JPH01256922A - Ophthalmologic diagnosing device - Google Patents

Abstract

PURPOSE:To simply and precisely measure the blood flow speed of a specific vein of an eyeground via the laser speckle method by making the laser beam spot diameter adjustable and making the laser beam spot position movable within the preset angle of view narrower than the observation angle of view. CONSTITUTION:After the laser light flux from a laser light source 1 is set to a proper beam diameter by an opening 6, the beam spot diameter on the eyeground 18b of an eye 18 under test is adjusted by a laser focusing lens 7. A movable mirror 9 is installed at the position nearly conjugate to the cornea or pupil, thus a beam can be moved on the eyeground without largely changing the laser beam incident position on the cornea of the eye 18 under test. The laser light is arranged in the same light path as the observation photographing light flux, the radiation of the laser light by the movable mirror 9 to the desired position of the eyeground 18b can be performed in the observation photographing visual field by utilizing the horizontal and vertical swinging mechanism and the fixed view guiding mechanism of an eyeground camera, thus it is very convenient.

Description

Detailed Description of the Invention [Industrial Field of Application] The present invention irradiates the fundus with a laser beam of a predetermined diameter and detects the movement of a laser speckle pattern formed on an observation surface by scattered reflected light from the fundus tissue. The present invention relates to an ophthalmologic diagnostic device that detects changes in speckle light intensity and measures the blood flow state of fundus tissue based on the analysis results of the obtained speckle signal.

[Prior Art] Conventionally, as a method of measuring the blood flow state of the fundus using laser light, there is a method disclosed in Japanese Patent Application Laid-Open No. 55-75668.55-75669.
．． Methods described in JP-A-55-75670, JP-A-56-125033, JP-A-58-118730, etc. are known. Since both of these methods determine the blood flow velocity based on the Doppler effect of laser light, it is necessary to detect the Doppler shift frequency, so the incident laser light is divided into two at equal angles to the optical axis and delivered to the subject's eye. This requires a configuration in which the laser beams scattered by the blood cells in the fundus are extracted from two different directions and detected, which requires a very complicated optical system configuration. is complex and requires precision.

Furthermore, the fact that the angle of incidence or the angle of detection must be known is extremely troublesome and difficult to obtain reproducible and reliable results in clinical applications, where each patient's eyes are different. is difficult. This is because the laser Doppler method is originally a precise and sensitive measurement method, so although it may be effective in the industrial field where the target is stable and stationary, it is also effective in the industrial field where the target environment and conditions are unstable and have large variations. This shows that in the other medical fields, they are instead affected by various influences, which significantly reduces the reproducibility of measurement results.

Furthermore, in actual measurement results, the Doppler shift frequency cannot be obtained as a single frequency, and various frequency components exist over a wide band from the low frequency side to the high frequency side, and as a result, it is difficult to obtain a reliable absolute velocity.

Furthermore, in order to irradiate the fundus with laser light, it must enter the fundus from a direction nearly perpendicular to the eye, so it is difficult to cause the Doppler effect and it is quite difficult to detect a beat signal. This is difficult because the laser Doppler method detects a single beat component. It is preferable to apply the laser speckle method, which is an effective method.

It is generally known that when a scattering object is irradiated with laser light, the scattered light forms a random speckle pattern due to the interference phenomenon of coherent light. Furthermore, if the scattering object moves, the speckle pattern also moves, so if this movement is detected as a time change in light intensity at an observation point, the movement of the object can be measured from the degree of signal change. The present invention applies this to the measurement of the state of blood flow in tissues such as the fundus of the eye.

[Problems to be Solved by the Invention] Examples of applying the speckle phenomenon to blood flow measurement include JP-A No. 1i0-199430 and JP-A No. 60-203235.
, JP-A No. 60-203236, etc. However, these methods are intended for measurements on the skin surface, and are almost impossible to use for measuring fundus blood flow due to laser irradiation, detection optical systems, and light intensity.

Therefore, the inventor of this application has already filed an application for a diagnostic method and apparatus for measuring blood flow in the eye region using the speckle method (Japanese Patent Laid-Open No. 62-275431).

However, in this method, a laser beam spot is irradiated over a range wider than the diameter of a single blood vessel in the eye, and the speckle pattern on the Fraunhofer diffraction surface is created in which scattered light from multiple blood vessels included in the irradiation area is superimposed. By detecting the 8th shift, the stability and reproducibility of measurement results is improved. Therefore, this method is excellent in that it evaluates the average blood flow activation state within any irradiated area of the fundus, but it is not suitable for evaluating the blood flow velocity of a specific blood vessel within the irradiated area. was inappropriate.

Therefore, the inventor of this application has also filed an application for a diagnostic device based on the speckle method using a new detection optical system that can evaluate the blood flow velocity of a specific blood vessel (Japanese Patent Application No. 1983-
75778.62-75779). However, in these methods, in order to select a specific blood vessel, a detection aperture (pinhole or slit) is placed on the image of the blood vessel to be measured on the enlarged image plane.
Therefore, as a means of selecting the position while visually observing the fundus image, an index is provided on the observation eyepiece, and by aligning this with the position of the target blood vessel within the eyepiece field of view, it is indirectly linked to the index. The detected aperture is set at the position of the corresponding katsura vessel image on the enlarged image plane. Therefore, it has been found that the interlocking mechanism is complicated and the device is expensive, while the mechanical adjustment of the index and detection aperture during manufacturing is troublesome. In addition, the mechanical play included in the interlocking mechanism directly leads to errors in positioning, and there are other problems such as poor operational responsiveness. In order to further identify the target blood vessel, we first perform positioning using a fixation target to irradiate the area containing the blood vessel with laser light, and then select a single blood vessel using the index on the eyepiece. This requires a two-step operation, and during this process, the detection position may shift due to eyeball movements of the subject's eye. It also had poor operability, meaning you had to start over from the beginning.

On the other hand, since ZARPIM irradiates an area on the fundus of the eye that is wider than the diameter of the blood vessels, the amount of scattered light from tissues other than the blood vessels within that area is greater than that from the vascular blood flow.
Blood vessels and surrounding tissues cannot be clearly separated on the magnified image plane. In order to improve this, filtering is performed in terms of spatial frequency, but as a result, there are also problems in that the optical system becomes complicated and the number of detection targets is significantly reduced.

Therefore, the present invention solves the above-mentioned problems and provides an ophthalmologic diagnostic method and device that can easily and accurately measure the blood flow velocity of a specific blood vessel in the fundus of the eye using the laser speckle method. The purpose is to

[Means for Solving the Problems] In order to solve the above-mentioned problems, the present invention irradiates the fundus with a laser beam of a predetermined diameter and forms a laser speckle pattern on the observation surface by scattered reflected light from the fundus tissue. In an ophthalmological diagnostic device that detects the movement of the eye as a speckle light intensity change and measures the blood flow state of the fundus tissue based on the analysis result of the obtained speckle signal, the ophthalmologic diagnostic device includes a laser light source;
an optical system that irradiates a laser beam from the laser light source onto the target blood vessel to be measured as a minute beam spot diameter that is the same as or smaller than the diameter of the blood vessel; and an optical system that focuses the scattered and reflected light from the fundus. an optical system that forms a same-magnification or enlarged image of the irradiated portion on a conjugate image plane; a multiple detection amount lobe pattern having a plurality of minute apertures arranged on the conjugate image plane; The boiling motion of speckles is detected through the multiple detection amount roboturn, and the blood flow velocity in the target blood vessel is determined by analyzing and evaluating the speckle signal according to the intensity change of the total amount of light that has passed through this multiple detection amount roboturn. The laser beam spot diameter can be adjusted, and the laser beam spot position can be moved within a predetermined field angle range that is narrower than the observation field angle.

[Function] In this configuration, the diameter of the laser beam spot is set to be equal to or smaller than the diameter of the target blood vessel from the beginning, and by irradiating the spot toward the blood vessel C to be measured, the position can be determined by a single operation of the spot. Easy to match.

Furthermore, almost all of the scattered reflected light from the laser is scattered light from the blood flow, and there is almost no light from the surrounding tissue, so the focused laser speckle light flux is divided into all parts of the light flux at the speckle detection surface. It can be used for signal detection. Moreover, if this was simply detected using a single aperture with a large diameter, the speckle light intensity distribution would be averaged out within the aperture, and the signal S/N ratio would drop markedly. By using multiple detection aperture patterns, signals can be detected without averaging and the amount of detected light can be increased, making effective use of the above-mentioned advantages of speckle light flux, which is mostly composed of scattered light from blood flow. has succeeded in making use of it.

In addition, the detection of speckle movement using this multiple detection aperture is
Magazine "Optics" Volume 11, No. 3 (June 1982) 29
As described on page 1-297, the target speckle movement is a translational movement, that is, a movement in which the speckle pattern on the detection surface translates while preserving its shape as the object moves. In this case, the cross-correlation component between any two adjacent apertures along the six speckle movement directions is detected, and the signal component is different from that detected with a single aperture. It was a problem. Alternatively, in order to reduce this effect, it is necessary to make the distance between each of the plurality of apertures sufficiently wide, and as a result, there is a problem that it becomes impossible to arrange a large number of minute apertures in the detection light beam. However, the inventors of this application found from various experiments that the movement of the laser speckle pattern due to the scattered reflection of the fundus causes boiling motion, that is, the speckles on the detection surface do not translate relative to the movement of the object and remain in the same position. He invented a wavering motion of light and dark spot patterns that constantly change their shape, and used multiple detection aperture patterns to detect this boiling motion. Boiling is not affected by cross-correlation between adjacent apertures, and a large number of minute apertures can be arranged, so the characteristics of multiple apertures can be effectively utilized. Of course, since the amount of detected light can be greatly increased and the measurement time can be shortened, errors caused by positional deviation during measurement due to eye movement, re-measurement, and burden on the subject can be greatly reduced, which greatly improves safety. can give favorable results.

On the other hand, similar to the detection of speckle movement using multiple detection aperture patterns, the inventor of this application filed a patent application
1-226107, in which the multiple apertures are randomly arranged random patterns, and in this case the signal component greatly depends on the randomness of the random pattern to be produced and the numerical aperture, so the same signal is generated in many random patterns. Obtaining the ingredients poses a major obstacle in terms of reproducibility and stability, and cannot be said to be good in terms of product production. Furthermore, the uneven intensity distribution of the speckle light beam on the detection surface due to the random arrangement cannot be ignored. The above-mentioned problem is solved by regularly arranging a plurality of apertures.

However, based on this method, the entire area within the laser beam spot irradiated to the fundus of the eye is subject to photodetection, so if a beam that is thicker than the blood vessel diameter is irradiated, specks due to scattering from surrounding tissues other than the vascular blood flow may be detected. The light flux is also detected, and since this light does not contain blood flow information, it becomes noise and deteriorates the S/N ratio of the signal. Therefore, a laser beam with a size similar to or smaller than the diameter of the target blood vessel in the fundus is required, but since the diameter of blood vessels varies depending on the individual and the location of the same person, it is difficult to perform accurate measurements unless there is an adjustment mechanism. I can't. Furthermore, making it possible to adjust the laser beam spot diameter over a wider range than necessary would complicate the mechanism, and the beam spot would become distorted into an ellipse due to lens aberrations, or the beam intensity distribution would become uneven. This device allows for appropriate adjustment by specifying the beam diameter adjustment range.

In addition, in order to align the laser beam spot with the desired blood vessel measurement site, the beam spot must be movable, but if the beam spot moves outside the field of view of the fundus and is lost, the current I don't know where the beam spot is located,
It takes time to bring the beam spot back into the field of view,
It is also unfavorable in terms of clinical application. In this device, by limiting the beam spot movable range in advance,
The above problems have been resolved. Furthermore, it is inconvenient to operate if the runner is not informed of the restricted movable range of the beam spot. In this device, a reticle indicating the range is set in the observation optical system, and this reticle can also be used by the measurer to correct the diopter.

In addition, in order to accurately irradiate the laser beam toward the blood vessels in the fundus, it is preferable that the fundus be seen as enlarged as much as possible. It's better to take it. Therefore, it is necessary to be able to switch the viewing angle depending on the purpose. In a normal fundus camera, a lens for changing the angle of view is prepared in the light-receiving optical system, and the lens is switched stepwise.However, in this device, the speckle pattern is detected through the light-receiving optical system, so the image is formed as the angle of view is changed. It is inconvenient if speckle detection conditions change due to changes in magnification. Therefore, the above objective is achieved by using a zoom type optical system as the observation eyepiece (eyepiece).

In addition, it is naturally necessary to adjust the light intensity of the laser beam spot that illuminates the fundus, but switching the light intensity from positioning before measurement to during measurement cannot be performed smoothly if done manually. . Furthermore, since the amount of light for measurement differs from person to person, depending on factors such as the reflectance of the fundus and the lens, it is necessary to switch between several types. However, this 1
A simple method using one mechanism is not suitable for high-speed switching at the start and end of measurement. This problem is solved by using two light amount adjustment mechanisms, each of which adjusts the amount of light independently.

[Embodiments] The first basic invention is particularly directed to the fundus of the eye, and the following embodiments will be explained based on the optical system of a fundus camera.

In Fig. 1, for example, red He-Ne (wavelength 6
32.8r+m) The laser beam from the laser light source 1 passes through two light intensity adjustment filters 2 and 3 and a condenser lens 4, and becomes a parallel beam at the collimating lens 5.
After an appropriate beam diameter is set using the aperture 6, the beam spot diameter on the fundus 18b of the eye 18 to be examined is adjusted using the laser focusing lens 7. The beam is further reflected by a movable mirror 9 via a mirror 8, and further reflected by a relay lens 1.
As shown in FIG. 2, the light beam for fundus observation and photographing enters the fundus after being reflected by a mirror 13 installed in a part of the annular opening 12a of the ring slit 12 in the fundus camera illumination optical system. is guided along the same optical path. For this reason, the laser beam is transmitted through the relay lens 14, 15 and the perforated mirror 16.
The light passes through the objective lens 1f, reaches the fundus 18b from the cornea 18a of the eye 18 to be examined, and is irradiated onto the blood vessel to be measured.

In the above laser irradiation optical system, there is a shutter 19 near the exit of the laser light source 1, which opens and closes as necessary. The movable mirror 9 is used to enable eight movements of the beam spot position on the fundus, and the eight movements are made by, for example, manipulating the manipulator 2° to move the movable mirror 9 in both the X and 7 directions relative to the optical axis. The method of independently changing the tilt angle of the mirrors, that is, the method normally used in coagulators, etc., can be used as is.

Furthermore, as is well known, the position of the movable mirror 9 is set at a position substantially conjugate with the cornea or pupil, so that the eye to be examined can be
The beam can be moved on the fundus of the eye without significantly changing the laser beam incident position on the cornea.

The measurement area of the fundus is illuminated by an illumination optical system used as a fundus camera to facilitate observation. This observation optical system includes an observation light source 21 arranged on the same optical axis as the photographing light source 22, a condenser lens 23, a condenser lens 24,
It is composed of a filter 25 and a mirror 26. Since the laser beam is placed on the same optical path as this observation/photographing light beam, by using the left/right and up/down swing mechanisms of the fundus camera and the fixation 8 guide mechanism, the movable mirror 9 can direct the laser beam to the desired position on the fundus 18b. This is very convenient because it allows irradiation to be carried out within the field of view for observation and photographing.

Note that the filter 25 disposed between the condenser lens 24 and the mirror 26 is configured as a wavelength separation filter having spectral characteristics as shown in FIG. 3, so that the red component contained in the observation and photographing light is cut. . Appropriate spectral characteristics are used depending on the wavelength of the laser light source used.

Speckle light generated when the laser light is scattered by blood cells moving in the blood vessels to be measured in the fundus of the eye and other reflected light for observation and photography are both received by the objective lens 17 and passed through the perforated mirror 16. Focusing lens 27,
An image is formed once on the surface of the space 31a via the relay lens 28, mirror 29, and flip-up mirror 30, and then a fundus image is formed on the surface of the reticle 35 via the flip-up mirror 32, relay lens 33, and flip-up mirror 34. be done. This image is observed by a zoom eyepiece 36 that can change magnification.

Here, the zoom type eyepiece 36 is designed to allow the observer to perform diopter correction using the reticle 35 as a reference.

When taking a photograph, the flip-up mirror 32 is flipped up to 32' in the direction of the arrow using 32a as a fulcrum, and the observation photographing light beam including the laser speckle light from the fundus that has been reflected by the flip-up mirror 30 is passed through the imaging lens 37. An image is formed on the photographic film 38 and photographed. As mentioned above, it is usually possible to observe and photograph the fundus as a fundus camera, and if it is irradiated with laser light, the condition can be observed and photographed, so measurement points can be directly confirmed and recorded. , is extremely useful.

On the other hand, in the case of blood flow measurement, a flip-up mirror 30 linked to a measurement switch 9, which will be described later, is placed at a position 30 with 30a as a fulcrum.
At the same time, another flip-up mirror 34, which is also linked to the measuring switch 9, is flipped up to position 34' using 34a as a fulcrum. Therefore mirror 2
The laser speckle light and the observation photographing light reflected from the fundus 9 are once imaged on the spatial plane 31b, which is another imaging point optically equivalent to the imaging point 31a. Furthermore, a wavelength separation mirror 39 fixedly installed behind it at an angle of about 45 degrees to the optical axis has the same spectral characteristics as shown in FIG. Reflects most of the speckle light caused by laser light. The reflected speckle light is sent to the relay lens 40 and the microscope objective lens 4.
An enlarged image of only the blood vessel region irradiated with the laser beam spot on the fundus through the laser beam spot 1 is formed on the surface of a multi-detection amount robot pattern 42 having a plurality of minute apertures set in a conjugate plane with the fundus. The speckle light flux that has passed through this detection opening 42 is collected by a condenser lens 43, and is sent to a photodetector (photomultiple) 45 via an interference filter 44 that allows only the light with a wavelength of 832.8 nm from the red He-Ne laser to pass through. Light detected. A shutter 46 is arranged in front of the photomultiplex 45, and the output signal from the photomultiplex 45 obtained when the shutter is opened 1 is sent to the analysis section 5°.

Note that the observation photographing light beam other than the red component that has passed through the wavelength separation mirror 39 and the speckle light beam that has passed through the wavelength separation mirror 39 form a fundus image on the surface of the reticle 35 via the mirror 47 and the relay lens 48, and the same as described above is used for zooming. formula eyepiece 36
observed by. In this way, the fundus can be observed even when measuring blood flow, which is very effective in preventing mistakes such as measuring without noticing that the target position has shifted.

An example of the configuration of the analysis section 50 is as shown in FIG.
1. Filter 52, analog-to-digital converter (A/D)
53, microcomputer (MC) 54, memory 55
, a keyboard 56, a CRT 57, and a printer 58.

The basic principle of evaluating blood flow velocity based on the above configuration will be explained below.

In FIG. 5, a laser beam 60 is irradiated onto a measuring portion 61' of a blood vessel 61 to be measured on the fundus retina. The irradiated measurement section 61'' is placed on the surface of a multiple detection aperture pattern 42 having a plurality of minute apertures shown in FIG. It is formed as an enlarged image.

This situation is illustrated in FIG. Since only the laser speckle light reflected by the wavelength separation mirror 39 in FIG. A blood vessel image is not formed.Here, the respective lights scattered by a large number of blood cells etc. moving within the blood vessel in the measuring section 61'' overlap and interfere with each other with random phases on the conjugate magnified image plane which is the observation surface. As a result, it is known that a laser speckle pattern that looks like a spatially random arrangement of speckles is formed. Furthermore, it is known that if blood cells, which are objects that cause scattering, move at their initial velocity, each minute speckle forming the speckle pattern moves in proportion to the object's speed.

Therefore, in this case as well, the enlarged image 62 is actually in the form of an image plane speckle pattern, and is observed as a pattern in which image plane speckles 63 are randomly present, rather than an enlarged image of each blood cell. Since each individual image surface speckle 63 moves according to the blood flow velocity, this is
As shown in FIG. 8(B), detection is performed by a photodetector 67 through a condensing lens 66 through a multi-detection amount robot pattern 65 consisting of a plurality of minute apertures 65', as shown in FIG. 8(B). Temporal changes in light intensity can be extracted.

Since the degree of change over time corresponds to the blood flow velocity, the blood flow velocity can be measured by examining the power spectrum or autocorrelation function of this light intensity change signal. In addition, in FIG. 7, the image plane speckle 63 and the multiple detection amount lobe pattern 65 are both on the enlarged image plane M.

In the example of the conventional prior application (Japanese Patent Application No. 82-75778), as shown in FIG. In addition, reflected laser speckle light from the surrounding tissue 69 reaches the magnified image plane. Therefore, in order to extract blood flow information, it is necessary to align a minute detection aperture such as a pinhole on the image of the blood vessel to be measured on the enlarged image plane, and this operation is very troublesome. That is, in order to correctly align the detection aperture on the magnified image of the blood vessel to be measured while observing the fundus, an index linked to the detection aperture is provided on the observation eyepiece, and this is positioned over the target blood vessel within the field of view. It was necessary to devise measures such as

Therefore, the interlocking mechanism is complicated and the device is expensive, while the mechanical adjustment of the index and detection aperture during manufacturing is troublesome. Further, there were other problems such as mechanical play included in the interlocking mechanism leading to errors in setting the position of the interlocking mechanism, and poor operational responsiveness. In order to further identify the target blood vessel, we first perform positioning using a fixation target to irradiate the area containing the blood vessel with laser light, and then select a single blood vessel using the index on the eyepiece. This requires a two-step operation, and while this is being performed, the detection position may shift due to eyeball movement of the subject's eye, resulting in poor operability as the procedure must be repeated from the beginning.

On the other hand, since the laser beam irradiates an area on the fundus of the eye that is wider than the diameter of the blood vessel, as shown in FIG. Because it occurs with a greater amount of light than the
Blood vessels and surrounding tissues cannot be clearly separated on the magnified image plane. In order to improve this, filtering is performed in terms of spatial frequency, but as a result, the optical system becomes complicated and the amount of detected light is significantly reduced.

In this embodiment, as shown in FIG.
The measured blood vessel is irradiated with a beam spot diameter of 0 equal to or smaller than the blood vessel diameter, and as a result, the speckle light flux on the conjugate magnified image plane originally contains scattered light from surrounding tissue. As shown in FIG. 6, only image plane speckles 63 have a motion reflecting the blood flow velocity in the blood vessels.

Therefore, the image plane speckle pattern formed on the conjugate enlarged image plane M allows signal detection with blood flow information no matter where the detection aperture is located within the light beam. Therefore, in order to select the part of the blood vessel to be measured from among the various blood vessels present in the fundus visual field, it is sufficient to simply irradiate the part with a minute spot of a laser beam. This is the first
This can be achieved by controlling the movable mirror 9 shown in the figure by operating a manipulator 20. In addition, the beam spot diameter must be equal to or less than the diameter of the blood vessel at the site to be measured; on the other hand, since the diameter of the blood vessel varies from person to person and varies from place to place even in the same person, correct measurement is required without a mechanism to adjust the laser beam spot diameter. can't do it. However, making it possible to adjust the laser beam spot diameter over a wider range than necessary requires a complicated mechanism, and lens aberrations may cause the beam spot to become distorted into an ellipse or cause unevenness in the beam intensity distribution.

Therefore, in this embodiment, the beam spot diameter adjustment range is defined to enable appropriate adjustment. In other words, this adjustment can be carried out continuously or stepwise by moving the laser focusing lens back and forth on the optical axis, or by changing the aperture 6 into a turret type with apertures of various diameters. I can do it.

Normally, the diameter of the fundus retinal blood vessels is approximately 150 μm (diameter) at most, so the beam spot diameter should be adjusted to a maximum diameter of 200 μm or less (preferably. This is because unnecessary speckle light will be detected.

The problem here is that by moving the beam spot to a desired position on the fundus, the position of the corresponding spot enlarged image (enlarged image of the measuring section) on the conjugate enlarged image plane also moves. If the detection aperture were moved accordingly, a complicated mechanism similar to the conventional one would be required. Therefore, in the present invention, as shown in FIG. 10, a predetermined angle of view that is included in the observation and photographing field of view 70 of the fundus camera and that is smaller than the angle of view of the field of view 70 (angle of view 73 shown in FIG. 11) is used. (11th
The configuration is set in advance so that the laser beam spot can be freely moved only within the range 71 of the angle of view 72) in the figure. As an example, if a range of about 3 mm in diameter is measured near the center of the visual field, the laser beam can be moved relatively easily. Therefore, if you want to measure a part that is outside the movable range 71, all you have to do is position the target part within the movable range 71 using fixation guidance, etc., and this is an easy operation and is not troublesome at all. do not have.

Next, FIG. 12(A) is also shown on the conjugate enlarged image plane M.

As shown in (B), the movable range 7 of the enlarged laser beam spot image (enlarged image of the measurement unit) corresponding to the above-mentioned laser beam movable range 71 within the visual field of the fundus surface S
2 is determined, a plurality of detection aperture pattern 65 is installed that sufficiently includes at least this movable range 72 and has a plurality of microscopic apertures 65' arranged over an area wider than this range 72. With this configuration, for example, in FIGS. 12(A) and 12(B), the irradiation laser beam 60a is
From part 61a to Bib, 61 like 0b, 60c
When moving in order to irradiate different parts, corresponding enlarged images 62a, 62b, and 62c of each measurement part are formed in order on the conjugate enlarged image plane M. However, since the magnified spot image at any position is always located somewhere within the multiple detection aperture pattern 65 installed on the same conjugate magnified image plane, the speckle movement caused by one of the minute apertures 65' at that position Changes in light intensity can be detected, and no positioning on the detection plane, that is, on the conjugate magnified image plane M is required. In other words, with this method, all you have to do is move the laser beam within a predetermined range while looking at the fundus and select an arbitrary part, and that part is determined as the measurement position.
The operation is straightforward and extremely practical.

Furthermore, another major feature here is that the enlarged spot image (for example 62a) formed on the conjugate enlarged image plane M in FIG.
) is used for detecting the movement of each image plane speckle 63 in the image plane. Therefore, for example, for a speckle signal (FIG. 8(A)) of a change in light intensity when one point in the enlarged spot image 62 is detected with a single opening 64 as shown in FIG. 7(A), As shown in FIG. 7(B), a completely similar enlarged spot image 62 is formed by a large number of minute apertures 65' (
However, the aperture diameter is assumed to be equal to the diameter of the single aperture 64 in FIG. The speckle signal (FIG. 8(B)) of the change in light intensity when detected by the multiple detection amount robot pattern 65 consisting of ) has a significantly higher intensity, that is, it has the advantage of increasing the amount of detected light.

When using a laser as an ophthalmological diagnostic device, from the standpoint of safety, there is a need to keep the irradiation laser intensity as low as possible and to do it in a short time. Furthermore, in order to avoid being affected by the eye movement of the subject's eye during measurement or the shaking of the entire measurement system, and to reduce the burden on the subject,
Short-time measurements are essential. However, since the laser reflectance of the fundus is generally low and cannot be controlled artificially, the best method is to improve the detection light intensity sensitivity. It can be said that the present invention has great practicality in this respect as well, by applying a multi-detection quantity robot pattern having a plurality of minute apertures.

To summarize the above characteristics, a laser beam is irradiated only to the blood flow part of the blood vessel, a speckle pattern is formed using only the scattered light of the blood flow signal component, and multiple micro-apertures are formed across the entire surface of the pattern. Detection is performed at multiple points using a multiple detection amount robot pattern. Moreover, the detection aperture is installed over a wide range on the detection surface, so that even if the spot image moves arbitrarily due to the beam spot's vertical movement, light can always be detected at that position. This facilitates the selection and positioning of blood vessels to be measured, and makes it possible to measure the blood flow velocity of a single blood vessel using the speckle method while greatly increasing the amount of detected light.

It appears to be the same method as the conventional laser Doppler method in that it controls the position of the laser beam spot that irradiates the fundus and hits the target blood vessel.
There are significant differences both in terms of operation and principle. In the former case, it is difficult to make two divided laser beams intersect on the target blood vessel at a known intersection angle, or to check the incident beam angle and scattering detection angle for the target blood vessel even if there is only one beam. In contrast, the speckle method of the present invention does not require any troublesome restrictive operations related to incidence and detection angles, and can directly utilize the observation of the fundus camera, the incidence of the photographing light beam, and the light receiving optical path. can do.

Laser Doppler method measures the Doppler shift frequency according to the speed at which blood cells cross the interference fringes formed by crossing two divided laser beams on the target blood vessel,
Alternatively, the Doppler shift frequency can be measured by heterodyne detection by interfering the Doppler-shifted light generated by a single beam scattered by blood cells with light without Doppler shift from a stationary scatterer prepared separately. In contrast, the laser speckle method is a method for determining the blood flow velocity, whereas the laser speckle method is a method in which the light intensity distribution of the speckle pattern, which is created by the random phase overlap of light scattered by a large number of blood cells, changes over time according to the blood flow velocity. In this method, blood flow velocity is determined by measuring the degree of the change as a frequency component of the signal.

Conventionally, as an example of detecting the movement of a speckle pattern accompanying the six movements of an object using a multiple detection amount robot pattern having multiple minute apertures, the target is a situation where speckles are translated, and there are multiple There is a problem in that interference occurs between signals detected from each aperture, resulting in cross-correlation components. That is, there are two types of movement forms of speckle patterns: translational movement and boiling movement. Translational movement means that each image surface speckle 63 moves in translation in a fixed direction without changing its shape, as shown in Fig. 13. In other words, boiling motion is the movement of individual image plane speckles 6 as shown in Figure 14.
3 does not translate, but changes its shape while remaining in almost the same position, and the pattern of bright and dark speckles fluctuates randomly, as if disappearing and then emerging again.

Therefore, when multiple apertures are used for translational speckles, there is a situation where one speckle 63 crosses adjacent two-opening holes 65' and 65# or 65'' and 65~ as shown in FIG. 15(A). The probability increases, and a cross-correlation component between the two apertures occurs.In the correlation function of the output signal of the photodetector in FIG. 16(A), the original autocorrelation component 74 of each aperture alone is In addition, the cross-correlation component 7 between adjacent apertures
5, cross-correlation components 76 between every single aperture, etc. are superimposed. For example, when the spread of the correlation function is expressed as the delay time τ at the time when the correlation value attenuates to 1/e, as the correlation time,
In the case of a single aperture, the value would be τC, but with multiple apertures, the value would be C', resulting in undesirable effects. Therefore, in order to reduce the cross-correlation component, it is necessary to take measures such as making the intervals between multiple apertures sufficiently long so that they are uncorrelated with each other. The disadvantage is that a sufficient number of multiple apertures cannot be arranged, resulting in detection no different from conventional single aperture detection.

However, the inventors of this application have discovered from various experiments that the movement of the laser speckle pattern due to the scattered reflection of blood flow in the blood vessels of the fundus is a boiling movement. When multiple apertures are used for boiling motion speckle, Figure 15 (B)
As a result, the probability of a single speckle 63 crossing two adjacent apertures without changing its shape is extremely low.
The correlation function of the output signal of the photodetector is only the autocorrelation component 74 of each aperture, as shown in FIG. 16(B). This is exactly the same as the autocorrelation component 74 in the case of single aperture detection, and the correlation time τC can be determined correctly. As described above, by applying a multiple detection aperture pattern having a plurality of minute apertures to boiling motion speckles, it is possible to obtain the great advantage that exactly the same signal components as in the case of a single detection aperture can be obtained with an increased amount of light. In this case, since the autocorrelation component 74 does not depend on the number of minute apertures involved in detection, this is also a preferable characteristic in that errors due to variations in numerical apertures do not occur. Moreover, the fact that the speckle movement caused by the blood flow in the fundus of the eye is a boiling movement is also essential to the present invention.

Since there is no influence of cross-correlation, each micro-aperture 65' can be arranged relatively close to each other, and as a result, as shown in FIG. 17(A).
This is very practical because a large number of apertures 65' can be installed within the range of the speckle pattern light beam 62 to be detected. However, if the minute apertures 65' are too close together (so that they are too spaced apart) and one speckle 63 straddles the adjacent double aperture 65' as shown in FIG. 17(C), a cross-correlation component will occur. occurs, so Figure 17 (D
), it is necessary that the interval between each minute aperture 65' be at least larger than the speckle size, based on the average speckle size on the detection surface. Of course, the speckle size depends on the optical system conditions (magnification and F value), so it is important to use the average speckle size on the multiple detection aperture pattern as a reference.

Moreover, the spacing is too large, and as shown in FIG. 17(B), there is only one minute aperture 65' within the range of the speckle pattern light beam 62 (that is, the enlarged spot image 62 here) to be detected.
If there is no aperture, the conventional single detection aperture will be used. Therefore, the size of the range of the luminous flux should be considered so that at least two or more micro apertures 65' contributing to detection are included in the range. The spacing between each opening must be determined.

On the other hand, since the diameter of each minute aperture 65' results in the same signal component as in the case of a single detection aperture, point detection as in the conventional speckle detection method is preferable. However, in reality, it is necessary to obtain a certain amount of light with a finite size, and it is already known theoretically and experimentally that it is sufficient to use an aperture with a diameter smaller than the average size of the target speckle. . Therefore, in the present invention, the opening diameters of the respective minute openings 65' are all the same, for example, the 17th
As shown in Figures (A) and (D), it is necessary that the speckle size is equal to or smaller than the average speckle size of the detection surface.

If the aperture diameter is larger than the average speckle size, a plurality of speckles 63 exist within one aperture 65' as shown in FIG. As a result of averaging, the DC component increases while the signal component decreases, resulting in a decrease in the S/N ratio, which is not desirable in practice.

If multiple detection aperture patterns are arranged randomly, the output signal component of the photodetector will depend greatly on the randomness and numerical aperture of the random pattern to be fabricated, and therefore will not depend on the position where the speckle light flux on the multiple detection aperture pattern reaches. It is difficult to obtain the same signal components, which poses a major obstacle in terms of reproducibility and stability, and it cannot be said to be good in terms of product production. Furthermore, the uneven intensity distribution of the speckle light beam on the detection surface due to the random arrangement cannot be ignored. Therefore, this embodiment is characterized in that a plurality of minute apertures are regularly arranged. Examples of regular arrangement include a lattice arrangement in which the centers of the four-opening holes 65' adjacent to each other are located at the four vertices of a square, as shown in FIG. 18(A), and an arbitrary arrangement as shown in FIG. 18(B). A triangular arrangement is preferred in which the centers of the three-opening rods 65' that are adjacent to each other are located at three vertices of an equilateral triangle. Furthermore, instead of each aperture, a plurality of optical fibers having a core diameter equal to the front diameter can be used to form a plurality of detection aperture patterns on the entrance surface 77 of the optical fiber bundle 77' as shown in FIG. In this way, the photodetector 67 and the like can be installed remotely.

In this embodiment, as shown in FIG. 7(B), all speckle light intensity changes detected from the individual minute apertures 65' of the plurality of detection aperture patterns 65 are collected by the condenser lens 66 into one light beam. A detector 67 detects the light as a change in total light intensity. Output signal obtained from the follower (Fig. 8 (B)
) naturally differs instantaneously from the output signal when light is detected by each aperture alone (for example, FIG. 8(A)). However, by statistically processing it, exactly the same information can be obtained, and the multiple detection aperture pattern also has the various advantages mentioned above, such as an increase in the amount of light.

Therefore, in this embodiment, statistical processing of the output signal can be said to be a particularly important process. However, commonly used correlation processing and frequency analysis can be effectively used as processing methods, and are extremely advantageous in practice in that they do not require any special processing. As a specific example, the autocorrelation function and power spectrum of the output signal are measured.

As shown in FIGS. 20(A) and 20(B), when the blood flow is fast, the boiling motion of the speckles is also fast, and the time change in the detected light intensity is also fast, so the autocorrelation shows curve 78.
The power spectrum shows a curve 79, and conversely, when the blood flow is slow, the time change of the speckle light intensity is also slow.
Curve 78' for autocorrelation and 79' for power spectrum
becomes. In autocorrelation, the reciprocal of the correlation time τC is proportional to the speed, and in the power spectrum, the reciprocal of the correlation time
dB) at which the power is attenuated (which can be defined as the cutoff frequency fc) is proportional to the speed, so the corresponding speeds can be evaluated as shown in FIGS. 21(A) and 21(B), respectively. The proportionality coefficient in this case is strongly dependent on the scattering object, and should be calibrated in advance based on other blood flow measurement methods (e.g., fluorescence fundus photography) or by flowing blood through a glass tube equivalent to a blood vessel. etc., it is possible to determine a value that is almost reliable. Even if the proportionality coefficient contains some error from the true value, this method is extremely superior in terms of data reproducibility and stability, and is extremely useful in clinical applications.

In the above example, it is naturally necessary to adjust the light intensity of the laser beam spot irradiated to the fundus, but switching the light intensity from positioning before measurement to during measurement is not possible manually. It is difficult to do so. Furthermore, since the amount of light for measurement differs for each individual due to the effects of fundus reflectance, cloudiness of the crystalline lens, etc., it is necessary to switch between several types.

However, this can be easily done with one mechanism by starting the measurement,
Not suitable for high-speed switching at the end. Therefore, in this embodiment, this problem is solved by using two light amount adjustment filters 2 and 3, each of which performs each function independently.

That is, the laser light from the laser light source 1 passes through the first light amount adjustment filter 2 as shown in FIGS. 1 and 22. The filter 2 is, for example, a switch type that rotates at a constant angle around the shaft 2', and is operated by a solenoid 83 or the like that is linked to the measurement switch 49. Measurement switch 49
When it is OFF, the filter section 2a is selected as the filter, and the amount of light is reduced to a relatively weak level that can be recognized by the measurer. When the switch is turned on, filter section 2
b is selected, and the amount of light necessary for measurement is set to be strong with some margin. Of course, the filter returns to the filter section 2a at the same time as the measurement ends. This allows instantaneous settings at the start and end of measurement.

Next, the light that passed through filter 2 continues to filter 3
pass through. The filter 3 can be switched in six stages from filter section 3a to filter section 3f, for example, by rotating about shaft 3'. This switching is performed manually, and in both cases, the amount of light required for measurement can be adjusted finely depending on the object to be measured, in combination with the filter section 2b of the filter 2.

Furthermore, as mentioned above, the range of the laser beam spot is limited within the visual field range 71, but the movable range 71 of the laser beam spot within the fundus visual field is inconvenient for operation unless the operator is informed of it in some way. be. Therefore, in this embodiment, the pattern of the reticle 35, which is an index for correcting the diopter of the observer in FIG. A double circle pattern 81 is added, and this double circle indicates the movable range of the laser beam spot within the field of view 82. Furthermore, by matching the focusing conditions of this double circle, it can be used as an index for diopter correction, which is very convenient in practice.

In order to accurately irradiate the laser beam toward the blood vessels in the fundus, it is preferable that the fundus be seen as enlarged as much as possible, but on the other hand, in order to see some of the various blood vessel conditions on the fundus, it is necessary to have a wide angle of view. It's better. Therefore, it is necessary to be able to switch the viewing angle depending on the purpose. In a normal fundus camera, a lens for changing the angle of view is prepared in the light receiving optical system, and the lens is switched in stages, but in this embodiment, the speckle pattern is detected through the light receiving optical system, so the angle of view is changed as the angle of view is changed. This is disadvantageous because the speckle detection conditions change depending on the change in imaging magnification.

Therefore, the above objective is achieved by using a zoom type optical system as the observation eyepiece (eyepiece) 36 shown in FIG.

In this way, it is possible to continuously change the viewing field magnification without changing the conditions of the speckle detection optical system. At this time, the minimum magnification can be determined according to the angle of view of the standard observation field of the fundus camera, as shown in Figure 23 (B), and the maximum magnification can be determined at the maximum magnification as shown in Figure 23 (C). However, the angle of view may be determined in such a way that the double circle 81 of the reticle, which indicates the eight movable range of the fundus irradiation laser beam spot, is included in the field of view 82'. This results in
Regardless of the magnification or standard of the zoom eyepiece, the daytime adjustment of the laser beam spot can always be performed while observing within the field of view.

Conventionally, when installing a wavelength separation mirror to detect laser speckle light from the fundus using a fundus camera, etc., a wavelength separation mirror that reflects only the wavelength component of the laser speckle light during measurement is used to observe and photograph the fundus. A commonly used method is to insert the light into the optical path of the light-receiving optical system.

However, in this embodiment, which has the feature that the entire range of the speckle light flux is effectively used for light detection with a multi-detection quantity robot pattern, the optical axis of the speckle light flux detection optical system is particularly accurately defined, and it is difficult to perform measurement operations. However, stability is required to prevent the wavelength separation mirror from shifting. Therefore,
A wavelength separation mirror insertion/removal method that requires mechanical switching each time it is operated is insufficient to meet this requirement.

In other words, in the case of flipping up, the vibration of the wavelength separation mirror when it is flipped up becomes the vibration of the propagation optical path of the speckle beam, and appears as a fluctuation of the speckle pattern on the detection surface, resulting in a speckle signal. This results in noise. In particular, this effect is significant because an enlarged imaging system is used to detect the speckle light flux. Moreover, the reproducibility of the position when it is flipped up and fixed is not sufficient. Furthermore, since the wavelength separation mirror that defines the optical axis of the speckle light flux detection optical system is movable, it is difficult to perform sufficient axis alignment and adjustment during manufacturing, and there is also the problem that each product varies. be.

Furthermore, there remains the concern that if the product is used many times and its durability deteriorates over time, it may become mechanically unstable.

Therefore, as already mentioned in Fig. 1, the main function of this device is speckle light flux detection, and the wavelength separation mirror 39 is fixedly installed, and no removable mirror is installed in the speckle light flux propagation optical path. It is structured. Observation and photography are performed via an insert/remove mirror 30 installed in front of the wavelength separation mirror 39, and when this mirror flips up during measurement and leaves the optical path, a wavelength separation mirror 30, which is precisely adjusted and fixed in advance, is placed behind it. The speckle light flux can be correctly sent to the detection optical system via the mirror 39. Because it is fixedly installed, the optical axis of the speckle beam remains stable at all times, allowing for good measurements.

In addition, if the operation of the insertion/removal mirror 30 for observation and photographing at the start and end of measurement is done manually, it will not be carried out quickly and smoothly.
It can also be expected that there will be human error. Since the measurement is carried out in a short period of time, it is undesirable that there is wasted time in the operation of the insertion/removal mirror 30, and it is also susceptible to the effects of displacement of the measurement site due to eye movement.

Therefore, as shown in FIG.
The above problem can be solved by removing the light from the optical path and inserting it into the optical path when the switch 49 is OFF. At this time, the other insertion/removal mirror 34 is also linked to the measurement switch, and when the switch 49 is turned on during measurement, the mirror 34 leaves the optical path, and the wavelength component light that has passed through the wavelength separation mirror 39 contributes to observation. do. When not measuring,
Since the switch 49 is in the OFF state, the mirror 34 is inserted into the optical path, and as a result, the light arriving via the lens 33 is reflected, contributing to observation. Therefore, fundus observation is possible during measurement.

Furthermore, since measurement usually starts and ends with the consciousness of the person measuring it,
The subject, other signal analyzers, and measurement collaborators are not informed of the exact timing of the start and end of the measurement, and only have a rough idea of when the measurement will start and end. Particularly when the measurement is performed in less than one second, it may end without you noticing. The patient's cooperation is required to maintain a quiet posture during the measurement, and the signal analyzer may not be able to determine whether the input signal is the correct signal being measured or some other noise. is possible.

Also, there is insufficient coordination among surrounding measurement collaborators.

Therefore, in this embodiment, the measurement switch is turned on and the measurement starts, or twice at the start and then again when the measurement switch is turned off and the measurement ends, or continuously from the start of the measurement and at the end of the measurement. Until then, an electronic beep will sound to notify each measurement. by this,
It becomes possible to convey the exact timing of the start and end of measurement to the subject and the people around them, which solves the above problems, making it a very practical device.

Signal Processing Using Photon Correlation Method FIG. 24 shows an embodiment in which the photon correlation method is used for signal analysis. After the signal from the photodetector 45 is input to a photon counting unit 91, it is amplified by a multiplier tube 92, and the correlation is examined by a correlator 93. After that, it is processed by the microcomputer 94, and the result is displayed on the CRT95.
Alternatively, it is output to the printer 96.

In this embodiment, the configuration shown in FIG. 25(B) is used. That is, all speckle light intensity changes detected from the individual apertures 65' of the multiple detection aperture pattern 65 are collected by the condenser lens 66, and one photodetector 67 detects the photon correlation as a total light intensity change. . Naturally, the output photoelectron pulse signal obtained from the photon counting unit and (PCU) is instantaneously different from the output photoelectron pulse signal when light is detected by each aperture alone.

Moreover, by statistically processing it, exactly the same information can be obtained as a photon correlation function.

When using the photon correlation method, it is possible to obtain the same effect as obtaining autocorrelation of light intensity using a plurality of detection aperture patterns.

That is, when a single detection aperture 64 is used as shown in FIG. 25(A), a long measurement time is required to sufficiently converge and obtain stable data (FIG. 26(A)).
As shown in FIG. 25(B), a plurality of minute openings 65'
When detected with the multiple detection aperture pattern 65 having
The photon correlation curve data becomes data as shown in FIG. 26(B), which improves the convergence stability of the plot, improves the measurement accuracy by increasing the amount of detected light, and shortens the measurement time.

The accuracy of signal processing using the photon correlation method depends on the integrated light amount determined by the detected light amount and detection time, so increasing the detected light amount as in this example makes it possible to shorten the detection time and obtain extremely favorable results. It will be done.

When the blood flow is fast, the speckle boiling motion is fast and the detected light intensity changes quickly over time. Therefore, the photon correlation function is similar to the autocorrelation function curve 78 shown in FIG.
If the attenuation is slow, it will be curve 7.
8', the attenuation is relatively gentle.

Since the reciprocal of the correlation time τC is proportional to the speed, the second
The corresponding speed can be evaluated as illustrated in FIG. 1(A).

Also, in this example, if the micro apertures of the multiple detection aperture pattern are arranged randomly, the correlation time of the photon correlation data largely depends on the randomness and numerical aperture of the random pattern to be manufactured, so the specifications on the multiple detection aperture pattern It is difficult to obtain the same signal components regardless of the location where the light speed reaches, which is a major obstacle in terms of reproducibility and stability, and it cannot be said to be good in terms of product production. Furthermore, the uneven intensity distribution of the speckle light beam on the detection surface due to the random arrangement cannot be ignored. Therefore, in the embodiment, a plurality of micro-apertures are arranged regularly.

[Effects of the Invention] As explained above, in the present invention, the spot diameter of the laser beam can be adjusted, and the spot position can be moved within a predetermined viewing angle range narrower than the viewing angle.
It becomes possible to measure the blood flow velocity of a specific blood vessel with ease and accuracy.

[Brief explanation of the drawing]

Fig. 1 is a block diagram showing the overall configuration of the device of the present invention, Fig. 2 is an explanatory diagram showing the structure of the ring slit used in the device shown in Fig. 1, and Fig. 3 is a characteristic diagram of the filter used in the device shown in Fig. 1. FIG. 4 is a block diagram showing the configuration of the signal processing section, FIG. 5 is an explanatory diagram showing the irradiation state of the fundus blood vessels, FIG. 6 is an explanatory diagram showing an enlarged image of the measuring section, and FIG. 7 (A). (B) is a perspective view showing the configuration of the signal detection section, FIG. 8 (A)
, (B) is a waveform diagram showing the signal waveform obtained from FIGS. 7(A) and (B), FIG. 9 is an explanatory diagram explaining the state of irradiation of blood vessels with a laser beam, and FIG. 10 is an observation and photographing field of view. FIG. 11 is an explanatory diagram showing a state in which the observation and photographing field of view is restricted, and FIGS. 12 (A) and (B) are a perspective view and a plan view showing speckles on a conjugate magnified image plane, respectively.
FIGS. 13 and 14 are explanatory diagrams showing the translational movement and boiling movement of speckles, and FIGS. 15(A) and (B) are explanatory diagrams showing passage through the detection aperture due to the translational movement and boiling movement of speckles, Figures 16 (A) and (B) are characteristic diagrams showing intensity changes due to passage through the detection aperture, Figures 17 (A) -
(E) is an explanatory diagram showing the relationship between speckle and detection aperture;
Figures 1a (A) and (B) are explanatory diagrams showing the pattern arrangement of the detection aperture, Figure 19 is a perspective view showing another configuration of the detection aperture, and Figure 20 (A) and (B) are speckles. Characteristic diagrams showing the autocorrelation function and power spectrum of the signal, Figures 21 (A) and (B) are diagrams showing how to determine the speed from the characteristics of Figures 20 (A) and (B), and Figure 22 The figure is a perspective view showing the mechanism for adjusting the light amount, Figures 23 (A) to (C) are explanatory diagrams showing the field of view image on the reticle, and Figure 24 is the signal analysis section when using the photon correlation method. A block diagram showing the structure, Figures 25 (A) and (B) are perspective views showing light amount detection using the photon correlation method, and Figures 26 (A) and (B) show the characteristics of photon correlation using the photon correlation method. This is the characteristic diagram shown. 1... Laser light source 9... Movable mirror 18...
・Eye to be examined 42...Detection aperture 50...Analysis section 63...Speckle 65...Detection aperture Fig. 7 Fig. 8 Fig. 11 (A) (B) (A) CB) Fig. 21 Figure 24 Figure 25

Claims (1)

[Claims] 1) The fundus of the eye is irradiated with a laser beam of a predetermined diameter, and the movement of the laser speckle pattern formed on the observation surface by the scattered reflected light from the fundus tissue is detected as a change in the intensity of the speckle light. Based on the analysis results of the speckle signal obtained by
An ophthalmologic diagnostic device that measures the blood flow state of fundus tissue includes a laser light source and a laser beam from the laser light source that irradiates a target blood vessel to be measured with a minute beam spot diameter that is the same as or smaller than the diameter of the blood vessel. an optical system, an optical system that condenses light scattered and reflected from the fundus to form a same-sized or enlarged image of the irradiated area on an image plane conjugate to the fundus, and an optical system arranged on the conjugate image plane; a plurality of detection aperture patterns having a plurality of microscopic apertures, and a boiling movement of speckles occurring on the image plane is detected through the plurality of detection aperture patterns, and the method is configured to detect boiling motion of speckles occurring on the image plane according to intensity changes of the total amount of light passing through the plurality of detection aperture patterns. means for measuring the blood flow velocity of the target blood vessel by analyzing and evaluating the speckle signal obtained, the laser beam spot diameter is adjustable, and the laser beam spot position is adjusted to a predetermined field of view narrower than the observation field angle. An ophthalmologic diagnostic device characterized by being movable within an angular range. 2) A predetermined pattern indicating the movable angle of view range is displayed on a reticle pattern for diopter correction, and the predetermined pattern also enables diopter correction. The ophthalmological diagnostic device described in section. 3) The fundus observation eyepiece is a zoom type that can continuously change magnification, its minimum magnification is determined to match the standard observation angle of view of the fundus camera, and its maximum magnification is determined by the laser that irradiates at least the fundus at maximum magnification. Claim 1, characterized in that the movable range of the beam spot is determined according to the angle of view that can be included within the field of view of the eyepiece.
The device according to paragraph 2 or paragraph 2. 4) Adjust the light intensity of the laser beam spot using first and second filters, and set the first filter to the light intensity necessary for measurement when the measurement switch is on, and to a low light intensity when it is off. and the second filter
4. The device according to claim 1, wherein the amount of light passing through the filter can be further adjusted.