JP4469977B2 - Teeth optical interference tomography device - Google Patents

Description

The present invention relates to a tooth optical interference tomography apparatus, and is extremely effective when applied to a caries detection apparatus that obtains a tomographic image of a tooth and inspects the characteristics of the tooth.

  An optical coherence tomography method (hereinafter referred to as “OCT method”) is not only invasive to a living body but also has a high resolution, so that it is not only used for tomography of the retina of the eye, but also the retina. Application to tomography of other organs other than those is also attempted (for example, see Non-Patent Document 1 below), and for example, it is considered to detect the characteristics of teeth (for example, Non-Patent Document 2 below). Etc.).

Japanese Patent No. 3471788 Ken Takemi, “Microscopic diagnosis by optical coherence tomography for clinical application”, Optronics, Optronics, Inc., July 10, 2002, No. 247, p.179-183 Edited by Brett E. Bouma et al., Handbook of Optical Coherence Tomography, (USA), Marcel Dekker Inc., 2002, p. 591-612 Yuzo Yoshikuni, “Development Trends of Wavelength Tunable Lasers and Their Expectations for System Applications,” Applied Physics, Japan Society of Applied Physics, 2002, Vol. 71, No. 11, pp. 1362-1366 Togaku Tsuji et al., “High-speed, high-resolution OFDR-OCT using SSG-DBR laser”, Proceedings of the 28th Optical Symposium, Japan Society of Applied Physics, Japan Optical Society, 19 June 2003, p. .39-40

  However, in the optical coherence tomograph apparatus (hereinafter referred to as “OCT apparatus”) using the conventional OCT method as described above, a broadband broadband is required as a light source. For this reason, a super-luminescent diode is used. In addition, since a natural light source of a fiber amplifier is used, the wavelength range available at present is 0.85 μm or 1.31 μm (for example, see p594 of Non-Patent Document 2). However, in order to obtain a deeper penetration distance, measurements at longer wavelengths are required. Therefore, there is a need for a light source that can be easily obtained and emits light in a longer wavelength region, and a method that enables OCT measurement using the light source.

  In addition, the OCT signal becomes weaker as the penetration degree becomes deeper, and in order to enable measurement at a deeper penetration distance, it is necessary to use an OCT method with higher measurement sensitivity.

  Also, for example, when trying to detect the initial caries in the tooth enamel, the microcrystals (apatite hydroxide) that make up the enamel have a birefringence, so a fault using a device that cannot measure polarization properties The image may become unclear, and it is necessary to be an OCT for tomographic imaging that enables OCT measurement with higher sensitivity in a wavelength region longer than that in the past and also enables measurement of polarization characteristics.

  Further, there is a need for a spectroscopic OCT that enables measurement of the composition ratio of constituents, which can provide more accurate diagnostic knowledge than conventional intensity measurement alone.

  Further, in the conventional OCT, since the reference mirror has to be moved mechanically, the measurement speed is limited, and the tooth to be measured moves during the measurement time, resulting in a distortion (artifact) in the tomographic image. (See, for example, p596 of Non-Patent Document 2). For this reason, there has been a need for an OCT method that can realize the above-described performance and also achieve a faster measurement speed improvement.

For this reason, the present invention can detect even a minute initial caries by providing a high-sensitivity, high-speed, and tooth optical interference tomography device capable of distinguishing tissue composition. It is an object of the present invention to provide a dental caries detection device that can perform detailed examination of teeth.

The tooth optical interference tomography device according to the present invention for solving the above-mentioned problem is a variable wavelength light generation means using a variable wavelength light generation device that emits light while discretely switching wavelengths as a light source, A polarization characteristic measuring means for measuring the polarization characteristics of the teeth, wherein the polarization characteristic measuring means controls the polarization direction of the light generated from the variable wavelength light generating means and divides the light into measurement light and reference light. Means, measurement light irradiation means for irradiating the teeth in the oral cavity with the measurement light divided by the main division means, signal light capturing means for capturing the signal light irradiated and reflected on the teeth, and the signal light The signal light captured by the capturing means is separated into two or more polarization direction components, and combined with the reference light divided by the main dividing means, respectively, and the combined signals having different polarization directions Based on the light intensity of Te, along with determining the polarization characteristics of the tooth, characterized in that based on the input information and a calculation control means for controlling the variable wavelength light generating apparatus.

The tooth optical interference tomography device according to the present invention for solving the above-mentioned problem is a variable wavelength light generation means using a variable wavelength light generation device that emits light while discretely switching wavelengths as a light source, A polarization characteristic measuring means for measuring the polarization characteristics of the teeth, wherein the polarization characteristic measuring means divides the light generated from the variable wavelength light generating means into measurement light and reference light, and the main division Measurement light irradiating means for irradiating the teeth in the oral cavity with the measurement light divided by the means, signal light capturing means for capturing the signal light reflected by the teeth, and captured by the signal light capturing means and multiplexing means for multiplexing the divided said reference light by said main dividing means and said signal light, so that the variable wavelength light generating apparatus wavelength region for the purpose of the light to be generated from the inputted based on the information Controls the said variable-wavelength light generating apparatus, based on the intensity of the combined beam in the variable wavelength light generating apparatus light which is generated from the wavelength region and the multiplexing means, the operation control for obtaining the characteristics of the tooth characterized Tei Rukoto and means.

The tooth optical interference tomography device according to the present invention for solving the above-mentioned problem is a variable wavelength light generation means using a variable wavelength light generation device that emits light while discretely switching wavelengths as a light source , Main splitting means for splitting light generated from the variable wavelength light generating means into measurement light and reference light; measurement light irradiating means for irradiating teeth in the oral cavity with the measurement light split by the main splitting means; Signal light capturing means for capturing the signal light irradiated and reflected on the teeth, and combining the signal light captured by the signal light capturing means and the reference light divided by the main dividing means and means, so that said tunable wavelength light generating apparatus wavelength region for the purpose of the light to be generated from, as well as controlling the variable wavelength light generating apparatus based on the input information, from the variable-wavelength light generating apparatus Generated It was based on the intensity of the light multiplexed by a wavelength region and the multiplexing means of light, characterized by Tei Rukoto and an arithmetic control means for determining the characteristics of the tooth.

According to the tooth optical interference tomography apparatus of the present invention, in the above-described tooth optical interference tomography apparatus , the calculation control unit controls the variable wavelength light generation unit so as to generate light in a plurality of different wavelength regions. The characteristic of the tooth is obtained by obtaining the intensity of the light combined by the combining means for each wavelength region .

The tooth optical interference tomography device according to the present invention is the above-described tooth optical interference tomography device , wherein the calculation control means is capable of performing all of the measurement within a measurement time within a predetermined wavelength of the wavelength variable range and the scanning wave number for the emitted light. The variable wavelength light generator is controlled so as to be scanned .

The tooth optical interference tomography device according to the present invention is the above-described tooth optical interference tomography device , wherein the calculation control means sets the variable wavelength region of the measurement light that can be generated from the variable wavelength light generator as a first wavelength region. And the second wavelength region, and the intensity of the signal light Ls1 with respect to the measurement light in the first wavelength region and the signal light with respect to the measurement light in the second wavelength region in the tooth thickness direction The intensity of Ls2 is obtained, the distribution of the light absorption coefficient for each wavelength region is obtained from the intensity distribution of the signal lights Ls1 and Ls2, and the composition of the tooth enamel and the moisture are obtained from the distribution of the light absorption coefficient. The characteristic of the tooth is obtained by obtaining the distribution of the ratio of the abundance per unit volume .

The tooth optical interference tomography device according to the present invention is the above-described tooth optical interference tomography device , wherein the calculation control unit obtains the intensity of the light combined by the combining unit for each wavelength region. Obtaining the light absorption coefficient of the tooth, and based on the light absorption coefficient, the abundance per unit volume of the enamel or dentin composition of the tooth or the moisture per unit volume of the tooth enamel or dentin The characteristic of the tooth is obtained by obtaining the abundance of .

The tooth optical interference tomography apparatus according to the present invention is the above-described tooth optical interference tomography apparatus, wherein the width of the variable range of the wave number of the emitted light is 4.7 × 10 ⁻² μm ⁻¹ or more, and the frequency of the emitted light is The width is 13 GHz or less, the frequency interval of the emitted light is 15 GHz or less, the wave number interval of the emitted light is 3.1 × 10 ⁻⁴ μm ⁻¹ or less, and the time interval of the emitted light is 530 μs or less. The arithmetic and control means controls the variable wavelength light generator so that all wave numbers of incident light are switched discretely .

According to the optical interference tomography device for teeth according to the present invention, a light source that emits light in a longer wavelength region than that required in the past, which is necessary for obtaining a deeper penetration distance, and OCT measurement using the light source are possible. In addition, since the sensitivity of the measurement is increased, the measurement can be performed at a deeper penetration distance.

In addition, since the OCT measuring apparatus is capable of measuring polarization characteristics, a tomographic image can be made clear even in a tooth tissue having a birefringence, such as enamel.
In addition, the OCT apparatus capable of spectroscopic analysis enables measurement of the composition ratio of constituents, and more accurate diagnostic knowledge can be obtained than a conventional OCT apparatus that merely measures intensity.
In addition, since the OCT apparatus realizes a measurement speed faster than that of the conventional OCT apparatus, it is possible to reduce tomographic image distortion (artifact) due to the movement of the tooth to be measured during the measurement time.

  Therefore, according to the caries detection device according to the present invention, even a minute initial caries can be easily detected.

It has been reported that the absorption coefficient of light due to scattering in the teeth is 60 cm ^{−1 at} a wavelength of 632 nm, 15 cm ⁻¹ at 1053 nm for enamel, 280 cm ⁻¹ at 620 nm and 260 cm ⁻¹ at 1053 nm for dentin. (See, for example, p593 of Non-Patent Document 2). In order to reduce the light absorption coefficient due to scattering, longer wavelengths are more advantageous. FIG. 6 shows the wavelength dependence of the absorption coefficient of enamel and water. Enamel is more advantageous at longer wavelengths because its absorption coefficient decreases exponentially with increasing wavelength. In order to enable OCT measurement through a layer having a thickness of 5 mm, the absorption coefficient needs to be 30 cm ⁻¹ or less at the current sensitivity of −120 dB of OCT. In order to allow the dentin to be measured through the enamel having a thickness of 5 mm based on this standard, the wavelength needs to be 0.9 μm or more as can be seen from FIG.

  FIG. 6 shows the absorption coefficient of light with 100% water. Actually, the composition ratio of water in the enamel is 1 to 2%, and in order to compare with the effect of enamel absorption, the water absorption graph of FIG. It is necessary to shorten it. Thus, in the tooth tissue, the influence of water absorption is small. However, there is an initial absorption peak at about 1.45 μm and a strong absorption coefficient peak for water at about 3 μm, and it is advantageous to use this peak to detect water accumulated in the caries. However, since the influence of thermal noise increases as the wavelength becomes smaller, a wavelength of 5 μm or less is desirable.

  There are three known OCT methods: OCDR (Optical Coherence Domain Reflectometer), FD (Frequency Domain), and OFDR (Optical Frequency Domain Reflectometer). . When the intensity of light irradiation on the teeth is limited due to the safety of the living body, the OFDR method is known to be 100 to 1000 times more sensitive than the OCDR method, and several tens of times more sensitive than the FD method. The OFDR method is the best method.

  The OCDR method and the FD method use a low-coherence light source with a wide spectral width, but the OFDR method uses a high-coherence light source with a narrow spectral width and good coherence to scan the oscillation wavelength. Is used to perform OCT measurement (see, for example, Non-Patent Document 4 above). For this reason, if data analysis is performed by dividing the wavelength region, there is an advantage that spectroscopic OCT becomes possible (for example, see Patent Document 1 above).

  As the variable wavelength light generation means of the OFDR method, the wavelength may be changed continuously or discretely, but the wavelength can be switched discretely without changing the wavelength while acquiring one data. With a simple light source, the wavelength dependency of the characteristics can be accurately determined.

In the variable wavelength light generating means for changing the wavelength in a discrete manner, the width of the variable range of the wave number (2π / wavelength) is 4.7 × 10 ⁻² so that the resolution for obtaining the tomographic image of the tooth is at least 80 μm or more. Since μm ⁻¹ or more is necessary and the measurable distance based on the coherence of light is 10 mm or more, the frequency width of the emitted light must be 13 GHz or less, and the distance at which the overlap of the OCT images does not occur is 10 mm or more. In order to be (5mm or more on one side), the wave number interval must be 3.1 × 10 ⁻⁴ μm ⁻¹ or less, and in order to reduce the influence of human movement such as heartbeat, the measurement must be fast. The wave number switching time is preferably 530 μs or less.

  Here, when the wavelength is scanned discretely, the wavelength may be gradually increased as shown in FIG. 12A, or the wavelength may be gradually decreased as shown in FIG. 12B. However, as shown in FIG. 12C, the wavelength may be changed irregularly. In short, all the predetermined wavelengths may be scanned within the measurement time. The “predetermined wavelength” is preferably a set of wave numbers arranged at equal intervals when viewed in terms of wave numbers, but is not necessarily limited to this. For example, calculation processing is performed when creating a tomographic image. By considering the above, it is possible to apply even in the case of a set of wavelengths whose wave number intervals are not constant.

  The wavelength region of the light for the above measurement is preferably 1.2 to 5.0 μm because the light emitting means and the light receiving means can be easily obtained, and particularly 1.3 to 1.6 μm (better 1.35 to 1.6 [mu] m, further 1.4 to 1.6 [mu] m, and most preferably 1.5 to 1.6 [mu] m) is very preferable because light emitting means and light receiving means for optical communication can be used. In other words, the variable wavelength region of the variable wavelength light generating means is preferably 1.2 to 5.0 μm and wider than 1.3 to 1.6 μm, and particularly 1.3 to 1.6 μm within 1.35. If it is wider than ~ 1.6 μm (more preferably within 1.35 to 1.6 μm, wider than 1.4 to 1.6 μm, and further within 1.4 to 1.6 μm, 1.5 to 1.6 μm) Widely, and most preferably within 1.5-1.6 μm).

  The main inspection target of the dental inspection device is whether the tooth structure is normal, especially there are no caries, and if the teeth are treated, whether the treatment status is normal, such as fillings, periodontal disease Whether the state of gums is normal. When used for dental diagnosis, OCT tomographic images of various sites in the oral cavity are also useful. The present invention is mainly intended for teeth, but can also be used for these.

  The tooth structure is enamel on the outermost side, dentine on the inner side, and pulp on the inner side. Enamel is known to exhibit strong birefringence in structure, and even if circularly polarized light is incident, scattered light is polarized. Therefore, the boundary between enamel and dentin can be clearly distinguished by measuring the polarization characteristics of the scattered light. In addition, since the prosthetic material associated with the dental treatment has low birefringence, the boundary with the enamel can be clearly discriminated by measuring the polarization. Since this birefringence exhibits wavelength dependence so that the refractive index depends on the wavelength, OFDR-OCT that can be analyzed spectroscopically by dividing the wavelength region is optimal.

Here, an embodiment of an optical interference tomography apparatus for teeth according to the present invention (an example in the case of measuring the polarization characteristics of teeth) will be described with reference to FIG. FIG. 10 is a schematic configuration diagram of a tooth optical interference tomography apparatus.

  As shown in FIG. 10, the direction of linearly polarized light emitted from the variable wavelength light generating device 11 serving as variable wavelength light generating means is determined by a polarizing plate 200, and the sample optical path is determined by a beam splitter 201 serving as main dividing means. Light is split into a reference optical path.

  The light (measurement light) in the sample optical path is circularly polarized by the wave plate 202 and is irradiated onto the teeth 100 through the probe 30 which is a tube-shaped measurement light irradiation means having flexibility. Reflected light (signal light) from the inside of the teeth is collected by the probe 30 which is also a signal light capturing means, and returns to the beam splitter 201 through the wave plate 202. The beam splitter 201 also serves as a multiplexing unit. Scattered light is polarized depending on the birefringent nature of the sample.

  The light (reference light) that is split by the beam splitter 201 and enters the reference optical path passes through the attenuation element 203 and the wave plate 204, is reflected by the reference mirror 205, returns to the reference optical path, enters the beam splitter 201, and enters the signal light. Is combined. The attenuation factor of the attenuation element 203 is set so that the signal-to-noise ratio is optimized. The wave plate 204 is set so that the light returning to the beam splitter 201 becomes circularly polarized light.

  A polarization beam splitter 206 as polarization separation means separates light from the beam splitter 201 into a horizontal polarization component and a vertical polarization component. The horizontal polarization component is detected by the photodetector 207, and the vertical polarization component is detected by the photodetector 208, amplified and A / D converted, and input to a computer (not shown) which is an arithmetic control means.

  From the detected light intensity of the horizontal and vertical polarization components and the phase relationship between them, the polarization characteristics of the teeth are calculated by the computer as a function of the position inside the teeth.

  That is, the optical interference tomography device for teeth according to the present embodiment includes a polarization property measuring unit for measuring the polarization property of the tooth, and the polarization property measuring unit is changed from the variable wavelength light generation unit (variable wavelength light generation device 11). Main splitting means (beam splitter 201) for controlling the polarization direction of the generated light to split it into measurement light and reference light, and the measurement light split by the main splitting means (beam splitter 201) to the tooth 100 in the oral cavity Measuring light irradiating means (probe 30) for irradiating, signal light capturing means (probe 30) for capturing signal light irradiated and reflected on the tooth 100, and signal light captured by the signal light capturing means (probe 30) Multiplexing means (beam splitter 201) for multiplexing the reference light split by the main splitting means (beam splitter 201) and multiplexed by the multiplexing means (beam splitter 201) Is separated into two or more polarization direction components, a polarization separation means (polarization beam splitter 206), and an arithmetic control means for obtaining the polarization characteristics of the tooth 100 based on the intensities of the light having different polarization directions separated by the polarization beam splitter 206 (The computer).

Further, an embodiment of a tooth optical interference tomography device according to the present invention (an example in the case where a tooth decay is detected by reflection intensity) will be described with reference to FIGS. FIG. 1 is a schematic configuration diagram of a caries detection device that is a tooth optical interference tomography device, and FIG. 2 is a schematic configuration diagram of the probe of FIG.

  As shown in FIG. 1, for example, variable wavelength light generating means for emitting light while changing the wavelength, such as a super-periodic structure diffraction grating distributed reflection semiconductor laser light generating device (see, for example, Non-Patent Document 3). The light output port of a certain variable wavelength light generator 11 is optically connected to the light receiving port of the first coupler 12 composed of a directional coupler or the like that divides light into two (for example, 90:10).

  The light transmission port on one side (the division ratio 90% side) of the first coupler 12 is a second coupler 13 which is a main dividing means including a directional coupler that divides light into two (for example, 70:30). It is optically connected to the light receiving port. The light receiving port of the second coupler 13 is optically connected to a light emitting port of an aiming light source 14 that is a visible light source that emits light in the visible region for visually confirming the irradiation position of the measuring light. ing.

  The light transmission port on one side of the second coupler 13 (the division ratio side of 70%) is optically connected to the light reception port of the optical circulator 15. The light transmission port on the other side (division ratio 30% side) of the second coupler 13 is a third coupler 16 that is a multiplexing means including a directional coupler that divides light into two (for example, 50:50). It is optically connected to the light receiving port. The optical circulator 15 is optically connected to the light receiving port of the third coupler 16 and is connected to the proximal end side of a tube-like probe 30 having flexibility. The probe 30 has a structure as shown in FIG.

  As shown in FIG. 2, the outer cylinder 31 made of a flexible resin or the like has light transmission at least at the tip side and is closed at the tip side. Inside the outer cylinder 31, an inner cylinder 32 made of a resin having flexibility and having a closed tip is inserted and supported so as to be able to slide and rotate in the circumferential direction with respect to the outer cylinder 31. . The inner cylinder 32 is filled with a filler 33 made of a flexible resin or the like, and an optical fiber 34 is arranged and supported so as to be coaxial. The proximal end side of the optical fiber 34 is optically connected to the optical circulator 15.

  An entrance / exit light window 32 a is formed in a part of the peripheral wall on the distal end side of the inner cylinder 32. A reflection mirror 35 is disposed on the tip side inside the inner cylinder 32. An optical system member 36 such as a condensing imaging lens is disposed between the distal end of the optical fiber 34 inside the inner cylinder 31 and the reflection mirror 35. An observation mirror 37 for visual confirmation is disposed outside the distal end side of the outer cylinder 31. The observation mirror 37 is fixedly supported by a bracket 38 attached to the outer peripheral surface on the distal end side of the outer cylinder 31. In addition, a movement support tool (not shown) that facilitates support and movement in the oral cavity is attached to the distal end side of the outer peripheral surface of the outer cylinder 31.

  That is, the measurement light incident from the proximal end side of the optical fiber 34 is shaped into a thin parallel beam by the optical system member 36, and is then passed through the reflection mirror 35 from the input / output light window 32 a of the inner cylinder 32. The signal light transmitted through the cylinder 31 and emitted, reflected on the teeth 100 and reflected (backscattered) is transmitted through the outer cylinder 31 and enters the inner cylinder 32 from the input / output light window 32a of the inner cylinder 32. The optical fiber 34 enters the optical fiber 34 through the reflection mirror 35 and the optical system member 36 and enters the optical circulator 15. In FIG. 2, 39 is a rotate bearing.

  In this embodiment, the reflection mirror 35, the optical system member 36, and the like constitute a communication means, and the probe 30, the optical circulator 15 and the like serve as both the measurement light irradiation means and the signal light acquisition means. Is configured.

  As shown in FIG. 1, the light transmission ports on one side and the other side of the third coupler 16 are optically connected to the light reception port of a first differential amplifier 17 having a light detection function. The Log output unit of the first differential amplifier 17 is electrically connected to the Log input unit of the second differential amplifier 18 that corrects and calculates fluctuations in the input signal intensity.

  On the other hand, the light transmission port on the other side (the division ratio 10% side) of the first coupler 12 is optically connected to the light reception port of the photodetector 19. The output unit of the photodetector 19 is electrically connected to the input unit of the Log amplifier 20. The Log output unit of the Log amplifier 20 is electrically connected to the Log input unit of the second differential amplifier 18.

  An output unit of the second differential amplifier 18 synthesizes a coherent interference waveform, that is, a backscattering intensity distribution (see, for example, Non-Patent Document 4). Is electrically connected. The output unit of the calculation control device 21 is electrically connected to the input unit of the display device 22 such as a monitor or a printer that displays the calculation results. The arithmetic and control unit 21 can control the variable wavelength light generator 11 based on the input information.

  In this embodiment, the first differential amplifier 17, the second differential amplifier 18, the photodetector 19, the log amplifier 20, the arithmetic control device 21, the display device 22, and the like constitute arithmetic control means. Yes.

  Next, a tooth inspection method (caries detection method) using the caries detection device which is such a tooth optical interference tomography device according to this embodiment will be described.

  By inserting the distal end side of the probe 30 into a human oral cavity, positioning and supporting the probe 30 at a predetermined location in the oral cavity using the moving support tool, and operating the arithmetic control device 21, the variable The light for measuring the target wavelength region (wavelength variable range: 1500 to 1550 nm, spectrum frequency width: 10 MHz or less, scanning wave number (A scan number): 400) is generated from the wavelength light generator 11 and aiming light -Visual light is emitted from the source 14.

  The light generated from the variable wavelength light generator 11 is divided into two (90:10) by the first coupler 12. The light on one side (90% side) divided into two by the first coupler 12 is divided into two (70:30) by the second coupler 13. The other side (10% side) light (correction light) divided into two by the first coupler 12 is sent to the photodetector 19.

  The light (measurement light) on one side (70% side) divided into two by the second coupler 13 passes through the optical fiber 34 of the probe 30 through the optical circulator 15 together with the visual recognition light, as described above. In this manner, the tooth 100 is irradiated by being emitted from the distal end side of the probe 30.

  At this time, the probe 30 has flexibility, and can slide and rotate the inner cylinder 32 in the circumferential direction with respect to the outer cylinder 31, and the visible light is emitted together with the measurement light. It is possible to easily irradiate the tooth 100 at the target position in the oral cavity with the measurement light.

  The light (signal light) irradiated and reflected (backscattered) on the teeth 100 is incident on the probe 30 again as described above, and sent to the third coupler 16 via the optical circulator 15. The other side (30% side) light (reference light) divided by the second coupler 13 is sent to the third coupler 16 to be combined with the signal light.

  The light combined by the third coupler 16 is sent to the first differential amplifier 17. The first differential amplifier 17 outputs a Log output signal to the second differential amplifier 18. The photodetector 19 converts the other side (10% side) light (corrected light) divided by the first coupler 12 into an electrical signal and outputs the electrical signal to the Log amplifier 20. The Log amplifier 20 outputs a Log output signal to the second differential amplifier 18. The second differential amplifier 18 performs an input intensity correction calculation, and then outputs the information signal to the analog / digital converter.

  The analog / digital converter converts the input information signal into a digital signal and outputs the digital signal to the arithmetic and control unit 21. The arithmetic and control unit 21 performs arithmetic processing based on various pieces of input information, obtains a coherence interference waveform, that is, the intensity of the signal light, and obtains characteristics of the tooth 100 based on the intensity and the like (details will be described later). The result is displayed on the display device 22.

  Based on the characteristic data of the tooth 100 thus obtained, it is possible to determine the characteristic of the tooth that enables detection of initial caries and the like.

  Here, FIG. 11 shows an example of a tooth measurement result using the above-described tooth optical interference tomography apparatus. In this measurement, the first differential amplifier 17 is directly connected to the arithmetic control device 21 without using the photodetector 19, the log amplifier 20, and the second differential amplifier 18 in FIG. 17 proportional outputs were directly input. As the variable wavelength light generator 11, a super periodic structure diffraction grating distributed reflection semiconductor laser light generator having a wavelength range of 1530 to 1570 nm, a wavelength interval of 0.1 nm, and a wavelength scanning speed of 0.1 nm / 10 μs was used.

  The extracted canine teeth were used as samples. The photograph is shown in FIG. In FIG. 11, (A) to (E) are OCT images of respective cross sections along the lines (a) to (e) shown in the photograph (P).

  As shown in (A), the penetration degree is about 4 mm in optical distance. As you go to the tip, the dentin in the surface enamel becomes observable. In (B), the enamel is thick and the dentin is not clearly visible. As you go to (C), (D), (E) and the tip, the dentin inside the enamel becomes more clearly visible. In (D) and (E), the signal of the internal dentin is observed more strongly than the signal of the enamel on the surface. This corresponds to the fact that the dentin is more scattering than the enamel.

  In (C) and (D), a crack penetrating through the dentin is observed from the middle of the enamel as the middle. This crack is not observed on the surface. Thus, according to OCT, it is possible to observe an internal lesion that does not appear on the surface.

  Here, this principle and the like will be described in more detail.

  The enamel that forms the outermost part of the tooth 100 is the hardest part in the human body, and is a composition of about 96% mineral, about 2% organic, and about 2% water by weight. This inorganic part is repeatedly attacked by the acid produced by caries (decalcification), and finally becomes caries. As shown in FIG. 3, the initial caries are not generated on the surface 101 a of the enamel 101 of the tooth 100, but are formed as small pits 102 in the interior 101 b of the enamel 101.

  The enamel 101 of the healthy tooth 100 has a very low moisture content in the composition as described above, but the enamel 101 of the initial carious tooth 100 having the pits 102 as described above Since saliva enters the pit 102, the amount of water per unit volume increases.

  Therefore, in the present embodiment, by determining the ratio of the abundance per unit volume of the composition of the enamel 101 of the tooth 100 and moisture, the characteristics of the tooth 100 such as the initial caries can be easily detected. It is.

  By the way, among the OCT methods, an OFDR-OCT apparatus to which an optical frequency domain reflectometry method (hereinafter referred to as “OFDR method”) using a variable wavelength light as a light source is used to capture a tomographic image of a living body. At the same time, the wavelength dependence of the light absorption coefficient of the living body can be measured, and for example, it has been proposed to use it for measuring the oxygen saturation of the living body (for example, see Patent Document 1).

  In such a conventional OFDR-OCT apparatus, water that occupies a high proportion of several tens of percent in most human tissues including bones absorbs infrared light strongly, so that near infrared light is used for measurement. Is used, and light in a wavelength region of 650 to 1100 nm, which is called a living body window, is used.

  On the other hand, in this embodiment, paying attention to the characteristics greatly different from other human tissues of the enamel 101 of the tooth 100 that hardly contain water, the composition of the enamel 101 of the tooth 100 is absorbed. The composition of the enamel 101 of the tooth 100 by using light in the wavelength region of 1.2 μm or more, which has not been used in the conventional OFDR-OCT apparatus without being absorbed in water. The characteristics of the tooth 100 (presence / absence of the pit 102 and its position) are obtained from the ratio of the amount of moisture per unit volume.

  Next, the calculation method of the signal light intensity distribution and the light absorption coefficient distribution of the enamel 101 of the tooth 100 and the abundance distribution of the composition and moisture per unit volume will be described more specifically.

  FIG. 4 is an explanatory diagram of the measurement principle of the light absorption coefficient in a minute region. In FIG. 4, a case is considered where measurement light Lm having a wavelength in a certain range having a center wavelength λ is incident along the Z-axis direction as the thickness direction (depth direction) of the tooth 100. In the OCT method, the signal light Ls reflected (backscattered) along the optical axis of the incident measurement light Lm can be measured with a resolution of tens of μm in the Z-axis direction.

  As shown in FIG. 4, a certain position along the optical axis (Z axis) of the incident measurement light Lm is defined as z1, and a position different from the position z1 by a slight distance Δz (for example, about several tens of μm) is defined as z2. . It should be noted that at the positions z1 and z2, the light scattering ability is equal, and light attenuation is caused only by light absorption.

  When the light absorption coefficient of a minute region between the positions z1 and z2 is expressed by μ (z1, λ) as a function of the position z1 and the center wavelength λ of the measurement light Lm, the signal light reflected (backscattered) at the position z1. The ratio between the intensity I (z1) of Ls1 (the intensity of the OCT signal) and the intensity I (z2) of the signal light Ls2 reflected (backscattered) at the position z2 is expressed as follows according to Bare-Lambert law of light absorption: It can be expressed by equation (1).

I (z1) / I (z2) = exp [2 · μ (z1, λ) · Δz] (1)

  The coefficient 2 in the above equation (1) is a value for considering both the incident measurement light Lm and the reflected (backscattered) signal lights Ls1 and Ls2. Based on this equation (1), the light absorption coefficient μ (z1, λ) between the positions z1, z2 in the measurement light Lm of the center wavelength λ is obtained.

  Here, when obtaining the distribution of the ratio of the composition of the enamel 101 of the tooth 100 and moisture per unit volume, it is very preferable to measure the distribution of the light absorption coefficient in a plurality of different wavelength regions. As the number of types of the center wavelength λ of the measurement light Lm increases, the number of parameters that can be determined increases and the accuracy can be improved. However, in the present embodiment, for convenience of explanation, the measurement lights Lm1 and Lm2 in the respective wavelength regions of two different types of center wavelengths λ1 and λ2 are used.

  Specifically, the variable wavelength region (1500 to 1550 nm) of the measurement light Lm that can be generated from the variable wavelength light generator 11 is a first wavelength region (1500 to 1525 nm) and a second wavelength region (1525 to 1550 nm). And the intensity (solid line) of the signal light Ls1 with respect to the measurement light Lm1 in the first wavelength region at the position in the thickness direction (depth direction) of the tooth 100, as shown in FIG. The intensity (dotted line) of the signal light Ls2 with respect to the measurement light Lm2 in the second wavelength region is obtained.

  Then, the distribution of the light absorption coefficient for each wavelength region is obtained from the distribution of the intensity of the signal light Ls1, Ls2, and from the distribution of the light absorption coefficient, per unit volume of the composition of the enamel 101 of the tooth 100 and moisture. The distribution of the abundance ratio is obtained.

  That is, in the conventional OFDR-OCT apparatus, the measurement light Lm is emitted over the entire variable wavelength region that can be generated from the variable wavelength light generation apparatus 11, and the light from the intensity of the reflected light (backscattered light) of the measurement light Lm is emitted. Although the absorption coefficient has been obtained, in this embodiment, the variable wavelength region that can be generated from the variable wavelength light generator 11 is changed to a plurality of measurement light beams Lm1, Lm2,. .. Are emitted, and the light absorption coefficient is obtained from the intensity of the signal light that is each reflected light (backscattered light) of each of the measurement lights Lm1, Lm2,.

  FIG. 6 is a graph showing the relationship between the center wavelength λ of the enamel 101 and water measurement light Lm and the light absorption coefficient. As can be seen from FIG. 6, the enamel 101 and water have greatly different light absorption coefficients depending on the wavelength.

For example, if the center wavelength λ1 of the first measurement light Lm1 is 1512.5 nm and the center wavelength λ2 of the second measurement light Lm2 is 1537.5 nm, the light absorption coefficient of the enamel 101 is the first and second light absorption coefficients. While the measurement light Lm1 and Lm2 show substantially the same value (3.8 cm ⁻¹ ), the first measurement light Lm1 (center wavelength λ1) has a light absorption coefficient of water that is the second measurement light Lm2 ( The value becomes larger than the central wavelength λ2).

Here, assuming that the existing ratio (concentration) of the composition of the enamel 101 at the position z1 is C _E (z1) and the existing ratio (concentration) of water at the position Z1 is C _H2O (z1), these values Can be obtained from the following equations (2) and (3).

μ _H2O (λ1) · C _H2O (z1) + μ _E (λ1) · C _E (z1) = μ (z1, λ1) (2)
μ _H2O (λ2) · C _H2O (z1) + μ _E (λ2) · C _E (z1) = μ (z1, λ2) (3)

Μ _H2O (λ1) is a light absorption coefficient of water of the first measurement light Lm1 having the center wavelength λ1, μ _H2O (λ2) is a light absorption coefficient of water of the second measurement light Lm2 having the center wavelength λ2, μ _E (λ1) is the light absorption coefficient of the enamel composition of the first measuring light Lm1 having the center wavelength λ1, and μ _E (λ2) is the enamel composition of the second measuring light Lm2 having the center wavelength λ2. These are light absorption coefficients, and these values can be obtained from the graph of FIG.

  Further, μ (z1, λ1) is an optical absorption coefficient of the first measurement light Lm1 having the center wavelength λ1 at the position z1, and μ (z1, λ2) is the second measurement light Lm2 having the center wavelength λ2 at the position z1. It is a light absorption coefficient, and these values are obtained by actual measurement.

  In this way, reflection of the first and second measurement lights Lm1 and Lm2 having the center wavelengths λ1 and λ2 within a narrow region between the position z1 and the position z2 that is Δz (about several tens of μm) away from the position z1. Based on the intensity of light (backscattered light), that is, the intensity of the signal lights Ls1 and Ls2, the respective light absorption coefficients between them are obtained, thereby obtaining the composition of the enamel 101 and moisture per unit volume. The proportion of abundance can be determined.

  At this time, as described above, since a small amount of water is present in the composition of the enamel 101 of the healthy tooth 100 (about 2%), the amount of water present as described above The ratio is a measurement result including moisture in the composition of enamel 101, but does not cause a substantial problem.

  FIG. 7 shows the distribution of the ratio of the abundance of the composition of the enamel 101 and the moisture in the thickness direction (depth direction) of the tooth 100 obtained as described above. As can be seen from FIG. 7, the position where the ratio of the amount of moisture present is high (about 1) is the location where the pits 102 are present. In addition, in the surface (interface) portion S of the tooth 100 and the pit 102, strong reflected light (backscattered light) generated by Fresnel reflection occurs, and the measurement error of the light absorption coefficient becomes very large. Calculation of percentage is omitted.

  Thus, in the present embodiment, paying attention to the phenomenon that a liquid mainly composed of water such as saliva enters the pit 102 formed in the enamel 101 of the tooth 100 with less moisture, As described above, the distribution of the ratio of the abundance between the composition of the enamel 101 and the water is obtained, so that the characteristics of the tooth 100 such as the presence / absence of the pit 102 and its size and position are clearly understood. It was.

  Then, the incident measurement light Lm is scanned (B scan) to perform two-dimensional measurement. Specifically, for example, the inner cylinder 32 is slid and rotated in the circumferential direction with respect to the outer cylinder 31 of the probe 30. By moving the measuring light Lm linearly along the surface of the tooth 100, a tomographic image of the pit 102 can be obtained, and further, two-dimensional measurement is repeated while shifting the scanning position little by little. By displaying two-dimensional images in parallel, a three-dimensional (three-dimensional) tomographic image of the pit 102 can be obtained, and the presence or absence of the pit 102 and the characteristics of the tooth 100 such as its size and position can be grasped more clearly. can do.

  Even in the decalcification progress state before the pits 102 are formed on the enamel 101 of the tooth 100, the inorganic components are melted from the enamel 101 and microscopic gaps are generated in the enamel 101. As a result, water permeates into the gap at the molecular level and the amount of water per unit volume of the enamel 101 increases, so that the gap can be detected in the same manner as described above.

  Further, by using measurement light in a wavelength region shorter than 1.2 μm, only the abundance per unit volume of the composition of the enamel 101 is obtained without obtaining the water amount per unit volume of the enamel 101 of the tooth 100. Although it is possible to determine the characteristics of the tooth 100 based on the above, there is a difficulty in accuracy, which is not preferable.

  Further, without obtaining the ratio of the abundance per unit volume of the composition of enamel 101 of the tooth 100 and moisture as shown in FIG. 7, the signal for each wavelength region as shown in FIG. Detection can be easily performed by directly obtaining the characteristics of the tooth 100 based on the light intensity. In the case of such a simple detection, it is particularly desired to obtain the tomographic data by performing the above-described B-scan.

  By the way, when performing tomographic imaging of the retina of the eye, if the intensity of the measurement light is strong, the eye may be adversely affected. Therefore, the optical coherence domain reflectometry method, which is a normal OCT method ( Hereinafter, in the OCDR-OCT apparatus to which the “OCDR method” is applied, the optical coupling efficiency with the optical fiber is poor and sufficient intensity as the measurement light cannot be obtained, and the S / N ratio is improved. Even if a super luminescent diode (hereinafter referred to as “SLD”), which is a kind of light emitting diode, is difficult to produce, no problem occurs.

  On the other hand, in the present embodiment, the variable wavelength light generator 11 which is a semiconductor laser light generator that has high optical coupling efficiency with the optical fiber 31 and can improve the S / N ratio with a sufficiently strong measurement light is used as the measurement light. Therefore, if the intensity of the measurement light generated from the variable wavelength light generator 11 is sufficiently increased to improve the S / N ratio, it is not necessary to use measurement light in a plurality of different wavelength regions. It is possible to detect the characteristics of the tooth 100 such as the initial caries. However, if measurement light in a plurality of different wavelength regions as described above is not used, accuracy is difficult, which is not preferable.

  In the present embodiment, it is assumed that the absorption of the measurement light Lm by the tooth 100 is dominant, the attenuation of the signal light Ls due to reflection scattering can be ignored, and the backscattering ability is the same at all positions. The case of obtaining the characteristics of the tooth 100 using the measurement lights Lm1 and Lm2 in the wavelength regions of two different central wavelengths λ1 and λ2 has been described. For example, depending on the composition of the enamel 101 of the tooth 100 and factors other than moisture When the amount of absorption of the measurement light Lm cannot be ignored, when the attenuation of the signal light Ls due to reflection / scattering cannot be ignored, or when the backscattering ability varies depending on the position, measurement of wavelength regions having different center wavelengths λ is performed. What is necessary is just to respond | correspond to the increase number of the parameter which should be determined by increasing the number of light Lm suitably.

  In the present embodiment, the second coupler 13 and the third coupler 16 are used by constructing a Mach-Zehnder type interferometer using the optical circulator 15, but the Michelson type interferometer is used. By constructing, it is also possible to apply a main dividing / combining means that combines the main dividing means and the combining means.

  In the present embodiment, the optical circulator 15 is applied. However, for example, when the optical circulator 15 does not operate with visible light, a coupler 25 is used instead of the optical circulator 15 as shown in FIG. It is also possible to apply.

  In the present embodiment, the optical circulator 15 is used to apply the probe 30 capable of performing the measurement light exit guide and the signal light entrance guide by the same optical fiber 34. However, for example, the optical circulator is used. 9 is omitted, as shown in FIG. 9, two optical fibers 34A and 34B are provided in parallel inside the inner cylinder 36, and the one optical fiber 34A guides the emission of the measurement light, and the other light It is also possible to apply the probe 30 configured to guide the incidence of signal light with the fiber 34B.

  At this time, the optical fibers 34A and 34B are slightly disadvantageous in practical use, although the optical axes of the optical fibers 34A and 34B slightly deviate from each other, resulting in a difference in the optical axes of the outgoing measurement light and the incoming signal light. There is nothing.

The tooth optical interference tomography device according to the present invention is used in the manufacturing industry of precision instruments and the like by producing the device.

It is a schematic block diagram of embodiment (an example in the case of detecting a decayed tooth by reflection intensity) of the optical interference tomography apparatus of the tooth | gear which concerns on this invention. It is a schematic block diagram of the probe of FIG. It is explanatory drawing of an initial cavity. It is explanatory drawing of the measurement principle of the light absorption coefficient in a micro area | region. It is a graph showing the relationship between the position of the tooth thickness direction (depth direction) and signal light intensity. It is a graph showing the relationship between the center wavelength of the measurement light of enamel and water, and a light absorption coefficient. It is a graph showing distribution of the ratio of the abundance of an enamel composition and moisture in the tooth thickness direction (depth direction). It is a schematic block diagram of other embodiment of the optical interference tomography apparatus of the tooth | gear which concerns on this invention. It is a schematic block diagram of the probe of other embodiment of the optical interference tomography apparatus of the tooth | gear which concerns on this invention. It is a schematic block diagram of embodiment (an example in the case of measuring the polarization characteristic of a tooth | gear) of the optical interference tomography apparatus of the tooth | gear which concerns on this invention. It is an OFDR-OCT image of a canine of a person who has extracted a tooth. It is explanatory drawing of the scanning method of the wavelength of the light radiate | emitted from a variable wavelength light generator.

Explanation of symbols

DESCRIPTION OF SYMBOLS 11 Variable wavelength light generator 12 1st coupler 13 2nd coupler 14 Aiming light source 15 Optical circulator 16 3rd coupler 17 1st differential amplifier 18 2nd differential amplifier 19 Optical detector 20 Log Amplifier 21 Arithmetic control device 22 Display device 25 Coupler 30 Probe 31 Outer tube 32 Inner tube 33 Filler 34, 34A, 34B Optical fiber 35 Reflection mirror 36 Optical system member 37 Observation mirror 38 Bracket 39 Rotate bearing 100 Tooth 101 Enamel 101a Surface 101b Inside 102 Pit 200 Polarizing plate 201 Beam splitter 202 Wave plate 203 Attenuating element 204 Wave plate 205 Reference mirror 206 Polarizing beam splitter 207 Photo detector 208 Photo detector

Claims (8)

Variable wavelength light generating means using a variable wavelength light generating device that emits light while switching wavelengths discretely as a light source ;
Main splitting means for splitting light generated from the variable wavelength light generating means into measurement light and reference light;
Measuring light irradiation means for irradiating the teeth in the oral cavity with the measuring light divided by the main dividing means,
Signal light capturing means for capturing the signal light irradiated and reflected on the teeth;
A multiplexing unit that combines the signal light captured by the signal light capturing unit and the reference light divided by the main dividing unit;
As the wavelength region of interest of said light to be generated from the variable-wavelength light generating apparatus, to control the variable wavelength light generating apparatus based on the input information was generated from the variable-wavelength light generating apparatus based on the intensity of the light multiplexed by a wavelength region and the multiplexing means of light, the optical interference tomographic apparatus teeth, characterized in Tei Rukoto and an arithmetic control means for determining the characteristics of the tooth. Variable wavelength light generating means using a variable wavelength light generating device that emits light while switching wavelengths discretely as a light source;
Polarization measuring means for measuring the polarization characteristics of teeth; and
With
The polarization characteristic measuring means is
Main splitting means for controlling the polarization direction of the light generated from the variable wavelength light generating means and splitting it into measurement light and reference light;
Measuring light irradiation means for irradiating the teeth in the oral cavity with the measuring light divided by the main dividing means,
Signal light capturing means for capturing the signal light irradiated and reflected on the teeth;
Multiplexing means for separating the signal light captured by the signal light capturing means into two or more polarization direction components, and respectively combining the reference light divided by the main dividing means;
And a calculation control means for determining the polarization characteristics of the teeth based on the intensity of the light of the combined signals having different polarization directions, and controlling the variable wavelength light generator based on the input information . An optical interference tomography device for teeth. Variable wavelength light generating means using a variable wavelength light generating device that emits light while switching wavelengths discretely as a light source;
Polarization measuring means for measuring the polarization characteristics of teeth; and
With
The polarization characteristic measuring means is
Main splitting means for splitting light generated from the variable wavelength light generating means into measurement light and reference light;
Measuring light irradiation means for irradiating the teeth in the oral cavity with the measuring light divided by the main dividing means,
Signal light capturing means for capturing the signal light irradiated and reflected on the teeth;
A multiplexing unit that combines the signal light captured by the signal light capturing unit and the reference light divided by the main dividing unit;
As the wavelength region of interest of said light to be generated from the variable-wavelength light generating apparatus, to control the variable wavelength light generating apparatus based on the input information was generated from the variable-wavelength light generating apparatus based on the intensity of the light multiplexed by a wavelength region and the multiplexing means of light, the optical interference tomographic apparatus teeth, characterized in Tei Rukoto and an arithmetic control means for determining the characteristics of the tooth. In any one of Claims 1-3,
The arithmetic and control unit controls the variable wavelength light generator so as to generate light of a plurality of different wavelength regions, and obtains the intensity of the light combined by the multiplexing unit for each wavelength region. To obtain the characteristics of the teeth
An optical interference tomography device for teeth. In any one of Claims 1-3,
The arithmetic and control means controls the variable wavelength light generator so that all of the emitted light is scanned within a measurement time within a predetermined wavelength variable range and scanning wavenumber of the target wavelength.
An optical interference tomography device for teeth. In any one of Claims 1-5,
The arithmetic control means is
Dividing the variable wavelength region of the measurement light that can be generated from the variable wavelength light generator into two, a first wavelength region and a second wavelength region, in the tooth thickness direction, the first wavelength region Obtaining the intensity of the signal light Ls1 with respect to the measurement light and the intensity of the signal light Ls2 with respect to the measurement light in the second wavelength region,
Obtain the distribution of the light absorption coefficient for each wavelength region from the intensity distribution of the signal light Ls1, Ls2, respectively.
From the distribution of the light absorption coefficient, by obtaining the distribution of the ratio of the abundance per unit volume of the tooth enamel composition and moisture,
The characteristic of the tooth is obtained
An optical interference tomography device for teeth. In any one of Claims 1-5 ,
The arithmetic control means is
By determining the intensity of the light combined by the combining means for each wavelength region, the light absorption coefficient of the tooth is determined,
Based on the optical absorption coefficient, by determining the abundance of water per unit volume of the abundance, or enamel or dentin of the tooth per unit volume of the enamel or dentin of the composition of the tooth,
A tooth optical interference tomography apparatus characterized in that the characteristic of the tooth is obtained. In any one of Claims 1-7 ,
The width of the variable range of the wave number of the emitted light is 4.7 × 10 ⁻² μm ⁻¹ or more, the frequency width of the emitted light is 13 GHz or less, the frequency interval of the emitted light is 15 GHz or less, and the wave number interval of the emitted light is 3.1 × 10 ⁻⁴ μm ⁻¹ or less, the time interval of the emitted light is set to 530 μs or less, and the arithmetic control unit generates the variable wavelength light so that all wave numbers of the emitted light are switched discretely within the measurement time. Tooth optical interference tomography device characterized in that the device is controlled .

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