JP5411180B2 - Component concentration measuring method and apparatus - Google Patents

Description

  The present invention relates to a method and an apparatus for measuring a concentration of a component by a photoacoustic method, which can be applied to the measurement of the concentration of glucose present in blood plasma or the concentration of other components present in plasma.

In order to prevent diabetes, it is necessary to accurately measure the glucose concentration present in plasma. There is a photoacoustic method as a method for continuously monitoring the blood glucose level of a diabetic patient, which is summarized as follows.
(A) The photoacoustic method provides continuous blood glucose monitoring.
(B) It is painless for a diabetic patient, does not require a blood sample, and does not cause discomfort to the diabetic patient.
(C) Compared with other optical technologies, there is no deterioration in efficiency due to scattering media.
(D) High sensitivity characteristics can be obtained by combining optics and acoustics.

  There are two types of photoacoustic methods: a pulse method and a continuous-wave (hereinafter referred to as CW) method. The pulse method has a drawback that high optical power must be used to obtain high sensitivity. On the other hand, the CW method has a drawback that the signal intensity changes when the characteristic at the reflection surface changes, that is, there is no reproducibility. However, since high optical power may cause a problem in terms of safety for the human body, it is preferable to adopt the CW method (see Patent Document 1, Patent Document 2, and Patent Document 3). In the pulse method and the CW method, the component concentration is quantified by utilizing the fact that the amplitude of the acoustic wave is proportional to the component concentration.

JP 2008-125542 A JP 2008-125543 A JP 2008-145262 A

  In the conventional pulse method and CW method, during the measurement of glucose concentration in plasma several times, it is highly possible that other plasma parameters other than the glucose concentration (for example, body temperature, concentration of other components, etc.) will also change. There was a problem that glucose selectivity was poor and it was difficult to obtain an accurate glucose concentration.

  The present invention has been made to solve the above problems, and an object of the present invention is to provide a component concentration measuring method and apparatus capable of measuring a component concentration such as blood glucose concentration with high accuracy.

  The component concentration measuring method according to the present invention includes a first light irradiation step of irradiating an object to be measured whose concentration of a component to be measured is known with intensity-modulated light, and the object to be measured by the first light irradiation step. A first photoacoustic signal detecting step for detecting the photoacoustic signal generated from the output signal and outputting an electric signal; and a modulation frequency at which the amplitude of the electric signal obtained in the first photoacoustic signal detecting step is maximized. A first frequency measuring step for measuring as a first frequency, a first phase measuring step for measuring a phase of an electric signal when the amplitude is maximum as a reference phase, and two light beams having different wavelengths from each other. A second light irradiation step for irradiating the object to be measured with signals having different frequencies and at a frequency of 1 and irradiating the object to be measured, and gradually changing the optical power of at least one of the two intensity-modulated lights. And a second photoacoustic signal detection step for detecting a photoacoustic signal generated from the object to be measured by the second light irradiation step and outputting an electrical signal, and a second photoacoustic signal detection step. A second phase measuring step for searching for a first inflection point at which the phase of the obtained electrical signal is 0; and a second phase measuring step for measuring a difference in optical power between two intensity-modulated lights at the first inflection point. 1 light power measurement step, a third light irradiation step for irradiating the object to be measured with intensity-modulated light after an arbitrary period of time, and light generated from the object to be measured by the third light irradiation step A third photoacoustic signal detection step for detecting an acoustic signal and outputting an electrical signal; and a second modulation frequency at which the phase of the electrical signal obtained in the third photoacoustic signal detection step is the reference phase. Search as frequency And measuring the intensity of the two light beams having different wavelengths and irradiating the object to be measured with signals of the second frequency and different phases, respectively, and irradiating at least one of the two intensity-modulated lights. A fourth light irradiation step for gradually changing the optical power of the intensity-modulated light, and a fourth photoacoustic for detecting a photoacoustic signal generated from the object to be measured and outputting an electric signal by the fourth light irradiation step. A signal detection step, a third phase measurement step for searching for a second inflection point where the phase of the electrical signal obtained in the fourth photoacoustic signal detection step is 0, and the second inflection point. A second optical power measurement step for measuring the difference between the optical powers of the two intensity-modulated lights in FIG. 2, and the difference between the optical powers measured in the second optical power measurement step and the first optical power measurement step. light And a concentration deriving step for deriving the concentration of the component to be measured after the arbitrary time elapses from the amount of change from the difference in power.

In one configuration example of the component concentration measurement method of the present invention, when the first optical power measurement step changes only the optical power of one of the two intensity-modulated lights, the intensity of one of the two intensity-modulated lights is changed. For the modulated light, a difference between the initial optical power before the change and the optical power at the first inflection point is measured, and the second optical power measurement step includes one of the two intensity-modulated lights. In the case of changing only the optical power, the difference between the initial optical power before the change and the optical power at the second inflection point is measured for the one intensity-modulated light.
Further, in one configuration example of the component concentration measurement method of the present invention, the concentration derivation step includes the change amount of the optical power, the light absorption coefficient of the measurement target in the initial state, and the light absorption of the measurement target after the arbitrary time elapses. A change amount of the component concentration of the measurement target is obtained from the coefficient change amount, and a component concentration of the measurement target after the arbitrary time elapses is derived from the change amount of the component concentration and the known component concentration. It is what.
In addition, in one configuration example of the component concentration measuring method of the present invention, a light absorption spectrum measurement step of further measuring a light absorption spectrum of the object to be measured and obtaining the light absorption coefficient and the amount of change in the light absorption coefficient from the measured spectrum. It is characterized by providing.
Further, in one configuration example of the component concentration measuring method of the present invention, the first and second optical power measuring steps include the difference in driving voltage between the two light irradiating means for emitting the two intensity-modulated lights. The concentration derivation step refers to calibration data indicating the relationship between the drive voltage difference and the amount of change in optical power, and the drive voltage difference measured in the second optical power measurement step and the first The amount of change in optical power is obtained from the amount of change from the drive voltage difference measured in step 1 of optical power measurement.

  The component concentration measuring apparatus of the present invention comprises a light irradiating means for irradiating light to the object to be measured, a light power controlling means for controlling light power, and a photoacoustic signal generated from the object to be measured by this light irradiation. A photoacoustic signal detecting means for detecting and outputting an electric signal; a frequency measuring means for measuring the frequency of the electric signal; a phase measuring means for measuring the phase of the electric signal; and the optical power of the two intensity-modulated lights. An optical power measuring means for measuring the difference; and a concentration deriving means for deriving the concentration of the component to be measured from the amount of change in the optical power after an arbitrary time has elapsed, the light irradiation means at the first time Irradiate intensity-modulated light to an object to be measured whose concentration of a component to be measured is known, and intensity-modulate two light beams having different wavelengths at a second time with signals having a first frequency and different phases. The covered Irradiate a fixed object, irradiate the measured object with intensity-modulated light at a third time, and emit light of two waves having different wavelengths at a second time with signals having a second frequency and different phases. Intensity modulation is performed to irradiate the object to be measured, and the optical power control means gradually changes the optical power of at least one of the two intensity modulation lights at the second and fourth times, The frequency measuring means measures, as the first frequency, a modulation frequency at which the amplitude of the electric signal obtained at the first time becomes maximum, and the phase of the electric signal obtained at the third time is referred to A phase modulation frequency is searched as the second frequency, and the phase measuring means measures the phase of the electrical signal when the amplitude of the electrical signal obtained at the first time is maximum as the reference phase. The first inflection point at which the phase of the electrical signal obtained at the second time becomes 0 is searched, and the second inflection at which the phase of the electrical signal obtained at the fourth time becomes 0 The optical power measuring means searches for a point, measures the optical power difference between the two intensity-modulated lights at the first inflection point at the second time, and the second power at the fourth time. The difference between the optical powers of the two intensity-modulated lights at the inflection point is measured, and the concentration deriving means calculates the difference between the optical power measured at the fourth time and the optical power measured at the second time. The component concentration of the measurement target at the fourth time is derived from the amount of change from the difference.

  According to the present invention, by appropriately selecting the wavelengths of the two intensity-modulated lights, it is possible to cause a difference between different substances in the gradient of the relationship between the concentration and the optical power. Therefore, in the present invention, by appropriately selecting the light wavelength so that the sensor response to a specific measurement target is maximized, the selectivity of the measurement target can be improved, and in a measurement object including a plurality of different substances. It becomes possible to measure the concentration of the component to be measured with high accuracy.

It is a figure which shows the absorption spectrum of glucose aqueous solution and albumin aqueous solution in the light wavelength in the range of a near-infrared spectrum. It is a block diagram which shows the structure of the conventional component density | concentration measuring apparatus. It is a figure which shows the relationship between the blood glucose level obtained by the component density | concentration measuring apparatus of FIG. 2, and the maximum amplitude of a photoacoustic signal. It is a block diagram which shows the structure of the component concentration measuring apparatus which concerns on embodiment of this invention. It is a block diagram which shows the structure of the information processing apparatus of the component concentration measuring apparatus which concerns on embodiment of this invention. It is a figure explaining a mode that an acoustic wave is generated by two light of a different wavelength. It is a figure which shows the amplitude and phase of a photoacoustic signal when optical power is changed when the blood glucose concentration of a to-be-measured object is a reference blood glucose concentration. It is a figure which shows the amplitude and phase of a photoacoustic signal when optical power is changed when the blood glucose concentration and albumin concentration of a to-be-measured object change. It is a figure which shows the result of having measured glucose concentration and albumin concentration with the component concentration measuring apparatus which concerns on embodiment of this invention. It is a flowchart which shows operation | movement of the component concentration measuring apparatus which concerns on embodiment of this invention. It is the figure which shows the measurement result of the component concentration measuring apparatus which concerns on embodiment of this invention, and the enlarged view of a measurement result. It is a figure which shows an optical power balance-phase characteristic.

[Principle of the Invention]
In the present invention, in order to accurately measure the blood glucose concentration, an optical component that is a new component concentration measuring method that improves the glucose selectivity according to the light wavelength by utilizing the principle that the amplitude of the photoacoustic signal depends on the light absorption coefficient. A Power Balance Shift method (hereinafter abbreviated as OPBS method) is proposed.

The measurement method of the present invention searches for a phase inflection point at a location where the amplitude of the photoacoustic signal is minimal, while increasing or decreasing the power of two light beams having different optical wavelengths and a phase difference of π, and from the result, blood It is a method of measuring the concentration of molecules dissolved in the inside. This method can be expanded to detect not only glucose components in plasma but also other plasma components (such as albumin and cholesterol).
The following theoretical formula is used to explain the concept of the detection method. The photoacoustic signal intensity S can be expressed as:

Here, K is a constant, β is a coefficient of thermal expansion of the object to be measured, v is a sound velocity, n is an experimental system parameter depending on the setup, C _p is a specific heat of the object to be measured, α is a light absorption coefficient of the object to be measured, P is the optical power.
Further, when two differential signal settings are used, the photoacoustic signal intensity S can be expressed by the following equation.

In Equation (2), P ₁ and P ₂ are optical powers, and α ₁ and α ₂ are optical absorption coefficients of the object to be measured for light having optical powers P ₁ and P ₂ , respectively.
The problem is that parameters such as constant K, thermal expansion coefficient β, sound velocity v, and specific heat C _p depend on temperature or the concentration of the mixture, so that the photoacoustic signal intensity S cannot be used directly for the calculation of blood glucose concentration. It is. In order to suppress such dependency, in the measurement method disclosed in Patent Document 1, normalization is performed with a signal of one wavelength.

On the other hand, in the measurement method of the present invention, the optical power P ₁ that minimizes the photoacoustic signal intensity S is searched for while changing the power (for example, P ₁ ) of one of the two optical beams. Theoretically, the minimum value of the photoacoustic signal intensity S is 0. However, since there is noise experimentally, it does not become 0. The photoacoustic signal intensity S at this time is approximately 100 times smaller than the photoacoustic signal intensity when a one-wavelength light beam is used. For the sake of simple explanation, here, the noise is ignored and the photoacoustic signal intensity S is set to zero. When S = 0, a new theoretical formula can be written as
α ₁ P ₁ −α ₂ P ₂ = 0 (3)

When the concentration of the component to be measured is changed, for example, the blood glucose concentration is changed by C _{g, and the} light absorption coefficients α ₁ and α ₂ are changed by δα ₁ and δα ₂ due to the change in concentration, respectively, Equation (3) is It changes from the established state to the state of Equation (4).
(Α ₁ + δα ₁ C _g ) P ₁ − (α ₂ + δα ₂ C _g ) P ₂ ≠ 0 (4)

When the power (for example, P ₁ ) of _one light beam is changed to return to the state of S = 0, the following equation is established.
(Α ₁ + δα ₁ C _g ) (P ₁ + δP ₁ ) − (α ₂ + δα ₂ C _g ) P ₂ = 0 (5)
The following equation is obtained from equation (5).

In equations (5) and (6), δP ₁ is the amount of change in the optical power P ₁ . From equation (6), it can be seen that in the present invention, the blood glucose concentration can be measured from the optical power variation δP ₁ , the known optical absorption coefficients α ₁ and α _2, and the optical absorption coefficient variations δα ₁ and δα _2. .

[Embodiment]
Hereinafter, embodiments of the present invention will be described with reference to the drawings. FIG. 1 is a diagram showing absorption spectra of an aqueous glucose solution and an aqueous albumin solution at a light wavelength within the near-infrared spectrum. In FIG. 1, 100 indicates an absorption spectrum of an aqueous glucose solution, and 101 indicates an absorption spectrum of an aqueous albumin solution. The horizontal axis in FIG. 1 is the light wavelength, and the vertical axis is the relative absorption coefficient.

  According to FIG. 1, although there is a portion where the absorption spectra of the aqueous glucose solution and the aqueous albumin solution overlap, the value of the optical wavelength at which the maximum value of the absorption coefficient is obtained and the value of the optical wavelength at which the minimum value of the absorption coefficient is obtained are glucose. And albumin are different. That is, it can be seen that glucose and albumin have different light wavelength dependencies of absorption spectra. Therefore, it is possible to improve the selectivity of one specific compound by carefully selecting the light wavelength in consideration of such light wavelength dependency.

  FIG. 2 is a block diagram showing a configuration of a conventional component concentration measuring apparatus. This component concentration measuring apparatus fixes a laser diode 200 that emits laser light, a laser driver 201 that drives the laser diode 200, an optical fiber 202 that guides laser light emitted from the laser diode 200, and an optical fiber 202. An optical fiber holder 203 that irradiates a human body or a measured object 210 that is a part of the human body with laser light, and a photoacoustic signal generated from the measured object 210 by the photoacoustic effect are detected, and an electric power proportional to the sound pressure is detected. An acoustic sensor 204 that converts the signal into the signal; an amplifier 205 that amplifies the electrical signal output from the acoustic sensor 204; a function generator 206 that generates a reference signal; an output signal from the amplifier 205 and an output from the function generator 206; The reference signal is used as an input to detect a measurement signal of a desired frequency from the output signal of the amplifier 205. The in-amplifier 207, the voltage-current converter 208 that supplies a drive current to the laser driver 201, the function generator 206, and the lock-in amplifier 207 are controlled, and the measurement signal detected by the lock-in amplifier 207 is processed to obtain blood glucose. And an information processing device 209 including a computer for deriving the density.

  FIG. 3 shows the results of examining the relationship between the blood glucose concentration and the maximum amplitude of the photoacoustic signal for each of the two light wavelengths of 1382 nm and 1610 nm using the component concentration measuring apparatus shown in FIG. In FIG. 3, 300 indicates the first measurement result when the light wavelength is 1382 nm, and 301 indicates the first measurement result when the light wavelength is 1610 nm. Reference numeral 302 denotes a second measurement result when the light wavelength is 1382 nm, and 303 denotes a second measurement result when the light wavelength is 1610 nm. The horizontal axis in FIG. 3 is the blood glucose concentration, and the vertical axis is the maximum amplitude of the photoacoustic signal.

  The maximum amplitude of the photoacoustic signal shows a similar trend with a linear relationship inversely proportional to blood glucose concentration. However, even if the similar component concentration measurement is executed twice, the signal variation is large, the noise is large in the low concentration region, and the accuracy is lowered. Therefore, it can be seen that it is difficult to obtain an accurate glucose concentration in the conventional component concentration measuring device, particularly in the low concentration range.

  In this embodiment, a new component concentration measurement method based on amplitude information of a photoacoustic signal is provided. FIG. 4 is a block diagram showing the configuration of the component concentration measuring apparatus according to the present embodiment. The component concentration measuring apparatus includes a laser diode 1-1 serving as a first light irradiation unit that emits laser light, a laser diode 1-2 serving as a second light irradiation unit that emits laser light, and a laser diode 1- The laser driver 2 that drives 1, 1-2, the optical fibers 3-1 and 3-2 that guide the laser light emitted from the laser diodes 1-1 and 1-2, and the laser diodes 1-1 and 1-2. An optical coupler 4 for combining the laser light emitted from the optical fiber 5, an optical fiber 5 for guiding the laser light combined by the optical coupler 4, a photoacoustic cell 6 which is a case for housing the object to be measured 13, and a laser light An optical window 7 made of glass, and an acoustic sensor 8 serving as a photoacoustic signal detecting means for detecting a photoacoustic signal generated from the object 13 to be measured by the photoacoustic effect and converting it into an electric signal proportional to sound pressure; , Sound sensor The amplifier 9 for amplifying the electric signal output from the output 8, the function generator 10 for generating the reference signal, the output signal of the amplifier 9 and the reference signal output from the function generator 10 as inputs, and the output of the amplifier 9 The lock-in amplifier 11 that detects a measurement signal of a desired frequency from the signal, the function generator 10 and the lock-in amplifier 11 are controlled, and the measurement signal detected by the lock-in amplifier 11 is processed to derive a blood glucose concentration. It is comprised from the information processing apparatus 12 which consists of computers.

Examples of the laser diodes 1-1 and 1-2 include, for example, a distributed feedback semiconductor laser (DFB-LD). An example of the acoustic sensor 8 is a microphone.
FIG. 5 is a block diagram illustrating a configuration of the information processing apparatus 12. The information processing apparatus 12 includes a function generator control unit 120 that controls the function generator 10, a frequency measurement unit 121 that measures the frequency of the measurement signal, a phase measurement unit 122 that measures the phase of the measurement signal, and optical power. An optical power control unit 123 for controlling, an information recording unit 124 for recording frequency and phase information of a measurement signal, an optical power measurement unit 125 for measuring a difference in optical power between two intensity-modulated lights, and a blood glucose concentration. It has a glucose concentration deriving unit 126 for deriving and a storage unit 127 for storing information.

  In the present embodiment, an advantage as defined in the prior art (see Patent Document 1) using light of two wavelengths is utilized. Two rectangular laser beams with different wavelengths are intensity-modulated by signals of the same frequency and opposite phase, respectively, to generate a rectangular wave beam, and the rectangular wave beam is combined and irradiated on a measurement object (for example, plasma). Then, the two laser beams are absorbed by the measurement object with their respective light absorption coefficients. As a result, an acoustic wave is generated by the photoacoustic effect (an effect in which the optical energy absorbed by the object to be measured is changed into thermal energy and an acoustic wave is generated by volume expansion due to the thermal energy) (FIG. 6). In FIG. 6, 60 indicates a photoacoustic signal by a rectangular wave beam emitted from the laser diode 1-1, and 61 indicates a photoacoustic signal by a rectangular wave beam emitted by the laser diode 1-2. The acoustic wave is proportional to a difference in intensity between two signals αP (α is a light absorption coefficient of the object to be measured and P is an optical power) by each of the two laser beams.

  In the present embodiment, first, the level (signal amplitude) of the reference photoacoustic signal is determined based on the known reference blood glucose concentration. When the blood glucose concentration changes from the reference blood glucose concentration, the amplitude of the photoacoustic signal from the two lights varies with the light wavelength and the light absorption coefficient. At this time, the light wavelength is determined in advance based on the absorption spectrum shown in FIG. The power of the two lights is changed to balance the light absorption effect due to the change in blood glucose concentration, and the amplitude of the photoacoustic signal is returned to the level of the reference photoacoustic signal determined at the reference blood glucose concentration.

When one light is irradiated on the object to be measured, the intensity S (signal amplitude) of the generated photoacoustic signal can be expressed as in the above equation (1). In addition, when two light beams having different wavelengths are intensity-modulated with signals having the same frequency and opposite phases, and the object to be measured is irradiated, the intensity S of the generated photoacoustic signal is expressed by the above equation (2). Can be represented. However, since the parameters such as the constant K, the thermal expansion coefficient β, the sound velocity v, and the specific heat C _p depend on the temperature or the concentration of the mixture as described above, the photoacoustic signal intensity S is directly used for calculating the blood glucose concentration. It is not possible. In order to suppress such dependency, in the measurement method disclosed in Patent Document 1, standardization is performed using a signal of one wavelength.

In the measurement method of the present embodiment, the optical power P ₁ or P ₂ that minimizes the photoacoustic signal intensity S is searched for by (α ₁ P ₁ −α ₂ P ₂ ) in Expression (2). The minimum value of the photoacoustic signal intensity S is determined within the noise range. In practice, in order to change the optical power output, the range of adjustment of the laser voltage in FIG. 4 is considered, and the two wavelengths (λ ₁ , λ ₂ ) and the two light absorption coefficients (α ₁ , α ₂ ) are set. On the other hand, the optical power is determined as (P ₁ , P ₂ ).

FIG. 7 is a diagram showing the amplitude and phase of the photoacoustic signal when the optical powers P ₁ and P ₂ are changed when the blood glucose concentration of the object to be measured is the reference blood glucose concentration of 0 g / dL. The horizontal axis in FIG. 7 is the optical power balance, and the vertical axis is the amplitude and phase of the photoacoustic signal. In FIG. 7, 70 indicates the amplitude of the photoacoustic signal, and 71 indicates the phase of the photoacoustic signal. In the optical power balance, the relationship between the powers of two lights is expressed by a laser voltage, and a point where αP in each of the two laser lights is equal is zero. In the optical power balance is smaller than zero region becomes lower in the optical power P ₂ than the light power P _1, whereas, in a large area from the optical power balance is zero, the optical power than the optical power P ₁ P ₂ Is higher.

  As is clear from FIG. 7, at the point where the optical power balance is 0, that is, at the point where αP is equal in each of the two laser beams, the amplitude of the photoacoustic signal shows a minimum value, and the phase of the photoacoustic signal is inflected. Indicates a point. Theoretically, the amplitude of the photoacoustic signal at the point where the optical power balance is zero should be zero, but it is experimentally a minimum value including noise and does not become zero.

  Changing the power of one of the two lights changes the amplitude and phase of the photoacoustic signal. The amplitude of the photoacoustic signal increases on both sides of the point where the optical power balance is zero. That is, when the power of one light is reduced, the difference in intensity between the two signals αP is increased, and the amplitude of the photoacoustic signal is increased. Also, when the power of one light is reduced, the phase of the photoacoustic signal approaches the phase of the photoacoustic signal when excited by the other light alone. That is, it approaches a state excited by only one light.

In summary, when one of the two light powers P ₂ is reduced or the light power P ₁ is raised to make the light power balance smaller than 0, the amplitude of the photoacoustic signal increases. The phase differs by -90 degrees between the two phases. That is, the phase of the photoacoustic signal at this time is the same as the phase of the photoacoustic signal (that is, −90 degrees) when excited by the light alone having the optical power P ₁ . Further, when the optical power P ₂ is increased or the optical power P ₁ is decreased to make the optical power balance larger than 0, the amplitude of the photoacoustic signal increases, and the phase of the photoacoustic signal is equal to that of the optical power P ₂ . The phase is the same as the phase of the photoacoustic signal when excited by light alone (ie, +90 degrees).

  Here, there are two important characteristics in that the optical power balance is zero. First, since the photoacoustic signals of two wavelengths are phase shifted by ± 90 degrees, the point where the optical power balance is 0 coincides with the phase 0 point which is the inflection point. Furthermore, on both sides of the inflection point, the phase of the photoacoustic signal has a clear positive or negative value, and the relationship between the phase and the optical power is linear, so that the phase can be measured quickly and the measurement is simple. Yes, it is possible to measure accurately, and the position of the inflection point 0 can be obtained with good accuracy from the positive and negative values of the phase. From these two features, blood glucose concentration can be measured with high accuracy by phase measurement.

  Here we focus on the amplitude signal, not the phase, to provide an easy-to-understand explanation. The phase point at which the amplitude of the photoacoustic signal is the minimum and becomes the inflection point of the phase of the photoacoustic signal is located at a point where αP in each of the two lights is equal. When the blood glucose concentration of the object to be measured is the reference blood glucose concentration of 0 g / dL, the above equation (3) is established at this phase point.

Experimentally, since noise exists, it is difficult to make the amplitude of the photoacoustic signal zero. However, the minimum value of the amplitude of the photoacoustic signal is a value two orders of magnitude lower than the amplitude of the photoacoustic signal when light of one wavelength is used. Then, conceptually, the amplitude of the photoacoustic signal can be reduced to a level at which the noise level can be completely ignored. Next, the component concentration of the object to be measured is changed. For example, the blood glucose concentration is changed from 0 to C _g [g / dL].

FIG. 8A shows the amplitude and phase of the photoacoustic signal when the optical powers P ₁ and P ₂ are changed when the blood glucose concentration of the object to be measured changes by C _g [g / dL]. FIG. 8B is a diagram showing the amplitude and phase of the photoacoustic signal when the optical powers P ₁ and P ₂ are changed when the albumin concentration of the object to be measured changes. 8A and 8B, the horizontal axis represents the optical power balance, and the vertical axis represents the amplitude and phase of the photoacoustic signal.

  In FIG. 8A, 80 indicates the amplitude of the photoacoustic signal when the blood glucose concentration changes, 81 indicates the phase of the photoacoustic signal when the blood glucose concentration changes, and 82 indicates the blood glucose concentration changes. The amplitude of the previous photoacoustic signal is shown, and 83 shows the phase of the photoacoustic signal before the blood glucose concentration changes. In FIG. 8B, 84 indicates the amplitude of the photoacoustic signal when the albumin concentration changes, 85 indicates the phase of the photoacoustic signal when the albumin concentration changes, and 86 indicates before the albumin concentration changes. , 87 indicates the phase of the photoacoustic signal before the albumin concentration changes.

When the blood glucose concentration or albumin concentration changes, the amplitude and phase of the photoacoustic signal change according to the change in αP. For example, when the blood glucose concentration changes by C _{g and the} light absorption coefficients α ₁ and α ₂ change by δα ₁ and δα ₂ due to the concentration change, the above equation (4) is established. There are two methods for detecting the component concentration (blood glucose concentration or albumin concentration). Either a method of searching for an optical power balance at which the amplitude of the photoacoustic signal is minimum, or a method of searching for an optical power balance at the inflection point of the phase of the photoacoustic signal.

Around the point where the optical power balance is 0, the relationship between the phase of the photoacoustic signal and the optical power is close to linear, so that the position of the phase 0 can be accurately evaluated. The optical power difference for moving the phase of the photoacoustic signal to a new inflection point after the component concentration change can be obtained by measurement. 8A and 8B, a new inflection point (amplitude 80 of the photoacoustic signal after the change in the component concentration) is obtained by increasing the optical power P ₂ or decreasing the optical power P _1. , 84 can be shifted to the minimum).

In the measurement method of the present embodiment, the optical power is changed in order to search for a point where the phase of the photoacoustic signal is zero. More specifically, in order to change the optical power, the drive voltages of the laser diodes 1-1 and 1-2 are changed. When the phase of the photoacoustic signal is 0, the above equations (5) and (6) are established. As apparent from the equation (6), the blood glucose concentration C _g can be obtained from a new light power balance shift value δP ₁ unique to glucose and relative light absorption coefficients δα ₁ and δα ₂ . When the component concentration to be measured is the albumin concentration, it can be obtained in the same manner. The light absorption coefficients α ₁ and α ₂ and the light absorption coefficient change amounts δα ₁ and δα ₂ can be obtained from light absorption spectrum measurement.

  FIG. 9 shows the results of measuring the glucose concentration and the albumin concentration by the measurement method of the present embodiment. Here, two lights having wavelengths of 1382 nm and 1610 nm are used. In FIG. 9, 90 indicates the characteristic of glucose concentration, and 91 indicates the characteristic of albumin concentration. The horizontal axis in FIG. 9 is the density, and the vertical axis is the shift amount of the optical power balance. As can be seen from FIG. 9, in the measurement method of the present embodiment, there is a difference in the slope of characteristics between the two components glucose and albumin. Furthermore, if the wavelengths of the two lights are appropriately selected, the difference in characteristics between the two components can be further increased. It is possible to use two or more different wavelengths of light for further optimization.

  The measurement method of this embodiment is an efficient method for non-invasively measuring the components of a solution based on photoacoustic measurement. This measurement method can optimize a very selective approach to one specific composition by choosing two optical wavelengths. Obviously, many solutes can be detected in different solvents given the various optical wavelengths available. In addition, it is possible to measure any difference in absorption coefficient (density) by adjusting the corresponding optical power and obtaining the power balance.

The choice of optical wavelength is not limited by the absorption coefficient. If α ₁ = 2α ₂ , P ₁ = 0.5P ₂ and equation (3) is still valid. Furthermore, measurement methods based on phase 0 inflection points converge quickly and provide very accurate measurements. In measuring the glucose concentration, the inventors selected the near-infrared NIR optical region as the wavelength range and measured at wavelengths of 1382 nm and 1610 nm. In order to measure the phase of the photoacoustic signal, it is necessary to store several measurement points. If the noise is completely ignored, it is possible to determine the linear slope from the 2-point measurement data, whereas the parabola (second-order polynomial) requires 3-point measurement data. From this point of view, the linear characteristic of the photoacoustic signal can obtain an inflection point with a phase of 0 earlier.

  However, based on the dependency between noise and the required measurement accuracy, the number of measurement points should be drastically increased. If the change of the photoacoustic signal is linear, the position of phase 0 can be obtained with very accurate accuracy from the measurement data of two points, and the position can be searched within a small range. On the other hand, when the change in the photoacoustic signal is parabolic, the binary search algorithm (binary search) is the best method. However, the time required to obtain the position where the phase is 0 is very long. From an experimental perspective, it is difficult to estimate the quantitative increase in measurement time in relation to the sensor response time. However, if the linear behavior of the phase of the photoacoustic signal is used, it can be measured more quickly and an accurate component concentration value can be provided.

  Next, the operation of the component concentration measuring apparatus according to the present embodiment will be described in more detail. FIG. 10 is a flowchart showing the operation of the component concentration measuring apparatus. In the present embodiment, as shown in FIGS. 3 and 9, an experiment was performed with a liquid sample at a high concentration level in order to clearly distinguish the signal and the noise level.

  First, in order to determine the reference level in the initial state at time t0, only one laser diode 1-1 is operated. The device under test 13 is introduced into the photoacoustic cell 6. When a drive current is supplied from the laser driver 2, the laser diode 1-1 emits laser light. At this time, when a rectangular-wave drive current is supplied from the laser driver 2, the laser diode 1-1 emits intensity-modulated light. The wavelength of light is, for example, 1384 nm. This intensity-modulated light is guided by the optical fiber 3-1, passes through the optical coupler 4, is further guided by the optical fiber 5, and irradiates the object to be measured 13 in the photoacoustic cell 6 through the optical window 7. (FIG. 10, step S1).

  The acoustic sensor 8 detects a photoacoustic signal generated from the device under test 13, and the amplifier 9 amplifies the electrical signal output from the acoustic sensor 8. The lock-in amplifier 11 detects a measurement signal having a frequency determined by the reference signal output from the function generator 10 among the signals included in the output of the amplifier 9.

  The function generator control unit 120 of the information processing device 12 changes the frequency of the drive current supplied from the laser driver 2 to the laser diode 1-1 by changing the frequency of the reference signal generated by the function generator 10. Then, the optical modulation frequency sweep is performed to gradually change the optical modulation frequency and to gradually change the frequency of the measurement signal detected by the lock-in amplifier 11 (the same frequency as the optical modulation frequency) (step S2 in FIG. 10). Thus, the acoustic resonance peak is searched.

Next, when the maximum amplitude of the measurement signal is found, the frequency measurement unit 121 of the information processing apparatus 12 measures the frequency of the measurement signal at the maximum amplitude (reference frequency F0), and the phase measurement unit 122 The phase of the measurement signal at the time of amplitude (reference phase P0) is measured (step S3 in FIG. 10).
The information recording unit 124 of the information processing apparatus 12 stores the reference frequency F0 measured by the frequency measuring unit 121 and the reference phase P0 measured by the phase measuring unit 122 in the storage unit 127 (step S4 in FIG. 10).

Next, a standard blood glucose measurement method is performed on the object to be measured 13 to obtain a reference blood glucose concentration C _g0 [g / dL] (step S5 in FIG. 10). Thus, the reference phase P0 and the reference frequency F0 are obtained with the reference blood glucose concentration C _g0 [g / dL]. To perform a standard blood glucose measurement method, insert a glucose sensor into the body of the blood glucose meter, set the needle on a dedicated machine (or body), collect blood from a finger, etc., and absorb the blood into the glucose sensor. Let A standard blood glucose measurement method uses a known concentration of glucose solution as a standard calibration solution for machine operation confirmation. It is used to check whether the machine is operating normally during initial operation, or to check whether the blood glucose level is abnormal (ie, whether the machine is operating normally).

  Next, the two laser diodes 1-1 and 1-2 are operated, and the two lights are combined to perform measurement. When a drive current is supplied from the laser driver 2, the laser diodes 1-1 and 1-2 emit laser light. At this time, the laser driver 2 supplies the light emitted from the laser diodes 1-1 and 1-2 by supplying a rectangular-wave drive current having the same frequency and an opposite phase to the laser diodes 1-1 and 1-2. Intensity modulation is performed using signals of the same frequency and opposite phase. At this time, the wavelength of the light emitted from the laser diode 1-1 is, for example, 1384 nm, and the wavelength of the light emitted from the laser diode 1-2 is, for example, 1610 nm. The power of the two lights is the same. The intensity-modulated lights emitted from the laser diodes 1-1 and 1-2 are respectively guided by the optical fibers 3-1 and 3-2, combined by the optical coupler 4, and further guided by the optical fiber 5, and the optical window. 7 to irradiate the object 13 to be measured in the photoacoustic cell 6 (step S6 in FIG. 10).

  Subsequently, the function generator control unit 120 of the information processing apparatus 12 supplies the laser diodes 1-1 and 1-2 from the laser driver 2 by changing the frequency of the reference signal generated by the function generator 10. The frequency of the drive current is changed, the optical modulation frequency is set to the reference frequency F0, and the frequency of the measurement signal detected by the lock-in amplifier 11 is set to the reference frequency F0. The optical power control unit 123 of the information processing apparatus 12 gradually changes the power of light emitted from the laser diode 1-1 by changing the magnitude of the drive current supplied from the laser driver 2 to the laser diode 1-1. The changing optical power sweep is performed (step S7 in FIG. 10).

The phase measurement unit 122 of the information processing apparatus 12 searches for an inflection point of the phase of the measurement signal, that is, a point where the phase becomes 0 (step S8 in FIG. 10). When the inflection point of the phase is found, the optical power measurement unit 125 of the information processing apparatus 12 measures the difference between the optical powers of the two lights at the inflection point (step S9 in FIG. 10). The optical power measurement unit 125 measures a reference drive voltage difference V _OPBS0 that is a difference between the drive voltage supplied to the laser diode 1-1 and the drive voltage supplied to the laser diode 1-2 as a difference in optical power.

  Since the two optical powers at the time of step S6 are the same, when only the power of one of the two lights is changed, the initial optical power and the inflection point at the time of step S6 for this one light. The difference (drive voltage difference) from the optical power at may be obtained. Further, in the optical power sweep in step S7, the power of light emitted from the two laser diodes 1-1 and 1-2 may be changed.

  Next, measurement at time t after an arbitrary time has elapsed from time t0 will be described. First, only one laser diode 1-1 is operated to perform measurement using only one light. The intensity-modulated light emitted from the laser diode 1-1 is guided by the optical fiber 3-1, passes through the optical coupler 4, is further guided by the optical fiber 5, passes through the optical window 7, and passes through the optical acoustic cell 6. The object to be measured 13 is irradiated (step S10 in FIG. 10).

The function generator control unit 120 of the information processing device 12 changes the frequency of the drive current supplied from the laser driver 2 to the laser diode 1-1 by changing the frequency of the reference signal generated by the function generator 10. The optical modulation frequency is set to the reference frequency F0. Further, the function generator control unit 120 changes the optical modulation frequency from the reference frequency F0 by changing the frequency of the reference signal generated by the function generator 10.
The phase measurement unit 122 of the information processing device 12 searches for a point where the phase of the measurement signal becomes the reference phase P0, and the frequency measurement unit 121 of the information processing device 12 measures the frequency F1 at this point. Thus, the frequency F1 corresponding to the reference phase P0 is searched (step S11 in FIG. 10). The frequency F1 is located in the vicinity of the reference frequency F0.

  Next, the two laser diodes 1-1 and 1-2 are operated, and the two lights are combined to perform measurement. The laser driver 2 supplies a rectangular-wave drive current having the same frequency and an opposite phase to the laser diodes 1-1 and 1-2, thereby allowing the light emitted from the laser diodes 1-1 and 1-2 to be the same frequency. Intensity modulation is performed using signals of opposite phase. Similarly to the above, the wavelength of light emitted from the laser diode 1-1 is, for example, 1384 nm, and the wavelength of light emitted from the laser diode 1-2 is, for example, 1610 nm. The power of the two lights is the same. The intensity-modulated lights emitted from the laser diodes 1-1 and 1-2 are respectively guided by the optical fibers 3-1 and 3-2, combined by the optical coupler 4, and further guided by the optical fiber 5, and the optical window. 7 to irradiate the object 13 to be measured (step S12 in FIG. 10).

  Subsequently, the function generator control unit 120 of the information processing apparatus 12 supplies the laser diodes 1-1 and 1-2 from the laser driver 2 by changing the frequency of the reference signal generated by the function generator 10. The frequency of the drive current is changed, the optical modulation frequency is set to the frequency F1, and the frequency of the measurement signal detected by the lock-in amplifier 11 is set to the frequency F1. The optical power control unit 123 of the information processing apparatus 12 gradually changes the power of light emitted from the laser diode 1-1 by changing the magnitude of the drive current supplied from the laser driver 2 to the laser diode 1-1. The changing optical power sweep is performed (step S13 in FIG. 10).

The phase measurement unit 122 of the information processing apparatus 12 searches for an inflection point of the phase of the measurement signal, that is, a point where the phase becomes 0 (step S14 in FIG. 10). When the inflection point of the phase is found, the optical power measurement unit 125 of the information processing apparatus 12 measures the difference between the optical powers of the two lights at the inflection point (step S15 in FIG. 10). The optical power measurement unit 125 measures a drive voltage difference V _OPBS1 that is a difference between a drive voltage supplied to the laser diode 1-1 and a drive voltage supplied to the laser diode 1-2 as a difference in optical power.

  Since the two optical powers at the time of step S12 are the same, when only the power of one of the two lights is changed, the initial optical power and the inflection point at the time of step S12 for this one light. The difference (drive voltage difference) from the optical power at may be obtained. Further, in the optical power sweep in step S13, the power of the light emitted from the two laser diodes 1-1 and 1-2 may be changed.

The storage unit 127 of the information processing apparatus 12 stores in advance calibration data indicating a relationship between the difference between the drive voltage difference V _OPBS1 and the reference drive voltage difference V _OPBS0 (V _OPBS1 −V _OPBS0 ) and the optical power change amount _δP . It is remembered. Such calibration data can be obtained by actually measuring in advance. When the blood glucose concentration at time t is C _g (t), (V _OPBS1 -V _OPBS0 ) is the difference between the blood glucose concentration C _g (t) and the reference blood glucose concentration C _g0 (C _g (t) -C _g 0).

The glucose concentration deriving unit 126 of the information processing apparatus 12 refers to the storage unit 127 to obtain the optical power change amount _δP corresponding to the drive voltage difference (V _OPBS1 −V _OPBS0 ) (step S16 in FIG. 10). From the change amount δP, the light absorption coefficients α ₁ and α _2, and the light absorption coefficient change amounts δα ₁ and δα ₂ , the change amount of blood glucose concentration C _g = (C _g (t) −C _g 0). Is calculated (step S17 in FIG. 10).

The light absorption coefficients α ₁ and α ₂ and the light absorption coefficient change amounts δα ₁ and δα ₂ can be obtained from light absorption spectrum measurement. The light absorption spectrum measuring means (not shown) measures the light absorption spectrum of the device under test 13 with respect to the light having the light power P ₁ emitted from the laser diode 1-1 in the initial state at time t0, and from this spectrum, the light absorption coefficient α Calculate ₁ The light absorption spectrum measuring means measures the light absorption spectrum of the device under test 13 with respect to the light having the optical power P ₁ emitted from the laser diode 1-1 at time t, calculates the light absorption coefficient from the spectrum, By obtaining the difference between this light absorption coefficient and the light absorption coefficient α _{1 in the} initial state, the light absorption coefficient change amount Δα ₁ can be calculated. Similarly, the light absorption coefficient α ₂ and the light absorption coefficient change amount δα ₂ can be obtained by measuring the light absorption spectrum of the object 13 with respect to the light having the light power P ₂ emitted from the laser diode 1-2. it can.

Finally, the glucose concentration deriving unit 126 determines the blood at time t from the blood glucose concentration variation C _g = (C _g (t) −C _g 0) calculated in step S17 and the known reference blood glucose concentration C _g0. The glucose concentration C _g (t) is calculated (step S18 in FIG. 10).
This is the end of the processing of the component concentration measuring apparatus of the present embodiment.

  FIG. 11A shows the measurement results of this embodiment at different blood glucose concentrations (0 [g / dL] -18.6 [g / dL]), and FIG. 11B shows the results of FIG. FIG. 11A and 11B, the horizontal axis represents the laser drive voltage, and the vertical axis represents the amplitude and phase of the photoacoustic signal. 11A and 11B, 110 and 111 indicate the amplitude and phase of the photoacoustic signal when the blood glucose concentration is 0 [g / dL], and 112 and 113 are when the blood glucose concentration is 5 [g / dL]. The amplitude and phase of a photoacoustic signal are shown, 114 and 115 show the amplitude and phase of a photoacoustic signal at a blood glucose concentration of 10 [g / dL], and 116 and 117 are photoacoustics at a blood glucose concentration of 15 [g / dL]. The amplitude and phase of the signal are shown. 118 and 119 show the amplitude and phase of the photoacoustic signal at a blood glucose concentration of 18.6 [g / dL].

According to FIGS. 11A and 11B, the phase of the photoacoustic signal shows a linear behavior and the amplitude of the photoacoustic signal shows a parabolic shape, regardless of the noise level. Looking at the amplitude of the photoacoustic signal, it can be seen that the signal level changes greatly, but does not depend on the blood glucose concentration. Considering equation (2), variations such as constant K, coefficient of thermal expansion β, speed of sound v, specific heat C _p depend on temperature, blood glucose concentration, or other parameters. Furthermore, fitting is necessary to find the minimum value of the amplitude of the photoacoustic signal. However, if there is not enough experimental data, an accurate fitting result cannot be obtained. In addition, the measurement accuracy deteriorates when there is noise.

  On the other hand, looking at the phase change of the photoacoustic signal, it has a linear shape, and the position where the phase is 0 can be immediately discriminated by the two measurement data. Therefore, if the phase of the photoacoustic signal is used, a short-time sensor response can be obtained, and the position of phase 0 can be easily found. Furthermore, with respect to noise, since the influence of noise can be made relatively small by increasing not only two measurement points but also phase measurement points, it is important for high-sensitivity measurement. Therefore, after determining what kind of measurement accuracy is necessary, the number of phase data measurements is determined. In the case of a parabolic characteristic such as the amplitude characteristic of a photoacoustic signal, if an attempt is made to measure the amplitude, a complicated and asymmetric analysis using a quadratic polynomial fitting is required, which may increase the measurement time.

  Considering the advantages as described above, the search measurement of phase 0 is performed in the present embodiment. FIG. 9 shows the results of measuring the glucose concentration and the albumin concentration by the measurement method of the present embodiment. In both cases of glucose and albumin, a linear response was obtained with respect to the component concentration. The slope of the characteristic in FIG. 9 is proportional to the light absorption coefficient at the wavelength used. Compared with the results of FIG. 3, it can be seen that the measurement method of the present embodiment has good linearity and can be used for low concentration measurement with high accuracy.

  Furthermore, when comparing the results of glucose and albumin in FIG. 9 with respect to the selectivity of the sensor, the result of slope −25 [V / g / dL] in the glucose solution and slope −11 [V / g / dL] in the albumin solution is obtained. Obtained. This result is the same as that expected from the result of the absorption spectrum shown in FIG. Note that the difference between the slopes of glucose and albumin can be further increased by optimizing the wavelengths of the two lights.

  As described above, in the present embodiment, it is possible to cause a difference in sensor response (gradient) with respect to the concentration between different substances by appropriately selecting the wavelengths of the two lights. That is, by appropriately selecting the light wavelength so that the sensor response to blood glucose is maximized, the glucose selectivity of the component concentration measuring device can be improved, and blood glucose can be measured in a measurement object containing a plurality of different substances. This produces an effect that quantification is possible.

In this embodiment, the intensity of the two lights is modulated with a signal having the same frequency and opposite phase, but the intensity of the light may be modulated with a signal having a phase difference other than 180 °.
The phase of the signal for intensity-modulating the two lights was manually changed to obtain the optical power balance-phase characteristic shown in FIG. The horizontal axis in FIG. 12 is the optical power balance, and the vertical axis is the phase of the photoacoustic signal. Here, light having a wavelength of 1438 nm and light having a wavelength of 1610 nm are used. The optical power balance is expressed by a laser driving voltage of a laser diode that generates light of 1438 nm. The laser drive voltage of the laser diode that generates light of 1610 nm is 1.4V. In FIG. 12, 400 is a characteristic when the phase difference of each signal that modulates the intensity of two lights is 180 °, 401 is a characteristic when the phase difference is 180.1 °, and 402 is a phase difference of 180.3 °. 403 is a characteristic when the phase difference is 180.8 °, 404 is a characteristic when the phase difference is 185.8 °, 405 is a characteristic when the phase difference is 195.8 °, and 406 is a phase difference. Shows the characteristics when the angle is 178.8 °.

  When the phase difference between the signals for intensity-modulating the two lights is 180 °, the phase of the photoacoustic signal remains −90 ° when the drive voltage of one laser diode is changed. Since the signal intensity is low at around 0.51 V, the characteristics are distorted by the influence of the phase noise of the photoacoustic signal. Immediately after that, a 180 ° phase shift suddenly occurs in the photoacoustic signal and becomes discontinuous. As in the case before the phase shift, distortion due to noise occurs, and the phase of the photoacoustic signal is stabilized at 90 °.

  When the phase difference between the signals for intensity-modulating the two lights is around 180 ° (180.1 °, 180.8 °, 180.3 °), the phase of the photoacoustic signal is at the discontinuous point of the phase change described above. It is changing continuously. There is a region where the phase of the photoacoustic signal is stable at ± 90 ° in the voltages before and after the discontinuity. When the phase difference between the signals for intensity-modulating the two lights is not 180 ° (185.8, 195.8), the phase of the photoacoustic signal is ± only in the continuous transition region of the photoacoustic signal phase at all voltages. There is no region that is stable at 90 °. As described above, the measurement was performed by increasing the phase difference of the signal from 180 °. However, when the phase difference of the signal was decreased to 178.8 °, a symmetrical tendency was obtained. That is, the slope of the phase change from the phase 90 ° to the phase −90 ° with respect to the drive voltage is reversed. In conclusion, for accurate and appropriate solution component concentration measurement, the phase difference between the signals for intensity modulation of the two lights can be about 180 ° ± 30 °.

  Practically, it is not appropriate to use 180 ° as the phase difference between the signals for intensity-modulating the two lights because the influence of noise is large in the phase transition region of the photoacoustic signal. Also, since the slope of the phase change of the photoacoustic signal is steep, the bisection method is the only approach for determining the point where the phase of the photoacoustic signal is 0, and it takes time to measure and the convergence is poor. . In order to increase the measurement speed, a linear transition region is most appropriate. That is, this is a method for estimating the position of phase 0 relatively from a plurality (2, 3) of measurement points around the phase 0 point of the photoacoustic signal. In such a method, measurement accuracy can be improved more reliably by repeating measurement and estimation.

  When the phase difference between the signals for intensity-modulating the two lights is close to 180 °, a linear transition region around the phase 0 point of the photoacoustic signal is obtained. Since the slope of the phase change of the photoacoustic signal is relatively steep, the phase changes greatly even in a small component concentration range, and appropriate measurement accuracy can be obtained. On the other hand, as the phase difference between the signals for intensity-modulating the two lights increases from 180 °, the slope of the phase change of the photoacoustic signal becomes gentle. In such a case, even in a large component concentration range, the concentration can be quantified because there is a phase 0 point in the phase transition region. Therefore, the phase difference between the signals for intensity-modulating the two lights is appropriately set according to a desired density range in order to determine measurement accuracy and dynamic range.

  When the phase difference between the signals for intensity-modulating the two lights is greatly different from 180 °, for example, when 180 ° is different from 10 °, 20 ° or 30 °, the linear transition region of the phase of the photoacoustic signal is wide. The measurement range can be used, and measurement of a wide concentration range can be supported. Therefore, for a large density range, the phase difference between the signals for intensity-modulating the two lights may be set to 190 °, 200 ° or 210 °, for example. However, since the gradient of the phase change of the photoacoustic signal becomes gentle and the measurement accuracy decreases, the concentration range and the measurement accuracy are in a trade-off relationship.

  In this embodiment, blood glucose and albumin have been described as examples of components to be measured. However, the present invention is not limited to this, and the present invention can be widely applied to the analysis of components contained in a measurement object. It is.

  The information processing apparatus 12 according to the present embodiment can be realized by, for example, a computer including a CPU, a storage device, and an interface, and a program that controls these hardware resources. A program for operating such a computer is provided in a state of being recorded on a recording medium such as a flexible disk, a CD-ROM, a DVD-ROM, or a memory card. The CPU writes the read program into the storage device, and executes the processing described in this embodiment in accordance with this program.

  The present invention can be applied to a technique for monitoring the concentration of components such as blood glucose.

  1-1, 1-2 ... laser diode, 2 ... laser driver, 3-1, 3-2, 5 ... optical fiber, 4 ... optical coupler, 6 ... photoacoustic cell, 7 ... optical window, 8 ... acoustic sensor, DESCRIPTION OF SYMBOLS 9 ... Amplifier, 10 ... Function generator, 11 ... Lock-in amplifier, 12 ... Information processing apparatus, 13 ... Device under test, 120 ... Function generator control part, 121 ... Frequency measurement part, 122 ... Phase measurement part, 123 ... Optical power control unit 124... Information recording unit 125. Optical power measurement unit 126 126 Glucose concentration deriving unit 127 127 Storage unit.

Claims (10)

A first light irradiation step of irradiating an object to be measured whose concentration of a component to be measured is known with intensity-modulated light;
A first photoacoustic signal detection step for detecting a photoacoustic signal generated from the measurement object by the first light irradiation step and outputting an electric signal;
A first frequency measurement step for measuring, as the first frequency, the modulation frequency at which the amplitude of the electrical signal obtained in the first photoacoustic signal detection step is maximum;
A first phase measuring step for measuring the phase of the electrical signal when the amplitude is maximum as a reference phase;
Two waves of different wavelengths are intensity-modulated by signals of the first frequency and different phases, respectively, and irradiated on the object to be measured, and the optical power of at least one of the two intensity-modulated lights A second light irradiation step for gradually changing
A second photoacoustic signal detection step for detecting a photoacoustic signal generated from the object to be measured by the second light irradiation step and outputting an electric signal;
A second phase measurement step for searching for a first inflection point where the phase of the electrical signal obtained in the second photoacoustic signal detection step is 0;
A first optical power measurement step for measuring a difference in optical power between two intensity-modulated lights at the first inflection point;
A third light irradiation step of irradiating the measured object with intensity-modulated light after an arbitrary period of time;
A third photoacoustic signal detection step for detecting a photoacoustic signal generated from the object to be measured by the third light irradiation step and outputting an electric signal;
A second frequency measurement step for searching for a modulation frequency at which the phase of the electrical signal obtained in the third photoacoustic signal detection step is the reference phase as a second frequency;
Two waves of different wavelengths are intensity-modulated with signals of the second frequency and different phases, respectively, and irradiated on the object to be measured, and the optical power of at least one of the two intensity-modulated lights A fourth light irradiation step for gradually changing
A fourth photoacoustic signal detection step for detecting a photoacoustic signal generated from the object to be measured by the fourth light irradiation step and outputting an electric signal;
A third phase measurement step for searching for a second inflection point where the phase of the electrical signal obtained in the fourth photoacoustic signal detection step is 0;
A second optical power measurement step for measuring a difference in optical power between the two intensity-modulated lights at the second inflection point;
From the amount of change between the difference in optical power measured in the second optical power measurement step and the difference in optical power measured in the first optical power measurement step, the concentration of the component to be measured after the arbitrary time has elapsed is calculated. And a concentration deriving step for deriving the concentration. In the component concentration measuring method according to claim 1,
In the first optical power measurement step, when only the optical power of one of the two intensity-modulated lights is changed, the first optical power before the change of the one intensity-modulated light and the first optical power are changed. Measure the difference from the optical power at the inflection point,
In the second optical power measurement step, when only the optical power of one intensity modulated light of the two intensity modulated lights is changed, the initial optical power before the change of the one intensity modulated light and the second optical power are changed. A method for measuring a concentration of a component, characterized by measuring a difference from an optical power at an inflection point. In the component concentration measuring method according to claim 1 or 2,
The concentration deriving step calculates a change amount of the component concentration of the measurement target from the change amount of the optical power, the light absorption coefficient of the measurement target in the initial state, and the change amount of the light absorption coefficient of the measurement target after the arbitrary time elapses. A component concentration measuring method characterized in that the component concentration to be measured after the arbitrary time elapses is derived from the change amount of the component concentration and the known component concentration. In the component concentration measuring method according to claim 3,
The component concentration measurement method further comprises a light absorption spectrum measurement step of measuring a light absorption spectrum of the object to be measured and obtaining the light absorption coefficient and the change amount of the light absorption coefficient from the measured spectrum. In the component concentration measuring method according to any one of claims 1 to 4,
In the first and second optical power measurement steps, a difference in driving voltage between two light irradiating units that emit two intensity-modulated lights is measured as a difference in optical power,
In the concentration derivation step, the drive voltage difference measured in the second optical power measurement step and the first optical power measurement are referred to with reference to calibration data indicating the relationship between the drive voltage difference and the amount of change in optical power. A component concentration measuring method, characterized in that the amount of change in optical power is obtained from the amount of change from the drive voltage difference measured in steps. A light irradiating means for irradiating the object to be measured;
Optical power control means for controlling optical power;
Photoacoustic signal detection means for detecting a photoacoustic signal generated from the object to be measured by this light irradiation and outputting an electrical signal; and
Frequency measuring means for measuring the frequency of the electrical signal;
Phase measuring means for measuring the phase of the electrical signal;
Optical power measuring means for measuring the difference between the optical powers of the two intensity-modulated lights; and
A concentration deriving means for deriving the concentration of the component to be measured from the amount of change in optical power after an arbitrary period of time;
The light irradiating means irradiates an object to be measured whose concentration of a component to be measured is known at a first time with an intensity-modulated light, and outputs two waves of light having different wavelengths at a second time. Intensity modulation is performed with signals having different frequencies and different phases, and the object to be measured is irradiated, intensity-modulated light is irradiated to the object to be measured at a third time, and two wavelengths having different wavelengths are irradiated at a fourth time. Irradiating the object to be measured with intensity modulation of light of a wave with signals of a second frequency and different phases,
The optical power control means gradually changes the optical power of at least one of the two intensity-modulated lights at the second and fourth times,
The frequency measuring means measures, as the first frequency, a modulation frequency at which the amplitude of the electric signal obtained at the first time becomes maximum, and the phase of the electric signal obtained at the third time is referred to Search for a modulation frequency to be a phase as the second frequency,
The phase measuring means measures the phase of the electrical signal when the amplitude of the electrical signal obtained at the first time is maximum as the reference phase, and the phase of the electrical signal obtained at the second time is Search for the first inflection point that becomes 0, search for the second inflection point where the phase of the electrical signal obtained at the fourth time becomes 0,
The optical power measuring means measures a difference in optical power between two intensity-modulated lights at the first inflection point at the second time, and 2 at the second inflection point at the fourth time. Measure the difference in optical power of the two intensity-modulated lights,
The concentration deriving means calculates the component concentration of the measurement target at the fourth time from the amount of change between the difference in optical power measured at the fourth time and the difference in optical power measured at the second time. A component concentration measuring apparatus characterized by deriving. In the component concentration measuring apparatus according to claim 6,
When the optical power measuring means changes only the optical power of one intensity modulated light of the two intensity modulated lights at the second time, the initial optical power before the change of the one intensity modulated light and the When the difference from the optical power at the first inflection point is measured and only the optical power of one of the two intensity-modulated lights is changed at the fourth time, this one intensity-modulated light A component concentration measuring apparatus for measuring a difference between an initial optical power before a change and an optical power at the second inflection point. The component concentration measuring apparatus according to claim 6 or 7,
The concentration deriving means changes the component concentration of the measurement target from the change amount of the optical power, the light absorption coefficient of the measurement target at the first time, and the light absorption coefficient change amount of the measurement target at the fourth time. A component concentration measuring apparatus that obtains an amount and derives the component concentration of the measurement target at the fourth time from the amount of change in the component concentration and the known component concentration. In the component concentration measuring apparatus according to claim 8,
And a light absorption spectrum measuring means for measuring the light absorption spectrum of the object to be measured and obtaining the light absorption coefficient and the amount of change in the light absorption coefficient from the measured spectrum. In the component concentration measuring apparatus according to any one of claims 6 to 9,
The optical power measuring means measures the difference in driving voltage between the two light irradiating means that emits two intensity-modulated lights as a difference in optical power,
The concentration deriving means refers to calibration data indicating the relationship between the drive voltage difference and the amount of change in optical power, and the drive voltage difference measured at the fourth time and the drive voltage measured at the second time A component concentration measuring device that obtains the amount of change in optical power from the amount of change from the difference.

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