Non-invasive measurement device

Abstract

This non-invasive measurement device comprises: a light source unit 10 that emits excitation light; an irradiation optical system 11 that irradiates skin tissue 2 with the excitation light; a detection optical system 12 that collects Raman-scattered light from the skin tissue 2; a spectroscope 13 that separates the Raman-scattered light; and an analysis device 14 that analyzes a spectroscopic signal obtained as a result of the separation by the spectroscope so as to measure blood sugar level. The irradiation optical system 11 forms the excitation light into minute beam spots in a dot pattern and irradiates the skin tissue 2 obliquely with the excitation light.

Description

Non-invasive measurement devices

[0001] The present invention relates to a non-invasive measurement device.

[0002] While blood glucose measurement requires invasive blood sampling, non-invasive infrared light-based methods exist. Raman spectroscopy blood glucose measurement is a non-invasive method that selectively detects the specific chemical structure of glucose. The measurement principle is to detect light generated by the Raman scattering process, and the glucose concentration is estimated based on the intensity of the light. However, the resulting intensity is very weak, which has been a major obstacle to measuring glucose in interstitial fluid deep in the skin. To extract signals from deep tissue, methods using Raman scattering with excitation light in the 700-1200 nm wavelength range, known as the biological window, have been investigated. In particular, spatially offset Raman spectroscopy (SORS), which can measure deep substances non-invasively, has been investigated to obtain Raman signals from interstitial fluid, which has the same glucose concentration as blood.

[0003] For example, Patent Document 1 discloses spatially offset Raman spectroscopy in which incident radiation is supplied to an entrance region on the surface of a sample, light scattered within the sample is collected from a collection region on the sample surface spaced apart from the entrance region, and the Raman characteristics of the collected light are detected.

[0004] Special Publication No. 2008-523375

[0005] However, even when spatially offset Raman spectroscopy is used to obtain Raman signals from interstitial fluid with the same glucose concentration as that in blood, the intensity of the obtained signal is still weak, making it impossible to accurately measure blood glucose levels from the Raman signal.

[0006] The present invention has been made in view of the above, and has an object to accurately measure blood glucose levels from Raman signals.

[0007] The non-invasive measurement device according to the present invention comprises a light source unit that emits excitation light, an irradiation optical system that irradiates skin tissue with the excitation light, a detection optical system that collects Raman scattered light from the skin tissue, a spectroscope that disperses the Raman scattered light, and an analysis device that measures blood glucose levels by analyzing the spectroscopic signals dispersed by the spectroscope, wherein the irradiation optical system forms the excitation light into a minute, point-like beam spot and irradiates the excitation light at an angle relative to the skin tissue.

[0008] According to the present invention, by forming excitation light into a minute point-like beam spot and irradiating the skin tissue with the excitation light at an angle, it is possible to accurately measure blood sugar levels from Raman signals.

[0009] FIG. 1 is a diagram for explaining the configuration of a non-invasive blood glucose measuring device according to an embodiment of the present invention; FIG. 2 is a diagram showing the trajectory of returned light on a cross section of a sample when excitation light irradiated from a non-invasive blood glucose measuring device according to an embodiment of the present invention is incident thereon; FIG. 3 is a diagram showing returned light on the surface of a sample when excitation light irradiated from a non-invasive blood glucose measuring device according to an embodiment of the present invention is incident thereon; FIG. 4 is a diagram showing the intensity distribution of light within skin tissue when the angle of incidence is changed; FIG. 5 is a diagram showing the amount of light detected by a photodetector relative to the angle of incidence; FIG. 6 is a diagram showing the spread of light detected by a photodetector relative to the angle of incidence; FIG. 7 is a diagram showing the spread of light detected by a photodetector relative to the angle of incidence; and FIG. 8 is a diagram showing another example of the detection optical system of a non-invasive blood glucose measuring device.

[0010] 1 is a diagram showing the configuration of a noninvasive blood glucose measuring device 1. The noninvasive blood glucose measuring device 1 analyzes the concentration of substances contained in skin tissue based on Raman scattered light generated in a sample 2, which is skin tissue, when irradiated with excitation light. The noninvasive blood glucose measuring device 1 includes a light source unit 10, an irradiation optical system 11, a detection optical system 12, a spectrometer 13, and an analyzer 14. The light source unit 10 and the irradiation optical system 11 are arranged so that the irradiation light is incident obliquely on the surface of the sample 2. The detection optical system 12 is also arranged so that the light is collected from a position on the surface of the sample 2 that is a predetermined distance away from the position where the irradiation light is incident.

[0011] The light source unit 10 is a light source that emits excitation light. Because Raman scattered light generated in the skin tissue 2 by irradiation with excitation light tends to be weak, the light source unit 10 is preferably a light source that emits high-intensity excitation light. Furthermore, because the concentration of glucose contained in the sample 2 is calculated based on the wavelength of the excitation light, the light source unit 10 is preferably a light source that emits excitation light of a single wavelength. Examples of such light sources include semiconductor lasers and solid-state lasers. The light source unit 10 may also be an LED (light emitting diode). The wavelength of the excitation light is a single wavelength selected from, for example, 785 nm, 830 nm, 860 nm, and 1064 nm.

[0012] An irradiation optical system 11 is coupled to the light source unit 10. The irradiation optical system 11 includes a lens system that forms excitation light into a minute, point-like beam spot in order to selectively collect Raman scattered light from the dermis layer in the skin tissue 2. The irradiation optical system 11 includes a collimating lens 110, a beam expander 111, a ring-shaped body forming unit 112, a band-pass filter 113, and an objective lens 114.

[0013] The collimator lens 110 prevents the light emitted from the light source unit 10 from spreading, and shapes the light into parallel light with a predetermined beam diameter.

[0014] The beam expander 111 expands the beam diameter of the incident light. By expanding the beam diameter, it is possible to reduce the spot diameter when the beam is focused. Specific examples of the beam expander 111 include a pair of single lenses and a pair of cylindrical lenses. The pair of single lenses is used as a beam expander for adjusting the thickness of the light emitted from the laser block. The pair of cylindrical lenses is used as a beam expander for pre-compensating the ring light to make it a perfect circle on the surface of the sample 2.

[0015] The ring-shaped forming unit 112 is an optical element, such as a ring-shaped light-blocking plate, that blocks the central portion of the incident excitation light and transmits the peripheral portion of the excitation light, thereby forming a ring-shaped beam spot. By forming the ring-shaped ring-shaped light, the spot diameter of the beam spot at the focal position can be made smaller. Note that the ring-shaped forming unit 112 may use an axicon lens instead of a light-blocking plate. The axicon lens forms the incident excitation light into a ring-shaped light by interference.

[0016] The bandpass filter 113 is a filter that selectively transmits light of a specific wavelength. A laser line filter is used as a specific example of the bandpass filter 113. The laser line filter is a narrow bandpass filter with the laser wavelength as its center wavelength, and is used to cut the base of the wavelength band of the laser light to sharpen the spectrum of the laser light and reduce background light.

[0017] The objective lens 114 focuses the ring light formed by the ring forming part 112 to form a minute beam spot on the surface of the sample 2 .

[0018] The detection optical system 12 is an optical system for collecting Raman scattered light generated by irradiation of the sample 2 with excitation light through the irradiation optical system 11, further removing the excitation light, and focusing the light on a spectrometer 13. The detection optical system 12 is disposed at a position on the surface of the sample 2 a predetermined distance away from the position where the excitation light is irradiated onto the sample 2 from the irradiation optical system 11, so as to detect Raman scattered light from the dermis layer in the skin tissue of the sample 2. The detection optical system 12 includes an objective lens 120, an edge filter 121, and a condenser lens 122.

[0019] The objective lens 120 is a single lens that collects Raman scattered light generated in the sample 2. It is preferable to use a lens with little aberration, a high numerical aperture, and a long working distance for the objective lens 120. Reducing aberration allows the collected light to be transmitted efficiently. Furthermore, a high numerical aperture allows more light to be collected. The objective lens 120 is set so that the return light from the shallow layer of the dermis has the highest intensity when the angle of incidence of the excitation light incident on the sample 2 through the irradiation optical system 11 is 45°±5°. To satisfy these conditions, a 1-2 inch aspherical single lens can be used.

[0020] The objective lens 120 simultaneously collects the excitation light and Rayleigh scattered light in addition to the Raman scattered light among the light returned from the sample 2. Therefore, an edge filter 121 is provided at the next stage. The edge filter 121 is a long-pass edge filter that removes the excitation light and Rayleigh scattered light so that they are not transmitted to the spectroscope 13.

[0021] The Raman scattered light separated from the excitation light and the Rayleigh scattered light by the edge filter 121 is focused onto the entrance of the spectrometer 13 using another single lens, the focusing lens 122. It is preferable that the focusing lens 122 has little aberration and a numerical aperture that matches the numerical aperture of the spectrometer 13. This allows all of the transmitted light to be focused onto the entrance of the spectrometer 13.

[0022] The spectrometer 13 has a spectroscopic element that separates the incident light into wavelengths and a photodetector 130. The spectroscopic element separates the Raman scattered light supplied from the sample 2 via the detection optical system 12 into wavelengths and guides them to the light-receiving surface of the photodetector 130. Methods for separating light into wavelengths include, for example, a dispersive spectrometer that utilizes the diffraction of light and a Fourier transform spectrometer that utilizes the coherence of light. A dispersive spectrometer is composed of a collimating mirror, a focusing mirror, and a diffraction grating, and disperses light by utilizing the diffraction and interference caused by the diffraction grating. A Fourier transform spectrometer uses an interferometer to measure the interference waveform of light. The measured interference waveform is Fourier transformed to measure the spectrum of light for each wavelength. The interferometer may be a Michelson interferometer composed of a reference mirror, a sample mirror, and an optical branching filter.

[0023] The photodetector 130 has a plurality of light-receiving elements on its light-receiving surface. Raman scattered light is incident on the light-receiving surface of the photodetector 130. When the photodetector 130 receives the Raman scattered light from the spectroscopic element, the light-receiving elements convert the light of each wavelength into an electrical signal and output a light detection signal indicating the intensity distribution for each wavelength. For example, a photodiode, a CCD (Charge Coupled Device), a CMOS (Complementary Metal Oxide Semiconductor), or the like is used as the photodetector 130. The spectroscopic signal output from the photodetector 130 is input to the analysis device 14.

[0024] The analyzer 14 is a computer such as a personal computer, and includes a processor that processes data according to a control program, a main memory that functions as a work area for the processor, an auxiliary memory for storing data for a long period of time, etc. The analyzer 14 calculates the concentration of glucose contained in the sample 2 based on the spectroscopic signal input from the photodetector 130 of the spectrometer 13.

[0025] The sample 2 is a living body's skin tissue, with an epidermis layer 20 on the surface and a dermis layer 21 below the epidermis layer 20. The epidermis layer 20 is approximately 0.1 to 0.3 mm thick and does not contain nerves or blood vessels. The dermis layer 21 is approximately 1 to 2 mm thick and contains capillaries, nerves, and lymphatic vessels. The dermis layer 21 contains interstitial fluid, which is the bodily fluid between cells. Glucose in the blood diffuses into the interstitial fluid through the capillary walls and is transported from the interstitial fluid to the tissue cells. Blood glucose levels can be determined by measuring the glucose contained in this interstitial fluid. Between the epidermis layer 20 and the dermis layer 21 is a melanin layer 22 that produces melanin pigment.

[0026] In the above configuration, the operation of detecting Raman scattered light from the sample 2 will be described. Figures 2A and 2B show the trajectory of return light 32 caused by scattered light generated inside the sample 2 when excitation light irradiated from the noninvasive blood glucose measuring device 1 is incident on the sample 2. Figure 2A is a diagram showing a cross section of the sample 2, and Figure 2B is a diagram showing the surface of the sample 2.

[0027] In FIG. 2A , excitation light 31 is incident on the sample 2 from the light source unit 10 through the irradiation optical system 11 of the noninvasive blood glucose measuring device 1 at an incident angle θ inclined by θ in the −y direction with respect to the z axis, which is perpendicular to the surface of the sample 2. The excitation light 31 irradiated onto the sample 2 travels inside the sample 2 and generates scattered light due to interactions between the excitation light and substances such as blood and fat in the skin tissue. The scattered light includes Raman scattered light, which is scattered light shifted to a longer wavelength than the wavelength of the excitation light, and Rayleigh scattered light, which has the same wavelength as the excitation light. Raman scattered light has the property of exhibiting unique characteristics for each irradiated substance. In the dermis layer 21, components contained in the interstitial fluid generate both Raman scattered light and Rayleigh scattered light.

[0028] The scattered light travels through the skin tissue while diffusing and returns to the skin surface. In Figure 2A, the arrows indicate the path from the incident position 33 of the excitation light 31 to the point where the scattered light scattered within the skin tissue is emitted to the skin surface as returned light. Here, the distance the scattered light travels within the skin tissue increases depending on the depth of the skin tissue it passes through, and the position where it is emitted from the skin surface as returned light 32 becomes farther away from the incident position 33 of the excitation light. In Figure 2B, the light scattered within the skin tissue returns to a position on the circumference of the skin surface, which is the x-y plane of the skin surface, centered on the incident position 33 of the excitation light 31 onto the skin surface, a distance that corresponds to the depth it has traveled within the skin tissue. Therefore, if the layer of skin tissue through which the scattered light passes is shallow, the position where the returned light 32 is emitted from the skin surface is not very far from the incident position 33. In contrast, when the layer of skin tissue through which the scattered light passes is deep, the position at which the return light 32 leaves the skin surface is farther away from the incident position 33 than when the layer of skin tissue through which the scattered light passes is shallower. Scattered light has unique characteristics depending on the substance contained in the layer through which it passes. Therefore, by changing the detection position of the return light 32, Raman scattered light corresponding to the substance in the layer to be detected can be obtained. In Figures 2A and 2B, scattered light generated by interaction with components contained in the interstitial fluid of the dermis layer 21 is shown. The scattered light generated by interaction with components contained in the interstitial fluid is emitted from a position a distance d away from the incident position 33 of the excitation light 31 in the x-y plane. To detect glucose contained in the interstitial fluid, a detection position 34 by the detection optical system 12 is set to coincide with the emission position of the return light 32 indicated by the solid arrow, and the detection optical system 12 collects the return light 32 as collected light 35.

[0029] The excitation light 31 is irradiated onto the dermis layer 21. However, since the melanin layer 22 is located immediately above the dermis layer 21, the excitation light is also irradiated onto the melanin layer 22. The melanin layer 22 absorbs the excitation light 31 and generates autofluorescence. The autofluorescence from the melanin layer 22 travels through the skin tissue and is emitted from the skin surface. Here, the autofluorescence from the melanin layer 22 is emitted from a position on the circumference of the skin surface, which is a plane perpendicular to the incident direction and centered on the incident position 33 of the excitation light 31 on the skin surface, at a predetermined radius. However, because the melanin layer 22 is located shallower in the skin tissue than the dermis layer 21, the radius of the circumference is shorter than that of scattered light generated by interaction with components contained in interstitial fluid, and the excitation light 31 is emitted from near the incident position 33. Therefore, the detection optical system 12 does not collect the autofluorescence from the melanin layer 22, and the autofluorescence is removed from the collected light 35.

[0030] The excitation light and Rayleigh scattered light are removed from the collected light 35 by the edge filter 121, and only the Raman scattered light is input to the spectroscope 13. The spectroscope 13 separates the supplied Raman scattered light by wavelength, guides it to the light-receiving surface of the photodetector 130, and outputs a photodetection signal indicating the intensity distribution for each wavelength. The signal output from the photodetector 130 is input to the analysis device 14.

[0031] When the analyzer 14 receives the spectrum from the photodetector 130, it performs a determination process to determine whether the waveform is specific to glucose based on the waveform pattern in the spectrum. If the determination process determines that the waveform is glucose, it acquires the glucose peak. The glucose concentration is measured based on the glucose peak in the acquired Raman spectrum.

[0032] Next, FIGS. 3 to 6 show scattered light detected when the incident angle θ of the excitation light 31 incident on the sample 2 from the light source unit 10 through the irradiation optical system 11 is changed.

[0033] FIG. 3 shows the light intensity distribution within the skin tissue of sample 2 when the incident angle θ is changed. In FIG. 3, for incident angles θ of 35°, 45°, and 55°, the light spread on the x-y plane parallel to the skin surface in the epidermis layer at a depth t of 0.1 mm from the skin surface and the light spread on the x-y plane in the dermis layer at a depth t of 1.75 mm from the skin surface are shown. The circular areas shown in each figure represent the areas where light spreads, and the center of the circle indicates high light intensity. Here, the light intensity distribution in the epidermis layer shows that power is concentrated in the center of the circle, with little light spread. In contrast, the light intensity distribution in the dermis layer shows a greater spread of light than in the epidermis layer, with the light being scattered over a wider area.

[0034] 4 shows the amount of light detected by the photodetector 130 as a function of the angle of incidence θ of the excitation light 31 incident on the sample 2, for light rays that have returned after passing through the epidermis and dermis layers of skin tissue. Specifically, for light rays that have returned after passing through the epidermis layer at a depth t of 0.1 mm from the skin surface, light rays that have returned after passing through the dermis layer at a depth t of 0.15 mm from the skin surface, light rays that have returned after passing through the dermis layer at a depth t of 0.25 mm from the skin surface, and light rays that have returned after passing through the dermis layer at a depth t of 1.75 mm from the skin surface, the graph shows the angle of incidence θ on the x-axis and the total number of light rays that reach the photodetector 130 and the power absorbed by the photodetector 130 on the y-axis.

[0035] Regarding the graph of the total number of rays and power of light rays that have passed through the epidermis layer at a depth t of 0.1 mm from the skin surface and returned, the total number of rays and power gradually increase but do not change significantly until the incident angle θ exceeds 50°. However, once the incident angle θ exceeds 50°, the total number of rays and power increase sharply. In other words, once the incident angle θ exceeds 50°, the effect of the returned light from the epidermis layer becomes significant.

[0036] Next, regarding the graph of the total number of rays and power of light rays returning after passing through the dermis layer at a depth t of 0.15 mm from the skin surface, the total number of rays and power increase until the incident angle θ reaches 50°, at which point they reach a peak. This peak is circled on the graph for easy understanding. When the incident angle θ exceeds 50°, the total number of rays and power decrease rapidly. In other words, at an incident angle θ of 50°, a large amount of scattered light from the dermis layer at a depth t of 0.15 mm can be detected.

[0037] Next, regarding the graph of the total number of rays and power of light rays returning after passing through the dermis layer at a depth t of 0.25 mm from the skin surface, the total number of rays and power increase until the incident angle θ reaches 45°, at which point they reach a peak. To make this peak easier to see, the graph is circled as above. When the incident angle θ exceeds 45°, the total number of rays and power decrease rapidly. In other words, at an incident angle θ of 45°, a large amount of scattered light from the dermis layer at a depth t of 0.25 mm can be detected.

[0038] Next, regarding the graph of the total number of rays and power of light rays that have returned after passing through the dermis layer at a depth t of 1.75 mm from the skin surface, there is no significant increase even when the incident angle θ increases, and the total number of rays and power remain roughly the same without any clear peak, and when the incident angle θ exceeds 50°, the total number of rays and power decrease sharply. In other words, there is no clear point at which a large amount of scattered light from the dermis layer at a depth t of 1.75 mm can be detected due to changes in the incident angle θ.

[0039] Next, Fig. 5 shows the spread of light detected by the photodetector 130 relative to the angle of incidence θ of the excitation light 31 incident on the sample 2, for light rays that have returned after passing through the epidermis and dermis layers of skin tissue. Specifically, the graph shows the size of the detected image of light detected by the photodetector 130 when the angle of incidence θ is 35°, 45°, and 55° for light rays that have returned after passing through the epidermis layer at a depth t of 0.1 mm from the skin surface and light rays that have returned after passing through the dermis layer at a depth t of 1.75 mm from the skin surface. In the graph, the x-axis represents the y-coordinate value shown in Figs. 2A and 2B, and the y-axis represents incoherent irradiance.

[0040] Regarding the graph of the size of the detected image for a ray of light returning after passing through the epidermis layer at a depth t of 0.1 mm from the skin surface, the width of the y coordinate value of the detected image, i.e., the size of the detected image, increases as the incident angle θ increases.

[0041] Similarly, in the graph of the size of the detected image for a light beam that has passed through the dermis at a depth t of 1.75 mm from the skin surface and returned, the width of the y coordinate value of the detected image, i.e., the size of the detected image, increases as the incident angle θ increases. However, the intensity of the detected light decreases as the incident angle θ increases.

[0042] Fig. 6 shows the light spread in Fig. 5 with the y-axis displayed logarithmically to confirm the spread of weak light. Fig. 6 shows the light beam returning after passing through the epidermis layer at a depth t of 0.1 mm from the skin surface and the light beam returning after passing through the dermis layer at a depth t of 1.75 mm from the skin surface, with the incident angle θ set to 40°, 45°, and 50°.

[0043] Regarding the graph of the size of the detected image for a ray of light returning after passing through the epidermis layer at a depth t of 0.1 mm from the skin surface, the size of the detected image expands in the -y direction of the y coordinate value, as shown in the circled area.

[0044] In contrast, in the graph of the size of the detected image for a ray of light returning after passing through the dermis layer at a depth t of 1.75 mm from the skin surface, the size of the detected image expands in the +y direction of the y coordinate value, as shown in the circled area.

[0045] That is, if scattered light is detected in the −y direction, which is the tilt direction side of the excitation light 31, it will be significantly affected by the returning light from the epidermis layer. Conversely, if scattered light is detected in the +y direction, which is the opposite side to the tilt direction of the excitation light 31, it will be possible to detect a large amount of scattered light from the dermis layer.

[0046] Furthermore, in the graph of the size of the detected image for the light beam returning after passing through the epidermis layer at a depth t of 0.1 mm from the skin surface, the light intensity drops sharply in the +y direction of the y coordinate value, and the size of the detected image does not expand. Therefore, when scattered light is detected in the +y direction, which is opposite to the tilt direction of the excitation light 31, the influence of the returning light from the epidermis layer is reduced.

[0047] As described above, when the incident angle θ is 45°±5°, it is possible to detect a large amount of scattered light from the dermis layer without being affected by the returning light from the epidermis layer. Furthermore, by detecting the returning light from a position in the direction opposite to the tilt direction of the excitation light 31, it is possible to detect a large amount of scattered light from the dermis layer without being affected by the returning light from the epidermis layer.

[0048] In this embodiment, the detection optical system 12 including the objective lens 120, the edge filter 121, and the condenser lens 122 is used to collect Raman scattered light generated by irradiating skin tissue with excitation light. Because scattered light travels through the skin tissue while diffusing, not all of the returning light necessarily returns from the surface of the skin tissue perpendicular to the surface; some returning light exits obliquely from the surface of the skin tissue. Therefore, returning light exiting from a position on the skin tissue other than the detection position, i.e., returning light that has passed through a layer other than the target layer of the skin tissue, may be collected by the objective lens 120. Furthermore, excitation light or other external light reflected by the surface of the skin tissue may also be collected by the objective lens 120. Therefore, a diffuse reflection prevention wall may be provided in the detection optical system 12 to prevent diffusely reflected light from entering the objective lens 120.

[0049] FIG. 7 shows a configuration in which a diffused reflection prevention wall 123 is provided on the objective lens 120. The diffused reflection prevention wall 123 is provided between the objective lens 120 and a sample 2, which is skin tissue. The diffused reflection prevention wall 123 has a conical shape with a bowl-shaped inner wall surface. An opening 1230 is provided in the portion of the diffused reflection prevention wall 123 facing the objective lens 120, and the periphery of the objective surface of the objective lens 120 is covered by the inner wall of the diffused reflection prevention wall 123 through the opening 1230. The inner wall diameter narrows as it moves away from the objective surface of the objective lens 120, and an opening 1231 is provided at the apex of the cone, which is the center of the inner wall surface. The opening 1231 is positioned so as to face the surface of the sample 2. Return light from the skin tissue enters the objective lens 120 through the opening 1231, which has a diameter smaller than that of the opening 1230. Therefore, scattered light from outside the detection position is less likely to enter the interior of the diffuse reflection intrusion prevention wall 123, preventing it from entering the objective lens 120. Furthermore, the inner wall surface of the diffuse reflection intrusion prevention wall 123 is painted a dark color and has a roughened surface. Therefore, even if scattered light from outside the detection position enters the interior of the diffuse reflection intrusion prevention wall 123, the scattered light is absorbed by the inner wall, reducing the amount of scattered light entering the objective lens 120. Furthermore, the height of the cone of the diffuse reflection intrusion prevention wall 123, which is the distance between the objective lens 120 and the opening 1231, is matched to the focal length of the objective lens 120. When detecting light returning from skin tissue, the detection optical system 12 is positioned so that the tip of the diffuse reflection intrusion prevention wall 123 is in contact with or close enough to contact the surface of the skin tissue. This makes it possible to eliminate, for example, changes in focal length caused by individual differences such as finger thickness when using the skin tissue of a finger as the measurement site, prevent the intrusion of scattered light from outside, and improve detection accuracy.

[0050] It should be noted that the present invention is susceptible to various embodiments and modifications without departing from the broad spirit and scope of the present invention. Furthermore, the above-described embodiment is intended to illustrate one example of the present invention and does not limit the scope of the present invention. The above-described embodiments and modifications can be combined in any manner. Furthermore, even if some of the constituent elements of the embodiment are omitted as necessary, they will still fall within the scope of the technical idea of ​​the present invention.

[0051] The present invention can be widely applied to non-invasive blood glucose measuring devices that acquire Raman signals from interstitial fluid that has the same glucose concentration as that in blood by spatially offset Raman spectroscopy.

[0052] 1 Non-invasive blood glucose measuring device, 2 Sample (skin tissue), 10 Light source unit, 11 Irradiation optical system, 12 Detection optical system, 13 Spectrometer, 14 Analysis device, 20 Epidermis layer, 21 Dermis layer, 22 Melanin layer, 31 Excitation light, 32 Return light, 33 Incident position, 34 Detection position, 35 Collected light, 110 Collimating lens, 111 Beam expander, 112 Ring-forming unit, 113 Band-pass filter, 114 Objective lens, 120 Objective lens, 121 Edge filter, 122 Condenser lens, 123 Diffuse reflection prevention wall, 130 Photodetector, 1230, 1231 Aperture.

Claims

1. A non-invasive measurement device comprising: a light source unit that emits excitation light; an irradiation optical system that irradiates the skin tissue with the excitation light; a detection optical system that collects Raman scattered light from the skin tissue; a spectroscope that disperses the Raman scattered light; and an analyzer that analyzes the spectral signal dispersed by the spectroscope to measure a blood glucose level, wherein the irradiation optical system forms the excitation light into a dot-like minute beam spot and irradiates the skin tissue with the excitation light while being inclined with respect to the skin tissue.

2. The non-invasive measurement device according to claim 1, wherein the detection optical system collects the Raman scattered light from the skin tissue at a position in a direction opposite to the inclination direction of the excitation light at the position where the excitation light is irradiated.

3. The non-invasive measurement device according to claim 1 or 2, wherein an incident angle of the excitation light incident on the skin tissue is 45° ± 5°.

4. The non-invasive measurement device according to any one of claims 1 to 3, wherein the blood glucose level is obtained by measuring glucose contained in the interstitial fluid of the dermis layer from the Raman scattered light.

5. The non-invasive measurement device according to any one of claims 1 to 4, wherein the irradiation optical system includes: an annular formation unit that shields a central portion of the excitation light and transmits the excitation light in a peripheral portion of the central portion to form annular light; and an objective lens that condenses the annular light to form the minute beam spot.

6. The non-invasive measurement device according to claim 5, wherein the annular formation unit is an axicon lens.

7. The non-invasive measurement device according to any one of claims 1 to 6, wherein the detection optical system includes: an objective lens that collects the Raman scattered light; and a stray reflection entry prevention wall that is located between the objective lens and the skin tissue, covers the periphery of the objective surface of the objective lens, and prevents entry of scattered light from the outside into the objective lens.

8. The non-invasive measurement device according to claim 7, wherein the stray reflection entry prevention wall includes: a mortar-shaped inner wall surface whose inner wall diameter becomes narrower as it moves away from the objective surface; and an opening provided at the center of the inner wall surface.

9. The non-invasive measurement device according to claim 8, wherein the inner wall surface is painted dark and the surface is roughened.

10. The non-invasive measurement device according to claim 8 or 9, wherein a distance between the objective lens and the opening is equal to a focal length of the objective lens.

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