JP4688094B2 - Optical coherence tomography device - Google Patents

Description

  The present invention relates to an optical coherence tomography (tomographic measurement using low-coherence light as a probe), which is one of non-destructive tomographic techniques.

  Conventionally, an X-ray imaging apparatus, an intraoral camera, a dental camera, X-ray CT, MRI, and the like have been used for imaging a jaw and mouth region in dental diagnosis.

  An image obtained by the X-ray imaging apparatus is a transmission image to the last, and information on the X-ray traveling direction of the subject is superimposed and detected. Therefore, the internal structure of the subject cannot be known three-dimensionally. Also, because X-rays are harmful to the human body, the annual exposure dose is determined, and only qualified surgeons can handle the device, and only in rooms surrounded by shielding materials such as lead and lead glass. I can not use it.

  Since the intraoral camera images only the surface of the intraoral tissue, internal information such as teeth cannot be obtained. X-ray CT is harmful to the human body as well as an X-ray imaging apparatus, has low resolution, and is large and expensive. MRI has poor resolution, the apparatus is large and expensive, and the internal structure of teeth without moisture cannot be photographed.

  By the way, an optical coherence tomography apparatus (hereinafter referred to as an OCT apparatus) is harmless to the human body and can obtain three-dimensional information of a subject with high resolution, and is therefore applied in the field of ophthalmology such as tomographic measurement of the cornea and retina. (For example, see Patent Documents 1 to 4).

  Here, a conventional OCT apparatus will be described. FIG. 11 is a diagram showing a configuration of a conventional OCT apparatus. In the OCT unit 1 constituting the OCT apparatus shown in FIG. 11, the light emitted from the light source 2 is collimated by the lens 3 and then divided into reference light 6 and measurement light 5 by the beam splitter 4. The measurement light 5 passes through the galvanometer mirror 8 and is collected by the objective lens 9 on the sample 10 to be measured. After being scattered and reflected, the measurement light 5 passes through the objective lens 9, the galvanometer mirror 8, and the beep splitter 4 again. It is condensed on the detector 14. On the other hand, the reference light 6 is reflected by the reference mirror 13 through the objective lens 12, and again passes through the objective lens 12 and the beam splitter 4, and then enters the condenser lens 7 in parallel with the measurement light 5 and enters the photodetector. 14 is condensed.

  The light source 2 is a temporally low coherence light source. Lights emitted at different times from a temporally low coherence light source are extremely unlikely to interfere with each other. For this reason, an interference signal appears only when the distance of the optical path through which the measurement light 5 passes and the distance of the optical path through which the reference light 6 passes are substantially equal. As a result, when the intensity of the interference signal is measured by the photodetector 14 while moving the reference mirror 13 in the optical axis direction of the reference light 6 and changing the optical path length difference between the measurement light 5 and the reference light 6, the sample to be measured 10 is obtained. The reflectance distribution in the depth direction (z-axis direction) can be obtained. That is, the structure in the depth direction of the sample 10 to be measured is obtained by optical path length difference scanning.

  In addition to scanning in the depth direction (z-axis direction) of the sample to be measured by the reference mirror 13, a two-dimensional cross-sectional image of the sample to be measured 10 is obtained by scanning in the lateral direction (x-axis direction) by the galvanometer mirror 8. It is done. Such an OCT apparatus can measure with a high resolution of several μm. Therefore, the OCT apparatus can obtain a high-resolution image inside the living body in a non-destructive and non-contact manner.

Regarding the application of the OCT apparatus to the dental field, an example in which a tomogram of a tooth is photographed using the OCT apparatus is disclosed (for example, see Non-Patent Documents 1 to 5).
JP 2003-329577 A JP 2002-310897 JP 11-325849 A JP 2001-059714 A Laser Research October 2003 Issue: Technology Development of Optical Coherence Tomography Centered on Medical Care Journal of Biomedical Optics, October 2002, Vol.7 No.4: Imaging caries lesions and lesion progression with polarization sensitive optical coherence tomography APPLIED OPTICS, Vol.37, No.16, 1 June 1998: Imaging of hard- and soft-tissue structure In the oral cavity by optical coherence tomography OPTICS EXPRESS, Vol.3, No.6,14 September 1998: Dental OCT OPTICS EXPRESS, Vol.3, No.6,14 September 1998: In vivo OCT Imaging of hard and soft tissue of the oral cavity

  However, the OCT apparatus is not used in actual dental practice. The use of an OCT device for dental diagnosis is not practical at least at present, and there is no OCT device for dental measurement as a product. This is because the OCT apparatus requires two-dimensional mechanical scanning including the depth direction in order to obtain one tomographic image, so that it takes time to image, and the apparatus is complicated and expensive, and durability is also high. It is because it is inferior. That is, there is a problem that it is difficult to apply the OCT apparatus to actual dental measurement.

  In view of the above problems, an object of the present invention is to provide an optical coherence tomography apparatus that has a simple structure, can be imaged at high speed, and can be applied to dental measurement.

  To achieve the above object, a Fourier-domain optical coherence tomography device for dental measurement according to the present invention includes a light source, a reference light for irradiating a reference mirror with a light source emitted from the light source, and a measurement for irradiating a sample to be measured. A light splitting unit that divides the light into light, an interference unit that interferes with the measurement light reflected by the sample to be measured and the reference light reflected by the reference mirror to form interference light, and a diffraction element that splits the interference light And a light detection unit that measures the spectrum of the interference light dispersed by the diffraction element, and an irradiation direction of the measurement light of the sample to be measured by performing Fourier inverse transform on the spectrum detected by the light detection unit And the sample to be measured is a living oral cavity region tissue or an artificial composition of the oral cavity region.

  To achieve the above object, a Fourier-domain optical coherence tomography device for dental measurement according to the present invention includes a light source, a reference light for irradiating a reference mirror with a light source emitted from the light source, and a measurement for irradiating a sample to be measured. A light splitting unit that divides the light into light, an interference unit that interferes with the measurement light reflected by the sample to be measured and the reference light reflected by the reference mirror to form interference light, and a diffraction element that splits the interference light And a light detection unit that measures the spectrum of the interference light dispersed by the diffraction element, and an irradiation direction of the measurement light of the sample to be measured by performing Fourier inverse transform on the spectrum detected by the light detection unit And a polarization operation unit for operating at least one polarization state of the light source light, the reference light, the measurement light, or the interference light. Samples stomatognathic region of body tissue or an artificial composition Stomatognathic region.

  In order to achieve the above object, an optical coherence tomography device according to the present invention includes a light source and light that divides the light source light emitted from the light source into reference light for irradiating a reference mirror and measurement light for irradiating a sample to be measured. A splitting unit, an interference unit that interferes with the measurement light reflected by the sample to be measured, and the reference light reflected by the reference mirror to make interference light, a light detection unit that measures the interference light, and the light A calculation unit for obtaining information of the sample to be measured based on the interference light measured by the detection unit, and a nonlinear optical element provided on the optical path of the reference light.

  According to the present invention, it is possible to provide an optical coherence tomography apparatus that has a simple structure, can be imaged at high speed, and can be applied to dental measurement.

  A Fourier domain optical coherence tomography device for dental measurement according to the present invention includes a light source, a light splitting unit that divides the light source light emitted from the light source into reference light that irradiates a reference mirror and measurement light that irradiates a sample to be measured. The measurement light reflected by the sample to be measured and the reference light reflected by the reference mirror are caused to interfere with each other to form interference light, a diffraction element that splits the interference light, and the diffraction light is split by the diffraction element. A light detection unit that measures the spectrum of the interference light, and a calculation unit that obtains information in the irradiation direction of the measurement light of the sample to be measured by performing inverse Fourier transform on the spectrum detected by the light detection unit. And the sample to be measured is a living tissue of the oral cavity region or an artificial composition of the oral region of the oral cavity.

  In the Fourier domain optical coherence tomography device for dental measurement according to the present invention, the interference light is spectrally separated by a diffraction element, so that the light detection unit can detect the spectrum of the interference light. The calculation unit obtains information on the irradiation direction of the measurement light of the sample to be measured by performing Fourier inverse transform on the spectrum of the interference light. Therefore, information in the depth direction of the sample to be measured can be obtained without mechanical scanning in the irradiation direction of the measurement light, that is, the depth direction of the sample to be measured. As a result, the structure of the apparatus is simplified and imaging can be performed at high speed. As a result, an OCT apparatus that can be applied to dental diagnosis while a patient is placed on a treatment chair generally used in dentistry is obtained.

  It is preferable that the light splitting unit and the interference unit have both functions using a beam splitter or a fiber coupler.

  The Fourier-domain optical coherence tomography device for dental measurement according to the present invention according to the present invention divides a light source and light source light emitted from the light source into reference light for irradiating a reference mirror and measurement light for irradiating a sample to be measured. A light splitting unit; an interference unit that interferes with the measurement light reflected by the sample to be measured; and the reference light reflected by the reference mirror to make interference light; a diffraction element that splits the interference light; and the diffraction Information on the irradiation direction of the measurement light of the sample to be measured is obtained by Fourier-transforming the spectrum detected by the light detection unit and the spectrum detected by the light detection unit. A calculation unit; and a polarization operation unit that operates at least one polarization state of the light source light, the reference light, the measurement light, or the interference light. Stomatognathic regions tissue or an artificial composition Stomatognathic region.

  In the optical coherence tomography device for dental measurement according to the present invention, the polarization operation unit has at least one polarization state of light irradiated from the light source to the light splitting unit, the reference light, the measurement light, or the interference light. Since the operation is performed, an image reflecting the polarization property or birefringence property of the sample to be measured is obtained. As a result, it is possible to observe oral tissues rich in specific polarization characteristics or birefringence characteristics such as initial caries, dentin, enamel, gingiva, and alveolar bone.

  The Fourier domain optical coherence tomography apparatus for dental measurement according to the present invention further includes a cylindrical lens or a cylindrical mirror that makes the cross section of the measurement light linear along a single axial direction in a plane perpendicular to the irradiation direction of the measurement light. It is preferable to provide.

  A cylindrical lens is a lens that acts as a lens in only one of two directions perpendicular to the optical axis, and at least one surface of the cross section in the direction acting as a lens becomes a lens-specific curve. The cross section in the direction not acting as a lens is a straight line parallel to each other on both surfaces.

  A cylindrical mirror is a mirror that acts as a lens in only one of two directions perpendicular to the optical axis. Only the surface of the cross section in the direction acting as a lens becomes a lens-specific curve and acts as a lens. The cross section in the direction not to be straight is a straight line.

  Since the cylindrical lens or the cylindrical mirror makes the cross section of the measurement light linear along a single axial direction in a plane perpendicular to the irradiation direction of the measurement light, the measurement light is the one axis of the sample to be measured. Distributed irradiation in the direction. That is, the measurement light is condensed on the line in the uniaxial direction in the sample to be measured. Therefore, the cross section of the sample to be measured in the uniaxial direction can be measured without performing mechanical scanning in the uniaxial direction.

  The spectrum expanded in the one-axis direction is a light beam spread in two dimensions, that is, the direction of the spectrum and the expanded direction perpendicular to the spectrum direction, and is preferably detected by a two-dimensional light detection unit. Further, the two-dimensional light detection unit is preferably a two-dimensional CCD image sensor or a two-dimensional CMOS image sensor.

  In the optical coherence tomography device according to the present invention, it is preferable that at least one of the light source light, the measurement light, the reference light, the interference light, and the light split into the spectrum is guided by an optical fiber. In that case, the optical fiber is preferably an optical fiber whose cross section is bundled into a one-dimensional linear shape or a two-dimensional circular shape.

  The optical coherence tomography device for dental measurement according to the present invention projects the measurement light or visible light pattern on the surface of the object to be measured, and monitors the surface image of the measurement site with a two-dimensional imaging device, or further tomographic measurement image It is preferable that the recording is performed in synchronization.

  Hereinafter, embodiments of the present invention will be described with reference to the drawings.

(Embodiment 1)
FIG. 1 is a diagram illustrating an example of a configuration of a Fourier domain optical coherence tomography apparatus (hereinafter referred to as an FD-OCT apparatus) in the first embodiment.

  The FD-OCT apparatus detects a spectrum obtained by separating the interference light between the measurement light reflected by the sample to be measured and the reference light reflected by the reference mirror, and from this spectrum, the measurement light irradiation direction of the sample to be measured is detected. The OCT apparatus is characterized in that information is obtained using Fourier inverse transform.

  As shown in FIG. 1, the FD-OCT apparatus includes an OCT unit 100, a measurement head 201, and a computer 27. The OCT unit 100 is provided with a light source 16, a fiber coupler 19, a reference mirror 24, a diffraction element 25, and a CCD camera 26. The measurement head 201 is provided with a galvanometer mirror 20 and an objective lens 21. The computer 27 is connected so as to be able to communicate with the light source 16, the CCD camera 26, and the galvanometer mirror 20. The computer 27 is, for example, a personal computer, and includes at least a calculation unit such as a CPU and a recording unit such as a hard disk.

  Note that the configurations of the OCT unit 100, the measurement head 201, and the computer 27 are not limited to the configurations shown in FIG. For example, the function of the computer 27 can be incorporated in the OCT unit 100.

  In the present embodiment, the sample 22 to be measured is a living tissue of the oral cavity region or an artificial composition of the oral region of the jaw.

  The light source 16 is a temporally low coherent light source. That is, the light source has a wavelength distributed in a narrow range. The light source 16 is preferably a superluminescent diode, for example.

  The fiber coupler 19 is an example of an optical interferor that functions as a light splitting unit and an interference unit. An optical interferometer is an input / output interchangeable optical component that causes two input lights to interfere and output in two directions. Examples of the optical interferometer include a beam splitter and a half mirror in addition to the fiber coupler 19.

  The diffraction element 25 is a reflection-type or transmission-type optical member having a diffraction spectroscopy function. The diffraction element 25 is preferably a grating element, a diffraction grating, a prism, or the like, for example. The diffraction element 25 may be a slice of an optical recording medium. The optical recording medium is, for example, a CD, DVD, MO, or the like.

  The CCD camera 26 is an example of a light detection unit. As the light detection unit, for example, a one-dimensional photodetector, a two-dimensional photodetector, or the like can be used. The one-dimensional photodetector is preferably a linear CCD, and the two-dimensional photodetector is preferably a CCD imaging device or a CMOS imaging device. The two-dimensional photodetector includes a two-dimensional imaging device.

  The measurement head 201 is preferably configured so that an operator can operate it by hand. By transmitting light between the OCT unit 100 and the measurement head 201 by the optical fiber 18, the movable range of the measurement head 201 is widened.

  When the OCT apparatus is applied to dentistry, it is assumed that the OCT apparatus is used on the chair side of a chair in which a patient is usually sitting during medical examination. In this case, in order to position the measurement head, in the aerial optical system (the optical path to the measurement head is not an optical fiber but in the air), the entire OCT unit must be precisely positioned in the patient's oral cavity. Moreover, it is unrealistic for an operator to operate a relatively heavy OCT unit.

  Since the measurement head 201 is configured to be operated by an operator by hand, the operator can easily use the chair side in a dental practice. The operator can use the OCT apparatus in a state where the positional relationship between the patient and the measurement head is free.

  Next, the operation of the FD-OCT apparatus shown in FIG. 1 will be described. In the following description, the coordinate system is defined as follows. As shown in FIG. 1, in the sample 22 to be measured, the optical axis direction of the measurement light 28, that is, the depth direction of the sample 22 to be measured is z, the tomographic plane is the zy plane (the galvano mirror 20 scan direction is y). ), And at locations other than the sample 22 to be measured, directions x, y, and z corresponding to the x, y, and z of the sample 22 to be measured are assumed to be x, y, and z. Optically means that even if the spatial direction changes with a mirror, lens, optical fiber, etc., the light traveling direction is z, the direction scanned with a galvanometer mirror is y, and both z and y are perpendicular to each other. The correct direction is x.

  The light emitted from the light source 16 is collimated by the lens 17 and then divided into reference light 29 and measurement light 28 by the fiber coupler 19. The measurement light 28 passes through the optical fiber 18 and the galvanometer mirror 20 and is collected by the objective lens 21 on the sample 22 to be measured. After being scattered and reflected there, the objective lens 21, the galvanometer mirror 20, the optical fiber 18, and the fiber coupler 19 are again passed. Then, the light is guided to the diffraction element 25 by the condenser lens 30.

  On the other hand, the reference light 29 is reflected by the reference mirror 24 through the optical fiber 18 and the objective lens 23, is again caused to interfere with the measurement light 28 by the fiber coupler 19 through the objective lens 23, and is parallel to the measurement light 28. Is incident on the condenser lens 30 and guided to the diffraction element 25.

  The measurement light 28 and the reference light 29 are simultaneously dispersed by the diffraction element 25 and overlapped in the spectral region, thereby forming spectral interference fringes on the CCD camera 26, that is, a combined power spectrum of the measurement light 28 and the reference light 29. . A spectral correlation fringe measured by the CCD camera 26 is subjected to inverse Fourier transform in the computer 27, whereby a combined correlation between the measurement light 28 and the reference light 29 is obtained. From this binding correlation, information on the structure, composition, or optical characteristics can be obtained through the reflectance characteristics in the depth direction (z-axis direction) of the sample 22 to be measured.

  Therefore, it is not necessary to move the reference mirror 24 to adjust the optical path length of the measurement light 28 and the optical path length of the reference light 29 and perform scanning in the z-axis direction. That is, information regarding the structure in the depth direction (z-axis direction) of the sample 22 to be measured can be obtained without performing a mechanical operation in the z-axis direction.

  As described above, the FD-OCT apparatus that obtains internal information in the z-axis direction of the sample to be measured based on the spectrum obtained by separating the interference light between the measurement light 28 and the reference light 29 by the diffraction element 25 is provided. Device.

  In order to obtain a two-dimensional cross-sectional image of the sample 22 to be measured, it is necessary to scan in the y-axis direction in addition to the z-axis direction. In the present embodiment, scanning in the y-axis direction is performed by driving the galvanometer mirror 20.

  As a scanning method in the y-axis direction, in addition to the method of driving the carbano mirror 20, a method using a cylindrical lens, a method of driving a lens, a method of driving an optical fiber, and a method of driving the sample 22 to be measured Alternatively, a method in which the operator moves the measurement head 201 can be used.

  Here, as a modified example of the scanning method in the y-axis direction, a method of driving a lens will be described.

  FIG. 5 is a conceptual diagram for explaining an example of a method for driving a lens. A linear actuator 31 is connected to one end of the lens 30, and the other end of the lens 30 is fixed to the apparatus. As a result of the linear actuator 31 being driven in the z direction, the lens 30 reciprocates in an arc shape around the rotation shaft 32. As a result of the circular reciprocation of the lens 30, the optical axis of the lens 30 moves in the zy plane, and the sample to be measured is scanned in the y direction.

  In order to obtain a three-dimensional structure of the sample 22 to be measured, it is necessary to perform scanning in the x-axis direction in addition to operations in the z-axis direction and the y-axis direction. Scanning in the x-axis direction can be performed in the x-axis direction by driving the carbano mirror 20 in the x-axis direction as well as scanning in the y-axis direction. Also in the scanning in the x-axis direction, the same method as the example of the scanning method in the y-axis direction can be used. A suitable one of the above-described examples of the scanning method in the y-axis direction can be combined to perform scanning in the y-axis and x-axis directions.

  In the FD-OCT apparatus, as described above, the structure in the z-axis direction of the sample 22 to be measured is obtained from the spectral interference fringes, so that one-dimensional mechanical scanning is performed to obtain a tomographic image of the sample 22 to be measured. Two-dimensional mechanical scanning is sufficient to obtain 3D information. As a result, the structure of the apparatus is simplified, and imaging can be performed at high speed. As a result, excellent characteristics such as non-invasiveness and high resolution such as the basic characteristics that the OCT apparatus can quantitatively acquire the three-dimensional internal information of the sample to be measured are utilized in the dental field.

  As described above, according to the present embodiment, the measurement head 201 can be operated by hand, and a FD-OCT apparatus having a simple structure capable of high-speed imaging can be obtained. An FD-OCT apparatus is obtained.

(Embodiment 2)
FIG. 2 is a diagram illustrating an example of a configuration of the FD-OCT apparatus according to the second embodiment. In FIG. 2, the same parts as those of the FD-OCT apparatus shown in FIG.

  The FD-OCT apparatus shown in FIG. 2 is different from the FD-OCT apparatus shown in FIG. 1 in that a cylindrical lens 33 is provided, a beam splitter 34 is used instead of the fiber coupler 19, and This is the scanning direction of the galvanometer mirror 20.

  In the first embodiment, the method of driving the carbano mirror 20 is used as the scanning method in the y-axis direction. However, in the present embodiment, instead of scanning in the y-axis direction by the carbano mirror 20, a cylindrical lens 33 is used. Employs optical expansion in the y-axis direction.

  The cylindrical lens 33 is a normal lens in the direction in which it functions as a lens and the cross section in the plane including the optical axis, and this cross-sectional shape is the same regardless of the position in the direction in which it does not function as a lens. The cylindrical lens 33 is arranged so that the direction of functioning as a lens is the y direction. That is, the light spread in the y direction by the cylindrical lens 33 is distributed and irradiated in the y direction of the sample 22 (the y direction on the cylindrical lens 33 and the y direction of the sample 22 are optically the same direction). And not necessarily in the same spatial direction). The cylindrical lens 33 is a y-direction light expansion means. The cross section of the measurement light is linear along the y-axis direction.

  Note that the same function as that of the cylindrical lens 33 can be realized by using a cylindrical mirror.

  Since the measurement light is light that is spatially expanded in the y-axis direction, when this light is guided by an optical fiber, the optical fiber 18 is an optical fiber in which cross sections are bundled on a one-dimensional line, or a cross section. Need to be an optical fiber bundled in a two-dimensional circle.

  Further, the direction of the groove of the diffraction element 25 is preferably the y-axis direction.

  Since the measurement light 28 is distributed and irradiated in the y-axis direction of the sample 22 to be measured, a cross-section of the sample 22 to be measured in the y-axis direction can be seen by the CCD camera 26 one-shot without mechanical scanning in the y-axis direction. Can be obtained at Therefore, the carbano mirror 20 can obtain the three-dimensional structure of the sample 22 to be measured simply by scanning in the x-axis direction.

  As a result, the apparatus is simple and inexpensive, and an FD-OCT apparatus applicable to dental measurement is obtained.

  Note that the FD-OCT apparatus according to the present embodiment is preferably used for dental measurement, but is not limited to dental measurement, and can be used for measurement in other fields. In the present embodiment, the FD-OCT apparatus has been described. However, the FD-OCT apparatus is not necessarily required, and a conventional OCT apparatus may be used. In other words, the CCD camera 26 may be a photo detector using a simple mirror instead of the diffraction element 25.

(Embodiment 3)
FIG. 3 is a diagram illustrating an example of a configuration of the FD-OCT apparatus according to the third embodiment. In FIG. 3, the same parts as those in the FD-OCT apparatus shown in FIG.

  The FD-OCT apparatus shown in FIG. 3 is different from the FD-OCT apparatus shown in FIG. 1 in that a light source light polarization operation device 35, a reference light polarization operation device 36, and an interference light polarization operation device 37 are provided. .

  FIG. 4A is a schematic diagram showing the configuration of the light source light polarization operation device 35. In FIG. 4, a polarizer 67 is a member that allows only a specific polarization component to pass, and a half-wave plate 68 and a quarter-wave plate 69 respectively set the wavelength of the transmitted light to a half wavelength, or 1 / It is a member shifted by 4 wavelengths. First, the polarizer 67 imparts basic polarization characteristics to the light source light or the measurement light. Further, the direction of polarization can be manipulated by rotating the half-wave plate 68 and the quarter-wave plate 69 around the optical axis at an appropriate angle. By using these wave plates 68 and 69, the polarization state of the light source light or the measurement light can be arbitrarily set.

  FIG. 4B is a schematic diagram showing the configuration of the reference light polarization manipulator 36 or the interference light polarization manipulator 37. These are constituted by the half-wave plate 70 and the quarter-wave plate 71, and the polarization state of the reflected light can be examined by adjusting these angles.

  In general, the polarization state of light can be represented by a four-component vector (four-dimensional vector) Si. When light in a certain polarization state enters the object, the polarization state of the transmitted light and reflected light changes due to the interaction with the object. That is, the four-dimensional vector S0 representing the polarization state of the reflected light is different from the four-dimensional vector Si representing the polarization state of the incident light. Therefore, the “characteristic that changes the polarization state (for example, birefringence characteristic)” of the object can be expressed by a 4 × 4 matrix M (Muller matrix). That is, when light having a polarization state represented by a vector Si is incident on a material having birefringence characteristics represented by a matrix M, the vector S0 representing the polarization state of light emitted from the material is S0 = M * It is calculated | required by Si.

  Therefore, in order to measure the matrix M representing the birefringence characteristics of a substance, the substance is passed through light having a polarization state represented by any four vectors, and the four vector components of the light after passing are detected. do it. This Mueller matrix can be measured at each measurement point of the object.

  In the present embodiment, the light source light polarization operation unit 35 in the measurement optical path is operated so that the measurement light becomes at least four types of independent polarization states, and the reference light polarization operation unit 36 or the interference light polarization operation unit 37 is operated. By operating and observing the interference light due to the four vector components in these polarization states, 16 types of Mueller matrix images can be obtained from these relationships. This is an image representing “characteristic that changes the polarization state = characteristic unique to the subject” of each part of the tomographic image of the subject.

  Note that only one of the reference light polarization manipulator 36 and the interference light polarization manipulator 37 may be provided as necessary.

  Since living tissues including oral tissues have unique polarization characteristics and birefringence characteristics, the birefringence characteristics of tooth buds and periodontal tissues can be detected according to this embodiment. Collagen in particular has a large birefringence characteristic, and for example, it becomes possible to discriminate between enamel that does not contain collagen and dentin that contains a large amount of collagen. In addition, it is possible to acquire 16 types of images reflecting the unique polarization characteristics and birefringence characteristics of each part of the oral tissue, and it is possible to visualize not only normal tissues but also lesion tissues such as caries and congestion.

  In addition, although this Embodiment demonstrated the example in the case of providing a polarization controller in the FD-COT apparatus shown in FIG. 1, this invention is not limited to this. For example, the present invention can also be applied to the FD-OCT apparatus in Embodiment 2 shown in FIG.

(Embodiment 4)
Since the OCT apparatus according to the fourth embodiment can apply the OCT apparatus according to the first to third embodiments or the conventional OCT apparatus to the parts other than the parts described below, the description thereof is omitted.

  FIG. 6 is a diagram showing a configuration of a light source of the OCT apparatus in the present embodiment. For example, the light source 16 of the FD-OCT apparatus shown in FIG. 1 is a single single wavelength light source. On the other hand, in the present embodiment, two or more light sources 56a, 56b, and 56c having different wavelengths are provided instead of the light source 16. The light sources 56a, 56b, and 56c are preferably superluminescent diodes in the wavelength range of 300 nm to 3000 nm. Of the light sources 56 a, 56 b and 56 c, the light source used for measurement is switched by driving the rotating mirror 57. That is, the light sources 56a, 56b, and 56c are arranged at positions corresponding to specific angles of the rotating mirror 57. A galvanometer mirror can be used as the rotating mirror 57.

  By the way, in the oral tissue, the oral lesion tissue, or the oral prosthesis, the wavelength dependency of the light absorption coefficient, transmission coefficient, and reflection coefficient varies. For example, in light having a wavelength near 800 nm, the transmission coefficient of cementum and alveolar bone is high, and the reflection coefficient of enamel and dentin is relatively large. In addition, soft tissue such as gingiva is highly permeable to light having a wavelength of around 1300 nm, so using such light as a light source is ideal for observing the alveolar bone under the gingiva and the deep gingival phototissue. It is. Further, since carious tissue has different fluorescence characteristics in the visible light region from normal tissues, it is necessary to use a light source that matches the wavelength of the fluorescence. Therefore, it has been difficult to visualize the entire structure of the oral tissue with a conventional OCT apparatus using a single wavelength light source.

  The OCT apparatus according to the present embodiment includes two or more light sources 56a, 56b, and 56c having different wavelengths. By appropriately selecting the wavelength of the light source, the OCT apparatus can be used in various ways such as oral tissues and oral lesion tissues. It is possible to visualize the fine structure of a substance having a light absorption coefficient, transmission coefficient, and reflection coefficient.

(Embodiment 5)
FIG. 7 is a diagram illustrating an example of a configuration of the FD-OCT apparatus according to the fifth embodiment. In FIG. 7, the same parts as those in the FD-OCT apparatus shown in FIG. The present embodiment can be applied not only to the FD-OCT apparatus shown in FIG. 2 but also to the FD-OCT apparatuses according to the first to third embodiments or a conventional OCT apparatus.

  The FD-OCT apparatus shown in FIG. 7 is different from the FD-OCT apparatus shown in FIG. 2 in that a pilot light source 59 and half mirrors 58 and 60 are provided.

  The pilot light source 59 is provided for irradiating the pilot light projected onto the sample 22 to be measured. The pilot light is light emitted to the imaging range so that an operator who performs measurement using the OCT apparatus can check the imaging region and imaging range during imaging or before and after imaging. The pilot light emitted from the pilot light source 59 is preferably visible light.

  The pilot light emitted from the pilot light source 59 is guided on the same optical axis as the light source light emitted from the light source 16 by the half mirror 58. The pilot light is projected onto the measurement sample 22 together with the light source light and the measurement light. This projection can be visually observed, and the same place as the measurement light is irradiated, so that the operator can recognize the measurement range.

  The cross section of the pilot light may be point-like or linear. In the case of pilot light having a point-like cross section, the pilot light is preferably arranged with its optical axis on the central optical axis of the measurement light. If the cross section is linear pilot light, the cross section of the pilot light is preferably arranged along the y direction of the measurement light. Since the OCT apparatus shown in FIG. 7 adopts a cylindrical lens 33, the pilot light passes through the cylindrical lens 33 and is irradiated to the measurement sample 22 in a cross-sectional line shape.

  The pilot light reflected by the sample 22 to be measured passes through the beam splitter 34 together with the measurement light, and is irradiated to the diffraction element 25 together with the interference light. Since a half mirror 60 that reflects at least visible light and transmits at least infrared light is provided in front of the diffraction element 25, interference light passes through the diffraction element 25, and pilot light is reflected by the half mirror 60. It is detected by the CCD camera 26.

  The CCD camera 26 serves as both a 2D-CCD camera having a visible light sensitivity band as a two-dimensional imaging device and a 2D-CCD camera as a two-dimensional photodetector used in FD-OCT employing a cylindrical lens. It is preferable. Thereby, the visible light image of a tooth is obtained by the CCD camera 26 with the interference spectral image by FD-OCT.

  Further, a 2D-CCD camera having a sensitivity band in visible light as a two-dimensional imaging device is provided separately from the CCD camera 26, and a 2D- having a sensitivity band in visible light by using a half mirror 60 or the like from the measurement optical axis. By guiding pilot light to the CCD camera, a visible light image of the sample 22 to be measured can be obtained. Note that a 2D-CCD camera having a sensitivity band in visible light can use a two-dimensional imaging device having a sensitivity band in the range of 300 nm to 3000 nm.

  In addition, a liquid crystal shutter may be interposed between the half mirror 60 and the diffraction element 25 so that only a visible light image is captured except during OCT imaging. Further, the half mirror 60 may be driven by a rotating or linear driving mechanism so that the pilot light is separated from the optical system during imaging by OCT.

  The optical fiber 18 is preferably composed of an image fiber capable of transmitting an image.

  Further, the computer 27 can monitor an image of the measurement site by pilot light. Further, a visible light image of the measurement site by pilot light can be recorded in synchronization with the tomographic measurement image.

(Embodiment 6)
As for the OCT apparatus according to the sixth embodiment, the FD-OCT apparatus according to the first to third embodiments or the conventional OCT apparatus can be applied to the parts other than the parts described below, and the description thereof is omitted.

  FIG. 8 is a diagram showing a configuration near the reference mirror of the OCT apparatus in the present embodiment. In the OCT apparatus according to the present embodiment, the phase of the reference light is changed by inserting a phase modulation element on the reference optical path or moving the reference mirror in the optical axis direction. FIG. 8A is a diagram showing a configuration example of an apparatus when a phase modulation element is inserted on the reference optical path. In FIG. 8A, a phase modulation element 62 is inserted in front of the reference mirror 24. The phase modulation element 62 is driven by an electrical drive signal. As the phase modulation element 62, for example, a rapid scanning optical delay line (RSOD), an acousto-optic element, an electro-optic element, or the like is preferably used.

  FIG. 8B is a diagram illustrating a configuration example of the apparatus when the phase of the reference light is changed by moving the reference mirror in the optical axis direction. In FIG. 8B, the reference mirror 24 is provided with a piezoelectric element 63. The piezoelectric element 63 is driven by an electrical drive signal. As the piezoelectric element 63 vibrates in the optical axis direction of the reference light, the reference mirror 24 is vibrated in the same direction as the optical axis direction of the reference light. As a result, the phase of the reference light is changed.

  According to the present embodiment, since the phase of the reference light can be changed by the phase modulation element 62 or the piezoelectric element 63, the intensity of the five sets of diffraction spectral interference light in which the phase of the reference light is shifted by 90 degrees, for example. Distribution can be obtained. If the shape of the sample to be measured in the depth direction (z-axis direction) is measured using this diffraction spectral interference light intensity distribution, the measurement range in the depth direction can be doubled. The principle will be described in detail below.

  In general, the measurement range of FD-OCT is determined in principle by the resolution of the diffraction element, the objective lens, and the CCD camera, and as a result, the measurement range in the depth direction is determined. In FD-OCT, the light intensity distribution (one-dimensional or two-dimensional) on the ξ axis of a diffraction element obtained by a CCD is Fourier-transformed by a computer to convert it into a distribution on the time t axis (that is, the depth of the sample to be measured). converted into a reflection characteristic distribution on the z-axis). In this case, since the light intensity distribution is a power spectrum, the result of inverse Fourier transform is that the complex conjugate signal of the autocorrelation of the reference light and the cross correlation of the reference light and the z-direction object reflected light is based on the distribution in the depth z-axis direction. The image is superimposed as an image (artifact) caused by a defect in the apparatus regardless of the subject. For this reason, it is assumed that the measurement of the diffraction spectral interference image on the diffractive element can measure not only the light intensity distribution but also the phase distribution of the light. The measurement range is halved.

  Measuring the phase of diffracted spectrum interference light directly is a phenomenon that is too fast (a phenomenon that is less than a few femtoseconds when the wavelength of light is divided by the speed of light). not exist. Therefore, the phase of the diffracted spectral interference light is indirectly measured by modulating the spatial phase instead of the temporal phase phenomenon. That is, the phase modulation element 62 or the piezoelectric element 63 can obtain five sets of diffraction spectral interference light intensity distributions in which the phase of the reference light is shifted by 90 degrees, for example. This diffraction spectral interference light intensity distribution is subjected to inverse complex Fourier transform in the computer 27, whereby a complex conjugate signal of the autocorrelation of the reference light and the cross-correlation of the reference light and the sample reflected light (measurement light) in the z direction is obtained. Artifacts are removed, and the original measurement range in the depth direction of FD-OCT is realized.

  In addition, by appropriately selecting the phase modulation frequency and synchronously detecting the detection signal of the CCD, it is possible to remove noise and improve the resolution, and further expand the measurement range.

(Embodiment 7)
FIG. 9 is a diagram illustrating an example of a configuration of the FD-OCT apparatus according to the seventh embodiment. In FIG. 9, the same parts as those in the FD-OCT apparatus shown in FIG. This embodiment can be applied not only to the FD-OCT apparatus shown in FIG. 1 but also to the FD-OCT apparatus according to the first to third embodiments or a conventional OCT apparatus.

  The FD-OCT apparatus shown in FIG. 9 is different from the FD-OCT apparatus shown in FIG. 1 in that an electrical modulator 64 is provided between the computer 27 and the light source 16.

  The calculator 27 sends a light amount modulation signal together with an ON / OFF signal to the electrical modulator 64. The electrical modulator 64 sends a light amount control signal based on the light amount modulation signal to the light source 16. The output light amount of the light source 16 is controlled by a light amount control signal output from the electrical modulator 64.

  The data detected by the CCD camera 26 is demodulated by the computer 27 in accordance with the light quantity modulation signal. This modulation / demodulation improves the S / N ratio of the detected data.

  Modulation and demodulation methods can be AM modulation and FM modulation, for example. Further, an optical modulator may be provided on the optical path of the light emitted from the light source 16 instead of the electrical modulator 64. Further, an optical modulator may be provided on the optical path of the measurement light (object reflected light) 28 or the optical path of the reference light 29. Further, a modulator that applies modulation in synchronization with the positions of the sample 22 to be measured and the reference mirror 24 may be provided.

  In general, the measurement range of the OCT apparatus is principally determined by the resolution of the objective lens, but in reality, the range is narrowed due to the influence of noise. That is, since the measurement light is attenuated as it enters the sample 22 to be measured, the z-direction object reflected light is buried in noise as the measurement sample becomes deeper. This narrows the measurement range in the depth direction.

  According to the present embodiment, since the light source light, the measurement light, or the reference light is modulated and the detection signal is detected, the S / N ratio is improved and the measurable range is expanded.

(Embodiment 8)
FIG. 10 is a diagram illustrating an example of a configuration of the FD-OCT apparatus according to the eighth embodiment. In FIG. 10, the same parts as those in the FD-OCT apparatus shown in FIG. This embodiment can be applied not only to the FD-OCT apparatus shown in FIG. 1 but also to the FD-OCT apparatus according to the first to third embodiments or a conventional OCT apparatus.

  The FD-OCT apparatus shown in FIG. 10 is different from the FD-OCT apparatus shown in FIG. 1 in that a nonlinear optical element 65 is provided on the optical path of the reference light 29 and a filter 66 is provided. . The nonlinear optical element 65 is an optical element that generates a harmonic in the vibration waveform of light, and is preferably, for example, a beta barium baudate.

  Further, a filter 66 is provided on the optical path of the interference light after passing through the fiber coupler 19 to cut the wavelength component of the light source light and pass the wavelength component of the half wavelength of the light source light.

  In general, living organisms emit more or less fluorescence, most of which is second harmonic fluorescence. In particular, in the tooth germ tissue, there are many cases in which this fluorescence characteristic changes due to lesions such as caries.

  According to the present embodiment, harmonics are generated in the vibration waveform of the reference light 29 by providing the nonlinear optical element 65 on the reference light. The second harmonic component in the reference harmonic and the second harmonic component of the measurement light 28 (z-direction object reflected light) are caused to interfere with each other by the fiber coupler 19. As a result, the fluorescence characteristic of the sample 22 to be measured can be detected more clearly. As a result, for example, the tomographic image discrimination of lesions such as caries is improved.

  The principle will be described in detail below.

  Living organisms have remarkable second-harmonic fluorescence characteristics due to two-photon absorption. This is the fluorescence emitted when electrons bound to biological constituent atoms receive energy equivalent to two photons of measurement light, jump to a high potential energy level, and then return to the original level. That is. In the case of a living body, the potential energy level of the second harmonic fluorescence is close to a continuous band, and the level exists in almost all wavelength bands. Further, what is characteristic of the second harmonic fluorescence is that fluorescence synchronized with the incident measurement light is generated, that is, coherency as an OCT apparatus is maintained. The second harmonic fluorescence is emitted from inside the sample 22 to be measured, and a part thereof returns as z-direction object reflected light (measurement light 28). On the other hand, since the nonlinear optical element 65 is provided on the path of the reference light 29, harmonics are generated in the vibration waveform of the reference light 29. By measuring the interference light between the reference light 29 and the z-direction object reflected light (measurement light 28), the fluorescence characteristics inside the sample 22 to be measured can be detected. Therefore, for example, it is effective for diagnosis of lesions such as caries accompanied by changes in fluorescence characteristics.

(Embodiment 9)
Since the FD-OCT apparatus or the conventional OCT apparatus according to the first to third embodiments can be applied to the OCT apparatus according to the ninth embodiment other than the parts described below, the description thereof is omitted.

  In the OCT apparatus, an image obtained by measurement is displayed on, for example, a monitor provided in the computer 27. However, if the OCT tomographic image of OCT is displayed as it is, it will give some uncomfortable images. In the OCT apparatus in the present embodiment, an easy-to-see image is provided by performing the following display.

  When displaying the image of the sample to be measured, it is preferable to display so that the light transmitting part that is the part where the measurement light enters and reaches the sample to be measured and the non-measurement light reaching part of the sample to be measured can be distinguished. .

  In addition, the one-shot measurement range in the OCT apparatus is smaller than the size of the teeth, and it is difficult to determine where the image was taken from which direction with only one image. Therefore, it is preferable to display a composite of a plurality of images.

  The distance in the depth direction (z-axis direction) of the sample to be measured in the OCT image is an optical distance, not an actual distance. For this reason, it is preferable to display the optical distance corrected to a spatial distance.

  Further, the deeper the object surface in the z direction is, the smaller the measured light quantity is, so the reflected light quantity in the z direction is also reduced. As a result, an image portion “dark” = “less reflected” is displayed on the image display. For this reason, it is preferable to display an image obtained by correcting the density in the depth direction based on the optical distance or the reflection amount integrated value.

  In addition, when displayed on a normal PC monitor screen, especially when enlarged display is performed, the measurement resolution is rougher than the resolution on the PC monitor screen, and the point display or point density display or rough floor is displayed. The key is displayed. Therefore, it is preferable to display a display in which the grayscale display based on the point density is corrected to a solid display.

  In addition, since an image obtained by measurement is a tomographic image of an object, it is difficult to grasp a spatial position and direction. Therefore, it is preferable to display a tomographic image in a three-dimensional manner on the pilot monitor image. Moreover, it is preferable to provide a user interface that allows an observer to arbitrarily select a display section.

  Moreover, the captured image may contain noise. Therefore, it is preferable to display a plurality of tomographic images or images obtained by integrating and averaging the plurality of images to remove temporal noise and improve the spatial resolution. Alternatively, an image obtained by integrating and averaging a plurality of tomographic images obtained by scanning in the vertical direction (x direction) and removing spatial noise in the x direction may be displayed. Further, an image obtained by integrating and averaging a plurality of tomographic images and removing noise in the x and / or y directions may be displayed.

  It can also be taken and displayed with an OCT device in combination with an intraoral camera image.

  In addition, in FIGS. 1-11 referred in order to demonstrate the said embodiment, what was represented in the figure and the ratio of the magnitude | size and length of a focal distance do not represent the ratio of an actual thing exactly | strictly. .

  INDUSTRIAL APPLICABILITY The present invention can be used particularly in the dental field as an optical coherence tomography apparatus capable of high-speed measurement and having a simple structure and is inexpensive.

1 is a diagram illustrating an example of a configuration of a Fourier domain optical coherence tomography apparatus (hereinafter referred to as an FD-OCT apparatus) in Embodiment 1. FIG. 10 is a diagram illustrating an example of a configuration of an FD-OCT apparatus according to Embodiment 2. FIG. 10 is a diagram illustrating an example of a configuration of an FD-OCT apparatus in Embodiment 3. FIG. (A) is the schematic which shows the structure of the light source light polarization operation device 35. FIG. FIG. 6B is a schematic diagram illustrating the configuration of the reference light polarization operation device 36 or the interference light polarization operation device 37. It is a conceptual diagram for demonstrating an example of the method of driving a lens. It is a figure which shows the structure of the light source of an OCT apparatus. 10 is a diagram illustrating an example of a configuration of an FD-OCT apparatus in Embodiment 5. FIG. FIG. 10 is a diagram showing a configuration near a reference mirror of an OCT apparatus in a sixth embodiment. FIG. 20 is a diagram illustrating an example of a configuration of an FD-OCT apparatus in a seventh embodiment. 218 is a diagram illustrating an example of a configuration of an FD-OCT apparatus in Embodiment 8. [FIG. It is a figure which shows the structure of the conventional OCT apparatus.

Explanation of symbols

1, 100, 101, 103, 104, 105, 106 OCT unit 2, 16, 56 Light source 3, 17, 23, 46 Lens 4, 34 Beam splitter 5, 28 Measuring light 6, 29 Reference light 8, 20, 49 Galvano Mirror 9, 12, 21 Objective lens 10, 22 Sample to be measured 13, 24 Reference mirror 14 Photo detector 18 Optical fiber 19 Fiber coupler 25 Diffraction element 26 CCD camera 27 Computer 30 Condensing lens 31 Linear actuator 32 Rotating shaft 33 Cylindrical lens 35 Light source light polarization controller 36 Reference light polarization controller 37 Interference light polarization controller 57 Rotating mirror 58, 60 Half mirror 59 Pilot light source 62 Phase modulation element 63 Piezoelectric element 64 Electrical modulator 65 Nonlinear optical element 66 Filter 201, 202 Measuring head

Claims (4)

A light source;
A light splitting unit that divides the light source light emitted from the light source into reference light for irradiating a reference mirror and measurement light for irradiating a sample to be measured;
A non-linear optical element which is provided on the optical path of the reference light and generates a harmonic component in the reference light;
An interference unit that interferes with the measurement light reflected by the sample to be measured and the reference light that is reflected by the reference mirror and includes a harmonic component generated by the nonlinear optical element;
A light detector for measuring the interference light;
Based on the interference light measured by the light detection unit, by calculating a reflectance characteristic of the harmonic component in the depth direction of the sample to be measured, that is, the irradiation direction of the measurement light, the depth direction of the sample to be measured optical coherence tomography device and a calculation unit for obtaining the information. Before SL calculation section by inverse Fourier transform of the spectrum obtained by spectroscopy of the interference light to obtain the information in the irradiation direction of the object to be measured sample, the optical coherence tomography apparatus according to claim 1. A cylindrical lens or a cylindrical mirror that makes the cross-section of the light source light linear along one axial direction in a plane perpendicular to the irradiation direction of the light source light;
A diffraction element that separates the interference light,
The interference unit causes the measurement light reflected by the sample to be measured and the reference light reflected by the reference mirror to interfere with each other to form linear interference light along the one-axis direction,
The optical coherence tomography apparatus according to claim 2, wherein the light detection unit is a two-dimensional image sensor that images the linear spectrum of the interference light dispersed by the diffraction element in one shot.   The optical coherence tomography apparatus according to claim 1, further comprising an optical fiber that guides the measurement light between the light splitting unit and the sample to be measured.

Images