JP2018066762A - Jones matrix oct device and program - Google Patents

Abstract

PROBLEM TO BE SOLVED: To provide a Jones matrix OCT which is improved in image stability, image quality, and an image depth area.SOLUTION: A Jones matrix OCT device comprises: a light source; a polarization separation unit for separating a luminous flux from the light source into P polarized light and S polarized light; a polarization delay unit for relatively applying a phase difference to the P polarized light and the S polarized light from the light separation unit and overlapping the polarized light and the S polarized light having a phase difference; an objective unit for irradiating a measurement target with a luminous flux from the polarization delay unit and receiving reflection light from the measurement target; a detection unit for making a luminous flux from the objective unit and the luminous flux from the light source interfere with each other and separating the same into vertical polarization and horizontal polarization to detect the same; and a calculation unit for generating a tomogram image by receiving a signal from the detection unit.SELECTED DRAWING: Figure 1

Description

  The present invention relates to an optical interferometry technique, and relates to an image processing technique for an optical coherence tomography (abbreviated as “OCT”) image, and in particular, a Jones matrix OCT apparatus using a Jones matrix, and The present invention relates to a program for image processing of measurement data obtained by the OCT.

  Conventionally, as one of the non-destructive tomographic techniques used in the medical field etc., there is an optical tomographic imaging method “optical coherence tomography” (OCT) using temporally low coherence light as a probe (patent) (patent) Reference 1). Since OCT uses light as a measurement probe, it has the advantage that it can measure the refractive index distribution, spectral information, polarization information (birefringence distribution), etc. of the measured object.

  OCT can visualize the cross-sectional structure and three-dimensional structure of a living body with a resolution of approximately 2 to 15 μm with non-invasive, high-contrast images, so that ophthalmology, dermatology, dentistry, gastroenterology and cardiology Widely used in such fields.

  The basic OCT 93 is based on a Michelson interferometer, and its principle will be described with reference to FIG. The light emitted from the light source 94 is collimated by the collimator lens 95 and then divided into reference light and object light by the beam splitter 96. The object light is condensed on the measurement object 98 by the objective lens 97 in the object arm, scattered and reflected there, and then returns to the objective lens 97 and the beam splitter 96 again.

  On the other hand, the reference light passes through the objective lens 99 in the reference arm, is reflected by the reference mirror 100, and returns to the beam splitter 96 through the objective lens 99 again. The object light and the reference light that have returned to the beam splitter 96 in this way are incident on the condenser lens 101 together with the object light and are collected on the photodetector 102 (photodiode or the like).

  The OCT light source 94 uses a light source of light having low temporal coherence (light emitted from the light source at different times and extremely unlikely 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 102 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 measured object 98, and the structure in the depth direction of the measured object 98 can be obtained by one-dimensional axial scanning. As described above, in the OCT 93, the structure in the depth direction of the measurement object 98 can be measured by optical path length scanning.

  In addition to such axial scanning, a two-dimensional tomographic 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).

  FD-OCT obtains the wavelength spectrum of the reflected light from the object to be measured with a spectrometer (spectrum spectrometer), and performs Fourier transform on the spectrum intensity distribution, thereby realizing the real space (OCT signal space). The FD-OCT can measure the tomographic structure of the object to be measured by performing scanning in the x-axis direction without performing scanning in the depth direction.

  SS-OCT obtains a three-dimensional optical tomographic image by changing the wavelength of the light source using 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. It is. As a means for changing the wavelength of the light source, a device using a monochromator can be used as SS-OCT.

  PS-OCT, like the Fourier domain OCT, acquires the wavelength spectrum of the reflected light from the object to be measured using a spectrum spectrometer, but captures the polarization information of the sample (object to be detected) This is an optical coherence tomography apparatus capable of measuring a fine structure and refractive index anisotropy (see Patent Document 3).

  More specifically, PS-OCT, for example, continuously modulates the polarization state of a linearly polarized beam simultaneously with a B-scan, and allows incident light and reference light to pass through a half-wave plate and a quarter-wave plate, respectively. As horizontal linearly polarized light, vertical linearly polarized light, 45 ° linearly polarized light, and circularly polarized light, the reflected light from the object to be measured and the reference light are superimposed and passed through a half-wave plate, a quarter-wave plate, etc. The light is incident on a spectrum spectrometer and interferes, and only a component having a specific polarization state of the object light is extracted and subjected to Fourier transform. This PS-OCT also does not require scanning in the depth direction.

  Doppler optical coherence tomography (Doppler) is an OCT that is suitable for ophthalmic practice by measuring blood flow in the living body, especially the fundus of the eye (retina blood flow) in a non-invasive manner, and also for imaging in the cancer and brain. OCT) is known (see Patent Document 4).

  Doppler OCT is a means by which the distribution of blood flow and the like can be measured using the FD-OCT and 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.

  Moreover, OCT (Jones matrix OCT) using a Jones matrix is known as an aspect of OCT (see Patent Document 5). In particular, OCT that uses fibers to form Doppler and polarized images is known (Non-Patent Document 1).

JP 2002-310897 A JP 11-325849 A JP 2004-028970 A JP 2009-165710 A Japanese Patent No. 434429

Makita, Shuichi; Hong, Youngjoo; Yamanari, Masahiro; Yatagai, Toyohiko; Yasuno, Yoshiaki, “Optical Coherence Angiography”, Optics Express Vol. 14, no. 17, pp. 7821-7840 (2006).

  OCT such as PS-OCT using the above-mentioned conventional polarized light is relatively shallow in measurable depth region and has low image stability and image quality compared to OCT not using polarized light, and is used for medical treatment. There was a limit to doing it.

  Further, by using Doppler OCT, it is possible to selectively visualize the fundus blood vessel, but only a relatively low-contrast pseudo-image image can be obtained with the image obtained by Doppler OCT as it is.

  An object of the present invention is to solve the above-described conventional problems, and an object thereof is to realize an improved Jones matrix OCT in the image stability, image quality, and image depth regions.

  In order to solve the above-described problems, the present invention provides a light source, a polarization separation unit that separates a light beam from the light source into P-polarized light and S-polarized light, and relative to the P-polarized light and S-polarized light from the polarization separation unit. A polarization delay unit that gives a phase difference and superimposes P-polarized light and S-polarized light having a phase difference; an objective unit that irradiates a measurement object with a light beam from the polarization delay unit and receives reflected light from the measurement object; From the following (Equation 1), the light beam from the objective unit and the light beam from the light source interfere with each other and receive a signal from the detection unit that separates and detects vertical polarization and horizontal polarization. There is provided a Jones Matrix OCT apparatus having a calculation unit for generating a tomographic image based on Equation (3).

In the above ( _Equation ¹⁾ , E ⁽¹⁾ _outA (z) is an OCT cross-sectional image signal corresponding to the horizontal polarization and the P polarization, and E ⁽²⁾ _outA (z) is the horizontal polarization and the An OCT cross-sectional image signal corresponding to S-polarized light, E ⁽¹⁾ _outB (z) is an OCT cross-sectional image signal corresponding to the vertically polarized light and the P-polarized light, and E ⁽²⁾ _outB (z) is It is an OCT cross-sectional image signal corresponding to the vertically polarized light and the S polarized light.

  The calculation unit can calculate and display a Doppler image by the following equation (Equation 4). By measuring the same point by a plurality of (m) B-scans, the following equation (Equation The Doppler shift Δφ (z, j) represented by 5) may be applied to Doppler measurement to improve the sensitivity of Doppler measurement.

In the above (Equation 5), m ₀ is a parameter of the B-scan at the start, W (z, j) is an intensity mask defined by (Equation 6) below, and ε ² is an OCT image Noise floor (minimum noise).

  The calculation unit calculates an OCT image based on scattered reflected light from the measurement target based on the following equation (Expression 7) using a coherent matrix component, and uses a plurality of B-scan data. A configuration may be provided in which a scattered OCT image is calculated based on the following equation (Equation 8) by taking a coherent average of the matrix components.

  The polarization delay unit includes a phase calibration unit having a calibration mirror, and the calibration mirror reflects the P-polarized light or the S-polarized interference light in the polarization delay unit, sends the reflected light to the detection unit, and performs the calculation. The unit may correct jitter in synchronization between wavelength scanning and A-scanning of the light source using a calibration signal generated by superimposing the interference light and the light beam from the light source in the detection unit.

  In order to solve the above-described problems, the present invention provides a light source, a polarization separation unit that separates a light beam from the light source into P-polarized light and S-polarized light, and relative to the P-polarized light and S-polarized light from the polarization separation unit. A polarization delay unit that gives a phase difference and superimposes P-polarized light and S-polarized light having a phase difference; an objective unit that irradiates a measurement object with a light beam from the polarization delay unit and receives reflected light from the measurement object; A detection unit that causes the light beam from the objective unit to interfere with the light beam from the light source and separates the light into vertical polarization and horizontal polarization, and a calculation unit that generates a tomographic image based on a signal from the detection unit; , A program installed in the calculation unit of a Jones matrix OCT apparatus having the calculation unit from the detection unit Receiving a signal, the following from equation (9) based on the equation (11), to function as an image processing means for generating a tomographic image, a program.

In ( _Equation 9), E ⁽¹⁾ _outA (z) is an OCT cross-sectional image signal corresponding to the horizontal polarization and the P polarization, and E ⁽²⁾ _outA (z) is the horizontal polarization and the S OCT cross-sectional image signal corresponding to polarized light, E ⁽¹⁾ _outB (z) is the OCT cross-sectional image signal corresponding to the vertical polarization and P-polarized light, and E ⁽²⁾ _outB (z) is the vertical It is an OCT cross-sectional image signal corresponding to polarized light and S polarized light.

  The program according to the present invention can display the calculation unit by calculating a Doppler image by the following equation (Equation 12) and measuring the same point by a plurality of (m) B-scans. The Doppler shift Δφ (z, j) expressed by the following equation (Equation 13) may be applied to Doppler measurement to function as an image processing means for improving the sensitivity of Doppler measurement.

In (Equation 13), m ₀ is a parameter of the B-scan at the start, W (z, j) is an intensity mask defined by (Equation 14) shown below, and ε ² is an OCT image Noise floor (minimum noise).

 The program according to the present invention calculates the OCT image based on the scattered reflected light from the measurement object based on the following equation (Equation 15) using a coherent matrix component, and a plurality of times. By using the B-scan data of the above and taking the coherent average of the matrix components, the scattered OCT image may be functioned as an image processing unit that is calculated based on the following equation (Equation 16).

  According to the Jones matrix OCT system and program according to the present invention, the image quality of an image obtained by the Jones matrix OCT is improved, and an improved image is also obtained for a deep region of a measurement object such as the choroid on the fundus. Can be obtained.

  In particular, a high-contrast image can be obtained by applying to OCT having a Doppler function. Therefore, for example, image information with higher resolution can be obtained non-invasively compared to image information obtained by fundus angiography in ophthalmic diagnosis, which has been obtained only invasively, such as the vasculature of the fundus vasculature. Enables quantitative evaluation.

1 is a diagram illustrating an overall configuration of an OCT apparatus according to an embodiment of a Jones matrix OCT system according to the present invention. It is the OCT tomographic image by the interference light respectively corresponding to the S wave and P wave which were received with the horizontal balance polarization | polarized-light detector and vertical balance polarization | polarized-light detector of the OCT apparatus of the said Example. It is a figure which shows the computer which performs the image process in the OCT system of the said Example. (A) is a tomographic image of a choroidal blood vessel obtained by the Jones matrix OCT system of the above embodiment, and (b) is a tomographic image of a choroidal blood vessel obtained by applying the present invention to an OCT having a Doppler function. It is an image, (c) is a comparative example, and is a contrast image of blood vessels of the same choroid obtained by angiography in the form of conventional indocyanine green fluorescent fundus angiography (ICGA). It is a figure explaining the conventional basic OCT.

  Jones matrix optical coherence tomography according to the present invention (Jones matrix OCT: OCT which obtains an image by processing interference signals obtained by OCT using Jones matrix means) and image processing of image data obtained by OCT A mode for executing a program to be executed will be described below with reference to the drawings by way of an example.

  FIG. 1 is a diagram showing an overall configuration of a Jones matrix OCT system 1 according to the present invention. As shown in FIG. 1, a Jones matrix OCT system 1 according to the present invention includes an optical image measurement device (OCT device) 2 and a computer 3 that performs image processing on image data obtained by the optical image measurement device (OCT device) 2. I have.

  In the present invention, the optical image measurement device (OCT device) 2 can be applied to general OCT data processing using polarized light. As an example, in this embodiment, as the optical image measurement device (OCT device) 2, A description will be given of a configuration using wavelength scanning OCT (Swept Source OCT (abbreviated as “SS-OCT” for short), hereinafter referred to as “SS-OCT”) that obtains a spectrum interference signal by scanning the wavelength of the light source.

  A program for performing image processing according to the present invention causes the computer 3 to function as image processing means using image data obtained by the wavelength scanning OCT2.

(OCT equipment)
In the present invention, the wavelength scanning OCT 2 divides the beam from the light source 6 and linearly polarizes one of the beams, and the S wave (S wave polarization, S wave component) and P wave (P wave polarization, P wave component). ) And irradiating the measurement object 7 as incident beams having different optical path lengths (B-scan) to obtain reflected light (object light), and using the other beam as reference light, both spectra OCT measurement is performed by interference.

  In this specification, the A-scan is a scan in the depth direction of the measurement object 7 (actually irradiation with a measurement beam), and the data obtained by the A-scan is A-line data or abbreviated. A-line. The B-scan is a scan in a direction perpendicular to the depth direction of the measurement object 7.

  Then, among the spectral interference components caused by the S wave and the P wave, the horizontal polarization component (H) and the vertical polarization component (V) are measured by a balanced polarization detector, thereby expressing the four polarization characteristics of the measurement object 7. It is characterized by a multi-contrast OCT configuration capable of obtaining Jones vectors (H image by S wave and P wave, and V image by S wave and P wave, respectively).

  The wavelength scanning OCT 2 includes a light source 6, an optical isolator 8 (a device having a function of transmitting only light in one direction and blocking light in the reverse direction), a fiber coupler 11 (optical coupler), a reference arm 12, and a probe arm 13 (measurement). Optical elements such as a target arm) and a polarization separation detection unit 14.

  The optical system of the wavelength scanning type OCT 2 may be of a type (free space type) in which the optical elements are coupled to each other by the fiber 17 but not coupled to the fiber 17.

  The light source 6 is a wavelength scanning light source 6 that scans a wavelength, and uses a super luminescent diode (SLD).

  The probe arm 13 has a polarization controller 20, a polarization delay unit 21 (polarization delay unit), a coupler 22, a fiber collimator 23, a biaxial galvano scanner 24, an objective lens 25, and an aspherical ophthalmic unit on its optical path. Lenses for use 26 are arranged in order.

  A phase calibration unit 31 branched from the optical path to the measurement object 7 is connected to the coupler 22. The phase calibration unit 31 includes a fiber collimator 32, a lens 33, and a mirror (calibration mirror) 34.

  The polarization delay unit 21 includes a fiber collimator 37, a 45 ° linear polarizer 38, a polarization beam splitter 39, and a first dove prism 40 disposed on each of two optical paths divided by the polarization beam splitter 39. , A second dove prism 41, a polarization beam splitter 42 for multiplexing, and a fiber collimator 43.

  The “dove prism” is a prism in which a right-angle prism is halved. When a parallel light beam is incident on the inclined surface of a right-angle triangle, it is reflected internally and emitted parallel to the incident direction. The position of the first dove prism 40 is fixed at a fixed position, and the optical path length from the polarization beam splitter 39 via the first dove prism 40 to the polarization beam splitter 42 for multiplexing is constant. is there.

  The second dove prism 41 is a prism for adjusting the optical path length, and is movably provided in the optical path direction. By appropriately moving in the optical path direction, the optical path length from the polarizing beam splitter 39 via the second dove prism 41 to the polarizing beam splitter 42 for multiplexing can be adjusted.

  The reference arm 12 includes a fiber collimator 46, a mirror 47, a mirror 48, a fiber collimator 49, and a polarization controller 50 arranged in this order on the optical path.

  The polarization separation detection unit 14 includes a beam splitter 54 for multiplexing, a fiber collimator disposed on the optical path from one probe arm 13 on the incident side of the beam splitter 54 for multiplexing and connected to the polarization controller 55. A fiber collimator 57 and a 45 ° linear polarizer 58 are connected in order to the optical path from the other optical path reference arm 12.

  The polarization separation detection unit 14 includes a first optical path 61 and a second optical path 62 on the emission side of the beam splitter 54 for multiplexing and for splitting and transmitting the beam. A first polarization beam splitter 63 is provided on the first optical path 61, and a second polarization beam splitter 64 is provided on the second optical path.

  A horizontal balance polarization detector 68 is connected to the optical path splitting by the first beam splitter 63 and sending the horizontal polarization component via a fiber collimator 67, and the fiber collimator 69 is connected to the optical path sending the vertical polarization component. The vertical balance polarization detector 70 is connected via

  Similarly, the optical path splitting by the second beam splitter 64 and sending the horizontal polarization component is connected to the horizontal balance polarization detector 68 via the fiber collimator 74, and the optical path of the horizontal beam sending the vertical polarization component Are connected to a vertical balance polarization detector 70 via a fiber collimator 75.

(Function)
In the wavelength scanning OCT 2 configured as described above, the beam from the light source 6 is separated by the fiber coupler 11 through the optical isolator 8 with a light quantity of 10% in the direction of the reference arm 12 and 90% in the direction of the probe arm 13. The optical isolator 8 permits only passage of the beam in one direction, and blocks the reflected light to protect the light source 6.

  The beam separated in the direction of the probe arm 13 passes through the polarization controller 20 and enters the polarization delay unit 21. In the polarization delay unit 21, after passing through the fiber collimator 37, linearly polarized beams that have been linearly polarized by the linear polarizer 38 and linearly polarized beams are oscillated in the vertical direction by the polarization beam splitter 39 (reflected polarization components). The polarization component is oscillating perpendicular to the incident surface) and the P wave (the polarized component oscillating in parallel with the incident surface by the transmitted polarization component).

  The S wave beam passes through the first dove prism 40 and is incident on the polarization beam splitter 42 for multiplexing, and the optical path length from the polarization beam splitter 39 to the polarization beam splitter 42 for multiplexing is constant.

  The P-wave beam passes through the second dove prism 41 for adjusting the optical path length and is incident on the polarization beam splitter 42 for multiplexing. For the P-wave beam, the optical path length from the polarizing beam splitter 39 to the combining polarizing beam splitter 42 can be adjusted by appropriately moving the second dove prism 41 along the optical path direction. .

  Here, the optical path length of the P wave is adjusted to be longer than the optical path length of the S wave. The S-wave beam and the P-wave beam having different optical path lengths are superimposed (combined) in a state where the phase difference is shifted in the multiplexing polarization beam splitter 42, passes through the fiber collimator 43, and then is polarized by the polarization delay unit 21. Is emitted toward the coupler 22.

  20% of the beam incident on the coupler 22 passes through the fiber collimator 23, the biaxial galvano scanner 24, the objective lens 25, and the aspherical ophthalmic lens 26 as a probe beam. The cornea, retina, choroid, etc.) and the remaining 80% is incident on the phase calibration unit 31.

  The measurement beam irradiated and reflected on the measurement object 7 and the calibration beam reflected by the calibration mirror 34 of the phase calibration unit 31 return to the probe arm 13 and pass from the coupler 22 through the polarization controller 55 to the polarization separation detection unit. 14, passes through the fiber collimator 56, and enters the polarization beam splitter 54 for multiplexing.

  On the other hand, the reference beam incident on the reference arm 12 from the coupler passes through the fiber collimator 46, the mirror 47, the mirror 48, the fiber collimator 49, and the polarization controller 50, enters the polarization separation detection unit 14, and the fiber collimator 57. Then, the linearly polarized light is converted into linearly polarized light by the linear polarizer 58 and is incident on the polarization beam splitter 54 for multiplexing.

  The measurement beam and the reference beam incident on the beam splitter 54 for multiplexing are superimposed and interfere with each other to form interference light. This interference light is transmitted between the first polarization beam splitter 63 and the second beam on the first optical path 61. To the second polarization beam splitter 64 on the optical path 62.

  The interference light incident on the first polarization beam splitter 63 on the first optical path 61 is separated into a horizontal polarization component and a vertical polarization component in the first polarization beam splitter 63, and a horizontal balance polarization detector 68 is respectively provided. Are detected by the vertical balance polarization detector 70.

  The interference light incident on the second polarization beam splitter 64 on the second optical path 62 is separated into a horizontal polarization component and a vertical polarization component in the second polarization beam splitter 64, and a horizontal balance polarization detector 68, respectively. Are detected by the vertical balance polarization detector 70.

  For the interference light corresponding to the P wave and the horizontal polarization component of the interference light detected by the horizontal balance polarization detector 68, a first tomographic image is obtained, and the interference light corresponding to the S wave and the interference light are obtained. For the horizontal polarization component, a second tomographic image is obtained.

  FIG. 2 is an OCT tomographic image in which the macular retina is the measurement object 7. The two OCT tomographic images are as shown in the upper and lower parts of FIG. The tomographic image in FIG. 2A is an image corresponding to the S wave and is formed on the shallow side of the measurement object 7. 2A is a second tomographic image corresponding to the P wave, and is formed on the deep side of the measurement object 7. FIG.

  In the vertical balance polarization detector, the first tomographic image is obtained by the interference light corresponding to the S wave received by one of the light receivers and the vertical polarization component of the interference light, and is received by the other light receiver. A second tomographic image based on the interference light corresponding to the P wave and the vertical polarization component of the interference light is obtained.

  These two OCT tomographic images are as shown in the upper and lower parts of FIG. The tomographic image shown in FIG. 2B is a first tomographic image corresponding to an S wave, and is formed on the shallow side of the measurement object 7. The tomographic image shown at the bottom of FIG. 2B is a second tomographic image corresponding to the P wave, and is formed on the deep side of the measurement object 7.

  As described above, the wavelength scanning OCT2 of the present invention constitutes a multi-contrast OCT, and four OCT tomographic images are obtained. Incidentally, a further feature of the apparatus of the present invention is that a phase calibration unit 31 is provided on the probe beam 13 side. This will be described below.

  In the polarization delay unit 21, a part (about 4.4%) of the P-wave beam out of the linearly polarized incident beam is reflected without passing through the polarization beam splitter 39 to be converted into an S-wave beam. A phenomenon occurs that cannot be separated incompletely.

  That is, the polarization beam splitter 39 of the polarization delay unit 21 ideally transmits the P wave and reflects the S wave, but in reality, a commercially available polarization beam splitter cannot completely separate linearly polarized light, P wave is mixed in the reflected light. Accordingly, the polarization delay unit 21 behaves like a Mahahender interferometer having optical path length differences zd with respect to the P wave, and generates a calibration signal.

  That is, when the P-wave beam transmitted through the polarization beam splitter 39 and the P-wave beam mixed in the reflected light are superimposed by the polarization beam splitter 42, the P-wave beam functions like a Maha-Zehnder interferometer to generate interference light. This functions as a calibration signal.

  When the polarization beam splitter is configured to reflect the P wave and transmit the S wave, it utilizes the fact that the S wave is partially reflected and mixed into the P wave, resulting in an optical path length difference of the S wave itself. Then, interference light may be generated and function as a calibration signal.

  The interference light that is split from the coupler 22 of the probe arm 13 and functions as a calibration signal enters the phase calibration unit 31, is reflected by the calibration mirror 34 through the fiber collimator 22 and the lens 33, and functions as the reflected calibration signal. The interference light that passes through the coupler 22 passes through the polarization controller 55 and enters the polarization separation detection unit 14.

  The interference light from the phase calibration unit 31 is superposed and interfered with the reference beam by the beam splitter 54 for multiplexing in the polarization separation detection unit 14, divided and sent to the first optical path 61 and the second optical path 62. The first polarizing beam splitter 63 and the second polarizing beam splitter 64 respectively further separate the light into a horizontal polarization component and a vertical polarization component. The horizontal polarization component and the vertical polarization component are detected by the horizontal balance polarization detector 68 and the vertical balance polarization detector 70, respectively.

  2A and 2B, the tomographic images respectively detected by the horizontal balance polarization detector 68 and the vertical balance polarization detector 70 by the interference light functioning as a calibration signal generated by the phase calibration unit 31 respectively. In FIG. 2, a calibration (calibration) line image (an image indicated as “Calibration signal” in FIG. 2) is formed at an intermediate position in the vertical direction.

(Image processing)
The configuration and program of the image processing means for processing the four OCT tomographic images obtained as described above using the computer 3 will be described below.

  Measurement data obtained by the wavelength scanning OCT having the above configuration is input to a computer 3 used as an image processing apparatus. The computer 3 is a normal computer, and includes an input unit 81, an output unit 82, a CPU 83, a storage device 84, and a data bus 85 as shown in FIG.

  The program according to the present invention is a program stored in the storage device 84 of the computer 3 and functions as a means for more vividly processing an image obtained by the wavelength scanning OCT 2 input to the computer 3. It is a program to let you. By mounting this program on the computer 3, the Jones matrix OCT system 1 according to the present invention is a system including means for processing an image more clearly.

  Conventionally, in a Jones matrix OCT (an OCT for obtaining an image by processing an interference signal obtained by OCT using a Jones matrix means), the OCT image has four elements of the Jones matrix (four OCT tomographic image signals). Was obtained by averaging the squared intensity.

  In addition, Doppler OCT is performed by averaging Doppler phase shift signals of four elements (four OCT tomographic image signals). With such conventional means, it is inevitable that the sensitivity decreases due to the power of the light source being separated into four OCT tomographic image signals, that is, four Jones matrix elements.

  In order to solve this problem, in the present invention, four elements of the Jones matrix (four OCT tomographic image signals) are in a complex amplitude state (a portion not including time in the case of complex display of the OCT tomographic image signals). It is characterized by adopting a novel signal processing that is coherently coupled (coherent coupling).

  In the present invention, by adopting such a configuration, the tomographic image signal is converted into four elements of the matrix (four OCT tomographic image signals) by the scattered OCT signal and the Doppler OCT signal measured in the wavelength scanning type OCT2. It can be obtained by coherent coupling.

Specifically, the coherent combination E _out (z) of the matrix elements is represented by four OCT tomographic image signals E ⁽¹⁾ _outA (z), E ⁽¹⁾ _outB (z), which are _{expressed by depth} (z). As a matrix of E ⁽²⁾ _outA (z) and E ⁽²⁾ _outB (z), it is expressed as shown in the following equation (Expression 7).

In (Equation 17), A is an OCT tomographic image signal corresponding to horizontal polarization, (1) corresponds to a P wave component, and (2) is an OCT tomographic image signal corresponding to an S wave component. It means that. Therefore, E ⁽¹⁾ _outA (z) is an OCT tomographic image signal corresponding to horizontal polarization and P wave component, and E ⁽²⁾ _outA (z) is an OCT tomographic image corresponding to horizontal polarization and S wave component. Signal.

In (Equation 17), B is an OCT tomographic image signal corresponding to vertical polarization, (1) corresponds to a P wave component, and (2) is an OCT tomographic image signal corresponding to an S wave component. It means that. Therefore, E ⁽¹⁾ _outB (z) is an OCT tomographic image signal corresponding to vertically polarized light and P wave component, and E ⁽²⁾ _outB (z) is an OCT tomographic image corresponding to vertically polarized light and S wave component. Signal.

  In this manner, OCT tomographic image signals corresponding to the horizontally polarized light and P wave component, the horizontally polarized light and S wave component, the vertically polarized light and P wave component, and the vertically polarized light and S wave component, respectively, are shown as a matrix.

_{Here, θ 1, θ 2, θ} 3 is not dependent on the depth direction, E ^(1), respectively E for ^{_{outA (z) (2) outA}} (z), E (1) outB (z), E ( ²⁾ The relative phase offset of _outB (z), and the assumption that the change in the depth direction of the birefringence component of the phase can thus be ignored is the normal non-polarization intensity OCT This is equivalent to the approximation that the birefringence of the object to be measured is sufficiently small, which is also assumed in OCT).

In this coherent coupling, θ ₁ , θ ₂ , and θ ₃ are expressed by the following equation (Equation 18).

Here, Σ with z attached below is the sum of pixels in the depth direction. Arg is a function that gives the phase component of the argument. That is, Arg [x] is a function that gives a phase component (also called declination) of a complex number x. E ⁽¹⁾ _outA (z) ^* represents the complex conjugate of E ⁽¹⁾ _outA (z) (the same _{applies to} the other expressions below). Here, ≡ means that the symbol on the left side is defined on the right side, and is used in the same meaning as = (equal).

For the coherent combining, the average value (superscript line E _out (z)) is coherent combining E _out of the matrix elements (z), theta _1, theta _2, by using the theta _3, the following equation (19) Defined by According to this equation, an intensity OCT image can be obtained by taking the average value of each element of the Jones matrix. Here, the upper line of E _out (z) means an average value.

  Since this synthesized signal is summed with the random phase of the complex OCT signal, it is enhanced compared to the conventional method that takes the sum of the square intensities including the random phase signal. Sensitivity can be obtained, and even when applied to Doppler phase shift measurement described later, higher resolution can be obtained.

  As described above, a scattered OCT image (an OCT image based on scattered reflected light from a measurement object (ordinary intensity OCT image)) based on a synthesized signal that can obtain enhanced sensitivity is a coherent matrix component. And obtained as the following equation (Equation 20).

  Further, by using a plurality of B-scan data and taking their coherent average, a higher-quality scattered OCT image can be obtained. This is shown by the following equation (Equation 21).

  FIG. 4 (a) shows an OCT image obtained as a result of the above formula for the choroidal blood vessel located deeper than the retina by the Jones matrix OCT system 1 according to the present invention.

  This image has clearer contrast than the contrast image of the same choroidal blood vessel obtained by angiography in the form of conventional indocyanine green fluorescence fundus angiography (ICGA) shown in FIG. It is clear that

  Next, a configuration in which the present invention is applied to Doppler measurement (Doppler OCT) will be described below.

  Particularly in Doppler measurement, the Doppler signal (Doppler phase shift) is a spectral interference obtained by two A-scans (first A-scan A1 and second A-scan A2 obtained by two B-scans). Since it is obtained by the phase difference of the signal, the jitter of the spectrum interference signal acquired by OCT (Jitter: fluctuation of the signal waveform in the time axis direction and image disturbance caused by the fluctuation) directly affects the data error in Doppler measurement. Affect.

  For this purpose, the A-line data at the same location at different times is obtained by repeating the B-scan.

In general, the raw data of the Doppler phase shift [Delta] [phi (z) of biological measurement object is obtained from Δφ (z) = (4πτ / λ C) nv z (z) + φ b.

Here, λ _C is the center wavelength of the light source, n is the refractive index of the measurement object, v _z (z) is the velocity optical axis direction (z) component of the measurement object that is the measurement object, and φ _b is the measurement object. This is a constant offset (bulk offset) due to the bulk motion (overall movement) of. τ is a time difference between two A-scans, and here is a time between different B-scans.

  In the present invention, the raw Doppler shift Δφ (z, j) is a coherent component conjugated to each other and is defined as the following equation (Equation 22).

Here, Δφ (z, j) represents a Doppler shift between the A-line in the j-th B-scan and the A-line in the j + 1-th B-scan. ^* Indicates a complex conjugate. The upper line of E _out (upper line) shows an average value.

The bulk offset Δφ _b (j) can be written by the following equation (Equation 23) using an integral value in the depth direction.

  Sensitivity can be improved by measuring the same point by a plurality of (m) B-scans. That is, it is as shown in the following equation (Equation 24).

Here, m ₀ is a parameter of the start B-scan, W (z, j) is an intensity mask, defined by the following equation (Equation 25), and ε ² is the noise floor (minimum noise) of the OCT image It is. This equation, window function W when the amount Doppler phase shift is greater than a predetermined magnitude epsilon ² is set to 1, if a predetermined smaller than the size is set to 0 so as not to added to treat the data as noise, that Meaning.

  FIG. 4B shows an OCT image obtained as a result of applying the Jones matrix OCT according to the present invention to the Doppler measurement and improving the sensitivity by the above formula for the blood vessels of the choroid existing deeper than the retina.

  This image has improved sensitivity compared to the same choroidal blood vessel contrast image obtained by angiography in the form of conventional indocyanine green fluorescence fundus angiography (ICGA) shown in FIG. 4 (c). It is clear.

  When m = 1, the above (Expression 24) becomes the following expression (Expression 26).

  When displaying a Doppler image, the following equation (Expression 27) indicating the square strength of the Doppler phase shift is used.

(Proofreading)
Since the wavelength scanning type light source scans while changing the wavelength with time, it is between the wavelength scanning of the light source (timing of wavelength change) and the timing of collecting data as a spectral interference signal by the photodetector. Jitter (caused by mismatch) (Jitter: fluctuation of signal waveform in time axis direction and image disturbance caused by the fluctuation) becomes a problem.

  This jitter results in a random shift in spectral sampling, resulting in jitter of spectral interference signals acquired with OCT.

  In the present invention, a calibration signal is generated using the polarization delay unit 21 (polarization delay unit), and jitter (wavelength scanning of the light source and A) occurring in spectrum sampling between A-lines in different B-scans. -Fluctuation of the time position of the phase in synchronization between lines), and using the calibration signal, the image processing program of the present invention causes the computer to function as correction means and corrects the phase, thereby stabilizing the phase. Is possible.

  The calibration signal is generated by the wavelength scanning type OCT2 according to the present invention as follows. As described above, the polarization beam splitter 39 of the polarization delay unit 21 ideally transmits the P wave and reflects the S wave.

However, in practice, a commercially available polarizing beam splitter cannot completely separate linearly polarized light, and P waves are mixed in reflected light. Thus, the polarization delay unit 21 acts like a Mahatsuenda interferometer having an optical path length difference z _d each other in the P-wave, to generate an interference light that serves as a calibration signal.

  This calibration signal is reflected by the calibration mirror 34 of the phase calibration unit 31 through the coupler 22, and again through the coupler 22 by the horizontal balance polarization detector 68 and the vertical balance polarization detector 70 of the polarization separation detection unit 14. Detected.

Horizontally balanced polarization detector 68 and vertically balanced polarization detector 70 in the tomographic image based on the calibration signal are detected (calibration (line image for calibration)), the depth of which corresponds to the optical path length difference z _d Although shown in position, this position is always exactly the middle of the multiplexed signal of the two input polarizations, as shown in FIGS. 2 (a) and 2 (b), respectively. This is used to correct jitter (phase position fluctuation) in synchronization between the wavelength scanning of the light source and the A-line.

Jitter correction is as follows. For example, assuming that the first A-line of the B-scan is the reference, its spectral intensity I _r (j), and other A-line data to be calibrated is I _c (j), these are given by ).

Here, E _r (j) and E _c (j) are spectral interference components due to light reflected and transmitted through the polarization delay unit, and j is a parameter of the B-line. * Represents convolution integral (convolution), and β _j is a relative shift amount of the spectrum.

  In estimating the shift amount of the spectrum, first, the complex conjugate of the Fourier transform of the reference A-line signal as the reference and the Fourier transform of the A-line signal to be calibrated is expressed by the following equation (Equation 29). Accumulate as follows.

  Here, F [] represents a Fourier transform, and the superscript * represents a complex conjugate. When (Equation 29) is subjected to inverse Fourier transform, the following equation (Equation 30) is obtained.

Here, a symbol in which “x” is shown in a circle is a correlation calculation (correlation). The following equation _(31), since the autocorrelation of _I r (j), to take a maximum at j = 0, the total maximum value, the place of _j = -β j.

Therefore, the amount of spectral shift β _j can be determined from the point where the signal of the above equation becomes maximum. In order to increase the accuracy of detection, the resolution in the frequency space can be improved by performing the zero transform for expanding the outside of the effective data with the data of 0 value and performing the Fourier transform in the Fourier transform. For example, if the area is enlarged 16 times, the shift amount can be detected in the frequency space with an accuracy of 1/16 pixel.

  In the case of SS-OCT, the light source performs a linear scan with respect to the frequency of light. In data processing, it is necessary to convert the data into wave number linear data. This procedure is called rescaling and is converted using a table. Jitter correction and rescaling can be performed with a single conversion by preparing and preparing the above-mentioned shift amount incorporated in the frequency-wavenumber rescaling table in advance.

  As mentioned above, although the form for implementing the Jones matrix OCT system which concerns on this invention and the program for image-processing the measurement data obtained by this OCT was demonstrated based on the Example, this invention is limited to such an Example. It goes without saying that various embodiments exist within the scope of the technical matters described in the claims.

  The Jones matrix OCT system according to the present invention and the program for image processing the measurement data obtained by the OCT are optimally used for non-invasive ophthalmic diagnostic apparatuses, and are used for pseudoangiography and quantitative evaluation of choroidal blood vessels. It is extremely useful for the very early diagnosis of diabetic retinopathy.

1 Jones Matrix OCT System 2 Wavelength Scanning OCT
3 Computer 6 Light Source 7 Measurement Object 8 Optical Isolator 11 Fiber Coupler 12 Reference Arm 13 Probe Arm 14 Polarization Separation Detection Unit 17 Fiber 20 Polarization Controller 21 Polarization Delay Unit 22 Coupler 23 Fiber Collimator 24 Biaxial Galvano Scanner 25 Objective Lens 26 Non Spherical ophthalmic lens 31 Phase calibration unit 32 Fiber collimator 33 Lens 34 Mirror (calibration mirror)
37 Fiber Collimator 38 45 ° Linear Polarizer 39 Polarizing Beam Splitter 40 First Dove Prism 41 Second Dove Prism 42 Polarizing Beam Splitter for Multiplexing 43 Fiber Collimator 46 Fiber Collimator 47 Mirror 48 Mirror 49 Fiber Collimator 50 Polarization Controller 54 Beam splitter for multiplexing 55 Polarization controller 56 Fiber collimator 57 Fiber collimator 58 45 ° linear polarizer 61 First optical path 62 Second optical path 63 First polarizing beam splitter 64 Second polarizing beam splitter 67 Fiber collimator 68 Horizontal Balance Polarization Detector 69 Fiber Collimator 70 Vertical Balance Polarization Detector 74 Fiber Collimator 75 Fiber Collimator 81 Input Unit 82 Output Unit 83 CP
84 Storage device 85 Data bus 93 OCT
94 Light source 95 Collimating lens 96 Beam splitter 97 Objective lens 98 Object to be measured 99 Objective lens 100 Reference mirror 101 Condensing lens 102 Photo detector

Claims (7)

A light source;
A polarization separation unit that separates a light beam from the light source into P-polarized light and S-polarized light;
A polarization delay unit that gives a phase difference relative to the P-polarized light and the S-polarized light from the polarization separation unit, and superimposes the P-polarized light and the S-polarized light having a phase difference;
An objective unit that irradiates the measurement object with the light beam from the polarization delay unit and receives reflected light from the measurement object;
A detection unit for causing a light beam from the objective unit and a light beam from the light source to interfere with each other and separating them into vertical polarization and horizontal polarization;
A calculation unit that receives a signal from the detection unit and generates a tomogram based on the following (Equation 1) to (Equation 3);
Having
Jones matrix OCT device.
In ( _Expression ¹⁾ , E ⁽¹⁾ _outA (z) is an OCT cross-sectional image signal corresponding to the horizontally polarized light and P polarized light, and E ⁽²⁾ _outA (z) is the horizontally polarized light and S OCT cross-sectional image signal corresponding to polarized light, E ⁽¹⁾ _outB (z) is the OCT cross-sectional image signal corresponding to the vertical polarization and P-polarized light, and E ⁽²⁾ _outB (z) is the vertical It is an OCT cross-sectional image signal corresponding to polarized light and S polarized light.
The calculation unit is
The Doppler image can be calculated and displayed by the following equation (Equation 4),
By measuring the same point by a plurality of (m) B-scans, the Doppler shift Δφ (z, j) expressed by the following equation (Equation 5) is applied to the Doppler measurement, and the sensitivity of the Doppler measurement Improve,
The Jones matrix OCT apparatus according to claim 1.
In the above (Equation 5), m ₀ is a parameter of the B-scan at the start, W (z, j) is an intensity mask defined by (Equation 6) below, and ε ² is an OCT image Noise floor (minimum noise).
The calculation unit is
While calculating the OCT image by the scattered reflected light from the measurement object based on the following equation (Equation 7) using a coherent matrix component,
Using a plurality of times of B-scan data and taking a coherent average of the matrix components, a scattered OCT image is calculated based on the following equation (Equation 8).
The Jones matrix OCT apparatus according to claim 1.
The polarization delay unit includes a phase calibration unit having a calibration mirror,
The calibration mirror reflects the P-polarized light or the S-polarized interference light in the polarization delay unit, and sends the reflected light to the detection unit.
The calculation unit corrects jitter in synchronization between wavelength scanning of the light source and A-scan using a calibration signal generated by superimposing the interference light and the light beam from the light source in the detection unit.
The Jones matrix OCT apparatus according to any one of claims 1 to 3. A light source, a polarization separation unit that separates a light beam from the light source into P-polarized light and S-polarized light, and a P that has a phase difference relative to the P-polarized light and the S-polarized light from the polarization separation unit. A polarization delay unit that superimposes polarized light and S-polarized light, an objective unit that irradiates a measurement target with a light beam from the polarization delay unit, receives reflected light from the measurement target, a light beam from the objective unit, and a light source The calculation unit of the Jones matrix OCT apparatus includes: a detection unit that causes the light beam to interfere and separates and detects vertical polarization and horizontal polarization; and a calculation unit that generates a tomogram based on a signal from the detection unit The program installed in
The calculation unit receives a signal from the detection unit and functions as an image processing unit that generates a tomographic image based on the following (Equation 9) to (Equation 11).
program.
In ( _Equation 9), E ⁽¹⁾ _outA (z) is an OCT cross-sectional image signal corresponding to the horizontal polarization and the P polarization, and E ⁽²⁾ _outA (z) is the horizontal polarization and the S OCT cross-sectional image signal corresponding to polarized light, E ⁽¹⁾ _outB (z) is the OCT cross-sectional image signal corresponding to the vertical polarization and P-polarized light, and E ⁽²⁾ _outB (z) is the vertical It is an OCT cross-sectional image signal corresponding to polarized light and S polarized light.
The calculation unit can calculate and display a Doppler image by the following equation (Equation 12), and measure the same point by a plurality of (m) B-scans to obtain the following equation (Equation 13) The Doppler shift Δφ (z, j) represented by 13) is applied to Doppler measurement to function as an image processing means for improving the sensitivity of Doppler measurement.
The program according to claim 5.
In (Equation 13), m ₀ is a parameter of the B-scan at the start, W (z, j) is an intensity mask defined by (Equation 14) shown below, and ε ² is an OCT image Noise floor (minimum noise).
The calculation unit calculates an OCT image based on scattered reflected light from the measurement object based on the following equation (Expression 15) using a coherent matrix component, and uses a plurality of B-scan data. By taking a coherent average of the matrix components, the scattered OCT image is made to function as an image processing unit that is calculated based on the following equation (Equation 16).
The program according to claim 5.