sound wave detected in the sound wave detection procedure.

  In the blood component concentration measuring apparatus control method according to the present embodiment, the intensity of two light beams having different wavelengths output from the first light source 101 and the second light source 105 is modulated with signals having the same period, that is, the same frequency. Therefore, there is a feature that it is not affected by non-uniformity in the frequency characteristics of the measurement system that detects sound waves, which is superior to existing technologies.

  As described above, the blood component concentration measuring apparatus control method according to the present embodiment can measure blood components with high accuracy.

  In the blood component concentration measuring apparatus according to the present embodiment, the light modulating unit may be a unit that modulates at the same frequency as a resonance frequency related to detection of a sound wave generated in a living body. By modulating each of the two light beams having different wavelengths at the same frequency as the resonance frequency related to the detection of the sound wave generated in the living body, the sound wave generated in the living body can be detected with high sensitivity.

  In the blood component concentration measuring apparatus control method according to the present embodiment, the light modulation procedure may be a procedure of modulating at the same frequency as the resonance frequency related to detection of sound waves generated in the living body. By modulating each of the two light beams having different wavelengths at the same frequency as the resonance frequency related to the detection of the sound wave generated in the living body, the sound wave generated in the living body can be detected with high sensitivity.

  In the blood component concentration measuring apparatus according to the present embodiment, the blood component concentration calculating means determines the pressure of the sound wave generated by irradiating the living body with the two waves of light having the different wavelengths, one of the two waves of light. It is also possible to divide by the pressure of the sound wave generated when the light of zero is zero. By such division, the blood component concentration can be measured with high accuracy.

  In the blood component concentration measurement device control method according to the present embodiment, the blood sample concentration calculation procedure includes the step of calculating the pressure of a sound wave generated by irradiating a living body with the two waves of light, and one wave of the two waves of light. It is also possible to divide by the pressure of the sound wave generated when the light of zero is zero. By such division, the blood component concentration can be measured with high accuracy.

  In the blood component concentration measuring apparatus according to the present embodiment, the light generating means combines the intensity-modulated two light beams having different wavelengths into one light beam and irradiates water with zero pressure of the sound wave generated. It is also possible to use a means for adjusting the relative intensity of each of the two light beams having different wavelengths.

  In the blood component concentration measurement apparatus according to the present embodiment, for example, in FIG. 1, the first light source 101 and the first light source 101 and The first light source 101 and the second light source 105 are irradiated such that the light output from the second light source 105 is combined with one light beam so that the photoacoustic signal detected by the ultrasonic detector 113 becomes zero. This is a case in which the relative intensity of the light output from is adjusted.

  When the light intensities of the first light source 101 and the second light source 105 are adjusted as described above, the relative intensities of the light from the first light source 101 and the second light source 105 can be easily adjusted equally. Therefore, the blood component concentration can be easily measured with high accuracy.

  In the blood component concentration measuring apparatus control method according to the present embodiment, the two light beams having different wavelengths, which have been modulated in intensity, are combined into one light beam between the light modulation procedure and the light emission procedure to form water. The blood component concentration measuring device control method may further include an intensity adjustment procedure for adjusting the relative intensity of each of the two waves of light so that the pressure of the sound wave emitted by being emitted toward zero becomes zero. .

  The blood component concentration measuring apparatus control method according to the present embodiment, for example, after the procedure of combining two light beams of different wavelengths whose intensity is modulated into one light beam, the two light beams of different wavelengths are combined into one light beam. The first light source 101 and the second light source described above are adjusted by adjusting the relative intensities of the two light beams so that the pressure of the sound wave generated by being combined with the light flux and emitted toward the water becomes zero. Since the relative intensity of the light output from the light source 105 can be easily adjusted equally, blood components can be easily measured with high accuracy.

  In the blood component concentration measuring apparatus according to the present embodiment, the sound wave detecting means may be a means for detecting by synchronous detection in synchronization with the modulation frequency. In the blood component concentration measurement apparatus according to the present embodiment, for example, a photoacoustic signal corresponding to each of the light output from the first light source 101 and the second light source 105 is detected by the ultrasonic detector 113 and converted into an electrical signal. This is an example in which the detected signal is detected by synchronous detection in synchronization with a signal for intensity-modulating each of the light output from the first light source 101 and the second light source 105 in the phase detection amplifier 114. In the phase detection amplifier 114, the detection accuracy of the photoacoustic signal corresponding to each of the lights output from the first light source 101 and the second light source 105 is improved, and the photoacoustic signal can be measured with higher accuracy.

  In the blood component concentration measuring apparatus control method according to the present embodiment, the sound wave detection procedure may be a procedure of detecting by synchronous detection in synchronization with the modulation frequency. In the blood component concentration measuring device control method according to the present embodiment, for example, the photoacoustic signal corresponding to each of the two light beams having different wavelengths is converted into a signal for intensity-modulating each of the two light beams having different wavelengths. This is a case of detecting by synchronous detection in synchronization. The detection accuracy of the photoacoustic signal corresponding to each of the light output from the first light source 101 and the second light source 105 is improved, and the photoacoustic signal can be measured with higher accuracy.

  In the blood component concentration measuring apparatus according to this embodiment, the light generating means and the light modulating means may be means for directly modulating each of the two semiconductor laser light sources with rectangular wave signals having the same frequency and opposite phases. it can. By adopting an apparatus configuration in which each of the two semiconductor laser light sources is directly modulated by rectangular wave signals having the same frequency and opposite phases, the apparatus configuration can be simplified.

  In the blood component concentration measurement device control method according to the present embodiment, the light generation procedure and the light modulation procedure are procedures for directly modulating each of the two semiconductor laser light sources with rectangular wave signals having the same frequency and opposite phases. be able to. The procedure can be simplified by adopting a procedure in which each of the two semiconductor laser light sources is directly modulated by rectangular wave signals having the same frequency and opposite phases.

  Next, details of the technology that is the basis of the blood component concentration measuring apparatus and the blood component concentration measuring apparatus control method according to the present embodiment will be described.

  With reference to FIG. 1, the structure of the blood component concentration measuring apparatus according to this embodiment will be described. The blood component concentration measuring apparatus according to the present embodiment shown in FIG. 1 includes a first light source 101, a second light source 105, a drive circuit 104, a drive circuit 108, a 180 ° phase shift circuit 107, a multiplexer 109, and an ultrasonic wave. It comprises a detector 113, a phase detection amplifier 114, an output terminal 115, and an oscillator 103.

  The oscillator 103 is connected to the drive circuit 104, the 180 ° phase shift circuit 107, and the phase detection amplifier 114 by signal lines, and transmits signals to the drive circuit 104, the 180 ° phase shift circuit 107, and the phase detection amplifier 114, respectively.

  The driving circuit 104 receives a signal transmitted from the oscillator 103, supplies driving power to the first light source 101 connected by the signal line, and causes the first light source 101 to emit light.

  The 180 ° phase shift circuit 107 receives the signal transmitted from the oscillator 103, and transmits a signal obtained by giving a phase change of 180 ° to the received signal to the drive circuit 108 connected by the signal line.

  The drive circuit 108 receives the signal transmitted from the 180 ° phase shift circuit 107, supplies drive power to the second light source 105 connected by the signal line, and causes the second light source 105 to emit light.

  Each of the first light source 101 and the second light source 105 outputs light having different wavelengths, and guides the light output from the light source to the multiplexer 109 by the light wave transmission means.

  The light output from the first light source 101 and the light output from the second light source 105 are input to the combiner 109, combined, and applied to a predetermined position of the living body test unit 110 as one light beam. A sound wave, that is, a photoacoustic signal is generated in the living body test part 110.

  The ultrasonic detector 113 detects the photoacoustic signal of the living body test part 110, converts it into an electrical signal, and transmits it to the phase detection amplifier 114 connected by a signal line.

  The phase detection amplifier 114 receives a synchronization signal necessary for synchronous detection transmitted from the oscillator 103 and also receives an electrical signal proportional to the photoacoustic signal transmitted from the ultrasonic detector 113 to perform synchronous detection and amplification. Then, filtering is performed and an electric signal proportional to the photoacoustic signal is output to the output terminal 115.

  The first light source 101 outputs light whose intensity is modulated in synchronization with the oscillation frequency of the oscillator 103. On the other hand, the second light source 105 outputs light whose intensity is modulated in synchronization with a signal having the oscillation frequency of the oscillator 103 and a phase change of 180 ° by the 180 ° phase shift circuit 107.

  As described above, in the blood component concentration measurement apparatus according to the present embodiment, the light output from the first light source 101 and the light output from the second light source 105 are intensity-modulated by signals of the same frequency. In the prior art, there is no influence of non-uniformity in the frequency characteristics of the measurement system, which is a problem when intensity modulation is performed with signals of a plurality of frequencies.

  On the other hand, the nonlinear absorption coefficient dependency existing in the measurement value of the photoacoustic signal, which is a problem in the prior art, uses light of a plurality of wavelengths that give the same absorption coefficient in the blood component concentration measurement apparatus according to this embodiment. What can be solved by measuring in this way will be described below.

数式（１）を解いて未知の血液成分濃度Ｍを求める。ここで、Ｃは、変化し制御或は予想困難な係数、即ち、音響的な結合状態、超音波検出器の感度、前記照射部と前記検出部の間の距離（以下ｒと定義する）、比熱、熱膨張係数、音速、変調周波数、更に、吸収係数にも依存する未知乗数である。 Background light absorption coefficients α ₁ ^(b) , α ₂ ^(b) and molar absorption α ₁ ⁽⁰⁾ , α ₂ ^{(0) of} the blood component to be measured for light of wavelength λ ₁ and wavelength λ ₂ , respectively. Is known, the simultaneous equations including the measured values s ₁ and s ₂ of the photoacoustic signals at the respective wavelengths are expressed as the following formula (1).

Equation (1) is solved to determine an unknown blood component concentration M. Here, C is a coefficient that changes and is difficult to control or predict, that is, acoustic coupling state, sensitivity of the ultrasonic detector, distance between the irradiation unit and the detection unit (hereinafter, defined as r), It is an unknown multiplier that also depends on specific heat, thermal expansion coefficient, sound speed, modulation frequency, and absorption coefficient.

If there is a difference between C in the first row and the second row of Equation (1), it can be other than the amount related to the irradiation light, that is, the difference due to the absorption coefficient. Here, if the combination of the wavelength λ ₁ and the wavelength λ ₂ is selected in parentheses in each row of the formula (1), that is, the absorption coefficients are equal to each other, the absorption coefficients become the same, and the first and second lines Are equal. However, strictly doing this is inconvenient because the combination of wavelength λ ₁ and wavelength λ ₂ will depend on the unknown blood component concentration M.

数式（２）の後段の変形にはｓ_１≒ｓ_２という性質を用いている。 Here, as for the ratio of the absorption coefficient (in each parenthesis) of the formula (1), the background (α _i ^(b) , i = 1, 2) has a term (Mα _i ^{(0 )} ) Is significantly larger than. Thus, it is sufficient to make the absorption coefficients of the background, α _i ^(b) equal, instead of making the absorption coefficients of each row exactly equal. That is, the two light beams having different wavelengths λ ₁ and λ ₂ may be selected so that the background absorption coefficients α ₁ ^(b) and α ₂ ^(b) are equal to each other. Thus, if C in the first and second lines can be equal, it is erased as an unknown constant, and the blood component concentration M to be measured is expressed by Equation (2).

The property of s ₁ ≈s ₂ is used for the subsequent deformation of Equation (2).

Here, looking at Equation (2), the difference in the absorption coefficient of the blood component to be measured at the wavelengths λ ₁ and λ ₂ appears in the denominator. When the difference is larger, the difference signal s ₁ -s _{2 of the} photoacoustic signal is larger, and the measurement becomes easier. In order to maximize this difference, the wavelength at which the absorption coefficient α ₁ ⁽⁰⁾ of the component to be measured is maximized is selected as the wavelength λ ₁ and α ₂ ⁽⁰⁾ = 0, that is, the component to be measured is It is preferable to select the wavelength λ ₂ as a wavelength that does not show the absorption characteristics. Here, from the previous condition, this second wavelength λ ₂ must be α ₂ ^(b) = α ₁ ^(b) , ie the background absorption coefficient must be equal to the absorption coefficient of the first wavelength λ _1. .

Furthermore, in Equation (2), the photoacoustic signal s ₁ appears only in the form of a difference s ₁ -s _{2 from} the photoacoustic signal s ₂ . Now, taking as an example the glucose as a component to be measured, as described above, the two intensities of the photoacoustic signal s ₁ and the photoacoustic signal s _2, only less than 0.1% difference.

However, it is sufficient that the photoacoustic signal s ₂ in the denominator of the formula (2) has an accuracy of about 5%. Therefore, rather than sequentially measuring the _two photoacoustic signals s ₁ and the photoacoustic signal s ₂ , the difference s ₁ -s ₂ between them is measured, and this measured value is determined by the individually measured photoacoustic signal s ₂ . It is much easier to maintain accuracy. Therefore, in the blood component concentration measuring apparatus according to the present embodiment, the light of the _two wavelengths λ ₁ and λ ₂ is irradiated with intensity modulated in opposite phases to each other, so that the photoacoustic signal s ₁ and A difference signal s ₁ -s ₂ between the photoacoustic signals generated by superimposing the photoacoustic signals s ₂ is measured.

  As described above, when measuring blood component concentrations, using two waves of different specific wavelengths, the photoacoustic signals generated by the two waves of different specific wavelengths in the living body are individually provided. Rather than measuring, the difference signal of the photoacoustic signal is measured, and further, the predetermined one photoacoustic signal is set to zero, the other photoacoustic signal is measured, and these are calculated by Equation (2), It can be seen that the blood component concentration can be easily measured.

  Next, sound pressure generated by light irradiation will be described with reference to FIG. FIG. 2 is an explanatory diagram of the basic direct photoacoustic method according to the present embodiment, and FIG. 2 shows the arrangement of observation points in the direct photoacoustic method together with a sound source distribution model. In FIG. 2, the irradiation light 201 enters the living body perpendicularly, and as a result, as described above, the sound source 202 is generated near the surface of the portion irradiated with the light.

  For a sound wave that travels out of the sound source 202 and propagates through the living body (for simplicity, the sound wave is uniform), the sound pressure p is at an observation point 203 that is on the extension line of the irradiation light 201 and is separated from the sound source by a distance r. Observe (r).

  For light with a wavelength of 1 μm or longer used in the blood component concentration measuring apparatus according to the present embodiment, the living body receives strong absorption by the background (water), so the sound source 202 is the surface of the portion irradiated with light. As a result, the generated sound wave can be regarded as a spherical wave.

  The wave equation describing the sound wave propagation shown in FIG. 2 is obtained from an equation of fluid dynamics. That is, the continuous equation and the Navier Stokes equation are assumed to be linear when the density change, the pressure change, and the flow velocity change are very small, and the state equations describing the relationship between the pressure and density in the fluid (water) are combined. It is calculated by solving. Here, the state equation includes temperature as a parameter, and a temperature change when the heat source Q is present is taken in via the state equation.

ここで、ｃは音速、βは熱膨張係数、Ｃ_ｐは定圧比熱である。 When neglecting heat conduction, the small pressure change p is described by the following inhomogeneous Helmholtz equation.

Here, c is the speed of sound, β is the thermal expansion coefficient, and C _p is the constant pressure specific heat.

  In the case of the blood component concentration measuring apparatus according to the present embodiment, light whose intensity is modulated at a constant period T is irradiated and a change in sound pressure synchronized with the constant period T is detected, so that the modulation frequency is f = 1 / T, When the modulation angular frequency is set to ω = 2πf, it is only necessary to pay attention to only the amount having time dependency exp (−iωt) for all the amounts. As a result, the time derivative is a product of -iω.

ここで、音波の波数ｋ＝ω／ｃ＝２πλ（λは音波の波長）を導入した。 Further, since the heat source Q is caused by non-emission relaxation following absorption of the irradiation light 201, the heat source Q is proportional to the absorption coefficient α, and the distribution thereof is irradiation light 201 in the medium (including scattered light if it is generated). Equal to the spatial distribution of. That is, if the light intensity at each point is written as I, Q = αI. As described above, the basic equation relating to the steady direct photoacoustic method is expressed as the following equation (4).

Here, the wave number k = ω / c = 2πλ (λ is the wavelength of the sound wave) was introduced.

The solution of the mathematical formula (4) under the boundary condition of p (r → ∞) → 0 is expressed as the following mathematical formula (5) in a sufficiently far distance (r) α ⁻¹ ).

Now, the observed sound pressure is calculated for some light distribution by Equation (5). First, as the light distribution model A204, a hemispherical distribution in which the intensity is attenuated by e− ^{αr ′} with respect to the radius r ′. This corresponds to the case where the scattering is extremely large and the irradiation light 201 is scattered in all directions as soon as it is incident.

On the other hand, the case where the scattering is zero is the model B205 and the model C206 in FIG. 2, which corresponds to the case where a Gaussian beam and a uniform circular beam having a radius w ₀ are respectively incident. The light intensity distribution of each of these models is shown in FIG.

ここで、Ｐ_０は照射光２０１の全パワーであり、またＦ（ξ）は、

と計算される。音源の分布の情報は、この形状関数Ｆ（ｋα^―１）に集約される。前記形状関数のグラフを、図３に示した。 Now, in addition to the condition r >> α ⁻¹ already used, r >> w ₀ and N≡w ₀ ² / (rλ) << 1 (N in the model A204 is changed to α ⁻¹ instead of w _0. The result of calculation according to Equation (5) is summarized as follows.

Here, P ₀ is the total power of the irradiation light 201, and F (ξ) is

Is calculated. Information on the distribution of the sound sources is collected in this shape function F (kα ⁻¹ ). A graph of the shape function is shown in FIG.

According to the above results, when ξ = kα− ¹ is small, that is, when the wavelength of the sound wave is very long compared to the absorption length (λ) α ⁻¹ ), the photoacoustic signal indicates the information on the absorption coefficient. Does not include anything. The reason is that when ξ << 1, F (ξ) ≈ξ and αF (ξ) ≈k. Therefore, it is understood that the blood component concentration cannot be measured by the photoacoustic method when the wavelength of the sound wave is very long compared to the absorption length, that is, when the modulation frequency is too low.

  Therefore, in the direct photoacoustic method performed on the living body, the modulation frequency should be set to ξ≈1, that is, f≈αc / (2π) or more, and the modulation is performed when the wavelength of the irradiation light is near 1.6 μm. When the frequency f is 150 kHz or more, or when the wavelength of irradiation light is near 2.1 μm, the modulation frequency f needs to be 0.6 MHz or more.

Next, since there is no difference in the results of model B205 and model C206, it can be seen that the light intensity distribution in the direction perpendicular to the optical axis does not affect the signal. However, this simplification is allowed only when N = w ₀ ² / (rλ) << 1 holds. This N is an amount called Fresnel number. When looking at the sound source from the observation point, N represents the width of the phase change caused by the sound wave from each point of the sound source due to the spread of the sound source in the direction perpendicular to the line of sight. Yes. If the Fresnel number N is sufficiently smaller than 1, it is equivalent to that the sound source does not spread in the direction perpendicular to the line of sight.

As a result, the beam diameter w ₀ of the irradiation light 201 has a very advantageous property that it does not affect the photoacoustic signal. There are two reasons for this.

The first is suppression of the influence of scattering in the living body. The model A204 assumes an extreme case where scattering is large, but scattering in a living body is actually not so great. In general, the scattering phenomenon is characterized by a scattering coefficient μ _s and anisotropy g. Here, the latter is the average value <cos θ> of the cosine of the scattering angle θ, and a value of approximately 0.9 has been reported as a value in a living body, particularly skin (for example, Applied Optics, 32, 1993, Pp. 435-447). That is, the scattering in an actual living body is mainly small-angle scattering <θ> ≈26 °.

Now, the rate at which light is reduced by scattering from an incident light beam during propagation of unit length is given by the reduced scattering coefficient μ ′ _{s =} μ _s (1−g), which is longer than the wavelength of light 1 μm. On the other hand, it is actually measured as approximately 1 mm ⁻¹ . (Refer nonpatent literature 3). This value is the value of the absorption coefficient α which is the rate at which light is reduced by absorption from the incident light beam during propagation of the unit length (0.6 mm ⁻¹ and around 2.1 μm at a light wavelength of around 1.6 μm). And 2.4 mm ⁻¹ ).

That is, now, in the living body, the irradiation light 201 is only subjected to scattering at most twice during the absorption length α ⁻¹ , and the scattering angle is small. As a result, the light distribution inside the living body (the sum of incident light flux and scattered light) gradually increases in beam diameter with depth, and looks like a pin head. An actual measurement example of such a light distribution has also been reported (see Applied Optics, 40, 2001, pages 5770-5777). At this time, the total amount of light distribution in the plane of depth z is still expected to attenuate according to exp (−αz). This is because a small number of scattering occurs at a small scattering angle.

  Therefore, when the photoacoustic signal does not depend on the beam diameter of the irradiation light 201, the beam diameter itself of the light distribution at each depth does not matter, and only the total amount in each depth plane is the shape function F (ξ ) May be affected. If this is exp (−αz), as a result, it is expected that there is no influence of scattering on the shape function, as it does not differ in the case of model B205 and model C206 without scattering.

It is the essence of the method in the present embodiment that the shape function is equivalent in the light irradiation of the _two wavelengths λ ₁ and λ ₂ . Therefore, it is highly undesirable that there is a difference in scattering at the _two wavelengths λ ₁ and λ ₂ . Actually, there is no actual measurement report on the wavelength dependence of scattering in the skin for light wavelengths of 1.3 μm or longer, but a constant reduced scattering coefficient μ ′ _s has been reported for blood (Journal). of Biomedical Optics, Vol. 4, 1999, pp. 36-46).

  Therefore, for example, even if there is a slight scattering effect on the shape function, its wavelength dependence is small, and there is a possibility that it will not cause any real harm. Further, as shown here, if the Fresnel number is set small, the influence of scattering on the shape function itself can be suppressed. Therefore, regardless of the wavelength dependence of scattering, the equivalence of the shape function is justified, and it can be seen that the method in this embodiment has high reliability.

The second is that the modulation frequency can be optimized. The light irradiation on the human body has an allowable limit of light intensity depending on the irradiation site, wavelength, irradiation time, and the like. If the beam diameter w ₀ is increased in a range where the Fresnel number N is small, the total power P ₀ of the irradiation light can be increased and the photoacoustic signal can be increased without exceeding the light intensity limit.

Here, if the limit of irradiation intensity is written as I _max , P ₀ = πw ₀ ² I _max , and the Fresnel number N depends on the total power P ₀ , N = f / (πcr) (P ₀ / I _max ) It is expressed. Considering that the distance r is an amount determined by the thickness of the living body test part 110 (for example, about 10 mm for the fingertip and about 40 mm for the wrist), N is kept constant and k, that is, the modulation frequency f (∝k If you raise the), forced to reduce the total power P _0. However, since the magnitude of the shape function | F (kα ⁻¹ ) | does not increase in proportion to k, the detected sound wave decreases. Thus, it can be seen that too high a modulation frequency is also undesirable.

ここで、音圧上界ｐ_ｓｕｐは以下の数式（９）となる。

数式（８）で、｜Ｆ（ξ）｜／ξは、ξについて単調に減少する関数であり、信号振幅のみの観点では、低い変調周波数が有利となる。 The sound pressure amplitude _{p a} given by the equation (6), with N and _{I max,} rewritten becomes as follows.

Here, the sound pressure upper bound p _sup is expressed by the following formula (9).

In Equation (8), | F (ξ) | / ξ is a function that decreases monotonously with respect to ξ, and a low modulation frequency is advantageous from the viewpoint of only the signal amplitude.

In this case, the rate of change related to α in Equation (8), ∂p _a / ∂α = − (p _sup N / α) ξd (| F (ξ) | / ξ) / dξ is maximized ξ = kα ^-1 gives the optimum modulation frequency. Such xi] is model A204 in 2.49, model B205, and an on Model C206 ^{2 1/2,} in such ξ | F (ξ) | values of / xi], respectively, 0.620, It is calculated as 1/3 ^1/2 . That is, the optimum modulation frequency exists as a compromise between the conflicting requirements of the signal strength and the sensitivity to the absorption coefficient α.

As described above, since the light distribution in an actual living body is considered to be close to the model B205 and the model C206, the optimum modulation frequency is 2πf = 1.41cα, and at that time, the maximum value p _sup N at f → 0. On the other hand, a signal amplitude of 57.7% is expected.

Next, the principle of the blood component concentration measuring apparatus according to the present embodiment will be described with reference to FIG. The first light source 101 shown in FIG. 1 is intensity-modulated in synchronization with the oscillator 103, and the light output from the first light source 101 has a waveform shown as light 211 of the _first light source (λ ₁ ) in the upper part of FIG. It becomes.

On the other hand, the intensity of the second light source 105 shown in FIG. Here, since the signal transmitted from the oscillator 103 is given a phase shift of 180 ° by the 180 ° phase shift circuit 107, the light output from the second light source 105 is opposite to the light output from the first light source 101. The intensity is modulated by the phase signal, and the waveform shown as the light 212 of the _second light source (λ ₂ ) is shown in the lower part of FIG.

  Here, in FIG. 3, the signal for intensity-modulating the first light source 101 and the second light source 105 is a signal having a period of 1 μsec, that is, a modulation frequency f of 1 MHz and an occupation ratio of 50%. Show.

  Here, in Equation (4), a sinusoidal change is assumed for the irradiation light 201, and FIG. 3 shows the case of irradiation with rectangular wave light, but this is consistent with the following reason.

  That is, Equation (3) is linear, and components of different frequencies can be handled as being independent of each other. In addition, when the amplitude of the sound wave increases, it is affected by the nonlinearity of the Navier Stokes equation itself. However, in the case of the photoacoustic signal in the blood component concentration measuring apparatus according to the present embodiment, the generated sound wave is weak and is a linear formula. (3) is applicable. In addition, the rectangular wave includes odd-order harmonic components, but the amplitude of the sine wave component of the basic period may be read as I in Equation (4). The light source is easier to modulate the intensity into a rectangular wave shape than the sine wave shape, and the rectangular wave has a fundamental period sine wave component that is 4 / π = 1.27 times that of a sine wave of the same amplitude. The efficiency is slightly better.

  The two light beams having different wavelengths output from the first light source 101 and the second light source 105 are combined by the multiplexer 109 and applied to the living body test unit 110. Here, each of the two light beams having different wavelengths can be considered to independently generate the sound pressure expressed by the equation (6).

  Here, it is clear from the linearity of Equation (3) that sound waves are linearly superimposed. Further, since each of the two light beams having different wavelengths is not so strong that the absorption is saturated, the heat generation Q by each of the two light beams having different wavelengths is also linearly superimposed. Here, even when the absorption is saturated, if the absorption has a non-uniform spread and the wavelength interval between the two light beams having different wavelengths is wider than the uniform width, the linear superposition of the heat generation still remains. To establish. Here, such a condition is well satisfied for water in which absorption is common to the two light beams having different wavelengths.

ここで、α_ｉＦ（ｋα_ｉ ^―１）（ｉ＝１、２）が差の形で重畳されているのは、前記異なる波長の２波の光の各々の入射光が互いに逆相で強度変調された結果である。これを、超音波検出器１１３により検出し変換して得られる電気信号の中の基本周期の正弦波成分の波形を図６に実線で示す。図６に実線で示す信号の振幅（ｒｍｓ値）が、発振器１０３に同期した位相検波増幅器１１４によって測定され、図６にＶ_ｄとして示す信号として、出力端子１１５に出力される。 As described above, the photoacoustic signals having the sound pressures represented by the mathematical formula (6) are generated independently from each other by the two light beams having different wavelengths, and the sound pressure obtained by superimposing these signals is the ultrasonic detector 113. Is detected. Therefore, the sound pressure superimposed as described above is expressed by the following mathematical formula.

Here, α _i F (kα _i ⁻¹ ) (i = 1, 2) is superimposed in the form of a difference because the incident lights of the two light beams having different wavelengths are in opposite phases and intensities. This is a modulated result. The waveform of the sine wave component of the fundamental period in the electrical signal obtained by detecting and converting this with the ultrasonic detector 113 is shown by a solid line in FIG. Amplitude (rms value) of the signal shown by the solid line in FIG. 6 is measured by the phase detector amplifier 114 which is synchronized with the oscillator 103, as a signal shown in FIG. 6 as V _d, is outputted to the output terminal 115.

次に、数式（２）により、測定対象とする血液成分濃度の算出の原理を説明する。既に、第１の光源１０１および第２の光源１０５の各々が出力する光に対応する光音響信号の差信号ｓ_１−ｓ_２が得られているので、次に光音響信号ｓ_２を測定すれば、数式（２）から、測定対象の血液成分濃度Ｍを算出できる。 The unknown constant C is expressed by the following mathematical formula using the mathematical formula (10) and the mathematical formula (1).

Next, the principle of calculating the blood component concentration to be measured will be described using Equation (2). Since the difference signal s ₁ -s _{2 of the} photoacoustic signal corresponding to the light output from each of the first light source 101 and the second light source 105 has already been obtained, the photoacoustic signal s ₂ is measured next. For example, the blood component concentration M to be measured can be calculated from Equation (2).

Therefore, the photoacoustic signal is measured in a state where only the light 212 of the second light source (λ ₂ ) shown in FIG. 5 is irradiated. That is, as shown in FIG. 5, the output of the first light source 101 is set to zero while maintaining the waveform of the light output from the second light source 105. This can be realized by means such as blocking the light output from the first light source 101 shown in FIG. 1 with a mechanical shutter, or lowering the output of the drive circuit 104 below the oscillation threshold value of the first light source 101.

When the value of the photoacoustic signal measured in the above state is detected by the ultrasonic detector 113 and converted into an electric signal, a waveform indicated by a broken line in FIG. 6 is obtained as a fundamental periodic sine wave component. Further, it rms amplitude value of the waveform shown by the broken line in FIG. 6 is measured by the phase sensitive amplifier 114 similarly to the method described above, as the signal shown in FIG. 6 as V _r, is output to the output terminal 115.

Here, the photoacoustic signal s ₂ has a phase opposite to the difference signal s ₁ -s _{2 of the} photoacoustic signal. Further, the photoacoustic signal s ₂ is orders of magnitude larger than the difference signal s ₁ -s ₂ of the photoacoustic signal. For example, it is 1000 times or more in the case of blood glucose level measurement of a healthy person. Therefore, the sensitivity and time constant of the phase detection amplifier 114 are switched between the two measurements of the photoacoustic signal s ₂ and the difference signal s ₁ -s ₂ between the photoacoustic signals.

If two measurement values V _{d and} V _r are obtained by the above measurement, each of them is substituted into each of s ₁ -s ₂ and s _{2 in} the formula (2), and blood to be measured The component concentration M is calculated.

Here, for conversion from the ratio V _d / V _r of the measured values to the blood component concentration M, if the specific absorbance α ₁ ⁽⁰⁾ / α ₁ ^(b) (α ₂ ⁽⁰⁾ is non-zero, α ₂ ⁽⁰⁾ / α ₁ ^(b) ) is required.

FIG. 7 shows a method for selecting the _two wavelengths λ ₁ and λ _{2 to be} measured so that the specific absorbance value and the background absorption coefficient are equal as described above.

  FIG. 4 is a diagram illustrating a wavelength selection method for each of the first light source 101 and the second light source 105 in the blood component concentration measurement apparatus according to the present embodiment in the case of measuring a blood glucose level.

  FIG. 7 shows the absorbance (OD) of water and an aqueous glucose solution (concentration: 1.0 M) over a light wavelength of 1.2 μm to 2.5 μm. The absorbance OD has a relationship of α = ODln10 with the absorption coefficient α. The vertical axis on the right side of FIG.

  In FIG. 7, the absorption by the glucose molecules is observed only in the vicinity of 1.6 μm and in the vicinity of 2.1 μm, but the absorption by the glucose molecules is much smaller than that of water.

  The difference in absorbance between water and glucose is shown on the upper side of FIG. 8, and the specific absorbance obtained by dividing this by the absorbance of water is shown on the lower side of FIG.

According to the specific absorbance shown in FIG. 8, a clear maximum of absorption by glucose molecules is observed at 1608 nm and 2126 nm. Here, as an example, the wavelength λ ₁ of the first light source 101 is set to 1608 nm (specific absorbance is 0.114 M ⁻¹ ) as the absorption wavelength by glucose molecules. This is indicated by a vertical solid line with a circle in FIG.

Here, the absorption coefficient α ₁ ^(b) of the background (water) at the wavelength of 1608 nm can be read as 0.608 mm ⁻¹ from FIG. Therefore, the wavelength λ ₂ where α ₂ ^(b) = α ₁ ^(b) is the wavelength 1381 nm or the wavelength 1743 nm from the water absorption spectrum of FIG. For each of the candidates for the wavelength λ ₂ of the second light source 105, the value of α ₂ ⁽⁰⁾ is checked by the spectrum of the specific absorbance shown in FIG. As a result, the specific absorbance is zero at the wavelength of 1381 nm, while the wavelength of 1743 nm is in the absorption band of glucose molecules, and the specific absorbance is 0.0601M- ¹ . Since it is easier to measure the absorbance difference α ₁ ⁽⁰⁾ −α ₂ ⁽⁰⁾ as much as possible, 1381 nm is selected as the wavelength λ ₂ of the _second light source 105 in the above case.

When 2126 nm is set to the wavelength λ ₁ (specific absorbance is 0.0890 M ⁻¹ ) of the first light source 101 in the absorption band on the long wavelength side, the absorption coefficient α at a wavelength of 2126 nm is determined by the same method as described above. ₁ ^(b) = 2.361 mm ⁻¹ as a wavelength showing an absorption coefficient equal to 1837 nm or 2294 nm, both of which deviate from glucose absorption (indicated by a vertical dotted line in FIG. 8). As the wavelength λ ₂ of the second light source 105, either 1837 nm or 2294 nm may be selected.
(Example)

  Here, specific examples in the first embodiment will be described.

(First embodiment (part 1))
In the blood component concentration measuring apparatus according to the embodiment shown in FIG. 1, it is also effective to use laser light sources as the first light source 101 and the second light source 105. In selecting a laser light source, it is first necessary to estimate the required output power level.

There is an allowable limit of light intensity in the irradiation of light to the human body. Generally, 1/10 of the intensity at which 50% of individuals are damaged is established in JIS C6802 as the maximum allowable amount. According to JIS C6802, In continuous irradiation of non-visible infrared light (wavelength of 0.8 μm or more) to the skin, 1 mW per mm ² is the maximum allowable amount.

In the first embodiment, the blood component to be measured is blood glucose, the wavelength of light to be irradiated is 1.6 μm band, and the modulation frequency f is 150 kHz or more according to the principle described above. Here, the wavelength λ = c / f of the photoacoustic signal generated in the living body test part 110 is 10 mm or less. If the living body test unit 110 is a fingertip, the distance r between the irradiation unit irradiated with light and the detection unit where the ultrasonic detector 113 contacts the living body test unit 110 is 10 mm, and the Fresnel number N = w _0. ^The beam diameter w _{0 where} ² / (rλ) is 0.1 is calculated as w ₀ ² ≦ 10 mm ² . Multiplying this by π to calculate the beam area of the light to be irradiated, further adding the above-mentioned maximum allowable amount, and calculating the maximum light power that can be irradiated, it is 31 mW.

  Here, when the irradiation light is in the 2.1 μm band, the maximum power is calculated in the same manner to be 8 mW, and this light output can be sufficiently supplied by the semiconductor laser.

  Since the semiconductor laser is small and has a long lifetime, and also has an advantage that the intensity can be easily modulated by modulating the injection current, in this embodiment, the first light source 101 and the second light source 105 are: A semiconductor laser is used.

  FIG. 9 shows a configuration example of the blood component concentration measuring apparatus according to the first embodiment. The configuration of the first embodiment of the blood component concentration measuring apparatus according to this embodiment shown in FIG. 9 is a forward propagation type that detects a sound wave propagated in the direction of irradiation light, and the basic of the blood component concentration measuring apparatus shown in FIG. The first light source 101, the second light source 105, the drive circuit 104, the drive circuit 108, the 180 ° phase shift circuit 107, the multiplexer 109, the ultrasonic detector 113, shown in FIG. Each of the phase detection amplifier 114, the output terminal 115, and the oscillator 103 includes a first semiconductor laser light source 501 and a lens 502, a second semiconductor laser light source 505 and a lens 506, a drive current source 504, and a drive current source 508 shown in FIG. , 180 ° phase shift circuit 507, multiplexer 509, ultrasonic detector 513 and acoustic coupler 512, phase detection amplifier 514, output terminal 515, and oscillator 503, respectively. , Each has the same function.

  However, the lights output from the first semiconductor laser light source 501 and the second semiconductor laser light source 505 shown in FIG. 9 are converged into parallel light beams by the lens 502 and the lens 506, respectively. Are combined and applied to the living body test portion 510 as one light beam. Further, the acoustic coupler 512 shown in FIG. 9 is provided between the ultrasonic detector 513 and the living body test unit 510, and the photoacoustic signal transmission efficiency between the ultrasonic detector 513 and the living body test unit 510. It has a function to improve.

  FIG. 9 also shows a calibration sample 511, but the function of the calibration sample 511 will be described later.

  The first semiconductor laser light source 501 is intensity-modulated by the drive current source 504 in synchronization with the oscillator 503, and its output light is condensed into a parallel light flux by the lens 502 and input to the multiplexer 509, and also the second semiconductor. The laser light source 505 is intensity-modulated by the drive current source 508 in synchronization with the oscillator 503, and the output light is condensed into a parallel light beam by the lens 506 and input to the multiplexer 509. Here, since the output of the oscillator 503 is transmitted to the drive current source 508 via the 180 ° phase shift circuit 507, the light output from the second semiconductor laser light source 505 is from the first semiconductor laser light source 501. The intensity of the output light is modulated by an antiphase signal.

  The light output from each of the first semiconductor laser light source 501 and the second semiconductor laser light source 505 input to the multiplexer 509 is combined and applied to the living body test portion 510 as one light beam.

  The light irradiated on the living body test unit 510 generates a photoacoustic signal in the living body test unit 510, and the generated photoacoustic signal is detected by the ultrasonic detector 513 via the acoustic coupler 512, and the photoacoustic signal It is converted into an electric signal proportional to the sound pressure.

  The signal detected by the ultrasonic detector 513 and converted into an electric signal proportional to the sound pressure of the photoacoustic signal is synchronously detected, amplified and filtered by the phase detection amplifier 514 synchronized with the oscillator 503, and output to the output terminal 515. Is output.

Here, as described above, the wavelength of the first semiconductor laser light source 501 is set to 1608 nm, and the wavelength of the second semiconductor laser light source 505 is set to 1381 nm. Further, the oscillation frequency of the oscillator 503, that is, the modulation frequency f is set to 207 kHz so that ξ = kα ₁ ^(b) = ^2½ .

  The light output of the first semiconductor laser light source 501 is 5.0 mW, and the output light of the second semiconductor laser light source 505 is 5.0 mW.

The beam diameter of the light irradiated on the living body test unit 510 is set to w ₀ = 2.7 mm so that the distance r from the irradiation unit to the detection unit is 10 mm and the Fresnel number N is 0.1. ing.

In the above state, the irradiation intensity of the light of the combined light of the outputs of the first semiconductor laser light source 501 and the second semiconductor laser light source 505 to the skin of the living body test portion 510 is 0.44 mW / mm ² , It is a safe level that is more than twice the maximum allowable value. However, since this is a dangerous level for the eyes, light reflected or scattered from the acoustic coupler 512 during measurement or when the living body test part 510 is not placed does not directly enter the eyes. In addition, it is indispensable to install a light shielding hood (not shown in FIG. 9) in the multiplexer 509 and the living body test portion 510.

  The ultrasonic detector 513 is a frequency flat type electrostrictive element (PZT) incorporating a FET (field effect transistor) amplifier, and the acoustic coupler 512 is an acoustic matching gel.

In the above configuration, first, when the light output of the first semiconductor laser light source 501 is set to zero and only the light output from the second semiconductor laser light source 505 is irradiated as shown in FIG. A voltage of V _r = 20 μV was obtained as an electrical signal corresponding to the photoacoustic signal s ₂ at the output terminal 515 of the phase detection amplifier 514 set to seconds.

Here, the phase difference θ between the synchronization signal transmitted from the oscillator 503 in the phase detection amplifier 514 and the signal obtained by detecting the photoacoustic signal by the ultrasonic detector 513 and converted into an electrical signal is the phase difference θ of the living body test unit 510. Since the distance r between the irradiation unit irradiated with light and the contact unit contacting the acoustic coupler 512 and the modulation frequency f change, it is necessary to search for an optimum phase difference for each measurement. The search for the phase difference is effectively performed by measuring the photoacoustic signal s ₂ having a large signal amplitude.

Here, in the case of the two-phase type phase detection amplifier 514, there is always the ability to automatically obtain the phase difference θ, and it is not necessary to manually search for the phase difference. That is, the photoacoustic signal s ₂ is measured by the R-θ mode capable of measuring the unknown phase and amplitude, the phase and amplitude are obtained, and the noise suppression ratio is improved by 3 dB when the phase is known by using the measured value of this phase. The difference signal s ₁ -s ₂ of the above-described photoacoustic signal is measured by the X measurement mode in which the amplitude can be measured in the above state.

Next, when the first semiconductor laser light source 501 is caused to emit light, an electric signal corresponding to the photoacoustic signal difference signal s ₁ -s ₂ is output to the output terminal 515 as V _d = 7.7 nV (since the phase is inverted, The measured value was −7.7 nV). Subsequently, again, the optical output of the first semiconductor laser light source 501 is set to zero, the sensitivity and time constant of the phase detection amplifier 514 are restored, the photoacoustic signal s ₂ is measured, and V _r = 22 μV is obtained. It was. The average of the front and rear two _{V r,} the value of _{V r} became 21MyuV.

As described above, it is desirable to measure the signal corresponding to the photoacoustic signal s ₂ , that is, V _r twice, before and after the measurement of the difference signal s ₁ -s ₂ of the photoacoustic signal.

By the above procedure, during measurement of the difference signal s ₁ -s ₂ , the unknown multiplier C derived from a change in the distance r due to a change in the pressing force of the subject's fingertip, a local temperature change due to light irradiation, and the like. Drift can be corrected.

The glucose concentration M was determined to be 3.2 mM (58 mg / dl) based on the above measured value, the specific absorbance value 0.114 M ^{−1 at} the wavelength of 1608 nm, and the mathematical formula (2).

Value for water, C _p = 1 (cal / g · deg) = 4.18 × 10 ³ (J / kg · K), β = 300 ppm / deg, c = 1.51 × 10 ³ (m / s) When I _max = 1 mW / mm ² , 0.17 Pa is obtained as the sound pressure upper bound p _sup . Thereto, the Fresnel number N ⁼ 0.1, ξ = ^{2 1/2} to involve decay ^{1/3 1/2} and, multiplied by the actual irradiation power ratio 0.22, the expected sound pressure amplitude is 2. 1 mPa.s.

On the other hand, the nominal sensitivity of the ultrasonic detector 513 is 66 mV / Pa and the output voltage of the output terminal 515 is predicted to be 140 μV, but the measured value of the photoacoustic signal s ₂ is 1 The reason for staying at / 7 is considered to be an imperfection of the acoustic coupler 512.

(First embodiment (part 2))
In this embodiment, for the purpose of improving the acoustic coupling state, an acrylic plate having a thickness of 6.6 mm is formed to have the same diameter as the ultrasonic detector 513 and 10 mmφ in order to make the acoustic coupler 512 resonant. . One surface of the acoustic coupler 512 is attached to the ultrasonic detector 513 via vacuum grease, and the other surface is in contact with the living body test portion 510 via an acoustic matching gel.

As a result of measuring twice in the same procedure as described above with the above configuration, the measured values of the photoacoustic signal s ₂ signal are 150 μV and 153 μV, and the measured value of the difference signal s ₁ -s ₂ of the photoacoustic signal is It was 59 nV. In the above measurement, the time constant of the phase detection amplifier 514 is 3 seconds. From these measured values, the glucose concentration M was determined to be 3.4 mM (61 mg / dl).

(First embodiment (part 3))
In the first embodiment (part 2), the resonance frequency of the acoustic coupler 512 and the modulation frequency f do not completely match. Therefore, in this embodiment, when measuring the photoacoustic signal s _{2 in} the previous stage, first, the frequency of the oscillator 503 is swept within a range of several percent, and the two-phase type phase detection amplifier 514 is operated in the R-θ mode. By operating and setting the modulation frequency f so that the output of the signal output terminal 515 is maximized, the resonance frequency of the acoustic coupler 512 and the modulation frequency f are completely matched.

Except for the above, 600 μV and 604 μV were obtained as the photoacoustic signal s ₂ by two measurements by the same procedure as in the previous example. The measured value of the difference signal s ₁ -s ₂ of the photoacoustic signal was 0.25 μV. In this case, the time constant of the phase detection amplifier 514 is 1 second.
From the above measured values, the glucose concentration M was determined to be 3.6 mM (65 mg / dl).

  In the first embodiment (part 1), the first embodiment (part 2), and the first embodiment (part 3), a frequency flat type electrostrictive element (PZT) is used as the ultrasonic detector 513. However, even in the case of a normal type electrostrictive element (PZT), it is possible to carry out sensitization measurement using resonance characteristics by searching for a modulation frequency f that maximizes the amplitude of the signal obtained at the output terminal 515. It is effective for downsizing and cost reduction.

(First embodiment (part 4))
In this embodiment, a calibration sample 511 is introduced as means for adjusting the powers of light output from the first semiconductor laser light source 501 and the second semiconductor laser light source 505 to be equal.

  As a configuration of the calibration specimen 511, water in which a scatterer such as latex particles is dispersed is sealed so as to simulate scattering in a glass container in which water is sealed or in a living body.

Here, in order to ensure the uniformity of the transmittance with respect to the wavelength λ ₁ and the wavelength λ ₂ of the glass (upper surface in FIG. 9) of the calibration sample 511, the upper surface of the calibration sample 511 is irradiated. It is effective to provide a pipe-shaped edge with a diameter through which the beam passes to prevent direct contact with the surface, or to perform cleaning by a predetermined product and procedure before using the calibration sample 511.

  The calibration procedure performed by mounting the calibration sample 511 as described above instead of the living body test portion 510 is as follows.

  First, the light output of the first semiconductor laser light source 501 is set to zero, and only the light output from the second semiconductor laser light source 505 is irradiated as shown in FIG. Here, the two-phase type phase detection amplifier 514 is operated in the R-θ mode, and the phase θ at this time is obtained and fixed. In the case of resonance type ultrasonic detection, the optimum modulation frequency f is searched at this stage in order to make the resonance frequency of the acoustic coupler 512 coincide with the modulation frequency f in the same manner as described above.

  Furthermore, while increasing the light output from the first semiconductor laser light source 501, it is observed that the signal output to the output terminal 515 of the phase detection amplifier 514 decreases, and the signal output to the output terminal 515 decreases. At the same time, the sensitivity and time constant of the phase detection amplifier 514 are switched, and when the output obtained at the output terminal 515 becomes zero, the optical output of the first semiconductor laser light source 501 is fixed.

  Using the calibration sample 511 according to the above procedure, the relative intensities of the light output from the first semiconductor laser light source 501 and the light output from the second semiconductor laser light source 505 are equal and opposite in phase. Can be calibrated to a state in which the intensity is modulated.

  With the calibration sample 511 mounted in place of the living body test portion 510, the usage method of turning on the blood component measuring instrument according to the present embodiment was established, and the above sequence was determined as POST (Power On Self Test).

(Second embodiment (part 1))
The second embodiment is a backward propagation type that detects sound waves propagating in a direction reverse to the irradiation light 201. As shown in FIG. 10, the configuration of the second embodiment is the same as that of the first embodiment of the blood component concentration measuring apparatus shown in FIG. 9, but the acoustic coupler 512 is located between the multiplexer 509 and the living body test portion 510. One surface of the acoustic coupler 512 is in contact with the living body test portion 510, and light combined by the multiplexer 509 is incident from the other surface of the acoustic coupler 512, and this incident light is input to the acoustic coupler It is changed so as to pass through 512 and irradiate the living body test portion 510. Here, the ultrasonic detector 513 is installed on the side where the combined light is incident on the acoustic coupler 512.

  Further, the operation of the blood component concentration measuring apparatus according to the second embodiment is different from that of the first embodiment, as shown in FIG. 10, the light output from the multiplexer 509 passes through the acoustic coupler 512. Then, the photoacoustic signal emitted to the living body test portion 510 and generated in the living body test portion 510 is propagated again through the acoustic coupler 512 and detected by the ultrasonic detector 513.

  In the above configuration, since the irradiation light 201 passes through the acoustic coupler 512, it is desirable that the light absorption is small and the acoustic impedance is close to that of the living body (water).

  In the present embodiment, the acoustic coupler 512 is made of quartz glass with little light absorption. The acoustic impedance of quartz glass is eight times that of water, and only about 1/5 of the generated sound pressure becomes a propagation wave in the quartz glass and is observed by the ultrasonic detector 513. Therefore, since it is disadvantageous in terms of sensitivity, it is essential that the acoustic coupler 512 itself has resonance characteristics to increase the sensitivity. That is, the thickness of the quartz glass (corresponding to the propagation length of the light beam in the glass in the figure) is set to 14 mm, which is a value of an approximate half wavelength of the wavelength λ of 27.85 mm with respect to the modulation frequency f of 200 kHz.

  Since the sound wave in the quartz glass can be regarded as a spherical wave at a distance from the living body test portion 510, the ultrasonic detector 513 is installed in a direction that forms an angle of 150 ° with the incident light beam. (If an ultrasonic detector having a hole through which an incident light beam passes is used, it can be placed in the rear 180 ° direction completely.)

  In the configuration of the present embodiment, the distance r between the irradiation unit irradiated with light in the living body test unit 510 and the detection unit in which the photoacoustic signal is detected by the ultrasonic detector 513 in the acoustic coupler 512 is: It is fixed to a constant value (r = 14 mm in this case) determined by the size of the acoustic coupler 512.

  Here, the first semiconductor laser light source 501, the second semiconductor laser light source 505, and the ultrasonic detector 513 are the same as those in the first embodiment. Further, as a safety measure, a sample detection switch (not shown in FIG. 10) is provided so that light is not irradiated when nothing is placed on the acoustic coupler 512.

Similarly to the first embodiment, the frequency of the oscillator 503 is swept when the photoacoustic signal s ₂ in the previous stage is measured, and a modulation frequency f that matches the resonance frequency of the acoustic coupler 512 is searched. Furthermore, 200 μV and 206 μV were obtained as the photoacoustic signal s ₂ by two measurements according to the same procedure as in the first example. As the difference signal s ₁ -s ₂ of the photoacoustic signal, 79 nV was measured with the time constant of the phase detection amplifier 514 as 1 second. From these measured values, the glucose concentration M was determined to be 3.4 mM (61 mg / dl).

(Second embodiment (part 2))
In this embodiment, the acoustic coupler 512 is made of low density polyethylene. Low density polyethylene has an acoustic impedance that differs only by 18% with respect to water and is very good at sonic coupling (less than 9% pressure loss). However, there is some light absorption and it is difficult to be too soft. However, it is excellent in that the close contact with the living body is good due to flexibility, and no filler such as an acoustic matching gel is required. Here, high-density polyethylene with higher rigidity is not suitable because it does not transmit light.

  In this embodiment, the thickness of the acoustic coupler 512 is 10 mm, which is substantially equal to the wavelength of the sound wave with respect to the modulation frequency f of 200 kHz, and the distance r between the irradiation unit and the detection unit is also a fixed value of 10 mm.

Similar to the second example (part 1), 300 μV and 289 μV were obtained as the photoacoustic signal s ₂ by two measurements. As the difference signal s ₁ -s ₂ of the photoacoustic signal, 117 nV is measured with the time constant of the phase detection amplifier 514 as 1 second, and the glucose concentration M is 3.5 mM (63 mg / dl) from these measured values. I asked.

  Here, although the pressure loss of the low density polyethylene is low, the measurement signal does not increase because the acoustic coupler 512 is deformed by the pressing of the living body test portion 510, the dimensions become unstable, and the sensitivity improvement due to resonance is not improved. This is enough.

(Second embodiment (part 3))
This embodiment is a case where the calibration means using the calibration sample 511 is introduced in the second embodiment (part 2). In this case, in the calibration sample 511, the material of the container that encloses water or water containing scatterers is the same as the material of the acoustic coupler 512.

  Here, the surface irradiated with light is a surface in contact with the acoustic coupler 512 shown in FIG. 10 of the calibration specimen 511. In order to ensure long-term clarity, a predetermined specimen is used before the calibration specimen 511 is used. Carry out cleaning with supplies and procedures.

  The calibration procedure performed by mounting the calibration sample 511 in place of the biological specimen 510 is the same as in the first embodiment (part 4).

(Third embodiment)
In the first embodiment and the second embodiment, an example of blood glucose concentration, that is, blood glucose level is shown. However, as components constituting blood, in addition to glucose, many components such as cholesterol, lipids, proteins and inorganic components are included. The third example shows an example in which the blood component concentration measuring device and the blood component concentration measuring device control method according to the first embodiment are applied to cholesterol. In addition, the blood component concentration measuring apparatus and the blood component concentration measuring apparatus control method in the present embodiment can be similarly applied to the embodiments described later.

  FIG. 11 shows the absorbance of water over a wavelength range of 1200 nm to 2500 nm. FIG. 12 shows the absorbance of cholesterol ranging from 1600 nm to 2600 nm. According to the spectrum shown in FIG. 12, a clear maximum of absorption by cholesterol molecules is observed at 2310 nm.

Here, the absorption coefficient α ₁ ^(b) of the background (water) at the wavelength of 2310 nm can be read as 1.19 mm ⁻¹ from FIG. Therefore, the wavelength λ _{2 at which} α ₂ ^(b) = α ₁ ^(b) is the wavelength 2120 nm or the wavelength 1880 nm from the water absorption spectrum of FIG. For each of the candidates for the wavelength λ ₂ of the second light source 105, the value of α ₂ ^(b) is confirmed by the absorption spectrum of FIG. As a result, it can be seen that cholesterol molecules have a large absorption at a wavelength of 1880 nm as compared to absorption at a wavelength of 2120 nm. Since it is easier to measure the absorbance difference α ₁ ⁽⁰⁾ −α ₂ ⁽⁰⁾ as much as possible, 2120 nm is selected as the wavelength of the second light source in the above case. As a result, the measurement was performed by setting the wavelength of the first light source to 2310 nm and the wavelength of the second light source to 2120 nm.

  FIG. 13 shows the configuration of a blood component concentration measuring apparatus according to the third embodiment. The configuration of the third embodiment of the blood component concentration measuring apparatus shown in FIG. 13 is a forward propagation type that detects a photoacoustic signal propagated in the direction of irradiation light, and the basic configuration of the blood component concentration measuring apparatus shown in FIG. It is a similar configuration.

  That is, the first semiconductor laser light source 801 and the lens 802, the second semiconductor laser light source 805 and the lens 806, the drive current source 804, the drive current source 808, the 180 ° phase shift circuit 807, the multiplexer 809 shown in FIG. The ultrasonic detector 813, the acoustic coupler 812, the phase detection amplifier 814, the output terminal 815, and the oscillator 803 are the first light source 101, the second light source 105, the drive circuit 104, the drive circuit 108, 180 ° shown in FIG. The phase shift circuit 107, the multiplexer 109, the ultrasonic detector 113, the phase detection amplifier 114, the output terminal 115, and the oscillator 103 have the same functions.

  The first semiconductor laser light source 801 is intensity-modulated by the drive current source 804 in synchronization with the oscillator 803, and the output light is condensed into a parallel light flux by the lens 802 and input to the multiplexer 809. The intensity of the second semiconductor laser light source 805 is also modulated in synchronization with the oscillator 803 by the drive current source 808, and the output light is condensed into a parallel light flux by the lens 806 and input to the multiplexer 809.

  Here, since the output of the oscillator 803 is transmitted to the drive current source 808 via the 180 ° phase shift circuit 807, the light output from the second semiconductor laser light source 805 is the first semiconductor laser light source 801. The intensity of the output light is modulated with an antiphase signal.

  The light output from each of the first semiconductor laser light source 801 and the second semiconductor laser light source 805 input to the multiplexer 809 is combined and applied to the living body test unit 810 as the subject as one light beam. The

  The light applied to the living body test unit 810 generates a photoacoustic signal in the living body test unit 810. The generated photoacoustic signal is detected by the ultrasonic detector 813 through the acoustic coupler 812. The ultrasonic detector 813 converts the photoacoustic signal into an electric signal proportional to the sound pressure.

  The signal converted into the electric signal is synchronously detected and amplified by the phase detection amplifier 814 synchronized with the oscillator 803, filtered, and then output to the output terminal 815.

The wavelength of the first semiconductor laser light source 801 is set to 2310 nm, and the wavelength of the second semiconductor laser light source 805 is set to 2120 nm. Further, the oscillation frequency of the oscillator 803, that is, the modulation frequency f is set to 207 kHz so that ξ = kα ₁ ^(b) = 2 ^1/2 .

  The light output of the first semiconductor laser light source 801 is 5 mW, and the light output of the second semiconductor laser light source 805 is also 5 mW.

Beam diameter of the light irradiated to the living body test region 810, as 10mm the distance r to the detection portion from the irradiation portion of the living body test region 810, w _{0 =} 2 as the Fresnel number N becomes 0.1. It is set to 7 mm.

In the above state, the irradiation intensity of the light of the combined output of the first semiconductor laser light source 801 and the second semiconductor laser light source 805 to the skin of the living body test part 810 is 0.44 mW / mm ² . This is a safe level that is less than half of the maximum allowable value. However, it is preferable to install a light shielding hood (not shown) that covers the multiplexer 809 and the living body test part 810 in consideration of leakage to the outside.

  The ultrasonic detector 813 is a frequency flat type electrostrictive element (PZT) incorporating a field effect transistor (FET) amplifier. An acoustic matching gel was used as the acoustic coupler 812.

In FIG. 13 having the above-described configuration, first, when the light output of the first semiconductor laser light source 801 is set to zero and only the output light of the second semiconductor laser light source 805 is irradiated, the phase in which the time constant is set to 0.1 second is used. A voltage of Vr = 40 μV was obtained as an electrical signal corresponding to the photoacoustic signal s ₂ at the output terminal 815 of the detection amplifier 814.

The phase difference θ between the synchronization signal from the oscillator 803 input to the phase detection amplifier 814 and the signal detected by the ultrasonic detector 813 and converted into an electrical signal is the irradiation of the living body test part 810 irradiated with light. Therefore, an optimum phase adjustment is required for each measurement because it varies depending on the distance r between the contact portion and the contact portion of the living body test portion 810 that contacts the acoustic coupler 812 and the modulation frequency f. Such phase adjustment, it is effective to implement by measuring a large photoacoustic signal s ₂ of the signal amplitude as a phase reference.

When the phase detection amplifier 814 is a two-phase type, it is possible to always have a function of automatically tracking the phase difference θ, and by using this function, the phase difference can be automatically adjusted. That is, when the R-θ mode is used to measure the unknown phase and amplitude, the phase and amplitude of the photoacoustic signal s ₂ are measured, and the noise suppression ratio is improved by 3 dB when the phase is known using this phase. The above-described photoacoustic signal difference signal s ₁ -s ₂ is measured in the X measurement mode in which the amplitude can be measured with

Next, when the first semiconductor laser light source 801 was caused to emit light, an output of about 10 nV was obtained at the output terminal 815 as the electric signal Vd corresponding to the difference signal s ₁ -s ₂ of the photoacoustic signal. Subsequently, again, the optical output of the first semiconductor laser light source 801 was set to zero, the sensitivity and time constant of the phase detection amplifier 814 were restored, and the photoacoustic signal s ₂ was measured. As a result, Vr = 42 μV. Output was obtained. The average of Vr of these two times before and after was Vr = 41 μV.

As described above, before and after the difference signal s ₁ -s ₂ of the photoacoustic signal measurement, the measurement of Vr is a signal corresponding to the photoacoustic signal s ₂ is performed twice desirable. According to the above procedure, during measurement of the difference signal s ₁ -s ₂ , the drift of the unknown multiplier C derived from the variation in the distance r due to the change in the pressing force of the fingertip of the subject and the local temperature change due to light irradiation, etc. I was able to correct it.

When the photoacoustic signal derived from cholesterol in the living body test part is measured by an ultrasonic detector using the above measurement system, an output value of several hundred nV can be obtained as the difference signal s ₁ -s ₂ of the photoacoustic signal. did it.

  Here, the blood component concentration measuring device of the living body and the method for controlling the blood component concentration measuring device of the living body have been described, but the same applies to a case where liquid is used instead of the living body. In other words, the liquid component concentration measuring apparatus and the liquid component concentration measuring apparatus control method according to the present embodiment are based on the description of the direct photoacoustic method that is the basis of the above-described embodiment and the component concentration calculating method shown in Equation (2). As is easily known, it can be performed on a measurement object other than a living body. In this case, in general, when two wavelengths having the same absorption coefficient for the liquid and different absorption coefficients for the target substance are used, the components in the liquid can be detected without being covered by the absorption of the liquid. Furthermore, in the configuration of the above-described embodiment or example, if a fruit is used instead of the living body test part, it functions as a fruit sugar content meter. This is because sucrose and fructose, which are the sweetness components of fruits, have absorption at a wavelength similar to that of glucose, which is a blood sugar component. Thus, in the range not departing from the spirit of the present embodiment, It goes without saying that the measuring apparatus and the measuring apparatus control method can be applied to various objects.

(Fourth embodiment)
FIG. 14 shows the configuration of a blood component concentration measuring apparatus according to the fourth embodiment. The fourth embodiment is a case where a contact thermometer 138 is further introduced into the blood component concentration measuring apparatus described in the first embodiment (part 1) to (part 4).

  In the configuration shown in FIG. 14, the wavelength of the first light source 101 is set to 1608 nm, while the wavelength of the second light source 105 is set to 1381 nm. These wavelength values are equal to the water temperature 39 shown in FIG. Based on the absorbance of water at ° C. The reference temperature of 39 ° C. is a value higher than the normal temperature as the body temperature, and strictly speaking, it is necessary to change the wavelength setting of the laser light source in accordance with the body temperature of the subject, that is, the temperature of the living body test part 110. This is because the light absorption characteristics of water change depending on the water temperature.

  The water absorbance depending on the water temperature is shown in FIG. FIG. 15 shows the absorbance of an absorption band of water having a maximum in the vicinity of a wavelength of 1450 nm for a water temperature in increments of 5 ° C. from 25 ° C. to 55 ° C. using the water temperature as a parameter. As shown in FIG. 14, the water absorption band shifts in the short wavelength direction as the water temperature increases, and accordingly, the absorption on the short wavelength side increases while the absorption on the long wavelength side decreases.

In order to see this property in detail, FIG. 16 is obtained when the temperature change of the absorbance of water at a certain wavelength is shown. At the wavelength of 1608 nm of the first light source 101 on the long wavelength side, the absorbance of water decreases at a rate of 1.366 × 10 ⁻³ mm ⁻¹ / ° C. with respect to the temperature. On the other hand, at a wavelength of 1381 nm of the second light source 105 on the short wavelength side, the absorbance of water increases at a rate of 1.596 × 10 ⁻³ mm ⁻¹ / ° C.

As a result, the difference in absorbance between the two wavelengths decreases with respect to the temperature at a rate of 2.962 × 10 ⁻³ mm ⁻¹ / ° C. and the specific absorbance at 1.001 × 10 ⁻² / ° C. When the specific absorbance value 0.114 M ⁻¹ of glucose at 1608 nm is used for this rate of change, the glucose concentration M is underestimated as 87.78 mM (1581 mg / dl) per 1 / ° C. of deviation from the reference temperature of the body temperature. It turns out that occurs.

  As a countermeasure against this error, a contact thermometer 138 is installed on the light irradiation side of the living body test part 110, the local body temperature in the vicinity of the light irradiation part is measured, and the temperature obtained by subtracting the reference temperature from this temperature measurement value. A value obtained by multiplying the difference by the correction coefficient 1581 mg / dl / ° C. is added to the calculated value of the glucose concentration M according to the above equation (2). The reason for installing the contact thermometer 138 on the light irradiation side is that the reason for the correction is the temperature of the irradiation side surface of the living body test part 110 where light absorption occurs. For example, if the temperature of the living body surface on the side in contact with the ultrasonic detector 113 is substituted, the body surface temperature that has undergone a change due to inevitable thermal contact with the ultrasonic detector 113 will be used. There is a risk of causing a large error.

  In addition, when a calibration sample is used in the blood component concentration measurement apparatus shown in FIG. 17, the correction based on the surface body temperature measurement may be performed as follows. FIG. 18 shows an embodiment in which a calibration sample 141 is further applied to the blood component concentration measuring apparatus shown in FIG.

  A thermometer 143 that measures the liquid temperature in the calibration sample 141 is attached to the calibration sample 141. In the above procedure, the reading of the thermometer 143 at the time when the output of the photoacoustic signal from the output terminal 115 becomes zero and the output of the drive circuit 104 is fixed is recorded as the calibration temperature. In the measurement for the living body test part 110 after the dredging, correction is performed using the calibration temperature instead of the reference temperature of the correction algorithm shown in the embodiment. That is, the local thermometer near the light irradiation part is measured by the contact thermometer 138, and the value obtained by multiplying the temperature difference obtained by subtracting the calibration temperature from the temperature measurement value by the correction coefficient 1581 mg / dl / ° C. What is necessary is just to add to the calculated value of the glucose concentration M by (2).

  When a constant temperature means (not shown in FIG. 18) for keeping the liquid temperature constant is installed in the calibration sample 141, the thermometer 143 and the living body test unit 110 installed in the calibration sample 141 are measured at the time of measuring the living body test unit 110. It is also possible to simultaneously operate the contact thermometer 138 and obtain the temperature difference from the difference in reading. In particular, in this case, if the thermometer 143 and the contact thermometer 138 are the same type of thermometer, an equilibrium configuration can be obtained in which the difference between the output values of the two is read with high accuracy using, for example, a bridge circuit. In such a balanced configuration, the accuracy of the absolute temperature is not required for the thermometer 143 and the contact thermometer 138, and therefore, it can be implemented using a simple temperature sensor such as a thermistor.

(5th Example)
FIG. 19 shows the configuration of a blood component concentration measuring apparatus according to the fifth embodiment. The fifth embodiment is a case where a contact thermometer 138 is further introduced into the blood component concentration measuring apparatus described in the second embodiment (part 1) to (part 3).

  In the present embodiment, the contact thermometer 138 for correction based on the surface body temperature measurement is preferably embedded in the surface of the acoustic coupler 142 that contacts the living body test part 110. In this case, it is desirable to use a contact thermometer 138 having an acoustic impedance approximate to the acoustic impedance of the acoustic coupler 142. This is for suppressing the propagation of the ultrasonic wave in the acoustic coupler 142 from being disturbed by the contact thermometer 138. Thereafter, the correction based on the surface body temperature value measured by the contact thermometer 138 is performed by the same calculation method as in the fourth embodiment. Further, a calibration procedure performed by mounting the calibration sample 141 in place of the biological specimen 110, correction based on surface body temperature measurement by the contact thermometer 138, and the like may be performed in accordance with the fourth embodiment. .

(Second Embodiment)
FIG. 20 is a schematic configuration diagram of a component concentration measuring apparatus according to the present embodiment. The component concentration measuring apparatus according to the present embodiment measures a sound wave that is not affected by the concentration of a target component such as glucose. Therefore, the configuration described in the first embodiment is further provided with a third light source 106 as calibration light generating means. This will be specifically described below.

  A component concentration measuring apparatus 150 shown in FIG. 20 includes a first light source 101 and a second light source 105 as measurement light generating means, a third light source 106 as calibration light generating means, an oscillator 103, and a light source. Drive circuits 104 and 108 as modulation means, 180 ° phase shift circuit 107, multiplexers 109a and 109b as optical multiplexing means, irradiation head 112 as light emitting means, and ultrasonic detection as acoustic wave detection means And a container 113. In the present embodiment, the light source switching 145, the phase detection amplifier 114, the output terminal 115, the lenses 139, 140, 144, the acoustic coupler 142 as acoustic matching means, the calculator 129, and the calibration sample 141a are further provided. , 141b will be described.

  In the component concentration measuring device control method according to the present embodiment, a control circuit (not shown) provided in the component concentration measuring device 150 shown in FIG. And a measurement sound wave detection procedure. In the standardized sound wave detection procedure, the control circuit included in the component concentration measurement device 150 causes the first light source 101, the oscillator 103, the drive circuit 104, the irradiation head 112, and the ultrasonic detector 113 to function. In the measurement sound wave detection procedure, the control circuit included in the component concentration measurement device 150 includes the first light source 101, the second light source 105, the oscillator 103, the drive circuits 104 and 108, the 180 ° phase shift circuit 107, the irradiation head 112, and the like. Each of the ultrasonic detectors 113 is caused to function.

  The component concentration measurement apparatus control method according to the present embodiment further includes a temperature change measurement procedure. In the temperature change measurement procedure, the control circuit included in the component concentration measurement apparatus 150 includes the first light source 101, the third light source 106, the oscillator 103, the drive circuits 104 and 108, the 180 ° phase shift circuit 107, the irradiation head 112, and the super head. Each of the sound wave detectors 113 is caused to function.

  In the component concentration measurement device control method according to the present embodiment, a control circuit (not shown) provided in the component concentration measurement device 150 further includes a calibration sound wave determination unit 151, a temperature calculation unit 152, an output unit 153, and a concentration calculation unit 154. It is preferable to cause each of the density correction means 155 to function.

  In the present embodiment, one of the measurement light generation means is described as the first light source 101, the other of the measurement light generation means is described as the second light source 105, and the calibration light generation means is described as the third light source 106. The first light source 101 and the second light source 105 constitute two measurement light generating means. The first light source 101 and the second light source 105 generate and output light having different wavelengths with equal absorption of water in blood obtained by mixing glucose as a target component with a liquid such as water. The third light source 106 generates and outputs light having the same wavelength as that of the light source for measurement, for example, the first light source 101, which is absorbed by glucose as a target component.

FIG. 21 is an example of output wavelengths of the first light source, the second light source, and the third light source. (A) shows the absorbance and extinction coefficient of water and an aqueous glucose solution, and (b) shows the absorbance of glucose. The difference in absorbance and the absorbance ratio comparing the absorbance of water are shown. λ ₁ is a wavelength output from the first light source, λ ₂ is a wavelength output from the second light source, and λ ₃ is a wavelength output from the third light source. As shown in FIG. 21 (a), the absorbance of water in the wavelength lambda ₁ and wavelength lambda ₂ are equal. The absorbance in water with wavelength λ ₃ is different from the absorbance in water with wavelengths λ ₁ and λ ₂ . On the other hand, as shown in FIG. 21 (b), the absorbance of glucose wavelength lambda ₁ and wavelength lambda ₃ are substantially equal. The absorbance of glucose at wavelength λ ₂ is different from the absorbance of glucose at wavelengths λ ₁ and λ ₃ . The third light source is used to measure the absorbance that is not affected by the glucose concentration. Furthermore, it is preferable that the specific absorbance of glucose of wavelength λ ₂ output from the second light source is zero.

  Here, a setting example of the relative output power of the first light source 101 and the second light source 105 illustrated in FIG. 20 will be described. The output powers of the first light source 101 and the second light source 105 are set using the calibration sample 141a. The calibration sample 141a is a sample used as a reference for absorption exhibited by a liquid in a solution obtained by mixing a target component with a liquid. When the solution is blood, the calibration sample 141a is a substance close to a liquid component in blood, for example, water. The first light source 101 and the second light source 105 are intensity-modulated in opposite phases, irradiated to the calibration sample 141a, and the intensity of the sound wave generated in the calibration sample 141a is measured. If the relative powers of the light emitted from the first light source 101 and the second light source 105 are equal, the intensity of the sound wave generated by the calibration sample 141a is zero. For this reason, if the relative power of the second light source 105 with respect to the first light source 101 is shifted so that the intensity of the sound wave generated by the calibration sample 141 a becomes zero, the first light source 101 and the second light source 105 can be changed. The relative output power can be set.

  An example of setting the relative output power of the first light source 101 and the third light source 106 will be described. The relative powers of the first light source 101 and the third light source 106 are set using the calibration sample 141b. The calibration sample 141b is a sample used as a reference for absorption exhibited by the target component in a solution obtained by mixing the target component with a liquid. When the target component is glucose, the calibration sample 141b is glucose with a low water content, for example, a crystal of glucose. The first light source 101 and the third light source 106 are intensity-modulated in opposite phases, irradiated to the calibration sample 141b, and the intensity of the sound wave generated by the calibration sample 141b is measured. If the relative powers of the light emitted from the first light source 101 and the third light source 106 are equal, the intensity of the sound wave generated by the calibration sample 141b becomes zero. Therefore, if the relative power of the third light source 106 is shifted so that the intensity of the sound wave generated in the calibration sample 141b becomes zero, the relative output powers of the first light source 101 and the third light source 106 can be reduced. Settings can be made.

  The drive circuits 104 and 108 intensity-modulate the light from the first light source 101 and the light from the second light source 105 and the third light source 106 in opposite phases with each other at a constant frequency oscillated by the oscillator 103. Output. For example, the drive circuit 104 modulates the intensity of the light from the first light source 101 at a predetermined frequency that is oscillated by the oscillator 103 and outputs the modulated light. The drive circuit 108 and the 180 ° phase shift circuit 107 have the intensity of the light from the second light source 105 at a predetermined constant frequency oscillated by the oscillator 103 and in the opposite phase to the light from the first light source 101. Modulate and output. In addition, the drive circuit 108 and the 180 ° phase shift circuit 107 are configured so that the light from the third light source 106 is in a phase that is opposite to the light from the first light source 101 at a predetermined constant frequency that the oscillator 103 oscillates. Output with intensity modulation.

  The irradiation head 112 emits the intensity-modulated light, the measurement combined light, and the calibration combined light from the first light source 101 toward the living body test unit 110 as light emitting means. As a result, the intensity-modulated light of the light from the first light source 101, the synthetic light for measurement, or the synthetic light for calibration is irradiated onto the blood existing in the living body test part 110. Here, the combined light for measurement is light obtained by combining the intensity-modulated light of the light from the first light source 101 and the intensity-modulated light of the light from the second light source 105.

  When the intensity-modulated light of the light from the first light source 101 is emitted toward the living body test unit 110, the light source switching 145 opens the drive circuit. That is, the drive circuit 108 is not connected to either the first light source 101 or the third light source 106. The lens 139 causes the intensity-modulated light from the first light source 101 to enter the multiplexer 109b. The multiplexer 109 b guides the intensity modulated light from the first light source 101 to the irradiation head 112. The irradiation head 112 emits the intensity-modulated light from the first light source 101 toward the living body test unit 110.

  When the synthetic light for measurement is emitted toward the living body test unit 110, the light source switch 145 connects the drive circuit 108 and the second light source 105. The lens 140 causes the intensity-modulated light from the second light source 105 to enter the multiplexer 109a. The multiplexer 109a causes the intensity-modulated light from the second light source 105 to enter the multiplexer 109b. The lens 139 causes the intensity-modulated light from the first light source 101 to enter the multiplexer 109b. The multiplexer 109 b multiplexes the intensity-modulated light from the first light source 101 and the intensity-modulated light from the second light source 105. Then, the irradiation head 112 supplies the measurement combined light, which combines the intensity-modulated light from the first light source 101 and the intensity-modulated light from the second light source 105, combined by the multiplexer 109b to the living body test unit 110. Exit toward.

  When the calibration composite light is emitted toward the living body test unit 110, the light source switch 145 connects the drive circuit 108 and the third light source 106. The lens 144 causes the intensity-modulated light from the third light source 106 to enter the multiplexer 109a. The multiplexer 109a causes the intensity-modulated light from the third light source 106 to enter the multiplexer 109b. The lens 139 causes the intensity-modulated light from the first light source 101 to enter the multiplexer 109b. The multiplexer 109 b multiplexes the intensity modulated light from the first light source 101 and the intensity modulated light from the third light source 106. The irradiation head 112 then applies the combined light for calibration, which combines the intensity-modulated light from the first light source 101 and the intensity-modulated light from the third light source 106 that are combined by the multiplexer 109b, to the living body test unit 110. Exit toward.

  The ultrasonic detector 113 detects normalization sound waves, measurement sound waves, and calibration sound waves as sound wave detection means. The normalization sound wave is a sound wave generated from the blood existing in the living body test part 110 by the intensity-modulated light of the light from the first light source 101. The measurement sound wave is a sound wave generated from blood by the measurement combined light emitted from the irradiation head 112. Here, since the synthetic light for measurement exhibits the same absorption of the liquid in blood, the acoustic wave for measurement reflects the difference in absorption with glucose as the target component. The calibration sound wave is a sound wave generated from blood by the calibration synthesized light emitted from the irradiation head 112. Here, since the calibration synthetic light has the same absorption exhibited by glucose, which is the target component in blood, the calibration acoustic wave reflects the difference in absorption of the liquid in blood. The ultrasonic detector 113 detects normalization sound waves, measurement sound waves, and calibration sound waves via an acoustic coupler 142 as an acoustic matching means in order to prevent reflection of sound waves on the surface of the ultrasonic detector. It is preferable.

  By providing the third light source 106, it is possible to measure sound waves that are not affected by the concentration of the target component such as glucose by irradiating the living body test part 110 with the calibration composite light. By observing changes in the intensity of sound waves generated by the calibration synthetic light, it is possible to monitor whether or not there has been a sudden change in the absorbance of a liquid such as water in the solution.

  The phase detection amplifier 114 outputs an electrical signal proportional to each photoacoustic signal of the normalization sound wave, the measurement sound wave, and the calibration sound wave to the output terminal 115. The output terminal 115 inputs an electrical signal proportional to the photoacoustic signals of the normalization sound wave, the measurement sound wave, and the calibration sound wave to the computer 129. The computer 129 performs various processes on the electrical signals proportional to the input photoacoustic signals of the normalization sound wave, the measurement sound wave, and the calibration sound wave. For example, the computer 129 includes a first light source 101, a second light source 105, a third light source 106, an oscillator 103, driving circuits 104 and 108, a 180 ° phase shift circuit 107, an irradiation head 112, and an ultrasonic detector 113. A control circuit (not shown) for controlling is provided. In the present embodiment, an example in which the computer 129 includes a calibration sound wave determination unit 151, a temperature calculation unit 152, an output unit 153, a concentration calculation unit 154, and a concentration correction unit 155 will be described.

  The calibration sound wave determination unit 151 determines whether the ratio between the intensity of the calibration sound wave detected by the ultrasonic detector 113 and the intensity of the normalization sound wave is within a predetermined range. The ratio between the intensity of the calibration sound wave and the intensity of the normalization sound wave is, for example, a value obtained by dividing the intensity of the calibration sound wave by the intensity of the normalization sound wave. The absorbance exhibited by the aqueous glucose solution is greatly affected by the temperature of the water. By detecting the difference in absorbance exhibited by the aqueous glucose solution excluding the influence of the glucose concentration, it is possible to confirm that the temperature change of the water in the aqueous glucose solution is within an allowable range. For this reason, it is preferable that the predetermined range is a range in which the temperature change of water is recognized as being within the allowable range.

  When the calibration sound wave determination unit 151 determines that the ratio between the intensity of the calibration sound wave detected by the ultrasonic detector 113 and the intensity of the normalization sound wave is within a predetermined range, the concentration calculation unit 154 It is preferable to calculate the concentration of glucose. In this case, the density correction unit 155 can output the density calculated by the density calculation unit 154 to the output unit 153 without correcting it.

  If the calibration sound wave determination unit 151 determines that the ratio between the intensity of the calibration sound wave detected by the ultrasonic detector 113 and the intensity of the normalization sound wave is outside the predetermined range, there is a temperature change. It is preferable to output a message to the output means 153. At this time, it is preferable that the concentration calculation means 154 does not calculate the glucose concentration.

  When the output unit 153 determines that the calibration sound wave determination unit 151 is out of the range, the output unit 153 outputs that the temperature has changed. The output unit 153 may display that the glucose concentration is measured again. The output means 153 can notify the subject that the blood component concentration cannot be measured normally. On the other hand, if the calibration sound wave determination means 151 determines that it is within the range, it may output that there is no temperature change. The output unit 153 may display the density calculated by the density calculation unit 154.

  The temperature calculation unit 152 calculates the temperature based on the intensity of the calibration sound wave detected by the ultrasonic detector 113. By detecting the difference in absorbance exhibited by the aqueous glucose solution excluding the influence of the glucose concentration, the temperature change of the water in the aqueous glucose solution can be calculated. The concentration calculation means 154 calculates the glucose concentration based on the measurement sound wave detected by the ultrasonic detector 113. The concentration correction unit 155 corrects the glucose concentration calculated by the concentration calculation unit 154 based on the temperature calculated by the temperature calculation unit 152. Hereinafter, the calculation method of the temperature calculation unit 152, the concentration calculation unit 154, and the concentration correction unit 155 will be specifically described.

ここで、α_ｉ ^（ｂ）は、第ｉ（ｉ＝１〜３）番目の波長λ_ｉにおける液体の吸光度である。例えば、α_１ ^（ｂ）は、第１の光源１０１の出力する波長λ_１における水の吸光度である。α_３ ^（ｂ）は、第３の光源１０６の出力する波長λ_３における水の吸光度である。ｄα_ｉ ^（ｂ）／ｄＴは、第ｉ（ｉ＝１〜３）番目の波長λ_ｉにおける液体の吸光度の温度に対する変化率である。例えば、α_１ ^（ｂ）／ｄＴは、第１の光源１０１の出力する波長λ_１における水の吸光度の温度に対する変化率である。α_２ ^（ｂ）／ｄＴは、第２の光源１０５の出力する波長λ_２における水の吸光度の温度に対する変化率である。α_３ ^（ｂ）／ｄＴは、第３の光源１０６の出力する波長λ_３における水の吸光度の温度に対する変化率である。また、Ｓ_１は、規格化用音波の強度である。Ｄ_１３は、校正用音波の強度である。Ｄ_１３／Ｓ_１｜_０は校正用音波Ｄ_１３と規格化用音波Ｓ_１の比の、初期値を表す。 FIG. 14 is a graph showing an example of a change in the absorbance of water. The change of the light absorbency for every wavelength when the light absorbency of water changes temperature in 5 degreeC increments of 25 degreeC or more and 55 degreeC is shown. It can be seen that the absorbance at each wavelength shifts to the short wavelength side as the temperature of the water rises. At this time, the temperature change ΔT of water is expressed by the following equation.

Here, α _i ^(b) is the absorbance of the liquid at the _i- th (i = 1 to 3) wavelength λ _i . For example, α ₁ ^(b) is the absorbance of water at the wavelength λ ₁ output from the first light source 101. α ₃ ^(b) is the absorbance of water at the wavelength λ ₃ output from the third light source 106. dα _i ^(b) / dT is the rate of change of the absorbance of the liquid at the _i- th (i = 1 to 3) wavelength λ _{i with} respect to temperature. For example, α ₁ ^(b) / dT is the rate of change of the absorbance of water at the wavelength λ ₁ output from the first light source 101 with respect to temperature. α ₂ ^(b) / dT is the rate of change of the absorbance of water at the wavelength λ ₂ output from the second light source 105 with respect to temperature. α ₃ ^(b) / dT is the rate of change of the absorbance of water at the wavelength λ ₃ output from the third light source 106 with respect to temperature. Further, S ₁ is the intensity of the sound wave for normalization. D ₁₃ is the intensity of the calibration sound wave. D ₁₃ / S ₁ | ₀ represents the initial value of the ratio between the calibration sound wave D ₁₃ and the normalization sound wave S ₁ .

If the target component is determined, the wavelength λ ₃ of the third light source is determined. If the rate of change of the absorbance of the liquid, eg, water, at the wavelength λ ₃ of the third light source and the absorbance of the liquid, eg, water, at the wavelength λ _{3 with} respect to the temperature is acquired, the intensity change of the calibration sound wave detected by the ultrasonic detector can be obtained. By measuring, the changed temperature of the blood, that is, the solution can be measured.

ここで、α_ｉ ^（０）は、第ｉ（ｉ＝１〜３）番目の波長λ_ｉにおける対象成分の単位量あたりの吸光度である。例えば、α_１ ^（０）は、第１の光源１０１の出力する波長λ_１におけるグルコースの単位量あたりの吸光度である。α_２ ^（０）は、第２の光源１０５の出力する波長λ_２におけるグルコースの単位量あたりの吸光度である。また、Ｄ_１２は、測定用音波の強度である。 In addition, the concentration M of the target component, for example, glucose, is expressed by the following equation using the temperature change ΔT of water.

Here, α _i ⁽⁰⁾ is the absorbance per unit amount of the target component at the _i- th (i = 1 to 3) wavelength λ _i . For example, α ₁ ⁽⁰⁾ is the absorbance per unit amount of glucose at the wavelength λ ₁ output from the first light source 101. α ₂ ⁽⁰⁾ is the absorbance per unit amount of glucose at the wavelength λ ₂ output from the second light source 105. _{Further, D 12} is the intensity of the measuring sound wave.

Therefore, the concentration M of the target component, for example, glucose is represented by the following formula.

  From the intensity change of the calibration sound wave detected by the ultrasonic detector and the measurement sound wave detected by the ultrasonic detector, the concentration of the target component corrected for the temperature change of the solution, that is, blood can be measured.

  The liquid component concentration measuring apparatus and the liquid component concentration measuring apparatus control method according to the present embodiment can be applied to the field of measuring the component concentration in a liquid, for example, sugar content measurement of fruits.

It is explanatory drawing which showed the structure of the blood component concentration measuring apparatus which concerns on 1 embodiment. It is explanatory drawing of the sound source distribution in a biological body. It is explanatory drawing of the shape function which concerns on the sound source distribution in a biological body. It is explanatory drawing which showed the photoacoustic signal of the blood component concentration measuring apparatus which concerns on 1 embodiment. It is explanatory drawing which showed the photoacoustic signal of the blood component concentration measuring apparatus which concerns on 1 embodiment. It is explanatory drawing which showed the photoacoustic signal of the blood component concentration measuring apparatus which concerns on 1 embodiment. It is explanatory drawing which showed the light absorption characteristic of water and glucose, and the light wavelength to be used. It is explanatory drawing which showed the light absorption characteristic of water and glucose. It is explanatory drawing which showed the structural example of the blood component concentration measuring apparatus which concerns on 1 embodiment. It is explanatory drawing which showed the structural example of the blood component concentration measuring apparatus which concerns on 1 embodiment. It is explanatory drawing which showed the light absorption characteristic of water. It is explanatory drawing which showed the light absorption characteristic of cholesterol. It is explanatory drawing which showed the structural example of the blood component concentration measuring apparatus which concerns on 1 embodiment. It is explanatory drawing of the calculation principle of the blood component density | concentration which concerns on 1 embodiment. It is explanatory drawing of the sensitivity characteristic of an ultrasonic detector. It is explanatory drawing of the calculation principle of the blood component density | concentration which concerns on 1 embodiment. It is explanatory drawing which showed the structure of the blood component concentration measuring apparatus which concerns on 1 embodiment. It is the figure which showed the Example of the blood component concentration measuring apparatus which concerns on 1 embodiment. It is the figure which showed the Example of the blood component concentration measuring apparatus which concerns on 1 embodiment. It is a schematic block diagram of the component concentration measuring apparatus which concerns on 2 embodiment. It is an example of the output wavelength of a 1st light source, a 2nd light source, and a 3rd light source, (a) shows the light absorbency and extinction coefficient of water and glucose aqueous solution, (b) shows the light absorbency of glucose, and the light absorbency of water. Comparison of absorbance difference and absorbance ratio are shown. It is explanatory drawing which showed the structural example of the conventional blood component concentration measuring apparatus. It is explanatory drawing which showed the structural example of the conventional blood component concentration measuring apparatus.

Explanation of symbols

DESCRIPTION OF SYMBOLS 101 1st light source 103 Oscillator 104 Drive circuit 105 2nd light source 106 3rd light source 107 180 degree phase shift circuit 108 Drive circuit 109, 109a, 109b Multiplexer 110 Living body test part 112 Irradiation head 113 Ultrasonic detector 114 Phase detection amplifier 115 Output terminal 129 Computer 138 Contact thermometer 139 Lens 140 Lens 141, 141a, 141b Calibration sample 142 Acoustic coupler (acoustic matching means)
143 Thermometer 144 Lens 145 Light source switching 150 Component concentration measurement device 151 Calibration sound wave determination means 152 Temperature calculation means 153 Output means 154 Concentration calculation means 155 Concentration correction means 201 Irradiation light 202 Sound source 203 Observation point 204 Model A
205 Model B
206 Model C
211 Light of the first light source (λ1) 212 Light of the second light source (λ2) 501 First semiconductor laser light source 502 Lens 503 Oscillator 504 Drive current source 505 Second semiconductor laser light source 506 Lens 507 180 ° phase shift circuit 508 Drive current source 509 multiplexer 510 living body test part 511 calibration sample 512 acoustic coupler 513 ultrasonic detector 514 phase detection amplifier 515 output terminal 601 first light source 604 drive power source 605 second light source 608 drive power source 609 Multiplexer 610 Living body test part 613 Ultrasonic detector 616 Pulse light source 617 Chopper plate 618 Motor 619 Acoustic sensor 620 Waveform observer 621 Frequency analyzer 801 First semiconductor laser light source 802 Lens 803 Oscillator 804 Drive current source 805 Second Semiconductor laser light source 806 Ren 807 180 ° phase shift circuit 808 drive current source 809 multiplexer 810 biological test unit 811 calibration sample 812 acoustic coupler 813 ultrasonic detector 814 phase sensitive amplifier 815 output terminal

Claims (12)

Two measuring light generating means for generating and outputting light having different wavelengths with the same absorption of the liquid in a solution in which the target component is mixed with the liquid;
Calibration light generating means for generating and outputting light having a wavelength equal to that of one of the measurement light generating means, the absorption exhibited by the target component;
The light from one of the measurement light generation means is intensity-modulated at a predetermined frequency and output, and the light from one of the measurement light generation means, the other of the measurement light generation means, and the Each of the light from the calibration light generating means, the light modulating means for outputting the intensity modulated in the opposite phase with each other at the constant frequency; and
The intensity modulated light of the light from one of the measurement light generating means, the intensity modulated light of the light from one of the measurement light generating means, and the intensity modulated light of the light from the other of the measurement light generating means are synthesized. Synthetic light for measurement, and light for intensity modulation from one of the light generation means for measurement and light for intensity calibration from the light generation means for calibration are emitted toward the solution. Light emitting means for
Normalized sound wave generated from the solution by intensity-modulated light from one of the measurement light generation means, measurement sound wave generated from the solution by the measurement synthetic light emitted from the light emission means, and A sound wave detecting means for detecting a calibration sound wave generated from the solution by the calibration synthetic light emitted by the light emitting means;
A concentration calculating means for calculating the concentration of the target component based on the measurement sound wave detected by the sound wave detecting means;
Equipped with a,
A component concentration measuring apparatus for detecting a temperature change of the liquid based on the normalization sound wave and the calibration sound wave detected by the sound wave detecting means . It is determined whether the ratio between the intensity of the calibration sound wave detected by the sound wave detection means and the intensity of the normalization sound wave is within a predetermined range, and the temperature change of the liquid in the solution is determined. The component concentration measuring apparatus according to claim 1, further comprising a calibration sound wave determination unit that determines whether or not the allowable range is satisfied. When the calibration sound wave determination means determines that it is out of the range, it further comprises an output means for outputting a change in temperature ,
When the output means outputs that the temperature has changed, the concentration calculation means stops calculating the concentration of the target component
Constituent concentration measuring apparatus according to claim 2, wherein the this. The apparatus further comprises temperature calculation means for calculating a temperature change of the liquid based on a change amount of a ratio between the intensity of the calibration sound wave detected by the sound wave detection means and the intensity of the normalization sound wave. The component concentration measuring apparatus according to claim 1. Concentration correction means for correcting the concentration of the target component calculated by the concentration calculation means based on the temperature change calculated by the temperature calculation means and the rate of change of the absorbance of the solution and the target component with respect to temperature. The component concentration measuring apparatus according to claim 4, characterized in that: The liquid is water;
The target component is glucose or cholesterol,
The component concentration measuring device according to claim 1, wherein the solution is blood. Two measuring light generating means for generating and outputting light having different wavelengths with the same absorption of the liquid in a solution in which the target component is mixed with the liquid, a calibration light generating means, a light modulating means, A control method for a component concentration measuring apparatus comprising a control circuit for controlling light emitting means, sound wave detecting means, and concentration calculating means,
The control circuit causes one of the measurement light generation means to output light, and causes the light modulation means to output the light from one of the measurement light generation means after intensity modulation at a predetermined constant frequency. The light emitting means emits intensity-modulated light from one of the measurement light generation means toward the solution, and the sound wave detection means emits light from one of the measurement light generation means. A normalization sound wave detection procedure for detecting a normalization sound wave generated from the solution by intensity-modulated light;
Before, after or simultaneously with the normalization sound wave detection procedure, the control circuit causes the two measurement light generation means to output light, and causes the light modulation means to emit light from one of the measurement light generation means. And the light from the other of the measurement light generating means are modulated in intensity with the constant frequency in opposite phases to each other and output to the light emitting means, and the intensity modulated light of the light from one of the measurement light generating means and The measurement combined light combined with the intensity-modulated light from the other of the measurement light generation means is emitted toward the solution, and the sound wave detection means is caused by the measurement combined light emitted from the light emission means. A measurement sound wave detection procedure for detecting a measurement sound wave generated from the solution;
Before, after or simultaneously with the measurement sound wave detection procedure, the control circuit generates light having the same wavelength as the one of the measurement light generation means in the calibration light generation means. And output the light from one of the measurement light generating means by intensity modulation at a predetermined constant frequency and output the light from one of the measurement light generating means and the calibration light. Light from the light generating means is intensity-modulated in phase opposite to each other at the constant frequency and output, and the light emitting means directs intensity-modulated light from one of the measurement light generating means to the solution. And radiating the combined light for calibration obtained by synthesizing the intensity-modulated light of the light from one of the measurement light generation means and the intensity-modulated light of the light from the calibration light generation means toward the solution, The measuring light generating means is connected to the sound wave detecting means. Sonic for normalization generated from the solution by the intensity-modulated light of the light from one of, and, by calibration synthetic light emission of the light emitting means to detect a calibration waves generated from the solution, the acoustic wave detection means A temperature change detection procedure for detecting a temperature change of the liquid based on the normalized sound wave and the calibration sound wave detected by
After the measurement sound wave detection procedure, the control circuit causes the concentration calculation unit to calculate the concentration of the target component based on the measurement sound wave detected by the sound wave detection unit;
A method for controlling a component concentration measuring apparatus, comprising: The component concentration measuring device further includes a sonic wave determination means for calibration,
The control circuit further determines whether or not the calibration sound wave determination means has a ratio between the intensity of the calibration sound wave detected by the sound wave detection means and the intensity of the normalization sound wave within a predetermined range. 8. A calibration sound wave determination procedure for determining whether or not a temperature change of the liquid in the solution is within an allowable range is provided after the temperature change detection procedure. 2. The component concentration measuring device control method according to 1. The component concentration measuring device further comprises output means,
When it is determined that the calibration sound wave determination means is out of the range, the control circuit further outputs an output procedure to the effect that the output means has a temperature change after the calibration sound wave determination procedure. Yes, and
9. The component concentration measurement according to claim 8, wherein, in the concentration calculation procedure, when the output unit outputs that the temperature has changed, the concentration calculation unit stops calculating the concentration of the target component. Device control method. The component concentration measuring device further includes a temperature calculating means,
The control circuit further causes the temperature calculation means to change the temperature of the liquid based on the amount of change in the ratio between the intensity of the calibration sound wave detected by the sound wave detection means and the intensity of the normalization sound wave. The component concentration measurement device control method according to claim 7, further comprising: a temperature calculation procedure to be calculated after the temperature change detection procedure. The component concentration measurement apparatus further includes a concentration correction unit,
The control circuit further causes the concentration correction unit to calculate the target component calculated by the concentration calculation unit based on a temperature change calculated by the temperature calculation unit and a change rate of the absorbance of the solution and the target component with respect to temperature. The component concentration measurement apparatus control method according to claim 10, further comprising: a concentration correction procedure for correcting the concentration of the component concentration after the concentration calculation procedure and the temperature change detection procedure . The liquid is water;
The target component is glucose or cholesterol,
The component concentration measuring device control method according to claim 7, wherein the solution is blood.

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