JP5750750B2 - Concentration meter - Google Patents

Description

  The present invention relates to a concentration quantification apparatus, probe, concentration quantification method, and program for non-invasively quantifying the concentration of a target component in an arbitrary layer among observation targets composed of a plurality of light scattering medium layers in a living body. It is about.

In recent years, Japan is in the age of satiety, and the number of diabetic patients continues to increase every year. For this reason, the number of patients with diabetic nephritis will continue to increase every year. As a result, the number of patients with chronic renal failure continues to increase by 10,000 each year, and the number of patients exceeds 280,000.
On the other hand, with the arrival of an aging society, the importance of metabolic rate management in individuals is rapidly increasing in response to increasing demand for preventive medicine. Among them, the blood glucose level can be measured based on the reaction of glucose metabolism at the very early stage of diabetes, whereby early treatment based on early diagnosis of diabetes can be performed.

Conventionally, blood sugar levels are measured by collecting blood from a vein such as an arm or fingertip, and measuring enzyme activity for glucose in the blood. However, in such a blood sugar level measuring method, blood sampling is complicated. In addition, there are various problems such as blood collection with pain and further risk of infection.
In addition, as a method for continuously measuring blood glucose level, an instrument that continuously measures glucose corresponding to blood glucose level with a needle inserted into a vein has been developed in the United States. However, since the needle is stuck in the vein, there is a risk that the needle may come off during the measurement of blood glucose level and there is a risk of infection.
Therefore, development of a blood glucose level measuring apparatus that can measure blood glucose level frequently without blood collection and that is free from the risk of infectious diseases is demanded. Furthermore, there is a demand for the development of a blood glucose level measuring device that can be easily and always worn and can be miniaturized.

In view of this, an apparatus that non-invasively measures blood glucose levels has been proposed by applying a spectroscopic analyzer that uses the principle of molecular absorption to a blood glucose level measuring apparatus (see, for example, Patent Document 1).
This device irradiates the skin with near-infrared continuous light and calculates the glucose concentration from the amount of light absorption. Specifically, the glucose concentration and the wavelength and light of the near-infrared light to be irradiated in advance. A calibration curve showing the relationship with the amount of absorption of light, irradiating the skin with near-infrared continuous light, scanning the wavelength of the return light from the skin with a monochromator, etc. The amount of light absorbed for each wavelength in the region is obtained, and the amount of light absorbed at each wavelength is compared with a calibration curve to calculate the glucose concentration in blood, that is, the blood sugar level.

In general, when performing near-infrared spectroscopic analysis of an aqueous solution or a sample with a high water content, their spectra, like the spectrum of water, have large fluctuations such as a shift of the spectrum accompanying a temperature change. When used for quantitative analysis, the effect of the temperature of the aqueous solution or sample cannot be ignored.
Therefore, when measuring the glucose concentration in the tissue near the surface of the living body using light absorption in the near-infrared region, the measurement surface of the probe tip of the optical fiber bundle for near-infrared light reception and emission and the living body There has also been proposed an apparatus for making the temperature of the contact portion with the tissue in the vicinity of the surface constant by using a heater and a surface temperature detecting means (see Patent Document 2).

Japanese Patent No. 3931638 JP 2001-299727 A

However, the conventional glucose concentration measuring apparatus has the following problems.
1stly, in the apparatus which heats the temperature of the contact part of the probe front-end | tip of the optical fiber bundle for near-infrared light receiving and emitting, and the structure | tissue near the surface of a biological body with a heater, the measurement surface which contact | connects the surface of a biological body It is necessary to supply energy for heating the body, and a control means for making the temperature of the contact portion with the tissue near the surface of the living body constant is required, which makes it difficult to reduce the size of the apparatus and increase the price.

  Secondly, the control means for keeping the temperature of the contact portion with the tissue near the surface of the living body constant at the time when the near-infrared light receiving and emitting part is brought into contact with the tissue near the surface, The measurement is performed after a predetermined time has elapsed since the start of near-infrared light irradiation, and the target temperature, the environmental temperature, and the near-infrared light receiving and emitting means are placed on the living body. It is determined from the living body temperature at the time of contact with the tissue in the vicinity of the surface. Therefore, after obtaining the environmental temperature at the time of contact with the tissue near the surface, until the temperature of the living body after a predetermined time has elapsed after contacting the tissue near the surface and starting irradiation with near infrared light In addition, it is necessary to obtain a sufficient thermal equilibrium state. It usually takes 60 to 120 seconds to obtain this thermal equilibrium state.

  As described above, the conventional glucose concentration measuring device cannot achieve downsizing of the device and greatly shorten the measuring time while ensuring the accuracy of quantifying the glucose concentration in the living body.

  The present invention has been made in order to solve the above-described problems, and accurately grasps each state such as a resting state and an active state of a living body, as well as a physical state, a mental state, and a pulse rate of the living body. By taking into account the periodic fluctuations of the device, it is possible to reduce the size of the device while significantly ensuring the accuracy of quantifying the glucose concentration in the observation target in the living body, and to greatly shorten the measurement time It is an object of the present invention to provide a concentration determination device, a probe, a concentration determination method, and a program.

In order to solve the above-described problems, the present invention employs the following concentration determination apparatus, probe, concentration determination method, and program.
That is, the concentration quantification device of the present invention is a concentration quantification device that quantifies the concentration of a target component in an arbitrary layer among observation targets configured by a plurality of layers of light scattering media in a living body. A pulse wave detecting means for detecting a pulse pressure in the vicinity of a part having a pulsation over a predetermined area, a temperature measuring means provided in the vicinity of the pulse wave detecting means for measuring a temperature over the predetermined area, and the temperature A heat insulating material covering the periphery of the measuring means, a body temperature specifying means for specifying the temperature of the part where the maximum pulse pressure is detected among the pulse pressures detected over the predetermined region, and the observation target Irradiating means for irradiating light, and light scattering medium layer selecting means for selecting backscattered light emitted from the arbitrary layer from a plurality of types of backscattered light emitted from the observation object by irradiating the light And any layer above The light receiving means for receiving the backscattered light emitted from the light receiving means, the light intensity obtaining means for obtaining the intensity of the backscattered light emitted from the arbitrary layer received by the light receiving means, and the light intensity obtaining means. Light absorption coefficient calculation means for calculating the light absorption coefficient of the arbitrary layer based on light intensity, and the concentration of the target component in the arbitrary layer based on the light absorption coefficient calculated by the light absorption coefficient calculation means And a concentration correction unit that corrects the concentration of the target component calculated by the concentration calculation unit based on the body temperature specified by the body temperature specifying unit. .
That is, the concentration quantification device of the present invention is a concentration quantification device that quantifies the concentration of a target component in an arbitrary layer among observation targets configured by a plurality of layers of light scattering media in a living body. A pulse wave detecting means for detecting a pulse pressure in the vicinity of a portion having a pulsation over a predetermined area, a temperature measuring means provided in the vicinity of the pulse wave detecting means for measuring a temperature over the predetermined area, and the predetermined Body temperature specifying means for specifying the temperature of the part of the pulse pressure detected over the region as the body temperature of the living body, irradiation means for irradiating the observation object with light, and the light Light scattering medium layer selection means for selecting backscattered light emitted from the arbitrary layer from a plurality of types of backscattered light emitted from the observation object by irradiation, and backscattered emitted from the arbitrary layer Receiving light A light intensity acquisition means for acquiring the intensity of backscattered light emitted from the arbitrary layer received by the light receiving means, and the light intensity acquisition means based on the light intensity acquired by the light intensity acquisition means. A light absorption coefficient calculating means for calculating a light absorption coefficient; a concentration calculating means for calculating a concentration of the target component in the arbitrary layer based on the light absorption coefficient calculated by the light absorption coefficient calculating means; and the concentration calculation. And a concentration correcting means for correcting the concentration of the target component calculated by the means based on the body temperature specified by the body temperature specifying means.

In the concentration quantification device of the present invention, the pulse wave detection means detects the pulse pressure in the vicinity of the part having pulsation in the living body over a predetermined area, the temperature measurement means measures the temperature over the predetermined area, The body temperature specifying means specifies the temperature of the part where the maximum pulse pressure is detected among the pulse pressures detected over a predetermined region as the body temperature approximating the deep body temperature of the living body.
Further, the concentration correction means corrects the calculated concentration of the target component based on the body temperature specified by the body temperature specifying means.
Thus, after specifying the body temperature approximated to the deep body temperature of the living body, the calculated concentration of the target component is corrected based on the body temperature specified by the pulse wave detecting means, the temperature measuring means, and the body temperature specifying means. The concentration of the target component in an arbitrary observation target layer calculated based on the backscattered light can be accurately detected according to the activity state of the living body. Therefore, the concentration of the target component in an arbitrary layer to be observed can be accurately measured in a short time in a non-invasive manner according to the activity state of the living body.

The concentration quantification device according to the present invention is a concentration quantification device for quantifying the concentration of a target component in an arbitrary layer among observation targets configured by a plurality of layers of light scattering media in a living body, wherein A pulse wave detecting means for detecting a pulse pressure in the vicinity of a moving part over a predetermined region, a surface temperature measuring means for measuring a temperature in the vicinity of the surface of the observation object, and a sensor for measuring a temperature in the vicinity of the surface temperature measuring means The difference between the temperature near the surface of the observation target measured by the internal temperature measuring means, the surface temperature measuring means, and the temperature near the surface temperature measuring means measured by the sensor internal temperature measuring means is calculated per unit time. A surface / internal temperature change rate calculating means for calculating a temperature change rate, a heat insulating material covering the periphery of the temperature measuring means, an irradiating means for irradiating the observation object with light, Light scattering medium layer selection means for selecting backscattered light emitted from the arbitrary layer from a plurality of types of backscattered light emitted from the observation object by irradiating light, and emitted from the arbitrary layer Based on the light receiving means for receiving the back scattered light, the light intensity acquiring means for acquiring the intensity of the back scattered light emitted from the arbitrary layer received by the light receiving means, and the light intensity acquired by the light intensity acquiring means. A light absorption coefficient calculating means for calculating the light absorption coefficient of the arbitrary layer, and a concentration for calculating the concentration of the target component in the arbitrary layer based on the light absorption coefficient calculated by the light absorption coefficient calculating means. If the temperature change rate calculated by the calculating means and the surface / internal temperature change rate calculating means is a set value, the concentration calculation is performed based on the temperature near the surface of the observation target measured by the surface temperature measuring means. Means characterized by being provided with, a density correcting means for correcting the concentration of the target component was calculated.
The concentration quantification device according to the present invention is a concentration quantification device for quantifying the concentration of a target component in an arbitrary layer among observation targets configured by a plurality of layers of light scattering media in a living body, wherein A pulse wave detecting means for detecting a pulse pressure in the vicinity of a moving part over a predetermined region, a surface temperature measuring means for measuring a temperature in the vicinity of the surface of the observation object, and a sensor for measuring a temperature in the vicinity of the surface temperature measuring means Temperature measuring means comprising an internal temperature measuring means, irradiation means for irradiating the observation object with light, and the arbitrary layer from a plurality of types of backscattered light emitted from the observation object by irradiating the light Light scattering medium layer selection means for selecting the backscattered light emitted more from the light receiving means, light receiving means for receiving the backscattered light emitted from the arbitrary layer, and the arbitrary layer received by the light receiving means. Backward A light intensity acquisition means for acquiring the intensity of turbulent light; a light absorption coefficient calculation means for calculating a light absorption coefficient of the arbitrary layer based on the light intensity acquired by the light intensity acquisition means; and the light absorption coefficient calculation. Based on the light absorption coefficient calculated by the means, a concentration calculating means for calculating the concentration of the target component in the arbitrary layer, and the surface temperature measuring means measures the concentration of the target component calculated by the concentration calculating means. Density correction means for correcting based on the difference between the temperature near the surface of the observation target and the temperature near the surface temperature measurement means measured by the sensor internal temperature measurement means. .

In the concentration quantification apparatus of the present invention, the pulse wave detection means detects the pulse pressure in the vicinity of the part having pulsation in the living body over a predetermined region, and the surface temperature measurement means measures the temperature near the surface of the observation target. Then, the temperature in the vicinity of the surface temperature measuring means is measured by the sensor internal temperature measuring means.
Further, the concentration correction means determines the concentration of the target component based on the difference between the temperature near the surface of the observation target measured by the surface temperature measurement means and the temperature near the surface temperature measurement means measured by the sensor internal temperature measurement means. To correct.
By correcting in this way, it is possible to accurately detect the concentration of the target component in an arbitrary observation target layer calculated based on the backscattered light. Therefore, the concentration of the target component in any layer to be observed can be accurately measured in a short time non-invasively.

（但し、Ｉ（ｔ）は前記受光手段が時刻ｔにて受光した光強度、Ｎ（ｔ）は前記短時間パルス光の時間分解波形のモデルの時刻ｔにおける光強度、Ｌｉ（ｔ）は前記複数の光散乱媒質の各々の層における伝搬光路長分布のモデルの時刻ｔにおける第ｉ層の光路長、μｉは第ｉ層の光吸収係数である）
から算出することを特徴とする。 The concentration quantification device according to the present invention uses the light as a short-time pulse light, and further propagates the optical path length of each of the layers of the plurality of light scattering media of the short-time pulse light irradiated onto the observation target. From the optical path length distribution storage means for storing the model of distribution, the time-resolved waveform storage means for storing the model of the time-resolved waveform of the short-time pulse light irradiated to the observation object, and the optical path length distribution storage means, From the optical path length acquisition means for acquiring the optical path length of each of the layers of the light scattering medium at the predetermined time of the model of the propagation optical path length distribution, and the short-time pulse from the time-resolved waveform storage means and a light intensity model acquiring means for acquiring the intensity of the light in the predetermined time in the model of the time-resolved waveform of the light, the light intensity obtaining unit, in a plurality of times t ₁ ~t _m of the optional layer Gets the intensity, the light absorption coefficient calculating means, a light absorption coefficient of the arbitrary layer, the following equation (1)

(Where I (t) is the light intensity received by the light receiving means at time t, N (t) is the light intensity at time t of the model of the time-resolved waveform of the short-time pulse light, and Li (t) is the light intensity (The optical path length of the i-th layer at time t in the model of the propagation optical path length distribution in each layer of the plurality of light scattering media, and μ i is the light absorption coefficient of the i-th layer)
It is characterized by calculating from.

In the concentration determination apparatus of the present invention, the light intensity acquisition unit acquires the light intensity at a plurality of times t _{1 to} t _m of an arbitrary layer, and the light absorption coefficient calculation unit calculates the light absorption coefficient of the arbitrary layer as described above. Is calculated from the equation (1).
Thus, by measuring the time-resolved backscattered light, it is possible to reduce the backscattered light from any layer other than any layer as noise, and to reduce the influence of the temperature of each layer on the concentration of the target component. it can. Therefore, the concentration of the target component can be measured with higher accuracy.

（但し、Ｉ（ｔ）は前記受光手段が時刻ｔにて受光した光強度、Ｎ（ｔ）は前記短時間パルス光の時間分解波形のモデルの時刻ｔにおける光強度、Ｌｉ（ｔ）は前記複数の光散乱媒質の層各々の層における伝搬光路長分布のモデルの時刻ｔにおける第ｉ層の光路長、ｎは前記観測対象となる層の数、μｉは第ｉ層の光吸収係数である）
から算出することを特徴とする。 The concentration quantification device according to the present invention uses the light as a short-time pulse light, and further propagates the optical path length of each of the layers of the plurality of light scattering media of the short-time pulse light irradiated onto the observation target. From the optical path length distribution storage means for storing the model of distribution, the time-resolved waveform storage means for storing the model of the time-resolved waveform of the short-time pulse light irradiated to the observation object, and the optical path length distribution storage means, From the optical path length acquisition means for acquiring the optical path length of each of the layers of the light scattering medium at the predetermined time of the model of the propagation optical path length distribution, and the short-time pulse from the time-resolved waveform storage means A light intensity model acquisition unit that acquires the light intensity at the predetermined time of the model of the time-resolved waveform of light, the light intensity acquisition unit between a predetermined time and at least a predetermined time τ Get the light intensity, the light absorption coefficient calculating means, a light absorption coefficient of the arbitrary layer, the following equation (2)

(Where I (t) is the light intensity received by the light receiving means at time t, N (t) is the light intensity at time t of the model of the time-resolved waveform of the short-time pulse light, and Li (t) is the light intensity The optical path length of the i-th layer at time t in the model of the propagation optical path length distribution in each layer of the plurality of light scattering media, n is the number of layers to be observed, and μi is the light absorption coefficient of the i-th layer. )
It is characterized by calculating from.

In the concentration quantification device of the present invention, the light intensity acquisition means acquires light intensity between a predetermined time and at least a predetermined time τ, and the light absorption coefficient calculation means calculates the light absorption coefficient of an arbitrary layer as described above. (2) is calculated.
Thus, by measuring the time-resolved backscattered light, it is possible to reduce the backscattered light from any layer other than any layer as noise, and to reduce the influence of the temperature of each layer on the concentration of the target component. it can. Therefore, the concentration of the target component can be measured with higher accuracy.

（但し、μａは前記任意の層である第ａ層における光吸収係数、ｇｊは前記観測対象を構成する第ｊ成分のモル濃度、εｊは第ｊ成分の光吸収係数、ｐは前記観測対象を構成する主成分の個数、ｑは前記短時間パルス光の種類数である）
から算出することを特徴とする。 In the concentration quantification apparatus of the present invention, the concentration calculation means calculates the concentration of the target component in the arbitrary layer by the following equation (3):

(Where μa is the light absorption coefficient in the a-th layer which is the arbitrary layer, gj is the molar concentration of the j-th component constituting the observation object, εj is the light absorption coefficient of the j-th component, and p is the observation object. (The number of constituent main components, q is the number of types of short-time pulsed light)
It is characterized by calculating from.

In the concentration determination apparatus of the present invention, the concentration calculation means calculates the concentration of the target component in an arbitrary layer from the above equation (3).
Thus, by calculating the concentration of the target component in any layer using the backscattered light that has been time-resolved, the influence of the temperature of each layer on the concentration of the target component can be reduced. Therefore, the concentration of the target component can be measured with higher accuracy.

  The concentration quantification device of the present invention comprises a plurality of sets of the pulse wave detection means and the temperature measurement means, and the body temperature specifying means is the pulse of the site where the maximum pulse pressure is detected among the plurality of pulse wave detection means. The temperature measured by the temperature measuring means in the vicinity of the wave detecting means is specified as a body temperature.

In the concentration quantifying device of the present invention, the temperature measured by the temperature measuring means in the vicinity of the pulse wave detecting means at the site where the maximum pulse pressure is detected among the plurality of pulse wave detecting means is specified as the body temperature.
In this way, by using a plurality of sets of pulse wave detection means and temperature measurement means, the body temperature of the site to be observed in the living body can be measured in a short time with a minimum operation.

  The concentration quantification apparatus according to the present invention provides the temperature measuring means with a temperature near the surface of the observation target measured by the surface temperature measuring means, and a temperature near the surface temperature measuring means measured by the sensor internal temperature measuring means. A surface / internal temperature change rate calculating means for calculating the difference between the two as a temperature change rate per unit time is provided.

In the concentration determination apparatus of the present invention, the surface / internal temperature change rate calculating means calculates the temperature near the surface of the observation target measured by the surface temperature measuring means and the vicinity of the surface temperature measuring means measured by the sensor internal temperature measuring means. Is calculated as a temperature change rate per unit time.
In this way, the concentration of the target component calculated by the concentration calculating means is the difference between the temperature near the surface of the observation target measured by the surface temperature measuring means and the temperature near the surface temperature measuring means measured by the sensor internal temperature measuring means. By correcting based on the temperature change rate per unit time calculated from the above, the influence of the temperature on the concentration of the target component in any observation target layer calculated based on this backscattered light can be made extremely small . Therefore, the concentration of the target component can be accurately measured noninvasively.

The concentration determination apparatus of the present invention is characterized in that the pulse wave detection means is provided with a pulsation discrimination means for determining whether or not the pulse wave detection means detects a pulse pressure.
In the concentration quantification device of the present invention, the pulse wave detecting means is provided with a pulsation discriminating means for determining whether or not the pulse wave detecting means detects the pulse pressure. Whether or not it is detected can be known with certainty, and there is no possibility of misidentifying detection of pulse pressure.

  The concentration quantification device of the present invention includes a body motion detecting unit for detecting the body motion of the living body, and a body motion determining unit for determining whether or not the body motion of the living body detected by the body motion detecting unit is within a predetermined range. And.

  In the concentration quantification device of the present invention, when the body movement detecting means detects the body movement of the living body, and the body movement determining means determines that the body movement of the living body detected by the body movement detecting means is within a predetermined range, Since the body temperature of the living body is specified, it is possible to accurately grasp the resting state of the living body. Therefore, it is possible to accurately detect the concentration of the target component in an arbitrary layer to be observed in a state where the living body is at rest.

The probe of the present invention has a pulse wave detection means for detecting a pulse pressure in the vicinity of a part having a pulsation in a living body over a predetermined area, and measures a temperature over the predetermined area provided in the vicinity of the pulse wave detection means. Irradiating means for irradiating light to an observation target composed of a plurality of light scattering medium layers in the living body, and a temperature measuring means for performing irradiation, and irradiating the light Light scattering medium layer selection means for selecting backscattered light radiated from an arbitrary layer from a plurality of types of backscattered light radiated from the observation target, and light reception for receiving backscattered light radiated from the arbitrary layer And means.
The probe of the present invention has a pulse wave detection means for detecting a pulse pressure in the vicinity of a part having a pulsation in a living body over a predetermined area, and measures a temperature over the predetermined area provided in the vicinity of the pulse wave detection means. Temperature measuring means, irradiating means for irradiating the observation object with light, and backscattered light emitted from the arbitrary layer from a plurality of types of backscattered light emitted from the observation object by irradiating the light A light scattering medium layer selecting means for selecting the light and a light receiving means for receiving the backscattered light emitted from the arbitrary layer.

In the probe of the present invention, the pulse wave detection means detects the pulse pressure in the vicinity of the part having pulsation in the living body over a predetermined area, the temperature measurement means measures the temperature over the predetermined area, and the irradiation means observes it. The target is irradiated with light, the backscattering light emitted from any layer is selected from the multiple types of backscattered light emitted from the observation target by the light scattering medium layer selection means, and the light receiving means emits the light from any layer. Receiving backscattered light.
Thus, the backscattered light radiated from the body temperature of the observation target and the arbitrary layer of the observation target in the living body can be collected efficiently and simply.

  The concentration quantification method of the present invention is a concentration quantification method for quantifying the concentration of a target component in an arbitrary layer among observation targets constituted by a plurality of light scattering medium layers in a living body, and comprises a pulse wave detection means. The pulsation pressure in the vicinity of the part having pulsation in the living body is detected over a predetermined area, the temperature is measured over the predetermined area by the temperature measuring means, and then the body temperature specifying means is applied to the predetermined area. The temperature of the part where the maximum pulse pressure is detected among the detected pulse pressures is specified as the body temperature of the living body, and then the observation object is irradiated with light by the irradiation means, and then the light scattering medium layer The selection means selects the backscattered light emitted from the arbitrary layer from the plurality of types of backscattered light emitted from the observation object by irradiating the light, and then the light receiving means selects the optional layer. Or Receiving the emitted backscattered light, then, by the light intensity acquisition means, obtain the intensity of the backscattered light emitted from the arbitrary layer received by the light receiving means, then, by the light absorption coefficient calculation means, Based on the light intensity acquired by the light intensity acquisition means, calculate the light absorption coefficient of the arbitrary layer, and then by the concentration calculation means based on the light absorption coefficient calculated by the light absorption coefficient calculation means, Calculating a concentration of the target component in an arbitrary layer, and then correcting the concentration of the target component calculated by the concentration calculation unit based on the body temperature specified by the body temperature specifying unit by a concentration correction unit; It is characterized by.

In the concentration quantification method of the present invention, the pulse wave detection means detects a pulse pressure in the vicinity of a part having a pulsation in a living body over a predetermined area, the temperature measurement means measures the temperature over the predetermined area, and the body temperature is specified. By the means, the temperature of the part where the maximum pulse pressure is detected among the pulse pressures detected over the predetermined region is specified as the body temperature approximating the deep body temperature of the living body.
Thus, after specifying a body temperature that approximates the deep body temperature of the living body, the concentration of the target component in any layer to be observed is corrected by correcting the calculated concentration of the target component based on this body temperature. It can be accurately detected according to the activity state. Therefore, the concentration of the target component in an arbitrary layer to be observed can be accurately measured in a short time in a non-invasive manner according to the activity state of the living body.

The program of the present invention provides a computer of a concentration quantification apparatus for quantifying the concentration of a target component in an arbitrary layer among observation targets constituted by a plurality of layers of light scattering media in a living body, and transmits the pulsation in the living body. A pulse wave detection procedure for detecting the pulse pressure in the vicinity of the site over a predetermined region, a temperature measurement procedure for measuring the temperature over the predetermined region,
The body temperature specifying procedure for specifying the temperature of the part of the pulse pressure detected over the predetermined region as the body temperature of the living body, the irradiation procedure for irradiating the observation target with light, and the light Light scattering medium layer selection procedure for selecting backscattered light emitted from the arbitrary layer from a plurality of types of backscattered light emitted from the observation object by irradiation, backscattered light emitted from the arbitrary layer Based on the light intensity acquired in the light intensity acquisition procedure, the light intensity acquisition procedure for acquiring the intensity of the backscattered light emitted from the arbitrary layer obtained in the light reception procedure, Light absorption coefficient calculation procedure for calculating the light absorption coefficient of the arbitrary layer, concentration calculation procedure for calculating the concentration of the target component in the arbitrary layer based on the light absorption coefficient calculated in the light absorption coefficient calculation procedure The above The concentration of the target component calculated by the degree calculation procedure, characterized in that to perform the density correction procedure that corrects, based on the body temperature specified by said body temperature specifying step.

In the program of the present invention, a pulse wave detection procedure for detecting a pulse pressure in the vicinity of a part having a pulsation in a living body over a predetermined region, a temperature measurement procedure for measuring a temperature over the predetermined region A body temperature specifying procedure for specifying the temperature of the part where the maximum pulse pressure is detected among the pulse pressures detected over a predetermined region as the body temperature of the living body is executed.
As described above, by performing the above-described pulse wave detection procedure, temperature measurement procedure, and body temperature specifying procedure, after specifying a body temperature that approximates the deep body temperature of the living body, the concentration of the target component is determined based on the specified body temperature. Correction can be made with high accuracy. Therefore, the concentration of the target component in an arbitrary layer to be observed can be accurately measured in a short time in a non-invasive manner according to the activity state of the living body.

It is a schematic block diagram which shows the structure of the blood glucose level measuring apparatus of the 1st Embodiment of this invention. It is sectional drawing which shows the structure of the light guide part of the blood glucose level measuring apparatus of the 1st Embodiment of this invention. It is a perspective view which shows the external appearance of the measurement site | part of the palm vicinity of a human body. It is a figure which shows the measurement result of the temperature in each measurement point of the palm vicinity of a human body. It is a schematic diagram which shows the wide circulation system of a human body. It is a schematic diagram which shows the cross section of the skin tissue of a human body. It is a figure which shows the propagation optical path length distribution of each layer which the simulation part computed. It is a figure which shows the time-resolved waveform which the simulation part computed. It is a figure which shows the light absorption wavelength characteristic by water. It is a figure which shows the absorption spectrum of the main component of skin. It is a figure which shows the relationship between the wavelength of the light irradiated to each of the epidermis layer of a skin, a dermis layer, and subcutaneous tissue, and an absorption coefficient. It is a figure which shows the temperature dependence of the absorbance spectrum of water. It is a figure which shows the temperature dependence of the light absorption spectrum difference. It is a figure which shows an example of the absorbance spectrum of glucose aqueous solution. It is a flowchart which shows the procedure which measures the blood glucose level of the 1st Embodiment of this invention. It is sectional drawing which shows the structure of the light guide part of the blood glucose level measuring apparatus of the 2nd Embodiment of this invention. It is a flowchart which shows the procedure which measures the blood glucose level of the 2nd Embodiment of this invention. It is sectional drawing which shows the structure of the light guide part of the blood glucose level measuring apparatus of the 3rd Embodiment of this invention. It is a flowchart which shows the procedure which measures the blood glucose level of the 3rd Embodiment of this invention. It is sectional drawing which shows the structure of the light guide part of the blood glucose level measuring apparatus of the 4th Embodiment of this invention. It is a flowchart which shows the procedure which measures the blood glucose level of the 4th Embodiment of this invention. It is a schematic block diagram which shows the structure of the blood glucose level measuring apparatus of the 5th Embodiment of this invention. It is a flowchart which shows the procedure which measures the blood glucose level of the 5th Embodiment of this invention. It is a schematic block diagram which shows the structure of the blood glucose level measuring apparatus of the 6th Embodiment of this invention. It is a flowchart which shows the procedure which measures the blood glucose level of the 6th Embodiment of this invention.

A concentration quantification apparatus, a probe, a concentration quantification method, and an embodiment for executing a program according to the present invention will be described.
In the present invention, a blood glucose level measuring device is used as a concentration determination device, the skin of the palm of a human body (living body) is used as an observation target, glucose is used as a target component, and short-time pulse light of a specific wavelength is used as light of a specific wavelength. I will explain.

[First Embodiment]
FIG. 1 is a schematic block diagram showing a configuration of a blood glucose level measuring apparatus according to a first embodiment of the present invention, and FIG. 2 is a cross-sectional view showing an outline of a configuration of a light guide unit of the blood glucose level measuring apparatus.
This blood glucose level measuring apparatus 1 non-invasively determines the concentration of glucose (target component) contained in the dermis layer (arbitrary layer) of a plurality of layers constituting the skin (observation target) such as the palm of a human body (living body). A simulation unit 2, an optical path length distribution storage unit (optical path length distribution storage unit) 3, a time-resolved waveform storage unit (time-resolved waveform storage unit) 4, and an irradiation unit (irradiation unit) 5. , A light guide unit 6, a light scattering medium layer selection unit (light scattering medium layer selection unit) 7, a light receiving unit (light receiving unit) 8, a light intensity acquisition unit (light intensity acquisition unit) 9, and an optical path length acquisition unit (Optical path length acquisition means) 10, non-absorption light intensity acquisition section (light intensity model acquisition means) 11, light absorption coefficient calculation section (light absorption coefficient calculation means) 12, concentration calculation section (concentration calculation means) 13, , Body temperature specifying part (body temperature specifying means) 14 and density correcting part (concentration correcting means) And a 15.

The simulation unit 2 performs a simulation of irradiating light to a skin model having a light absorption coefficient of zero.
The optical path length distribution storage unit 3 stores a model of a propagation optical path length distribution in each layer constituting the skin of short-time pulse light irradiated to the skin. Here, the propagation optical path length distribution of the skin model having zero light absorption coefficient is stored.
The time-resolved waveform storage unit 4 stores a model of a time-resolved waveform of short-time pulse light that is applied to the skin. Here, the time-resolved waveform of the skin model having zero light absorption coefficient is stored.
The irradiation unit 5 irradiates the skin with pulsed light for a short time. Here, the short-time pulsed light means pulsed light having a pulse width of about 100 psec or less. Note that pulse light having a pulse width in the range of 0.1 psec to several psec may be used as the short-time pulse light.

As shown in FIG. 2, the light guide 6 is in close contact with the skin 31 and guides short-time pulsed light emitted from the irradiation unit 5 toward the skin 31, and the irradiation light guide 21. A light receiving light guide 22 that is integrally provided outside the light source and collects a plurality of types of backscattered light radiated from the skin 31 and guides it to the light scattering medium layer selector 7, and an integrated irradiation light guide 21 and a heat insulating material 23 provided outside the light receiving light guide 22 and a pulse wave which is provided on the skin side surface of the heat insulating material 23 and detects a pulse pressure in the vicinity of a portion having a pulsation in the human body over a predetermined region. A sensor (pulse wave detecting means) 24, a temperature sensor (temperature measuring means) 25 provided in the vicinity of the pulse wave sensor 24 and measuring the temperature over the predetermined area, the irradiation light guide path 21, the light receiving light guide path 22, and the heat insulation. It is comprised by the base 26 which fixes the material 23
The light guide unit 6, the irradiation unit 5, and the light receiving unit 8 constitute a probe of the present invention.

The irradiation light guide path 21 and the light receiving light guide path 22 may be made of a material that has a small absorption loss of short-time pulsed light to be guided. For example, quartz glass, plastic such as polymethyl methacrylate (PMMA) or polyethylene is preferably used. It is done.
The heat insulating material 23 may be any material having a heat insulating property with a sufficiently small heat capacity so as not to affect the temperature change of the skin 31. It is preferable that the space | interval of this heat insulating material 23 and the skin 31 is separated so that the heat insulating material 23 may not receive the temperature change of this skin 31 directly, and 0.5 mm-1.0 mm are preferable. In this heat insulating material 23, the heat response time until the temperature sensor 25 reaches 90% of the temperature reached value is suppressed within 0.2 seconds by making the heat capacity sufficiently small so as not to affect the temperature change of the skin 31. Can do.

The pulse wave sensor 24 detects, in a non-contact manner, the pulse pressure in the vicinity of a part having a pulsation in the human body, that is, the pulse pressure of an artery such as the radial artery over the surface (predetermined region) of the skin 31 including the artery. To do.
In this pulse wave sensor 24, the detected analog signal is converted into a digital signal by an A / D converter, and the pulse wave signal converted into a digital signal by an FFT (Fast Fourier Transform) processor is subjected to FFT processing, and thereafter The number of beats is obtained from the pulse wave signal subjected to the FFT processing in the number-of-beats calculation unit, and this number of beats is output as a signal.

In addition, when calculating | requiring a beat rate from a pulse wave signal, you may obtain | require the peak interval of the waveform of the taken-in pulse wave signal, and may calculate the reciprocal number as a beat number.
In addition, what is originally required is the heart rate, that is, the heart rate per unit time. However, since heart rate = pulse rate, the pulse rate may be obtained. Therefore, a configuration may be adopted in which an electrocardiogram is detected to directly obtain the heart rate.
Although the pulse rate and the heart rate should be distinguished medically, it is not necessary to distinguish them in the present invention, so they are described as “beat number” including both.
Since the periphery of the pulse wave sensor 24 is covered with the heat insulating material 23, there is no possibility of being affected by the temperature of the irradiation light guide path 21 and the light receiving light guide path 22.
The pulse wave sensor 24 may be provided with pulsation discrimination means for discriminating whether or not the pulse pressure is detected as necessary.

The temperature sensor 25 measures the temperature in a non-contact manner over the surface (predetermined region) of the skin 31 including the artery.
In the temperature sensor 25, the detected analog signal is converted into a digital signal by the A / D converter and output.
Since the temperature sensor 25 is covered with the heat insulating material 23, the temperature sensor 25 can measure the temperature of the skin 31 without being affected by the temperature of the irradiation light guide path 21 and the light reception light guide path 22.

  In the light guide unit 6, when the irradiation light guide path 21 guides the short-time pulse light irradiated by the irradiation unit 5 and irradiates the skin 31 with the short-time pulse light, the short-time pulse light is applied to the skin 31. The skin 31 emits a plurality of types of backscattered light. These multiple types of backscattered light are guided to the light scattering medium layer selection unit 7 by the light receiving light guide path 22.

The light scattering medium layer selection unit 7 selects the backscattered light mainly emitted from the dermis layer from the multiple types of backscattered light emitted from the skin condensed and guided by the light guide unit 6.
The light receiving unit 8 receives light obtained by backscattering the short-time pulsed light by the skin.

The light intensity acquisition unit 9 acquires the light reception intensities at different times of the backscattered light emitted from the dermis layer received by the light receiving unit 8.
Here, it is preferable that the plurality of times include the peak time of the propagation optical path length distribution of each layer constituting the skin.
As described above, by including the peak time of the propagation optical path length distribution of each layer, an arbitrary layer, for example, the dermis layer can be efficiently selected from a plurality of skin layers.

The optical path length acquisition unit 10 acquires from the optical path length distribution storage unit 3 the optical path length of each layer of the skin at a predetermined time of the model of the propagation optical path length distribution. Here, the optical path length at a certain time is acquired from the optical path length distribution storage unit 3.
The non-absorption light intensity acquisition unit 11 acquires the light intensity at a predetermined time of the model of the time-resolved waveform of the short-time pulsed light from the time-resolved waveform storage unit 4. Here, the light intensity at a certain time is acquired from the time-resolved waveform storage unit 4.
The light absorption coefficient calculation unit 12 calculates the light absorption coefficient in the dermis layer of the skin irradiated with the short-time pulse light.

（但し、Ｉ（ｔ）は受光部８が時刻ｔにて受光した光強度、Ｎ（ｔ）は特定波長λｋの短時間パルス光の時間分解波形の無吸収モデルの時刻ｔにおける光強度、Ｌｉ（ｔ）は皮膚の各々の層における伝搬光路長分布のモデルの時刻ｔにおける第ｉ層の光路長、μｉは第ｉ層の光吸収係数である）
から算出する。
ここで、第１層は表皮層、第２層は真皮層、第３層は皮下組織を示し、μ_１は表皮層の光吸収係数、μ_２は真皮層の光吸収係数、μ_３は皮下組織の光吸収係数を示す。 In the light absorption coefficient calculation unit 12, the light absorption coefficient of an arbitrary layer in the skin is expressed by the following equation (4).

(Where I (t) is the light intensity received by the light receiving unit 8 at time t, N (t) is the light intensity at time t of the non-absorption model of the time-resolved waveform of the short-time pulse light of the specific wavelength λk, Li (T) is the optical path length of the i-th layer at time t of the model of the propagation optical path length distribution in each layer of the skin, and μ i is the light absorption coefficient of the i-th layer)
Calculate from
Here, the first layer is the epidermis layer, the second layer is the dermis layer, the third layer is the subcutaneous tissue, μ ₁ is the light absorption coefficient of the epidermis layer, μ ₂ is the light absorption coefficient of the dermis layer, and μ ₃ is the subcutaneous Shows the light absorption coefficient of tissue.

（但し、μａは皮膚の任意の層である第ａ層における光吸収係数、ｇｊは皮膚を構成する第ｊ成分のモル濃度、εｊは第ｊ成分の光吸収係数、ｐは皮膚を構成する主成分の個数、ｑは特定波長λｋの種類数である）
から算出する。
ここで、第１層は表皮層、第２層は真皮層、第３層は皮下組織を示し、μ_１は表皮層の光吸収係数、μ_２は真皮層の光吸収係数、μ_３は皮下組織の光吸収係数を示す。 The concentration calculation unit 13 calculates the concentration of glucose contained in the dermis layer from the light absorption coefficient in the dermis layer.
In the concentration calculation unit 13, the glucose concentration in an arbitrary layer of the skin is expressed by the following equation (5).

(Where μa is the light absorption coefficient in the a layer, which is an arbitrary layer of the skin, gj is the molar concentration of the jth component constituting the skin, εj is the light absorption coefficient of the jth component, and p is the main component constituting the skin. The number of components, q is the number of types of specific wavelength λk)
Calculate from
Here, the first layer is the epidermis layer, the second layer is the dermis layer, the third layer is the subcutaneous tissue, μ ₁ is the light absorption coefficient of the epidermis layer, μ ₂ is the light absorption coefficient of the dermis layer, and μ ₃ is the subcutaneous Shows the light absorption coefficient of tissue.

  The body temperature specifying unit 14 detects the pulse pressure data detected by the pulse wave sensor 24 in a non-contact manner across the surface of the skin 31 including the artery, and the temperature sensor 25 applies the non-pressure across the surface of the skin 31 including the artery. Based on the temperature data measured by contact, the maximum pulse pressure of the pulse pressure is specified, and the temperature of the part corresponding to the specified maximum pulse pressure is specified as a body temperature that approximates the deep body temperature of the living body. This body temperature is specified at regular intervals and stored in storage means such as a memory.

The concentration correction unit 15 corrects the glucose concentration of the dermis layer calculated by the concentration calculation unit 13 using the body temperature of the living body specified by the body temperature specifying unit 14.
The concentration correction unit 15 corrects the glucose concentration of the dermis layer calculated by the concentration calculation unit 13 using the difference between the body temperature of the living body specified by the body temperature specifying unit 14 and the reference temperature, thereby It is possible to accurately detect the glucose concentration in the dermis layer according to the activity state of the human body. Thereby, the glucose concentration in the dermis layer can be accurately measured in a short time in a non-invasive manner according to the activity state of the human body.

  In the blood glucose level measuring apparatus 1 configured as described above, the light guide unit 6 is slid in an arbitrary direction along the skin 31 while the tip of the light guide unit 6 is applied to the skin 31 of the human body. The pulse wave sensor 24 detects the pulse pressure in the vicinity of the part having pulsation in the human body, that is, the pulse pressure of the artery without contact. At this time, the temperature of the skin 31 in the vicinity of the artery detected by the pulse wave sensor 24 is measured by the temperature sensor 25 in a non-contact manner.

On the other hand, the short-time pulse light emitted from the irradiation unit 5 is irradiated to the skin 31 via the irradiation light guide path 21. A plurality of types of backscattered light is radiated from the skin 31, and these backscattered light are collected by the light receiving light guide path 22 and guided to the light scattering medium layer selection unit 7.
The light scattering medium layer selection unit 7 selects only the backscattered light mainly emitted from the dermis layer from the multiple types of backscattered light emitted from the skin 31. The light receiving unit 8 receives only backscattered light mainly emitted from the dermis layer.

Furthermore, the light intensity acquisition unit 9 acquires the light intensity of the backscattered light emitted from the dermis layer received by the light receiving unit 8 at time t.
On the other hand, the optical path length acquisition unit 10 acquires the optical path length of each layer of the skin 31 at the time t of the propagation optical path length distribution in the skin model from the optical path length distribution storage unit 3, and the non-absorbing light intensity acquisition unit 11 The light intensity at time t of the time-resolved waveform of the short-time pulse light in the skin model is acquired from the waveform storage unit 4.

Next, the light absorption coefficient calculation unit 12 includes the light intensity acquired by the light intensity acquisition unit 9, the optical path length of each layer of the skin acquired by the optical path length acquisition unit 10, and the light intensity acquired by the non-absorption light intensity acquisition unit 11. Based on the above, the light absorption coefficient of the dermis layer of the skin is calculated.
Next, the concentration calculation unit 13 calculates the concentration of glucose contained in the dermis layer of the skin 20 based on the equation (5) based on the light absorption coefficient calculated by the light absorption coefficient calculation unit 12.
In this way, the concentration of glucose contained in the dermis layer is calculated.

On the other hand, in the body temperature specifying unit 14, pulse pressure data detected by the pulse wave sensor 24 across the surface of the skin 31 including the artery and the surface of the skin 31 including the artery by the temperature sensor 25. Based on the temperature data measured in a non-contact manner, the maximum pulse pressure of the pulse pressure is specified, and the temperature corresponding to the specified maximum pulse pressure is specified as a body temperature that approximates the deep body temperature of the living body. To do.
Next, the concentration correction unit 15 corrects the glucose concentration of the dermis layer calculated by the concentration calculation unit 13 using the body temperature specified by the body temperature specifying unit 14.

  As described above, the concentration of glucose contained in the dermis layer calculated based on the backscattered light emitted from the dermis layer is measured by the pulse pressure data detected by the pulse wave sensor 24 and the artery measured by the temperature sensor 25. Is corrected using the body temperature specified by the pulse pressure body temperature specifying unit 14 based on the temperature data of the surface of the skin 31 including the dermal layer, thereby accurately detecting the glucose concentration in the dermis layer according to the activity state of the living body. be able to. Therefore, the glucose concentration contained in the dermis layer can be accurately measured noninvasively according to the activity state of the living body.

Next, the body temperature measurement principle of this embodiment will be described based on the results of experiments conducted by the inventors.
Here, the temperature distribution around the radial artery was measured using a radiation thermometer having an opening diameter of about 5 mm. As a result, it was found that the temperature just above the radial artery was 1 ° C. higher than the temperature at the periphery, and a body temperature close to normal heat was measured.

FIG. 3 is a perspective view showing the appearance of a measurement site in the vicinity of the palm of a human body.
The temperature was measured by providing measurement points at 5 mm intervals on a virtual line orthogonal to the radial artery / ulna artery moved 10 mm from the radial styloid process to the central side, and the temperature was measured for each of these measurement points. .

FIG. 4 is a diagram showing the temperature measurement results at each measurement point in the vicinity of the palm of the human body. FIG. 4 (a) shows the temperature measurement results when the arm is dry, and FIG. 4 (b) shows the measurement. It is a measurement result of temperature when temperature measurement is performed after the part is once wetted with water.
As is clear from FIG. 4, the temperature on the radial artery and the temperature on the ulnar artery are both higher than the temperature at the peripheral portion and are closer to the deep body temperature. In addition, the difference between the temperature on the artery and the periphery is more prominent when wet with water than when it is dry, especially the temperature measured on the radial artery. A value almost equal to that at the time of drying is obtained without being substantially affected by the above.

  Examining the above experimental results from a medical point of view, arteries such as radial arteries are routes that carry a heat source called blood, so the temperature of the epidermis directly above such arteries is higher than that of the surrounding area. It is thought that it is close enough to deep body temperature. In addition, in the portion directly above the radial artery, a fast pulsation with time response is observed as the blood from the heart is pumped. Find the part where this pulsation occurs and measure the temperature of this part. Thus, a body temperature sufficiently close to the deep body temperature can be obtained.

The site for detecting the pulsation may be anywhere as long as it is directly above the arteries other than the small arteries, that is, the aorta, middle artery, and small artery. For example, the radial artery mentioned above can be cited as the site of the middle artery, and the side trunk of the finger can be cited as the small artery.
FIG. 5 is a schematic diagram showing a wide circulation system of a human body.
This wide-area circulatory system is a physical vascular pathway that distributes blood from the heart to each part of the human body and returns blood from each part of the human body. On the other hand, the microscopic blood vessels involved in the exchange between body fluid and tissue, the accompanying lymph capillaries, and the circulatory unit including the interstitial or parenchyma surrounding them are called the microcirculatory system. . In the microcirculatory system, at the end of the arterial system, after the small arteries branch into reticulated capillaries, these capillaries reassemble to form small veins that are connected to the veins.

Thus, if the body temperature is measured at the peripheral part of the radial artery, etc., it is not under the unusual situation of being constantly exposed to water, but as long as you live normally, even if it is wet, etc. It is estimated that body temperature close to deep body temperature can be measured with accuracy.
For example, when the use of monitoring the change in the body temperature of the person during sleep is assumed, the body temperature can be measured without any problem by applying the above measurement principle.
In the present embodiment, in order to measure a body temperature sufficiently close to the deep body temperature of the human body, the pulse wave sensor 24 and the temperature sensor 25 of the light guide unit 6 described above are disposed on the surface of the radial artery of the human body. The temperature detected at the part of the human body is the body temperature.

Conventionally, when measuring deep body temperature, temperature measurement was performed on the rectum, tongue, armpit, etc. using a small device such as a thermometer or a table-type device, but at these points, one point at a certain time It was only possible to measure. Furthermore, since the device is generally large, it has not been possible to continuously measure body temperature while being carried around.
On the other hand, in this embodiment, body temperature sufficiently close to the deep body temperature of the human body can be measured relatively easily. Therefore, the body temperature sufficiently close to the deep body temperature measured in the present embodiment is not only useful for calculating calorie consumption, but also the body temperature itself is extremely important from the viewpoint of clinical medicine.

Next, the operation of the blood sugar level measuring apparatus 1 will be described.
The blood glucose level measuring apparatus 1 needs to calculate the propagation optical path length distribution and the time-resolved waveform in each layer of the skin model before measuring the blood glucose level.
FIG. 6 is a schematic diagram showing a cross-section of human skin tissue. The skin 31 includes approximately 20% of water and the rest is made of protein, and a skin layer 32 having a thickness of approximately 0.3 mm and a skin layer 32 below the skin layer 32. A dermis layer (arbitrary layer) 33 having a thickness of about 1.2 mm containing approximately 60% water and containing proteins, lipids and glucose, and formed under the dermis layer 33, and generally containing 90% or more of lipids; The remaining part is composed of water and a subcutaneous tissue 34 having a thickness of about 3.0 mm.

Here, a method of calculating the propagation optical path length distribution and time-resolved waveform of the skin model will be described.
First, the simulation unit 2 generates a skin model. The skin model is generated by determining the light scattering coefficient, light absorption coefficient, and thickness of each layer of the skin. Here, since the scattering coefficient and thickness of each layer of the skin have little difference between individuals, it is preferable to determine by taking a sample in advance.
The light absorption coefficient of the skin model used here is zero. The reason is that the amount of light absorption is calculated using this skin model.

  When the simulation unit 2 generates a skin model, the simulation unit 2 performs a simulation of irradiating the skin model with light. At this time, it is necessary to determine the distance between the position of the irradiation unit 5 and the position of the light receiving unit 8. The simulation is preferably performed using the Monte Carlo method. The simulation by the Monte Carlo method is performed as follows, for example.

ただし、μｓは、皮膚モデルの第ｓ層（表皮層、真皮層、皮下組織層の何れか）の散乱係数を示す。 First, the simulation unit 2 performs calculation for irradiating the skin model with a photon (light beam) as a model of light to be irradiated. Photons irradiated to the skin model move in the skin model. At this time, the photon determines the distance L and the direction θ to the next advancing point by the random number R. The simulation unit 2 calculates the distance L to the point at which the photon advances next using Equation (6).

Here, μs represents the scattering coefficient of the s-th layer (any one of the epidermis layer, dermis layer, and subcutaneous tissue layer) of the skin model.

ただし、ｇは、散乱角度の余弦（ｃｏｓ）の平均である非等方性パラメータを示し、皮膚の非等方性パラメータは、略０．９である。 In addition, the simulation unit 2 calculates the direction θ up to the point where the photon advances next by Expression (7).

However, g shows the anisotropic parameter which is the average of the cosine (cos) of a scattering angle, and the anisotropic parameter of skin is about 0.9.

  The simulation unit 2 can calculate the movement path of photons from the irradiation unit 5 to the light receiving unit 8 by repeating the calculations of the above formulas (6) and (7) every unit time. The simulation unit 2 calculates the movement distance for a plurality of photons. For example, the simulation unit 2 calculates the movement distance for 108 photons.

FIG. 7 is a diagram illustrating the propagation optical path length distribution of each layer calculated by the simulation unit.
In FIG. 7, the horizontal axis represents the elapsed time from the photon irradiation, and the vertical axis represents the logarithmic display of the optical path length.
The simulation unit 2 classifies each movement path of the photons that have reached the light receiving unit 8 for each layer through which the movement path passes. And the simulation part 2 calculates the propagation | transmission optical path length distribution of each layer of skin as shown in FIG. 7 by calculating the average length of the movement path | route of the photon which arrived for every unit time for every classified layer.

FIG. 8 is a diagram illustrating a time-resolved waveform calculated by the simulation unit.
In FIG. 8, the horizontal axis represents the elapsed time from photon irradiation, and the vertical axis represents the number of photons detected by the light receiving unit 8.
The simulation unit 2 calculates the time-resolved waveform of the skin model as shown in FIG. 5 by calculating the number of photons that have reached the light receiving unit 8 per unit time.
Through the processing described above, the simulation unit 2 calculates the propagation optical path length distribution and time-resolved waveform of the skin model for a plurality of wavelengths. At this time, the simulation unit 2 may calculate the propagation optical path length distribution and the time-resolved waveform for a wavelength at which the difference in absorption spectrum of the skin main components (water, protein, lipid, glucose, etc.) becomes large.

FIG. 9 is a diagram showing the light absorption wavelength characteristics of water (by Kubo City, “Introduction to Medical Lasers”, 1st Edition, Ohmsha, published on June 25, 1985, page 70, ISBN4-274. -03065-2).
In FIG. 9, the horizontal axis represents the wavelength of light to be irradiated (μm), the vertical axis represents the penetration depth of the light to be irradiated into the skin (cm), and when light is incident toward water, the light intensity at the time of incidence Shows the penetration depth that advances until it decreases to 1/10 for light of each wavelength in the infrared region.
For example, light having a wavelength band near 3.0 μm has a penetration depth as shallow as about 2 × 10 ⁻³ cm and is easily absorbed by water, and light having a wavelength band of 2.0 μm or less has a penetration depth of about 2 × 10 ⁻³ cm. It can be seen that it is deeper than 10 ⁻² cm and is hardly absorbed by water.

FIG. 10 is a diagram showing an absorption spectrum of the main component of the skin. In FIG. 7, the horizontal axis represents the wavelength of light to be irradiated, and the vertical axis represents the absorption coefficient.
As can be seen from FIG. 10, the absorption coefficient of glucose is maximized when the wavelength is 1600 nm, and the absorption coefficient of water is maximized when the wavelength is 1450 nm.
Therefore, the simulation unit 2 may calculate the propagation optical path length distribution and the time-resolved waveform for a wavelength at which the difference in the absorption spectrum of the main component of the skin becomes large, such as 1450 nm and 1600 nm.

  When the simulation unit 2 calculates the propagation optical path length distribution and the time-resolved waveform of the skin model for a plurality of wavelengths, the simulation unit 2 stores the information on the propagation optical path length distribution in the optical path length distribution storage unit 3 and the time-resolved waveform information as the time-resolved waveform. Store in the storage unit 4.

FIG. 11 is a diagram showing the relationship between the wavelength of light applied to each of the epidermis layer 32, the dermis layer 33 and the subcutaneous tissue 34 of the skin 31 and the absorption coefficient. In the figure, A represents the absorption coefficient of the epidermis layer 32. , B indicates the absorption coefficient of the dermis layer 33, and C indicates the absorption coefficient of the subcutaneous tissue 34, respectively.
According to FIG. 11, it can be seen that the absorption spectrum of the dermal layer 33 has a maximum value near the wavelength of 1450 nm and contains about 60% of the water absorption coefficient. Also, in the absorption spectrum of the skin layer 32, there is a maximum value of about 1/3 the size of the dermis layer 23 in the vicinity of the wavelength of 1450 nm, and it is understood that the moisture content is about 20% of the water absorption coefficient. . On the other hand, in the absorption spectrum of the subcutaneous tissue 34, it can be seen that there is only a maximum value of about 1/10 the size of the dermis layer 23 in the vicinity of the wavelength of 1450 nm, and it contains only about several percent of the water absorption coefficient.

  As described above, the dermis layer 33 contains about 60% of the water absorption coefficient, and the skin layer 32 contains about 20% of the water absorption coefficient. It can be seen that the subcutaneous tissue 34 contains only about several percent of the water absorption coefficient. Therefore, it can be understood that in order to non-invasively measure the blood glucose level from the skin, it is only necessary to select the dermis layer 33 containing glucose as a measurement target and measure the amount of glucose contained in the dermis layer 33.

Incidentally, it is known that the absorption coefficient of water has temperature dependency.
FIG. 12 is a diagram showing the temperature dependence of the water absorbance spectrum, where A is the water absorbance spectrum at 41 ° C., and B is the water absorbance spectrum at 21 ° C.
Here, an optical cell having a cell length of 1 mm is used, a temperature control unit type optical cell holder is used, temperature is adjusted within a range of ± 0.1 ° C. using a constant temperature circulation tank, and an ultraviolet-visible near-infrared spectrophotometer is used. The absorbance spectrum of water at 41 ° C. and 21 ° C. was measured using a total Lambda 900S (manufactured by Perkin Elmer).
According to FIG. 12, the maximum value of the absorbance spectrum of water is near the wavelength of 1450 nm at 21 ° C., and the maximum value shifts to the shorter wavelength side from 1450 nm as the temperature becomes higher than 21 ° C.

FIG. 13 is a diagram showing the temperature dependence of the difference in the absorbance spectrum of water. In the wavelength region of 1000 to 2000 nm, the absorbance spectrum of distilled water at 25 ° C. and 21 ° C. based on the absorbance spectrum of distilled water at 21 ° C. It shows the difference from the absorbance spectrum of distilled water at.
According to FIG. 13, it can be seen that the absorbance spectrum of distilled water has a different temperature effect depending on the wavelength. Therefore, when the glucose concentration in the aqueous glucose solution is determined by absorbance, if the temperature is corrected according to the wavelength used, an accurate measurement value excluding the influence of the absorbance spectrum of water can be obtained.

  As described above, if the amount of glucose contained in the dermis layer 33 is measured using short-time pulsed light, and the amount of glucose is corrected by the temperature of the dermis layer 33, the blood sugar level is accurately measured non-invasively from the skin. be able to.

  FIG. 14 is a diagram showing an example of the absorbance spectrum of an aqueous glucose solution, in which A is the absorbance spectrum of a high-concentration glucose aqueous solution of 9.4 g / dl measured with the reference side as distilled water (21.5 ° C.). B represents a correction value obtained by temperature correction and volume correction of the measurement value of the absorbance spectrum of the glucose aqueous solution.

In this glucose aqueous solution, since the glucose concentration was 9.4 g / dl, which is about 100 times that of a healthy person, the volume increase at this time is about 6%.
At this concentration, the volume ratio of glucose and water is as large as 6: 100 and cannot be ignored. Thus, since the volume of water on the sample side is reduced compared to the volume of water on the reference side, the absorbance spectrum difference is greatly negative in the vicinity of 1400-1500 nm where the absorbance of water is large and in the wavelength region of 1900 nm or more. It has become. This volume reduction corresponds to a reduction of 0.057 mm for a cell length of 1 mm.

The following can be understood from FIG.
For example, at a wavelength of 1600 nm, the absorbance is about 0.035 with a glucose amount of 9.4 g / dl, so the absorption coefficient is about 0.08 / mm. On the other hand, since the normal value of the glucose amount in the dermis layer is about 100 mg / dl, the absorption coefficient corresponding to this normal value is about 0.0008 / mm.

Next, the influence of temperature will be described with reference to FIG.
In FIG. 13, at a wavelength of 1600 nm, the absorbance difference is about −0.008 with respect to an increase of 4 ° C., so the change in absorption coefficient is about −0.02 / mm.
Here, assuming that the absorbance change is linear with respect to the temperature change, when the temperature rises by 1 ° C., the amount of change in the absorption coefficient becomes −0.004 / mm. That is, this 1 ° C. rise corresponds to a decrease in the glucose concentration of 500 mg / dl.
This is a correction value in the result of obtaining the absorbance spectra of a cell length of 1 mm and sample temperatures of 21 ° C. and 41 ° C. using an absorptiometer.

  Note that the change in absorbance due to the glucose concentration in the actual skin is based on the scattering coefficient in the skin, and the results obtained from the absorbance spectra of the cell length of 1 mm and the sample temperatures of 21 ° C. and 41 ° C. were obtained using an absorptiometer. It is about an order of magnitude larger. Therefore, the correction value of the absorption coefficient with respect to the temperature change in the actual skin is also increased by about one digit.

Next, a procedure for measuring a blood sugar level using the blood sugar level measuring apparatus 1 will be described with reference to FIG.
First, the person to be measured applies the blood sugar level measuring device 1 to the skin such as the wrist, and operates the blood sugar level measuring device 1 by pressing a measurement start switch (not shown) or the like.
Here, the light guide 6 is slid in an arbitrary direction along the skin 31 with the tip of the light guide 6 placed on the skin 31 of the human body, and the pulse wave sensor 24 causes pulsation in the human body. The pulse pressure in the vicinity of the site, that is, the pulse pressure of the artery is detected without contact. At the same time, the temperature of the skin 31 in the vicinity of the artery detected by the pulse wave sensor 24 is measured in a non-contact manner by the temperature sensor 25 (step S1).

Next, the body temperature specifying unit 14 detects the pulse pressure data detected by the pulse wave sensor 24 over the surface of the skin 31 including the artery and the surface of the skin 31 including the artery with the temperature sensor 25. Based on the temperature data measured in a non-contact manner, the maximum pulse pressure of the pulse pressure is specified, and the temperature corresponding to the specified maximum pulse pressure is specified as a body temperature that approximates the deep body temperature of the living body. (Step S2).
On the other hand, the irradiation unit 5 irradiates the dermis layer 33 constituting the skin 31 with short-time pulsed light on the skin 31 (step S3).

Next, the light guiding unit 6 collects a plurality of types of backscattered light radiated from the skin 31, that is, backscattered light emitted from the subcutaneous tissue 32, the dermis layer 33, and the epidermis layer 34, and a light scattering medium layer. The light is guided to the selection unit 7.
In the light scattering medium layer selection unit 7, the dermis layer 33 emits the backscattered light emitted from each of the subcutaneous tissue 32, the dermis layer 33, and the epidermis layer 34 collected and guided by the light guide unit 6. Backscattered light is selected (step S4).

Next, the backscattered light per unit time emitted from the dermis layer 33 is received by the light receiving unit 8 (step S5). At this time, the light receiving unit 8 records the received light intensity for each unit time from the start of irradiation (for example, times t _{1 to} t _m every 1 picosecond) in the internal memory.

When the light intensity acquisition unit 9 is notified that the light receiving unit 8 has completed the light reception, the light intensity acquisition unit 9 acquires the received light intensity at different times of the backscattered light emitted from the dermis layer 33 (step S6). ). That is, the light intensity of the backscattered light at each of the plurality of times t _{1 to} t _m is acquired.
Here, it is preferable that the times t ₁ to t _m at which the light intensity acquisition unit 9 acquires the light intensity include the time at which the peak of the backscattered light emitted from the dermis layer 33 is obtained. That is, it is preferable that the time when the light path length of the dermis layer 33 is maximized is added to the time when the irradiation unit 5 irradiates the pulsed light for a short time.

（但し、Ｉ（ｔ）は受光部５が時刻ｔにて受光した光強度、Ｎ（ｔ）は短時間パルス光の時間分解波形のモデルの時刻ｔにおける光強度、Ｌｉ（ｔ）は皮膚の各々の層における伝搬光路長分布のモデルの時刻ｔにおける第ｉ層の光路長、μｉは第ｉ層の光吸収係数である）
から算出する（ステップＳ７）。
ここでは、第１層は表皮層、第２層は真皮層、第３層は皮下組織を示し、μ_１は表皮層の光吸収係数、μ_２は真皮層の光吸収係数、μ_３は皮下組織の光吸収係数を示す。 Next, in the light absorption coefficient calculation unit 12, the received light intensity at different times of the backscattered light emitted from the dermis layer 33 acquired by the light intensity acquisition unit 9, that is, backscattering at each of a plurality of times t _{1 to} t _m. Based on the light intensity of light, the light absorption coefficient of the dermis layer 33 is expressed by the following equation (8).

(Where I (t) is the light intensity received by the light receiving unit 5 at time t, N (t) is the light intensity at time t of the time-resolved waveform model of the short-time pulse light, and Li (t) is the skin (The optical path length of the i-th layer at time t in the model of the propagation optical path length distribution in each layer, μi is the light absorption coefficient of the i-th layer)
(Step S7).
Here, the first layer is the epidermis layer, the second layer is the dermis layer, the third layer is the subcutaneous tissue, μ ₁ is the light absorption coefficient of the epidermis layer, μ ₂ is the light absorption coefficient of the dermis layer, and μ ₃ is the subcutaneous Shows the light absorption coefficient of tissue.

（但し、μａは皮膚の任意の層である第ａ層における光吸収係数、ｇｊは皮膚を構成する第ｊ成分のモル濃度、εｊは第ｊ成分の光吸収係数、ｐは皮膚を構成する主成分の個数、ｑは短時間パルス光の種類数である）
から算出する（ステップＳ８）。 Next, the concentration calculation unit 13 calculates the concentration of glucose contained in the dermis layer 33 based on the light absorption coefficient μ of the dermis layer 33 calculated by the light absorption coefficient calculation unit 12 using the following equation (9).

(Where μa is the light absorption coefficient in the a layer, which is an arbitrary layer of the skin, gj is the molar concentration of the jth component constituting the skin, εj is the light absorption coefficient of the jth component, and p is the main component constituting the skin. The number of components, q is the number of types of short-time pulsed light)
(Step S8).

Next, the concentration correction unit 14 uses the difference between the body temperature approximated to the deep body temperature of the living body specified by the body temperature specifying unit 14 and the reference temperature as the glucose concentration of the dermis layer 33 calculated by the concentration calculation unit 13. The following correction formula:
The measured value of glucose concentration is corrected by a value corresponding to the absorption coefficient of water (step S9).
For example, when the temperature of the dermis layer 33 is increased by T ° C., the amount of change in the light absorption coefficient is −0.004 / mm × T. Therefore, the amount of decrease in glucose concentration when the temperature of the dermis layer 33 is increased by T ° C. is 500 mg / dl × T.

The blood glucose level measuring apparatus 1 described above has a built-in computer system, and the processing operation of each step described above is stored in a computer-readable recording medium in the form of a program. Therefore, the above-described processing operation can be performed by the computer reading and executing this program.
Here, examples of the computer-readable recording medium include a magnetic disk, a magneto-optical disk, a CD-ROM, a DVD-ROM, and a semiconductor memory.
Alternatively, the computer program may be distributed to the computer via a communication line, and the computer that has received the distribution may execute the program.

The program may be for realizing a part of the above steps.
Furthermore, what can implement | achieve the function mentioned above in combination with the program already recorded on the computer system, what is called a difference file (difference program) may be sufficient.

  As described above, according to the present embodiment, the pulse wave sensor 24 detects the pulse pressure of the artery without contact, and at the same time, the temperature sensor 25 measures the temperature of the skin 31 near the artery without contact, Since the glucose concentration in the dermis layer of the skin calculated based on the backscattered light emitted from the dermis layer is corrected based on the body temperature approximated to the deep body temperature of the living body specified by the body temperature specifying unit 14, It is possible to accurately detect the glucose concentration in the dermis layer calculated based on the scattered light according to the activity state of the human body. Therefore, the glucose concentration in the dermis layer can be accurately measured in a short time in a non-invasive manner according to the activity state of the living body.

[Second Embodiment]
FIG. 16: is sectional drawing which shows the outline of a structure of the light guide part of the blood glucose level measuring apparatus (concentration determination apparatus) of the 2nd Embodiment of this invention, and the light guide part of the blood glucose level measuring apparatus 41 of this embodiment 42 differs from the light guide unit 6 of the blood sugar level measuring apparatus 1 of the first embodiment in that the temperature sensor 25, the surface temperature sensor (surface temperature measuring means) 43 that measures the temperature of the surface of the skin 31, and the heat insulation. The temperature of the dermis layer 33 measured by the surface temperature sensor 43 is used instead of the sensor internal temperature measurement sensor (sensor internal temperature measurement means) 44 that is provided in the material 23 and directly measures the temperature of the surface temperature sensor 43 itself. Surface / internal temperature change rate calculation unit (surface / internal temperature change rate calculation) that calculates the difference between the temperature in the vicinity of the surface temperature sensor 43 measured by the sensor internal temperature measurement sensor 44 as a temperature change rate per unit time Means) 45 The pulse wave sensor 24 and the surface / internal temperature change rate calculation unit 45 are connected to the concentration correction unit 15, and the simulation unit 2 to the concentration calculation unit 13 other than the light guide unit 42 are the first. Since it is completely the same as the blood glucose level measuring apparatus 1 of the embodiment, the description is omitted.

Next, a procedure for measuring a blood glucose level using the blood glucose level measuring device 41 will be described with reference to FIG.
First, the person to be measured applies the blood sugar level measuring device 41 to the skin such as the wrist, and operates the blood sugar level measuring device 41 by pressing a measurement start switch (not shown) or the like.
Here, the light guide 42 is slid in an arbitrary direction along the skin 31 while the tip of the light guide 42 is applied to the skin 31 of the human body, and the pulsation in the human body is caused by the pulse wave sensor 24. The pulse pressure in the vicinity of the site, that is, the pulse pressure of the artery is detected without contact (step S11).
Next, the surface temperature sensor 43 measures the surface of the skin 31, that is, the surface temperature just above the artery, and the sensor internal temperature measurement sensor 44 measures the temperature near the surface temperature sensor 43 (step S12).

Next, the difference between the surface temperature immediately above the artery measured by the surface temperature sensor 43 and the temperature near the surface temperature sensor 43 measured by the sensor internal temperature measurement sensor 44 is calculated by the surface / internal temperature change rate calculation unit 45. The temperature change rate per unit time is calculated, and it is determined whether or not the temperature change rate per unit time is within a set value (step S13).
Here, if the rate of temperature change per unit time is within the set value, the irradiating unit 5 irradiates the skin 31 with a short-time pulsed light on the dermis layer 33 constituting the skin 31 (step S14). .
On the other hand, if the rate of temperature change per unit time exceeds the set value, the fact is notified by voice or other notification means, and the arterial pulse pressure is detected again without contact (step S11).

  After irradiating pulsed light for a short time, from the procedure (step S15) of selecting the backscattered light emitted by the dermis layer 33 from the backscattered light emitted from the subcutaneous tissue 32, the dermis layer 33 and the epidermis layer 34, respectively. The procedure up to the procedure (step S19) for calculating the concentration of glucose contained in the dermis layer 33 by the concentration calculator 13 is exactly the same as the procedure (steps S4 to S8) shown in FIG. 15 of the first embodiment.

The concentration correction unit 15 uses the surface temperature directly above the artery measured by the surface temperature sensor 43 for the glucose concentration of the dermis layer 33 calculated by the concentration calculation unit 13, and the following correction formula:
The measured value of glucose concentration is corrected by a value corresponding to the absorption coefficient of water (step S20).
For example, when the temperature of the dermis layer 33 (= surface temperature immediately above the artery) is increased by T ° C., the amount of change in the light absorption coefficient is −0.004 / mm × T. Therefore, the amount of decrease in glucose concentration when the temperature of the dermis layer 33 is increased by T ° C. is 500 mg / dl × T.

  As described above, according to the present embodiment, the pulse wave sensor 24 detects the pulse pressure of the artery in the human body without contact, and then measures the temperature immediately above the artery and the temperature near the surface temperature sensor 43. The rate of change in temperature per unit time calculated from the difference between the temperature immediately above the artery measured by the surface temperature sensor 43 and the temperature in the vicinity of the surface temperature sensor 43 measured by the sensor internal temperature measurement sensor 44 is set. When the value is within the range, the glucose concentration in the dermis layer 33 is corrected using the temperature immediately above the artery measured by the surface temperature sensor 43, so that the glucose concentration in the dermis layer is accurately determined according to the activity state of the human body. It can be detected well. Therefore, the glucose concentration in the dermis layer can be accurately measured in a short time in a non-invasive manner according to the activity state of the living body.

[Third Embodiment]
FIG. 18 is a cross-sectional view schematically showing the configuration of the light guide unit of the blood sugar level measuring device (concentration quantifying device) according to the third embodiment of the present invention, and the light guide unit of the blood sugar level measuring device 51 of the present embodiment. 52 differs from the light guide 6 of the blood sugar level measuring apparatus 1 of the first embodiment in that the pulse wave sensor 24 and the temperature sensor 25 are disposed on the surface of the heat insulating material 23 opposite to the pulse wave sensor 24 and the temperature sensor 25. In the vicinity, a temperature is adjusted to a predetermined temperature, for example, 36.0 ° C. by using a heating means such as a heater, and an internal heat retaining part (temperature adjusting means) 53 for keeping the temperature is provided. Since the simulation unit 2 to the concentration correction unit 15 are exactly the same as those in the blood glucose level measurement apparatus 1 of the first embodiment, description thereof is omitted.

Next, a procedure for measuring a blood glucose level using the blood glucose level measuring device 51 will be described with reference to FIG.
First, the person to be measured places the blood sugar level measuring device 51 on the skin such as the wrist and operates the blood sugar level measuring device 51 by pressing a measurement start switch (not shown) or the like.
Here, the temperature of the vicinity of the temperature sensor 25 is adjusted to a predetermined temperature, for example, 36.0 ° C., by the internal heat retaining unit 53 to retain the temperature (step S21).
Next, the light guide 52 is slid in any direction along the skin 31 with the tip of the light guide 52 placed on the skin 31 of the human body, and the pulse wave sensor 24 causes pulsation in the human body. The pulse pressure in the vicinity of the part, that is, the pulse pressure of the artery is detected without contact. At the same time, the temperature of the skin 31 immediately above the artery detected by the pulse wave sensor 24 is measured by the temperature sensor 25 without contact (step S22).

Next, the body temperature specifying unit 14 detects the pulse pressure data detected by the pulse wave sensor 24 over the surface of the skin 31 including the artery and the surface of the skin 31 including the artery with the temperature sensor 25. Based on the temperature data measured in a non-contact manner, the maximum pulse pressure of the pulse pressure is specified, and the temperature corresponding to the specified maximum pulse pressure is specified as a body temperature that approximates the deep body temperature of the living body. (Step S23).
On the other hand, the irradiation unit 5 irradiates the dermis layer 33 constituting the skin 31 with pulsed light for a short time on the skin 31 (step S24).

  Thereafter, the concentration calculator 13 calculates the backscattered light emitted from the dermis layer 33 from the backscattered light emitted from the subcutaneous tissue 32, the dermis layer 33, and the epidermis layer 34 (step S25). Until the concentration of glucose in the dermis layer 33 is corrected using the temperature of the dermis layer 33 measured by the temperature sensor 24 (step S30), the procedure shown in FIG. 15 of the first embodiment (steps S3 to S9). ) Is exactly the same.

Also in this embodiment, the same effect as the blood sugar level measuring apparatus 1 of the first embodiment can be obtained.
In addition, the internal heat retaining unit 53 adjusts the vicinity of the pulse wave sensor 24 and the temperature sensor 25 to a predetermined temperature and retains the temperature, and then the pulse wave sensor 24 detects the arterial pulse pressure in a non-contact manner. The temperature of the nearby skin 31 is measured in a non-contact manner, and the concentration of glucose in the dermis layer of the skin calculated based on the backscattered light emitted from the dermis layer is determined by the body temperature specifying unit 14. Since the correction is based on the body temperature approximated to the deep body temperature, the temperature fluctuations in the pulse wave sensor 24 and the temperature sensor 25 can be suppressed by adjusting the temperature in the vicinity of the pulse wave sensor 24 and the temperature sensor 25 to a predetermined temperature. The measurement accuracy of the pulse wave sensor 24 and the temperature sensor 25 can be improved.

[Fourth Embodiment]
FIG. 20 is a cross-sectional view schematically illustrating the configuration of the light guide unit of the blood sugar level measuring device (concentration quantifying device) according to the fourth embodiment of the present invention, and the light guide unit of the blood sugar level measuring device 61 of the present embodiment. 62 differs from the light guide part 42 of the blood glucose level measuring device 41 of the second embodiment in that the inner surface of the third embodiment is arranged on the surface of the heat insulating material 23 opposite to the pulse wave sensor 24 and the surface temperature sensor 43. The heat retaining unit (temperature adjusting means) 53 is provided, and the simulation unit 2 to the concentration correction unit 15 other than the light guide unit 62 are exactly the same as the blood sugar level measuring device 41 of the second embodiment. Therefore, the description is omitted.

  The surface temperature sensor 43, the internal temperature sensor 44, and the internal heat retaining portion 53 constitute a deep thermometer using a heat flow compensation method. In this deep thermometer, when sufficient time has passed, the tissues of the epidermis layer 32 and the dermis layer 33 reach thermal equilibrium, and the temperature of the epidermis layer 32 and the temperature of the dermis layer 33 coincide. Therefore, the temperature of the dermis layer 33 can be measured.

Next, a procedure for measuring a blood glucose level using the blood glucose level measuring device 61 will be described with reference to FIG.
First, the person to be measured applies the blood sugar level measuring device 41 to the skin such as the wrist, and operates the blood sugar level measuring device 41 by pressing a measurement start switch (not shown) or the like.
Here, the internal heat retaining unit 53 adjusts the temperature of the vicinity of the pulse wave sensor 24, the surface temperature sensor 43, and the internal temperature sensor 44 to a predetermined temperature, for example, 36.0 ° C., and retains the temperature (step S31).

Next, the light guide 62 is slid in an arbitrary direction along the skin 31 with the tip of the light guide 62 placed on the skin 31 of the human body, and the pulse wave sensor 24 causes pulsation in the human body. The pulse pressure in the vicinity of the part, that is, the pulse pressure of the artery is detected without contact (step S32).
Next, the surface temperature sensor 43 measures the temperature of the skin 31 immediately above the artery, and the sensor internal temperature measurement sensor 44 measures the temperature near the surface temperature sensor 43 (step S33).

Next, the difference between the surface temperature immediately above the artery measured by the surface temperature sensor 43 and the temperature near the surface temperature sensor 43 measured by the sensor internal temperature measurement sensor 44 is calculated by the surface / internal temperature change rate calculation unit 45. The temperature change rate per unit time is calculated, and it is determined whether or not the temperature change rate per unit time is within a set value (step S34).
Here, if the rate of temperature change per unit time is within the set value, the irradiating unit 5 irradiates the dermis layer 33 constituting the skin 31 with the pulsed light for a short time on the skin 31 (step S35). .
On the other hand, if the rate of change in temperature per unit time exceeds the set value, the fact is notified by voice or other notification means, and temperature adjustment and heat retention (step S31) are performed again.

  After irradiating pulsed light for a short time, from the procedure of selecting the backscattered light emitted by the dermis layer 33 from the backscattered light emitted from the subcutaneous tissue 32, the dermis layer 33 and the epidermis layer 34 (step S36), The procedure up to the procedure (step S40) for calculating the concentration of glucose contained in the dermis layer 33 by the concentration calculator 13 is exactly the same as the procedure (steps S15 to S19) shown in FIG. 17 of the second embodiment.

The concentration correction unit 15 uses the surface temperature directly above the artery measured by the surface temperature sensor 43 for the glucose concentration of the dermis layer 33 calculated by the concentration calculation unit 13, and the following correction formula:
The measured value of glucose concentration is corrected by a value corresponding to the absorption coefficient of water (step S41).
For example, when the temperature of the dermis layer 33 (= surface temperature immediately above the artery) is increased by T ° C., the amount of change in the light absorption coefficient is −0.004 / mm × T. Therefore, the amount of decrease in glucose concentration when the temperature of the dermis layer 33 is increased by T ° C. is 500 mg / dl × T.

Also in this embodiment, the same effect as the blood glucose level measuring device 41 of the second embodiment can be obtained.
Moreover, the internal heat retaining unit 53 retains the vicinity of the pulse wave sensor 24, the surface temperature sensor 43, and the sensor internal temperature measurement sensor 44, and then detects the pulse pressure of the artery in a non-contact manner by the pulse wave sensor 24. The surface temperature immediately above the artery and the temperature in the vicinity of the surface temperature sensor 43 are measured. The surface temperature directly above the artery measured by the surface temperature sensor 43 and the temperature in the vicinity of the surface temperature sensor 43 measured by the sensor internal temperature measurement sensor 44 are measured. When the rate of temperature change per unit time calculated from the difference from the temperature is within the set value, the glucose concentration in the dermis layer 33 is corrected using the surface temperature directly above the artery measured by the surface temperature sensor 43. Therefore, the pulse wave sensor 24, the surface temperature sensor 4 and the internal temperature sensor 44 are adjusted to a predetermined temperature and kept warm to maintain the pulse wave sensor 24, the surface temperature sensor 4 and so on. And it is possible to suppress the variation of temperature in the interior temperature measuring sensor 44 each sensor, the pulse wave sensor 24, thereby improving the surface temperature sensor 43 and the sensor internal temperature measuring sensor 44 each measurement accuracy.

[Fifth Embodiment]
FIG. 22 is a schematic block diagram showing a configuration of a blood sugar level measuring device (concentration quantifying device) according to the fifth embodiment of the present invention. The blood sugar level measuring device 71 of the present embodiment is the blood sugar level of the first embodiment. The difference from the measuring apparatus 1 is that the light intensity acquisition unit 9 and the light absorption coefficient calculation unit 12 are different from each other in that the light intensity acquisition unit (light intensity acquisition means) 72 and the light absorption coefficient calculation unit (light absorption coefficient) having different functions. (Calculating means) 73.

The light intensity acquisition unit 72 acquires the light intensity between a predetermined time and at least a predetermined time τ of backscattered light emitted from the dermis layer received by the light receiving unit 8.
The light absorption coefficient calculation unit 73 calculates the light absorption coefficient in the dermis layer of the skin irradiated with the short-time pulse light having the specific wavelength λk.

（但し、Ｉ（ｔ）は受光部５が時刻ｔにて受光した光強度、Ｎ（ｔ）は特定波長λｋの短時間パルス光の時間分解波形のモデルの時刻ｔにおける光強度、Ｌｉ（ｔ）は皮膚の各々の層における伝搬光路長分布のモデルの時刻ｔにおける第ｉ層の光路長、ｎは皮膚の観測対象となる層の数、μｉは第ｉ層の光吸収係数である）
から算出する。
ここで、第１層は表皮層、第２層は真皮層、第３層は皮下組織を示し、μ_１は表皮層の光吸収係数、μ_２は真皮層の光吸収係数、μ_３は皮下組織の光吸収係数を示す。 In this light absorption coefficient calculation unit 73, the light absorption coefficient of an arbitrary layer in the skin is expressed by the following equation (10).

(Where I (t) is the light intensity received by the light receiving unit 5 at time t, N (t) is the light intensity at time t of the model of the time-resolved waveform of the short-time pulse light of the specific wavelength λk, and Li (t ) Is the optical path length of the i-th layer at time t in the model of the propagation optical path length distribution in each layer of the skin, n is the number of layers to be observed on the skin, and μ i is the light absorption coefficient of the i-th layer)
Calculate from
Here, the first layer is the epidermis layer, the second layer is the dermis layer, the third layer is the subcutaneous tissue, μ ₁ is the light absorption coefficient of the epidermis layer, μ ₂ is the light absorption coefficient of the dermis layer, and μ ₃ is the subcutaneous Shows the light absorption coefficient of tissue.

Next, a procedure for measuring a blood glucose level using the blood glucose level measuring device 71 will be described with reference to FIG.
In this procedure, the pulse pressure of the artery is detected by the pulse wave sensor 24 in a non-contact manner, and the temperature of the skin 31 immediately above the artery is measured in a non-contact manner by the temperature sensor 25 (step S51). The procedure up to the procedure for selecting the emitted backscattered light (step S54) is the same as the procedure (steps S1 to S4) shown in FIG.

After selecting the backscattered light, the light receiving unit 8 receives the backscattered light for a predetermined time τ emitted from the dermis layer 33 (step S55). At this time, the light receiving unit 8 records the received light intensity between the start of irradiation and at least a predetermined time τ in the internal memory.
Next, when the light intensity acquisition unit 72 is informed that the light receiving unit 8 has completed the light reception, the light intensity acquisition unit 72 starts at least the predetermined time τ from the start of the backscattered light emitted from the dermis layer 33. The received light intensity during the period is acquired (step S56).

（但し、Ｉ（ｔ）は受光部５が時刻ｔにて受光した光強度、Ｎ（ｔ）は特定波長λｋの短時間パルス光の時間分解波形のモデルの時刻ｔにおける光強度、Ｌｉ（ｔ）は皮膚の各々の層における伝搬光路長分布のモデルの時刻ｔにおける第ｉ層の光路長、ｎは皮膚の観測対象となる層の数、μｉは第ｉ層の光吸収係数である）
から算出する（ステップＳ５７）。 Next, in the light absorption coefficient calculation unit 73, the dermis layer is based on the received light intensity from the start of irradiation of the backscattered light emitted from the dermis layer 33 acquired by the light intensity acquisition unit 72 to at least a predetermined time τ. The light absorption coefficient of 33 is expressed by the following equation (11).

(Where I (t) is the light intensity received by the light receiving unit 5 at time t, N (t) is the light intensity at time t of the model of the time-resolved waveform of the short-time pulse light of the specific wavelength λk, and Li (t ) Is the optical path length of the i-th layer at time t in the model of the propagation optical path length distribution in each layer of the skin, n is the number of layers to be observed on the skin, and μ i is the light absorption coefficient of the i-th layer)
(Step S57).

（但し、μａは皮膚の任意の層である第ａ層における光吸収係数、ｇｊは皮膚を構成する第ｊ成分のモル濃度、εｊは第ｊ成分の光吸収係数、ｐは皮膚を構成する主成分の個数、ｑは短時間パルス光の種類数である）
から算出する（ステップＳ５８）。 Next, the concentration calculator 13 calculates the concentration of glucose contained in the dermis layer 33 based on the light absorption coefficient μ of the dermis layer 33 calculated by the light absorption coefficient calculator 73 by the following equation (12).

(Where μa is the light absorption coefficient in the a layer, which is an arbitrary layer of the skin, gj is the molar concentration of the jth component constituting the skin, εj is the light absorption coefficient of the jth component, and p is the main component constituting the skin. The number of components, q is the number of types of short-time pulsed light)
(Step S58).

Next, the concentration correction unit 15 uses the temperature of the skin 31 immediately above the artery measured by the temperature sensor 25 to determine the glucose concentration of the dermis layer 33 calculated by the concentration calculation unit 13 as follows:
The measured value of glucose concentration is corrected by a value corresponding to the absorption coefficient of water (step S59).
For example, when the temperature of the skin 31 immediately above the artery is increased by T ° C., the amount of change in the light absorption coefficient is −0.004 / mm × T. Therefore, the amount of decrease in glucose concentration when the temperature of the dermis layer 33 is increased by T ° C. is 500 mg / dl × T.

Also in this embodiment, the same effect as the blood sugar level measuring apparatus 1 of the first embodiment can be obtained.
Moreover, the light intensity acquisition unit 72 acquires the light intensity between a predetermined time and at least the predetermined time τ, and the light absorption coefficient calculation unit 73 calculates the light absorption coefficient of the dermis layer 33 from the above equation (2). ), And the glucose concentration of the dermis layer 33 calculated in this way is corrected using the temperature of the skin 31 immediately above the artery measured by the temperature sensor 25, and is calculated based on this backscattered light. The concentration of glucose in the dermis layer to be detected can be accurately detected according to the activity state of the human body. Therefore, the glucose concentration in the dermis layer can be accurately measured in a short time in a non-invasive manner according to the activity state of the living body.

[Sixth Embodiment]
FIG. 24 is a schematic block diagram showing a configuration of a blood sugar level measuring apparatus (concentration quantifying apparatus) according to the sixth embodiment of the present invention. The blood sugar level measuring apparatus 81 of the present embodiment is the blood sugar level of the first embodiment. The difference from the measuring apparatus 1 is that a body motion detection unit (body motion detection means) 82 that detects body motion of the living body, and whether the body motion of the living body detected by the body motion detection unit 82 is within a predetermined range. This is the point that a body motion discriminating section (body motion discriminating means) 83 for discriminating is provided.

The body motion detection unit 82 is configured by a sensor that detects a body motion in the motion of a living body, such as an acceleration sensor.
In the body motion detection unit 82, the detected analog signal is converted into a digital signal by an A / D conversion unit, and the body motion signal converted into the digital signal by an FFT (fast Fourier transform) processing unit is subjected to FFT processing.

The body motion determining unit 83 determines whether the living body is in a resting state or an active (exercising) state based on the body motion signal subjected to the FFT processing.
As this discrimination method, for example, it is determined whether or not the highest amplitude level of the frequency component is equal to or less than a threshold value for the body motion signal after the FFT processing. If the amplitude level exceeds the threshold value, the active state is determined.

Next, a procedure for measuring a blood glucose level using the blood glucose level measuring device 81 will be described with reference to FIG.
First, the person to be measured puts the blood sugar level measuring device 81 on the skin such as the wrist and operates the blood sugar level measuring device 81 by pressing a measurement start switch (not shown) or the like.
Here, the movement of the living body is detected while the body motion detector 82 is applied to the human skin 31 (step 61).

Next, the body motion determination unit 83 determines whether or not the highest amplitude level of the frequency component of the body motion signal detected by the body motion detection unit 82 is equal to or less than a threshold value (step S62). Here, if the amplitude level is equal to or lower than the threshold value, it is determined that the subject is in a resting state, and the light guide unit 6 is arbitrarily placed along the skin 31 while the tip of the light guide unit 6 is applied to the skin 31 of the human body. The pulse wave sensor 24 detects the pulse pressure in the vicinity of the part having pulsation in the human body, that is, the pulse pressure of the artery without contact. At the same time, the temperature of the skin 31 near the artery detected by the pulse wave sensor 24 is measured by the temperature sensor 25 in a non-contact manner (step S63).
On the other hand, if the amplitude level exceeds the threshold value, it is determined that the active state is established, and the movement of the living body is detected again (step 61).

  In this procedure, the pulse pressure of the artery is detected by the pulse wave sensor 24 in a non-contact manner, and the temperature of the skin 31 immediately above the artery is measured in a non-contact manner by the temperature sensor 25 (step S63). The procedure up to correcting the glucose concentration (step S71) is the same as the procedure shown in FIG.

  As described above, according to the present embodiment, the movement of the living body is detected by the body movement detection unit 82, and then the body movement determination unit 83 determines whether the living body is in a resting state. It is possible to accurately grasp the resting state. Therefore, it is possible to accurately detect the concentration of the target component in an arbitrary layer to be observed in a state where the living body is at rest.

As described above, each embodiment of the present invention has been described with reference to the drawings. However, the specific configuration is not limited to the above-described one, and various design changes and the like can be made without departing from the scope of the present invention. Is possible.
For example, in each of the above embodiments, the blood glucose level measuring device as the concentration quantification device, the skin of a human palm as the observation target, glucose as the target component, short-time pulsed light of a specific wavelength as light of a specific wavelength, Although the case where the concentration of glucose contained in the dermis layer of the skin is measured has been described, the present invention is not limited to this, and the concentration determination method is not limited to any layer to be observed formed from a plurality of light scattering medium layers. You may use for the other apparatus which quantifies the density | concentration of the target component in.

In each of the above-described embodiments, the pulse wave sensor 24 and the temperature sensor 25 or the pulse wave sensor 24, the surface temperature sensor 43, and the internal temperature sensor 44 are provided as a set. It is good also as a structure.
For example, the pulse wave sensor 24 and the temperature sensor 25 or the pulse wave sensor 24, the surface temperature sensor 43, and the internal temperature sensor 44 may be provided in a plurality of sets, for example, in a matrix of 2 rows and 2 columns.

DESCRIPTION OF SYMBOLS 1 ... Blood glucose level measuring apparatus (concentration determination apparatus), 3 ... Optical path length distribution storage part (optical path length distribution storage means), 4 ... Time-resolved waveform storage part (time-resolved waveform storage means), 5 ... Irradiation part (irradiation means) 7, light scattering medium layer selection section (light scattering medium layer selection means), 8 light receiving section (light receiving means), 9 light intensity acquisition section (light intensity acquisition means), 10 optical path length acquisition section (optical path length acquisition) Means), 11 ... Non-absorption light intensity acquisition part (light intensity model acquisition means), 12 ... Light absorption coefficient calculation part (light absorption coefficient calculation means), 13 ... Concentration calculation part (concentration calculation means), 14 ... Body temperature specifying part (Body temperature specifying means), 15 ... concentration correction unit (concentration correction means), 24 ... pulse wave sensor (pulse wave detection means), 25 ... temperature sensor (temperature measurement means), 31 ... skin, 33 ... dermis layer (arbitrary) Layer), 41 ... blood glucose level measuring device (concentration quantification device), 43 ... surface temperature sensor (table) Temperature measuring means), 44 ... sensor internal temperature measuring sensor (sensor internal temperature measuring means), 45 ... surface / internal temperature change rate calculating section (surface / internal temperature change rate calculating means), 51 ... blood glucose level measuring device (concentration determination) Devices), 53... Internal heat retaining unit (temperature adjusting means), 61... Blood glucose level measuring device (concentration quantifying device), 71... Blood glucose level measuring device (concentration quantifying device), 72. 73 ... Light absorption coefficient calculating unit (light absorption coefficient calculating means), 81 ... Blood glucose level measuring device (concentration quantifying device), 82 ... Body motion detecting unit (body motion detecting unit), 83 ... Body motion determining unit (body motion) Determination means), S1 to S9, S11 to S20, S21 to S30, S31 to S41, S51 to S59, S61 to S71...

Claims (6)

A concentration quantification device for quantifying the concentration of a target component in an arbitrary layer among observation targets composed of layers of a plurality of light scattering media in a living body,
Pulse wave detection means for detecting a pulse pressure in the vicinity of a part having a pulsation in the living body over a predetermined region;
Surface temperature measuring means for measuring the temperature near the surface of the observation target, sensor internal temperature measuring means for measuring the temperature near the surface temperature measuring means, and temperature near the surface of the observation target measured by the surface temperature measuring means And a surface / internal temperature change rate calculating means for calculating a difference between the temperature in the vicinity of the surface temperature measuring means measured by the sensor internal temperature measuring means as a temperature change rate per unit time. Means,
A heat insulating material covering the periphery of the temperature measuring means;
Irradiating means for irradiating the observation object with light;
A light scattering medium layer selection means for selecting backscattered light emitted from the arbitrary layer from a plurality of types of backscattered light emitted from the observation object by irradiating the light;
A light receiving means for receiving backscattered light emitted from the arbitrary layer;
A light intensity acquisition means for acquiring the intensity of backscattered light emitted from the arbitrary layer received by the light receiving means;
A light absorption coefficient calculating means for calculating a light absorption coefficient of the arbitrary layer based on the light intensity acquired by the light intensity acquiring means;
Based on the light absorption coefficient calculated by the light absorption coefficient calculation means, a concentration calculation means for calculating the concentration of the target component in the arbitrary layer;
If the temperature change rate the surface and internal temperature change rate calculation detection means has calculated set value, based on the temperature in the vicinity of the surface of the surface temperature measuring means and the observation target that has been determined, the density calculating means calculates A concentration quantification device comprising: a concentration correction means for correcting the concentration of the target component. The light is a short-time pulsed light, and
Optical path length distribution storage means for storing a model of the propagation optical path length distribution in each of the layers of the plurality of light scattering media of the short-time pulse light irradiated to the observation object;
Time-resolved waveform storage means for storing a model of the time-resolved waveform of the short-time pulsed light irradiated to the observation object;
An optical path length acquisition unit that acquires an optical path length of each of the layers of the plurality of light scattering media at a predetermined time of the model of the propagation optical path length distribution from the optical path length distribution storage unit;
A light intensity model obtaining means for obtaining the light intensity at the predetermined time of the time-resolved waveform model of the short-time pulsed light from the time-resolved waveform storage means;
The light intensity acquisition means acquires the light intensity at a plurality of times t _{1 to} t _m of the arbitrary layer,
The light absorption coefficient calculating means calculates the light absorption coefficient of the arbitrary layer by the following formula (1):

(Where I (t) is the light intensity received by the light receiving means at time t, N (t) is the light intensity at time t of the model of the time-resolved waveform of the short-time pulse light, and Li (t) is the light intensity 2. An optical path length of an i-th layer at a time t in a model of a propagation optical path length distribution in each layer of a plurality of light scattering media, and μ i is a light absorption coefficient of the i-th layer). concentration Determination device according to. The light is a short-time pulsed light, and
Optical path length distribution storage means for storing a model of the propagation optical path length distribution in each of the layers of the plurality of light scattering media of the short-time pulse light irradiated to the observation object;
Time-resolved waveform storage means for storing a model of the time-resolved waveform of the short-time pulsed light irradiated to the observation object;
An optical path length acquisition unit that acquires an optical path length of each of the layers of the plurality of light scattering media at a predetermined time of the model of the propagation optical path length distribution from the optical path length distribution storage unit;
A light intensity model obtaining means for obtaining the light intensity at the predetermined time of the time-resolved waveform model of the short-time pulsed light from the time-resolved waveform storage means;
The light intensity acquisition means acquires a light intensity between a predetermined time and at least a predetermined time τ,
The light absorption coefficient calculating means calculates the light absorption coefficient of the arbitrary layer by the following equation (2):

(Where I (t) is the light intensity received by the light receiving means at time t, N (t) is the light intensity at time t of the model of the time-resolved waveform of the short-time pulse light, and Li (t) is the light intensity The optical path length of the i-th layer at time t in the model of the propagation optical path length distribution in each layer of the plurality of light scattering media, n is the number of layers to be observed, and μi is the light absorption coefficient of the i-th layer. The concentration quantification apparatus according to claim 1 , wherein The light is a short-time pulse light,
The concentration calculation means calculates the concentration of the target component in the arbitrary layer by the following formula (3)

(Where μa is the light absorption coefficient in the a-th layer which is the arbitrary layer, gj is the molar concentration of the j-th component constituting the observation object, εj is the light absorption coefficient of the j-th component, and p is the observation object. 3. The concentration quantification apparatus according to claim 1 , wherein the concentration quantification apparatus is calculated from the number of main components to be configured, q being the number of types of the short-time pulsed light). The pulse wave detecting means,該脈wave detecting means according to any one of claims 1 to 4, characterized in that provided beating discriminating means for discriminating whether the detected pulse pressure Concentration determination device. Body movement detecting means for detecting body movement of the living body, and body movement determining means for determining whether or not the body movement of the living body detected by the body movement detecting means is within a predetermined range. The concentration quantification apparatus according to any one of claims 1 to 5 .

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