JP5674094B2 - Concentration determination apparatus, concentration determination method, and program - Google Patents

Description

  The present invention relates to a concentration quantification apparatus, a concentration quantification method, and a program for non-invasively quantifying the concentration of a target component in an arbitrary layer among observation targets configured by a plurality of light scattering medium layers.

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, in the conventional method of measuring glucose concentration in blood or tissue near the surface of a living body from the amount of absorption of near-infrared light, the temperature change of the absorption coefficient with respect to the near-infrared light of water contained in the blood or living body There is a problem that the rate is high and it is difficult to accurately measure the glucose concentration in blood or in the living body.
For example, if the temperature of the contact portion between the measurement surface of the probe tip of the optical fiber bundle for near-infrared light receiving and emitting and the tissue near the surface of the living body is made constant by using a heater and surface temperature detection means, it is certain. In addition, although the temperature of the contact portion is constant, if the temperature of the living body itself changes, the temperature of the tissue near the surface of the living body also changes, and the absorption coefficient for near infrared light also changes. It is difficult to accurately measure the glucose concentration.

  The present invention has been made to solve the above-described problem, and even when the temperature of the observation target itself, which is an object to be measured such as skin, changes, the temperature change of the observation target Concentration quantification that can measure the concentration of the target component contained in this observation target non-invasively and accurately by correcting the temperature based on the rate of change in the absorption coefficient of water in the near-infrared light An object is to provide an apparatus, a concentration determination method, and a program.

In order to solve the above problems, the present invention employs the following concentration determination apparatus, concentration determination method, and program.
That is, 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 composed of a plurality of layers of light scattering media, and irradiates the observation target with light. Irradiating means for irradiating, and a light scattering medium layer for selecting time-resolved 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 selecting unit; a light receiving unit that receives backscattered light emitted from the arbitrary layer; a temperature measuring unit that measures a temperature of the arbitrary layer among the observation targets; and the arbitrary unit received by the light receiving unit. Light intensity acquisition means for acquiring the intensity of backscattered light emitted from the layer, and light absorption coefficient calculation means for calculating the light absorption coefficient of the arbitrary layer based on the light intensity acquired by the light intensity acquisition means; The light Based on the light absorption coefficient calculated by the yield coefficient calculating means, a concentration calculating means for calculating the concentration of the target component in the arbitrary layer, and the temperature measuring means for calculating the concentration of the target component calculated by the concentration calculating means. Density correction means for correcting based on the temperature measured by the temperature measurement means, the density based on the difference between the temperature of the arbitrary layer measured by the temperature measurement means and the ambient temperature of the temperature measurement means The correcting means corrects the density of the target component.
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 light scattering medium layers,
Irradiation means for irradiating the observation target with light, and a light scattering medium for selecting backscattered light radiated from the arbitrary layer from a plurality of types of backscattered light radiated from the observation target by irradiating the light A layer selecting unit; a light receiving unit that receives backscattered light radiated from the arbitrary layer; a temperature measuring unit that measures a temperature of the arbitrary layer among the observation targets; and the arbitrary light received by the light receiving unit. Light intensity acquisition means for acquiring the intensity of backscattered light emitted from the layer, and light absorption coefficient calculation means for calculating the light absorption coefficient of the arbitrary layer based on the light intensity acquired by the light intensity acquisition means And, 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, and the concentration of the target component calculated by the concentration calculation means, Above A degree measuring unit density correcting means for correcting, based on the temperature measured by, and characterized in that it comprises a.

In the concentration determination apparatus of the present invention, when the backscattered light radiated from an arbitrary layer is received by the light receiving means, the temperature of the arbitrary layer is measured by the temperature measuring means. Further, the concentration of the target component in an arbitrary layer calculated by the concentration calculating unit is corrected based on the temperature measured by the temperature measuring unit by the concentration correcting unit.
In this way, by correcting the concentration of the target component calculated by the concentration calculation means based on the temperature of an arbitrary layer measured by the temperature measurement means, any arbitrary observation target calculated based on the backscattered light is corrected. The influence of temperature on the concentration of the target component in the layer can be reduced. Therefore, the influence of the temperature of an arbitrary layer on the concentration of the target component can be reduced, and the concentration of the target component can be accurately measured noninvasively.

（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無吸収モデルの時刻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 A light intensity model acquisition means for acquiring the light intensity at the predetermined time of the model of the time-resolved waveform of light,
The light intensity acquisition unit acquires the light intensity of the arbitrary layer at a plurality of times t _{1 to} t _m , and the light absorption coefficient calculation unit 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 non-absorption model of the time-resolved waveform of the short-time pulsed light, and Li (t) Is the optical path length of the i-th layer at time t of the propagation optical path length distribution model 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 a layer other than an arbitrary layer as noise, and to reduce the influence of temperature on the concentration of the target component. 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無吸収モデルの時刻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 means for acquiring the light intensity at the predetermined time of the model of the time-resolved waveform of light,
The light intensity acquisition means acquires a temporal change in 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 the arbitrary layer as follows: Formula (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 non-absorption model of the time-resolved waveform of the short-time pulsed light, 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 plurality of light scattering media, n is the number of layers to be observed, μi is the light absorption coefficient of the i-th layer Is)
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 a layer other than an arbitrary layer as noise, and to reduce the influence of temperature on the concentration of the target component. 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 an arbitrary layer using the backscattered light that has been time-resolved, the influence of temperature on the concentration of the target component can be reduced. Therefore, the concentration of the target component can be measured with higher accuracy.

  In the concentration quantification device of the present invention, the temperature measuring means includes a surface temperature measuring means for measuring the temperature near the surface of the observation target, and an internal temperature measuring means for measuring the temperature near the surface temperature measuring means, The concentration correction unit is configured to measure the concentration of the target component calculated by the concentration calculation unit, the temperature near the surface of the observation target measured by the surface temperature measurement unit, and the surface temperature measurement measured by the internal temperature measurement unit. It is a density correction means for correcting based on the difference from the temperature in the vicinity of the means.

In the concentration determination apparatus of the present invention, the temperature in the vicinity of the surface to be observed is measured by the surface temperature measuring means, and the temperature in the vicinity of the surface temperature measuring means is measured by the internal temperature measuring means.
Further, the concentration correction means calculates the concentration of the target component calculated by the concentration calculation means, 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 internal temperature measurement means. It corrects based on the difference.
As described above, the concentration of the target component calculated by the concentration calculating unit is calculated by using the concentration correcting unit to measure the temperature near the surface of the observation target measured by the surface temperature measuring unit and the vicinity of the surface temperature measuring unit measured by the internal temperature measuring unit. By correcting based on the difference from the temperature, the influence of the temperature on the concentration of the target component in any layer to be observed calculated based on the backscattered light can be made extremely small. Therefore, the influence of the temperature of an arbitrary layer on the concentration of the target component can be made extremely small, and the concentration of the target component can be accurately measured noninvasively.

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

In the concentration quantification device 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 internal temperature measuring means. The difference from the temperature is calculated as the rate of temperature change per unit time.
In this way, the concentration of the target component calculated by the concentration calculating means is determined from 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 internal temperature measuring means. By correcting based on the calculated temperature change rate per unit time, the influence of the temperature on the concentration of the target component in an arbitrary observation target layer calculated based on the backscattered light can be extremely reduced. Therefore, the concentration of the target component can be accurately measured noninvasively.

The concentration determination apparatus of the present invention is characterized in that a temperature adjusting means is provided in the temperature measuring means.
In the concentration determination apparatus of the present invention, the temperature measuring means is adjusted to a predetermined temperature, for example, 36.0 ° C. by the temperature adjusting means, and held.
This eliminates temperature fluctuations in the temperature measuring means and improves the measurement accuracy of the temperature measuring means.

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 composed of a plurality of layers of light scattering media, wherein the concentration is determined by the irradiation means. Irradiate light, then measure the temperature of the arbitrary layer among the observation objects by the temperature measurement means, and then radiate from the observation object by irradiating the light by the light scattering medium layer selection means. Time-resolved measurement and selection of backscattered light emitted from the arbitrary layer from a plurality of types of backscattered light, and then receiving backscattered light emitted from the arbitrary layer by the light receiving means; The light intensity acquisition means acquires the intensity of the backscattered light emitted from the arbitrary layer received by the light receiving means, and then the light intensity acquired by the light intensity acquisition means by the light absorption coefficient calculation means. Based on the light absorption coefficient of the arbitrary layer, and then the concentration calculating means calculates the concentration of the target component in the arbitrary layer based on the light absorption coefficient calculated by the light absorption coefficient calculating means. And then correcting the concentration of the target component calculated by the concentration calculation means based on the temperature measured by the temperature measurement means, and measuring the temperature measurement means by the concentration correction means. The density correction unit corrects the density of the target component based on the difference between the temperature of the arbitrary layer and the ambient temperature of the temperature measurement unit.
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 configured by a plurality of light scattering medium layers,
The irradiation unit irradiates the observation target with light, then the temperature measurement unit measures the temperature of the arbitrary layer of the observation target, and then the light scattering medium layer selection unit irradiates the light. Thus, the back scattered light emitted from the arbitrary layer is selected from a plurality of types of back scattered light emitted from the observation target, and then the back scattered light emitted from the arbitrary layer is received by the light receiving means. Then, the light intensity acquisition means acquires the intensity of the backscattered light emitted from the arbitrary layer received by the light receiving means, and then the light intensity acquisition means acquires the light intensity acquisition means. Based on the light intensity, the light absorption coefficient of the arbitrary layer is calculated, and then the concentration calculation unit calculates the light absorption coefficient calculated by the light absorption coefficient calculation unit based on the light absorption coefficient. Calculating the concentration of the components, followed by density correcting means, the concentration of the target component the concentration calculating means is calculated, is corrected based on the temperature measured by said temperature measuring means, characterized in that.

In the concentration determination method of the present invention, the temperature measurement unit measures the temperature of an arbitrary layer of the observation target, and the concentration calculation unit calculates the purpose of the arbitrary layer based on the light absorption coefficient calculated by the light absorption coefficient calculation unit. The component concentration is calculated, and the concentration correction unit corrects the target component concentration calculated by the concentration calculation unit based on the temperature measured by the temperature measurement unit.
In this way, by correcting the concentration of the target component in an arbitrary layer calculated by the concentration calculation unit based on the temperature measured by the temperature measurement unit by the concentration correction unit, the target component in the arbitrary layer to be observed is corrected. Also in the concentration, the influence of the temperature of an arbitrary layer on the concentration can be reduced. Therefore, the influence of the temperature of an arbitrary layer on the concentration of the target component can be reduced, and the concentration of the target component can be measured noninvasively with high accuracy and efficiency.

The program of the present invention is an irradiation procedure for irradiating the observation target with light on a computer of a concentration quantification apparatus that quantifies the concentration of a target component in an arbitrary layer among the observation targets configured by a plurality of light scattering medium layers. , A temperature measurement procedure for measuring the temperature of the arbitrary layer among the observation objects, backscattering radiated from the arbitrary layer from a plurality of types of backscattered light radiated from the observation object by irradiating the light Light scattering medium layer selection procedure for selecting light by time-resolved measurement, light receiving procedure for receiving backscattered light radiated from the arbitrary layer, and the rear radiated from the arbitrary layer obtained in the light receiving procedure Light intensity acquisition procedure for acquiring the intensity of scattered light, light absorption coefficient calculation procedure for calculating the light absorption coefficient of the arbitrary layer based on the light intensity acquired in the light intensity acquisition procedure, calculation of the light absorption coefficient Based on the light absorption coefficient calculated in order, a concentration calculation procedure for calculating the concentration of the target component in the arbitrary layer, and the concentration of the target component calculated by the concentration calculation procedure can be obtained by the temperature measurement procedure. A density correction procedure for correcting the temperature based on the temperature, and the temperature of the arbitrary layer measured by the temperature measurement procedure and the surroundings of the means for measuring the temperature of the arbitrary layer by the temperature measurement procedure. Based on the difference between the temperature and the temperature, the density correction procedure corrects the density of the target component.
The program of the present invention is an irradiation procedure for irradiating the observation target with light on a computer of a concentration quantification apparatus that quantifies the concentration of a target component in an arbitrary layer among the observation targets configured by a plurality of light scattering medium layers. , A temperature measurement procedure for measuring the temperature of the arbitrary layer among the observation objects, backscattering radiated from the arbitrary layer from a plurality of types of backscattered light radiated from the observation object by irradiating the light Light scattering medium layer selection procedure for selecting light, light receiving procedure for receiving backscattered light emitted from the arbitrary layer, intensity of backscattered light emitted from the arbitrary layer obtained in the light receiving procedure Based on the light intensity acquisition procedure to be acquired, the light intensity acquired in the light intensity acquisition procedure, the light absorption coefficient calculation procedure for calculating the light absorption coefficient of the arbitrary layer, and the light absorption coefficient calculation procedure. Based on the absorption coefficient, the concentration calculation procedure for calculating the concentration of the target component in the arbitrary layer, the concentration of the target component calculated by the concentration calculation procedure is based on the temperature obtained by the temperature measurement procedure And a density correction procedure for correcting the density.

In the program of the present invention, the temperature measurement procedure for measuring the temperature of an arbitrary layer of the observation target and the concentration of the target component calculated by the concentration calculation procedure were obtained by the temperature measurement procedure in the computer of the concentration determination device. A density correction procedure for correcting based on the temperature is executed.
In this way, a temperature measurement procedure for measuring the temperature of any layer of the observation target, and a concentration correction procedure for correcting the concentration of the target component calculated by the concentration calculation procedure based on the temperature obtained by the temperature measurement procedure , The influence of temperature in the backscattered light emitted from any layer can be reduced, and the concentration of the target component in any layer to be observed calculated based on this backscattered light can be reduced. In addition, the influence of the temperature of any layer can be reduced. Therefore, the influence of the temperature of an arbitrary layer on the concentration of the target component can be reduced, and the concentration of the target component can be measured noninvasively with high accuracy and efficiency.

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 schematic diagram which shows the cross-sectional structure of skin. 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.

An embodiment for carrying out a concentration determination apparatus, a concentration determination method, and a program according to the present invention will be described.
In the present invention, a blood glucose level measuring device will be described as an example of a concentration determination device, the skin of a human palm as an observation target, glucose as a target component, and short-time pulsed light of a specific wavelength as light of a specific wavelength will be described as examples. .

[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 device 1 is a device that non-invasively quantifies the concentration of glucose (target component) contained in a dermis layer (arbitrary layer) of a plurality of layers constituting skin (observation target) such as a palm. Yes, 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, an irradiation unit (irradiation unit) 5, and 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 unit). ) 10, a non-absorption light intensity acquisition unit (light intensity model acquisition unit) 11, a light absorption coefficient calculation unit (light absorption coefficient calculation unit) 12, a concentration calculation unit (concentration calculation unit) 13, and a density correction unit ( Density correction means) 14.

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 sensor 23 provided on the outer side of the light receiving light guide 22 and a temperature sensor (temperature measurement) that measures the temperature of the dermis layer among a plurality of layers that are provided on the skin side surface of the heat insulator 23 and constitute the skin 31. Means) 24 and a base 25 for fixing the irradiation light guide 21, the light receiving light guide 22, and the heat insulating material 23.

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 a material having a heat insulating property with a sufficiently large heat capacity within a range not affecting 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. With this heat insulating material 23, by making the heat capacity sufficiently small so as not to affect the temperature change of the skin 31, the thermal response time until it reaches 90% of the temperature arrival value can be suppressed within 0.2 seconds.
The temperature sensor 24 measures the temperature of the dermis layer at a depth of 0.3 mm to 1.5 mm from the surface of the skin 31 in a non-contact manner.

In the light guide section 6, the irradiation light guide path 21 guides the short-time pulse light irradiated by the irradiation section 5 and irradiates the skin 31. In this case, a plurality of types of backscattered light are emitted from the skin 31 by irradiating the skin 31 with the short-time pulsed 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. After the irradiation with the short-time pulse light, the temperature sensor 24 measures the temperature of the dermis layer among the plurality of layers constituting the skin.
Since the temperature sensor 24 is covered with the heat insulating material 23, the temperature sensor 24 is not affected by the temperature of the irradiation light guide path 21 and the light reception light guide path 22, and is near the dermis layer of the plurality of layers constituting the skin. The temperature can be measured.

The light scattering medium layer selection unit 7 selects the backscattered light emitted by the dermis layer from a plurality of 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, 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, 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 wavelengths)
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 concentration correction unit 14 corrects the glucose concentration of the dermis layer calculated by the concentration calculation unit 13 using the temperature of the dermis layer measured by the temperature sensor 24.
The concentration correction unit 14 corrects the glucose concentration of the dermis layer calculated by the concentration calculation unit 13 by using the difference between the temperature of the dermis layer measured by the temperature sensor 24 and the reference temperature. It is possible to suppress the influence of the temperature of the dermis layer on the concentration of glucose in the dermis layer. Thereby, it becomes possible to suppress the influence of the temperature of the dermis layer on the concentration of glucose in the dermis layer, and it is possible to measure the concentration of glucose noninvasively with high accuracy.

In the blood glucose level measuring apparatus 1 configured as described above, the short-time pulse light having a continuous wavelength or a specific wavelength emitted from the irradiation unit 5 is irradiated to the skin 31 via the irradiation light guide 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.
Further, the temperature sensor 24 measures the temperature of the dermis layer of the plurality of layers constituting the skin 31 before and after the collection of the backscattered light.
The light scattering medium layer selection unit 7 selects only the backscattered light emitted from the dermis layer from the multiple types of backscattered light emitted from the skin 31. The light receiving unit 8 receives only the backscattered light 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.

Next, the concentration correction unit 14 corrects the glucose concentration of the dermis layer calculated by the concentration calculation unit 13 using the temperature of the dermis layer measured by the temperature sensor 24.
Thus, by correcting the dermal layer glucose concentration at an arbitrary temperature using the dermal layer temperature measured by the temperature sensor 24, the influence of the temperature on the dermal layer glucose concentration can be reduced. it can.

  As described above, the concentration of glucose contained in the dermis layer calculated based on the backscattered light radiated from the dermis layer is corrected using the temperature of the dermis layer measured by the temperature sensor 24. The influence of the temperature of the dermis layer on the concentration of glucose contained can be suppressed. Therefore, the glucose concentration contained in the dermis layer can be accurately measured noninvasively.

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. 3 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 has 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. 4 is a diagram illustrating the propagation optical path length distribution of each layer calculated by the simulation unit.
In FIG. 4, the horizontal axis represents the elapsed time from the photon irradiation, and the vertical axis represents the logarithm 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. 4 by calculating the average length of the movement path | route of the photon which arrived for every unit time for every classified layer.

FIG. 5 is a diagram illustrating a time-resolved waveform calculated by the simulation unit.
In FIG. 5, 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. 6 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. 6, the horizontal axis represents the wavelength of light (μm), the vertical axis represents the penetration depth of the light to the skin (cm), and the light intensity at the time of incidence when entering the water. 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. 7 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.
FIG. 7 shows that 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 wavelengths where the difference in the absorption spectrum of the main component of the skin becomes large, such as 1400 nm, 1450 nm, 1500 nm, 1600 nm, 1680 nm, 1720 nm, and 1740 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. 8 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 this figure, the absorption spectrum of the dermis layer 33 has a maximum value near the wavelength of 1450 nm, and its absorption coefficient value is about 60% of the absorption coefficient of water, so that 60% of the absorption of the dermis layer 33 is It is thought to be due to moisture. 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 near the wavelength of 1450 nm, and the absorption coefficient value is about 20% of the absorption coefficient of water. It is believed that 20% of the absorption of layer 32 is due to moisture. On the other hand, in the absorption spectrum of the subcutaneous tissue 34, there is only a maximum value of about 1/10 the size of the dermis layer 23 near the wavelength of 1450 nm, and the absorption coefficient value is about several percent of the absorption coefficient of water. It is thought that about several percent of absorption of 34 is due to moisture.

  From the above, 60% of the absorption of the dermis layer 33 is considered to be due to moisture, and 20% of the absorption of the epidermis layer 32 is considered to be due to moisture. It is thought to be due to moisture. 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. 9 is a diagram showing the temperature dependence of the water absorbance spectrum, in which 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. 9, it can be seen that 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. 10 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. 10, the absorbance spectrum of distilled water has different temperature effects 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. 11 is a diagram showing an example of the absorbance spectrum of an aqueous glucose solution. In the figure, A is the absorbance spectrum of a high-concentration glucose aqueous solution of 9.4 g / dl measured using 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. 10, 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.

  In addition, the absorbance change accompanying the glucose concentration in the actual skin is obtained by extending the optical path length due to scattering in the skin, and using an absorptiometer, obtain the absorbance spectra for a cell length of 1 mm and sample temperatures of 21 ° C. and 41 ° C., respectively. The result 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 temperature of the dermis layer 33 constituting the skin 31 is measured by the temperature sensor 24 (step S1).
On the other hand, the irradiation unit 5 irradiates the dermis layer 33 constituting the skin 31 with the pulsed light for a short time on the skin 31 (step S2).

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 S3).

Next, the light receiving unit 8 receives backscattered light emitted from the dermis layer 33 per unit time (step S4). 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 notifies the light intensity acquisition unit 9 that the light reception unit 8 has completed 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 S5). ). 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（ステップＳ６）。
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 non-absorption model of the time-resolved waveform of the short-time pulse light, 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, μ i is the light absorption coefficient of the i-th layer)
(Step S6).
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, 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 12.
Here, for example, a case where the concentration of glucose is calculated from the relationship between the light absorption coefficient of the main component at a single wavelength and the light absorption coefficient of each layer will be described.
Equation (9) below shows that the light absorption coefficient of the skin is the sum of the product of the function and the coefficient relating to the wavelength of each of water, glucose, protein, lipid and other components.

Equation (10) shows the light absorption coefficient of each layer of the skin epidermis, dermis, and subcutaneous tissue from the light absorption coefficient of the skin shown in Equation (9), and the wavelength of water, glucose, protein, lipid and other components. This is the sum of the product of the light absorption coefficient and the concentration.

Formula (11) is a formula representing the above formula (10) as a matrix, and in formula (11), “C” represents water, glucose, protein for each layer of the skin epidermis, dermis, and subcutaneous tissue, It is a coefficient matrix which shows the density | concentration of each of a lipid and another component.

By transforming Equation (11), Equation (12), which is a simultaneous linear equation for solving “C”, can be derived.

Expression (13) is an expression in which Expression (12) is expressed as a matrix.

Here, with respect to the light absorption coefficient of the dermis (L2), a linear simultaneous equation represented by Expression (14) is established for the relationship between the skin principal component absorption coefficient and its concentration at a plurality of wavelengths (λ).

This equation (14) can be transformed to the equation (15).

このように、真皮（Ｌ２）に係わる係数行列「Ｃ」は、式（１６）から算出することができる。 By transforming Equation (15), Equation (16) that is a simultaneous linear equation for solving “C” can be derived.
In equation (16), ε _W is the light absorption coefficient of water, ε _G is the light absorption coefficient of glucose, ε _P is the light absorption coefficient of the protein, ε _L is the light absorption coefficient of the lipid, and μ _L1 is the light absorption coefficient of the epidermis. , mu _L2 light absorption coefficient of the dermis, the mu _L3 is a light absorption coefficient of the subcutaneous tissue. The light absorption coefficient of the skin main component and the skin layer may be further added.

Thus, the coefficient matrix “C” related to the dermis (L2) can be calculated from the equation (16).

（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Ｓ７）。 Furthermore, here, the explanation is based on the difference at a plurality of wavelengths.
Here, the concentration of glucose contained in the dermis layer 33 is expressed by the following equation (17).

(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 S7).

Next, the concentration correction unit 14 uses the temperature of the dermis layer 33 measured by the temperature sensor 24 to calculate 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 S8).
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, when the skin is irradiated with pulsed light for a short time, the temperature of the dermis layer is measured, and the skin is calculated based on the backscattered light emitted from the dermis layer. Since the glucose concentration in the dermis layer is corrected based on the measured temperature of the dermis layer, the influence of the temperature on the glucose concentration in the dermis layer calculated based on the backscattered light can be reduced. Therefore, the influence of the temperature of the dermis layer on the glucose concentration can be reduced, and the glucose concentration can be accurately measured noninvasively.

[Second Embodiment]
FIG. 13: 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 6 of the blood sugar level measuring apparatus 1 of the first embodiment in that the temperature sensor 24, the surface temperature sensor (surface temperature measuring means) 43 that measures the temperature of the surface of the skin 31, and the heat insulation. Instead of an internal temperature sensor (internal temperature measuring means) 44 that is provided in the material 23 and directly measures the temperature of the surface temperature sensor 43 itself, the temperature of the dermis layer 33 measured by the surface temperature sensor 43 and the internal temperature are also measured. A surface / internal temperature change rate calculating unit (surface / internal temperature change rate calculating means) 45 is provided for calculating a difference between the temperature near the surface temperature sensor 43 measured by the sensor 44 as a temperature change rate per unit time. , Surface and internal temperature change The calculation unit 45 is connected to the concentration correction unit 14, and the simulation unit 2 to the concentration correction unit 14, which are components other than the light guide unit 42, are exactly the same as the blood glucose level measurement device 1 of the first embodiment. Therefore, 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 surface temperature sensor 43 measures the vicinity of the surface of the skin 31, that is, the temperature of the dermis layer 33, and the internal temperature sensor 44 measures the temperature near the surface temperature sensor 43 (step S11).

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 S12).
Thereafter, from the procedure (step S13) of selecting 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, the concentration calculation unit 13 performs the dermis. The procedure (step S17) for calculating the concentration of glucose contained in the layer 33 is exactly the same as the procedure (steps S3 to S7) shown in FIG. 12 of the first embodiment.

Next, the difference between the temperature of the dermis layer 33 measured by the surface temperature sensor 43 and the temperature in the vicinity of the surface temperature sensor 43 measured by the internal temperature sensor 44 by the surface / internal temperature change rate calculation unit 45 is expressed as a unit. The temperature change rate per time is calculated, and it is determined whether or not the temperature change rate per unit time is within a set value (step S18).
Here, if the rate of temperature change per unit time is within the set value, the temperature correction by the density correction unit 14 as the next procedure is performed. If the rate exceeds the set value, a notification means such as sound is used. Then, the surface temperature sensor 43 measures the temperature near the surface of the skin 31, that is, the temperature of the dermis layer 33, and the internal temperature sensor 44 measures the temperature near the surface temperature sensor 43 (step S11).

The concentration correction unit 14 uses the temperature of the dermis layer 33 measured by the surface temperature sensor 43 to calculate 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 S19).
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.

  As described above, according to the present embodiment, when the skin is irradiated with pulsed light for a short time, the temperature of the dermis layer 33 and the temperature in the vicinity of the surface temperature sensor 43 are measured and measured by these surface temperature sensors 43. When the temperature change rate per unit time calculated from the difference between the measured temperature of the dermis layer 33 and the temperature in the vicinity of the surface temperature sensor 43 measured by the internal temperature sensor 44 is within a set value, Since the glucose concentration is corrected using the temperature of the dermis layer 33 measured by the surface temperature sensor 43, the influence of the temperature on the glucose concentration in the dermis layer can be reduced. Therefore, the influence of the temperature of the dermis layer on the glucose concentration can be reduced, and the glucose concentration can be accurately measured noninvasively.

[Third Embodiment]
FIG. 15: 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 3rd Embodiment of this invention, and the light guide part of the blood glucose level measuring apparatus 51 of this embodiment 52 differs from the light guide 6 of the blood glucose level measuring apparatus 1 of the first embodiment in that the surface of the heat insulating material 23 opposite to the temperature sensor 24 is placed on the surface near the temperature sensor 24 using a heating means such as a heater. The temperature is adjusted to a predetermined temperature, for example, 36.0 ° C., and an internal heat retaining portion (temperature adjusting means) 53 is provided to retain the temperature. About the simulation unit 2 to the concentration correction unit 14 other than the light guide unit 52 Since it is completely the same as the blood glucose level measuring apparatus 1 of the first embodiment, the 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 24 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 temperature sensor 24 measures the temperature near the surface of the skin 31, that is, the temperature of the dermis layer 33 (step S22).

On the other hand, the irradiation unit 5 irradiates the dermis layer 33 constituting the skin 31 with short-time pulse light on the skin 31 (step S23).
Thereafter, the concentration calculation unit 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, respectively (step S24). 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 S29), the procedure (steps S3 to S8) shown in FIG. 12 of the first embodiment is performed. ) Is exactly the same.

Next, the concentration correction unit 14 uses the temperature of the dermis layer 33 measured by the temperature sensor 24 to calculate the glucose concentration of the dermis layer 33 as follows.
The measured value of glucose concentration is corrected by a value corresponding to the absorption coefficient of water (step S29).
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.

  As described above, according to the present embodiment, the temperature of the dermis layer 33 is measured by the temperature sensor 24 after the temperature of the vicinity of the temperature sensor 24 is adjusted to a predetermined temperature by the internal heat retaining unit 53 and the temperature sensor 24 measures the temperature. By adjusting the vicinity to a predetermined temperature and keeping the temperature, fluctuations in temperature in the temperature sensor 24 can be suppressed, and the measurement accuracy of the temperature sensor 24 can be improved.

[Fourth Embodiment]
FIG. 17: 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 4th Embodiment of this invention, and the light guide part of the blood glucose level measuring apparatus 61 of this 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 heat retaining part (temperature adjustment of the third embodiment) is provided on the surface of the heat insulating material 23 opposite to the surface temperature sensor 43. Means) 53 is provided, and the simulation unit 2 to the concentration correction unit 14 other than the light guide unit 62 are exactly the same as the blood glucose level measurement apparatuses 1 and 41 of the first and second embodiments. 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 temperature of the surface temperature sensor 43 and the vicinity of the internal temperature sensor 44 is adjusted to a predetermined temperature, for example, 36.0 ° C., by the internal heat retaining unit 53 to retain the temperature (step S31).

Next, the surface temperature sensor 43 measures the temperature near the surface of the skin 31, that is, the temperature of the dermis layer 33, and the internal temperature sensor 44 measures the temperature near the surface temperature sensor 43 (step S32).
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 S33).

  Thereafter, 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 S34), the concentration calculator 13 causes the dermis to perform Up to the procedure (step S38) for calculating the concentration of glucose contained in the layer 33 is exactly the same as the procedure (steps S13 to S17) shown in FIG. 14 of the second embodiment.

Next, the difference between the temperature of the dermis layer 33 measured by the surface temperature sensor 43 and the temperature in the vicinity of the surface temperature sensor 43 measured by the internal temperature sensor 44 by the surface / internal temperature change rate calculation unit 45 is expressed as a unit. The temperature change rate per time is calculated, and it is determined whether or not the temperature change rate per unit time is within a set value (step S39).
Here, if the rate of temperature change per unit time is within the set value, the temperature correction by the density correction unit 14 as the next procedure is performed. If the rate exceeds the set value, a notification means such as sound is used. Then, the surface temperature sensor 43 measures the temperature near the surface of the skin 31, that is, the temperature of the dermis layer 33, and the internal temperature sensor 44 measures the temperature near the surface temperature sensor 43 (step S32).

The concentration correction unit 14 uses the temperature of the dermis layer 33 measured by the surface temperature sensor 43 to calculate 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 S40).
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.

  As described above, according to the present embodiment, when the skin is irradiated with pulsed light for a short time, the internal heat retaining unit 53 adjusts the vicinity of the surface temperature sensor 43 and the internal temperature sensor 44 to a predetermined temperature and retains the temperature. Thereafter, the temperature of the dermis layer 33 and the temperature in the vicinity of the surface temperature sensor 43 are measured. The temperature of the dermis layer 33 measured by the surface temperature sensor 43 and the vicinity of the surface temperature sensor 43 measured by the internal temperature 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 of the dermis layer 33 is corrected using the temperature of the dermis layer 33 measured by the surface temperature sensor 43. Therefore, temperature fluctuations in the temperature sensor 24 can be suppressed, and the influence of temperature on the glucose concentration in the dermis layer can be reduced. Therefore, the influence of the temperature of the dermis layer on the glucose concentration can be reduced, and the glucose concentration can be accurately measured noninvasively.

[Fifth Embodiment]
FIG. 19 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, and 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、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 the light absorption coefficient calculation unit 73, the light absorption coefficient of an arbitrary layer in the skin is expressed by the following equation (18).

(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 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 propagation optical path length distribution model 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 light scattering medium layer selection unit 7 selects 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 S43). Since the procedure is the same as the procedure shown in FIG.

After selecting the backscattered light, the light receiving unit 8 receives the backscattered light until a predetermined time τ emitted from the dermis layer 33 (step S44). 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 S45).

（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Ｓ４６）。 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 (19).

(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 S46).

（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Ｓ４７）。 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 73 using the following equation (20).

(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 S47).

Next, the concentration correction unit 14 uses the temperature of the dermis layer 33 measured by the temperature sensor 24 to calculate 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 S48).
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.

  Also in the present embodiment, similarly to the first embodiment, temperature fluctuations in the temperature sensor 24 can be suppressed, and the influence of temperature on the glucose concentration in the dermis layer can be reduced. Therefore, the influence of the temperature of the dermis layer on the glucose concentration can be reduced, and the glucose concentration can be accurately measured noninvasively.

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 embodiments described above, the case where the short-time pulsed light with the specific wavelength is used has been described. However, continuous light with the specific wavelength may be used instead of the short-time pulsed light with the specific wavelength.
In this case, the simulation unit 2, the optical path length distribution storage unit 3, the time-resolved waveform storage unit 4, the optical path length acquisition unit 10, and the non-absorption light intensity acquisition unit 11 are not necessary, and temperature correction is possible.
Furthermore, when the concentration quantification apparatus of each of the above embodiments is applied to, for example, a portable concentration measuring apparatus for skin main components, it can be effectively used for examination, diagnosis and treatment of skin diseases.

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 ... Concentration correction part (Concentration correction means), 24 ... temperature sensor (temperature measurement means), 31 ... skin (observation object), 33 ... dermis layer (arbitrary layer), 41 ... blood glucose level measurement device (concentration determination device), 43 ... surface temperature Sensor (surface temperature measuring means), 44... Internal temperature sensor (internal temperature measuring means) 45 ... surface / internal temperature change rate calculating part (surface / internal temperature change rate calculating means), 51 ... blood sugar level measuring device (concentration quantifying device), 53 ... internal heat retaining part (temperature adjusting means), 61 ... blood sugar level measuring Device (concentration quantification device), 71 ... blood glucose level measurement device (concentration quantification device), 72 ... light intensity acquisition unit (light intensity acquisition unit), 73 ... light absorption coefficient calculation unit (light absorption coefficient calculation unit), S1 to S8 , S11 to S19, S21 to S29, S31 to S40, S41 to S48 ... step

Claims (9)

A concentration quantification device for quantifying the concentration of a target component in an arbitrary layer among observation targets composed of a plurality of light scattering medium layers,
Irradiating means for irradiating the observation object with light;
A light scattering medium layer selection means for selecting time-resolved backscattered light emitted from the arbitrary layer from a plurality of types of backscattered light emitted from the observation target by irradiating the light; and
A light receiving means for receiving backscattered light emitted from the arbitrary layer;
Temperature measuring means for measuring the temperature of the arbitrary layer of the observation target;
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;
A concentration correction unit for correcting the concentration of the target component calculated by the concentration calculation unit based on the temperature measured by the temperature measurement unit;
Equipped with a,
A density wherein the density correction means corrects the density of the target component based on the difference between the temperature of the arbitrary layer measured by the temperature measurement means and the ambient temperature of the temperature measurement means A quantitative device. 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 means for acquiring an optical path length of each of the layers of the plurality of light scattering media at the predetermined time of the model of the propagation optical path length distribution from the optical path length distribution storage means;
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 non-absorption model of the time-resolved waveform of the short-time pulsed light, and Li (t) Is the optical path length of the i-th layer at time t of the propagation optical path length distribution model in each layer of the plurality of light scattering media, and μ i is the light absorption coefficient of the i-th layer)
The concentration quantification apparatus according to claim 1, wherein the concentration quantification apparatus is calculated from: 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 means for acquiring an optical path length of each of the layers of the plurality of light scattering media at the predetermined time of the model of the propagation optical path length distribution from the optical path length distribution storage means;
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 non-absorption model of the time-resolved waveform of the short-time pulsed light, 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 plurality of light scattering media, n is the number of layers to be observed, μi is the light absorption coefficient of the i-th layer Is)
The concentration quantification apparatus according to claim 1, wherein the concentration quantification apparatus is calculated from: 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. (The number of constituent main components, q is the number of types of short-time pulsed light)
The concentration quantification apparatus according to claim 1, wherein the concentration quantification apparatus is calculated from: The temperature measuring means comprises surface temperature measuring means for measuring the temperature near the surface of the observation target, and internal temperature measuring means for measuring the temperature near the surface temperature measuring means,
The concentration correction unit is configured to measure the concentration of the target component calculated by the concentration calculation unit, the temperature near the surface of the observation target measured by the surface temperature measurement unit, and the surface temperature measurement measured by the internal temperature measurement unit. 5. The concentration quantifying apparatus according to claim 1, wherein the concentration quantifying device corrects based on a difference from a temperature in the vicinity of the means.   The temperature measurement unit is configured to determine a difference between a temperature near the surface of the observation target measured by the surface temperature measurement unit and a temperature near the surface temperature measurement unit measured by the internal temperature measurement unit, as a temperature per unit time. 6. The concentration quantification apparatus according to claim 5, further comprising surface / internal temperature change rate calculation means for calculating the change rate.   7. The concentration determination apparatus according to claim 1, wherein the temperature measuring means is provided with a temperature adjusting means. A concentration quantification method 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,
By irradiating means, the observation object is irradiated with light,
Next, the temperature measurement means measures the temperature of the arbitrary layer among the observation objects,
Next, the light scattering medium layer selection means selects time-resolved backscattered light emitted from the arbitrary layer from a plurality of types of backscattered light emitted from the observation target by irradiating the light. ,
Next, the light receiving means receives backscattered light emitted from the arbitrary layer,
Next, the light intensity acquisition means acquires the intensity of backscattered light emitted from the arbitrary layer received by the light receiving means,
Next, the light absorption coefficient calculation means calculates the light absorption coefficient of the arbitrary layer based on the light intensity acquired by the light intensity acquisition means,
Next, the concentration calculation means calculates the concentration of the target component in the arbitrary layer based on the light absorption coefficient calculated by the light absorption coefficient calculation means,
Then, the concentration correction unit corrects the concentration of the target component calculated by the concentration calculation unit based on the temperature measured by the temperature measurement unit ,
The concentration quantification, wherein the concentration correction unit corrects the concentration of the target component based on the difference between the temperature of the arbitrary layer measured by the temperature measurement unit and the ambient temperature of the temperature measurement unit. Method. In the computer of the concentration quantification device that quantifies the concentration of the target component in any layer among the observation objects composed of multiple layers of light scattering media,
Irradiation procedure for irradiating the observation object with light,
A temperature measurement procedure for measuring the temperature of the arbitrary layer of the observation target;
A light scattering medium layer selection procedure for selecting time-resolved 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 procedure for receiving backscattered light emitted from the arbitrary layer;
A light intensity acquisition procedure for acquiring the intensity of backscattered light emitted from the arbitrary layer obtained in the light reception procedure;
A light absorption coefficient calculation procedure for calculating a light absorption coefficient of the arbitrary layer based on the light intensity acquired in the light intensity acquisition procedure;
A 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;
A concentration correction procedure for correcting the concentration of the target component calculated by the concentration calculation procedure based on the temperature obtained by the temperature measurement procedure ;
Based on the difference between the temperature of the arbitrary layer measured in the temperature measurement procedure and the ambient temperature of the means for measuring the temperature of the arbitrary layer in the temperature measurement procedure, the object in the concentration correction procedure A program for correcting the concentration of a component .

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