Non-invasive measurement device

Abstract

A non-invasive measurement device according to the present invention comprises: a light source unit 10 which emits excitation light; an irradiation optical system 11 which irradiates skin tissue 2 with the excitation light; a detection optical system 13 which collects Raman scattered light from the skin tissue 2; a spectroscope 14 which disperses the Raman scattered light; and an analysis device 15 which analyzes a spectral signal obtained as a result of the dispersion by the spectroscope 14 and measures a blood glucose level. The irradiation optical system 11 shapes the cross section of the excitation light into a ring and irradiates the skin tissue 2 with the excitation light, and the detection optical system 13 is positioned to detect light returned from a part of the skin tissue 2 that is the central portion of a ring-shaped beam spot irradiated on the skin tissue 2.

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 signal dispersed by the spectroscope, wherein the irradiation optical system forms a cross section of the excitation light into a ring shape and irradiates it onto the skin tissue, and the detection optical system is positioned to detect return light from the skin tissue at the center of the ring-shaped beam spot irradiated onto the skin tissue.

[0008] According to the present invention, the cross section of excitation light is formed into a ring shape and irradiated onto skin tissue, and the return light from the skin tissue is detected at the center of the ring-shaped beam spot irradiated onto the skin tissue, thereby making it possible to accurately measure blood glucose 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 detected when the ring width of an annular beam profile is changed. FIG. 5 is a diagram showing the intensity distribution of light detected when the ring width of an annular beam profile is changed. FIG. 6 is a diagram showing the intensity distribution of light detected when the ring width of an annular beam profile is changed. FIG. 7 is a diagram showing the intensity distribution of light detected when the inner diameter of the ring of annular beam profile is changed. 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 dichroic mirror 12, a detection optical system 13, a spectroscope 14, and an analyzer 15. The light source unit 10 and the irradiation optical system 11 are arranged so that irradiation light is incident perpendicularly on the surface of the sample 2 via the dichroic mirror 12. The detection optical system 13 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 illumination optical system 11 is coupled to the light source unit 10. The illumination optical system 11 includes a lens system that forms the excitation light into a ring-shaped ring-shaped ring beam in order to selectively collect Raman scattered light from the dermis layer in the skin tissue 2. The illumination optical system 11 includes a collimating lens 110, a beam expander 111, a ring-shaped ring forming unit 112, a band-pass filter 113, and a projection 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 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 ring-shaped ring light. Note that the ring forming unit 112 may be an axicon lens instead of a light blocking plate. The axicon lens forms the incident excitation light into ring 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 projection lens 114 is used to project the ring light formed by the ring forming unit 112 onto the surface of the sample 2 .

[0018] The dichroic mirror 12 reflects the excitation light and guides it to the sample 2 , while transmitting the Raman scattered light, which has a wavelength different from that of the excitation light, and directing it to the detection optical system 13 .

[0019] The detection optical system 13 is an optical system that collects Raman scattered light generated by irradiating the sample 2 with excitation light through the irradiation optical system 11, removes the excitation light, and focuses the light on a spectrometer 14. The detection optical system 13 is disposed at a position on the surface of the sample 2 that is 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 13 includes an objective lens 130, an edge filter 131, and a condenser lens 132.

[0020] The objective lens 130 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 as the objective lens 130. By reducing aberration, the collected light can be transmitted efficiently. Furthermore, by increasing the numerical aperture, more light can be collected. To satisfy these conditions, a 1-2 inch aspherical single lens can be used.

[0021] The objective lens 130 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 131 is provided at the next stage. The edge filter 131 is a long-pass edge filter that removes the excitation light and Rayleigh scattered light so that they are not transmitted to the spectroscope 14.

[0022] The Raman scattered light separated from the excitation light and the Rayleigh scattered light by the edge filter 131 is focused onto the entrance of the spectrometer 14 using another single lens, the focusing lens 132. It is preferable that the focusing lens 132 has little aberration and a numerical aperture that matches the numerical aperture of the spectrometer 14. This allows all of the transmitted light to be focused onto the entrance of the spectrometer 14.

[0023] The spectrometer 14 has a spectroscopic element that separates the incident light into wavelengths and a photodetector 140. The spectroscopic element separates the Raman scattered light supplied from the sample 2 via the detection optical system 13 into wavelengths and guides them to the light-receiving surface of the photodetector 140. Methods for separating light into wavelengths include, for example, a dispersive spectrometer that utilizes the diffraction properties 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 light spectrum for each wavelength. The interferometer may be a Michelson interferometer composed of a reference mirror, a sample mirror, and an optical branching filter.

[0024] The photodetector 140 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 140. When the photodetector 140 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 140. The spectroscopic signal output from the photodetector 140 is input to the analyzing device 15.

[0025] The analyzer 15 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 15 calculates the concentration of glucose contained in the sample 2 based on the spectroscopic signal input from the photodetector 140 of the spectrometer 14.

[0026] 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.

[0027] 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.

[0028] In FIG. 2A , excitation light 31 incident on the sample 2 from the light source unit 10 of the noninvasive blood glucose measuring device 1 through the irradiation optical system 11 is incident perpendicular to the surface of the sample 2. Here, the excitation light 31 is ring-shaped light whose cross section is formed by the ring-shaped ring-forming unit 112. Therefore, since a ring-shaped beam spot is formed on the surface of the sample 2, in FIG. 2B , the incident position 33 of the excitation light 31 is ring-shaped, and there are two incident positions 33 in the y-axis direction shown in FIG. 2A . The excitation light 31 irradiated on 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 with the same wavelength as the excitation light. Raman scattered light has the property of revealing unique characteristics of each irradiated substance. In the dermis layer 21, components contained in the interstitial fluid generate Raman scattered light and Rayleigh scattered light.

[0029] 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, it is possible to obtain Raman scattered light corresponding to the substance in the layer to be detected. In Figures 2A and 2B, the solid arrows indicate scattered light generated by interaction with components contained in the interstitial fluid in the dermis layer 21. 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 13 is set to coincide with the emission position of the return light 32 indicated by the solid arrow, and the detection optical system 13 collects the return light 32 as collected light 35.

[0030] As described above, the excitation light 31 has an annular beam profile, and from the ring-shaped incident position 33, return light 32 that has passed through layers at the same depth is emitted toward the center of the ring. Therefore, at the center of the ring, these return lights 32 are superimposed, and return light 32 with a high optical intensity is detected. If the inner diameter of the ring is 0.6 mm or less and the ring width is 0.5 mm or less, the smaller the ring width, the higher the sensitivity. By appropriately setting the ring radius, scattered light from a target layer can be selectively detected.

[0031] 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 13 does not collect the autofluorescence from the melanin layer 22, and the autofluorescence is removed from the collected light 35.

[0032] The edge filter 131 removes the excitation light and Rayleigh scattered light from the collected light 35, and only the Raman scattered light is input to the spectrometer 14. The spectrometer 14 separates the supplied Raman scattered light by wavelength, guides the separated light to the light-receiving surface of the photodetector 140, and outputs a photodetection signal indicating the intensity distribution for each wavelength. The signal output from the photodetector 140 is input to the analyzing device 15.

[0033] When the analyzer 15 receives the spectrum from the photodetector 140, 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.

[0034] Next, FIGS. 3 to 6 show the intensity of scattered light detected when the diameter and ring width of the annular beam profile are changed.

[0035] 3 shows the light intensity distribution detected by the photodetector 140 for light rays returning after passing through the epidermal and dermal layers of skin tissue when the inner diameter of the annular beam profile is kept constant and the ring width is narrowed. The detected light intensity distribution is shown here when the ring width is set to 0.45 mm and when the ring width is set to a narrower value of 0.335 mm. Comparing the two, the detected light intensity 41 from the epidermal layer does not change significantly even when the ring width is narrowed, whereas the detected light intensity 42 from the dermal layer increases as the ring width is narrowed.

[0036] Conversely, the light intensity distribution when the ring width is increased is shown in Figure 4. This figure shows the intensity distribution of detected light when the ring width is set to 0.525 mm. Compared to Figure 3, the detected light intensity 42 from the dermis layer is reduced by increasing the ring width.

[0037] Figure 5 shows the light intensity distribution when the ring width is further narrowed compared to Figure 3. Here, the detected light intensity distribution is shown when the ring width is set to 0.26 mm and 0.12 mm, which are narrower than 0.335 mm. In both cases, the detected light intensity 42 from the dermis layer is higher than when the ring width is 0.335 mm. Furthermore, the detected light intensity 42 from the dermis layer is higher when the ring width is 0.12 mm than when the ring width is 0.26 mm. This shows that the narrower the ring width, the more scattered light from the dermis layer can be obtained.

[0038] Next, Figure 6 shows the light intensity distribution when the ring width is kept constant but the inner diameter of the ring is changed. This figure shows the intensity distribution of detected light when the ring width is kept constant at 0.12 mm and the inner diameter is increased to 1.00 mm and 1.49 mm, compared to the case where the ring width is 0.12 mm and the inner diameter is 0.51 mm shown in Figure 5 . In both cases, the detected light intensity 42 from the dermis layer is reduced compared to when the inner diameter of the ring is 0.51 mm. Furthermore, the detected light intensity 42 from the dermis layer is reduced when the ring width is 1.49 mm compared to when the inner diameter of the ring is 1.00 mm.

[0039] For these reasons, even if the ring width is narrow, if the inner diameter of the ring is large, it becomes difficult to obtain scattered light from the dermis layer. In other words, the ring width and inner diameter of the ring are important factors in obtaining a large amount of scattered light from the dermis layer. Therefore, by setting the ring forming part 112 of the irradiation optical system 11 so as to adjust either or both of the ring width and the inner diameter of the ring, it is possible to obtain a large amount of scattered light from the dermis layer.

[0040] In this embodiment, the detection optical system 13 including the objective lens 130, the edge filter 131, and the condenser lens 132 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 130. Furthermore, excitation light or other external light reflected by the surface of the skin tissue may also be collected by the objective lens 130. Therefore, a diffuse reflection prevention wall may be provided in the detection optical system 13 to prevent diffusely reflected light from entering the objective lens 130.

[0041] FIG. 7 shows a configuration in which an objective lens 130 is provided with a diffused reflection prevention wall 133. Here, the objective lens 130 is provided between the dichroic mirror 12 and the sample 2, which is skin tissue, and the diffused reflection prevention wall 133 is provided between the objective lens 130 and the sample 2. The diffused reflection prevention wall 133 has a conical shape with an inner wall surface formed in a mortar shape. An opening 1330 is provided in the portion of the diffused reflection prevention wall 133 facing the objective lens 130, and the objective surface of the objective lens 130 is covered by the inner wall of the diffused reflection prevention wall 133 through the opening 1330. The inner wall diameter of the inner wall surface narrows as it moves away from the objective surface of the objective lens 130, and an opening 1331 is provided at the apex of the cone, which is the center of the inner wall surface. The opening 1331 is positioned to face the surface of the sample 2. The return light from the skin tissue is incident on the objective lens 130 through the opening 1331, which is smaller in diameter than the opening 1330. Therefore, scattered light from outside other than the detection position is less likely to enter the interior of the diffuse reflection intrusion prevention wall 133, preventing it from entering the objective lens 130. Furthermore, the inner wall of the diffuse reflection intrusion prevention wall 133 is painted dark and has a roughened surface. Therefore, even if scattered light from other than the detection position enters the interior of the diffuse reflection intrusion prevention wall 133, the scattered light is absorbed by the inner wall, reducing the amount of scattered light entering the objective lens 130. Furthermore, the height of the cone of the diffuse reflection intrusion prevention wall 133, which is the distance between the objective lens 130 and the opening 1331, is matched to the focal length of the objective lens 130. When detecting the return light from the skin tissue, the detection optical system 13 is positioned so that the tip of the diffuse reflection intrusion prevention wall 133 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.

[0042] 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.

[0043] 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.

[0044] 1 Non-invasive blood glucose measuring device, 2 Sample (skin tissue), 10 Light source unit, 11 Irradiation optical system, 12 Dichroic mirror, 13 Detection optical system, 14 Spectrometer, 15 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, 41 Detected light intensity from epidermis layer, 42 Detected light intensity from dermis layer, 110 Collimating lens, 111 Beam expander, 112 Ring forming unit, 113 Band-pass filter, 114 Projection lens, 130 Objective lens, 131 Edge filter, 132 Condenser lens, 133 Diffuse reflection prevention wall, 140 Photodetector, 1330, 1331 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 a cross-sectional shape of the excitation light in a ring shape and irradiates the skin tissue, and the detection optical system is disposed at a position to detect return light from the skin tissue at the central portion of the ring-shaped beam spot irradiated on the skin tissue.

2. The non-invasive measurement device according to claim 1, wherein Raman scattered light from the dermis layer of the skin tissue is collected by adjusting at least one of the inner diameter and the ring width of the ring-shaped beam spot.

3. The non-invasive measurement device according to claim 1 or 2, wherein the inner diameter of the ring-shaped beam spot is 0.6 mm or less.

4. The non-invasive measurement device according to any one of claims 1 to 3, wherein the ring width of the ring-shaped beam spot is 0.5 mm or less.

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

6. The non-invasive measurement device according to any one of claims 1 to 5, wherein the irradiation optical system includes an annular formation unit that shields a central portion of the excitation light and transmits excitation light in a peripheral portion of the central portion to form annular light.

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

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

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

10. The non-invasive measurement device according to claim 9, wherein the inner wall surface is painted dark and has a rough surface.

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

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