JP2009136323A - Component concentration measuring apparatus - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To stabilize the intensity of the light emitted from a light emitting means even if the light emitting means vibrates because the intensity of sonic waves generated in a measured object depends also on the intensity of the light to be applied to the measured object though the concentration of a target component is measured based on the intensity of sonic waves generated in the measured object, and the concentration of the target component can not be accurately measured if the intensity of the light applied to the measured object varies. <P>SOLUTION: The component concentration measuring apparatus 235 related to this invention is characterized by an unnecessary mode eliminating means 237 for eliminating an unnecessary mode of an optical fiber 236 for guiding light to an irradiation head 112. By eliminating the unnecessary mode of the optical fiber 236, the intensity of the light emitted from the irradiation head 112 can be stabilized even if the bending state of the optical fiber 236 varies. <P>COPYRIGHT: (C)2009,JPO&INPIT

Description

  The present invention relates to a non-invasive component concentration measuring apparatus, and more particularly to a non-invasive measuring apparatus for measuring the concentration, that is, a blood glucose level, of glucose as a blood component in a non-invasive manner.

  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. 12 is a diagram illustrating 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. This waveform observer 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 tube surface of the waveform observer 620, and the signals are integrated and averaged for measurement. be able to.

  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. 12, a light pulse having a sub-microsecond pulse width is repeatedly generated at a maximum of 1 kHz, and the 1024 light pulses are averaged to obtain an electric signal proportional to the sound pressure. Is not obtained.

  Therefore, an invention using a light source that is continuously intensity-modulated for the purpose of improving accuracy is disclosed (for example, see Patent Document 1). 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 addition, for the purpose of improving accuracy, an invention for correcting the temperature dependence of the absorbance of a solvent using a contact thermometer is disclosed (for example, see Patent Document 2).

  Therefore, a second conventional example using a light source that is continuously intensity-modulated has been disclosed for the purpose of improving accuracy. FIG. 13 shows the configuration of a second conventional apparatus (see, for example, Patent Document 1). 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 2 will be described with reference to FIG. In FIG. 13, 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.

  When irradiating the object to be measured with irradiation light for measuring the component concentration, the optical fiber is used to guide the surface of the object to be measured to bring the light emitting means into close contact with the surface of the object to be measured. When the surface of the object to be measured is soft, it is difficult to keep the position of the light emitting means constant. If the position of the light emitting means varies, the optical fiber directly connected to the light emitting means vibrates.

It is known that the optical wavelength and light intensity of an optical fiber become unstable due to the bending state of the optical fiber. When the optical fiber vibrates, the intensity of the intensity-modulated light emitted from the light emitting means varies due to the vibration of the optical fiber. When the surface of the object to be measured is soft, the intensity of the intensity-modulated light emitted from the light emitting means may vary.
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.

  In the photoacoustic method, the concentration of the target component is measured based on the intensity of sound waves generated from the object to be measured. The intensity of the sound wave generated by the object to be measured also depends on the intensity of light applied to the object to be measured. If the intensity of light applied to the object to be measured varies, the concentration of the target component cannot be measured accurately.

  Therefore, an object of the present invention is to stabilize the light intensity emitted from the light emitting means even when the light emitting means vibrates.

  The component concentration measuring apparatus according to the present invention is characterized by removing an unnecessary mode of the optical fiber that guides light to the light emitting means. The unnecessary mode of the optical fiber varies the intensity of light emitted from the optical fiber depending on the bending state of the optical fiber. By removing the unnecessary mode of the optical fiber, the intensity of light emitted from the light emitting means can be stabilized even when the bending state of the optical fiber changes.

  The component concentration measuring apparatus according to the present invention comprises two measuring light generating means for generating and outputting light of different wavelengths with the same absorption of the liquid in a solution in which the target component is mixed with the liquid, and the measurement 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 and the light from the other of the measurement light generation means are output. A light modulation means for outputting an intensity modulated in phase opposite to each other at the constant frequency; and a light intensity modulated light from one of the measurement light generation means; and from one of the measurement light generation means A light combining means for measuring the light intensity modulated light and the light intensity modulating light from the other of the light generating means for measurement to output the combined light for measurement; and the intensity modulation output from the light combining means for measurement Light and the synthetic light for measurement, The solution by means of a light emitting means that emits toward the liquid, a normalization sound wave generated from the solution by the intensity-modulated light emitted from the light emitting means, and the synthetic light for measurement emitted from the light emitting means A sound wave detection means for detecting a measurement sound wave generated from the optical fiber, an optical fiber for guiding the intensity-modulated light and the measurement composite light output from the measurement light combining means to the light emitting means, and a middle part of the optical fiber. And an unnecessary mode removing means for removing an unnecessary mode of the optical fiber.

  The unnecessary mode removing unit suppresses light fluctuation caused by the vibration of the light emitting unit. Even when the light emitting means vibrates, the intensity of the light emitted from the light emitting means is stabilized. Even when the light emitting means vibrates, a stable concentration of the target component can be measured.

In the component concentration measuring apparatus according to the present invention, it is preferable that the unnecessary mode removing unit includes one or more struts that bend the optical fiber with a predetermined radius of curvature.
The predetermined radius of curvature is a radius of curvature that eliminates an unnecessary mode of the optical fiber. By bending the optical fiber with the curvature radius of the support column, unnecessary modes of the optical fiber can be removed.

In the component concentration measuring apparatus according to the present invention, the support column may be a cylinder or a cylinder having the radius of curvature as the radius, and the unnecessary mode removing unit may be one in which the optical fiber is wound around the support column. preferable.
The optical fiber can be bent with a radius of curvature that eliminates unwanted modes. Since there is one support column, the unnecessary mode can be removed with a simple configuration.

In the component concentration measuring apparatus according to the present invention, the support column is two cylinders or cylinders having the radius of curvature as the radius, and the unnecessary mode removing unit is configured such that the optical fiber is wound around the support in an 8-shape. It is preferable that it is attached.
The optical fiber can be bent with a radius of curvature that eliminates unwanted modes. The optical fiber can be wound without twisting.

In the component concentration measuring apparatus according to the present invention, it is preferable that the optical fiber is a single mode optical fiber, and the radius of curvature is a radius of curvature for removing a cladding mode of the single mode optical fiber.
By removing the cladding mode of the single mode optical fiber, the unnecessary mode can be removed.

In the component concentration measuring apparatus according to the present invention, it is preferable that the optical fiber is a multimode optical fiber, and the radius of curvature is a radius of curvature that removes a cladding mode and a higher-order mode of the multimode optical fiber.
By removing the cladding mode and the higher order mode of the multimode optical fiber, the unnecessary mode can be removed. Since the optical fiber is a multimode optical fiber, unnecessary modes can be removed at low cost. Further, since the optical fiber is a multimode optical fiber, the beam diameter can be increased.

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, the above-described component concentration measuring apparatus can be applied to a blood component concentration measuring apparatus. Thereby, even if it is a case where a biological test part vibrates, the intensity | strength of the light from a light-projection means can be stabilized.

  According to the present invention, an unnecessary mode of an optical fiber that guides light to a light emitting means can be removed. By removing the unnecessary mode of the optical fiber, the intensity of light from the light emitting means can be stabilized even when the object to be measured vibrates. In addition, since it is not necessary to protect the optical fiber excessively against vibration and bending, the configuration of the component concentration measuring device can be simplified.

  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, the component concentration measuring device will be described as a blood component concentration measuring device. If the living body test portion 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, the component concentration measuring device can be implemented. 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 such a measuring apparatus can be applied to various objects.

(First embodiment)
The blood component concentration measuring apparatus according to the present embodiment is configured such that the measurement light generating means for generating two light beams having different wavelengths and the two light beams having different wavelengths are electrically transmitted by signals having the same frequency and opposite phases. Light modulating means for intensity-modulating the light, light emitting means for combining the intensity-modulated two light beams of different wavelengths into one light beam, and emitting the light toward the living body, and generated in the living body by the emitted light A blood component concentration measuring device comprising: a sound wave detecting means for detecting a sound wave; 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 measurement light generating means sets the wavelength of one wave of light to a wavelength at which the blood component exhibits characteristic absorption, and the other one wave of light. Can be set to a wavelength at which water exhibits absorption equal to that of the one-wave light.

  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, the first light source 101 as a part of the measuring light generating means is intensity-modulated by a drive circuit 104 as a part of the light modulating means in synchronization with the oscillator 103 as a part of the light modulating means. Has been.

  On the other hand, the second light source 105 as a part of the measuring light generating means is intensity-modulated in synchronism with the oscillator 103 by a driving 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 multiplexer 109 serving as a light emitting unit, and is subjected to a test as a light beam. It irradiates to living body test part 110 as. 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.

  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 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 measuring apparatus according to the present embodiment, the measurement light generating means combines the two light beams of different wavelengths, which have been intensity-modulated, into one light beam and irradiates water with the pressure of the sound wave generated. It is possible to adjust the relative intensity of each of the two light beams having different wavelengths so that becomes zero.

  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 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 according to this embodiment, the measurement light generating means and the light modulating means are means 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. 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.

  Next, the details of the technology that is the basis of the blood component concentration measuring apparatus 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 wavelengths λ ₁ and 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.

  A sound wave propagating from the sound source 202 and propagating through the living body (for the sake of simplicity, the sound wave is uniform) is on the extension line of the irradiated light and is at the observation point 203 at a distance r from the sound source at its sound pressure p ( r) is observed.

  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 irradiation light absorption, the heat source Q is proportional to the absorption coefficient α, and its distribution is a space of irradiation light (including scattered light if it is generated) in the medium. Equal to the distribution. 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 a case where the scattering is remarkably large and as soon as the irradiation light enters, the light is scattered in all directions.

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, 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 a 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 around 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 . 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 at all, the beam diameter 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) can be affected. If this is exp (−αz), it is expected that the model B205 and the model C206 without scattering result in no influence of scattering on the shape function.

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 equality of the shape function is justified, and it can be seen that the apparatus 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 this 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. 4, 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. 4 shows a case where rectangular wave light is irradiated, 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 conditions are 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 of the photoacoustic signal.

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. 7 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 range 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.

(Second Embodiment)
FIG. 9 is a schematic configuration diagram of a component concentration measuring apparatus according to the present embodiment. The component concentration measuring apparatus 235 according to this embodiment includes an optical fiber 236 that guides the light output from the multiplexer 109 to the irradiation head 112, and an unnecessary mode removing unit 237 that removes an unnecessary mode of the optical fiber 236. It is characterized by. This will be specifically described below.

  The component concentration measuring apparatus 235 shown in FIG. 9 has a configuration for detecting a photoacoustic signal generated in the living body test unit 110. For example, the component concentration measuring device 235 includes a first light source 101 and a second light source 105 as measurement light generation means, a first light source 101 as a light modulation means, a second light source 105, a drive circuit 104, A drive circuit 108, a 180 ° phase shift circuit 107 and an oscillator 103, a multiplexer 109 as a measuring light combining means, an irradiation head 112 as a light emitting means, and an ultrasonic detector 113 as a sound wave detecting means. Prepare. Further, the component concentration measuring device 235 may include lenses 139 and 140, an acoustic coupler 142, a phase detection amplifier 114, and an output terminal 115. About the structure and operation | movement for detecting the photoacoustic signal which generate | occur | produces in the above-mentioned living body test part 110, it is the same as that of 1st Embodiment.

  The component concentration measuring apparatus 235 further includes an optical fiber 236 and an unnecessary mode removing unit 237 that stabilize the intensity of light from the irradiation head 112 even when the living body test part 110 vibrates. Hereinafter, configurations and operations of the optical fiber 236 and the unnecessary mode removing unit 237 will be described. In the present embodiment, one of the measurement light generation means is described as the first light source 101 and the other of the measurement light generation means is described as the second light source 105. One of the measurement light generation means may be the second light source 105 and the other of the measurement light generation means may be the first light source 101.

  The optical fiber 236 guides the intensity-modulated light and the measurement combined light output from the multiplexer 109 to the irradiation head 112. By providing the optical fiber 236, the living body test part 110 can be irradiated with intensity-modulated light and measurement combined light without causing the first light source 101 and the second light source 105 to vibrate. The optical fiber 236 may be a single mode optical fiber or a multimode optical fiber. Further, a plastic optical fiber having high flexibility may be used.

  When the first light source 101 and the second light source 105 are guided to the irradiation head by an optical fiber and combined by the irradiation head 112, the optical fiber 236 is irradiated with the first light source 101 and the second light source 105. An optical fiber provided between the head 112 and the head 112 is obtained.

  The unnecessary mode removing unit 237 is provided in the middle of the optical fiber 236 and removes the unnecessary mode of the optical fiber 236. The unnecessary mode removing unit 237 removes unnecessary mode light from the light output from the multiplexer 109 while passing the light output from the multiplexer 109. The intensity-modulated light and the measurement combined light output from the multiplexer 109 enter the steady mode after passing through the unnecessary mode removing means 237.

  The unnecessary mode removing unit 237 includes, for example, one or more struts that bend the optical fiber 236 with a predetermined radius of curvature. The predetermined radius of curvature is a radius of curvature that eliminates unnecessary modes of the optical fiber 236. By bending the optical fiber 236 with a predetermined radius of curvature, unnecessary modes of the optical fiber can be removed.

  The predetermined radius of curvature varies depending on the properties of the optical fiber 236 and the propagating light. For example, when the optical fiber 236 is a single mode optical fiber, the predetermined radius of curvature is a radius of curvature that removes the cladding mode of the single mode optical fiber. When the optical fiber 236 is a multimode optical fiber, the predetermined radius of curvature is a radius of curvature that removes the cladding mode and higher-order modes of the multimode optical fiber. The radius of curvature for removing the unnecessary mode of the optical fiber 236 can be obtained by detecting the unnecessary mode when the optical fiber 236 is bent.

  FIG. 10 is a schematic diagram showing a first example of unnecessary mode removing means. The first example of the unnecessary mode removing means includes a single column 238a that bends the optical fiber 236 with a predetermined radius of curvature R. The support 238a is a column or a cylinder. In the first example of the unnecessary mode removing unit, the optical fiber 236 is wound around the support 238a. Since the radius R of the support column 238a is a radius of curvature for removing the unnecessary mode of the optical fiber 236, the unnecessary mode of the optical fiber 236 can be removed. In addition, since the number of the support columns 238a is one, the configuration can be simplified.

  FIG. 11 is a schematic diagram illustrating a second example of the unnecessary mode removing unit. The second example of the unnecessary mode removing unit includes two columns 238b and 238c that bend the optical fiber 236 with a predetermined radius of curvature R. The support columns 238b and 238c are columns or cylinders. In the second example of the unnecessary mode removing means, the optical fiber 236 is wound around the support columns 238b and 238c in the shape of figure 8. Since the radius R of the support columns 238b and 238c is a radius of curvature for removing the unnecessary mode of the optical fiber 236, the unnecessary mode of the optical fiber 236 can be removed. Further, the optical fiber 236 can be wound without being twisted.

  The detection of the normalization sound wave will be described. In this case, the second light source 105 does not emit intensity modulated light. The first light source 101 is driven by the drive circuit 104 and outputs intensity-modulated light. The lens 139 causes the intensity-modulated light from the first light source 101 to enter the multiplexer 109. The multiplexer 109 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. The ultrasonic detector 113 detects normalization sound waves generated in the blood existing in the living body test part 110.

  The detection of sound waves for measurement will be described. The first light source 101 is driven by the drive circuit 104 and outputs intensity-modulated light. The lens 139 causes the intensity-modulated light from the first light source 101 to enter the multiplexer 109. The second light source 105 is driven by the drive circuit 108 and outputs intensity modulated light. The lens 140 causes the intensity-modulated light from the second light source 105 to enter the multiplexer 109. The multiplexer 109 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 combined by the multiplexer 109 and the intensity-modulated light from the second light source 105, to the living body test unit 110. Exit toward. The ultrasonic detector 113 detects sound waves for measurement generated in blood existing in the living body test part 110.

  Here, in the detection of the normalization sound wave and the measurement sound wave, the unnecessary mode removing unit 237 removes the unnecessary mode. For this reason, even if the living body test unit 110 moves, the living body test unit 110 vibrates, or the surface of the living body test unit 110 vibrates, the intensity modulation emitted from the irradiation head 112 is modulated. The intensity of the light and the synthetic light for measurement can be stabilized.

  The phase detection amplifier 114 outputs an electric signal proportional to the photoacoustic signal of the normalization sound wave detected by the ultrasonic detector 113 to the output terminal 115. By analyzing the electrical signal proportional to the photoacoustic signal of the normalization sound wave and the measurement sound wave, the concentration of glucose as the target component can be measured. Since the intensity of the intensity-modulated light and measurement synthetic light emitted from the irradiation head 112 are stable, the concentration of glucose as the target component can be stabilized.

  The component concentration measuring apparatus and the 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 a schematic block diagram of the component concentration measuring apparatus which concerns on 2nd Embodiment. It is a schematic diagram which shows the 1st example of an unnecessary mode removal means. It is a schematic diagram which shows the 2nd example of an unnecessary mode removal means. 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 107 180 degree phase shift circuit 108 Drive circuit 109 Multiplexer 110 Living body test part 112 Irradiation head 113 Ultrasonic detector 114 Phase detection amplifier 115 Output terminal 139 Lens 140 Lens 142 Acoustic coupler 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) 235 Component concentration measuring device 236 Optical fiber 237 Unnecessary mode removing means 238a, 238b, 238c Post 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

Claims (7)

Two measuring light generating means for generating and outputting light of different wavelengths with the same absorption of the liquid in a solution in which the target component is mixed with the liquid;
The light from one of the measurement light generation means is intensity-modulated at a predetermined constant frequency and output, and the light from one of the measurement light generation means and the other of the measurement light generation means A light modulation means for intensity-modulating and outputting light in opposite phases to each other at the constant frequency;
Outputs intensity-modulated light of light from one of the measurement light generation means, and intensity of light from one of the measurement light generation means and intensity of light from the other of the measurement light generation means A measuring light combining means for combining the modulated light and outputting a measuring combined light;
A light emitting means for emitting the intensity-modulated light output from the measuring light combining means and the measuring combined light toward the solution;
Sound wave detection for detecting a normalization sound wave generated from the solution by the intensity-modulated light emitted from the light emitting unit and a measurement sound wave generated from the solution by the synthetic light for measurement emitted from the light output unit Means,
An optical fiber that guides the intensity-modulated light output from the measurement light combining means and the measurement combined light to the light emitting means;
Unnecessary mode removing means provided in the middle of the optical fiber to remove unnecessary modes of the optical fiber;
A component concentration measuring device. The component concentration measuring apparatus according to claim 1, wherein the unnecessary mode removing unit includes one or more struts that bend the optical fiber with a predetermined radius of curvature. The support column is a cylinder or a cylinder having the radius of curvature as a radius,
The component concentration measuring apparatus according to claim 2, wherein the unnecessary mode removing means is one in which the optical fiber is wound around the support column. The support column is two columns or cylinders having the radius of curvature as a radius,
The component concentration measuring apparatus according to claim 2, wherein the unnecessary mode removing unit is configured such that the optical fiber is wound in a figure 8 shape around the support column. The optical fiber is a single mode optical fiber;
The component concentration measuring apparatus according to any one of claims 1 to 4, wherein the radius of curvature is a radius of curvature for removing a cladding mode of a single mode optical fiber. The optical fiber is a multimode optical fiber;
5. The component concentration measuring apparatus according to claim 1, wherein the radius of curvature is a radius of curvature for removing a cladding mode and a higher-order mode of a multimode optical fiber. The liquid is water;
The target component is glucose or cholesterol,
The component concentration measuring device according to any one of claims 1 to 6, wherein the solution is blood.

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