JP2016198280A - Ophthalmologic apparatus and control method of the same - Google Patents

Abstract

PROBLEM TO BE SOLVED: To provide an ophthalmologic apparatus capable of rapidly and accurately obtaining a position and a flow rate of a target blood vessel by a simple method without any new structure involved.SOLUTION: An ophthalmologic apparatus comprises eyeground tomographic image acquisition means for splitting light from a source into reference light and measurement light, emitting the measurement light to an eyeground, multiplexing the reflected measurement light and the reference light to obtain interference light, and acquiring an eyeground tomographic image based on the interference light. The ophthalmologic apparatus also comprises scanning control means capable of scanning control of the measurement light to acquire a two-dimensional tomographic image (B-scan image) extending in the eyeground depth direction at two or more different scanning velocities.SELECTED DRAWING: Figure 3

Description

The present invention relates to an ophthalmologic apparatus mainly used in ophthalmology for acquiring a tomographic image using optical coherence tomography (OCT) and a control method thereof.

Optical coherence tomography (OCT) is a method widely used as a means for acquiring high-resolution tomographic images of living tissue in ophthalmology because it can be measured non-invasively and non-contactly. In OCT, an image in a one-dimensional depth direction (in the eyeball, in the direction of the eye axis) is usually called an A-scan image, a two-dimensional image is called a B-scan image, and a three-dimensional image is called a C-scan image. An image is also described as an A-scan (image) image, a two-dimensional image as a B-scan (image) image, and a three-dimensional image as a C-scan (image) image.

In optical coherence tomography (OCT), a time domain method called a time domain method, a time domain OCT that acquires a tomographic image while mechanically changing the optical path length of reference light by moving a mirror, and a spectroscope called a Fourier domain method are used. There are spectral domain OCT that uses spectral information to detect tomographic images, and optical frequency swept OCT that detects spectral interference signals using a wavelength scanning light source to acquire tomographic images.

In the Fourier domain OCT, the time variation of the phase obtained by the Fourier transform of the acquired spectrum interference signal corresponds to the moving speed of the test object as a Doppler signal. It is used as a means for obtaining. Since the Fourier domain OCT for acquiring the Doppler signal is also referred to as Doppler OCT, it will be described as Doppler OCT below.

For example, Doppler OCT acquires an OCT image of a predetermined part twice or more at a predetermined time Δt in order to measure a blood flow velocity in a blood vessel in a predetermined part of the test object, and acquires two or more acquired OCTs. A method has been employed in which a blood flow velocity is calculated by obtaining a phase difference Δφ at a predetermined time Δt from two adjacent OCT images having consecutive images.

However, in this conventional method, in order to measure the blood flow velocity, it is necessary to acquire the OCT image by performing OCT imaging of the same predetermined portion a plurality of times as described above. There was a problem.

Patent Document 1 discloses two tomographic images (B-scan) acquired by dividing the measurement light applied to the fundus of the subject's eye into two independent beams and scanning (B-scan) with the same fundus position shifted by Δt. A method of obtaining a blood flow velocity by obtaining a phase change Δφ at Δt from pixel data at an arbitrary same position for a B scan image) is disclosed.

JP 2013-7601 A

The method described in Patent Document 1 using two independent beams can shorten the measurement time as compared with the method of scanning the same position a plurality of times that has been conventionally performed. However, the method of generating two independent beams and scanning the same position is complicated in configuration and control method, and is difficult to implement. Further, the tomographic images acquired by the two beams are not overlapped in the depth direction. Therefore, there is a problem that the blood flow velocity at the deep position of the fundus cannot be obtained because a certain difference in optical path length is required between the two beams.

The present invention has been made to solve the above-described problems, and does not require a new configuration, and can quickly and accurately determine the target blood vessel position and blood flow velocity by a simple method. It is an object of the present invention to provide an ophthalmic apparatus that can be obtained well.

In order to achieve the above object, an ophthalmologic apparatus according to claim 1 divides light from a light source into reference light and measurement light, irradiates the branched measurement light to the fundus, and combines the reflected measurement light and reference light. In order to acquire a two-dimensional tomographic image (B-scan image) extending in the depth direction of the fundus in an ophthalmologic apparatus having a fundus tomographic image acquisition means for acquiring a tomographic image of the fundus of the eye to be examined from the interference light obtained by waves. It is characterized by comprising scanning control means capable of performing scanning control of the measuring light at at least two different speeds.

In the case of Doppler OCT that quantitatively measures the state of blood flow of a target blood vessel in the fundus, adjacent B scans are obtained by repeatedly acquiring B scan images at the same position including the target blood vessel. By detecting the position of the blood vessel from the phase difference of the pixel data at the same position of the image and scanning a narrow region centered on the detected position of the blood vessel at a low scanning speed, the A scan images constituting the B scan image By reducing the distance, it is possible to obtain the (vertical) blood flow velocity at the detected position of the target blood vessel from the phase difference between adjacent A-scan images. By adopting such a scanning control means, the position of the target blood vessel and the (vertical) blood flow velocity can be acquired by a simple method without requiring a new configuration.

In order to achieve the above object, an ophthalmologic apparatus according to claim 2 is the ophthalmologic apparatus according to claim 1, which illuminates the fundus of the eye to be examined and emits reflected light or scattered light from the fundus. Fundus front image acquisition means for receiving a front image of the fundus of the eye to be received by receiving a three-dimensional image element and a position for acquiring an arbitrary B-scan image from the acquired front image of the fundus of the eye to be examined B scan designating means is provided.

Although it is possible to acquire a fundus front image from an OCT C-scan image, a fundus front image acquisition unit such as a fundus camera or SLO can be provided to acquire a fundus front image in real time. By providing a means for displaying a front image on a monitor or the like and designating a position including a target blood vessel, the blood flow velocity of the target blood vessel can be easily measured. Furthermore, since the front image of the fundus can be acquired in real time, it can also be employed for tracking control when photographing OCT.

In order to achieve the above object, an ophthalmologic apparatus according to claim 3 is the ophthalmologic apparatus according to claim 1 or 2, wherein at least of at least two different scanning speeds controlled by the scanning control means. The two scanning speeds are different by 10 times or more.

For example, by setting the different scanning speeds to 300 A scan / B scan (high speed scan) and 8,000 A scan / B scan (low speed scan), the blood of the target blood vessel is avoided while making the entire measurement time longer. The flow velocity can be measured accurately.

In order to achieve the above object, an ophthalmologic apparatus according to claim 4 is the ophthalmologic apparatus according to claim 3, wherein a B scan image acquired at a slow scanning speed of at least two different scanning speeds is used. On the other hand, a first phase difference acquisition means for acquiring a phase difference between adjacent A-scan images (substantially the same position) is provided.

By reducing the scanning speed so that the distance between adjacent A-scan images is substantially the same position, the (vertical) blood flow velocity at the detected position of the target blood vessel is determined from the phase difference between the adjacent A-scan images. It is possible to ask.

In order to achieve the above object, an ophthalmic apparatus according to claim 5 is the ophthalmic apparatus according to claim 4, wherein a B-scan image acquired at a fast scanning speed among at least two different scanning speeds is used. On the other hand, there is provided a second phase difference acquisition means for acquiring a phase difference between the same positions of adjacent B-scan images acquired successively.

By scanning at a relatively high speed, a B-scan image can be acquired in a short time, and the target blood vessel position can be easily and quickly detected from the phase difference between the same positions of adjacent B-scan images. .

In order to achieve the above object, an ophthalmologic apparatus according to claim 6 is the ophthalmologic apparatus according to any one of claims 2 to 5, wherein the B-scan designation means includes at least the same blood vessel. It is characterized by designating two different positions.

At least two different positions including the target blood vessel are designated, and at each position, the target blood vessel position (x0, z0), (x1, z1) and the distance Δd between the two positions are the target. The blood flow gradient is obtained, and the absolute blood flow rate can be calculated from the obtained gradient and the (vertical) blood flow velocity at each position.

In order to achieve the above object, an ophthalmologic apparatus according to a seventh aspect is the ophthalmologic apparatus according to any one of the second to sixth aspects, wherein a scanning direction of the scanning control means is a target blood vessel. The direction is substantially perpendicular to the direction.

By scanning so that the scanning direction intersects the target blood vessel substantially perpendicularly and obtaining a B-scan image, it is possible to calculate an absolute blood flow with high accuracy.

As described above, according to the present invention, by adopting the scanning control means as described above, the absolute blood flow rate of the target blood vessel can be acquired by a simple method without requiring a new configuration. is there.

It is the figure which showed the structure of the optical system of one Example of the ophthalmologic apparatus which concerns on this invention. It is the figure which showed the structure of the whole apparatus of one Example of the ophthalmic apparatus which concerns on this invention. It is the figure which showed the operation procedure (flowchart) of one Example of the ophthalmologic apparatus which concerns on this invention. It is a figure explaining the flow until acquisition of the three-dimensional tomogram by OCT. It is the figure which showed an example of the (a) B scan image and emphasis image (created from three-dimensional OCT data) of the fundus by OCT displayed on the monitor of the ophthalmologic apparatus according to the present embodiment. It is the figure which showed an example of the front image of the fundus by SLO displayed on the monitor of the ophthalmologic apparatus concerning a present Example.

Hereinafter, an ophthalmologic apparatus according to an embodiment of the present invention will be described with reference to the drawings.
[One Embodiment]
FIG. 1 is a diagram illustrating details of an optical system of an ophthalmologic apparatus 1 according to the present invention. FIG. 2 is a diagram showing a configuration of the entire apparatus of an embodiment of the ophthalmologic apparatus 1 according to the present invention.

The ophthalmic apparatus 1 includes the following two optical systems. Irradiating the fundus of the eye E using an interference optical system (hereinafter referred to as OCT optical system) 100 for non-invasively acquiring a tomographic image of the fundus of the eye E to be examined non-invasively using interference light technology. A scanning laser ophthalmoscope (SLO) optical system (hereinafter referred to as an SLO optical system) 200 that acquires a fundus SLO image for observation.

The configuration of each optical system will be described below.
(OCT optical system 100)
The OCT optical system 100 includes a light source 101 to an ADC 116 that performs A / D conversion of interference light. In this embodiment, SS-OCT using a wavelength sweep type light source is adopted as the light source 101, which is one of Fourier domain type OCT. SS-OCT is superior in that it can acquire interference signals (tomographic image data) at a higher speed than other OCT methods because of its measurement principle. The OCT optical system 100 is not limited to the SS-OCT of this embodiment, and may be a spectral domain OCT (SD-OCT) which is another Fourier domain type OCT.

The light output from the light source 101 passes through the fiber and is split into measurement light input to the collimator lens 103 and reference light input to the collimator lens 110 by the fiber coupler 102. The measurement light input to the collimator lens 103 is applied to the fundus of the eye E through the focus lens 104, the galvanometer mirror 105, the lens 106, the dichroic mirror 107, and the objective lens 109. Then, the measurement light reflected from the fundus of the eye E passes through the objective lens 109, the dichroic mirror 107, the lens 106, the galvanometer mirror 105, the focus lens 104, the collimator lens 103, and the fiber coupler 102, contrary to the time of irradiation. , And input to one input section of the fiber coupler 114.

The reference light branched to the fiber coupler 102 and input to the collimator lens 110 is reflected by the prism 112, passes through the collimator lens 111, and is input to the other input portion of the fiber coupler 114.

Measurement light and reference light input to the fiber coupler 114 are combined in the fiber coupler 114 and input to the balance detector 115 as interference light and converted into an electrical signal (interference signal). The two interference lights output from the fiber coupler 114 are interference lights having a phase difference of 180 °, and the two interference lights are input to the balance detector 115 and differentially amplified. Here, when the influence of noise components such as common noise is low, a simple one-input detector or the like may be employed.

The interference signal output from the balance detector 115 is sampled as a digital signal by the ADC 116, input to the calculation unit 500 including a CPU and a memory, and Fourier-transformed to obtain A scan data that is a tomographic signal in the depth direction. It is stored in the memory in the calculation unit 500.

The prism 112 is moved by the control unit 113 on the optical axis so that the reference optical path length can be changed and adjusted. Usually, before the OCT imaging, the control unit 113 moves so that the reference optical path length and the measurement optical path length are the same, and is fixed during the measurement.

The galvanometer mirror 105 scans horizontally (X-axis direction) and vertical (Y-axis direction) with respect to the eye E, and a control signal is input from the calculation unit 500. A three-dimensional tomographic image of the fundus of the eye E can be acquired by scanning the galvanometer mirror 105 in the X-axis direction and the Y-axis direction.

In this embodiment, the dichroic mirror 107 passes light having a long wavelength of, for example, 900 nm or more (light from the OCT light source 101), and reflects light having a short wavelength shorter than 900 nm (for example, light from a light source of 840 nm, SLO). Is set to The dichroic mirror 107 is not limited to the above specifications, and may be set as appropriate depending on the wavelength of the light source to be used.

As described above, light having two different wavelengths (OCT light and SLO light) irradiated and reflected on the eye E using the dichroic mirror 107 is appropriately divided to enable each measurement.

FIG. 4 shows how tomographic images (B-scan images) are acquired by the OCT optical system 100. 4A shows an example of the fundus retina of the eye E, and FIG. 4B shows a plurality of two-dimensional tomograms (B-scan images) of the fundus retina 401 obtained from the tomogram acquisition unit 100. An example is shown. FIG. 4C shows an example of a three-dimensional tomographic image of the fundus generated in the present embodiment. 4A to 4C, the x-axis indicates the B-scan scan direction, and the y-axis indicates the C-scan direction. Furthermore, the z-axis in FIGS. 4B and 4C indicates the depth direction of the A scan signal, that is, the depth direction of the fundus. FIG. 5 shows an OCT image actually taken and displayed on the monitor.

(SLO optical system 200)
The SLO optical system 200 includes a light source 201 to an ADC 210 that is an A / D converter. Usually, an SLO light source uses an infrared laser diode of 800 to 900 nm to acquire a fundus image non-invasively. In this embodiment, a 840 nm laser diode is used as the SLO light source. The SLO light source is not limited to the laser diode of this embodiment, but may be another light source such as an LED.

Measurement light output from the SLO light source 201 (hereinafter referred to as SLO measurement light in order to distinguish other measurement light) is reflected by the mirror 204. Here, the light irradiated to the fundus and the reflected light reflected from the fundus follow the same path. Therefore, in order to split the irradiation light and the reflected light, a half mirror or a beam splitter that reflects and transmits the mirror 204 at a predetermined ratio is employed. In order to reduce noise light caused by unintentional scattering and reflection in the optical system, a polarizing beam splitter may be employed for the mirror 204.

Therefore, a part of the SLO measurement light is reflected by the mirror 204 and input to the focus lens 203, and then passes through the scanning device 208 and the lens 202 and is input to the dichroic mirror 107. The input SLO measurement light is reflected by the dichroic mirror 107, passes through the objective lens 109, and is irradiated on the fundus of the eye to be examined. The focus lens 203 is controlled to move on the optical axis so that the SLO measurement light applied to the fundus is focused on the fundus.

The SLO measurement light reflected from the fundus is input in the reverse path through the objective lens 109, the dichroic mirror 107, the lens 202, the scanning device 208, and the focus lens 203, and a part of the light passes through the mirror 204. The light is input to the lens 205, passes through the pinhole 206 after being condensed, received by the photodetector 207, converted into an electrical signal, and then input to the ADC 210.

Here, similarly to the galvanometer mirror 105 in the OCT optical system 100 described above, the scanning device 208 scans the SLO measurement light with respect to the fundus of the eye to be examined in the X-axis direction and the Y-axis direction. The apparatus 208 can acquire the front image data of the fundus by scanning the irradiation position of the SLO measurement light. The scanning device 208 is not limited to a galvanometer mirror, and may use a polygon mirror or a configuration in which a galvanometer mirror and a polygon mirror are combined. The photodetector 207 is, for example, an avalanche photodiode or a photomultiplier tube.

As described above, the front surface image of the fundus of the eye E can be acquired by XY scanning the fundus, sampling the reflected light by the ADC 210, and processing the signal by the arithmetic unit 500. An SLO image actually taken and displayed on the monitor is shown in FIG.

(Operating procedure)
Next, an operation procedure of the ophthalmologic apparatus according to the present embodiment will be described.

FIG. 3 is a flowchart for explaining the operation procedure in this embodiment.
First, in S10, the head (also referred to as a head portion) on which the two optical systems are arranged is aligned with the eye of the eye to be examined (hereinafter referred to as alignment). The alignment is performed using a joystick (not shown) provided in the main body. In this embodiment, in order to facilitate alignment, a fixation lamp (not shown) is used to irradiate the subject's eye with a fixation lamp. To match). As the fixation optical system, a fixation optical system provided in a general ophthalmic apparatus can be used.

Next, in S12, focus adjustment of the SLO optical system is performed. The focus lens 203 is moved and controlled on the optical axis so that the SLO light emitted from the light source 201 is focused on the fundus (retina) of the eye to be examined. Then, the eye refractive power of the eye to be examined is also calculated from the value of the control signal obtained at this time and stored in the storage unit in the main body.

Next, in S14, the eye to be examined is fixed using a fixation optical system (not shown). As described above, the fixation method may be a method employed in a general ophthalmologic apparatus, or may be implemented by arranging a more effective fixation optical system.

In S16, SLO imaging is performed to obtain a front image of the fundus.

When the front image of the fundus is acquired in S16, the front image of the fundus is displayed on a monitor (not shown). The display method may be displayed in real time, or may be a method of switching and displaying images by the examiner's operation. In S18, the front image of the fundus displayed on the monitor is viewed, and two or more imaging positions (fundus position) of the OCT imaging are determined. The determination may be made by the examiner looking at the monitor, or if the imaging position is determined in advance (for example, a plurality of positions including specific blood vessels set in advance), the front image of the fundus is calculated by SLO. The imaging position of OCT imaging may be determined by analyzing by the unit 500 and automatically detecting the imaging position. FIG. 6 shows an example of the imaging position of the OCT imaging.

When the imaging position of the OCT imaging is determined, OCT focus adjustment is performed in S20. Focus adjustment is performed by moving the focus lens 104 on the optical axis and controlling the position of the focus lens 104 so that the OCT light is focused on the fundus (retina), which is the imaging target position.

When the OCT focus adjustment is completed, the OCT zero point adjustment is performed in S22. Zero point adjustment refers to the optical path length (referred to as measurement optical path length) of the measurement optical system (referred to as measurement optical path length) in the OCT optical system (which irradiates the eye to be examined to obtain reflected light) and the optical path length (referred to as reference optical path length) of the above-described reference optical system ), The prism 112 in the reference optical system is controlled to move on the optical axis. In this embodiment, the prism 112 is employed, but the present invention is not limited to the prism, and a mirror or the like may be employed.

When the OCT zero point adjustment is completed in S22, "m = 0" is set in S24. In this embodiment, since the B-scan images are acquired at two or more fundus positions, the first fundus position is counted as the 0th position.

In S26, the OCT control is performed by the arithmetic unit 500 so that the OCT B scan rate is 300 A scan / B scan. The B scan rate is not limited to this, and may be set as appropriate in consideration of the entire measurement time. In this embodiment, the purpose is to detect the position of the blood vessel, so that a high-accuracy image is not required and the overall measurement time is reduced. A B scan rate of 300 A scan / B scan, which is a small A scan image to be acquired, that is, a rate at which a relatively coarse image is acquired) is employed.

In S28, “n = 0” is set. In this embodiment, in order to acquire B scan images at the same fundus position of a plurality of sheets, the first B scan image is counted as the 0th B scan image.

In step S30, OCT imaging is started. In the case of SS-OCT, which is a Fourier domain type OCT as in this embodiment, scanning in the depth direction (Z direction) is not necessary, so by scanning once in the X direction or Y direction by the galvanometer mirror 105, Since a plurality of A-scan images (also referred to as one-dimensional tomographic images) in the scanned range can be acquired, in this embodiment, hereinafter, OCT imaging is treated as B scan imaging and consent. An image obtained by OCT imaging is called a B scan image or a B scan image. In this embodiment, the B-scan image is a two-dimensional tomographic (image) image in the depth direction of the fundus. Then, the operations of S30 to S40 described below are repeated until a preset number of sheets (that is, a preset number n (final)) is acquired.

SLO imaging (S16) is performed in parallel with OCT imaging (agreement with B-scan imaging). Then, the movement of the fundus is detected (S32) from the SLO image acquired in the SLO imaging (S16), the movement of the fundus is calculated in S34, and the OCT is calculated in S36 based on the calculated movement of the fundus. The position of the scan is corrected (changed), and the next OCT image is taken (S38, n = n + 1).

OCT scan position correction (sometimes referred to as eye tracking) using the SLO image of S36 is not necessarily required. In the case of the present embodiment, since the object to be examined is a human eye, even if the subject is fixed, microscopic movements of the fixation always occur, and motion artifacts are likely to enter the acquired OCT image. If eye tracking is performed and the test object is not a living body and motion artifacts need not be considered, eye tracking as described above is not necessary. The process of S36 can be deleted.

If the number of shots reaches a predetermined number (S40, n> n (final)), the process proceeds to S42. If the number of shots is less than the predetermined number, the process returns to S30 and OCT imaging is continued.

In S42, the current OCT scan rate is confirmed. If the OCT scan rate is 300 A scan / B scan, the process proceeds to S 44.

(Acquisition of blood vessel position coordinates)
When a predetermined number of B-scan images at the same fundus position are acquired by 300 A scan / B scan, a phase difference (ΔφB (n) (x, z)) is calculated between two adjacent B-scan images in S44. . (N indicates the number of the acquired B-scan image.)

A plurality of phase differences (ΔφB (n) (x, z)) calculated in S44 are averaged in S46, and the position coordinates (d (m) (x, z)) of the blood vessel are obtained from the result. calculate. (M represents the number of the fundus measurement position.)

After calculating the position coordinates (d (m) (x, z)) of the blood vessel, the position coordinates are stored in the storage section, and in S48, the calculation section so that the OCT scan rate is 8,000 A scan / B scan. OCT control is performed by 500. The B scan rate is not limited to this, and may be appropriately set in consideration of the entire measurement time. However, the OCT scan rate is set so that the positions of adjacent A scan images are substantially the same. It is desirable to set.

In step S50, an OCT scan area is set. Since the purpose is to obtain the blood flow velocity of the blood vessel at the coordinate position obtained in S46, the region is centered on the blood vessel position, and the region where the OCT scan rate is narrower than that in the case of 300 A scan / B scan. It is desirable to set. This is because the OCT scan rate of the 8,000 A scan / B scan is a very slow scan rate, and the larger the area to be scanned, the longer the measurement takes. By appropriately setting the region width to such an extent that the blood flow velocity can be calculated, a highly accurate blood flow velocity can be calculated, and the overall measurement time can be shortened.

In S48, the OCT scan rate is changed, and when the scan area width is set in S50, the process returns to S28, and OCT imaging of S30 to S40 is performed at the scan rate of 8,000 A scan / B scan. When a predetermined number of images have been shot (n> n (final)), the process proceeds to S42.

In this case, since the scan rate is 8,000 A scan / B scan, the flow proceeds to S 52, and the blood flow velocity is calculated from the B scan image captured by the 8,000 A scan / B scan.

First, a phase difference (ΔφA (n) (z)) between two adjacent A scan images with respect to one B scan image taken in S52 is calculated. ((N indicates the number of the acquired B-scan image)

Next, averaging processing or the like is performed on the values of the plurality of phase differences (ΔφA (n) (z)) calculated in each B-scan image in S54, and (vertical) at the mth fundus position. The blood flow velocity (V (m) (z)) is calculated. (M represents the number of the fundus measurement position.)

In S56, OCT imaging is performed at the next measurement position (m = m + 1). If the OCT imaging is not performed at the predetermined number of measurement positions in S58, the measurement position is changed and the process returns to S26 to continue the OCT imaging at the next measurement position. When OCT imaging is completed at a predetermined number of fundus positions (m> m (final)), the OCT imaging is completed, and the blood vessel gradient is calculated from the blood vessel position information obtained at each fundus measurement position. The absolute blood flow of the target blood vessel can be calculated from the gradient and each (vertical) blood flow velocity. In addition, since the various methods already known by patent document 1 etc. can be employ | adopted for the calculation method which calculates a blood flow velocity etc. from phase difference information, the detail is abbreviate | omitted here.

The embodiments of the present invention have been described in detail above. However, these are merely examples, and the present invention is not construed as being limited by specific descriptions in the embodiments. The present invention can be carried out in a mode in which various changes, modifications, improvements, etc. are added based on the knowledge, and such a mode is within the scope of the present invention as long as it does not depart from the gist of the present invention. It should be understood that it is included in.

For example, in the above embodiment, the SLO is employed for acquiring the fundus front image, but the present invention is not limited to the SLO, and a fundus camera may be employed. Furthermore, if an emphasis image (FIG. 5B) created from an OCT C-scan image is used to specify a position for obtaining a blood flow, fundus front image acquisition means such as an SLO or a fundus camera is obtained. Is not necessarily required. In such a case, it is also possible to perform eye tracking using, for example, an anterior ocular segment image obtained by separately arranging an anterior ocular segment imaging optical system. The anterior ocular segment imaging optical system can be arranged together with the fundus frontal imaging optical system, and eye tracking can be performed using both captured images.

Further, in the above embodiment, the features of the present invention have been described by taking Doppler OCT for measuring blood flow of blood vessels as an example. However, the method according to the present invention is not limited to Doppler OCT, and even in normal OCT. Is available. For example, when there is a lesion in the fundus and it is desired to perform OCT imaging of the lesion site with high accuracy, the method according to the present invention can be used to provide information that is measured without taking time and is effective for diagnosis. . In other words, in the above-described embodiment, a lesion position is detected at a scan rate of 300 A scan / B scan, and imaging is limited to the periphery of the lesion at a scan rate of 8,000 A scan / B scan, so that a high-accuracy tomography relating to the lesion Images can be acquired in a short time.

DESCRIPTION OF SYMBOLS 1 ... Ophthalmic apparatus of a present Example, 100 ... OCT optical system, 101 ... Light source, 102 ... Fiber coupler, 105 ... Galvano mirror, 107 ... Dichroic mirror, 112 ... Prism, 200 ... SLO optical system, 201 ... SLO light source, 204 DESCRIPTION OF SYMBOLS ... Beam splitter, 208 ... SLO scanning apparatus, 300 ... Anterior eye imaging | photography optical system, 303 ... CCD camera, 500 ... Operation part, E ... Eye to be examined

Claims (7)

The tomographic image of the fundus of the eye to be examined is obtained from the interference light obtained by splitting the light from the light source into reference light and measurement light, irradiating the branched measurement light on the fundus, and combining the reflected measurement light and reference light An ophthalmologic apparatus having fundus tomographic image acquisition means for
An ophthalmologic apparatus comprising scan control means capable of performing scan control of measurement light for acquiring a two-dimensional tomographic image (B scan image) extending in the depth direction of the fundus at at least two different speeds. A fundus front image acquisition unit that illuminates the fundus of the eye to be examined, receives reflected light or scattered light from the fundus with a two-dimensional image element, and acquires a front image of the fundus of the eye to be examined;
The ophthalmologic apparatus according to claim 1, further comprising a B-scan designation unit that designates a position at which an arbitrary B-scan image is acquired from the acquired frontal image of the fundus of the eye to be examined. The ophthalmologic apparatus according to claim 1 or 2, wherein at least two of the at least two different scanning speeds controlled by the scanning control means differ by 10 times or more. A first phase difference acquisition unit that acquires a phase difference between adjacent A scan images (substantially the same position) with respect to a B scan image acquired at a slow scanning speed of at least two different scanning speeds. The ophthalmic apparatus according to 3. 2nd phase difference acquisition means which acquires the phase difference between the same positions of the adjacent B scan image acquired continuously with respect to the B scan image acquired at high scanning speed among at least two different scanning speeds The ophthalmic apparatus according to claim 4. The ophthalmologic apparatus according to any one of claims 2 to 5, wherein the B-scan designating unit designates at least two different positions including the same blood vessel. The ophthalmologic apparatus according to any one of claims 2 to 6, wherein the scanning direction of the scanning control means is a direction that intersects the target blood vessel substantially perpendicularly.