Patent Description:
In recent years, attention has been drawn to optical coherence tomography (OCT) which is used to form images representing the surface morphology and the internal morphology of an object using light beams emitted from a laser light source or the like. Since OCT does not have invasiveness to human body as X-ray CT (Computed Tomography) does, development of application of OCT in medical field and biology field is particularly expected. For example, in the ophthalmologic field, apparatuses for forming images of the fundus, the cornea, or the like have been in practical use. Such an apparatus using OCT imaging (OCT apparatus) can be used to observe a variety of sites of a subject's eye. In addition, because of the ability to acquire high precision images, the OCT apparatus is applied to the diagnosis of various eye diseases.

In case of performing OCT measurement on a predetermined site inside the subject's eye, the measurement light for scanning the predetermined site is made incident on the eye from the pupil, and the measurement light is deflected, for example, around the scan center position arranged near the pupil. For example, an A-scan image is formed from acquired A-scan data, and a tomographic image (B-scan image) is obtained by arranging a plurality of A-scan images in a B-scan direction (for example, Patent Document <NUM>).

<CIT> discloses an OCT apparatus with image correction capabilities. It addresses distortion in tomographic images caused by the geometry of scanning and alignment eccentricity. Specifically, it discloses a controller (image generating apparatus) that includes an acquisition unit for acquiring tomographic data, an eccentric amount acquisition unit, and an image generation unit with a calculation unit and a correction unit (paragraph [<NUM>]). It teaches correction of distortion in OCT images by addressing changes in incident angle and optical path length (paragraphs [<NUM>]-[<NUM>]). D1 further discloses "actual shape correction" where data arranged on X-Y coordinates is corrected by rearranging data on the polar coordinates system having the pivot point as its center (paragraphs [<NUM>]-[<NUM>]). <CIT> discloses an ophthalmic device with optical systems for scanning a subject's eye. It teaches deflection of measurement light around a scan center position and using mechanical movement of the optical system to scan wide areas by combining multiple scan areas (paragraphs [<NUM>]-[<NUM>]). It also discusses image processing of the collected optical data.

However, in the conventional method, the contour shape of the acquired tomographic image is transformed into a rectangle. Therefore, the wider the angle of view is, the greater the difference between the shape of the predetermined site in the tomographic image and the actual shape becomes.

Further, in case of measuring an intraocular distance between two points in the acquired tomographic image, for example, the intraocular distance is obtained by multiplying the number of pixels between the two points by the pixel size, the pixel size being specific to the apparatus. Therefore, the error in the intraocular distance increases depending on the depth position.

The present invention has been made in view of such circumstances, and an object of the present invention is to provide a new technique for specifying an actual shape or the like of a predetermined site inside an eye.

According to the present invention, a new technique for specifying an actual shape or the like of a predetermined site inside an eye can be provided.

Referring now to the drawings, exemplary embodiments of an ophthalmologic information processing apparatus, an ophthalmologic apparatus, an ophthalmologic information processing method, and a program according to the present invention are described below. Any of the contents of the documents cited in the present specification and arbitrary known techniques may be applied to the embodiments below.

An ophthalmologic information processing apparatus according to embodiments corrects an OCT image of a subject's eye, two-dimensional scan data of the subject's eye, or three-dimensional scan data of the subject's eye. The OCT image, the two-dimensional scan data, or the three-dimensional scan data is acquired by scanning inside the subject's eye using an optical scanner. The optical scanner is arranged at a position optically conjugate with a predetermined site in the subject's eye. Examples of the predetermined site include a pupil. The OCT image is a two-dimensional image or a three-dimensional image. Examples of the OCT image include a tomographic image of a fundus and a three-dimensional image of the fundus. The OCT image or the scan data is acquired using optical coherence tomography (OCT). The OCT image is formed by arranging a plurality of A-scan images acquired by scanning inside the subject's eye with measurement light deflected around the predetermined site of the subject's eye. A scan center position of the scanning is arranged at the predetermined site. The two-dimensional scan data or the three-dimensional scan data is formed by arranging a plurality of A-scan data acquired by scanning inside the subject's eye with measurement light deflected around the predetermined site of the subject's eye. The scan center position of the scanning is arranged at the predetermined site.

The ophthalmologic information processing apparatus specifies a transformation position along an A-scan direction (traveling direction of the measurement light passing through the predetermined site of the subject's eye). Here, the transformation position corresponds to a pixel position in the OCT image, or a scan position in the two-dimensional scan data or the three-dimensional scan data. The ophthalmologic information processing apparatus transforms the pixel position or the scan position into the transformation position specified based on the pixel position or the like. The transformation position is a position in a predetermined coordinate system. The predetermined coordinate system is defined by two or more coordinate axes including an coordinate axis in the same axial direction as the scan direction of at least one A-scan.

In some embodiments, the ophthalmologic information processing apparatus specifies the transformation position based on a parameter representing optical characteristics of the subject's eye. In some embodiments, the ophthalmologic information processing apparatus specifies at least one of a component of a first axis direction of the transformation position and a component of a second axis direction of the transformation position, the second axis direction intersecting the first axis direction, in a predetermined coordinate system, based on a scan radius in the A-scan direction, a scan angle, a depth range that can be measured using OCT and the pixel position or the scan position.

This allows to correct the shape of the intraocular site such as the fundus represented by the OCT image or the scan data to a shape along the actual scan. In particular, the actual shape can be easily grasped from the OCT image or the scan data, which is acquired using a wide-angle imaging system or an observation system. Further, morphology information representing the morphology of the subject's eye can also be acquired as the information representing the actual morphology, using the corrected OCT image, the corrected two-dimensional scan data, or the corrected three-dimensional scan data.

An ophthalmologic information processing method according to the embodiments includes one or more steps for realizing the processing executed by a processor (computer) in the ophthalmologic information processing apparatus according to the embodiments. A program according to the embodiments causes the processor to execute each step of the ophthalmologic information processing method according to the embodiments.

The term "processor" as used herein refers to a circuit such as, for example, a central processing unit (CPU), a graphics processing unit (GPU), an application specific integrated circuit (ASIC), and a programmable logic device (PLD). Examples of PLD include a simple programmable logic device (SPLD), a complex programmable logic device (CPLD), and a field programmable gate array (FPGA). The processor realizes, for example, the function according to the embodiments by reading out a computer program stored in a storage circuit or a storage device and executing the computer program.

In this specification, an image acquired using OCT may be collectively referred to as an "OCT image". Also, the measurement operation for forming OCT images may be referred to as OCT measurement.

Hereinafter, the case where a tomographic is acquired as the OCT image of the subject's eye will be described. However, the same applies to the case where the three-dimensional image, the two-dimensional scan data, or the three-dimensional scan data is acquired using OCT.

Further, hereinafter, the case where the ophthalmologic apparatus according to the embodiments has the function of the ophthalmologic information processing apparatus according to the embodiments will be described. However, the ophthalmologic information processing apparatus according to the embodiments may be configured to acquire the OCT image, the two-dimensional scan data, or the three-dimensional scan data from an external ophthalmologic apparatus.

Hereinafter, in the embodiments, the case of using the swept source type OCT method in the measurement or the imaging (photographing) using OCT will be described. However, the configuration according to the embodiments can also be applied to an ophthalmologic apparatus using other type of OCT (for example, spectral domain type OCT or time domain OCT).

The ophthalmologic apparatus according to some embodiments includes any one or more of an ophthalmologic imaging apparatus, an ophthalmologic measuring apparatus, and an ophthalmologic therapy apparatus. The ophthalmologic imaging apparatus included in the ophthalmologic apparatus according to some embodiments includes, for example, any one or more of a fundus camera, a scanning laser ophthalmoscope, a slit lamp ophthalmoscope, a surgical microscope, and the like. Further, the ophthalmologic measuring apparatus included in the ophthalmologic apparatus according to some embodiments includes any one or more of an eye refractivity examination apparatus, a tonometer, a specular microscope, a wave-front analyzer, a perimeter, a microperimeter, and the like, for example. Further, the ophthalmologic therapy apparatus included in the ophthalmologic apparatus according to some embodiments includes any one or more of a laser therapy apparatus, a surgical apparatus, a surgical microscope, and the like, for example.

The ophthalmologic apparatus according to the following embodiments includes an OCT apparatus and a fundus camera. The OCT apparatus can perform OCT measurement. Alternatively, the configuration according to the following embodiments may be applied to a single-functional OCT apparatus.

Hereinafter, an ophthalmologic apparatus capable of performing OCT measurement on a fundus of the subject's eye will be described as an example. However, the ophthalmologic apparatus according to the embodiments may be capable of performing OCT measurement on an anterior segment of the subject's eye. In some embodiments, a measurement site of the OCT measurement and / or a range of the OCT measurement are changed by moving a lens for changing focal position of the measurement light. In some embodiments, the ophthalmologic apparatus has a configuration capable of performing OCT measurement on the fundus, OCT measurement on the anterior segment, and OCT measurement on the whole eyeball including the fundus and anterior segment, by adding one or more attachments (objective lens, front lens, etc.). In some embodiments, in the ophthalmologic apparatus for measuring fundus, OCT measurement is performed on the anterior segment, by making the measurement light incident on the subject's eye, the measurement light having been converted into a parallel light flux by arranging a front lens between the objective lens and the subject's eye.

As shown in <FIG>, the ophthalmologic apparatus <NUM> includes a fundus camera unit <NUM>, an OCT unit <NUM>, and an arithmetic control unit <NUM>. The fundus camera unit <NUM> is provided with an optical system and a mechanism for acquiring front images of a subject's eye E. The OCT unit <NUM> is provided with a part of an optical system and a mechanism for performing OCT. Another part of the optical system and the mechanism for performing OCT are provided in the fundus camera unit <NUM>. The arithmetic control unit <NUM> includes one or more processors for performing various kinds of arithmetic processing and control processing. In addition to these elements, an arbitrary element or a unit, such as a member (chin rest, forehead pad, etc.) for supporting a face of the subject, a lens unit (for example, an attachment for an anterior segment OCT) for switching the target site of OCT, and the like, may be provided in the ophthalmologic apparatus <NUM>. In some embodiments, the lens unit is configured to be manually inserted and removed between the subject's eye E and an objective lens <NUM> described later. In some embodiments, the lens unit is configured to be automatically inserted and removed between the subject's eye E and the objective lens <NUM> described later, under the control of the controller <NUM> described later.

In some embodiments, the ophthalmologic apparatus a includes a display apparatus <NUM>. The display apparatus <NUM> displays a processing result (for example, an OCT image or the like) obtained by the arithmetic control unit <NUM>, an image obtained by the fundus camera unit <NUM>, operation guidance information for operating the ophthalmologic apparatus <NUM>, and the like.

The fundus camera unit <NUM> is provided with an optical system for imaging (photographing) a fundus Ef of the subject's eye E. An image (called fundus image, fundus photograph, etc.) of the fundus Ef to be obtained is a front image such as an observation image, a photographic image, or the like. The observation image is obtained by moving image shooting using near infrared light. The photographic image is a still image using flash light. Furthermore, the fundus camera unit <NUM> can obtain the front image (anterior segment image) by photographing (imaging) an anterior segment Ea of the subject's eye E.

The fundus camera unit <NUM> includes an illumination optical system <NUM> and an imaging (photographing) optical system <NUM>. The illumination optical system <NUM> projects illumination light onto the subject's eye E. The imaging optical system <NUM> detects returning light of the illumination light from the subject's eye E. Measurement light from the OCT unit <NUM> is guided to the subject's eye E through an optical path in the fundus camera unit <NUM>. Returning light of the measurement light is guided to the OCT unit <NUM> through the same optical path.

Light (observation illumination light) emitted from the observation light source <NUM> of the illumination optical system <NUM> is reflected by a reflective mirror <NUM> having a curved reflective surface, and becomes near-infrared light after penetrating a visible cut filter <NUM> via a condenser lens <NUM>. Further, the observation illumination light is once converged near an imaging light source <NUM>, is reflected by a mirror <NUM>, and passes through relay lenses <NUM> and <NUM>, a diaphragm <NUM>, and a relay lens <NUM>. Then, the observation illumination light is reflected on the peripheral part (the surrounding area of a hole part) of a perforated mirror <NUM>, penetrates a dichroic mirror <NUM>, and is refracted by an objective lens <NUM>, thereby illuminating the subject's eye E (fundus Ef or anterior segment Ea). Returning light of the observation illumination light reflected from the subject's eye E is refracted by the objective lens <NUM>, penetrates the dichroic mirror <NUM>, passes through the hole part formed in the center region of the perforated mirror <NUM>, penetrates a dichroic mirror <NUM>. The returning light penetrating the dichroic mirror <NUM> travels through a photography focusing lens <NUM> and is reflected by a mirror <NUM>. Further, this returning light penetrates a half mirror 33A, is reflected by a dichroic mirror <NUM>, and forms an image on the light receiving surface of an image sensor <NUM> by a condenser lens <NUM>. The image sensor <NUM> detects the returning light at a predetermined frame rate. It should be noted that the focus of the imaging optical system <NUM> is adjusted so as to coincide with the fundus Ef or the anterior segment Ea.

Light (imaging illumination light) emitted from the imaging light source <NUM> is projected onto the fundus Ef via the same route as that of the observation illumination light. Returning light of the imaging illumination light from the subject's eye E is guided to the dichroic mirror <NUM> via the same route as that of the observation illumination light, penetrates the dichroic mirror <NUM>, is reflected by a mirror <NUM>, and forms an image on the light receiving surface of the image sensor <NUM> by a condenser lens <NUM>.

A liquid crystal display (LCD) <NUM> displays a fixation target and a visual target used for visual acuity measurement. Part of light output from the LCD <NUM> is reflected by the half mirror 33A, is reflected by the mirror <NUM>, travels through the photography focusing lens <NUM> and the dichroic mirror <NUM>, and passes through the hole part of the perforated mirror <NUM>. The light flux (beam) having passed through the hole part of the perforated mirror <NUM> penetrates the dichroic mirror <NUM>, and is refracted by the objective lens <NUM>, thereby being projected onto the fundus Ef.

By changing the display position of the fixation target on the screen of the LCD <NUM>, the fixation position of the subject's eye E can be changed. Examples of the fixation position include a fixation position for acquiring an image centered at a macula, a fixation position for acquiring an image centered at an optic disc, a fixation position for acquiring an image centered at a fundus center between the macula and the optic disc, a fixation position for acquiring an image of a site (fundus peripheral part) far away from the macula, and the like. The ophthalmologic apparatus <NUM> according to some embodiments includes GUI (Graphical User Interface) and the like for designating at least one of such fixation positions. The ophthalmologic apparatus <NUM> according to some embodiments includes GUI etc. for manually moving the fixation position (display position of the fixation target).

The configuration for presenting the movable fixation target to the subject's eye E is not limited to the display device such LCD or the like. For example, the movable fixation target can be generated by selectively turning on a plurality of light sources of a light source array (light emitting diode (LED) array or the like). Alternatively, the movable fixation target can be generated using one or more movable light sources.

Further, the ophthalmologic apparatus <NUM> may be provided with one or more external fixation light sources. One of the one or more external fixation light sources can project fixation light onto a fellow eye of the subject's eye E. A projected position of the fixation light on the fellow eye can be changed. By changing the projected position of the fixation light on the fellow eye, the fixation position of the subject's eye E can be changed. The fixation position projected by the external fixation light source(s) may be the same as the fixation position of the subject's eye E using the LCD <NUM>. For example, the movable fixation target can be generated by selectively turning on a plurality of external fixation light sources. Alternatively, the movable fixation target can be generated using one or more movable external fixation light sources.

The alignment optical system <NUM> generates an alignment indicator for alignment of the optical system with respect to the subject's eye E. Alignment light emitted from an LED <NUM> travels through the diaphragms <NUM> and <NUM> and the relay lens <NUM>, is reflected by the dichroic mirror <NUM>, and passes through the hole part of the perforated mirror <NUM>. The alignment light having passed through the hole part of the perforated mirror <NUM> penetrates the dichroic mirror <NUM>, and is projected onto the subject's eye E by the objective lens <NUM>. Corneal reflection light of the alignment light is guided to the image sensor <NUM> through the same route as the returning light of the observation illumination light. Manual alignment or automatic alignment can be performed based on the received light image (alignment indicator image) thereof.

The focus optical system <NUM> generates a split indicator for adjusting the focus with respect to the subject's eye E. The focus optical system <NUM> is movable along an optical path (illumination optical path) of the illumination optical system <NUM> in conjunction with the movement of the photography focusing lens <NUM> along an optical path (imaging optical path) of the imaging optical system <NUM>. The reflection rod <NUM> can be inserted and removed into and from the illumination optical path. To conduct focus adjustment, the reflective surface of the reflection rod <NUM> is arranged in a slanted position on the illumination optical path. Focus light emitted from an LED <NUM> passes through a relay lens <NUM>, is split into two light beams by a split indicator plate <NUM>, passes through a two-hole diaphragm <NUM>, is reflected by a mirror <NUM>, and is reflected after an image is once formed on the reflective surface of the reflection rod <NUM> by a condenser lens <NUM>. Further, the focus light travels through the relay lens <NUM>, is reflected by the perforated mirror <NUM>, penetrates the dichroic mirror <NUM>, and is refracted by the objective lens <NUM>, thereby being projected onto the fundus Ef. Fundus reflection light of the focus light is guided to the image sensor <NUM> through the same route as the corneal reflection light of the alignment light. Manual focus or automatic focus can be performed based on the received light image (split indicator image) thereof.

The dichroic mirror <NUM> combines an optical path for fundus photography and an optical path for OCT. The dichroic mirror <NUM> reflects light of wavelength band used in OCT, and transmits light for fundus photography. The optical path for OCT (optical path of measurement light) is provided with, in order from the OCT unit <NUM> side to the dichroic mirror <NUM> side, a collimator lens unit <NUM>, an optical path length changing unit <NUM>, an optical scanner <NUM>, an OCT focusing lens <NUM>, a mirror <NUM>, and a relay lens <NUM>.

The optical path length changing unit <NUM> is movable in directions indicated by the arrow in <FIG>, thereby changing the length of the optical path for OCT. This change in the optical path length is used for correcting the optical path length according to the axial length, adjusting the interference state, or the like. The optical path length changing unit <NUM> includes a corner cube and a mechanism for moving the corner cube.

The optical scanner <NUM> is disposed at a position optically conjugate with the pupil of the subject's eye E. The optical scanner <NUM> deflects the measurement light traveling along the OCT optical path. That is, the optical scanner <NUM> deflects the measurement light for scanning inside the subject's eye E while changing the scan angle within a predetermined deflection angle range with the pupil (or the vicinity thereof) of the subject's eye E as the scan center position. The optical scanner <NUM> can deflect the measurement light in a one-dimensionally or two-dimensional manner.

In case that the optical scanner <NUM> deflects the measurement light in a one-dimensionally manner, the optical scanner <NUM> includes a galvano scanner capable of deflecting the measurement light in a predetermined deflection direction within a predetermined deflection angle range. In case that the optical scanner deflects the measurement light LS in a two-dimensionally manner, the optical scanner <NUM> includes a first galvano scanner and a second galvano scanner. The first galvano scanner deflects the measurement light so as to scan a photographing (imaging) site (fundus Ef or the anterior segment) in a horizontal direction orthogonal to the optical axis of the OCT optical system <NUM>. The second galvano scanner deflects the measurement light deflected by the first galvano mirror so as to scan the photographing site in a vertical direction orthogonal to the optical axis of the OCT optical system <NUM>. Examples of scan mode with the measurement light performed by the optical scanner <NUM> include horizontal scan, vertical scan, cross scan, radial scan, circle scan, concentric scan, helical (spiral) scan, and the like.

The OCT focusing lens <NUM> is moved along the optical path of the measurement light in order to perform focus adjustment of the optical system for OCT. The OCT focusing lens <NUM> can move within a moving range. The moving range includes a first lens position for placing the focal position of the measurement light at the fundus Ef or near the fundus Ef of the subject's eye E and a second lens position for making the measurement light projected onto the subject's eye E a parallel light beam. The movement of the photography focusing lens <NUM>, the movement of the focus optical system <NUM>, and the movement of the OCT focusing lens <NUM> can be controlled in conjunction with each other.

An example of the configuration of the OCT unit <NUM> is shown in <FIG>. The OCT unit <NUM> is provided with an optical system for acquiring OCT images of the subject's eye E. The optical system includes an interference optical system that splits light from a wavelength sweeping type (i.e., a wavelength scanning type) light source into measurement light and reference light, makes the measurement light returning from the subject's eye E and the reference light having traveled through the reference optical path interfere with each other to generate interference light, and detects the interference light. The detection result of the interference light obtained by the interference optical system (i.e., the detection signal) is an interference signal indicating the spectrum of the interference light, and is sent to the arithmetic control unit <NUM>.

Like swept source type ophthalmologic apparatuses commonly used, the light source unit <NUM> includes a wavelength sweeping type (i.e., a wavelength scanning type) light source capable of sweeping (scanning) the wavelengths of emitted light. The wavelength sweeping type light source includes a laser light source that includes a resonator. The light source unit <NUM> temporally changes the output wavelengths within the near-infrared wavelength bands that cannot be visually recognized with human eyes.

Light L0 output from the light source unit <NUM> is guided to the polarization controller <NUM> through the optical fiber <NUM>, and the polarization state of the light L0 is adjusted. The polarization controller <NUM>, for example, applies external stress to the looped optical fiber <NUM> to thereby adjust the polarization state of the light L0 guided through the optical fiber <NUM>.

The light L0 whose the polarization state has been adjusted by the polarization controller <NUM> is guided to the fiber coupler <NUM> through the optical fiber <NUM>, and is split into the measurement light LS and the reference light LR.

The reference light LR is guided to the collimator <NUM> through the optical fiber <NUM>. The reference light LR is converted into a parallel light beam by the collimator <NUM>. Then, the reference light LR is guided to the optical path length changing unit <NUM> via an optical path length correction member <NUM> and a dispersion compensation member <NUM>. The optical path length correction member <NUM> acts so as to match the optical path length of the reference light LR with the optical path length of the measurement light LS. The dispersion compensation member <NUM> acts so as to match the dispersion characteristics between the reference light LR and the measurement light LS.

The optical path length changing unit <NUM> is movable in directions indicated by the arrow in <FIG>, thereby changing the length of the optical path of the reference light LR. Through such movement, the length of the optical path of the reference light LR is changed. The change in the optical path length is used for the correction of the optical path length according to the axial length of the subject's eye E, for the adjustment of the interference state, or the like. The optical path length changing unit <NUM> includes, for example, a corner cube and a movement mechanism for moving the corner cube. In this case, the corner cube in the optical path length changing unit <NUM> changes the traveling direction of the reference light LR that has been made into the parallel light flux by the collimator <NUM> in the opposite direction. The optical path of the reference light LR incident on the corner cube and the optical path of the reference light LR emitted from the corner cube are parallel.

The reference light LR that has traveled through the optical path length changing unit <NUM> passes through the dispersion compensation member <NUM> and the optical path length correction member <NUM>, is converted from the parallel light beam to the convergent light beam by a collimator <NUM>, and enters an optical fiber <NUM>. The reference light LR that has entered the optical fiber <NUM> is guided to the polarization controller <NUM>. With the polarization controller <NUM>, the polarization state of the reference light LR is adjusted. The polarization controller <NUM> has the same configuration as, for example, the polarization controller <NUM>. The reference light LR whose the polarization state has been adjusted by the polarization controller <NUM> is guided to the attenuator <NUM> through the optical fiber <NUM>, and the light amount thereof is adjusted by the attenuator <NUM> under the control of the arithmetic control unit <NUM>. The reference light LR whose light amount has been adjusted by the attenuator <NUM> is guided to the fiber coupler <NUM> by the optical fiber <NUM>.

The configuration shown in <FIG> and <FIG> includes both the optical path length changing unit <NUM> that changes the length of the optical path of the measurement light LS (i.e., measurement optical path or measurement arm) and the optical path length changing unit <NUM> that changes the length of the optical path of the reference light LR (i.e., reference optical path or reference arm). However, any one of the optical path length changing units <NUM> and <NUM> may be provided. The difference between the measurement optical path length and the reference optical path length can be changed by using other optical members.

Meanwhile, the measurement light LS generated by the fiber coupler <NUM> is guided through the optical fiber <NUM>, and is made into the parallel light beam by the collimator lens unit <NUM>. The measurement light LS made into the parallel light flux is guided to the dichroic mirror <NUM> via the optical path length changing unit <NUM>, the optical scanner <NUM>, the OCT focusing lens <NUM>, the mirror <NUM>, and the relay lens <NUM>. The measurement light LS guided to the dichroic mirror <NUM> is reflected by the dichroic mirror <NUM>, refracted by the objective lens <NUM>, and projected onto the subject's eye E. The measurement light LS is scattered (and reflected) at various depth positions of the subject's eye E. The returning light of the measurement light LS including such backscattered light advances through the same path as the outward path in the opposite direction and is led to the fiber coupler <NUM>, and then reaches the fiber coupler <NUM> through the optical fiber <NUM>.

The fiber coupler <NUM> combines (interferes) the measurement light LS incident through the optical fiber <NUM> and the reference light LR incident through the optical fiber <NUM> to generate interference light. The fiber coupler <NUM> generates a pair of interference light LC by splitting the interference light generated from the measurement light LS and the reference light LR at a predetermined splitting ratio (for example, <NUM> : <NUM>). The pair of the interference light LC emitted from the fiber coupler <NUM> is guided to the detector <NUM> through the optical fibers <NUM> and <NUM>, respectively.

The detector <NUM> is, for example, a balanced photodiode that includes a pair of photodetectors for respectively detecting the pair of interference light LC and outputs the difference between the pair of detection results obtained by the pair of photodetectors. The detector <NUM> sends the detection result (i.e., interference signal) to the data acquisition system (DAQ) <NUM>. A clock KC is supplied from the light source unit <NUM> to the DAQ <NUM>. The clock KC is generated in the light source unit <NUM> in synchronization with the output timing of each wavelength sweeping (scanning) within a predetermined wavelength range performed by the wavelength sweeping type light source. For example, the light source unit <NUM> optically delays one of the two pieces of branched light obtained by branching the light L0 of each output wavelength, and then generates the clock KC based on the result of the detection of the combined light of the two pieces of branched light. The DAQ <NUM> performs sampling of the detection result obtained by the detector <NUM> based on the clock KC. The DAQ <NUM> sends the result of the sampling of the detection result obtained by the detector <NUM> to the arithmetic control unit <NUM>. For example, the arithmetic control unit <NUM> performs the Fourier transform etc. on the spectral distribution based on the detection result obtained by the detector <NUM> for each series of wavelength scanning (i.e., for each A-line). With this, the reflection intensity profile for each A-line is formed. In addition, the arithmetic control unit <NUM> forms image data by applying imaging processing to the reflection intensity profiles for the respective A-lines.

The arithmetic control unit <NUM> analyzes the detection signals fed from the DAQ <NUM> to form an OCT image or scan data of the fundus Ef (or the anterior segment Ea). The arithmetic processing therefor is performed in the same manner as in the conventional swept-source-type OCT apparatus.

Further, the arithmetic control unit <NUM> controls each part of the fundus camera unit <NUM>, the display apparatus <NUM>, and the OCT unit <NUM>.

Also, as the control of the fundus camera unit <NUM>, the arithmetic control unit <NUM> performs following controls: the operation control of the observation light source <NUM>, of the imaging light source <NUM> and of the LEDs <NUM> and <NUM>; the operation control of the LCD <NUM>; the movement control of the photography focusing lens <NUM>; the movement control of the OCT focusing lens <NUM>; the movement control of the reflection rod <NUM>; the movement control of the focus optical system <NUM>; the movement control of the optical path length changing unit <NUM>; the operation control of the optical scanner <NUM>, and the like.

For example, the arithmetic control unit <NUM> controls the display apparatus <NUM> to display the OCT image of the subject's eye E.

Further, as the control of the OCT unit <NUM>, the arithmetic control unit <NUM> controls: the operation of the light source unit <NUM>; the operation of the optical path length changing unit <NUM>; the operations of the attenuator <NUM>; the operation of the polarization controllers <NUM> and <NUM>; the operation of the detector <NUM>; the operation of the DAQ <NUM>; and the like.

As in the conventional computer, the arithmetic control unit <NUM> includes a processor, RAM, ROM, hard disk drive, and communication interface, for example. A storage device such as the hard disk drive stores a computer program for controlling the ophthalmologic apparatus <NUM>. The arithmetic control unit <NUM> may include various kinds of circuitry such as a circuit board for forming OCT images. In addition, the arithmetic control unit <NUM> may include an operation device (or an input device) such as a keyboard and a mouse, and a display device such as an LCD.

The fundus camera unit <NUM>, the display apparatus <NUM>, the OCT unit <NUM>, and the arithmetic control unit <NUM> may be integrally provided (i.e., in a single housing), or they may be separately provided in two or more housings.

<FIG> and <FIG> illustrate a configuration example of a control system of the ophthalmologic apparatus <NUM>. In <FIG> and <FIG>, a part of the components included in the ophthalmologic apparatus <NUM> is omitted.

The controller <NUM> executes various controls. The controller <NUM> includes a main controller <NUM> and a storage unit <NUM>.

The main controller <NUM> includes a processor and controls each part of the ophthalmologic apparatus <NUM>. For example, the main controller <NUM> controls components of the fundus camera unit <NUM> such as focusing drivers 31A and 43A, the image sensors <NUM> and <NUM>, the LCD <NUM>, the optical path length changing unit <NUM>, the optical scanner <NUM>, and a movement mechanism <NUM> for moving the optical system. Further, the main controller <NUM> controls components of the OCT unit <NUM> such as the light source unit <NUM>, the optical path length changing unit <NUM>, the attenuator <NUM>, the polarization controllers <NUM> and <NUM>, the detector <NUM>, and the DAQ <NUM>.

For example, the main controller <NUM> controls the LCD <NUM> to display the fixation target at a position on the screen of the LCD <NUM> corresponding the fixation position set manually or automatically. Moreover, the main controller <NUM> can change the display position of the fixation target displayed on the LCD <NUM> (in a continuous manner or in a phased manner). Thereby, the fixation target can be moved (that is, the fixation position can be changed). The display position of the fixation target and movement mode of the fixation target are set manually or automatically. Manual setting is performed using GUI, for example. Automatic setting is performed by the data processor <NUM>, for example.

The focusing driver 31A moves the photography focusing lens <NUM> in the direction along the optical axis of the imaging optical system <NUM>, and moves the focus optical system <NUM> in the direction along the optical axis of the illumination optical system <NUM>. With this, the focus position of the imaging optical system <NUM> is changed. The focusing driver 31A may include a dedicated mechanism for moving the photography focusing lens <NUM> and a dedicated mechanism for moving the focus optical system <NUM>. The focusing driver 31A is controlled when performing focus adjustment or the like.

The focusing driver 43A moves the OCT focusing lens <NUM> in the optical axis direction of the measurement optical path. As a result, the focus position of the measurement light LS is changed. For example, the focus position of the measurement light LS can be arranged at the fundus Ef or near the fundus Ef by moving the OCT focusing lens <NUM> to the first lens position. For example, the focus position of the measurement light LS can be arranged at a far point position by moving the OCT focusing lens <NUM> to the second lens position. The focus position of the measurement light LS corresponds to the depth position (z position) of the beam waist of the measurement light LS.

The movement mechanism <NUM> three-dimensionally moves at least the fundus camera unit <NUM> (optical system), for example. In a typical example, the movement mechanism <NUM> includes a mechanism for moving at least the fundus camera unit <NUM> in the x direction (left-right direction, horizontal direction), a mechanism for moving it in the y direction (up-down direction, vertical direction), and a mechanism for moving it in the z direction (depth direction, front-back direction). The mechanism for moving in the x direction includes a x stage movable in the x direction and a x movement mechanism for moving the x stage, for example. The mechanism for moving in the y direction includes a y stage movable in the y direction and a y movement mechanism for moving the y stage, for example. The mechanism for moving in the z direction includes a z stage movable in the z direction and a z movement mechanism for moving the z stage, for example. Each movement mechanism includes an actuator such as a pulse motor, and operates under the control of the main controller <NUM>.

The control for the movement mechanism <NUM> is used for alignment and tracking. Here, tracking is to move the optical system of the apparatus according to the movement of the subject's eye E. To perform tracking, alignment and focus adjustment are performed in advance. The tracking is a function of maintaining a suitable positional relationship in which alignment and focusing are matched by causing the position of the optical system of the apparatus and the like to follow the eye movement. In some embodiments, the movement mechanism <NUM> is configured to be controlled to change the optical path length of the reference light (that is, the difference of the optical path length between the optical path of the measurement light and the optical path of the reference light).

In the case of manual alignment, a user operates a user interface (UI) <NUM> described later to relatively move the optical system and subject's eye E so as to cancel the displacement of the subject's eye E with respect to the optical system. For example, the main controller <NUM> controls the movement mechanism <NUM> to relatively move the optical system and the subject's eye E by outputting a control signal corresponding to the operation content with respect to the user interface <NUM> to the movement mechanism <NUM>.

In the case of automatic alignment, the main controller <NUM> controls the movement mechanism <NUM> to relatively move the optical system and the subject's eye E so as to cancel the displacement of the subject's eye E with respect to the optical system. For example, the movement mechanism <NUM> is controlled so as to cancel a displacement between (a reference position of) the image of the subject's eye E acquired using imaging optical system <NUM> and a reference position of the optical system. In some embodiments, the main controller <NUM> controls the movement mechanism <NUM> to relatively move the optical system and the subject's eye E by outputting a control signal to the movement mechanism <NUM> so that the optical axis of the optical system substantially coincides with the axis of the subject's eye E and the distance of the optical system with respect to the subject's eye E is a predetermined working distance. Here, the working distance is a preset value which is called a working distance of the objective lens <NUM>, and it means the distance between the subject's eye E and the optical system when measuring (imaging) using the optical system.

The main controller <NUM> controls the fundus camera unit <NUM> etc. to control the fundus imaging (photography) and the anterior segment imaging. Further, the main controller <NUM> controls the fundus camera unit <NUM> and the OCT unit <NUM> etc. to control the OCT measurement. The main controller <NUM> is capable of performing a plurality of preliminary operations prior to OCT measurement. Examples of the preliminary operation include alignment, rough focus adjustment, polarization adjustment, and fine focus adjustment. The plurality of preliminary operations is performed in a predetermined order. In some embodiments, the plurality of preliminary operations is performed in an order described above.

It should be noted that the types and the orders of the preliminary operations are not so limited, and they may be optional. For example, the preliminary operations may further include small-pupil judgment. The small-pupil judgment is a preliminary operation to judge whether the pupil of the subject's eye E is small or not (whether the subject's eye E is microcoria or not). The small-pupil judgment may be performed between the rough focus adjustment and the optical path length difference adjustment. In some embodiments, the small-pupil judgment includes, for example, a series of processes as follows: acquiring a front image (anterior segment image) of the subject's eye E; specifying an image region corresponding to the pupil; calculating the size (e.g., diameter, circumference length) of the pupil region; judging whether the pupil of the subject's eye E is small or not based on the calculated size (threshold processing); and controlling the diaphragm <NUM> when judged that the pupil of the subject's eye E is small. In some embodiments, the calculation of the size of the pupil region includes processing of circularly or elliptically approximating the pupil region.

The rough focus adjustment is a kind of focus adjustment using the split indicator. The rough focus adjustment may be performed by determining the position of the photography focusing lens <NUM> based on information, which is obtained by associating the eye refractive power acquired in advance with the position of the photography focusing lens <NUM>, and a measured value of the refractive power of the subject's eye E.

The fine focus adjustment is performed on the basis of interference sensitivity of OCT measurement. For example, the fine focus adjustment can be performed by: monitoring interference intensity (interference sensitivity) of interference signal acquired by performing OCT measurement of the subject's eye E; searching the position of the OCT focusing lens <NUM> so as to maximize the interference intensity; and moving the OCT focusing lens <NUM> to the searched position.

To perform the optical path length difference adjustment, the optical system is controlled so that a predetermined position on the subject's eye E is a reference position of a measurement range in the depth direction. The control is performed on at least one of the optical path length changing units <NUM> and <NUM>. Thereby, the difference of the optical path length between the measurement optical path and the reference optical path is adjusted. By setting the reference position in the optical path length difference adjustment, OCT measurement can be performed with high accuracy over a desired measurement range in the depth direction simply by changing the wavelength sweep speed.

To perform the polarization adjustment, the polarization state of the reference light LR is adjusted for optimizing the interference efficiency between the measurement light LS and the reference light LR.

The storage unit <NUM> stores various types of data. Examples of the data stored in the storage unit <NUM> include image data of an OCT image, image data of a fundus image, scan data, image data of an anterior segment image, and subject's eye information. The subject's eye information includes information on the subject such as patient ID and name, and information on the subject's eye such as identification information of the left eye / right eye.

Further, the storage unit <NUM> stores an eyeball parameter 212A. The eyeball parameter 212A includes a parameter (standard value) defined by a known eyeball model such as a Gullstrand schematic eye. In some embodiments, the eyeball parameter 212A includes a parameter in which at least one of the parameters defined by a known eyeball model is replaced with the measured value of the subject's eye E. In this case, it means that the eyeball parameter 212A includes a parameter representing optical characteristics of the subject's eye E. Examples of the measured value include an axial length, a corneal thickness, a curvature radius of an anterior surface of cornea, a curvature radius of a posterior surface of cornea, an anterior chamber depth, a curvature radius of an anterior surface of a lens, a lens thickness, a curvature radius of a posterior surface of lens, a vitreous cavity length, a retinal thickness, and a choroid thickness. In some embodiments, the measured value is acquired by analyzing OCT data obtained by performing OCT measurement. The eyeball parameter 212A may include a parameter designated using the operation unit 240B described later.

In addition, the storage unit <NUM> stores various kinds of computer programs and data for operating the ophthalmologic apparatus <NUM>.

The image forming unit <NUM> performs signal processing such as the Fourier transform on sampling data obtained by sampling the detection signal from the detector <NUM> in the DAQ <NUM>. With this, the reflection intensity profile for each A-line is formed. The above signal processing includes noise removal (noise reduction), filtering, fast Fourier transform (FFT), and the like. The reflection intensity profile for the A-line is an example of the A-scan data. The image forming unit <NUM> can form the reflection intensity profile for each A-line, and form B-scan data (two-dimensional scan data) by arranging a formed plurality of reflection intensity profiles in the B-scan direction (intersecting direction of the A-scan direction).

In some embodiments, the image forming unit <NUM> (or the data processor <NUM> described later) forms three-dimensional scan data by arranging the plurality of reflection intensity profiles formed for each A-line in the B-scan direction (for example, x direction) and a direction intersecting both of the A-scan direction and the B-scan direction (for example, y direction).

Further, the image forming unit <NUM> can form A-scan image (OCT image, image data) of the subject's eye E, by applying imaging processing to the reflection intensity profile in the A-line. The image forming unit <NUM> can form a B-scan image by arranging the plurality of A-scan images formed for each A-line in the B-scan direction (intersecting direction of the A-scan direction).

In some embodiments, the image forming unit <NUM> extracts data at a predetermined depth position (scan position) in each A-scan data, and forms C-scan data by arranging the extracted plurality of data in the B-scan direction (intersecting direction of the A-scan direction). In some embodiments, the image forming unit <NUM> extracts a pixel at a predetermined depth position (scan position) in each A-scan image, and forms a C-scan image by arranging the extracted plurality of pixels in the B-scan direction (intersecting direction of the A-scan direction).

In some embodiments, the function of the image forming unit <NUM> is realized by a processor. Note that "image data" and an "image" based on the image data may not be distinguished from each other in the present specification.

The data processor <NUM> processes data acquired through photography of the subject's eye E or data acquired through OCT measurement.

For example, the data processor <NUM> performs various kinds of image processing and various kinds of analysis processing on the image formed by the image forming unit <NUM>. For example, the data processor <NUM> performs various types of image correction such as brightness correction. The data processor <NUM> performs various kinds of image processing and various kinds of analysis on images captured by the fundus camera unit <NUM> (e.g., fundus images, anterior segment images, etc.).

The data processor <NUM> performs known image processing such as interpolation for interpolating pixels in tomographic images to form three-dimensional image data of the fundus Ef. Note that image data of a three-dimensional image means image data in which the position of a pixel is defined by a three-dimensional coordinate system. Examples of the image data of the three-dimensional image include image data defined by voxels three-dimensionally arranged. Such image data is referred to as volume data or voxel data. When displaying an image based on volume data, the data processor <NUM> performs rendering (volume rendering, maximum intensity projection (MIP), etc.) on the volume data, thereby forming image data of a pseudo three-dimensional image viewed from a particular line of sight. The pseudo three-dimensional image is displayed on the display device such as a display unit 240A.

The three-dimensional image data may be stack data of a plurality of tomographic images. The stack data is image data formed by three-dimensionally arranging tomographic images along a plurality of scan lines based on positional relationship of the scan lines. That is, the stack data is image data formed by representing tomographic images, which are originally defined in their respective two-dimensional coordinate systems, by a single three-dimensional coordinate system. That is, the stack data is image data formed by embedding tomographic images into a single three-dimensional space.

The data processor <NUM> can form a B-mode image (longitudinal cross-sectional image, axial cross-sectional image) in an arbitrary cross section, a C-mode image (transverse section image, horizontal cross-sectional image) in an arbitrary cross section, a projection image, a shadowgram, etc., by performing various renderings on the acquired three-dimensional data set (volume data, stack data, etc.). An image in an arbitrary cross section such as the B-mode image or the C-mode image is formed by selecting pixels (voxels) on a designated cross section from the three-dimensional data set. The projection image is formed by projecting the three-dimensional data set in a predetermined direction (z direction, depth direction, axial direction). The shadowgram is formed by projecting a part of the three-dimensional data set (for example, partial data corresponding to a specific layer) in a predetermined direction. An image having a viewpoint on the front side of the subject's eye, such as the C-mode image, the projection image, and the shadowgram, is called a front image (en-face image).

The data processor <NUM> can build (form) the B-mode image or the front image (blood vessel emphasized image, angiogram) in which retinal blood vessels and choroidal blood vessels are emphasized (highlighted), based on data (for example, B-scan image data) acquired in time series by OCT. For example, the OCT data in time series can be acquired by repeatedly scanning substantially the same site of the subject's eye E.

In some embodiments, the data processor <NUM> compares the B-scan images in time series acquired by B-scan for substantially the same site, converts the pixel value of a change portion of the signal intensity into a pixel value corresponding to the change portion, and builds the emphasized image in which the change portion is emphasized. Further, the data processor <NUM> forms an OCTA image by extracting information of a predetermined thickness at a desired site from a plurality of built emphasized images and building as an en-face image.

An image (for example, a three-dimensional image, a B-mode image, a C-mode image, a projection image, a shadowgram, and an OCTA image) generated by the data processor <NUM> is also included in the OCT image.

Further, the data processor <NUM> determines the focus state of the measurement light LS in fine focus adjustment control by analyzing the detection result of the interference light obtained by the OCT measurement. For example, the main controller <NUM> performs repetitive OCT measurements while controlling the focusing driver 43A according to a predetermined algorithm. The data processor <NUM> analyzes detection results of interference light LC repeatedly acquired by the OCT measurements to calculate predetermined evaluation values relating to image quality of OCT images. The data processor <NUM> determines whether the calculated evaluation value is equal to or less than a threshold. In some embodiments, the fine focus adjustment is continued until the calculated evaluation value becomes equal to or less than the threshold. That is, when the evaluation value is equal to or less than the threshold, it is determined that the focus state of the measurement light LS is appropriate. And the fine focus adjustment is continued until it is determined that the focus state of the measurement light LS is appropriate.

In some embodiments, the main controller <NUM> monitors the intensity of the interference signal (interference intensity, interference sensitivity) acquired sequentially while acquiring the interference signal by performing the repetitive OCT measurements described above. In addition, while performing this monitoring process, the OCT focusing lens <NUM> is moved to find the position of the OCT focusing lens <NUM> in which the interference intensity is maximized. With the fine focus adjustment thus performed, the OCT focusing lens <NUM> can be guided to the position where the interference intensity is optimized.

Further, the data processor <NUM> determines the polarization state of at least one of the measurement light LS and the reference light LR by analyzing the detection result of the interference light obtained by the OCT measurement. For example, the main controller <NUM> performs repetitive OCT measurements while controlling at least one of the polarization controllers <NUM> and <NUM> according to a predetermined algorithm. In some embodiments, the main controller <NUM> controls the attenuator <NUM> to change an attenuation of the reference light LR. The data processor <NUM> analyzes detection results of interference light LC repeatedly acquired by the OCT measurements to calculate predetermined evaluation values relating to image quality of OCT images. The data processor <NUM> determines whether the calculated evaluation value is equal to or less than a threshold. The threshold is set in advance. Polarization adjustment is continued until the evaluation value calculated becomes equal to or less than the threshold. That is, when the evaluation value is equal to or less than the threshold, it is determined that the polarization state of the measurement light LS is appropriate. And the polarization adjustment is continued until it is determined that the polarization state of the measurement light LS is appropriate.

In some embodiments, the main controller <NUM> can monitor the interference intensity also in the polarization adjustment.

Further, the data processor <NUM> performs predetermined analysis processing on the detection result of the interference light acquired by the OCT measurement or the OCT image formed based on the detection result. Examples of the predetermined analysis processing include specifying (identification) of a predetermined site (tissue, lesion) of the subject's eye E; calculation of a distance, area, angle, ratio, or density between designated sites (distance between layers, interlayer distance); calculation by a designated formula; specifying of the shape of a predetermined site; calculation of these statistics; calculation of distribution of the measured value or the statistics; image processing based on these analysis processing results, and the like. Examples of the predetermined tissue include a blood vessel, an optic disc, a fovea, a macula, and the like. Examples of the predetermined lesion include a leukoma, a hemorrhage, and the like.

The data processor <NUM> performs coordinate transformation on the pixel positions in the OCT image or the scan positions in the scan data so that the site in the eye in the acquired OCT image (or the scan data) is drawn in actual shape. Further, the data processor <NUM> can obtain a distance between predetermined sites in the eye using the OCT image after coordinate transformation or the scan data after coordinate transformation.

<FIG> and <FIG> show diagrams of comparative examples of the embodiments. <FIG> schematically shows the path of the measurement light incident on the subject's eye E. <FIG> shows an example of the tomographic image obtained by scanning with the measurement light incident on the subject's eye E through the path shown in <FIG>.

The measurement light deflected by the optical scanner <NUM>, for example, is incident on the pupil of the subject's eye E, which is a scan center position, at various incident angles, as shown in <FIG>. The measurement light incident on the subject's eye E is projected toward each part in the eye around the scan center position Cs set at the center of the pupil, for example.

An A-scan image is formed from the interference data obtained using the measurement light LS1 in <FIG>, an A-scan image is formed from the interference data obtained using the measurement light LS2, and an A-scan image is formed from the interference data obtained using the measurement light LS3. The tomographic image IMG0 of the fundus shown in <FIG> is formed by arranging the plurality of A-scan images thus formed.

In this way, the A scan directions vary within the scan angle range centered on the scan center position Cs, and the shape of the site is deformed in the tomographic images in which the obtained plurality of A scan images are arranged in the horizontal direction. The wider the angle of view is, the greater the difference from the actual shape becomes.

Further, morphology information representing the morphology of the subject's eye E can be obtained from the positions of arbitrary pixels in the tomographic image. Examples of the morphology information include an intraocular distance (including a distance between layer regions), an area of region, a volume of region, a perimeter of region, a direction of site with reference to a reference position, an angle of site with reference to a reference direction, and a curvature radius of site.

For example, the intraocular distance as the morphology information can be obtained by measuring a distance between arbitrary two points in the tomographic image. In this case, the distance between the two points can be specified using the number of pixels in the tomographic image, and can be measured by multiplying the specified number of pixels by the pixel size specific to the apparatus. At this time, the same pixel size is adopted for all pixels in the tomographic image. However, as described above, the scan directions are different with the scan center position Cs as the center. Thereby, the pixel size in the horizontal direction of the tomographic image differs depending on the depth position in the scan direction. For example, in case that the depth range is <NUM> millimeters, when the same pixel size is adopted for all pixels in the tomographic image, there is a difference of about <NUM>% in the scan length of the B-scan between the upper portion and the lower portion of the tomographic image, and when the depth range is <NUM> millimeters, there is a difference of about <NUM>%.

Therefore, the data processor <NUM> according to the embodiments performs coordinate transformation on the pixel positions in the acquired OCT image or the scan positions in the scan data. Hereinafter, the intraocular distance will be described as an example of the morphology information representing the morphology of the subject's eye E.

Such the data processor <NUM> includes a position specifying unit <NUM>, a position transforming unit <NUM>, an interpolator <NUM>, and an intraocular distance calculator <NUM>.

The position specifying unit <NUM> is configured to specify a transformation position along a traveling direction of the measurement light passing through the scan center position Cs, the transformation position corresponding to a pixel position in the acquired OCT image (or the scan position in the scan data). In some embodiments, the position specifying unit <NUM> uses the eyeball parameter 212A for performing processing for specifying the transformation position.

<FIG> shows a diagram describing the operation of the position specifying unit <NUM> according to the embodiments. In <FIG>, parts similarly configured to those in <FIG> are denoted by the same reference numerals, and the description thereof is omitted unless it is necessary.

Here, the scan angle is φ, the scan radius is r, the depth range in which OCT measurement can be performed d, the length of the tomographic image in the depth direction is h, and the lateral length of the tomographic image is w. The scan angle φ corresponds to the deflection angle of the measurement light LS around the scan center position Cs. The scan radius r corresponds to the distance from the scan center position Cs to a zero optical path length position where the measurement optical path length and the reference optical path length are substantially equal. The depth range d is a value (known) specific to the apparatus, the value being uniquely determined by the optical design of the apparatus.

The position specifying unit <NUM> specifies the transformation position (X, Z) in a second coordinate system from the pixel position (x, z) in a first coordinate system. The first coordinate system is a coordinate system having the origin at the upper left coordinate position in the OCT image (B-scan image). The first coordinate system is defined by an x coordinate axis having the B-scan direction as the x direction and a z coordinate axis, which is orthogonal to the x coordinate axis, having the A-scan direction as the z direction. The pixel position (x, z) in the OCT image is defined in the first coordinate system. The second coordinate system is defined a Z coordinate axis (for example, second axis) and a X coordinate axis (for example, first axis). The Z coordinate axis has the traveling direction of the measurement light LS having the scan angle of <NUM> degrees with respect to the measurement optical axis passing through a predetermined site (for example, fovea) in the fundus Ef, as the Z direction. The X coordinate axis has the B-scan direction orthogonal to the Z coordinate axis at the predetermined site, as the X direction. In the second coordinate system, a predetermined Z position is set as the origin of the Z coordinate axis so that the position of the scan radius r becomes the deepest portion in the measurement optical axis passing through the predetermined site (for example, the fovea). Further, a predetermined X position in the measurement optical axis passing through the predetermined site (for example, the fovea) is set as the origin of the X coordinate axis so as to have a predetermined depth direction length d as described below. The transformation position (X, Z) is defined in the second coordinate system. The transformation position (X, Z) corresponds to the pixel position (x, z), and is a position along the traveling direction of the measurement light LS passing through the scan center position Cs (A-scan direction).

For the OCT image, the position specifying unit <NUM> specifies the transformation position (X, Z) based on the scan radius r of the A-scan direction, the scan angle φ, the depth range d in which the OCT measurement can be performed, and the pixel position (x, z). The position specifying unit <NUM> can specify at least one of the X component of the transformation position (component of the first axis direction) and the Z component of the transformation position (component of the second axis direction).

For the OCT image (tomographic image) in which the number of A-scan lines is N (N is a natural number), the transformation position (X, Z), which corresponds to the pixel position (x, z) in the n-th (n is a natural number) A-scan line, is specified as shown in Equations (<NUM>) and (<NUM>). [Equation <NUM>] <MAT>
[Equation <NUM>] <MAT>.

Here, the length h in the depth direction of the OCT image, the length w in the horizontal direction of the OCT image, and the x component of the pixel position are expressed by Equations (<NUM>) to (<NUM>). [Equation <NUM>] <MAT>
[Equation <NUM>] <MAT>
[Equation <NUM>] <MAT>.

In Equations (<NUM>) and (<NUM>), the x coordinate of the pixel position is expressed by Equation (<NUM>). Thus, the position specifying unit <NUM> can specify the transformation position (X, Z) from the pixel position (x, z), based on the scan radius r, the scan angle φ, and the depth range d.

In some embodiments, for the scan data, the position specifying unit <NUM> can specify the transformation position (X, Z) based on the scan radius r in the A-scan direction, the scan angle φ, the depth range d in which the OCT measurement can be performed, and the scan position, in the same way as above.

In some embodiments, the scan radius r is specified by analyzing the detection result of the interference light LC obtained using the OCT optical system <NUM>. This allows to specify the transformation position (X, Z) that more accurately reflects the eyeball optical characteristics of subject's eye E.

In some embodiments, the position specifying unit <NUM> specifies the scan angle φ by performing ray trace processing on the measurement light LS based on the corneal shape information of the subject's eye E. Examples of the corneal shape information include a corneal curvature radius (curvature radius of an anterior surface of cornea, curvature radius of a posterior surface of cornea) and corneal thickness. This allows to specify the transformation position (X, Z) that more accurately reflects the eyeball optical characteristics of subject's eye E.

The position transforming unit <NUM> transforms the pixel position (x, z) in the OCT image into the transformation position (X, Z) specified by the position specifying unit <NUM>. In some embodiments, for each of all pixel positions in the OCT image, the position specifying unit <NUM> specifies the transformation position and the position transforming unit <NUM> transforms the pixel position into the transformation position.

This allows to arrange the A-scan images, which are acquired by performing A-scan, in the A-scan direction as shown in <FIG>. Therefore, even if the angle of view is wide as in the tomographic image IMG1 shown in <FIG>, the tomographic image in which the shape of the predetermined site is similar to the actual shape can be obtained.

The interpolator <NUM> interpolates pixels between the transformation positions. For example, intervals between the A-scan images adjacent to each other in which the pixel positions have been transformed into the transformation position varies depending on the distance from the scan center position Cs. The interpolator <NUM> interpolates pixel(s) between the A-scan images using a pixel in the A-scan images adjacent to each other according to the depth position in the A-scan image. As interpolation processing on pixels performed by the interpolator <NUM>, a known method such as a nearest neighbor method, a bilinear interpolation method, or a bicubic interpolation method can be adopted. In some embodiments, the interpolator <NUM> interpolates pixels between the A-scan images adjacent to each other according to the distance from the scan center position Cs. For example, the interpolator <NUM> interpolates pixels between the A-scan images adjacent to each other by changing interpolation processing method according to the distance from the scan center position Cs.

In some embodiments, for the scan position in the scan data, the scan data is interpolated, in the same way as above.

The intraocular distance calculator <NUM> calculates an intraocular distance of the subject's eye E based on the image OCT image in which the pixel position has been transformed into the transformation position by the position transforming unit <NUM>.

The intraocular distance calculator <NUM> obtains the intraocular distance between predetermined sites in the subject's eye E based on the OCT image transformed by the position transforming unit <NUM>. For example, the intraocular distance calculator <NUM> specifies the predetermined sites in the eye by analyzing the transformed OCT image, and obtains the intraocular distance described above based on the distance between the specified sites. The distance between the two points can be specified using the number of pixels in the tomographic image, and can be measured by multiplying the specified number of pixels by the pixel size specific to the apparatus. At this time, the same pixel size is adopted for all pixels in the tomographic image.

Examples of the intraocular distance between the predetermined sites include a distance between designated sites (tissue, layer region), an axial length, a distance from a scan center position of the measurement light, which is set at the center of the pupil, or the like, to a retina. In case that the axial length is obtained as the intraocular distance, the intraocular distance calculator <NUM> obtains the axial length based on a distance from a site corresponding to a corneal apex to a site corresponding to the retina.

In some embodiments, the intraocular distance calculator <NUM> calculates an intraocular distance of the subject's eye, in the same way as above, based on the scan data in which the scan position has been transformed into the transformation position by the position transforming unit <NUM>.

The data processor <NUM> that functions as above includes, for example, a processor described above, a RAM, a ROM, a hard disk drive, a circuit board, and the like. In a storage device such as the hard disk drive, a computer program for causing the processor to execute the functions described above is stored in advance.

The user interface <NUM> includes the display unit 240A and an operation unit 240B. The display unit 240A includes the aforementioned display device of the arithmetic control unit <NUM> and the display apparatus <NUM>. The operation unit 240B includes the aforementioned operation device of the arithmetic control unit <NUM>. The operation unit 240B may include various types of buttons and keys provided on the case of the ophthalmologic apparatus <NUM> or the outside. For example, when the fundus camera unit <NUM> has a case similar to that of the conventional fundus camera, the operation unit 240B may include a joy stick, an operation panel, and the like provided to the case. Besides, the display unit 240A may include various types of display devices such as a touch panel and the like arranged on the case of the fundus camera unit <NUM>.

Note that the display unit 240A and the operation unit 240B need not necessarily be formed as separate devices. For example, a device like a touch panel, which has a display function integrated with an operation function, can be used. In such cases, the operation unit 240B includes the touch panel and a computer program. The content of operation performed on the operation unit 240B is fed to the controller <NUM> in the morphology of an electrical signal. Moreover, operations and inputs of information may be performed by using a graphical user interface (GUI) displayed on the display unit 240A and the operation unit 240B.

The optical system in the path from the interference optical system included in the OCT unit <NUM> to the objective lens <NUM>, or these optical systems and the image forming unit <NUM> is an example of the "acquisition unit" according to the embodiments that acquires a plurality of A-scan images or a plurality of A-scan data using OCT.

The operation of the ophthalmologic apparatus <NUM> according to the embodiments will be described.

<FIG> and <FIG> show an example of the operation of the ophthalmologic apparatus <NUM> according to the embodiments. <FIG> and <FIG> shows flowcharts of the example of the operation of the ophthalmologic apparatus <NUM> according to the embodiments. <FIG> shows a flowchart of an example of the operation of step S7 in <FIG>. The storage unit <NUM> stores computer programs for realizing the processing shown in <FIG> and <FIG>. The main controller <NUM> operates according to the computer programs, and thereby the main controller <NUM> performs the processing shown in <FIG> and <FIG>.

The main controller <NUM> performs alignment.

That is, the main controller <NUM> controls the alignment optical system <NUM> to project the alignment indicator onto the subject's eye E. At this time, a fixation target generated by the LCD <NUM> is also projected onto the subject's eye E. The main controller <NUM> controls the movement mechanism <NUM> based on the movement amount of the optical system to relatively to move the optical system with respect to the subject's eye E by the movement amount. The movement amount is specified based on the receiving light image obtained using the image sensor <NUM>, for example. The main controller <NUM> repeatedly executes this processing.

In some embodiments, the alignment rough adjustment and the alignment fine adjustment are performed after the alignment in step S1 is completed.

The main controller <NUM> controls the LCD <NUM> to display the fixation target for OCT measurement at a predetermined position on the LCD <NUM>. The main controller <NUM> can display the fixation target at a display position on the LCD <NUM> corresponding to a position of an optical axis of the optical axis on the fundus Ef.

Subsequently, the main controller <NUM> controls the OCT unit <NUM> to perform OCT provisional measurement, and to acquire a tomographic image for adjustment for adjusting the reference position of the measurement range in the depth direction. Specifically, the main controller <NUM> controls the optical scanner <NUM> to deflect the measurement light LS generated based on the light L0 emitted from the light source unit <NUM> and to scan a predetermined site (for example, fundus) of the subject's eye E with the deflected measurement light LS. The detection result of the interference light obtained by scanning with the measurement light LS is sent to the image forming unit <NUM> after being sampled in synchronization with the clock KC. The image forming unit <NUM> forms the tomographic image (OCT image) of the subject's eye E from the obtained interference signal.

Subsequently, the main controller <NUM> adjusts the reference position of the measurement range in the depth direction (z direction).

For example, the main controller <NUM> controls the data processor <NUM> to specify a predetermined site (for example, sclera) in the tomographic image obtained in step S2, and sets a position separated by a predetermined distance in the depth direction from the specified position of the predetermined site as the reference position of the measurement range. The main controller <NUM> controls at least one of the optical path length changing units <NUM> and <NUM> according to the reference position. Alternatively, a predetermined position determined in advance so that the optical path lengths of the measurement light LS and the reference light LR substantially coincide may be set as the reference position of the measurement range.

Next, the main controller <NUM> perform control of adjusting focusing and of adjusting polarization.

For example, the main controller <NUM> controls the OCT unit <NUM> to perform OCT measurement, after controlling the focusing driver 43A to move the OCT focusing lens <NUM> by a predetermined distance. The main controller <NUM> controls the data processor <NUM> to determine the focus state of the measurement light LS based on the detection result of the interference light acquired by the OCT measurement, as described above. When it is determined that the focus state is not appropriate based on the determination result of the data processor <NUM>, the main controller <NUM> controls the focusing driver 43A again and repeats this until it is determined that the focus state of the measurement light LS is appropriate.

Further, for example, the main controller <NUM> controls the OCT unit <NUM> to perform OCT measurement after controlling at least one of the polarization controllers <NUM> and <NUM> to change the polarization state of at least one of the light L0 and the measurement light LS by a predetermined amount. And then, the main controller <NUM> controls the image forming unit <NUM> to form the OCT image on the basis of the acquired detection result of the interference light. The main controller <NUM> controls the data processor <NUM> to determine the image quality of the OCT image acquired by the OCT measurement, as described above. When it is determined that the polarization state is not appropriate based on the determination result of the data processor <NUM>, the main controller <NUM> controls the polarization controllers <NUM> and <NUM> again and repeats this until it is determined that the polarization state of the measurement light LS is appropriate.

Subsequently, the main controller <NUM> controls the OCT unit <NUM> to perform OCT measurement. The detection result of the interference light acquired by the OCT measurement is sampled by the DAQ <NUM> and is stored as the interference signal in the storage unit <NUM> or the like.

Next, the main controller <NUM> controls the image forming unit <NUM> to form the data set group of the A-scan image data of the subject's eye E based on the interference signal acquired in step S5. The image forming unit <NUM> forms the tomographic image as shown in <FIG>, by arranging the formed A-scan images in the B-scan direction.

The main controller <NUM> corrects the tomographic image, which is formed in step S6, as described above using the eyeball parameter 212A stored in the storage unit <NUM>. This allows to acquire the tomographic image in which the A-scan images are arranged in the A-scan direction.

For example, the main controller <NUM> controls the display unit 240A to display the newly generated tomographic image (for example, the tomographic image IMG1 shown in <FIG>) and the tomographic image before correction (for example, the tomographic image IMG0 shown in <FIG>) on the same screen SCR of the display unit 240A (<FIG>). This allows to easily compare with the conventional tomographic images in which many imaging findings are accumulated, even if different morphologies are drawn in the conventional tomographic image by displaying the corrected tomographic image.

In some embodiments, the tomographic image before correction is displayed according to the measurement site. For example, when displaying a tomographic image or the like in the vicinity of the macula where the change in morphology is small before and after correction, the tomographic image before correction may be intentionally displayed.

This terminates the operation of the ophthalmologic apparatus <NUM> (END).

In step S7 in <FIG>, processing as shown in <FIG> is performed.

In step S7, the main controller <NUM> controls the position specifying unit <NUM> to specify the transformation position corresponding to the pixel position in the tomographic image formed in step S6. The position specifying unit <NUM> specifies the transformation position corresponding to the pixel position in the tomographic image, as described above.

Subsequently, the main controller <NUM> controls the position transforming unit <NUM> to transform the pixel position in the tomographic image into the transformation position calculated in step S11.

The main controller <NUM> determine whether or not the next pixel position should be transformed.

When it is determined that the next pixel position should be transformed (S13: Y), the operation of the ophthalmologic apparatus proceeds to step S11. When it is determined that the next pixel position should not be transformed (S13: N), the operation of the ophthalmologic apparatus proceeds to step S14.

Through steps S11 to S13, for each pixel position of the tomographic image, specifying the transformation position and transforming to the specified transformation position are performed.

When it is determined that the next pixel position should not be transformed in step S13 (S13: N), the main controller <NUM> controls the interpolator <NUM> to interpolate the pixels between the A-scan images adjacent to each other, A-scan images having been transformed into the transformation positions in step S12.

This terminates the processing of step S7 in <FIG> (END).

In the embodiments described above, the case has been described in which the two-dimensional OCT image (or the two-dimensional scan data) is corrected. However, the configuration according to the embodiments is not limited thereto. The ophthalmologic apparatus according to the embodiments can correct three-dimensional OCT data (or the three-dimensional scan data), as in the embodiments described above. Hereinafter, an ophthalmologic apparatus according to a modification example of the embodiments will be described focusing on differences from the embodiments.

The configuration of the ophthalmologic apparatus according to the modification example of the embodiments is similar to the configuration of the ophthalmologic apparatus <NUM> according to the embodiments. Therefore, the description thereof will be omitted.

The data processor according to the present modification example performs processing for specifying the transformation position in the three-dimensional space, or the like.

The position specifying unit <NUM> according to the present modification example specifies the transformation position along the traveling direction of the measurement light passing through the scan center position Cs, the transformation position corresponding to the pixel position in the acquired OCT image (or the scan position in the scan data). In some embodiments, the position specifying unit <NUM> specifies the transformation position using the eyeball parameter 212A.

<FIG> shows a diagram describing the operation of the position specifying unit <NUM> according to the present modification example. In <FIG>, parts similarly configured to those in <FIG> are denoted by the same reference numerals, and the description thereof is omitted unless it is necessary.

In <FIG>, a Y plane is defined in addition to the X plane and the Z plane in <FIG>. In addition to the parameters shown in <FIG>, the central angle in the C-scan direction is θ, and the length in the C-scan direction is lc.

The position specifying unit <NUM> specifies the transformation position (X, Y, Z) in a fourth coordinate system from the pixel position (x, y, z) in a third coordinate system. The third coordinate system is a coordinate system having the origin at the upper left coordinate position in the three-dimensional OCT image. The third coordinate system is defined by the x coordinate axis having the B-scan direction as the x direction, a y coordinate axis, which is orthogonal to the x coordinate axis, having the C-scan direction as the y direction, and the z coordinate axis, which is orthogonal to both of the x coordinate axis and the y coordinate axis, having the A-scan direction as the z direction. The pixel position (x, y, z) in the OCT image is defined in the third coordinate system. The fourth coordinate system is defined the Z coordinate axis, the X coordinate axis, and a Y coordinate axis. The Z coordinate axis has the traveling direction of the measurement light LS having the scan angle of <NUM> degrees with respect to the measurement optical axis passing through a predetermined site (for example, fovea) in the fundus Ef, as the Z direction. The X coordinate axis has the B-scan direction orthogonal to the Z coordinate axis at the predetermined site, as the X direction. The Y coordinate axis has the C-scan direction orthogonal to the Z coordinate axis at the predetermined site, as the Y direction. In the fourth coordinate system, a predetermined Z position is set as the origin of the Z coordinate axis so that the position of the scan radius r becomes the deepest portion in the measurement optical axis passing through the predetermined site (for example, the fovea). Further, a predetermined X position and Y position in the measurement optical axis passing through the predetermined site (for example, the fovea) are set as the origin of the X coordinate axis and the Y coordinate axis so as to have a predetermined depth direction length d as described below. The transformation position (X, Y, Z) is defined in the fourth coordinate system. The transformation position (X, Y, Z) corresponds to the pixel position (x, y, z), and is a position along the traveling direction of the measurement light LS passing through the scan center position Cs (A-scan direction).

The position specifying unit <NUM> can specify at least one of the X component, the Y component, and the Z component of the transformation position.

For the OCT image (tomographic image) in which the number of A-scan lines is N (N is a natural number) and the number of B-scan lines is M (M is a natural number), the transformation position (X, Y, Z), which corresponds to the pixel position (x, y, z) in the n-th (n is a natural number) A-scan line of the m-th (m is a natural number) B-scan line, is specified as shown in Equations (<NUM>) to (<NUM>). [Equation <NUM>] <MAT>
[Equation <NUM>] <MAT>
[Equation <NUM>] <MAT>.

Here, the x component and the y component of the pixel position are expressed by Equations (<NUM>) to (<NUM>) from the length h in the depth direction, the length w in the B-scan direction, and the length 1c in the C-scan direction of the three-dimensional OCT image. [Equation <NUM>] <MAT>
[Equation <NUM>] <MAT>
[Equation <NUM>] <MAT>
[Equation <NUM>] <MAT>
[Equation <NUM>] <MAT>.

In Equations (<NUM>) to (<NUM>), the x coordinate and the y coordinate of the pixel position are expressed by Equations (<NUM>) and Equation (<NUM>). Thus, the position specifying unit <NUM> can specify the transformation position (X, Y, Z) from the pixel position (x, y, z), based on the scan radius r, the scan angle φ, and the depth range d.

In some embodiments, for the scan data, the position specifying unit <NUM> can specify the transformation position (X, Y, Z), in the same way as above.

The position transforming unit <NUM> according to the present modification example transforms the pixel position (x, y, z) in the OCT image into the transformation position (X, Y, Z) specified by the position specifying unit <NUM>. In some embodiments, for each of all pixel positions in the OCT image, the position specifying unit <NUM> specifies the transformation position and the position transforming unit <NUM> transforms the pixel position into the transformation position.

In the embodiments described above, the case where the tomographic image is corrected in the ophthalmologic apparatus including the OCT unit <NUM>, and the like has been described. However, the configuration according to the embodiments is not limited thereto. For example, the ophthalmologic information processing apparatus, which realizes the function of the data processor <NUM> shown in <FIG>, may correct the tomographic image for the acquired OCT image (or the scan data), as described above. In this case, the OCT image (or the scan data) is acquired by an external OCT apparatus (ophthalmologic apparatus).

Claim 1:
An ophthalmologic information processing apparatus for correcting an image of a subject's eye (E) formed by arranging a plurality of A-scan images acquired by scanning inside the subject's eye (E) with measurement light (LS) deflected around a scan center position (C), the ophthalmologic information processing apparatus being characterising by comprising:
a specifying unit (<NUM>) configured to specify for each pixel position, defined in a first coordinate system, in the image a transformation position, defined in a second coordinate system, along a traveling direction of the measurement light (LS) passing through the scan center position (C), the transformation position corresponding to a pixel position in the image;
a transforming unit (<NUM>) configured to transform the pixel positions into the transformation positions specified by the specifying unit (<NUM>); and
an interpolator (<NUM>) configured to interpolate pixels between the transformation positions in the second coordinate system by interpolating these pixels between the A-scan images in which the pixel positions have been transformed into the transformation positions using pixels in the A-scan images adjacent to each other, according to a depth position in the A-scan image.