Ophthalmic devices

Abstract

To provide a technology which can efficiently perform a plurality of kinds of measurements for an eye to be examined.SOLUTION: An ophthalmologic apparatus includes: a first light source for outputting a first light radiated onto an eye to be examined; a second light source for outputting a second light radiated onto the eye to be examined; a first interferometer for performing an examination for a first range of the eye to be examined on the basis of a first interference light acquired from a reflection light of the first light; and a second interferometer for performing an examination for a second range of the eye to be examined different from the first range on the basis of a second interference light acquired from a reflection light of the second light. A center wavelength of the first light is different from a center wavelength of the second light. A first light path, which is a light path of the first light, is overlapped with at least a part of a second light path, which is a light path of the second light. The first range and the second range can be examined simultaneously.SELECTED DRAWING: Figure 1

Description

[Technical Field]

[0001] The present disclosure relates to an ophthalmic apparatus, and more particularly to an ophthalmic apparatus capable of performing multiple types of measurements on a subject's eye. [Background technology]

[0002] Ophthalmic devices have been developed that measure the refractive power of a subject's eye and the morphology of each part (e.g., the anterior segment, fundus, etc.). By comprehensively assessing the refractive power and the morphology of each part of the subject's eye, it is possible to accurately diagnose diseases of the subject's eye and test its visual function. To achieve this, multiple types of measurements are required for each part of the subject's eye or the entire subject's eye, and ophthalmic devices capable of performing multiple types of measurements have been developed. For example, Patent Document 1 discloses an ophthalmic device that measures the corneal shape, refractive power, axial length, etc. of the subject's eye. The ophthalmic device of Patent Document 1 measures the anterior corneal shape using a keratometry ring and measures the axial length using optical interference. Patent Document 2 also discloses an ophthalmic device that measures the anterior segment and fundus using optical coherence tomography (OCT). The ophthalmic device of Patent Document 2 splits light output from a single light source into two light beams with different wavelengths using a beam splitter. One of the split beams is used for anterior segment OCT measurement, and the other is used for fundus OCT measurement. [Prior art documents] [Patent documents]

[0003] [Patent Document 1] JP 2016-77774 A [Patent Document 2] Japanese Patent Application Publication No. 2017-502817 Summary of the Invention [Problem to be solved by the invention]

[0004] Conventional ophthalmic devices are capable of performing multiple types of measurements, but because each measurement is performed separately, the measurement time is long and the measurement efficiency of the subject's eye is insufficient. This causes problems such as an increased burden on the subject and a decrease in measurement accuracy due to changes in the condition of the subject's eye during measurement. This specification provides a technology that can efficiently perform multiple types of measurements on the subject's eye. [Means for solving the problem]

[0005] The ophthalmologic apparatus disclosed in the present specification includes a first light source that outputs a first light to be irradiated onto an eye to be examined, a second light source that outputs a second light to be irradiated onto the eye to be examined, a first interferometer that performs an inspection of a first range of the eye to be examined based on a first interference light obtained from a reflected light of the first light, and a second interferometer that performs an inspection of a second range of the eye to be examined that is different from the first range based on a second interference light obtained from a reflected light of the second light. The first light and the second light have different central wavelengths, and the first optical path, which is an optical path of the first light, and the second optical path, which is an optical path of the second light, at least partially overlap, so that the inspection of the first range and the inspection of the second range can be performed simultaneously.

[0006] In the above-described ophthalmic apparatus, the first light and the second light have different central wavelengths, allowing the first light and the second light to irradiate different regions of the subject's eye with strong light. Therefore, by setting the first light and the second light to have central wavelengths suitable for the first range and the second range, respectively, the first range and the second range can be accurately inspected. Furthermore, at least a portion of the first optical path, which is the optical path of the first light, and the second optical path, which is the optical path of the second light, overlap, allowing the first light and the second light to be simultaneously irradiated onto the subject's eye. Thus, the above-described ophthalmic apparatus allows simultaneous inspection of different regions of the subject's eye using the first interferometer and the second interferometer. In other words, the above-described ophthalmic apparatus allows efficient performance of multiple types of measurements on the subject's eye (i.e., inspection of the first range and inspection of the second range). Therefore, the measurement time can be shortened compared to conventional methods, and inspections of the subject's eye can be performed on subjects in substantially the same condition. [Brief explanation of the drawings]

[0007] [Figure 1] 1 is a schematic configuration diagram of an anterior segment OCT interferometer and a fundus OCT interferometer of an ophthalmic apparatus according to a first embodiment. [Figure 2] FIG. 2 is a schematic configuration diagram of a probe optical system of the ophthalmic apparatus according to the first embodiment. [Figure 3] FIG. 2 is a diagram schematically illustrating the optical path of light in an anterior segment OCT optical system. [Figure 4] FIG. 2 is a diagram schematically illustrating the optical path of light in a fundus OCT optical system. [Figure 5] FIG. 2 is a diagram showing a control system of the ophthalmologic apparatus according to the first embodiment. [Figure 6] 4 is a flowchart showing a process of performing an examination of an eye to be examined using the ophthalmologic apparatus of the first embodiment. [Figure 7] 1 is a flowchart showing the process of OCT measurement. [Figure 8] 1A and 1B are diagrams for explaining a procedure for processing an interference signal waveform obtained by an anterior segment OCT interferometer. [Figure 9]9 is a diagram for explaining a procedure for scanning the incident position of light on the subject's eye within a predetermined range and identifying the position of each part of the subject's eye from information obtained for each incident position (information obtained by the procedure shown in FIG. 8). [Figure 10] FIG. 1 is a diagram for explaining a radial scan method in anterior segment OCT measurement. [Figure 11] 1A and 1B are diagrams for explaining a procedure for processing an interference signal waveform obtained by a fundus OCT interferometer. [Figure 12] FIG. 1 is a diagram for explaining the derivation of scan resolution in OCT measurement. [Figure 13] FIG. 1 is a diagram for explaining the derivation of optical resolution in OCT measurement. [Figure 14] FIG. 13 is a diagram for explaining a procedure for scanning the incident position and incident angle of light on the test eye within a predetermined range and identifying the position of each part of the test eye from information obtained for each incident position and incident angle (information obtained by the procedure shown in FIG. 12). [Figure 15] 10A and 10B are diagrams for explaining correction of the distance between the anterior segment and the fundus. [Figure 16] 10A and 10B are diagrams for explaining acquisition of a peak indicating the depth position of a calibration mirror in anterior segment OCT measurement. [Figure 17] 10A and 10B are diagrams for explaining the acquisition of a peak indicating the depth position of a calibration mirror in fundus OCT measurement. [Figure 18] FIG. 10 is a schematic configuration diagram of an anterior segment OCT interferometer and a fundus OCT interferometer of an ophthalmic apparatus according to a second embodiment. [Figure 19] FIG. 10 is a diagram showing a control system of the ophthalmologic apparatus according to the second embodiment. [Figure 20] FIG. 1 shows the configuration of a sampling trigger / clock generator. DETAILED DESCRIPTION OF THE INVENTION

[0008] The main features of the embodiments described below are listed below. Note that the technical elements described below are independent technical elements that exhibit technical utility alone or in various combinations, and are not limited to the combinations described in the claims at the time of filing.

[0009] In an embodiment of the present technology, the first area may include an anterior segment of the subject's eye, and the second area may include a fundus of the subject's eye. With this configuration, it is possible to suitably examine a site of the subject's eye that is important in an ophthalmic examination.

[0010] In one embodiment of the present technology, the first light output from the first light source may have a central wavelength of 0.95 μm or more and 1.80 μm or less, and the second light output from the second light source may have a central wavelength of 0.40 μm or more and 1.15 μm or less. With this configuration, it is possible to irradiate light with wavelengths more suitable for each of the first range and the second range. Therefore, it is possible to accurately inspect the first range and the second range.

[0011] In one embodiment of the present technology, the first optical path and the second optical path may be arranged on an overlapping optical path, and the ophthalmic apparatus may further include a scanner that scans the first light output from the first light source and the second light output from the second light source. With this configuration, the scanner for scanning the first light irradiated onto the first range and the scanner for scanning the second light irradiated onto the second range can be shared. This makes it possible to avoid a complex configuration within the ophthalmic apparatus and reduce the number of components.

[0012] In one embodiment of the present technology, the second light source may be configured to output the second light in a pulsed manner when the scanning resolution for the inspection of the second range is greater than the optical resolution for the inspection of the second range. The inspection speed for the second range may be set slower than the inspection speed for the first range. That is, the time required for the inspection of the second range may be set longer than the time required for the inspection of the first range. In such a case, in a configuration in which the first light and the second light are simultaneously scanned by a shared scanner, if the scanner is scanned at the speed for the inspection of the first range, the relatively slow inspection speed for the second range may attenuate the interference signal obtained from the second interference light, making it difficult to inspect the second range accurately. In this way, when the scanning resolution in the scanning direction is greater than the optical resolution in the lateral direction, outputting the second light in a pulsed manner at the speed for the inspection of the second range can suppress attenuation of the interference signal obtained from the second interference light.

[0013] In one embodiment of the present technology, the second light source may be configured to continuously output the second light when the scanning resolution in the inspection of the second range is smaller than the optical resolution in the inspection of the second range. When the scanning resolution in the inspection of the second range is smaller than the optical resolution in the inspection of the second range, continuously outputting the second light makes it difficult for the interference signal to attenuate. Therefore, when the scanning resolution in the scanning direction in the inspection of the second range is smaller than the optical resolution in the lateral direction, adopting an aspect of continuously outputting the second light makes it easier to control the second light source.

[0014] In an embodiment of the present technology, the scanner may further include an objective lens disposed on the overlapping optical paths and disposed between the scanner and the subject's eye. The scanner may include a mirror that changes the traveling direction of the first light and the second light to a predetermined direction. The A-scan speed in the inspection of the first range is Fa, the scan range in the first direction perpendicular to the depth direction of the subject's eye in the inspection of the first range is Wa, the number of A-scans within the scan range in the first direction is Na, the magnification of the mirror in the inspection of the first range is Ma, the A-scan speed in the inspection of the second range is Fp, the scan range in the first direction in the inspection of the second range is Wp, the number of A-scans within the scan range in the first direction is Np, the magnification of the mirror in the inspection of the second range is Mp, the central wavelength of the second light is λp, the beam diameter of the second light is Dp, and the focal length of the objective lens is f obj , the second light source is expressed by the following formula: (Fp·λp·Mp) / Dp<(π / 4)·(Ma / f obj )·(Wa / Na)·Fa If the following condition is satisfied, the second light may be output in pulses. When such a condition is satisfied, the interference signal obtained from the second interference light is likely to attenuate. Therefore, by outputting the second light in a pulsed form, it is possible to suppress the attenuation of the interference signal.

[0015] In this specification, acquiring an interference signal containing positional information in the depth direction is referred to as an "A-scan." The term "A-scan" in this specification does not necessarily require dynamically changing a specific configuration in the ophthalmic device disclosed in this specification. Specifically, the term "A-scan" in this specification includes the following aspects. For example, the term "A-scan" includes an aspect in which, in swept-source OCT (SS-OCT), an interference signal containing positional information for each depth in the depth direction is acquired by sweeping the wavelength or wavenumber of light output from a light source; in spectral-domain OCT (SD-OCT), an interference signal containing positional information for each depth in the depth direction is acquired by dissolving a spectrum into wavelengths or wavenumbers; and in time-domain OCT (TD-OCT), an interference signal containing positional information for each depth in the depth direction of the subject's eye is acquired by changing the optical path length of a reference light. In this specification, the number of repetitions of A-scans performed per unit time is referred to as the "A-scan rate." For example, if 100,000 A-scans are performed per second, the A-scan rate is 100 kHz.

[0016] In an embodiment of the present technology, the scanner may further include an objective lens disposed on the overlapping optical paths and disposed between the scanner and the subject's eye. The scanner may include a mirror that changes the traveling direction of the first light and the second light to a predetermined direction. The A-scan speed in the inspection of the first range is Fa, the scan range in the first direction perpendicular to the depth direction of the subject's eye in the inspection of the first range is Wa, the number of A-scans within the scan range in the first direction is Na, the magnification of the mirror in the inspection of the first range is Ma, the A-scan speed in the inspection of the second range is Fp, the scan range in the first direction in the inspection of the second range is Wp, the number of A-scans within the scan range in the first direction is Np, the magnification of the mirror in the inspection of the second range is Mp, the central wavelength of the second light is λp, the beam diameter of the second light is Dp, and the focal length of the objective lens is f obj , the second light source is expressed by the following formula: (Fp·λp·Mp) / Dp>(π / 4)·(Ma / f obj)·(Wa / Na)·Fa The second light may be continuously output when the following condition is satisfied. When these conditions are satisfied, the interference signal obtained from the second interference light is less likely to attenuate, and therefore, by continuously outputting the second light, it is possible to easily control the second light source.

[0017] In one embodiment of the present technology, the second light source has a duty ratio D that satisfies the following formula: D <Fp / {π / (4λp)·(Dp / f obj )·(Ma / Mp)·(Wa / Na)·Fa} The second light may be configured to be output in pulses with a period 1 / Fp so as to satisfy the following: In such a configuration, by outputting light in accordance with the A-scan speed Fp of the subject's eye in the examination of the second range, attenuation of the interference signal obtained from the second interference light can be suitably suppressed. Note that the duty ratio can be expressed as to / (to+ts), for example, where to is the time during which the second light is output and ts is the time during which the second light is not output.

[0018] In an embodiment of the present technology, the scanner may further include an objective lens disposed on the overlapping optical paths and disposed between the scanner and the subject's eye. The scanner may include a mirror that changes the traveling direction of the first light and the second light to a predetermined direction. The A-scan speed in the inspection of the first range is Fa, the scan range in the first direction perpendicular to the depth direction of the subject's eye in the inspection of the first range is Wa, the number of A-scans within the scan range in the first direction is Na, the magnification of the mirror in the inspection of the first range is Ma, the A-scan speed in the inspection of the second range is Fp, the scan range in the first direction in the inspection of the second range is Wp, the number of A-scans within the scan range in the first direction is Np, the magnification of the mirror in the inspection of the second range is Mp, the central wavelength of the second light is λp, the beam diameter of the second light is Dp, and the focal length of the objective lens is f obj , the second interferometer has a duty ratio that satisfies the following formula: D <Fp / {π / (4λp)·(Dp / f obj)·(Ma / Mp)·(Wa / Na)·Fa} The second light may be detected in a pulsed manner with a period of 1 / Fp so as to satisfy the following condition. The inspection speed for the second range may be set slower than the inspection speed for the first range. That is, the time required for the second inspection may be set longer than the time required for the first range. In such a case, in a configuration in which the first light and the second light are simultaneously scanned by a shared scanner, if the scanner is scanned at the inspection speed for the first range, the inspection speed for the second range may be relatively slow, resulting in attenuation of the interference signal obtained from the second interference light, making it difficult to inspect the second range accurately. In the above configuration, by detecting the second interference light at an A-scan speed Fp for the inspection of the second range, attenuation of the interference signal obtained from the second interference light can be suppressed. Note that the duty ratio can be expressed, for example, as td / (td+tu), where td is the time during which the second light is detected and tu is the time during which the second light is not detected.

[0019] In one embodiment of the present technology, the ophthalmologic apparatus may further include a refractive power measurement optical system that measures the refractive power of the subject's eye using the second light output from the second light source. With this configuration, the light source used for testing the second range can be shared with the light source for measuring the refractive power of the subject's eye. This makes it possible to avoid a complex configuration within the ophthalmologic apparatus and reduce the number of components.

[0020] In one embodiment of the present technology, the first range may include an anterior segment of the subject's eye, and the second range may include a fundus of the subject's eye. The system may further include a computing device that calculates the shape of the anterior segment of the subject's eye based on the first interference light and calculates the shape of the fundus of the subject's eye based on the second interference light. The computing device may store a distance measured in advance from the first range to the second range, and may calculate the axial length of the subject's eye based on the calculated shape of the anterior segment and the shape of the fundus and the calculated distance measured in advance from the first range to the second range. With this configuration, multiple types of measurements of the subject's eye, including measurement of the axial length, can be performed with a single ophthalmic device.

[0021] In an embodiment of the present technology, the first interferometer may include a first calibration mirror, and the second interferometer may include a second calibration mirror. The calculation device may store a predetermined first reference position of the first calibration mirror, a predetermined second reference position of the second calibration mirror, and a distance from the first reference position to the second reference position. When calculating the shape of the anterior segment, the calculation device may calculate a first displacement amount between the measured position of the first calibration mirror and the first reference position. When calculating the shape of the fundus, the calculation device may calculate a second displacement amount between the measured position of the second calibration mirror and the second reference position. When calculating the axial length, the calculation device may correct the distance from the first range to the second range based on the first displacement amount and the second displacement amount. With this configuration, even if the difference between the optical path length of the first light in the first interferometer and the optical path length of the second light in the second interferometer fluctuates, the distance from the first range to the second range can be appropriately corrected, thereby enabling accurate calculation of the axial length. [Example]

[0022] An ophthalmic apparatus 1 according to Example 1 will be described below. As shown in Fig. 1, the ophthalmic apparatus 1 includes an anterior segment OCT interferometer 10 that takes tomographic images of the anterior segment of an eye E to be examined, and a fundus OCT interferometer 11 that takes tomographic images of the fundus of the eye E to be examined.

[0023] The anterior segment OCT interferometer 10 is used to capture tomographic images of the anterior segment of the subject's eye E using optical coherence tomography. The anterior segment OCT interferometer 10 employs a Fourier domain method in which light wave interference is performed in Fourier space. In particular, swept-source OCT (SS-OCT) is used to detect spectral interference signals using a wavelength-scanning light source that changes wavelength over time and performs scanning, thereby capturing tomographic images of the anterior segment of the subject's eye E. Examples of wavelength-scanning light sources that can be used include external resonator-type wavelength-swept light sources that use a wavelength-tunable filter such as a diffraction grating or prism, and various external resonator-type light sources that use a Fabry-Perot tunable filter with a variable resonator length. Other examples include a wavelength-tunable distributed Bragg reflector (DBR) laser and a wavelength-tunable vertical cavity surface-emitting laser (VCSEL) using a microelectromechanical systems (MEMS) mechanism. The shape of each part of the anterior segment of the subject's eye E (for example, the cornea, the anterior chamber, the crystalline lens, etc.) can be measured from the tomographic image captured by the anterior segment OCT interferometer 10. Note that the anterior segment OCT interferometer 10 is not limited to SS-OCT, and may be, for example, another OCT using the Fourier domain method (for example, spectral domain OCT) or a method other than the Fourier domain method (for example, a time domain method).

[0024] As shown in FIG. 1, the anterior segment OCT interferometer 10 includes an anterior segment light source 12, a measurement optical system 20, a calibration optical system 30, a reference optical system 40, and an interference optical system 50.

[0025] The anterior-segment light source 12 is a wavelength-swept light source, and the wavelength (wavenumber) of the output light changes periodically. The anterior-segment light source 12 outputs long-wavelength light, for example, light with a central wavelength of 0.95 μm or more and 1.80 μm or less. In this embodiment, the anterior-segment light source 12 outputs light with a central wavelength of 1.31 μm. Using long-wavelength light allows it to easily pass through strongly scattering tissues, such as opacified lenses, the ciliary body, the conjunctiva, and the sclera. Furthermore, because water absorbs light to a large extent, light does not easily reach the fundus, allowing for the irradiation of strong light. Therefore, by outputting light with a central wavelength of 0.95 μm or more from the anterior-segment light source 12, the degree of light reaching tissues composed of scattering substances can be increased. Furthermore, because light with a central wavelength of 0.95 μm or more and 1.80 μm or less is less dispersed by water, irradiating the subject's eye E with light in this range allows for the acquisition of high-quality anterior-segment OCT images. Furthermore, by emitting light having a central wavelength of 1.80 μm or less from the anterior-segment light source 12, the target site can be measured with high sensitivity using an indium-gallium-arsenide (InGaAs) light-receiving element. Therefore, by emitting light having a wavelength of 0.95 μm or more and 1.80 μm or less from the anterior-segment light source 12, a tomographic image of the anterior segment of the subject's eye E can be suitably captured.

[0026] The anterior-segment light source 12 is connected to a polarization control device 14 and an optical coupler 16. Therefore, the light output from the anterior-segment light source 12 is input to the optical coupler 16 via the polarization control device 14, and is split into measurement light and reference light at a ratio of, for example, 9:1 in the optical coupler 16, and then input to the measurement optical system 20 and the reference optical system 40, respectively.

[0027] The measurement optical system 20 includes an optical circulator 22, an optical coupler 24, and a probe optical system 26. The measurement light input from the anterior eye light source 12 to the measurement optical system 20 is input to the optical circulator 22. The measurement light input to the optical circulator 22 is input to the optical coupler 24, where it is split at a ratio of, for example, 99:1, and input to each of the probe optical system 26 and the calibration optical system 30.

[0028] 2, the probe optical system 26 includes an anterior segment OCT optical system 90 that performs tomographic imaging of the anterior segment of the subject's eye E, a fundus OCT optical system 92 that performs tomographic imaging of the fundus of the subject's eye E, a refraction measurement optical system 94 that measures the refractive power of the subject's eye E, an alignment optical system (not shown) that aligns the ophthalmic apparatus 1 to a predetermined positional relationship with respect to the subject's eye E, and an observation optical system (not shown) that observes the subject's eye E. Since the alignment optical system and the observation optical system can be those used in known ophthalmic apparatuses, detailed description thereof will be omitted. The fundus OCT optical system 92 and the refraction measurement optical system 94 will be described later.

[0029] The anterior segment OCT optical system 90 includes a fiber collimator 102, dichroic mirrors 104 and 110, a mirror 105 with a hole in the center (hereinafter referred to as the perforated mirror 105), a scanner 106, and an objective lens 112. Light input from the optical coupler 24 to the probe optical system 26 (i.e., the anterior segment OCT optical system 90) is emitted from the fiber collimator 102 and irradiated onto the dichroic mirror 104. The dichroic mirror 104 reflects light with wavelengths longer than 0.90 μm and transmits light with wavelengths shorter than 0.90 μm. The light output from the anterior segment light source 12 has a central wavelength of 1.31 μm and is therefore reflected by the dichroic mirror 104. The light reflected by the dichroic mirror 104 passes through the circular hole in the center of the perforated mirror 105 and is irradiated onto the scanner 106. The scanner 106 is, for example, a galvanometer, and the direction of light irradiation is changed to a predetermined direction by a galvanometer mirror 108 attached to the galvanometer. The light emitted from the scanner 106 is irradiated onto a dichroic mirror 110. The dichroic mirror 110 reflects light with wavelengths longer than 0.90 μm and transmits light with wavelengths shorter than 0.90 μm. Since the light emitted from the scanner 106 has a central wavelength of 1.31 μm, it is reflected by the dichroic mirror 110. Then, the light reflected by the dichroic mirror 110 is irradiated onto the anterior segment (e.g., cornea, anterior chamber, crystalline lens, etc.) of the subject's eye E via an objective lens 112. The measurement light reflected from the anterior segment of the subject's eye E is incident on the fiber collimator 102 via the objective lens 112, the dichroic mirror 110, the scanner 106, the aperture mirror 105, and the dichroic mirror 104. Then, as shown in FIG. 1, the measurement light is input again to the optical circulator 22 via the optical coupler 24. The measurement light input to the optical circulator 22 is input to the optical coupler 52 of the interference optical system 50.

[0030] As described above, the light separated by the optical coupler 24 of the measurement optical system 20 is input to the calibration optical system 30. As shown in FIG. 1, the calibration optical system 30 includes lenses 32, 35, and 38, a dichroic mirror 34, a mirror 39, and a calibration mirror 36. The light input from the optical coupler 24 to the calibration optical system 30 is emitted from a fiber collimator (not shown) and is irradiated onto the dichroic mirror 34 via the lens 32. The dichroic mirror 34 transmits light with wavelengths longer than 0.90 μm and reflects light with wavelengths shorter than 0.90 μm. The light emitted from the fiber collimator (not shown) has a center wavelength longer than 0.90 μm (i.e., 1.31 μm), and therefore is transmitted through the dichroic mirror 34. The light transmitted through the dichroic mirror 34 is incident on the calibration mirror 36 via the lens 35. The calibration light reflected by the calibration mirror 36 is again incident on the fiber collimator via the lens 35, the dichroic mirror 34, and the lens 32, and is input to the optical circulator 22 via the optical coupler 24. The calibration light input to the optical circulator 22 is input to the optical coupler 52 of the interference optical system 50. The lens 38 and the mirror 39 of the calibration optical system 30 are used in the fundus OCT interferometer 11, which will be described later.

[0031] On the other hand, the reference light separated by the optical coupler 16 is input to the reference optical system 40 as described above. The reference optical system 40 includes an optical circulator 42, lenses 44 and 47, an attenuator 46, and a reference mirror 48. The reference light input from the anterior-segment light source 12 to the reference optical system 40 is input to the optical circulator 42. The reference light input to the optical circulator 42 is emitted from a fiber collimator (not shown) and is incident on the reference mirror 48 via the lenses 44 and 47. The optical path length of the reference light output from the anterior-segment light source 12 is adjusted by a zero-point adjustment mechanism (not shown). Note that a zero-point adjustment mechanism used in a known ophthalmic device can be used, and therefore a detailed description thereof will be omitted. The reference light reflected by the reference mirror 48 is again incident on the fiber collimator and input to the optical circulator 42. The reference light input to the optical circulator 42 is input to an optical coupler 52 of the interference optical system 50.

[0032] The interference optical system 50 includes an optical coupler 52, a light-receiving element 54, and a signal processor 56. The optical coupler 52 combines the measurement light reflected from the subject's eye E with the reference light generated by the reference optical system 40, and the combined measurement interference light is input to the light-receiving element 54. The optical coupler 52 also combines the calibration light generated by the calibration optical system 30 with the reference light generated by the reference optical system 40, and the combined calibration interference light is input to the light-receiving element 54. The light-receiving element 54 may be, for example, an InGaAs-based element, and the light-receiving element 54 measures the interference between the measurement interference light and the calibration interference light for each wavelength. An interference signal corresponding to the intensity of the measured interference light is then input to the signal processor 56. The signal processor 56 samples the acquired interference signal. A known data acquisition device (so-called DAQ) may be used as the signal processor 56. The sampled interference signal is input to a calculation device 200, which will be described later. The arithmetic unit 200 performs processing such as Fourier transform on the interference signal, and a tomographic image of the anterior segment along the scanning line is obtained.

[0033] Scanning in the anterior segment OCT optical system 90 will be described with reference to FIG. 3. FIG. 3 shows the optical path along which light emitted from the fiber collimator 102 is irradiated to the subject's eye E. Only some optical components arranged on the optical path (i.e., the dichroic mirror 104, the scanner 106, and the objective lens 112) are illustrated, and other optical components are omitted. As shown in FIG. 3, in the anterior segment OCT optical system 90, the scanner 106 is arranged at the back focus of the objective lens 112. Therefore, the light scanned by the scanner 106 is irradiated to the subject's eye E parallel to the optical axis. That is, the anterior segment OCT optical system 90 performs telecentric scanning, and a distortion-free image can be obtained when tomographically imaging the subject's eye E. Furthermore, the end face of the fiber collimator 102 is arranged at a position conjugate with the anterior segment of the subject's eye E. This allows the light emitted from the fiber collimator 102 to be focused on the anterior segment of the subject's eye E. Therefore, the anterior segment OCT optical system 90 can suitably perform tomographic imaging of the anterior segment of the subject's eye E.

[0034] Next, the fundus OCT interferometer 11 will be described. The fundus OCT interferometer 11 is used to photograph the fundus of the subject's eye E by optical coherence tomography. The fundus OCT interferometer 11 employs a Fourier domain method that performs interference of light waves in Fourier space, and in particular, uses spectral domain OCT (SD-OCT) that detects spectral information using a fixed wavelength light source that outputs light of a wide wavelength band and a spectroscope to perform tomographic imaging of the fundus of the subject's eye E. From the tomographic image captured by the fundus OCT interferometer 11, the shape of each part of the fundus of the subject's eye E (e.g., retina, choroid, etc.) can be measured. Note that the fundus OCT interferometer 11 is not limited to SD-OCT and may be, for example, another OCT using the Fourier domain method (e.g., SS-OCT) or a method other than the Fourier domain method (e.g., a time domain method).

[0035] As shown in FIG. 1, the fundus OCT interferometer 11 includes a fundus light source 62, a measurement optical system 66, a calibration optical system 30, a reference optical system 70, and an interference optical system 80.

[0036] The fundus light source 62 is a fixed-wavelength light source. The fundus light source 62 outputs light having a different central wavelength from the light output from the anterior-segment light source 12. For example, the fundus light source 62 can output light having a central wavelength of 0.40 μm or more and 1.15 μm or less. Alternatively, the fundus light source 62 may output light having a half-width in a wavelength range different from the half-width wavelength range of the light output from the anterior-segment light source 12. In this embodiment, the fundus light source 62 outputs light having a central wavelength of 0.83 μm. Light having a central wavelength of 0.40 μm or more and 1.15 μm or less has high transmittance within the eye. Therefore, by outputting light having a central wavelength of 0.40 μm or more and 1.15 μm or less from the light source, the light can be sufficiently irradiated to the fundus of the subject's eye E. Furthermore, silicon-based light-receiving elements have high sensitivity to light having a central wavelength of 0.40 μm or more and 0.95 μm or less. Furthermore, light with a central wavelength of 0.95 μm or more and 1.15 μm or less is less dispersed by water, and therefore, a high-quality fundus OCT image can be obtained by irradiating light in this range onto the subject's eye E. Therefore, by outputting light with a central wavelength of 0.40 μm or more and 1.15 μm or less from the light source, a tomographic image of the fundus of the subject's eye E can be suitably captured.

[0037] An optical coupler 64 is connected to the fundus light source 62. Therefore, the light output from the fundus light source 62 is input to the optical coupler 64, where it is split into measurement light and reference light at a ratio of, for example, 9:1, and then input to the measurement optical system 66 and the reference optical system 70, respectively.

[0038] The measurement optical system 66 includes an optical coupler 68 and the probe optical system 26. The measurement light input from the fundus light source 62 to the measurement optical system 66 is input to the optical coupler 68, where it is split at a ratio of, for example, 99:1, and input to each of the probe optical system 26 and the calibration optical system 30.

[0039] As shown in FIG. 2 , the fundus OCT optical system 92 of the probe optical system 26 includes a fiber collimator 114, lenses 116, 118, and 120, dichroic mirrors 104 and 110, a perforated mirror 105, a scanner 106, mirrors 122, 124, and 126, and an objective lens 112. Light input from the optical coupler 68 to the probe optical system 26 (i.e., the fundus OCT optical system 92) is emitted from the fiber collimator 114 and irradiated onto the dichroic mirror 104 via the lenses 116 and 118. As described above, the dichroic mirror 104 transmits light with wavelengths shorter than 0.90 μm, and therefore the light output from the fundus light source 62 (with a central wavelength of 0.83 μm) is transmitted through the dichroic mirror 104. Here, the optical path of the fundus OCT optical system 92 overlaps with the optical path of the anterior segment OCT optical system 90.

[0040] The light transmitted through the dichroic mirror 104 passes through a circular hole in the center of the perforated mirror 105 and is irradiated onto the scanner 106. The light irradiated onto the scanner 106 has its irradiation direction changed to a predetermined direction and is then irradiated onto the dichroic mirror 110. As described above, the dichroic mirror 110 transmits light with wavelengths shorter than 0.90 μm, so the light transmitted through the lens 118 (with a central wavelength of 0.83 μm) is also transmitted through the dichroic mirror 110. Here, the optical path of the fundus OCT optical system 92 is again different from the optical path of the anterior segment OCT optical system 90.

[0041] The light transmitted through the dichroic mirror 110 is reflected by mirrors 122, 124, and 126 via a lens 120. The light reflected by the mirror 126 is again irradiated onto the dichroic mirror 110 and, as described above, is transmitted through the dichroic mirror 110. Here again, the optical path of the fundus OCT optical system 92 overlaps with the anterior segment OCT optical system 90. The light transmitted through the dichroic mirror 110 is irradiated onto the fundus (e.g., retina, choroid, etc.) of the subject's eye E via an objective lens 112.

[0042] As described above, the optical path of the fundus OCT optical system 92 overlaps with the optical path of the anterior segment OCT optical system 90 from the dichroic mirror 104 via the scanner 106 to the dichroic mirror 110, and is different from the optical path of the anterior segment OCT optical system 90 until it is irradiated from the dichroic mirror 110 via the lens 120 back to the dichroic mirror 110, and overlaps with the optical path of the anterior segment OCT optical system 90 from the dichroic mirror 110 via the objective lens 112 to the subject's eye E.

[0043] The measurement light reflected from the fundus of the subject's eye E is incident on the fiber collimator 114 via the objective lens 112, the dichroic mirror 110, mirrors 126, 124, and 122, the lens 120, the dichroic mirror 110, the scanner 106, the apertured mirror 105, the dichroic mirror 104, and lenses 118 and 116. Then, as shown in FIG. 1 , the measurement light is input to the optical coupler 64 again via the optical coupler 68. The measurement light input to the optical coupler 64 is input to the interference optical system 80.

[0044] As described above, the light separated by the optical coupler 68 of the measurement optical system 66 is input to the calibration optical system 30. As shown in FIG. 1 , the light input to the calibration optical system 30 is emitted from a fiber collimator (not shown) and irradiated onto the dichroic mirror 34 via the lens 38 and the mirror 39. As described above, the dichroic mirror 34 reflects light with wavelengths shorter than 0.90 μm, so the light (with a central wavelength of 0.83 μm) transmitted through the mirror 39 is reflected by the dichroic mirror 34. The light reflected by the dichroic mirror 34 is incident on the calibration mirror 36 via the lens 35. The calibration light reflected by the calibration mirror 36 is again incident on the fiber collimator via the lens 35, the dichroic mirror 34, the mirror 39, and the lens 38, and is input to the interference optical system 80 via the optical couplers 68 and 64.

[0045] On the other hand, the reference light separated by the optical coupler 64 is input to the reference optical system 70 as described above. The reference optical system 70 includes lenses 72 and 74 and a reference mirror 76. The reference light input to the reference optical system 70 from the fundus light source 62 is emitted from a fiber collimator (not shown) and is incident on the reference mirror 76 via the lenses 72 and 74. The optical path length of the reference light output from the fundus light source 62 is adjusted by a zero-point adjustment mechanism (not shown). The reference light reflected by the reference mirror 76 is again incident on the fiber collimator and is input to the interference optical system 80 via the optical coupler 64.

[0046] The interference optical system 80 includes lenses 82 and 88, a diffraction grating 84, a prism 86, and a light-receiving element 89. In the interference optical system 80, the measurement light reflected from the subject's eye E and the reference light generated by the reference optical system 70 are combined, and the combined measurement interference light is input to the light-receiving element 89. In addition, the calibration light generated by the calibration optical system 30 and the reference light generated by the reference optical system 70 are combined, and the combined calibration interference light is input to the light-receiving element 89. Specifically, the measurement interference light and the calibration interference light are emitted from a fiber collimator (not shown) and pass through the diffraction grating 84 via the lens 82. This separates the interference light into wavelength spectra. Then, by entering the prism 86, each separated light is converted from spectral data linear with respect to wavelength to spectral data linear with respect to wavenumber (wavenumber spectrum). Then, each light converted into a wavenumber spectrum by the prism 86 is input to the light-receiving element 89 via the lens 88. The light receiving element 89 may be, for example, a line sensor (such as a CCD camera). The light receiving element 89 measures the interference of the measurement interference light and the calibration interference light for each wave number. An interference signal corresponding to the intensity of the measured interference light is input to a calculation device 200, which will be described later. The calculation device 200 performs processing such as Fourier transform on the interference signal, and obtains a tomographic image of the fundus along the scanning line.

[0047] Scanning in the fundus OCT optical system 92 will be described with reference to FIG. 4. FIG. 4 shows the optical path along which light emitted from the fiber collimator 114 is irradiated to the subject's eye E. Only some optical components arranged on the optical path (i.e., lenses 116, 118, and 120, the dichroic mirror 104, the scanner 106, and the objective lens 112) are illustrated, while other optical components are omitted. As shown in FIG. 4, in the fundus OCT optical system 92, two lenses 112 and 120 are arranged between the scanner 106 and the subject's eye E. The scanner 106 is also arranged at a position conjugate with the interior of the eyeball of the subject's eye E. As a result, the fundus OCT optical system 92 performs pivot scanning, with a pivot located within the eyeball of the subject's eye E. Normally, in the case of pivot scanning, the pivot is located at the pupil of the subject's eye E. Therefore, if there is an opacified area in the crystalline lens of the subject's eye E, it becomes difficult to sufficiently irradiate the fundus with light. In this embodiment, a pivot is provided between the fundus and the crystalline lens of the subject's eye E, so that light can be easily irradiated to the fundus regardless of the state of the crystalline lens of the subject's eye E.

[0048] In addition, the end face of the fiber collimator 114 is arranged at a position conjugate with the fundus of the subject's eye E. This allows the light emitted from the fiber collimator 114 to be focused on the fundus of the subject's eye E. The position of the lens 116 can be changed by driving a focus adjustment mechanism 140, which will be described later. This allows the light emitted from the fiber collimator 114 to be focused on the fundus of the subject's eye E in accordance with the refractive power of the subject's eye E. Therefore, the fundus OCT optical system 92 can perform tomography of the fundus of the subject's eye E in an appropriate manner.

[0049] Next, the refraction measurement optical system 94 will be described. The refraction measurement optical system 94 is an optical system used to measure the refractive power of the subject's eye E. In this embodiment, the fundus light source 62 used in the fundus OCT optical system 92 is used to measure the refractive power of the subject's eye E. The refraction measurement optical system 94 includes a fiber collimator 114, lenses 116, 118, 120, 130, and 134, dichroic mirrors 104 and 110, a holed mirror 105, a scanner 106, mirrors 122, 124, 126, and 128, an objective lens 112, an aperture 132, a ring lens 136, a sensor 138, a focus adjustment mechanism 140, and a fogging mechanism (not shown).

[0050] 1 is emitted from the fiber collimator 114 and passes through the same optical path as the above-described fundus OCT optical system 92 until it is irradiated onto the subject's eye E. That is, in the reflex measurement optical system 94, the light emitted from the fiber collimator 114 passes through lenses 116 and 118, the dichroic mirror 104, the apertured mirror 105, the scanner 106, the dichroic mirror 110, the lens 120, mirrors 122, 124, and 126, the dichroic mirror 110, and the objective lens 112, before it is irradiated onto the subject's eye E.

[0051] Light reflected from the subject's eye E is irradiated onto the perforated mirror 105 via the objective lens 112, the dichroic mirror 110, the mirrors 126, 124, and 122, the lens 120, the dichroic mirror 110, and the scanner 106. The light irradiated onto the perforated mirror 105 is reflected by a reflective surface disposed around the periphery of a circular hole in the perforated mirror 105 and then reflected by the mirror 128. The light reflected by the mirror 128 is detected by the sensor 138 via the lens 130, the aperture 132, the lens 134, and the ring lens 136. The ring lens 136 has a ring-shaped lens portion disposed on the lens 134 side and a light-shielding portion disposed on the sensor 138 side. The light-shielding portion is light-shielded except for the portion connected to the lens portion. The light incident on the ring lens 136 is emitted from the ring lens 136 as a ring-shaped light. The sensor 138 detects the ring-shaped light emitted from the ring lens 136. The sensor 138 is, for example, a CCD camera, and the image detected (captured) by the sensor 138 is input to a computing device.

[0052] In the reflex measurement optical system 94, the optical path from the fiber collimator 114 to the subject's eye E overlaps with that of the fundus OCT optical system 92. Therefore, scanning in the reflex measurement optical system 94 is a pivot scan, similar to the fundus OCT optical system 92. Therefore, similar to the fundus OCT optical system 92, light can be irradiated to the fundus regardless of the state of the crystalline lens of the subject's eye E. Furthermore, by scanning the fundus, a good image can be obtained even if there are factors that attenuate reflection from the fundus, such as fundus disease or blood vessels.

[0053] The refraction measurement optical system 94 also includes a focus adjustment mechanism 140. The focus adjustment mechanism 140 includes a drive device (not shown) that moves the fiber collimator 114, the lenses 116 and 134, the diaphragm 132, the ring lens 136, and the sensor 138 in the optical axis direction (Z-axis direction) as a whole. The focus adjustment mechanism 140 can move the positions of the fiber collimator 114 and the sensor 138 in the optical axis direction as a whole by driving the drive device. This allows the positions of the fiber collimator 114 and the sensor 138 to be moved to positions conjugate with the eye E to be inspected according to the refractive power of the eye E, enabling accurate refraction measurement.

[0054] Furthermore, in the ophthalmic apparatus 1 of this embodiment, the scanner 106 is disposed on an overlapping optical path where the optical path of the anterior segment OCT optical system 90 and the optical path of the fundus OCT optical system 92 overlap. That is, the scanner 106 performs scanning in both the anterior segment OCT optical system 90 and the fundus OCT optical system 92. This makes it possible to avoid a complex configuration within the ophthalmic apparatus 1 and to reduce the number of parts.

[0055] The configuration of the control system of the ophthalmic apparatus 1 of this embodiment will be described with reference to Fig. 5. As shown in Fig. 5, the ophthalmic apparatus 1 is controlled by a calculation device 200. The calculation device 200 is configured by a microcomputer (microprocessor) including a CPU, ROM, RAM, etc. The calculation device 200 is connected to the anterior eye light source 12, the light receiving element 54, and the signal processor 56 in the anterior eye OCT interferometer 10, the fundus light source 62 and the light receiving element 89 in the fundus OCT interferometer 11, the scanner 106, the sensor 138, the focus adjustment mechanism 140, and the touch panel 142.

[0056] The arithmetic device 200 controls the on / off of the anterior-segment light source 12 and drives the scanner 106 to scan the light irradiated onto the anterior segment of the subject's eye E. The arithmetic device 200 also receives an interference signal sampled by the signal processor 56. The arithmetic device 200 generates a tomographic image by Fourier transforming the interference signal, identifies the position of each part of the anterior segment of the subject's eye E (e.g., the cornea, the anterior chamber, the crystalline lens, etc.), and calculates the shape of each tissue in the anterior segment.

[0057] Similarly, the arithmetic unit 200 controls the on / off of the fundus light source 62 and drives the scanner 106 to scan the light irradiated onto the fundus of the subject's eye E. An interference signal corresponding to the intensity of the interference light detected by the light receiving element 89 is also input to the arithmetic unit 200. The arithmetic unit 200 generates a tomographic image by Fourier transforming the interference signal from the light receiving element 89, identifies the position of each part of the fundus of the subject's eye E (e.g., retina, choroid, etc.), and calculates the shape of each tissue of the fundus.

[0058] The arithmetic device 200 also receives an electrical signal (photographed image) detected by the sensor 138, and calculates the refractive power of the subject's eye E based on the input image. The data input to the arithmetic device 200 and the calculation results are stored in a memory (not shown). The arithmetic device 200 also calculates the axial length of the subject's eye E based on the calculated shape of the anterior segment and the shape of the fundus. The calculation of the axial length will be described later.

[0059] Furthermore, the arithmetic device 200 controls the touch panel 142. The touch panel 142 functions as a display unit that provides the examiner with various information related to the measurement results of the subject's eye E, and also functions as an input unit that accepts instructions from the examiner. For example, the touch panel 142 can display a tomographic image of the anterior segment of the subject's eye E generated by the arithmetic device 200, a tomographic image of the fundus, calculated refractive power, and other data acquired by scanning. Also, for example, the touch panel 142 can input various settings for the ophthalmologic apparatus 1. Note that although the ophthalmologic apparatus 1 of this embodiment includes the touch panel 142, it may be configured to include a monitor and an input device (e.g., a mouse, a keyboard, etc.) as long as it is capable of displaying and inputting the above information.

[0060] In this embodiment, the arithmetic device 200 simultaneously irradiates the anterior segment of the subject's eye E with light from the anterior-segment light source 12 and the fundus of the subject's eye E with light from the fundus light source 62. That is, the arithmetic device 200 acquires an interference signal obtained from the anterior segment and an interference signal obtained from the fundus by driving the scanner 106 in a state in which both the light from the anterior-segment light source 12 and the light from the fundus light source 62 are incident on the scanner 106. Therefore, in this embodiment, it is possible to measure the anterior segment and the fundus of the subject's eye E in approximately the same state. Hereinafter, with reference to FIGS. 6 to 15, a process for measuring the anterior segment, fundus, refractive power, and axial length of the subject's eye E will be described.

[0061] Fig. 6 is a flowchart showing an example of a process for performing various measurements on the subject's eye E using the ophthalmic apparatus 1. As shown in Fig. 6, first, when the examiner inputs an instruction to start the examination on the touch panel 142, the arithmetic device 200 aligns the subject's eye E with the ophthalmic apparatus 1 (S12). The alignment is performed using an alignment optical system (not shown) provided in the ophthalmic apparatus 1. Note that the alignment using the alignment optical system can employ a method used in known ophthalmic apparatuses, and therefore a detailed description thereof will be omitted.

[0062] Once the alignment of the subject's eye E and the ophthalmologic apparatus 1 is complete, the arithmetic device 200 performs a reflex measurement (S14). The reflex measurement is performed in the following procedure. First, the arithmetic device 200 adjusts the scanner 106. At this time, the arithmetic device 200 adjusts the diameter of the circle to be scanned and the irradiation position on the subject's eye E based on a preset value (initial setting value). The initial setting value can be set to a value smaller than the pupil diameter of the simulated eye, for example. Furthermore, if the subject has previously undergone an examination, the scanner 106 may be adjusted based on the measurement results from that examination.

[0063] After the adjustment of the scanner 106 is completed, the arithmetic device 200 turns on the fundus light source 62, captures an image detected by the sensor 138, and measures the refractive power by analyzing the image. At this time, a fogging mechanism (not shown) may be used to measure the refractive power in a state where the refractive accommodation power of the crystalline lens of the subject's eye E is eliminated. Note that the fogging mechanism may be one used in known ophthalmic devices, and therefore a detailed description thereof will be omitted.

[0064] When the refraction measurement is completed, the calculation device 200 adjusts the focus adjustment mechanism 140 based on the result of the refraction measurement in S14 (S16). For example, if the subject's eye E is a hyperopic eye or a myopic eye, the calculation device 200 drives the focus adjustment mechanism 140 to move the positions of the fundus light source 62 and the sensor 138 relative to the subject's eye E, and moves the sensor 138 to a position conjugate with the fundus of the subject's eye E.

[0065] Next, the calculation device 200 simultaneously performs anterior segment OCT measurement and fundus OCT measurement (S18). First, the anterior segment OCT measurement will be described with reference to FIG. 7. As shown in FIG. 7, the calculation device 200 adjusts the galvanometer mirror 108 to one scan angle within the scan angle range (S32). As a result, light from the anterior segment light source 12 is incident on the subject's eye E at an incident position and incident angle corresponding to the adjusted scan angle.

[0066] After completing the adjustment of the galvanometer mirror 108, the arithmetic unit 200 turns on the anterior-segment light source 12 and acquires the interference signal sampled by the signal processor 56 while varying the frequency of the light emitted from the anterior-segment light source 12 (S34). The interference signal sampled by the signal processor 56 is a signal whose signal intensity varies over time, as shown in FIG. 8 . This signal is an interference wave signal obtained by combining the reference light with the light reflected from each part of the subject's eye E (e.g., the anterior and posterior surfaces of the cornea, the anterior and posterior surfaces of the crystalline lens, etc.). The arithmetic unit 200 can then Fourier transform the signal input from the signal processor 56 to identify the depth position of each part of the subject's eye E from the signal. In the anterior-segment OCT measurement described above, the A-scan speed is set to, for example, approximately 100 kHz.

[0067] Next, the calculation device 200 determines whether the measurement in S34 has been performed for all scan angles (i.e., all incident positions and angles) previously set before the measurement (S36). If the measurement in S34 has not been performed for all scan angles (NO in S36), the calculation device 200 returns to S32 and repeats the processes in S32 to S36. This allows an interference signal to be obtained by A-scan for each scan angle at which the galvanometer mirror 108 is scanned. In this embodiment, the galvanometer mirror 108 is scanned at the A-scan speed (approximately 100 kHz) of the anterior segment OCT measurement. Note that in this specification, changing the scan angle of the galvanometer mirror 108 to change the incident position and incident angle of light from the light source is referred to as a "B-scan." In the anterior segment OCT measurement, the galvanometer mirror 108 is scanned so that the B-scan range (width of the B-scan) is, for example, approximately 16 mm.

[0068] When the measurement in S34 has been performed for all scan angles (YES in S36), the calculation device 200 identifies the position of each part of the subject's eye E (e.g., the anterior and posterior surfaces of the cornea, the anterior and posterior surfaces of the crystalline lens, etc.) from the interference signals obtained for each scan angle (S38). Specifically, when the process of S34 is performed for each scan angle, interference signal information (A-scan information) is acquired for each scan angle. Therefore, as shown in FIG. 9, two-dimensional information is obtained in which interference signal information (A-scan information) is arranged for the number of scan angles (n). Therefore, the calculation device 200 identifies the position of each part of the subject's eye E by calculating the boundary line of each part of the subject's eye E (e.g., the cornea, the anterior chamber, the iris, the crystalline lens, etc.) included in each interference signal information. Note that in anterior segment OCT measurement, A-scan information is acquired for approximately 800 scan angles, for example. Therefore, two-dimensional information in which approximately 800 pieces of interference signal information (A-scan information) are arranged is obtained.

[0069] In this embodiment, the anterior segment OCT measurement in S14 is performed using the radial scan method shown in Fig. 10. This allows tomographic images of the entire anterior segment to be acquired. That is, the B scan direction is set to the radial direction from the corneal apex of the subject's eye E, and the C scan direction is set to the circumferential direction, and the tomographic images are acquired. The arithmetic unit 200 stores the acquired (photographed) tomographic image data in memory.

[0070] As described above, the anterior segment OCT optical system 90 employs telecentric scanning. Therefore, it is possible to acquire a tomographic image without distortion in the anterior segment OCT measurement. Furthermore, in the anterior segment OCT optical system 90, the end face of the fiber collimator 102 is disposed at a position conjugate with the anterior segment of the subject's eye E, and the anterior segment light source 12 outputs light of a wavelength suitable for capturing a tomographic image of the anterior segment of the subject's eye E. Therefore, it is possible to suitably calculate the shape of the anterior segment of the subject's eye E in the anterior segment OCT measurement.

[0071] Next, fundus OCT measurement will be described. As described above, fundus OCT measurement is performed simultaneously with anterior segment OCT measurement (S18). Therefore, like anterior segment OCT measurement, fundus OCT measurement is also performed according to the flowchart shown in FIG. 7. Specifically, the calculation device 200 turns on the fundus light source 62 together with the anterior segment light source 12, causing light from the fundus light source 62 to be incident on the subject's eye E at an incident position and incident angle corresponding to the scanning angle adjusted in S32 described above, and capturing an interference signal of light separated into wavenumber spectra (S34). The calculation device 200 outputs pulsed light from the fundus light source 62 and further receives a signal from the light receiving element 89 in synchronization with the light output. That is, the calculation device 200 acquires the interference signal by repeatedly turning on and off the fundus light source 62 at a predetermined cycle. Pulse light irradiation in fundus OCT measurement will be described later. In fundus OCT measurement, the A-scan speed is set to, for example, approximately 10 kHz.

[0072] In the fundus OCT optical system 92, the fundus light source 62 emits light of a wavelength that reaches the fundus of the subject's eye E, and the optical components are arranged so that the light emitted from the fundus light source 62 is focused on the fundus of the subject's eye E. Therefore, as shown in FIG. 11 , the arithmetic unit 200 can identify the position of the fundus of the subject's eye E, such as the retina. This measurement is repeated until all preset scan angles are performed, as in S36 described above. In fundus OCT measurement, the galvanometer mirror 108 is scanned so that the B-scan range is, for example, 3.8 mm. In fundus OCT measurement, A-scan information is acquired for approximately 80 scan angles, for example. In this embodiment, the light output from the fundus light source 62 and the light output from the anterior-segment light source 12 are scanned simultaneously by the scanner 106. Therefore, in the alignment of S12 described above, the positions of the scanner 106, objective lens 112, lens 120, etc. are adjusted based on the focal length of the objective lens 112, etc., so that the B-scan range in the anterior segment OCT measurement and the B-scan range in the fundus OCT measurement are within the above-mentioned ranges (i.e., approximately 16 mm and approximately 3.8 mm, respectively).

[0073] In this embodiment, the A-scan speed (approximately 10 kHz) in fundus OCT measurement is slower than the A-scan speed (approximately 100 kHz) in anterior segment OCT measurement. Therefore, when the scanner 106 simultaneously scans with the respective light beams in each measurement, the number of A-scan information (number of A-scans) obtained by fundus OCT measurement is smaller than the number of A-scan information obtained by anterior segment OCT measurement. Generally, in OCT measurement using the Fourier domain method, if the B-scan (change in the incident position and incident angle of the measurement light due to a change in the mirror scanning angle (i.e., scanning of the measurement light in a direction perpendicular to the depth direction of the subject's eye)) performed together with the A-scan is performed at a relatively high speed, a change in the scanning angle during A-scan measurement at a certain scanning angle can cause a deviation from the original measurement position, resulting in attenuation or loss of the interference signal obtained from the interference light. That is, in this embodiment, if the scan resolution (B-scan range per A-scan (i.e., the interval between adjacent A-scans)) in the OCT measurement is larger than the optical resolution in the OCT measurement, a phenomenon in which the interference signal obtained from the interference light is attenuated (so-called fringe washout) may occur. Therefore, it is desirable to set the scan resolution smaller than the optical resolution. The derivation of the scan resolution and the optical resolution will be explained below.

[0074] 12 is a diagram for explaining the derivation of the scan resolution in the anterior segment OCT measurement and the fundus OCT measurement. As shown in FIG. 12, the B scan range of the anterior segment OCT measurement is denoted by Wa, and the focal length of the objective lens is denoted by f obj If the galvanometer mirror's angle of oscillation is θ, the total angle of the beam on the galvanometer mirror surface is 2θ, so Wa=f obj The relationship 2tanθ holds. Also, if the galvanometer mirror is a real image rather than a physical object, then if the magnification of the galvanometer mirror is Ma, then Wa=f obj The relationship of 2tanθ / Ma holds. On the other hand, the B-scan range of fundus OCT measurement is Wp, and the focal length of the objective lens is f obj , the focal length of the lens is f relay, the focal length of the test eye is f eye , if the swing angle of the galvanometer mirror is θ, then Wp=(f eye f relay / f obj ) 2tanθ. If the magnification of the real image R on the galvanometer mirror is Mp, then Wp=f eye Therefore, the scan resolution ΔWp of fundus OCT measurement is given by ΔWp=(f eye It can be expressed as 2tanθ / Mp) / Np.

[0075] In this embodiment, the scanner 106 is shared between the anterior segment OCT measurement and the fundus OCT measurement. Therefore, the swing angle θ of the galvanometer mirror 108 is the same for the anterior segment OCT measurement and the fundus OCT measurement. Therefore, the scan resolution ΔWp in the fundus OCT measurement is given by ΔWp=(f eye / f obj )·(Ma / Mp)·(Wa / Np). Furthermore, if the time required for one B-scan is T, the A-scan speed in the anterior segment OCT measurement is Fa, the number of A-scans is Na, and the A-scan speed in the fundus OCT measurement is Fp, the number of A-scans is Np, then the relationship T=Na / Fa=Np / Fp holds. Therefore, in this embodiment, the scan resolution ΔWp of the fundus OCT measurement is given by ΔWp=(f eye / f obj )·(Ma / Mp)·(Wa / Na)·(Fa / Fp).

[0076] Fig. 13 is a diagram for explaining the derivation of the optical resolution in fundus OCT measurement. When the light output from the light source is in a single mode, as shown in Fig. 13, the optical resolution 2d is expressed as 2d = (4λ / π) (F / D), where λ is the central wavelength of the light, F is the focal length of the lens Le, and D is the beam diameter of the light incident on the lens Le. Therefore, in this embodiment, the optical resolution 2dp in fundus OCT measurement is calculated by dividing the central wavelength of the light output from the fundus light source 62 by λp and the focal length of the eyeball of the subject's eye E by f eyeIf the beam diameter of the light incident on the cornea is Dp, then 2dp=(4λp / π)·(f eye / Dp).

[0077] As mentioned above, to prevent fringe washout, the scanning resolution must be smaller than the optical resolution. In other words, the relationship δWp<2dp must be established. Therefore, from the scanning resolution and optical resolution of fundus OCT measurement derived above, (Fp·λp·Mp) / Dp>(π / 4)·(Ma / f obj In this embodiment, as described above, the conditions for the anterior segment OCT measurement are Fa=100 kHz, Wa=16 mm, and Na=800 lines. Therefore, if other conditions are set to, for example, Ma=1, f obj = 70 mm, λp = 0.83 μm, Dp = 1 mm, and Mp = 1, the occurrence of fringe washout can be prevented by setting the A-scan speed Fp in fundus OCT measurement to Fp > 27036 Hz.

[0078] In this embodiment, the A-scan rate for fundus OCT measurement is set to approximately 10 kHz. Therefore, if light is continuously output from the fundus light source 62, fringe washout occurs in the interference signal obtained in fundus OCT measurement. In contrast, in this embodiment, as described above, the A-scan rate for fundus OCT measurement is set to approximately 10 kHz, and light from the fundus light source 62 is output in pulses at a period that matches the A-scan rate for fundus OCT measurement. In this embodiment, Fp / {π / (4λp)·(Dp / f obj The occurrence of fringe washout can be suitably prevented by outputting light from the fundus light source 62 in pulses with a period of 1 / Fp and a duty ratio of less than {(Ma / Mp)·(Wa / Na)·Fa}. That is, in this embodiment, the occurrence of fringe washout can be prevented by irradiating light output from the fundus light source 62 in pulses with a duty ratio of less than 10000 / 27036 and a period of 100 μs.

[0079] When measurements for all scan angles are completed, similar to S38 described above, the calculation device 200 identifies the position of each part of the subject's eye E (e.g., retina, choroid, etc.) from the interference signals obtained for each scan angle. The fundus OCT optical system 92, unlike the anterior segment OCT optical system 90, uses pivot scanning. Therefore, as shown in FIG. 14, two-dimensional information connecting pivots within the eyeball of the subject's eye E is obtained at each scan angle. Therefore, the calculation device 200 identifies the position of each part of the subject's eye E by calculating the average value of the position information for each part of the subject's eye E contained in each piece of interference signal information. The shape of the fundus of the subject's eye E can be calculated from these measurement results.

[0080] After the calculation of the shape of the anterior segment and the fundus is completed, the arithmetic device 200 corrects the distance between the anterior segment and the fundus of the subject's eye E based on the interference signal obtained from the calibration mirror 36 (S20). Correction of the distance between the anterior segment and the fundus will be described with reference to FIG. 15. The process of FIG. 15 is executed in advance before the examination of the subject's eye E. As shown in FIG. 15, the arithmetic device 200 first performs anterior segment OCT measurement and fundus OCT measurement on an object (e.g., a simulated eye) with a known optical path length (S52). When the anterior segment OCT measurement is performed on the simulated eye, peaks indicating the depth positions of each part of the anterior segment of the simulated eye obtained from the measurement interference light and peaks indicating the depth position of the calibration mirror 36 obtained from the calibration interference light can be obtained, as shown in FIG. 16. The arithmetic unit 200 superimposes a signal indicating the depth position of the calibration mirror 36 on the tomographic image of the anterior segment by acquiring peaks indicating the depth position of the calibration mirror 36 for each scanning angle, and calculates the difference D1 between the optical path length from the anterior segment light source 12 to a predetermined site in the anterior segment (for example, the anterior surface of the cornea) and the optical path length from the anterior segment light source 12 to the calibration mirror 36 (S54). In addition, the arithmetic unit 200 calculates the depth position of the interference signal of the calibration mirror 36 in the anterior segment OCT measurement as a reference position Z Ai and stores it in memory (S56).

[0081] Similarly, the calculation device 200 performs fundus OCT measurement on the simulated eye, and acquires peaks indicating the depth positions of each part of the fundus of the simulated eye obtained from the measurement interference light and peaks indicating the depth position of the calibration mirror 36 obtained from the calibration interference light, as shown in Fig. 17. The calculation device 200 acquires peaks indicating the depth position of the calibration mirror 36 for each scanning angle, thereby superimposing a signal indicating the depth position of the calibration mirror 36 on the tomographic image of the fundus, and calculates a difference D2 between the optical path length from the fundus light source 62 to a predetermined part of the fundus (e.g., the retina) and the optical path length from the fundus light source 62 to the calibration mirror 36 (S58). Furthermore, the calculation device 200 calculates the depth position of the interference signal of the calibration mirror 36 in the fundus OCT measurement as a reference position Z Pi and stores it in memory (S60).

[0082] Then, the calculation device 200 calculates the reference position Z Ai and reference position Z Pi The distance between them (optical path difference) L i The depth position of each part of the simulated eye is known. Therefore, for example, the reference position Z is calculated based on the distance from the anterior surface of the cornea of ​​the simulated eye to the retina and the calculated optical path length differences D1 and D2. Ai and reference position Z Pi Optical path difference L i The calculation device 200 can calculate the calculated optical path length difference L i is stored in memory (S64).

[0083] In the OCT measurement using the ophthalmologic apparatus 1, the optical path length of each interferometer 10, 11 may vary. For example, if the optical path length of the anterior segment OCT interferometer 10 when measuring the subject's eye E varies from the optical path length of the anterior segment OCT interferometer 10 when measuring the simulated eye, the depth position of the interference signal of the calibration mirror 36 when measuring the subject's eye E may differ from the reference position Z of the interference signal of the calibration mirror 36 when measuring the simulated eye. Ai In this embodiment, even if the optical path length of the interferometer fluctuates in this way, the reference position Z Ai , Z PiBased on this, the optical path length difference between the interferometers 10 and 11 during OCT measurement can be calculated. Specifically, the depth positions of the calibration mirror 36 when the anterior segment OCT measurement and the fundus OCT measurement of the eye E to be examined in S18 of FIG. 6 are respectively Z A , Z P Then, the depth position Z A From depth position Z P The distance (optical path difference) L is L = L i +(Z A -Z Ai )+(Z P -Z Pi ) can be calculated as follows. In this way, by calculating the optical path length difference L during actual measurement, the distance from the anterior segment to the fundus is corrected (S20).

[0084] When the correction of the distance between the anterior segment and the fundus is completed, the calculation device 200 calculates the axial length of the eye E (S22). A , Z P The axial length of the subject's eye E is calculated based on the optical path length difference L, the depth positions of each part of the anterior segment calculated in the anterior segment OCT measurement, and the depth positions of each part of the fundus calculated in the fundus OCT measurement.

[0085] When all measurements (refractive index measurement, anterior segment OCT measurement, and fundus OCT measurement) are completed, the arithmetic device 200 outputs the analysis results to the touch panel 142 (S24). The ophthalmic apparatus 1 of this embodiment can perform various measurements, including anterior segment OCT measurement, fundus OCT measurement, and refractive index measurement, thereby enabling comprehensive analysis of the condition of the subject's eye E. For example, by measuring the subject's eye before cataract surgery, the analysis results can include calculation of intraocular lens (IOL) power, corneal aberration, and crystalline lens opacity. Furthermore, by measuring the subject's eye before cataract surgery, the error in the postoperative refractive power of the subject's eye E predicted before surgery can be evaluated, which can be used to improve the accuracy of IOL power calculation. Furthermore, by measuring the subject's eye with glaucoma, the progression of glaucoma can be predicted from the retinal thickness distribution at the fundus, and angle-closure disease can be screened for. Furthermore, by measuring the subject's eye that is severely myopic, the condition of the subject's eye can be examined in detail and comprehensively.

[0086] As described above, the ophthalmic apparatus 1 of this embodiment has a different central wavelength for the light irradiated onto the anterior segment and the light irradiated onto the fundus. Therefore, it is possible to irradiate the anterior segment and the fundus with light of wavelengths suitable for each examination. Therefore, the ophthalmic apparatus 1 of this embodiment can accurately measure the anterior segment and the fundus. Furthermore, in the ophthalmic apparatus 1 of this embodiment, the optical path of the light output from the anterior segment light source 12 and the optical path of the light output from the fundus light source 62 partially overlap. Therefore, both types of light can be simultaneously irradiated onto the subject's eye E. Therefore, the ophthalmic apparatus 1 of this embodiment can efficiently perform examinations of different areas of the subject's eye E (i.e., the anterior segment and the fundus). Therefore, multiple types of measurements can be performed on the subject's eye E in approximately the same condition. In this way, the ophthalmic apparatus 1 reduces the burden on the subject and accurately performs multiple types of measurements on the subject's eye E.

[0087] Furthermore, the ophthalmologic apparatus 1 of this embodiment shares the scanner 106 for all of the multiple types of measurements (anterior segment OCT measurement, fundus OCT measurement, and reflex measurement), which makes it possible to avoid a complicated configuration of the optical system within the ophthalmologic apparatus 1 and reduce the number of parts.

[0088] In addition, in the ophthalmologic apparatus 1 of this embodiment, the light source used for the reflex measurement is also used as the light source (fundus light source 62) for the fundus OCT measurement, which prevents the internal configuration of the ophthalmologic apparatus 1 from becoming complicated and reduces the number of parts.

[0089] Furthermore, in the ophthalmologic apparatus 1 of this embodiment, the scanner 106 is shared between anterior segment OCT measurement and fundus OCT measurement, and the A-scan speed (approximately 10 kHz) in fundus OCT measurement is set slower than the A-scan speed (approximately 100 kHz) in anterior segment OCT measurement. Therefore, if the fundus light source 62 were to continuously irradiate light, fringe washout would likely occur in the fundus OCT measurement if the scanner 106 were to scan at the A-scan speed for the anterior segment OCT measurement. However, in this embodiment, pulsed light is output from the fundus light source 62 at a predetermined cycle. This makes the scan resolution for fundus OCT measurement smaller than the optical resolution, thereby effectively preventing fringe washout in the fundus OCT measurement.

[0090] Furthermore, in the ophthalmologic apparatus 1 of this embodiment, the distance between the anterior segment and the fundus is corrected using a calibration mirror 36. In this embodiment, when examining the anterior segment and the fundus, two interferometers, an anterior segment OCT interferometer 10 and a fundus OCT interferometer 11, are used. Therefore, compared to when only a single interferometer is used, the difference in optical path length between the two interferometers 10 and 11 is relatively likely to fluctuate, and deviations are likely to occur in the calculation of the distance between the anterior segment and the fundus. However, in this embodiment, the calibration mirror 36 is used to correct the distance between the anterior segment and the fundus at a predetermined reference position Z Ai , Z Pi Based on this, it is possible to suitably correct the distance between the anterior segment and the fundus during measurement of the subject's eye E. Therefore, in this embodiment, it is possible to calculate an accurate axial length.

[0091] (Correspondence) The anterior segment light source 12 and the fundus light source 62 are examples of a "first light source" and a "second light source," respectively. The anterior segment OCT interferometer 10 and the fundus OCT interferometer 11 are examples of a "first interferometer" and a "second interferometer," respectively. The measurement interference light in the anterior segment OCT measurement and the measurement interference light in the fundus OCT measurement are examples of a "first interference light" and a "second interference light," respectively. The range including the anterior segment and the range including the fundus are examples of a "first range" and a "second range," respectively. The refractive power measurement optical system 94 is an example of a "refractive power measurement optical system." The calibration mirror 36 is an example of a "first calibration mirror" and a "second calibration mirror." [Example]

[0092] Next, an ophthalmic apparatus 2 according to Example 2 will be described. Note that, among the ophthalmic apparatus 2 of Example 2, components common to the ophthalmic apparatus 1 of Example 1 are given common reference numerals, and descriptions of their operations, processes, etc. will be omitted.

[0093] 18, the ophthalmologic apparatus 2 includes an anterior segment OCT interferometer 210 that performs tomographic imaging of the anterior segment of the subject's eye E, and a fundus OCT interferometer 211 that performs tomographic imaging of the fundus of the subject's eye E. The anterior segment OCT interferometer 210 uses a polarization-sensitive OCT (PS-OCT), which is an optical frequency sweep OCT that uses a wavelength sweep light source and is capable of capturing the polarization characteristics of the subject's eye E. The configuration, control, etc. of the fundus OCT interferometer 211 are similar to those of the fundus OCT interferometer 11 of the first embodiment.

[0094] As shown in FIG. 18, the anterior segment OCT interferometer 210 includes an anterior segment light source 12, a measurement optical system 220, a calibration optical system 230, a reference optical system 240, and an interference optical system 250.

[0095] The anterior-segment light source 12 is connected to a polarization control device 14 and a fiber coupler 215, and the fiber coupler 215 is connected to a PMFC (polarization-maintaining fiber coupler) 216 and a sampling trigger / clock generator 300. Therefore, the light output from the anterior-segment light source 12 is input to the PMFC 216 and the sampling trigger / clock generator 300 via the polarization control device 14 and the fiber coupler 215. The sampling trigger / clock generator 300 uses the light from the anterior-segment light source 12 to generate sampling triggers and sampling clocks for signal processors 256 and 257, which will be described later.

[0096] The measurement optical system 220 includes a PMFC 221 connected to the PMFC 216, two measurement optical paths S1 and S2 branching from the PMFC 221, a polarization beam combiner / splitter 225 connecting the two measurement optical paths S1 and S2, an SMFC (single-mode fiber coupler) 226 connected to the polarization beam combiner / splitter 225, two measurement optical paths S3 and S4 branching from the SMFC 226, and a probe optical system 26 connected to the measurement optical path S3. A calibration optical system 230 is connected to the measurement optical path S4 via the polarization control device 15. An optical path length difference generating unit 222 and a circulator 223 are disposed in the measurement optical path S1. A MEMS attenuator 224 and a circulator 227 are disposed in the measurement optical path S2. Therefore, the optical path length difference Δl between the measurement optical paths S1 and S2 is generated by the optical path length difference generating unit 222. The optical path length difference Δl may be set to be longer than the measurement range in the depth direction of the eye E to be examined. This makes it possible to prevent overlapping of interference light beams with different optical path length differences. The optical path length difference generation unit 222 may use, for example, an optical fiber or an optical system such as a mirror or a prism. In this embodiment, a 1-m PM fiber is used for the optical path length difference generation unit 222. The measurement optical system 220 further includes PMFCs 228 and 229. The PMFC 228 is connected to a circulator 223. The PMFC 229 is connected to a circulator 227.

[0097] One of the beams split by the PMFC 216 (i.e., the measurement beam) is input to the measurement optical system 220. The PMFC 221 splits the measurement beam input from the PMFC 216 into a first measurement beam and a second measurement beam. The first measurement beam split by the PMFC 221 is guided to a measurement optical path S1, and the second measurement beam is guided to a measurement optical path S2. The first measurement beam guided to the measurement optical path S1 passes through an optical path length difference generating unit 222 and a circulator 223 and is input to a polarization beam combiner / splitter 225. The second measurement beam guided to the measurement optical path S2 passes through a MEMS attenuator 224 and a circulator 227 and is input to the polarization beam combiner / splitter 225. The PM fiber 304 is connected to the polarization beam combiner / splitter 225 in a state rotated 90 degrees in the circumferential direction with respect to the PM fiber 302. As a result, the second measurement light input to the polarization beam combiner / splitter 225 has a polarization component orthogonal to that of the first measurement light. Because the optical path length difference generating unit 222 is provided in the measurement light path S1, the first measurement light is delayed relative to the second measurement light by the distance of the optical path length difference generating unit 222 (i.e., an optical path length difference Δl is generated). The polarization beam combiner / splitter 225 superimposes the input first measurement light and second measurement light.

[0098] The light output from the polarizing beam combiner / splitter 225 (light obtained by superimposing the first and second measurement beams) is input to the SMFC 226, which splits the input light into third and fourth measurement beams. The third measurement beam split by the SMFC 226 is guided to the measurement optical path S3, and the fourth measurement beam is guided to the measurement optical path S4. As in the first embodiment, the third measurement beam guided to the measurement optical path S3 is input to the anterior segment OCT optical system 90 of the probe optical system 26 and irradiated onto the subject's eye E. Here, the SM (single mode) fiber used in the measurement optical path S3 generates birefringence due to stress or bending, so the polarization state of the light irradiated onto the subject's eye E may change depending on the state of the SM fiber. The reflected light from the subject's eye E passes through the anterior segment OCT optical system 90 in the opposite direction to the incident path, is input to the SMFC 226, and is then input to the polarizing beam combiner / splitter 225.

[0099] On the other hand, the fourth measurement light guided to the measurement optical path S4 is input to the calibration optical system 230 and irradiated onto the glass block 33 via the lens 32. The glass block 33 is arranged so that a portion of the fourth measurement light irradiated via the lens 32 passes through. For example, as shown in FIG. 18, the glass block 33 is arranged so that approximately half of the lower portion of the fourth measurement light irradiated via the lens 32 passes through while approximately half of the upper portion of the fourth measurement light irradiated via the lens 32 does not pass through. As a result, a portion of the fourth measurement light irradiated via the lens 32 (light passing downward in FIG. 18) passes through the glass block 33 and is irradiated onto the quarter-wave plate 37. On the other hand, the other portion of the fourth measurement light irradiated via the lens 32 (light passing upward in FIG. 18) is irradiated onto the quarter-wave plate 37 without passing through the glass block 33. A phase difference of 90 degrees occurs in the light irradiated onto the quarter-wave plate 37 when it passes through the quarter-wave plate 37. The light that passed through the quarter-wave plate 37 is irradiated onto the calibration mirror 36 via the dichroic mirror 34 and the lens 35. The light reflected from the calibration mirror 36 passes through the lens 35 and is subjected to a further 90-degree phase difference at the quarter-wave plate 37. A portion of the light passes through the glass block 33 and is irradiated onto the lens 32, while another portion is irradiated onto the lens 32 without passing through the glass block 33. The light then passes through the lens 32 and is input to the SMFC 226. Therefore, the SMFC 226 receives light that has passed through the glass block 33 and light that has not passed through the glass block 33. The light that passed through the glass block 33 may be light that passed through the glass block 33 once (i.e., light that passed through the glass block 33 on only one of the outward and return paths while traveling back and forth between the lenses 32 and 35), light that passed through the glass block 33 twice (i.e., light that passed through the glass block 33 on both the outward and return paths while traveling back and forth between the lenses 32 and 35), or both. The calibration mirror 36 and the glass block 33 are positioned so that the measurement light passing through the measurement optical path S4 is measured near the Nyquist frequency of the measurement range. The light input to the SMFC 226 is input to the polarization beam combiner / splitter 225.The polarization state of the light passing through the SM fiber used in the measurement optical path S4 may change due to bending or temperature changes. In this embodiment, the polarization state of the light passing through the quarter-wave plate 37 is controlled by appropriately rotating the quarter-wave plate 37. This allows the polarization beam combiner / splitter 225 to split the light into polarization components with uniform intensities. In other words, the intensities of the interference signals input to the balanced photodetectors 286, 287, 296, and 297 described below can be made uniform.

[0100] The reflected light from the subject's eye E and the reflected light from the calibration mirror 36, which are input to the polarization beam combiner / splitter 225, are split into two mutually orthogonal polarization components by the polarization beam combiner / splitter 225. For convenience, these are referred to as horizontally polarized reflected light (horizontally polarized component) and vertically polarized reflected light (vertically polarized component). The horizontally polarized reflected light is guided to the measurement optical path S1, and the vertically polarized reflected light is guided to the measurement optical path S2. The optical path of the horizontally polarized reflected light is changed by the circulator 223, and the horizontally polarized reflected light is input to the PMFC 228. The PMFC 228 splits the input horizontally polarized reflected light and inputs it to PMFCs 251 and 261 of the interference optical system 250, which will be described later. Therefore, the horizontally polarized reflected light input to the PMFCs 251 and 261 includes a reflected light component due to the first measurement light and a reflected light component due to the second measurement light. The optical path of the vertically polarized reflected light is changed by the circulator 227 and input to the PMFC 229. The PMFC 229 splits the input vertically polarized reflected light and inputs it to the PMFCs 252 and 262. Therefore, the vertically polarized reflected light input to the PMFCs 252 and 262 includes a reflected light component due to the first measurement light and a reflected light component due to the second measurement light.

[0101] The reference optical system 240 includes a polarization beam combiner / splitter 241 connected to the PMFC 216, reference delay lines (242, 243, 244, 245, 246) connected to the polarization beam combiner / splitter 241, a PMFC 247 connected to the polarization beam combiner / splitter 241, two reference optical paths R1 and R2 branching from the PMFC 247, a PMFC 248 connected to the reference optical path R1, and a PMFC 249 connected to the reference optical path R2. The PM fiber 310 is connected to the polarization beam combiner / splitter 241 in a state rotated 90 degrees in the circumferential direction with respect to the PM fiber 308. An optical path length difference generating unit 232 is disposed in the reference optical path R1. No optical path length difference generating unit is provided in the reference optical path R2. Therefore, the optical path length difference Δl′ between the reference optical path R1 and the reference optical path R2 is generated by the optical path length difference generating unit 232. For example, an optical fiber is used for the optical path length difference generation unit 232. The optical path length Δl' of the optical path length difference generation unit 232 may be the same as the optical path length Δl of the optical path length difference generation unit 222. By making the optical path length differences Δl and Δl' the same, the depth positions of the multiple interference beams, which will be described later, relative to the test object become the same. In other words, there is no need to align the multiple tomographic images to be acquired.

[0102] The other light (i.e., reference light) split by the PMFC 216 is input to the reference optical system 240. The reference light input from the PMFC 216 passes through a polarization beam combiner / splitter 241 and is input to reference delay lines (242, 243, 244, 245, 246). The reference delay lines (242, 243, 244, 245, 246) are configured by a lens 242, an attenuator 243, a Faraday rotator 244, a lens 245, and a reference mirror 246. The reference light that has passed through the polarization beam combiner / splitter 241 passes through, for example, an SM fiber 306, is output from a fiber collimator (not shown), and is input to the Faraday rotator 244 via the lens 242 and the attenuator 243. When the reference light passes through the Faraday rotator 244, its polarization direction is rotated, for example, by 45 degrees, and the reference light is incident on a reference mirror 246 via a lens 245. The polarization direction of the reference light reflected by the reference mirror 246 is again rotated by 45 degrees, incident on a fiber collimator, and input to the polarization beam combiner / splitter 241. That is, the light that passes through the reference delay lines (242, 243, 244, 245, 246) is input to the polarization beam combiner / splitter 241 with its polarization direction rotated by 90 degrees. Here, the reference mirror 246 is movable in a direction toward or away from the lens 242. In this embodiment, before starting measurement, the position of the reference mirror 246 is adjusted so that the signal from the subject's eye E falls within the measurement range of the OCT in the depth direction.

[0103] The reference light reflected by reference mirror 246 and input to polarization beam combiner / splitter 241 has its polarization direction rotated by 90 degrees, and is therefore input to PMFC 247 via PM fiber 310. PMFC 247 splits the input reference light into a first reference light and a second reference light. The first reference light is guided to reference optical path R1, and the second reference light is guided to reference optical path R2. The first reference light passes through optical path length difference generating unit 232 and is input to PMFC 248. The reference light input to PMFC 248 is split into a first split reference light and a second split reference light. The first split reference light passes through delay line 264 and is input to PMFC 251. The second split reference light passes through delay line 265 and is input to PMFC 252. The second reference light is input to PMFC 249 and split into a third split reference light and a fourth split reference light. The third branched reference light passes through a delay line 266 and is input to the PMFC 261. The fourth branched reference light passes through a delay line 267 and is input to the PMFC 262.

[0104] The interference optical system 250 includes a first interference light generation section 280 , a second interference light generation section 290 , a first interference light detection section 285 , and a second interference light detection section 295 .

[0105] The first interference light generation unit 280 includes PMFCs 251 and 252. As described above, the horizontally polarized reflected light is input to the PMFC 251 from the measurement optical system 220, and the first branched reference light (light having an optical path difference Δl) is input to the PMFC 251 from the reference optical system 240. The horizontally polarized reflected light includes a reflected light component due to the first measurement light (light having an optical path difference Δl) and a reflected light component due to the second measurement light (light without an optical path difference Δl). Therefore, the PMFC 251 combines the reflected light component due to the first measurement light (light having an optical path difference Δl) of the horizontally polarized reflected light with the first branched reference light to generate the first interference light (horizontally polarized component).

[0106] Furthermore, PMFC 252 receives vertically polarized reflected light from measurement optical system 220 and second branched reference light (light having optical path difference Δl) from reference optical system 240. The vertically polarized reflected light includes a reflected light component due to the first measurement light (light having optical path difference Δl) and a reflected light component due to the second measurement light (light without optical path difference Δl). Therefore, in PMFC 252, the reflected light component due to the first measurement light (light having optical path difference Δl) of the vertically polarized reflected light is combined with the second branched reference light to generate second interference light (vertically polarized component).

[0107] The second interference light generation unit 290 has PMFCs 261 and 262. As described above, the horizontally polarized reflected light is input to PMFC 261 from the measurement optical system 220, and the third branched reference light (light without optical path length difference Δl) is input to PMFC 261 from the reference optical system 240. Therefore, in PMFC 261, the reflected light component of the horizontally polarized reflected light that is due to the second measurement light (light without optical path length difference Δl) is combined with the third branched reference light to generate the third interference light (horizontally polarized component).

[0108] Furthermore, PMFC 262 receives vertically polarized reflected light from measurement optical system 220 and fourth branched reference light (light without optical path difference Δl) from reference optical system 240. Therefore, in PMFC 262, a reflected light component of the vertically polarized reflected light due to the second measurement light (light without optical path difference Δl) is combined with the fourth branched reference light to generate fourth interference light (vertically polarized component). The first interference light and second interference light correspond to the measurement light that has passed through measurement optical path S1, and the third interference light and fourth interference light correspond to the measurement light that has passed through measurement optical path S2.

[0109] The first interference light detection unit 285 detects the interference light (first interference light and second interference light) generated by the first interference light generation unit 280. The second interference light detection unit 295 detects the interference light (third interference light and fourth interference light) generated by the second interference light generation unit 290.

[0110] The first interference light detection unit 285 includes balanced photodetectors (hereinafter simply referred to as detectors) 286 and 287, and a signal processor 288 connected to the detectors 286 and 287. The PMFC 251 is connected to an input terminal of the detector 286, and the signal processor 288 is connected to an output terminal of the detector 286. The PMFC 251 splits the first interference light into two interference lights with a phase difference of 180 degrees, and inputs the two interference lights to the input terminals of the detector 286. The detector 286 performs differential amplification and noise reduction processing on the two interference lights with a phase difference of 180 degrees input from the PMFC 251, converts them into an electrical signal (first interference signal), and outputs the first interference signal to the signal processor 288. That is, the first interference signal is an interference signal HH between the horizontally polarized reflected light from the subject's eye E and the calibration mirror 36 due to the horizontally polarized measurement light and the reference light. Similarly, the PMFC 252 is connected to the input terminal of the detector 287, and the signal processor 288 is connected to the output terminal of the detector 287. The PMFC 252 splits the second interference light into two interference light beams with a phase difference of 180 degrees, and inputs the two interference light beams to the input terminals of the detector 287. The detector 287 performs differential amplification and noise reduction processing on the two interference light beams with a phase difference of 180 degrees, converts them into an electrical signal (second interference signal), and outputs the second interference signal to the signal processor 288. That is, the second interference signal is an interference signal HV of the vertically polarized reflected light from the eye E and the calibration mirror 36 due to the horizontally polarized measurement light, and the reference light.

[0111] The signal processor 288 samples the first interference signal and the second interference signal based on the sampling trigger and sampling clock input from the sampling trigger / clock generator 300. The first interference signal and the second interference signal sampled by the signal processor 288 are input to a calculation device 400, which will be described later. A known data acquisition device (so-called DAQ) can be used as the signal processor 288.

[0112] Similar to the first interference light detection unit 285, the second interference light detection unit 295 includes balanced photodetectors (hereinafter also simply referred to as detectors) 296 and 297 and a signal processor 298 connected to the detectors 296 and 297. The PMFC 261 is connected to the input terminal of the detector 296, and the signal processor 298 is connected to the output terminal of the detector 296. The PMFC 262 splits the third interference light into two interference lights with a phase difference of 180 degrees and inputs them to the input terminals of the detector 296. The detector 296 performs differential amplification and noise reduction processing on the two interference lights with a phase difference of 180 degrees, converts them into an electrical signal (third interference signal), and outputs the third interference signal to the signal processor 298. That is, the third interference signal is an interference signal VH of the horizontally polarized reflected light from the subject's eye E and the calibration mirror 36 due to the vertically polarized measurement light and the reference light. Similarly, a PMFC 262 is connected to the input terminal of the detector 297, and a signal processor 298 is connected to the output terminal of the detector 297. The PMFC 262 splits the fourth interference light into two interference lights whose phases differ by 180 degrees, and inputs the two interference lights to the input terminals of the detector 297. The detector 297 performs differential amplification and noise reduction processing on the two interference lights whose phases differ by 180 degrees, converts them into an electrical signal (fourth interference signal), and outputs the fourth interference signal to the signal processor 298. That is, the fourth interference signal is an interference signal VV of the vertically polarized reflected light from the eye E and the calibration mirror 36 and the reference light, caused by the vertically polarized measurement light.

[0113] The signal processor 298 samples the third interference signal and the fourth interference signal based on the sampling trigger and sampling clock input from the sampling trigger / clock generator 300. The third interference signal and the fourth interference signal sampled by the signal processor 298 are input to a calculation device 400, which will be described later. A known data acquisition device (so-called DAQ) can also be used for the signal processor 298. With this configuration, interference signals representing four polarization characteristics of the subject's eye E can be acquired.

[0114] Next, with reference to FIG. 19, the configuration of the control system of the ophthalmic apparatus 2 of this embodiment will be described. As shown in FIG. 19, the ophthalmic apparatus 2 is controlled by a calculation device 400. The calculation device 400 is composed of a calculation unit 402, a first interference light detection unit 285, and a second interference light detection unit 295. The first interference light detection unit 285, the second interference light detection unit 295, and the calculation unit 402 are connected to the measurement unit 5. The calculation unit 402 outputs a control signal to the measurement unit 5 and drives the scanner 106 to scan the incident position of light irradiated onto the anterior segment of the subject's eye E. The first interference light detection unit 285 acquires first sampling data based on a sampling clock 1 input from the measurement unit 5 in response to interference signals (interference signals HH and HV) input from the measurement unit 5, using a sampling trigger 1 as a trigger. The first sampling data is output to the calculation unit 402. The calculation unit 402 performs calculation processing such as Fourier transform processing on the first sampling data to generate HH tomographic images and HV tomographic images. The second interference light detection unit 295 acquires second sampling data based on the sampling clock 2 input from the measurement unit 5 in response to the interference signals (interference signals VH and VV) input from the measurement unit 5, using the sampling trigger 2 as a trigger, and outputs the second sampling data to the calculation unit 402. The calculation unit 402 performs calculation processing such as Fourier transform processing on the second sampling data to generate VH tomographic images and VV tomographic images. Here, the HH tomographic image, VH tomographic image, HV tomographic image, and VV tomographic image are tomographic images at the same position. Therefore, the calculation unit 402 can generate tomographic images of four polarization characteristics (HH, HV, VH, VV) representing the Jones matrix of the subject's eye E.

[0115] 20, the sampling trigger / clock generator 300 includes a fiber coupler 502, sampling trigger generators (540 to 552), and sampling clock generators (560 to 572). Light from the anterior eye light source 12 is input to the sampling trigger generator 540 and the sampling clock generator 560 via the fiber coupler 215 and the fiber coupler 502, respectively.

[0116] The sampling trigger generator 540 may generate a sampling trigger using, for example, an FBG (Fiber Bragg Grating) 544. As shown in FIG. 3 , the FBG 544 reflects only a specific wavelength of light incident from the anterior-segment light source 12 to generate a sampling trigger. The generated sampling trigger is input to a distributor 550. The distributor 550 distributes the sampling trigger into sampling trigger 1 and sampling trigger 2. The sampling trigger 1 is input to the calculation unit 402 via a signal delay circuit 552. The sampling trigger 2 is input directly to the calculation unit 402. The sampling trigger 1 serves as a trigger signal for the interference signals (first interference signal and second interference signal) input from the first interference light detection unit 285 to the calculation unit 402. The sampling trigger 2 serves as a trigger signal for the interference signals (third interference signal and fourth interference signal) input from the second interference light detection unit 295 to the calculation unit 402. The signal delay circuit 552 is designed so that sampling trigger 1 is delayed in time relative to sampling trigger 2 by the optical path length difference Δl of the optical path length difference generation unit 222. This makes it possible to make the frequency at which sampling of the interference signal input from the first interference light detection unit 285 starts the same as the frequency at which sampling of the interference signal input from the second interference light detection unit 295 starts. Here, only sampling trigger 1 may be generated. Because the optical path length difference Δl is known, when sampling the interference input from the second interference light detection unit 295, sampling can be started so as to be delayed in time by the optical path length difference Δl from sampling trigger 1.

[0117] The sampling clock generator 560 may be configured, for example, with a Mach-Zehnder interferometer. As shown in FIG. 20 , the sampling clock generator 560 generates sampling clocks of equal frequency using a Mach-Zehnder interferometer. The sampling clock generated by the Mach-Zehnder interferometer is input to a distributor 572. The distributor 572 distributes the sampling clock into sampling clock 1 and sampling clock 2. Sampling clock 1 passes through a signal delay circuit 574 and is input to the first interference light detector 285. Sampling clock 2 is input directly to the second interference light detector 295. The signal delay circuit 574 is designed to delay the time by the optical path length difference Δl of the optical path length difference generator 222. This allows sampling of interference light delayed by the optical path length difference generator 222 at the same timing. This prevents misalignment of multiple acquired tomographic images. In this embodiment, a Mach-Zehnder interferometer is used to generate the sampling clock. However, the sampling clock may be generated using a Michelson interferometer or an electrical circuit, or may be generated using a light source equipped with a sampling clock generator as the anterior segment light source 12.

[0118] In the ophthalmic apparatus 2 of this embodiment, when various measurements are performed on the subject's eye E, the depth position of each part of the anterior segment and the fundus and the axial length can be measured by performing the same process as in the first embodiment (see FIG. 6 , etc.). Thus, even when a PS-OCT is used as the anterior segment OCT interferometer 210, multiple types of measurements on the subject's eye E can be performed with high accuracy and efficiency. When simultaneously measuring the anterior segment and the fundus in the ophthalmic apparatus 2, the MEMS attenuator 224 may be adjusted to zero the amount of light passing through the MEMS attenuator 224, so that the light output from the polarization beam combiner / splitter 225 is set to the first measurement light alone. In this case, the amount of light irradiated from the anterior segment OCT interferometer 210 to the subject's eye E is approximately halved, thereby increasing the amount of light that the fundus OCT interferometer 211 can irradiate to the subject's eye E within the range of the amount of light that can be irradiated to the subject's eye E according to safety standards. In this case, in anterior segment OCT measurement, optical signals are detected only by detectors 286 and 296, and by averaging the tomographic images obtained from these, an interference signal can be obtained without depending on the polarization state of the light backscattered from the test eye E.

[0119] (Variation) In the above-described embodiments, an example was shown in which a range including the anterior segment and fundus of the subject's eye E was examined, but the present invention is not limited to this. For example, a range including other parts of the subject's eye E may be examined by appropriately adjusting the central wavelength of the light output from the light sources 12 and 62 or by appropriately changing the configuration of the probe optical system 26.

[0120] In addition, in each of the above-described embodiments, when the condition for causing fringe washout is satisfied in the fundus OCT interferometer 11 (i.e., when the scan resolution is greater than the optical resolution), pulsed light is output from the fundus light source 62. However, even when the condition for causing fringe washout is not satisfied, pulsed light may be output from the fundus light source 62, and even when the condition for causing fringe washout is satisfied, light may be output continuously from the fundus light source 62 as long as the accuracy of the acquired tomographic image is not affected. Note that in anterior segment OCT measurement, light from the anterior segment light source 12 may be output in pulsed form.

[0121] In addition, in the above-described embodiments, the fundus light source 62 outputs pulsed light at a predetermined duty ratio and a predetermined period to suppress fringe washout. However, instead of the above configuration, the fundus light source 62 may continuously output light while detecting interference light from the fundus of the subject's eye E at a predetermined duty ratio and a predetermined period. That is, a detection period during which the interference light is detected and a non-detection period during which the interference light is not detected may be set, or the detection period may be set at a period corresponding to the A-scan speed of the fundus OCT measurement. Specifically, for example, the interference light may be detected in pulses to achieve the values ​​of the above-described embodiments (i.e., a duty ratio of less than 10,000 / 27,036 and a period of 100 μs). When light is output in pulses from the fundus light source 62, the characteristics of the light may become unstable. However, with this configuration, the characteristics of the light output from the fundus light source 62 are stabilized and fringe washout can be suitably suppressed.

[0122] In addition, the axial length of the subject's eye generally varies from subject to subject. For this reason, in fundus OCT measurement, depending on the initial setting of the optical path length, the interference signal reflected from the retina may fall outside the imaging range of the fundus OCT. As a result, a tomographic image of the fundus cannot be captured, and it may be difficult to correct the distance from the anterior segment to the fundus using the calibration mirror 36. In such a case, the optical path length from the fundus light source 62 to the calibration mirror 36 may first be adjusted so that the interference signal reflected from the retina falls within the imaging range. For example, the amount of shift in the depth direction from the initial setting of the imaging range is set to Z. PREF Then, the reference position Z taking into account the depth shift is Ai and reference position Z Pi The optical path length difference L' is L'=L i -Z PREF In this way, by changing the imaging range in fundus OCT measurement and calculating the displacement of the optical path length difference at each reference position of the calibration mirror 36 that accompanies the change, the axial length of the subject's eye E can be suitably calculated.

[0123] In the above-described embodiments, the reflex measurement optical system 94 includes the perforated mirror 105 disposed between the dichroic mirror 104 and the scanner 106, and the ring lens 136 disposed between the lens 134 and the sensor 138. However, for example, the reflex measurement optical system 94 may include a half mirror instead of the perforated mirror 105, or a lens array having a number of minute lenses arranged in a matrix instead of the ring lens 136. In such a configuration, the light reflected from the subject's eye E is split by the half mirror, and approximately half of the reflected light is reflected by the half mirror. The light reflected by the half mirror is irradiated onto the lens array. When light is irradiated onto the lens array, the sensor 138 is irradiated with light whose optical axis is offset from the number of lenses corresponding to the distorted position among the many minute lenses. That is, by disposing a lens array between the lens 134 and the sensor 138, the same function as that of the optical system of the wavefront sensor can be imparted to the refraction measurement optical system 94. Therefore, the refraction measurement optical system 94 can measure not only the total refractive power of the subject's eye E but also the total aberration of the subject's eye E, and can measure information related to the refraction of the subject's eye E in more detail.

[0124] In addition, in each of the above-described embodiments, the refraction measurement (S14) is performed before the anterior segment OCT measurement and the fundus OCT measurement (S18 in FIG. 6) are performed. However, the refraction measurement may be performed simultaneously with the fundus OCT measurement by using the light output from the fundus light source 62. For example, if the subject has previously undergone refraction measurement or OCT measurement, the anterior segment OCT measurement, fundus OCT measurement, and refraction measurement can be performed simultaneously on the subject's eye E by adjusting the probe optical system 26 based on the results of the previous measurements. This configuration further improves the efficiency of the examination of the subject's eye E. Furthermore, the refraction measurement does not need to be performed. In this case, S14 can be omitted.

[0125] In the above-described embodiments, light output from the fundus light source 62 was used to measure the refractive power of the subject's eye E. However, a light source other than the fundus light source 62 may be used to measure the refractive power of the subject's eye E. In this case, light of a different wavelength from the light output from the anterior eye light source 12 and the fundus light source 62 may be output. For example, light with a central wavelength of 0.70 μm or more and 0.95 μm or less may be output. Light with a wavelength of 0.70 μm or more and 0.95 μm or less has high transmittance within the eye. Furthermore, light with a wavelength of 0.70 μm or more and 0.95 μm or less has low luminosity of the subject's eye E, making it less likely for the subject to feel dazzled. However, since light with a wavelength of 0.70 μm or more and 0.95 μm or less is close to visible light, it is suitable for evaluating visual function. Therefore, by using light with a wavelength of 0.70 μm or more and 0.95 μm or less to measure the refractive power, it is possible to sufficiently irradiate the fundus of the subject's eye E with light from the light source and to output light with a wavelength suitable for measuring the refractive power of the subject's eye E.

[0126] In addition, in each of the above-described embodiments, the same calibration mirror 36 is used to correct the distance from the anterior segment to the fundus. However, a calibration mirror may be provided for each of the anterior segment OCT interferometer 10 and the fundus OCT interferometer 11.

[0127] Although specific examples of the technology disclosed in this specification have been described in detail above, these are merely examples and do not limit the scope of the claims. The technology described in the claims includes various modifications and variations of the specific examples exemplified above. Furthermore, the technical elements described in this specification or drawings exhibit technical utility alone or in various combinations, and are not limited to the combinations described in the claims at the time of filing. [Explanation of symbols]

[0128] 1: Ophthalmology equipment 10: Anterior segment OCT interferometer 11: Fundus OCT interferometer 12: Anterior segment light source 20: Measurement optical system 26: Probe optical system 30: Calibration optical system 40:Reference optical system 50: Interference optical system 62: Fundus light source 66: Measurement optical system 70:Reference optical system 80: Interference optical system 90: Anterior segment OCT optical system 92: Fundus OCT optical system 94:Reflective measurement optical system 106: Scanner 200: Arithmetic device

Claims

[Claim 1] a first light source that outputs a first light to be irradiated onto the subject's eye; a second light source that outputs a second light to be irradiated onto the subject's eye; a first interferometer that performs an inspection of a first range including an anterior segment of the subject's eye based on a first interference light obtained from a reflected light of the first light; a second interferometer configured to inspect a second range including the fundus of the subject's eye and different from the first range, based on a second interference light obtained from a reflected light of the second light; a calculation device that calculates a depth position of a first portion of the anterior eye of the subject's eye based on the first interference light, and calculates a depth position of a second portion of the fundus of the subject's eye based on the second interference light; Equipped with a central wavelength of the first light is different from a central wavelength of the second light, a first optical path that is an optical path of the first light and a second optical path that is an optical path of the second light are at least partially overlapped with each other; The inspection of the first range and the inspection of the second range can be performed simultaneously, the first interferometer includes a first calibration mirror; the second interferometer includes a second calibration mirror; The computing device includes: calculating a first reference position, which is a depth position of the first calibration mirror, a second reference position, which is a depth position of the second calibration mirror, and a reference distance from the first reference position to the second reference position, based on a distance from the first portion to the second portion of the artificial eye when the artificial eye is measured; calculating a first displacement amount between a depth position of the first calibration mirror when the anterior segment of the subject's eye is measured and the first reference position; calculating a second displacement amount between a depth position of the second calibration mirror when the fundus of the subject's eye is measured and the second reference position; correcting an optical path difference from the depth position of the first calibration mirror to the depth position of the second calibration mirror when the subject's eye is measured, using the reference distance, the first displacement amount, and the second displacement amount; calculating an axial length of the subject's eye based on a depth position of the first calibration mirror and a depth position of the second calibration mirror when the subject's eye is measured, the optical path length difference after correction, and a depth position of the first portion and a depth position of the second portion of the subject's eye; Ophthalmology equipment.

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