JP2009165710A - Quantitative measuring instrument of fundus blood flow - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To achieve a quantitative measuring instrument of a fundus blood flow simply extracting the retinal blood vessels using the two-dimensional en-face image and two-dimensional tomographic blood flow image of light coherence tomography to quantitatively determine its blood flow. <P>SOLUTION: A structure of the retinal blood vessels is extracted by a Doppler light coherence angiographic means to enable the quantification of the blood flow of the retinal blood vessels. The Doppler light coherence angiographic means is constituted so that an en-face blood vessel image and a blood flow image are formed by light coherence tomography, the diameter, direction and position of the blood vessel are obtained from both the en-face blood vessel image and the blood flow image to determine the blood flow, and the frequency shift of complex OCT (optical coherence tomography) data is analyzed to measure the velocity of the fundus blood flow. <P>COPYRIGHT: (C)2009,JPO&INPIT

Description

  The present invention relates to a fundus blood flow quantitative measurement apparatus, and more specifically, retinal blood flow is quantified by extracting (segmentation) a retinal vascular structure by Doppler optical coherence angiography means (Doppler OCT). The present invention relates to a device for quantitatively measuring fundus blood flow.

  The measurement of blood flow in the retina is necessary for the diagnosis of abnormal blood circulation due to pathology. Conventionally, several blood flow measurement means applied to retinal blood vessels have been reported.

  For example, there are means for tracking the flow rate of the retinal blood vessels by monitoring the fluorescence intensity of the dye injected into the blood vessels. However, dye injection is a burden on the patient. In addition, although the blood flow rate of the retinal blood vessel can be measured by laser Doppler velocity measurement (LDV) means, there is an inconvenience that it is necessary to measure scattered light from two directions.

  In addition, means for observing an image of blood flow using an optical tomographic imaging apparatus “optical coherence tomography” (OCT) is also known.

  The basic OCT 43 is based on a Michelson interferometer, and its principle will be described with reference to FIG. The light emitted from the light source 44 is collimated by the collimator lens 45 and then divided into reference light and object light by the beam splitter 46. The object light is condensed on the measurement object 48 by the objective lens 47 in the object arm, scattered and reflected there, and then returns to the objective lens 47 and the beam splitter 46 again.

  On the other hand, the reference light passes through the objective lens 49 in the reference arm, is reflected by the reference mirror 50, and returns to the beam splitter 46 through the objective lens 49 again. The object light and the reference light that have returned to the beam splitter 46 in this way are incident on the condensing lens 51 together with the object light and are collected on the photodetector 52 (photodiode or the like).

  The light source 44 of the OCT uses a light source of light having low temporal coherence (light emitted from the light source at different times is extremely difficult to interfere with each other). In a Michelson interferometer using temporally low coherence light as a light source, an interference signal appears only when the distance between the reference arm and the object arm is approximately equal. As a result, when the intensity of the interference signal is measured by the photodetector 52 while changing the optical path length difference (τ) between the reference arm and the object arm, an interference signal (interferogram) for the optical path length difference is obtained.

  The shape of the interferogram shows the reflectance distribution in the depth direction of the measurement object 48, and the structure in the depth direction of the measurement object 48 can be obtained by one-dimensional axial scanning. Thus, in the OCT 43, the structure in the depth direction of the measurement object 48 can be measured by optical path length scanning.

  In addition to the scanning in the axial direction, a two-dimensional cross-sectional image of the object to be measured can be obtained by performing a two-dimensional scanning by adding a horizontal mechanical scanning. The scanning device that performs the horizontal scanning includes a configuration in which the object to be measured is directly moved, a configuration in which the objective lens is shifted while the object is fixed, and a pupil of the objective lens while the object to be measured and the objective lens are fixed. The structure etc. which rotate the angle of the galvanometer mirror in the surface vicinity are used.

  As the development of the basic OCT described above, a spectral domain OCT (SD-OCT) that obtains a spectrum signal using a spectroscope, and a wavelength scanning OCT (Swept Source) that obtains a spectrum interference signal by scanning the wavelength of the light source. OCT, abbreviated as “SS-OCT”). SD-OCT includes Fourier domain OCT (Fourier Domain OCT, abbreviated as “FD-OCT”; see Patent Document 2), and polarization-sensitive OCT (Polarization-Sensitive OCT, abbreviated as “PS-OCT”). Reference 3).

  In the Fourier domain OCT, the wavelength spectrum of the reflected light from the object to be measured is acquired with a spectrometer (spectrum spectrometer), and Fourier transform is performed on this spectrum intensity distribution, so that the real space (OCT signal space) is obtained. This Fourier domain OCT does not need to scan in the depth direction, and can measure the cross-sectional structure of the object to be measured by scanning in the x-axis direction.

  The wavelength scanning type OCT obtains a three-dimensional optical tomographic image by changing the wavelength of a light source by a high-speed wavelength scanning laser, rearranging interference signals using a light source scanning signal acquired in synchronization with a spectrum signal, and applying signal processing. Is. As a means for changing the wavelength of the light source, a device using a monochromator can be used as the wavelength scanning OCT.

  Measurement of blood flow distribution in the retina by Doppler optical coherence tomography (Doppler OCT) is known. This is a means by which the distribution of retinal blood flow can be measured using the Fourier domain OCT or the like. Similarly, by using the spectral domain OCT, cross-sectional retinal blood flow image formation can be obtained. Also, the vascular structure of the retina can be observed.

JP 2002-310897 A JP 11-325849 A JP 2004-028970 A

  In quantitative measurement of blood flow using Doppler OCT, for example, several methods have been reported regarding velocity measurement methods in the optical axis direction and lateral direction, and measurement from multiple angles. However, these measurements are difficult to apply to living human eye measurements. In the spectral domain OCT, there is a problem of the measurement range in the optical axis direction.

  The present invention uses a two-dimensional en-face image and a two-dimensional tomographic blood flow image of optical coherence tomography to easily extract retinal blood vessels in two steps and to obtain the blood flow volume quantitatively. It is an object to realize a quantitative measurement device for flow rate.

  In order to solve the above-mentioned problems, the present invention provides a quantitative measurement apparatus for fundus blood flow that can extract the retinal blood vessel structure by Doppler optical coherence angiography means and quantify the blood flow of the retinal blood vessel, the Doppler light The coherence angiography means has a configuration of optical coherence tomography, and is capable of measuring the flow velocity of the fundus blood flow by analyzing the frequency shift of the complex OCT data. Providing equipment.

  The optical coherence tomography forms an en-face blood vessel image and a blood flow image, and the diameter, direction and position of the blood vessel are obtained from the en-face blood vessel image and the blood flow image, the diameter of the blood vessel, The blood flow rate may be determined from the direction and position.

  The absolute value of the blood flow rate may be obtained from the value of the Doppler frequency shift and the diameter, direction, and position of the blood vessel.

  According to the apparatus for quantitatively measuring fundus blood flow according to the present invention, a retinal blood vessel is simply extracted in two steps using a two-dimensional en-face image and a two-dimensional tomographic blood flow image of optical coherence tomography. Blood flow can be determined.

  The best mode for carrying out the fundus blood flow quantitative measurement apparatus according to the present invention will be described below with reference to the drawings based on the embodiments.

  First, Fourier domain OCT (FD-OCT) used as Doppler optical coherence angiography means (Doppler OCT) necessary for implementing the quantitative measurement apparatus for fundus blood flow according to the present invention and wavelength scanning OCT (SS-) OCT) will be schematically described.

  FIG. 1 is a diagram illustrating an overall configuration of the FD-OCT 1. A broadband light source 2, a low coherence interferometer 3, and a spectrometer 4 (spectrometer) are provided. Since the FD-OCT 1 obtains resolution in the depth direction using the principle of low coherence interference, a broadband light source 2 such as an SLD (super luminescent diode) or an ultrashort pulse laser is used as a light source.

  The light emitted from the broadband light source 2 is first split into object light and reference light by the beam splitter 5. Of these, the object light is reflected by the galvanometer mirror 7 through the lens 6 and irradiates the measurement object 8 (biological sample such as the fundus), and is reflected and scattered there and then guided to the spectrometer 4. On the other hand, the reference light is reflected by the reference mirror 10 (plane mirror) through the lens 9 and then guided to the spectroscope 4 in parallel with the object light. These two lights are simultaneously dispersed by the diffraction grating 11 of the spectroscope 4 and interfere in the spectral region. As a result, the spectral interference fringes are measured by the CCD 12.

  By performing appropriate signal processing on the spectral interference fringes, it is possible to obtain a differential of the one-dimensional refractive index distribution in the depth direction at a certain point of the measured object 8, that is, a reflectance distribution. Furthermore, a two-dimensional tomographic image (FD-OCT image) can be obtained by driving the galvanometer mirror 7 and one-dimensionally scanning a measurement point on the measurement object 8.

  In normal OCT, in order to obtain a two-dimensional tomographic image, scanning in the depth (optical axis) direction (this scanning is called “A-scan”, and this direction is also referred to as “A-direction” and “A-scan direction”). And two-dimensional mechanical scanning is required for vertical operation (this scanning is called “B-scan”, and this direction is also called “B-direction”, “B-scan direction”). In FD-OCT1, since A-scan is unnecessary and backscattering data in the depth direction can be acquired by one measurement, only one-dimensional mechanical scanning of B-scan is required.

  A scan in a direction perpendicular to the plane formed by the A-direction and the B-direction is referred to as “C-scan”, and this direction is also referred to as “C-direction” or “C-scan direction”. In short, in FD-OCT1, two-dimensional scanning (B-scan and C-scan) is performed in the plane, so that high-speed tomographic measurement is possible, and two-dimensional and three-dimensional information inside the measurement object 8 can be obtained. it can.

  FIG. 2 is a diagram showing an overall configuration of the wavelength scanning OCT (SS-OCT) 24. The output light emitted from the wavelength scanning light source 25 is sent to the fiber coupler 27 through the fiber 26. In the fiber coupler 27, the output light is divided into object light that is irradiated onto the measurement object 29 through the fiber 28 and reference light that is irradiated onto the fixed reference mirror 31 through the fiber 30.

  The object light is irradiated and reflected on the measurement object 29 through the fiber 28, the lens 32, the scanning mirror 33 and the lens 34 having variable angles, and returns to the fiber coupler 27 through the same route. The reference light is irradiated and reflected on the fixed reference mirror 31 through the fiber 30, the lens 35, and the lens 36, and returns to the fiber coupler 27 through the same route.

  Then, the object light and the reference light are overlapped by the fiber coupler 27 and sent to the optical detector 38 (a point sensor such as a PD (photodiode) is used) through the fiber 37 to be detected as a spectrum interference signal. Is taken into the computer 39. Based on the detection output of the photodetector 38, cross-sectional images of the measured object 29 in the depth direction (A direction) and the scanning direction of the scanning mirror (B direction) are formed. A display 40 is connected to the computer 39.

  Here, the wavelength scanning light source 25 is a light source that scans while changing the wavelength with time, that is, a light source having a wavelength-dependent wavelength. This makes it possible to obtain the reflectance distribution in the depth direction of the object 29 to be measured and to obtain the structure in the depth direction without scanning (moving, A-scan) the fixed reference mirror 31. Scanning in the primary direction A two-dimensional tomographic image can be formed simply by performing (B-scan).

  In the fundus blood flow quantitative measurement apparatus according to the present invention, the above-described Fourier domain OCT (FD-OCT) or wavelength scanning OCT (SS-OCT) is used for Doppler optical coherence angiography (Doppler optical coherence angiography). It is applied as OCT for means (hereinafter referred to as “Doppler OCT”).

  Then, by such OCT, a three-dimensional complex OCT image of the retinal blood vessel image of the fundus oculi, which is the object to be measured, is taken into a computer, blood vessel structure extraction (segmentation) based on the data, and computer quantification based on the extracted data The blood flow velocity of the retinal blood vessels is quantified by means (specifically, functional means of quantification software), and the fundus blood flow is quantitatively measured using the blood vessel lumen diameter obtained from the extracted data. is there.

  In the following, the configuration of the quantifying means in the present invention, specifically, the contents for quantifying the blood flow volume of the retinal blood vessels performed by the computer quantifying means based on the data of the retinal blood vessel structure extracted by Doppler OCT (steps) ).

  In the fundus blood flow quantitative measurement apparatus according to the present invention, in order to quantify the blood flow, it is necessary to obtain an absolute blood flow velocity and a blood vessel lumen (diameter) by a computer quantification means. .

  In order to obtain an absolute blood flow velocity using Doppler OCT, first a Doppler angle is required. Therefore, first, blood vessel structure analysis is required to determine the direction and diameter of the blood vessel by the blood vessel extraction algorithm in the computer quantification means (specifically, the quantification software function means).

  In the three-dimensional image of the OCT image, a plane perpendicular to the optical axis (a plane perpendicular to the A scan direction, a plane including the B scan and C scan directions) is referred to as an en-face. Since the network of retinal blood vessels runs almost parallel to the retina near the surface of the retina, the en-face retinal blood vessel image can be regarded as a blood vessel structure (arrangement) of the retina. Therefore, the blood vessel structure analysis is performed in the following two steps.

  FIG. 3 is a diagram showing the arrangement of the blood vessel wall S with respect to the projection plane of the en-face projection image and the two planes A and P orthogonal to the projection plane and orthogonal to each other. As parameters for the arrangement of the blood vessel wall S, in the first step, the azimuth angle θ (azimuth angle θ) and the diameter D of the blood vessel are obtained. These values are obtained in a plane E perpendicular to the optical axis (z axis), which is a projection plane of the en-face image.

  In the second step, the zenith angle φ of the blood vessel is obtained. This value is obtained in the plane A including the optical axis (z-axis) and the blood vessel. The plane P is a plane perpendicular to the plane A and including the optical axis.

  The position of the blood vessel wall S is determined by these first and second steps. In order to determine the position of the blood vessel wall with normal means, all the three-dimensional data (one million images for a 100 × 100 × 100 image) are scanned and pixels having the characteristics of the blood vessel wall are extracted and connected. Since only three parameters θ, D, and φ need be obtained, the calculation time can be shortened.

From the second-order derivative of the En-face blood vessel image, the blood vessel curve structure can be obtained. The second derivative image H is the second derivative L of the Gaussian function G (r, σ) with an appropriate variance σ = [∂ / ∂x, ∂ / ∂y] ^T [∂ / ∂x, ∂ / ∂y] It is possible to obtain the En-face blood vessel image I by line filtering with G (r, σ) at high speed (H is the correlation between L and I).

In order to determine semi-automatically whether or not the image of interest is a blood vessel, a “blood vessel-likeness” is defined, and the determination is performed using the size of the value. The blood vessel likelihood F is defined from the eigenvalues Λ ₁ and Λ ₂ of the matrix H. F = {| Λ ₂ | (| Λ ₂ | + Λ ₁ )} ^1/2 , (when Λ ₂ <Λ ₁ ≦ 0), {| Λ ₂ | (| Λ ₂ | + αΛ ₁ )} ^{1 / 2} (when Λ ₂ <0 <| Λ ₂ | / α), and 0 (otherwise).

For example, when | Λ ₁ | ˜ | Λ ₂ | >> 0, F is zero, which corresponds to a spotted structure in the image. Further, when Λ ₂ <0 <Λ ₁ , F becomes a large value, which corresponds to a continuous curve structure. Therefore, this means separates a continuous curved structure and a spot-like structure, is sensitive to a discontinuous structure (spot-like structure) on the line, and is effective in extracting a blood vessel structure. . α is the extraction sensitivity of the discontinuous structure, and is set to 0.25, for example.

  In the fundus blood flow quantitative measurement device according to the present invention, it is assumed that the extracted blood vessel image is brighter than the surroundings. In a normal OCT en-face image, blood vessels are measured as dark as shadows. Therefore, it is necessary to use a bright image of a blood vessel portion by reversing light and dark as an image for analysis. Therefore, it is necessary to reverse the image intensity in advance.

Since the eigenvalues Λ ₁ and Λ ₂ of the matrix H represent the direction of the line structure in the image, the direction indicated by the eigenvector is the direction of the azimuth angle θ. Line structures with different scales can be extracted by changing the variance σ of the Gaussian function G (r, σ) that is the basis of the line filter used in the line filtering (multi-scale line filtering method). The scale here refers to the relationship between the size of the pixel and the structure, such as how many pixels the diameter of the vascular structure of interest is composed of.

The blood vessel likelihood F is normalized by the variance σ of the Gaussian function and expressed as σF (σ). The likelihood of a blood vessel at a certain point is represented by the standardized maximum value F ′ = max [σF (σ)] of the likelihood of a blood vessel. The diameter of the blood vessel is represented by D = 4σ _max , and at that time, F ′ = σ _max F (σ _max ) These processes can be performed by computer quantification means (specifically, software function means).

  In the En-face image, retinal blood vessels are extracted semi-automatically by correlating with a top hat function (filtering processing) and performing threshold processing on the correlation value. Since the above automatic extraction cannot be performed with respect to a portion where retinal blood vessels intersect, the operator needs to indicate a portion where retinal blood vessels intersect. With respect to the extracted blood vessel, the central portion of the blood vessel can be determined by connecting the pixels having the maximum value of the blood vessel likelihood F ′ with a line.

  When the azimuth angle θ of the blood vessel and the center line of the blood vessel are known, the blood flow image in the cross section (plane A in FIG. 3) parallel to the longitudinal direction of the blood vessel including the center line of the blood vessel and the axial direction thereof are three-dimensional blood. It is obtained from the flow image. The zenith angle φ of the blood vessel can be obtained by connecting the blood vessel center determined as described in the previous paragraph and calculating the angle between the line and the optical axis.

  Bidirectional blood flow is determined from a three-dimensional blood flow image from the depth (depth position in the optical axis direction) of the blood vessel cross section (see plane P in FIG. 3) and the diameter D of the lumen. The depth of the blood vessel cross section can be obtained from the maximum value of the correlation between the OCT image of the surface P and the circle of diameter D.

  The zenith angle φ of the blood vessel can be considered as a Doppler angle (hereinafter referred to as “Doppler angle φ”) assuming that all blood flow vectors are parallel. In Doppler flow velocity measurement, only the velocity component in the optical axis direction (z direction) is sensitive, and the flow perpendicular to the optical axis is insensitive. The Doppler angle is an angle formed by the optical axis and the blood flow vector and corresponds to the zenith angle φ. In Doppler flow velocity measurement, the blood flow velocity can be corrected from the Doppler angle φ and the Doppler frequency shift Δf.

The absolute blood flow velocity V is expressed by V = λ ₀ Δf / (2n · cosφ). Here, λ ₀ represents the center wavelength of the light source, and n represents the refractive index of the sample. Since the Doppler frequency shift is below the measurement limit, this equation cannot be used when φ is near 90 °.

With these blood vessel parameters, the blood vessel lumen ellipse S in the cross-sectional blood flow image on the plane P can be estimated. The blood flow volume J is integrated in the blood vessel lumen S by taking into account the Doppler frequency shift and the Doppler angle and integrating the lateral blood flow velocity as J = Σ _S (λ ₀ Δf) / (2n) tanφΔs To obtain). Here, Δs is the area of one pixel of the cross-sectional blood flow image.

  An embodiment of a fundus blood flow quantitative measurement apparatus according to the present invention will be described below. In the apparatus for quantitatively measuring the fundus blood flow in this embodiment, as the Doppler OCT, an axial scanning speed of 27.7 kHz and an ultra high resolution SD-OCT having an axial resolution of about 3 μm in the living body are used. The measurement of the preservation of the blood flow rate in the bifurcated portion of the retinal artery (see FIG. 5, in particular, the region in the white square in FIG. 5A).

  The reason for selecting the retinal artery instead of the retinal vein as the measurement target is as follows. Since almost all of the retinal vessels intersect perpendicularly with the incident beam, the Doppler frequency shift is small. This introduces an error in blood vessel extraction. To avoid this situation, it is better to target the artery rather than the vein because the arterial Doppler frequency shift is greater than that of the vein.

  The portions where the Doppler shift occurs are considered to have blood flow, and those portions constitute a blood flow image. In the blood flow image, since the frequency change due to the Doppler shift is very small, the frequency change of the light wave is observed as the phase change. Since the phase change has a limited range of 0 to 2π, a phase uncertainty of 2π occurs. In order to correct this and obtain the correct frequency shift amount, a 2π phase connection algorithm is used two-dimensionally for the portion where the phase jump has occurred.

  Removal of motion artifacts in the blood flow image was performed by means using a histogram (see Japanese Patent Application Laid-Open No. 2007-127425). The distortion of the image due to the bulk motion, which is a phenomenon in which the measurement object moves during measurement (bulk motion), was corrected numerically considering the continuity of the image (S. Makita, Y Hong, M. Yamanari, T. Yatagai, and Y. Yasuno, “Optical coherence angiography,” Opt. Express 14, 7821-7840 (2006)).

  By calculating the maximum projection (MIP) of the three-dimensional OCT logarithmic image around the retinal pigment epithelium (RPE), a high-contrast en-face blood vessel image as shown in FIG. 5B can be obtained. FIG. 4 shows an en-face retinal blood vessel image obtained by performing the multi-scale line filter according to the paragraphs 0045 to 0047, and FIG. 5 shows the result of blood vessel extraction based on the image. 4A shows an en-face blood vessel image, FIG. 4B shows a distribution of blood vessel likelihood F ′, FIG. 4C shows an azimuth distribution, and FIG. 4D shows a blood vessel diameter.

  In the above axial scan by SD-OCT, blood vessel extraction and the flow calculation time of the three extracted blood vessels (three blood vessels constituting the branch shown in FIGS. 5A and 5B) are calculated using a computer (AMD Athlon64 3500+, 2Gb RAM ) Was calculated in about 3 minutes.

  In addition, an image and a fundus photograph (FIG. 5A) extracted semi-automatically were obtained by scanning in the axial direction by SD-OCT. The analysis region (the region including the target branch portion) is within the white square in FIG. 5A. The lines along the blood vessels 61, 62, and 63 in the blood vessels 61, 62, and 63 in FIG. 5B (corresponding to the area within the white square in FIG. 5A) and the line orthogonal to the lines are automatically The extracted blood vessel center line and the vertical transverse line. In the blood vessel 61, although the extracted blood vessel image is almost horizontal, the calculated azimuth angle has a small inclination.

  However, it can be seen from the fundus photograph (FIG. 5A) that the blood vessel 62 has a small azimuth angle. This can be considered that the multi-scale filtering means is sensitive to the sub-pixel structure.

  5C, E, and G show blood vessel lumens and blood flow images in a cross section (see plane P in FIG. 3) orthogonal to the longitudinal direction of the blood vessel. 5D, F, and H show blood vessel lumen and blood flow images in a cross section parallel to the longitudinal direction of the blood vessel (see plane A in FIG. 3). The thick lines in FIGS. 5D, F, and H are auxiliary lines indicating the center of the blood vessel at the center of the blood vessel image.

  Since the size of voxels (small cubes on three-dimensional lattice points. Representing the shape by the set is called “voxel expression”) is anisotropic (not cubes). Accuracy and blood flow measurement accuracy depend on the azimuth. Blood flow parameters were determined for all 20 sets, and peak blood flow velocity Vpeak, blood vessel diameter D, and volumetric flow rate J were determined. This is shown in Table 1 (for blood vessels 61-63 in FIG. 5B).

  In Table 1, the sign of the flow velocity and the flow rate indicates the direction of the flow, with + indicating the flow into the branch and − indicating the flow out of the branch. FIG. 6 is a plot of the total flow rate of the blood vessel 1 and the sum of the flow rates of the blood vessels 2 and 3 in FIG. 5B.

  The inflow amount and outflow amount of blood are in a directly proportional relationship, and the correlation coefficient is 0.75. The inflow blood flow rate of 3.41 μl / min substantially coincides with the outflow blood flow rate of 3.23 μl / mim. Therefore, it was confirmed that the fundus blood flow quantitative measurement apparatus according to the present invention has accuracy and reliability.

  The best mode for carrying out the present invention has been described based on the embodiments. However, the present invention is not limited to such embodiments, and within the scope of the technical matters described in the claims. Needless to say, there are various embodiments.

  Since the fundus blood flow quantitative measurement device according to the present invention according to the present invention is configured as described above, not only quantitative measurement of the fundus blood flow, but also medical examination devices generally have a high-resolution resolution. The present invention can be applied to various technical fields required, for example, living organisms and structures of animals and plants.

It is a figure explaining the whole structure of FD-OCT used for Doppler OCT of this invention. It is a figure explaining the whole structure of SS-OCT used for Doppler OCT of the present invention. It is a figure explaining the principle of this invention. It is a figure explaining the Example of this invention. It is a figure explaining the Example of this invention. It is a figure explaining the Example of this invention. It is a figure explaining the conventional basic OCT.

Explanation of symbols

1 FD-OCT
2 Broadband light source
3 Low coherence interferometer
4 Spectrometer
5 Beam splitter
6, 9, 32, 34, 35, 36 Lens
7 Galvano mirror
8 Object to be measured
10 Reference mirror
11 Diffraction grating
12 CCD
DESCRIPTION OF SYMBOLS 13, 21 Image processing apparatus 14 CCD photodetector 15 OCT interference signal input part 16 OCT interference signal smoothing means 17 Envelope detection means 18 Reverse filter preparation means 19 Envelope correction means 20 OCT interference signal output part 22 Interference signal conversion envelope detection means
24 wavelength scanning OCT
25 wavelength scanning light source
26, 28, 37 fiber
27 Fiber coupler
29, 48 Object to be measured
30 fiber
31 Fixed reference mirror
33 Scanning mirror
38 Light detector
39 computers
40 display
43 OCT
44 Light source
45 Collimating lens
46 Beam splitter
47 Objective lens in the object arm
49 Objective lens in the reference arm
50 reference mirror
51 condenser lens
52 (Photodiode etc.) Photodetector
61-63 blood vessels

Claims (3)

An apparatus for quantitatively measuring the fundus blood flow that can extract the retinal blood vessel structure by Doppler optical coherence angiography means and quantify the blood flow of the retinal blood vessel,
The Doppler optical coherence angiography means has a configuration of optical coherence tomography, and can measure the flow velocity of the fundus blood flow by analyzing the frequency shift of complex OCT data. Quantitative measurement device.   The optical coherence tomography forms an en-face blood vessel image and a blood flow image, and the diameter, direction and position of the blood vessel are obtained from the en-face blood vessel image and the blood flow image, the diameter of the blood vessel, The apparatus for quantitatively measuring fundus blood flow according to claim 1, wherein the blood flow is determined from a direction and a position.   The apparatus for quantitatively measuring fundus blood flow according to claim 1, wherein the absolute value of the blood flow is obtained from a Doppler frequency shift value and a diameter, direction, and position of a blood vessel.