JPS62275431A - Ophthalmic diagnostic method and apparatus - Google Patents

Abstract

(57) [Summary] This bulletin contains application data before electronic filing, so abstract data is not recorded.

Description

[Detailed Description of the Invention] 3. Detailed Description of the Invention [Field of Industrial Application] The present invention provides an ophthalmological diagnosis method and apparatus, and more specifically, uses a laser speckle phenomenon to diagnose the fundus and iris of the ophthalmic region. The present invention relates to a method and apparatus for measuring the state of blood flow in tissues at various sites.

[Prior Art] 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 the light intensity at the observation point, the movement of the object can be measured from the degree of signal change. The invention applies this to the measurement of the state of blood flow in tissues such as the fundus of the eye.

Conventionally, the laser Doppler method has been used as a method to measure the blood flow state of the fundus of the eye.
LI'ED 0PTIfll:S) [Optical Society' Op America (Optical 5o
Society of America) Vol. 20, No. 11, No. 1 (January 1981), pages 117-120. This method uses a fundus camera to irradiate laser light onto blood vessels in the fundus, and the laser light is scattered by blood cells flowing through it, which undergoes a frequency shift proportional to the blood cell velocity due to the Doppler effect. The principle is to determine blood flow velocity by measuring the amount of deviation in optical frequency.

[Problems to be solved by the invention] However, in this method, since it is necessary to detect the Doppler shift frequency, the incident laser beam is divided into two parts at equal angles to the optical axis and guided to the eye to be examined. The optical system configuration is extremely complex and requires a configuration that accurately intersects at the target blood vessel position, or conversely, a configuration that extracts the laser light scattered by the fundus blood cells from two different directions and detects the light. is required. Furthermore, the angle of incidence or detection angle must be known, and the target blood vessel must be accurately irradiated with a laser beam focused to the same extent as the diameter of the blood vessel (usually several tens to hundreds of gmφ). Considering that the subject must be kept still during the measurement period, it is extremely difficult to handle clinically, and it is difficult to obtain reproducible and reliable 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 has to enter from a direction close to perpendicular to the fundus, which can cause the Doppler effect. The difficulty lies in the fact that the method detects a single heat component, and in biological tissue that causes various types of irregular light interference, laser spectroscopy, which is itself an irregular scattering interference effect of light, is difficult. It is preferable to apply the method.

On the other hand, as a blood flow meter using a laser Svenkle, conventionally,
No. 235 and JP-A No. 80-203238. This method uses an optical fiber to irradiate the surface of living tissue with laser light, and guides the light scattered by blood cells in the surface MI tissue again to a photomultiplier tube using another optical fiber.The frequency of speckle No. 43 detected is The principle is to output I[pressure proportional to Al1 or to determine the frequency gradient of the same speckle signal to determine the blood flow state.

However, although this blood flow meter can be applied to areas where laser beam irradiation can be easily detected, such as living tissues such as the skin 1a, it is difficult to apply the blood flow meter to areas that can be easily detected by laser beam irradiation, such as the skin 1a. cannot be easily applied to objects that are

Furthermore, it is necessary to employ a special probe to irradiate the tip of the optical fiber to a fixed position, and due to the shape of this probe itself, the situation of the measurement site where the optical fiber inside the probe is actually irradiated with laser light is required. or its position cannot be visually observed accurately.

In addition, since the light emitted from an optical fiber spreads out, there are planets where the emitting end is placed close to the target biological tissue.
These also pose great difficulties in measuring blood flow in the eye region.

Therefore, the present invention has been made to solve these conventional drawbacks, and provides an ophthalmological diagnostic method that can accurately and efficiently measure the blood flow state of various regional tissues in the ophthalmological field using laser speckle imaging. and related to equipment.

[Means for Solving the Problems] In order to solve these problems, the present invention irradiates the eye region with a laser beam of a predetermined diameter, and the laser beam is formed by diffusely reflected light from the blood cells of the biological tissue. The movement of the laser speckle pattern is detected as a change in speckle light intensity at an observation point, and the shape of the power spectrum distribution of the obtained speckle signal is evaluated from the calculation result of the average frequency to measure the blood flow state of the living tissue. The configuration was adopted.

[Function] According to such a configuration, instead of measuring the absolute speed of blood flow flowing through a single blood vessel, the blood flow state of multiple blood vessels included in a limited irradiation area is measured as a whole and on average. This improves the ease of handling the device 6 and improves the reliability and stability of the measurement results.

[Example] The present invention will be described in detail below according to an example shown in the drawings.

In the following example, ophthalmological diagnosis will be explained using a case where fundus blood flow is measured using a fundus camera, but the present invention is not limited to this, and can be applied to various ophthalmological diagnoses. It is.

Vertical Difference j FIG. 1 is a schematic diagram of the entire apparatus for carrying out the measuring method according to the present invention. A laser beam from a laser light source I such as He-Ne or argon passes through a light adjustment filter 2 for adjusting the light intensity via a codenser lens 1'. Furthermore, it is guided to the fundus illumination optical system of the fundus camera via a relay lens 3.4. Also relay lenses 3 and 4
Apertures 5 and 6 are installed between them, and are used to select the size and shape of the laser beam irradiation area on the fundus. Further, there is a shutter 7 at the exit of the laser light source 1, which opens and closes as necessary. The laser beam guided by the relay lens 4 passes through the annular opening 8a of the ring slit 8 in the fundus illumination optical system, as shown in FIG.
It is reflected by a mirror 9 installed in a part of the fundus, and is guided to the same optical path 1 through which the fundus observation and photographing light beam enters the fundus. For this reason, the laser beam is reflected by the perforated mirror 12 via the relay lens 10.11, focused once on the cornea 13a of the ophthalmologist 3 through the objective lens 13', and then diffused into the fundus of the eye. L3b is reached, forming an irradiation area that is wider than the blood vessel diameter as described above.

This irradiation area is used as a fundus camera! ! It is illuminated by beam optics to facilitate observation. This observation optical system includes an observation light source 22 arranged on the same optical axis as the photographing light source 24, a condenser lens 23, a condenser lens 25, a filter 27, and a mirror 26. Since the laser beam is placed on the same optical path as this observation and photographing light beam, the laser beam is directed to point 13b of the fundus using the left and right, up and down swing a structure and fixation guidance mechanism of the fundus camera.
It can be irradiated to the position of

Note that the filter 27 disposed between the codenser lens 25 and the mirror 26 is configured as a #wavelength filter as shown in FIG.

The red component contained in the observation and photographing light is canted.

Speckle light generated when the laser light is scattered by blood cells moving in the fundus blood vessels is received again by the objective lens 13', passes through the perforated mirror 12, and reaches the photographing lens 14 and the wavelength a mirror 15. This wavelength separation mirror 15
Like the filter 27, it has the spectral characteristics as shown in FIG. ! He-H
The speckle light (red) generated by e-laser light is
Most of it is reflected. This reflected light is transmitted to the beam splitter.
16, relay lens 17, and detection apertures 18 and 19, are collected again by a condenser lens 20, and detected by a photomultiplier tube (photomultiplier) 40. A shunter 21 is arranged in front of the photomultiplier tube 40, and the output signal from it obtained when it is open passes through an amplifier 41 and a filter 42, and is then inputted to a non-processing circuit 50 (FI).

This signal processing circuit 50 includes a CPU as shown in FIG.
51. Memory 52. It is composed of an analog-to-digital converter 53, a keyboard 54, a CRT 55, and a blink 56.

Further, the light passing through the wavelength separation mirror 15 can be observed through a relay lens 28, a flip-up mirror 29, a mirror 30, and an eyepiece 31, and can be photographed with a photographic film 32.

Further, a portion of the speckle light reflected by the beam splitter 16 shown in Fig. m1 passes through a relay lens 33, a mirror 34, and an eyepiece 35 so that the ITII person can observe the speckle pattern.

Next, the operation of the present invention configured as described above is illustrated in FIGS. 5 and 6.
This will be explained using the flowchart shown in the figure.

First, after turning on the power in step SL in FIG. 5, the subject is set in step S2, and the fundus 13b of the subject's eye 13 is observed through the observation optical systems 22 to 26 (step S3).
, following the force switching in step S4, the laser f-light source 1 is activated in step S5. At this time light ff151
Set the output level to the adjustment level with filter 2, set the size and shape of the laser irradiation area with aperture 5.6, open shutter 7 in step S6, set the measurement position, and then create a speckle pattern. Confirm (step S7.
Step S8), then set the size and shape of the detection apertures 18 and 19. Step S9). When measuring, set the laser output to the measurement level and close the shutter 2 of the photomultiplier tube 40.
1 and adjust the input sensitivity (step 5IO-S
12). Subsequently, the No. 48 processing parameters are set in step 513, and measurement is started in step S14.

During measurement, the output from the photomultiplier tube 40 becomes a speckle signal that changes over time as the blood cells move. The speckle signal is amplified by an amplifier 41.

Unnecessary frequency components are removed through a band filter 42 whose band is set as necessary. This output is sent to a microcomputer-signal processing circuit 50. Here, as shown in FIG. 4, the speckle signal is converted into a digital signal by the A/D converter 53 in the signal processing circuit 50, and then subjected to frequency analysis by running an analysis program prepared in advance. A power spectral distribution is obtained. This result is displayed once on the display 55 and the distribution shape is observed (step S15).If the movement of blood cells is fast, the detected speckle light intensity changes quickly, so the power spectrum of the speckle signal is extends to the higher frequency side. In the opposite case, the power spectrum mainly has many low frequency components. Therefore, by evaluating the shape of the power spectrum distribution, the +m flow state can be measured.

In the prior art, there is a method of examining the wave number gradient as one way of evaluating this. This can be applied to relatively simple components 4j such as those shown in FIGS. 7(a) and 7(b).

In examples like (C) to (e), only the gradient? It is difficult to make a power evaluation, and the conventional technology that focuses on a specific part of the power spectrum cannot perform a sufficient evaluation. In the present invention, the so-called general average frequency <f> defined by the following equation is adopted in consideration of the case where the power spectrum is complicatedly distributed.

Here, F is the frequency and P (f) is the power spectrum. In addition, in order to clearly show the difference in the shape of the power spectrum distribution, the following average frequency <fe> is also calculated using the logarithm of the power spectrum.

Here, log, oP (f) ll1n is the logarithm of the minimum value of P (f) in the data.

The calculations of <f> and <fe> are performed according to the flowchart shown in FIG. In steps T1 to T4, the data is loaded into the memory 52, fast Fourier transform is performed under the control of the CPU 51, and the power spectrum is converted to the CRT5.
5 and stores the data in the memory 52. When determining <f> in step T5, steps T7 to T
Calculate <f> in LO and display it on the CRT 55, and if <f) is not determined, reset in step T6 and return to step T1.

Next, when finding <fe> in step TIL,
<fe> is calculated in T12 to T15 and displayed on the CRT 55 in step TL6.

If <fe> is not determined in step Tit, and f> is determined in step TL7 after step T16.

<fe> is stored in the memory 52, and the result is output to the printer 56 in step T18. This corresponds to steps S16 to S18 in FIG.

In FIG. 5, it is determined in step $19 whether or not to take a photograph of the output result.

If the power spectrum during calculation of <f> and <fe> contains more components on the high frequency side, the modulus frequencies <f> and <fe> will be high, and in the opposite case they will be low. , The blood flow condition is measured from this result. <
The calculations of f > and < f e > are normalized by the total power amount, so they are not affected by changes in the overall light intensity during measurement.The above signal processing can be performed in a relatively short time, so the blood flow state It is possible to measure changes over time. Furthermore, since the power spectrum information of the entire area is reflected in the calculation, the state of blood flow can be measured with good stability.

In the present invention, the laser beam irradiation area at the measurement site of the eye to be examined, such as the fundus 13b, is set to be a wider area than the blood vessels, for example, 1 to 3 mm in diameter, so that a plurality of blood vessels are included in this area. ing. Therefore, unlike the prior art, there is no need to irradiate a specific blood vessel with a laser beam as small as the diameter of the blood vessel, and there is no need to divide the irradiated laser beam into two, so irradiation can be easily performed. In addition, even if the eye 13 to be examined moves and the measurement position shifts slightly during δ14 measurement, it will not have a major effect on the measurement results. Because the area of the vascular tissue still existing within the laser beam irradiation area is small compared to the area of the vascular tissue that still exists within the laser beam irradiation area, the speckle light from the vascular portion entering and exiting the irradiation area is scattered from the entire tissue. The influence on changes in the intensity of speckle light is reduced. This is because the measurement area can be equivalently considered to be constant.

There is also no need to extract scattered light from multiple different directions.
Since measurements can be made regardless of the angle of incidence or angle of acceptance, reproducible and reliable results can be easily obtained. This point is a major feature of the present invention.

The basic principle of this method is not to measure the absolute velocity of blood flowing through a single blood vessel, but to evaluate the overall and average blood flow status of multiple blood vessels included in the irradiation area. . In other words, by detecting the intensity fluctuations of the speckle pattern that are contributed by all scatterers in a wide laser beam irradiation area, it is possible to measure the average state of each blood flow state in the laser beam irradiation area. It is possible. This method of measuring the blood flow state by averaging over a wide range is a blood flow measurement method for month diagnosis that is different from the conventional technology. We aim to improve reliability and stability. In particular, capillaries and the like have small diameters, so this method is very effective in evaluating the overall state of blood flow compared to conventional techniques.

Note that if the irradiation area is made larger than necessary, there will be practical inconveniences such as difficulty in individually irradiating different vascular tissue areas such as the papilla and its surrounding areas in the fundus, and the detection aperture size will increase. Speckle particles formed on the fundus become very small, which is undesirable because it causes problems such as difficulty in extracting changes in light intensity.The size of speckle particles is determined by , also varies depending on the pupil diameter of the eye 13 to be examined. If the detection aperture 18.19 is too large compared to the speckle particle diameter,
), a large number of speckle particles 18a are integrated within the aperture, so it is not possible to extract the light intensity changes due to individual speckle particles well. Measurement becomes impossible because sufficient light reception cannot be obtained. A good example of a combination of laser beam irradiation area diameter and detection aperture diameter is, for example, when the former is 3 mm, the latter is 50 tiles,
As another example, when the former is 1■Φ, the latter is 70~1
00gmφ etc. depending on the irradiation area diameter at any time
This is due to the fundamental property of the laser speckle phenomenon that the diameter of the speckle particles generated on the surface changes.

The speckle pattern 18b obtained from the fundus 13b is
Although observation is also made on an image plane conjugate with the fundus 13b, in the present invention, in order to detect the contribution of scattered light from the entire region irradiated with laser light, a detection aperture is basically provided on the Fraunhofer diffraction surface for the fundus 13b. 18.19 is placed to detect speckle light. This is because, as is well known, the light at one point on the Fraunhofer diffraction surface is a superposition of scattered light from each point within the desired irradiation area.

Note that the measuring device according to the present invention can measure the blood flow state while also having the function of a fundus camera. The light emitted from the observation light source 22 in FIG. 1' is transmitted to the condenser lens 23 along the optical path shown in FIG. Photography light source 24.
It is reflected by a mirror 26 via a condenser lens 25,
With the conventional illumination optical system of the fundus camera, the fundus 1
3b is illuminated.

Furthermore, in the embodiment shown in FIG. 1, by inserting a wavelength separation filter 27 having spectral characteristics as shown in Em3UgJ between the condenser lens 25 and the mirror 26, the red light component contained in the observation and photographing light is removed. Therefore, the red light reflected by the wavelength separation mirror 15 and detected by the photomultiplier tube 40 becomes only the laser beam for measurement, and the wavelength separation mirror prevents the observation and photographing light from affecting the speckle light. The observation light transmitted through relay lens 28.15 is transmitted through relay lens 28.15 as shown in FIG. Flip-up mirror 29. mirror 30
It can be observed simultaneously with the measurement through the vi-ocular lens 3L, and if necessary, a photograph can be taken on the photographic film 32 using the photographing light that has also passed through the wavelength separation mirror 15. At this time, the flip-up mirror 29 is flipped up to 29' in the direction of the arrow using 29a as a fulcrum. Wavelength separation mirror 15
As shown in Fig. 3, since a small amount of red laser light also passes through, the laser light irradiation position on the fundus 13b can be sufficiently visually observed or photographed and recorded, so it is effective for recording measurement results. In consideration of the case where the wavelength separation mirror 15 is used only as a fundus camera, the wavelength separation mirror 15 can also be fixed by pushing it upward to 15' using the fulcrum <15a as indicated by the arrow in FIG.

In the embodiment, with a fundus camera, it is usually possible to observe and photograph the anterior segment of the eye by adjusting the photographic lens 14 shown in FIG. The state of blood flow in the eye can be measured according to the above example. Here, it can be applied to areas where vascular tissues are relatively exposed on the surface, such as the iris and sclera. As there have been few specific devices for measuring the state of blood flow in the iris, etc., this method has great significance, and is useful for diagnosing diseases of the anterior segment of the eye, medicinal efficacy, surgical effects, etc. This is very useful for examinations, and being able to measure blood flow conditions in various parts of the eye, such as the fundus and iris, under the same conditions for the same eye to be examined, and being able to compare the results with each other is very useful clinically, and is one of the features of the present invention. It is one.

Note that step S5 in FIG. S9. It is also possible to start after setting the laser light intensity, the size and shape of the laser irradiation area, the size and shape of the detection aperture 18, 19, signal processing parameters, etc. in S13 that are known in advance. Also, photography is not only done at the final stage.

It may be performed at any time and under various conditions. The printout shows the power spectrum and average frequency <f>, <f
e> can be performed all or selectively.

Practical Examples In addition to the embodiments of the present invention described above, various improvements and modifications can be made, which will be described below.

As shown in Fig. 11, detection apertures 18, 19, condensing lens 2
The speckle detection light obtained through the optical fiber 10
1, and the emitted light is detected by the photomultiplier tube 40 through the lens 102, thereby photomultiplying 'I? 4
0 can be installed separately from the main body, making it possible to commercialize the main body. This embodiment also allows for a mechanically stable structure.

Furthermore, as shown in FIG. 12, the speckle light detected in the basic embodiment is optically detected by a photoelectron multiplier 1740, and photon counting is performed by a photon counting unit 60.

After passing this output through a filter 42, it is sent to a microcomputer-signal processing circuit 50 to obtain a photon correlation output. If the correlation length of this correlation curve is determined, the phase IA length will be short if the blood flow is fast, and will be long if the blood flow is fast, so by examining this, the blood flow state can be measured. The correlation output is calculated according to, for example, a 7'o-gram prepared as shown in the flowchart of FIG. 13, for example. This control is almost the same as the control flow in FIG. 6, and instead of the power spectrum, steps R1-R4 are used.
As shown in , the difference is in how the photon correlation output is determined.

This photon correlation method requires high accuracy and is effective when the intensity of the laser speckle light from the fundus is so weak that the above-described signal processing method as shown in FIG. 6 cannot be used.

Further, as in the example shown in FIG. 11, when detecting through the optical fiber 101, the core diameter of the optical fiber is generally 10p.
A single-mode type with a diameter of approximately mφ is preferable, and a multi-mode type with a large core diameter increases mode noise due to interference between modes and time delay, making it impossible to measure speckle light intensity.

When an optical fiber with a core diameter of about 10 gmφ is used, the detected light intensity becomes weak. Therefore, even when the optical fiber IO1 is used for detection, it is effective to use the photon correlation method as signal processing as shown in FIGS. 12 and 13.

Furthermore, as the laser light source 1, Ar (
Argon) lasers can also be used. In this case, the spectral characteristics of the wavelength separation mirror 15 and the wavelength separation filter 27 in Fig. 1 are as shown in Fig. 14 (a) when the wavelength of the light source is 488 nm;
).

Furthermore, as shown in FIG. 1, the same effect can be obtained by changing the position of the wavelength separation mirror 15 by using an up-and-down sliding method instead of the flip-up type as shown in FIG. 15 for the wavelength separation mirror 15.

Further, the arrangement of the observation light source 22 and the photographing light source 24 is as shown in FIG. The condenser lens 103 is used as a photographing light source 24. Condenser lens 104 and mirror 105
They may be placed on different optical paths with the two sides in between. In the latter case, the position of the mirror 105 is changed from the time of visual measurement of the fundus 13b and the measurement of the fundus blood flow state, and the position of the mirror 105 is changed from 105 when taking the fundus photograph.
′ enables both functions.

Further, as shown in FIGS. 17 and 18, the light from the observation light source 22 is transmitted to the condenser lens 106. Imageify, <
1 o 7 , a microlens 108 may be used to guide the light to the condenser lens 103, in which case the observation light source 22
By placing the fan for cooling the heat of the observation light source 22 in a separate location from the conventional fundus camera body, the main body of this device can be made smaller and lighter.
This is considered to be preferable in terms of design in the sense that it reduces the intimidating feeling of the measuring machine to the subject.

In the embodiment shown in FIG. 1, a light amount adjusting filter 2 is used to adjust the amount of laser light emitted by the laser light source 1. However, if a linearly polarized laser is used as the laser light source 1, the light amount can be adjusted. By arranging a polarizing plate as the filter 2 and rotating the polarizing plate in a plane perpendicular to the optical axis, the amount of laser light can be adjusted. Of course, a linearly polarized laser may be used as the laser light source 1, and the amount of light may be adjusted using an ND filter or the like.

Further, the light from the laser light source l is guided to an optical fiber 110 through a lens 109 as shown in FIG. You can guide me. According to this, the laser light source 1 can be separated from the main body at the stage 2, which is preferable because the main body can be made smaller and lighter.

Furthermore, FIGS. 20 to 22 show an aperture 5 that determines the size and shape of the laser beam irradiation area in the embodiment shown in FIG.
, 6, or various mechanisms for varying the size and shape of the detection apertures 18 and 19 are illustrated.

FIG. 20 shows a desired shape and size on the circumference of a circle centered on the center of a circular plate 70 (it may be a fan-shaped plate or a polygonal plate instead of a circular plate, but for convenience of explanation, a circular plate is used). A circular plate 70 is installed in a plane perpendicular to the optical axis so that the opening 70a intersects with the optical axis, and the circular plate 70 is rotated about the center of the circular plate. In this way, a method will be described which makes it possible to select an open lower opening 0a having a desired shape and size.

Figure i21 is a schematic diagram of a variable mechanism for detecting apertures 18 and 19 and apertures 5 and 6, which are composed of a replacement part 71 and a holding part 72. When the replacement part 71 is inserted into the holding part 72, the desired shape,
Aperture force of size (arranged in the exchange part 71 so as to intersect on the optical axis, and installed the holding part 72 so that the various apertures are in a plane perpendicular to the optical axis, and have various shapes and sizes) By replacing the replacement parts 71゜71', . . . with a replacement part having the desired shape and size, the detection aperture 18゜1
9 and apertures 5 and 6.

FIG. 22 is a schematic diagram of a variable mechanism for the detection apertures 18, 19 and the apertures 5, 6, which are composed of a slide portion 73 and a holding portion 74. When the slide part 73 is inserted into the holding part 74 and slid, the slide part 73 is arranged so that an aperture of the desired shape and size intersects the optical axis, and the various apertures are in a plane perpendicular to the optical axis. The detection apertures 18 and 19 and the aperture 5.6 are varied by installing the holding part 74 as shown in FIG.

Also, as an example of a mechanism that changes the transmittance of a filter,
An example of a mechanism is to provide an ND filter with a desired transmittance for each aperture shown in FIGS. 20 to 22, thereby selecting an ND filter with a desired transmittance according to each variable mechanism.

In that case, the shape and size of each opening shall be the same.

Next, in the basic embodiment, a method is adopted in which the speckle pattern is detected on the Fraunhofer surface relative to the surface to be measured, but it is possible to configure the measuring instrument by a method that detects the speckle pattern on the image surface. can. An example of this configuration is shown in FIG. 23. The speckle light reflected by the wavelength separation mirror 15 is transmitted to the beam splitter
and condensing lens 113. beam splinter 11
4 and is imaged on the image plane. Here, an image of the surface to be measured is formed using only laser light, and a speckle pattern is superimposed on this image. Since this speckle pattern also moves with the movement of blood cells, it is extracted as a change in light intensity through the detection aperture 115, and is detected by a photomultiplier tube (photomultiplier) 40 through a condensing lens 116. This signal processing can be performed in the same way as in the basic embodiment. In this image plane detection method, the entire laser beam irradiation area is formed as an image on the detection surface, but since the detection aperture 115 is placed there, specific parts within the laser beam irradiation area can be will be sampled as the measurement site.

This point is different from the Fraunhofer diffraction surface detection method. That is, for the Fraunhofer diffraction surface detection method,
It becomes possible to measure blood flow in only a portion of the laser beam irradiation area. Furthermore, due to the fundamental principle of optics that an image is formed by light arriving from various angles converging on an image plane, the amount of light detected at one point is greater than on a diffraction surface. Therefore, highly sensitive measurements can be made.

As described above, since the measurement site is sampled by the detection aperture 115, in order to visually observe it, the speckle light reflected by the beam splitter 114 is sampled by the detection aperture 115.
An image is formed on an index plate 117 placed on an image plane equivalent to the plane of , and observed through an eyepiece lens 118. Index plate 117
There is a crosshair at the top, and the center intersection is the detection aperture 115
coincides with the center of the aperture. Further, the indicator plate 117 and the detection aperture 115 are interlocked. This results in
Detection aperture 1 allows you to determine the measurement position while observing the image plane
Reference numeral 15 is a type 2 opening so that the shape and size can be adjusted, and the center of the opening is always constant.

In addition, the speckle light reflected by the beam splitter 112 is passed through a condensing lens 119 to detect oil at detection apertures 120 and 121 placed on the Fraunhofer surface, and further passed through a condensing lens 122 and detected by photoelectron multiplication 'f40. This makes it possible to perform the Fraunhofer diffraction surface detection shown in the basic embodiment.

In the present invention, it is possible to use an apparatus as shown in FIG. 23 which combines the Fraunhofer diffraction surface detection method and an image plane detection method, or an apparatus characterized by using only the image plane detection method alone. If both methods are combined, it is very meaningful because mutually complementary data can be obtained (ie, the blood flow state of the entire laser beam irradiated area and the blood flow state of a part thereof).

Other 7fM standard methods include the detection aperture 18 and 19 in Figure 1,
Instead of the detection apertures 115, 120, 121, etc. in FIG. 23, as shown in FIG. 24, an optical spatial filter 123 may be placed on the same Fraunhofer plane or image plane to measure the blood flow state. 0 With the original detection aperture, in principle the light intensity of the speckle pattern is detected at one point, but on the other hand, for example, a lattice-shaped spatial filter 124 may be placed to detect the speckle pattern in a certain direction. By extracting the degree of movement, it is possible to detect a specific frequency determined by the grid spacing from the frequency component of the speckle pattern movement, and by obtaining this value with the microcomputer-signal processing circuit 50, the state of blood flow in a specific direction can be detected. Measurement becomes possible. Furthermore, by using a random pattern type spatial filter 125, the irregular moving component of the speckle pattern movement is extracted, and the frequency distribution of this signal is evaluated by the signal processing method as shown in FIGS. 4 and 6. Also, irregularities in blood flow can be investigated. If these measurement results are evaluated together with the measurement results obtained according to the original basic embodiment, the blood flow state of the object to be measured can be investigated in more detail, which can be said to be of great clinical use.

ti Yeongjeongyu 1” Next, the interlocking of various 33il alignment mechanisms will be explained.

In order to prevent damage to the eye to be examined due to laser light, the intensity of the laser light irradiated to the eye to be examined in this device conforms to various standards (I EC, ANS) for preventing damage to the eye to be examined.
I). However, in order to reduce the psychological burden on the examinee due to the laser light being irradiated onto the eyeballs, and to reduce the discomfort caused by the glare of the laser light, there is a need for blood flow measurement. In the s@ stage, it is desired that the intensity of the laser light irradiated to the eye to be examined is automatically set lower than that in the blood flow state measurement stage. Further, as a general characteristic of the laser light source 1, the output is not stable immediately after the power is turned on, which tends to adversely affect measurement. Therefore, it is preferable to keep the laser light source powered on at all times.On the other hand, when using this device as a fundus camera, the laser light is not necessary, and considering the psychological burden on the examinee, the laser light source is in the laser oscillation state. However, a function that does not irradiate the eye to be examined with laser light is desired.

In addition, in the preparation stage for measuring the blood flow state, there are times when the focus of the laser beam is on the central part of the subject's cornea 13a, and in that case, the strong laser beam is directed toward the photomultiplier tube 40 side. reflect. If a high voltage is applied to the photomultiplier tube 40 at this time, there is a risk that the photomultiplier tube 40 may be damaged or its lifespan may be shortened due to receiving strong light. To prevent this, it is desirable that the photomultiplier tube 40 detects light only when measuring the state of blood flow.

In consideration of the above problems, the present invention can include the following interlocking mechanism.

As shown in Fig. 25, this device has both the function of a fundus camera and the movement of a blood flow meter, but a switch 81 is installed to select the function, and this allows the following modes: Three modes can be selected: a blood flow state measurement preparation mode, and a blood flow state measurement mode, and the modes are switched in the order of ■-■→■, and ■→■→■.

This is done in consideration of the conventional problems. In addition, in all three conditions (■, ■, ■), by turning on the fundus photography mode changeover switch 82, the fundus visual observation and fundus photography functions of the fundus camera can be operated, and the laser light source 1 is always oscillating.

Make sure that this device is in mode ■ when you turn on the main power of the device. That is, even if the mode is set to ■ or ■, the mode automatically changes to mode ■ when the power is turned on. At this time, the laser beam J11 is blocked by the shutter 7, so that the laser beam does not enter the eye to be examined, and the shutter 21 located on the entire surface of the photomultiplier tube 40 is closed.

When the mode is changed from ■ to ■ using the function selection switch 81, the shutter 21 on the entire surface of the photomultiplier tube 40 remains closed, the shutter 7 on the entire surface of the laser is opened, the light amount adjustment filter 2 enters the optical path, and the blood flow state is monitored. Laser light that is weaker than that used during measurement and measurement enters the eye to be examined.

When you turn off the fundus photography mode switch 82 and press the push-button switch 83 on the top of the focusing joystick of the fundus camera, and change the mode from ■ to ■, the shutter 7 in front of the laser remains open and the light fil 1 is The fiber filter is changed and the intensity of the laser beam irradiated to the eye to be examined is set to the strong intensity used for blood flow measurement, and the shutter 21 in front of the photomultiplier tube 40 is opened, and speckle light is detected by the photomultiplier tube 40. It becomes possible.

When the preset measurement time has passed, the mode automatically changes to ■, the shutter 21 on the entire surface of the photomultiplier tube 40 is closed, and the light amount adjustment filter 2 is switched back to the measurement preparation level, and the laser beam is applied to the eye to be examined. The irradiation intensity is reduced.

In addition to the above-mentioned automatic measurement time, it is also possible to set the measurement time manually by switching a separately prepared switch 84. In that case, measurement is started by pressing the push-button switch 83 at the top of the joystick once with the fundus photography mode changeover switch 82 turned off, and measurement is ended by pressing it again.

By turning on the fundus photography mode changeover switch 82 and pressing the push-button switch 83 at the top of the joystick, fundus photography can be performed in any of the modes ■, ■, and ■. This method allows operations to be performed safely and easily.

In addition, the detection aperture 18.19 and the aperture 5°6 in Fig. 1 are connected by a mechanical interlocking mechanism using wires, etc., to achieve the most appropriate detection for the laser beam irradiation area determined by the aperture 5.6. The size and shape of the openings 18 and 19 may be set in advance so that they can be automatically switched and selected.

Furthermore, in the embodiments described above, when a visible light laser is used as the laser light source, when visually observing the fundus, the area irradiated by the laser light source on the fundus can be visually confirmed by the examiner. Therefore, the advantage is that no special detection or display device is required for observation of the fundus of the eye to be examined.

However, in general, when human eyes recognize visible light, their pupils constrict depending on the amount of change in illuminance over time. With pupil miosis, it becomes difficult to direct the laser light from the laser light source into the fundus due to optical geometry, and even if the fundus were to be irradiated with laser light, it would be scattered by the blood cells in the fundus canal. Because almost no light exits the eyeball through the pupil, speckle light cannot be detected with a photomultiplier tube.
Measurement becomes impossible. Therefore, a mydriatic agent is required, but since the pupil is forcibly dilated at the point where miosis would normally occur to protect the retina, and the laser beam is irradiated to the fundus of the eye, the patient is I feel the laser light is dazzling,
It's unpleasant. Furthermore, if dilation (pupil side) is used, the subject's visual function will be incomplete even after measuring the blood flow state of the fundus until the medicinal effect of the mydriatic agent disappears.

Furthermore, some subjects may have a disease for which it is not advisable to administer mydriatic agents, but in that case, it becomes difficult to measure the state of blood flow in the fundus of the eye.

An embodiment that solves this problem is shown in FIG. 26, and can measure the blood flow state while having the same functions. The light emitted from the observation light source 201 passes through the condenser lens 2
02, the light is reflected by a mirror 206 via a filter 203, a beam splitter (or half mirror) 204, and a condenser lens 205. Here, the semiconductor laser 200 in the observation light source is filtered by a filter 203 having spectral characteristics as shown in FIG. It cuts light at the emission wavelength and shorter wavelengths. Light from the photographing light source 207 is reflected by a mirror 206 via a condenser lens 208, a wavelength separation filter 209 having spectral characteristics shown in FIG. 28, a beam splitter (or half mirror) 204, and a condenser lens 205. The light from the photographing light source 207 that has passed through the beam splitter (or half mirror) 204 and the light from the observation light source 201 that has been reflected are stopped by an optical trap 210. The observation light and photographing light reflected by the mirror 206 pass through the ring slit 2.
11. The light is reflected by the perforated mirror 214 via the relay lenses 212 and 213, and is partially focused in a ring shape on the cornea 13a of the eye 13 to be examined via the objective lens 215, and then reaches the fundus 13b in a diffused state. The fundus 13b is illuminated.

The laser beam from the semiconductor laser light source 200 passes through a condenser lens 216 and a light amount conversion filter 217 for adjusting the light intensity. Furthermore, the light is guided to the fundus illumination optical system of the fundus camera via relay lenses 218 and 219. Further, apertures 220 and 221 are installed between the relay lenses 218 and 219, which allows the size and shape of the laser irradiation area on the fundus to be selected. The laser beam guided by the relay lens 219 is reflected by a prism 222 installed in a part of the annular opening of the ring slit 211 in the fundus illumination optical system, and is passed on the same optical path as the light flux for fundus observation and photography that enters the fundus. guided by. Therefore, the laser beam is reflected by the perforated mirror 214 via the relay lenses 212 and 213, and is reflected by the cornea 13 of the eye 13 via the objective lens 215.
After converging once on a, the fundus 13b is diffused.
, forming an irradiation area that is wider than the diameter of the blood vessel. Furthermore, since the laser light is incident on the same optical path as the observation and photographing light beam, the laser light can be irradiated to a desired position on the fundus by using the left/right/up/down swing mechanism of the fundus camera or the fixation guidance mechanism.

Speckle light generated when the laser light is scattered by blood cells moving in the blood vessels of the fundus of the eye and other reflected light for observation and photography are both received by the objective lens 15 again and passed through the perforated mirror 214 to be photographed. It reaches a beam splitter (or half mirror) 224 through a lens 223,
The beam splitter (or half mirror) 224 divides the optical path into two. Beam splitter (
Alternatively, the light reflected by the half mirror 224 is detected by the photomultiplier tube 40 via the relay lens 225, interference filter 226, detection apertures 227, 228, and lens 229.

Here, the interference filter 226 has a spectral characteristic of transmitting only light in the oscillation wavelength range of the semiconductor laser as shown in FIG. The light is separated from the viewing and photographing light, and only the speckle light is detected by the photomultiplier tube 40.

The observation light and speckle light transmitted through the beam splitter (or half mirror) 224 are transmitted to the relay lens 23.
0, flip-up mirror 231. Relay lens 232, mirror 233. It is detected by an infrared television camera 240 through an infrared television camera lens 234, and the state of the fundus is displayed on a television monitor 241. Observation can be performed at the same time as measurement, and if necessary, a photograph can be taken using photographic film 235 using photographing light from photographing light source 207. Further, a video signal of the fundus from an infrared television camera 240 is stored in a frame memory 243 via a television camera controller 242, and a microcomputer signal processing device performs playback, processing, and other operations of the recorded image information data as necessary. Do this using 50.

A television monitor 241 is used to observe the fundus in the infrared region.
In the fundus image above, the irradiation light of the semiconductor laser is observed.
The measurement position can be confirmed.

It goes without saying that modifications such as those shown in FIGS. 11 to 24 may be applied to the embodiment shown in FIG. 26.

In the embodiment shown in FIG. 26, the following effects can be obtained.

a) Since the fil detection and measurement laser light is not visible light, the subject is not dazzled.

b) Since the laser light for Igt detection and measurement is not visible light, there is no need to dilate the pupils, reducing the burden on the subject.

C) Since there is no need to dilate the pupils, there is no need to consider the effect of the mydriatic agent on the disease of the subject, that is, the subject is not selected.

d) Since it is possible to measure the blood flow state of the fundus when the pupils are not mydriatic as well as when the pupils are not mydriatic, it is possible to investigate the effect of mydriatic agents on the blood flow state.

Also, if a semiconductor laser is used as the laser light source,
Furthermore, the following effects are considered to occur.

e) The size of the device can be reduced by making the laser light source smaller and lighter.
It becomes possible to reduce the weight.

f) Since semiconductor lasers are inexpensive, it is possible to reduce the cost of the device.

In addition to the examples shown in FIGS. 6 and 13, the speckle signal evaluation of speckle-'1'- is also possible in the following example.

That is, evaluation using a differential average frequency defined by a linear equation can be considered.

Here f is the frequency, p (r) is the power spectrum, f
AL is the lowest frequency of the signal analysis band, and fA)4 is the highest frequency of the signal analysis band. These results are obtained, for example, according to a program as shown in the flowchart of FIG. 30, and the results are output to the printer 56 of FIG. 4.

[effect] As explained above, in the present invention.

1) Since it uses the principle that the intensity fluctuations of laser speckles are optically detected and the movement information of the scatterer can be measured by signal processing, it is possible to simply detect the scattered light from blood cells in the target blood vessel on the observation surface. Therefore, the optical system can be very simple and the device can be easily realized.

2) Since the measurement area of the blood flow state of the object is made wider than before, it is less susceptible to the effects of visual fixation movements of the subject's eye.

3) In signal processing, blood flow shape is determined not from a part of the power spectrum of the speckle signal, but from the shape of the entire power spectrum! Because the island is being evaluated, highly accurate measurements can be made.

4) Also, 4 as a fundus camera! It is useful for clinical diagnostic tests in dark areas in ophthalmology, internal medicine, etc., including the fact that it has [ as it is].

etc., various effects can be obtained.

[Brief explanation of the drawing]

FIG. 1 is a configuration diagram showing the configuration of an embodiment of the device of the present invention, FIG. 2 is an explanatory diagram showing the structure of a ring-shaped slit, and FIG. 3 is a characteristic diagram showing the characteristics of the wavelength separation filter of FIG. 1. Fourth
The figure is a block diagram showing the configuration of the signal processing circuit in FIG. 1, FIGS. 5 and 6 are flowcharts showing the flow of control,
Figures 7 (a) to (e) are explanatory diagrams showing the power spectrum distribution, Figures 8 (a) and (b) are explanatory diagrams showing the relationship between the detection aperture and the speckle pattern, and Figure 9 is an observation photograph. FIG. 10 is an optical layout diagram showing the measurement optical system, FIG. 11 is an optical layout diagram showing another example of the measurement optical system, and FIGS. 12 and 13 are other examples. Block diagram and flowchart showing the measurement method of 14th UgJ(a)
, (b) is a characteristic diagram showing the characteristics of other wavelength separation filters, FIGS. 15 to 19 are optical layout diagrams showing other embodiments of the optical system, and FIGS. 20 to 22 are diagrams showing the characteristics of other wavelength separation filters. FIG. 23 is a perspective view showing a configuration for changing the detection aperture;
FIG. 24 is an optical layout diagram showing another embodiment of the measurement optical system;
Fig. 25 is a configuration diagram showing the interlocking mechanism, Fig. 26 is an optical arrangement diagram showing another embodiment, Figs. 27 to 29 are characteristic diagrams showing the optical characteristics of the filter, and Fig. 30 is a speckle pattern diagram. It is a flowchart figure which shows another evaluation method. 1... Laser light source 2... Light amount adjustment filter 8... Ring slit 13... Eye. 13b... Fundus 18.19... Detection aperture 2
1... Shutter 22... Observation light source 24...
Photographing light source 27...Wavelength separation filter

Claims (1)

[Claims] 1) A laser beam of a predetermined diameter is irradiated to the eye area, and the movement of the laser speckle pattern formed by the diffusely reflected light from blood cells in the living tissue is observed as a change in the light intensity of the speckle at an observation point. 1. An ophthalmological diagnostic method characterized in that the shape of the power spectrum distribution of the detected speckle signal is evaluated from the calculation result of the average frequency, and the blood flow state of the living tissue is measured. 2) The ophthalmological diagnosis method according to claim 1, wherein the optical paths of the laser beam and the observation photographing light are made to coincide with each other. 3) The ophthalmological diagnosis method according to claim 1 or 2, wherein the wavelength bands of the laser beam and the observation photographing light can be separated from each other. 4) Claims 1 to 3, wherein the laser light, observation light, and illumination light reflected from the living tissue of the eye are received by the same light receiving optical system, and only the laser light is selectively extracted and detected from them. The ophthalmological diagnostic method described in any one of the preceding paragraphs. 5) The ophthalmological diagnostic method according to any one of claims 1 to 4, wherein the predetermined diameter is set in a wider range than the blood vessel diameter. 6) The ophthalmological diagnostic method according to claim 5, wherein the predetermined diameter is changed in steps. 7) When the biological tissue surface is taken as the object surface, the movement of the diffraction surface speckle pattern observed on the Fraunhofer diffraction surface with respect to the object surface is detected as a change in light intensity. The ophthalmological diagnostic method according to any one of items 6 to 6. 8) The ophthalmological diagnostic method according to any one of claims 1 to 7, wherein the detection aperture on the laser speckle detection surface is made variable. 9) The ophthalmological diagnostic method according to any one of claims 1 to 8, wherein the average frequency is calculated using a logarithm value of a power spectrum. 10) The ophthalmological diagnosis method according to any one of claims 1 to 9, wherein the shape of the laser beam irradiation area is set to a predetermined shape. 11) The ophthalmological diagnosis method according to any one of claims 1 to 10, wherein the detection aperture has a predetermined shape. 12) The ophthalmological diagnostic method according to any one of claims 1 to 11, wherein the speckle pattern can be visually observed. 13) Any one of claims 1 to 12, wherein movement of an image plane speckle pattern observed on an image plane conjugate to a biological tissue plane is detected as a change in light intensity. Ophthalmological diagnostic method described in. 14) The ophthalmological diagnostic method according to claim 13, wherein the speckle pattern is detected as a speckle light intensity change on each of the image plane and the Fraunhofer diffraction surface. 15) The ophthalmological diagnosis method according to claim 14, wherein a spatial filter is installed in place of the detection aperture on the Fraunhofer plane or the image plane. 16) The ophthalmological diagnostic method according to any one of claims 1 to 15, wherein a laser beam having a spectrum in the near-infrared region is used as the laser beam. 17) The ophthalmological diagnostic method according to any one of claims 1 to 16, wherein a laser beam having a spectrum in the near-infrared region is used as the observation light. 18) In front of the speckle detection aperture, insert an interference filter that passes only the wavelength band of the laser light, in front of the observation light source, insert a filter that passes light with a wavelength longer than the wavelength of the laser light, and in front of the photography light source. 18. The ophthalmological diagnostic method according to claim 16 or 17, further comprising inserting a filter that passes light outside the wavelength band of the laser beam. 19) A laser beam of a predetermined diameter is irradiated to the eye area, and the movement of the laser speckle pattern formed by the diffusely reflected light from blood cells in the living tissue is measured from the photoelectron probability distribution based on the photoelectron count of the speckle light at an observation point. An ophthalmologic diagnostic method characterized by detecting changes in the light intensity of speckles and measuring blood flow conditions based on a correlation curve obtained by photon correlation processing of the speckle signals. 20) a laser light source; a laser optical system that irradiates the eye area with laser light from the laser light source as a beam of a predetermined diameter; a means for measuring diffusely reflected light from living tissue irradiated in the eye area; and the measuring means. It detects the movement of the laser speckle pattern formed by the diffusely reflected light as a change in speckle light intensity, and detects the speckle pattern based on the shape of the power spectrum distribution of the obtained speckle signal. An ophthalmologic diagnostic device that measures blood flow conditions in living tissues based on photon correlation processing of optical signals. 21) The ophthalmological diagnostic apparatus according to claim 20, wherein an optical system for observing and photographing the eye area is disposed on the same optical path as the laser optical system. 22) The ophthalmological diagnostic apparatus according to claim 21, wherein the observation light source and the photographing light source of the observation and photographing optical system are arranged on different optical axes. 23) The ophthalmological diagnostic apparatus according to any one of claims 20 to 22, wherein a He-Ne laser or an argon laser is used as the laser light source. 24) The ophthalmological diagnostic apparatus according to any one of claims 20 to 23, further comprising a wavelength separation mirror that reflects only laser speckle light among the diffusely reflected light. 25) The ophthalmologic diagnostic apparatus according to claim 24, wherein the wavelength separation mirror is flip-up type or vertically slidable type so that it can be inserted into and removed from the optical path. 26) Claims 20 to 25, wherein the measuring means has a detection aperture for detecting a speckle pattern.
The ophthalmologic diagnostic device according to any one of the preceding paragraphs. 27) The ophthalmological diagnostic apparatus according to claim 26, wherein the detection aperture is arranged on a Fraunhofer diffraction surface with respect to an object surface when biological tissue is used as an object surface. 28) The ophthalmological diagnostic apparatus according to claim 26, wherein the detection aperture is arranged on an image plane conjugate with a biological tissue plane, and movement of image plane speckles is detected. 29) The ophthalmological diagnostic apparatus according to claim 26, wherein the detection aperture is arranged on each of the image plane and the Fraunhofer diffraction surface. 30) The ophthalmological diagnostic apparatus according to claim 27, further comprising a mechanism for visually observing and adjusting the position of the detection aperture when the detection aperture is disposed on the image plane. 31) The ophthalmological diagnostic apparatus according to claim 30, wherein the adjustment and visual observation are performed using an index plate that is interlocked with the detection aperture. 32) The ophthalmological diagnostic apparatus according to any one of claims 27 to 31, wherein the size and shape of the detection aperture are variable. 33) The ophthalmological diagnostic apparatus according to any one of claims 20 to 32, which is provided with an aperture mechanism that adjusts the size and shape of the laser beam irradiation area. 34) The ophthalmological diagnostic apparatus according to claim 33, wherein the size of the detection aperture is set to a predetermined value in conjunction with the adjustment of the aperture. 35) Any one of claims 20 to 34, in which the amount of light from the laser light source can be adjusted.
The ophthalmological diagnostic device described in section. 36) The ophthalmological diagnostic apparatus according to claim 35, wherein the light amount is adjusted using a light amount adjustment filter or a polarizing filter. 37) The ophthalmologic diagnostic apparatus according to claim 35 or 36, wherein the adjustment of the light amount is linked with a measurement start switch. 38) The ophthalmological diagnostic device according to any one of claims 20 to 37, wherein the laser light source is separated from the main body and coupled via an optical fiber. 39) Claims 22 to 3 in which the observation light source is separated from the main body and coupled via an optical fiber.
The ophthalmological diagnostic device according to any one of Items 8 to 8. 40) The ophthalmological diagnostic device according to any one of claims 20 to 39, wherein the laser speckle pattern is visually observed. 41) The ophthalmological diagnostic device according to claim 39, wherein a photodetector for detecting changes in speckle light intensity is separated from the main body and coupled via an optical fiber. 42) Claims 20 to 41, in which the laser beam and observation light are light beams having a spectrum in the near-infrared region
The ophthalmologic diagnostic device according to any one of the preceding paragraphs. 43) Insert an interference filter that passes only the wavelength band of the laser light in front of the speckle detection aperture, insert a filter that passes light with a wavelength longer than the wavelength of the laser light in front of the observation light source, and insert a filter that passes only the wavelength band of the laser light in front of the observation light source. 43. The ophthalmological diagnostic device according to claim 42, further comprising an interference filter that allows only light outside the wavelength band to pass through. 44) Claim 42 or 43, which includes a device for monitoring the fundus image and the laser irradiation structure.
The ophthalmological diagnostic device described in section. 45) The ophthalmological diagnostic device according to claim 44, wherein the information displayed on the monitor device is stored in a memory. 46) The ophthalmological diagnostic device according to any one of claims 42 to 45, which functions as a non-mydriatic fundus camera.