OCT imaging system

Abstract

An apparatus for performing a method for determining and / or correcting a confocal function in an OCT imaging system, characterized in that the apparatus determines the confocal function and / or corrects the influence of the confocal function and / or corrects the confocal function, achieves the object of the present invention to overcome the drawbacks of the prior art. A system and method are further disclosed.

Description

[Technical Field]

[0001] The present invention relates to an apparatus, a system and a method. [Background technology]

[0002] Optical coherence tomography (OCT) is an optical imaging technique for cross-sectional imaging of tissue structures.

[0003] OCT measures the scattering profile of a sample along the OCT beam, each scattering profile called an A-line.

[0004] The OCT beam is scanned laterally across the sample, and a two-dimensional cross-sectional image called a B-scan is constructed from multiple A-lines.

[0005] Time-domain OCT (TD-OCT) requires mechanical scanning of the optical path length between the sample arm and the reference arm.

[0006] However, in Fourier-domain OCT (FD-OCT), the optical path length difference between the sample arm and the reference arm is not mechanically scanned.

[0007] Instead, all A-lines are obtained in parallel for all points along the axis by Fourier transforming the wavelength scan of a scanning light source in scanned-source OCT (SS-OCT) or by resolving the spectrum of a superluminescent diode (SLD) with a spectrally resolved and / or line-scan camera in spectral-domain OCT (SD-OCT).

[0008] The A-line I(z) along z can be written as follows:

[0009]

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[0010] where the roll-off r(z), the confocal function h(z), the backscattering coefficient α, and the attenuation coefficient μ t Let's say.

[0011] To date, OCT technology has been used clinically primarily to image tissue structures and measure geometric parameters such as layer thickness and distances between structures.

[0012] In addition, OCTA evaluates speckle and phase variations to evaluate and display blood flow within blood vessels. Summary of the Invention [Problem to be solved by the invention]

[0013] SUMMARY OF THE INVENTION An object of the present invention is to overcome the drawbacks of the prior art.

[0014] The object of the invention is achieved by the features of the independent claims. [Means for solving the problem]

[0015] According to the present invention, the OCT beam is focused onto the sample using a lens, so that the A-line scattering profile along z is calculated as the confocal function h(z) of the imaging system, i.e., the focal position z f and the Rayleigh length Z R It was found to depend on

[0016]

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[0017] It was also found that when quantitative parameters such as the attenuation coefficient are calculated from the A-line, it is necessary to correct for the influence of the confocal function.

[0018] Determining the confocal function requires knowledge of the Rayleigh length and focus of the optical system. Particularly for in vivo imaging of the human retina, determining the focus position is complicated by specimen movement and contraction.

[0019] In accordance with the present invention, it has further been found that common properties of two A-lines of the same sample position acquired at different focal positions can be used to determine the confocal function.

[0020] By taking the ratio of two A-lines acquired at the focus offset, for example, by applying the following equation to the ratio, the depth of focus and the Rayleigh length can be determined (see Non-Patent Documents 1 to 4).

[0021]

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[0022] Here, Δz f =z f2 -z f1 is the axial displacement between the two focal depths, z f1 and z f2 are the focal positions of the first and second A-lines, respectively. The three parameters z f1 , Z R , and Δz f can all be set as fitting parameters. For simplicity, one or two of the fitting parameters can be fixed.

[0023] After that, the focal position corrected A-lines can be calculated and the tissue attenuation coefficients can be determined, for example, according to the procedure described in Non-Patent Document 6.

[0024] To acquire two A-lines with different focal depths, it is necessary to change the focus between the acquisition of the two A-lines, which is typically done by axially moving the sample or the objective lens of the system (see Non-Patent Documents 1 to 4).

[0025] For example, in human retinal imaging, the confocal function cannot be reliably determined if the eye contracts during focus changes, so it is crucial to introduce focus shifts very quickly, for example between two consecutive B-scans, or, even better, to extract the confocal parameters from a single B-scan.

[0026] In particular, the following two alternative technical solutions are disclosed:

[0027] Solution 1. By adding a dispersive element to the imaging system (or by appropriately modifying the design of an existing lens), we can create different focal positions as a function of wavelength. Splitting the full-spectrum scan into subspectra before Fourier transformation simultaneously yields low-resolution A-lines with different focal positions. By splitting these A-lines (e.g., according to Equation 1.3) and fitting the ratio with the corresponding focal parameters, we can determine the wavelength-dependent confocal function.

[0028] Solution 2: Alternatively, by adding an active optical element to the system that can change the refractive power in milliseconds, two A-scans with different focal depths can be acquired in milliseconds. Such fast focusing is necessary to reliably determine the confocal function using the ratio fitting method, even in the presence of eye movement and contraction shift.

[0029] This active element can be either a transmissive element such as a liquid lens, a motorized lens, or any other electro-optical transmissive element that allows fast switching of optical power. Alternatively, a deformable mirror in reflective mode can be used to introduce fast focus shifts.

[0030] By dividing the A-lines acquired at different focus positions and fitting the ratio with the corresponding focus parameters, the depth of focus and the Rayleigh length can be determined.

[0031] The resulting confocal parameters can be used to optimize the depth of focus for a desired task or to correct the A-scan for the confocal function, which can be achieved by dividing the original A-scan by the confocal function.

[0032] The corrected A-scans can be used to calculate attenuation coefficients and extract other quantitative tissue parameters.

[0033] In solution 1, by individually confocal function correcting the subspectral A-lines, and therefore the B-scans, different confocal function corrected B-scans at different wavelengths can be obtained with lower resolution than the full-spectrum B-scan.

[0034] To obtain a high-resolution, full-spectrum, confocal-function-corrected B-scan, the individual confocal-function-corrected sub-B-scans are converted to a spectrum using an inverse Fourier transform, the complex-valued sub-spectra are coherently added, and then the full spectrum is converted back to an image using a Fourier transform (see Figure 4).

[0035] Alternatively, the full-spectrum B-scan may be corrected with the average confocal function of the subspectra.

[0036] Although ratio fitting methods for manually defocused A-lines have already been reported in Non-Patent Documents 1 to 4, this specification proposes and discloses various methods for obtaining defocused A-lines without manual adjustment.

[0037] 1. From a single broadband OCT A-scan, the spectrum can be split into subbands that allow the calculation of A-lines at different wavelengths and subsequently at different focal depths for the purpose of fitting to a confocal function. Focal parameters can be determined without additional data acquisition procedures.

[0038] 2. As mentioned above, fast automated focus shifting between consecutive A-scans and B-scans allows for accurate extraction of focus parameters, resulting in the correction of the confocal function with all the advantages mentioned above.

[0039] By simultaneously acquiring A-scans at different focal points at high speed, the ratio fitting method can be applied to samples or objects with unstable focal positions, and the confocal function can be determined as a result.

[0040] Correction of the confocal function is essential to determine absolute tissue scattering data; otherwise, measurement results will be erroneous and likely to be less reproducible (see Figure 3).

[0041] The following details and technical features are options:

[0042] Online display of focus parameters (mainly focus position) can be used as online focus control to enable optimal image quality during OCT or OCTA data acquisition.

[0043] It is well known that the focal position has a significant impact on the reproducibility of capillary plexuses, especially in OCTA data.

[0044] This may be achieved in the form of software guidance for the user to manually optimize the focus setting, or by the implementation of automated focus adaptation or correction algorithms, or by means of providing the focus position to the user in a graphical user interface.

[0045] Preferably, a high-resolution confocal corrected image is reconstructed based on coherent addition of correction profiles of sub-spectral bands.

[0046] The present disclosure relates in particular to OCT, confocal functions, focus, Rayleigh length, active optical elements, dispersive elements, quantitative OCT, focus control and attenuation coefficients. [Brief explanation of the drawings]

[0047] [Figure 1] Figure 1 shows the integration of active optical or electro-optical elements to generate a focus shift at the retina for successive A-scans or B-scans. Alternatively, passive dispersive elements can be added to introduce a chromatic focus shift at the retina. In this example, the focus shift elements are implemented in the fixed beam before scanning, but they could also be implemented in the scanning beam. [Figure 2] Figure 2 shows two B-scans obtained from a full-spectrum B-scan by spectrally splitting the frequency data before the Fourier transform. A focus shift is observed between λ1 and λ7 due to chromatic aberration introduced in the imaging system. The image on the right shows the focal planes of seven different spectral windows. From these focal planes, the average focal plane of the full spectrum can be determined. The scale bar is 200 μm. [Figure 3] The left frame (A) in Figure 3 shows a series of A-scans acquired at the same location but with different focus settings, while the right frame (B) shows the same data but with the confocal function corrected using a ratio fitting method for manually shifted A-scans. Note that over a wide range, the data follow the Beer-Lambert law with exponential decay. The legend indicates the expected depth of focus from the sample surface (see non-patent document 4). [Figure 4] FIG. 4 is a schematic diagram of the processing steps. [Figure 5] FIG. 5 is a further schematic illustration of the in vivo image processing steps of retinal imaging. DETAILED DESCRIPTION OF THE INVENTION

[0048] 1 shows an apparatus including at least an interferometer 5, a light source 3, a detection unit 4, a reference mirror 2, and an element 1. Element 1 may be a dispersive element or an active lens system.

[0049] FIG. 1 also shows an optical system, in particular an OCT imaging system, comprising the device and further optical means for forming an optical path to the eye 21 .

[0050] The optical system is part of a system consisting of the optical system and the eye 21 .

[0051] FIG. 1 shows the integration of an active optomechanical or electro-optical element 1 to generate a focus shift on the retina 20 for successive A-scans or B-scans.

[0052] It is also possible to add a passive dispersive element that introduces chromatic aberration at the retina. In this example, the focus-shift element 1 is implemented in the fixed beam before scanning, but it could also be implemented in the scanning beam.

[0053] Figure 2 shows two B-scans 6, 7 obtained from a full-spectrum B-scan by spectrally splitting the frequency data before Fourier transformation.

[0054] Due to the introduction of chromatic aberration into the imaging system, a focus shift is observed between λ1 and λ7.

[0055] The focal planes obtained for the seven different spectral windows can be seen in image 8 on the right. From these focal planes, the average focal plane for the entire spectrum can be determined. The scale bar is 200 μm.

[0056] Figure 3 (see Non-Patent Document 4) shows (A) A series of A-scans acquired at the same location but with different focus settings. (B) The same data, but with the confocal function corrected using the ratio fitting method on a manually shifted A-scan. Note that over a wide range, the data follows the Beer-Lambert law with exponential decay. The legend indicates the expected focal depth from the sample surface.

[0057] FIG. 4 is a diagram of a confocal function corrected high resolution OCT image 22.

[0058] Herein, we disclose the development and evaluation of a method for determining the confocal function by spectral partitioning.

[0059] The principle of the spectral splitting method was demonstrated in homogeneous intralipid samples with different scatterer concentrations. The feasibility of the method of manually introducing focus shifts was demonstrated in different samples (including layered samples) and two subjects (see Non-Patent Document 4).

[0060] Some features described in this disclosure may be implemented in the following manner. 1) Depth of focus indicator in GUI. 2) 3D focus tracking (adding depth tracking and adapting focus for perfect overlay at different time points), which requires motorized focus control. 3) Focal depth control in angio-OCT. 4) For the reproducibility of measurements that depend on the tissue scattering coefficient, i.e., the intensity of backscattered light, correction of the confocal function is essential to obtain quantitative results, unless, of course, the measured OCT intensity can be normalized to the signal of a structure at a very similar z-position, and only the ratio of the measured intensities is relevant.

[0061] Several publications have shown that in glaucoma, tissue deterioration of the RNFL leads to changes in the attenuation coefficient, which is preceded by a thinning of the geometric thickness of the RNFL (see, for example, Non-Patent Document 5).

[0062] If this hypothesis is correct, systematic evaluation of reliably measured tissue attenuation coefficients may play an important role in the early diagnosis of glaucoma.

[0063] In OCT oximetry, the signal at different z-positions (before and after passing through the blood vessel) is evaluated with respect to its spectral range, so here too the wavelength-dependent position of the focal point is important.

[0064] The broadband OCT system hardware optionally allows the use of such algorithms as add-on software tools with additional diagnostic capabilities.

[0065] If the existing color focal shift is already sufficient, such a tool can be used.

[0066] A high-resolution OCT system with extended spectral range may be comprised of one or more technical features herein.

[0067] In particular, additional software modules can be implemented to take advantage of the expanded spectral range as described.

[0068] 4 and 5 show a schematic representation of a method comprising the following steps: creating OCT images 6, 7 for different λ with different focal positions; determining a confocal function; generating OCT images for different λ with different focal positions and confocal functions; correcting the confocal function to obtain confocal function corrected OCT images for different λ; applying an inverse fast Fourier transform to obtain spectra for different λ; coherently adding a complex spectrum to the spectrum for each different λ; generating a spectrum for the full range of λ; applying a fast Fourier transform to the full range λ spectrum to generate a confocal function corrected high-resolution OCT image 22; Equipped with.

[0069] The dashed ellipses indicate the spectra for different λ. [Explanation of symbols]

[0070] 1. Dispersive element or active lens system 2 Reference Mirrors 3 light source 4 Detection Unit 5 Interferometer 6 λ1 B-scan (OCT image) 7 λ7 B-scan (OCT image) 8 Focal plane images for seven different spectral windows λ1-λ7 9 OCT images for different λ with different focal positions 10. Determination of Confocal Function 11 OCT images for different λ with different focal positions and confocal functions 12. Confocal Function Correction 13 Confocal function corrected OCT images for different λ 14 Inverse FFT 15 Spectra for different λ 16 Coherent addition of complex spectra 17 Full-range λ spectrum 18 FFT 19 Confocal function corrected high resolution OCT images 20 Retina 21 Eyeball 22 Confocal function corrected high resolution OCT images 23 Confocal Lens Arrangement λ wavelength OCTA Optical Coherence Tomography Angiography GUI Graphical User Interface RNFL Retinal nerve fiber layer [Prior art documents] [Non-patent literature]

[0071] [Non-Patent Document 1] Dwork, Nicholas, Gennifer T. Smith, John M. Pauly, and Audrey K. Ellerbee Bowden. 2016. “Automated Estimation of OCT Confocal Function Parameters from two B-Scans.” In Conference on Lasers and Electro-Optics, AW10.4. San Jose, California: Optical Society of America [Non-Patent Document 2] Stefan, S., K. S. Jeong, C. Polucha, N. Tapinos, S. A. Toms, and J. Lee. 2018. “Determination of confocal profile and curved focal plane for OCT mapping of the attenuation coefficient”, Biomed Opt Express, 9: 5084-99 [Non-Patent Document 3] Dwork, N., G. T. Smith, T. Leng, J. M. Pauly, and A. K. Bowden. 2019. “Automatically Determining the Confocal Parameters From OCT B-Scans for Quantification of the Attenuation Coefficients”, IEEE Trans Med Imaging, 38: 261-68 [Non-Patent Document 4] Kubler, J., V. S. Zoutenbier, A. Amelink, J. Fischer and J. F. de Boer (2021). “Investigation of methods to extract confocal function parameters for the depth resolved determination of attenuation coefficients using OCT in intralipid samples, titanium oxide phantoms, and in vivo human retinas.” Biomedical Optics Express 12(11) [Non-Patent Document 5] Vermeer, K. A., J. van der Schoot, H. G. Lemij and J. F. de Boer (2012). “RPE-normalized RNFL attenuation coefficient maps derived from volumetric OCT imaging for glaucoma assessment.” Invest Ophthalmol Vis Sci 53(10): 6102-6108 [Non-Patent Document 6] Vermeer, K. A., J. Mo, J. J. Weda, H. G. Lemij and J. F. de Boer (2013). “Depth-resolved model-based reconstruction of attenuation coefficients in optical coherence tomography.” Biomed Opt Express 5(1): 322-337

Claims

1. An OCT imaging system, a confocal lens arrangement (23) for producing a retinal image; an apparatus arranged to carry out the method for determining and correcting a confocal function in said OCT imaging system; Equipped with the apparatus is provided for determining the confocal function of the OCT imaging system and correcting for the effect of the confocal function; the apparatus is arranged to change the focus between acquisitions of the two A-lines to acquire the two A-lines with different depths of focus; the apparatus is arranged to take a ratio of the two A-lines to determine the confocal function; the device has a dispersive element (1) that produces different focal positions as a function of wavelength (λ), and / or The device comprises an active optical element (1) whose refractive power can be changed, An OCT imaging system comprising:

2. An OCT imaging system as described in claim 1, wherein the apparatus is configured to apply a Fourier transform to the spectrum and create a confocal function corrected high resolution OCT image (22).

3. An OCT imaging system as described in claim 1, wherein the device comprises: determining the confocal function, and / or correcting for the influence of the confocal function, and / or Correcting the confocal function 10. An OCT imaging system, comprising:

4. An OCT imaging system as described in claim 2, wherein the device comprises: determining the confocal function, and / or correcting for the influence of the confocal function, and / or Correcting the confocal function 10. An OCT imaging system, comprising:

5. An OCT imaging system as described in Claim 3, characterized in that the device is configured to perform a step of reconstructing a high-resolution confocal corrected image (22).

6. An OCT imaging system as described in Claim 4, characterized in that the device is configured to perform a step of reconstructing a high-resolution confocal corrected image (22).

7. 6. The OCT imaging system of claim 5, wherein the reconstruction of the high-resolution confocal corrected image (22) is based on coherent addition of corrected profiles of wavelength sub-bands.

8. An OCT imaging system as described in Claim 6, characterized in that the reconstruction of the high-resolution confocal corrected image (22) is based on coherent addition of corrected profiles of wavelength subbands.

9. 9. An OCT imaging system according to any one of claims 3 to 8, comprising: The device comprises: generating OCT images (6, 7) for different wavelengths (λ) with different focal positions; determining a confocal function of the OCT image; generating OCT images for different wavelengths (λ) with different focal positions and confocal functions; correcting the confocal function to obtain confocal function corrected OCT images for different wavelengths (λ); applying an inverse Fourier transform to obtain spectra at different wavelengths (λ); coherently adding a complex spectrum to the spectrum for each different wavelength (λ); generating a spectrum of a full range of wavelengths (λ); applying a Fourier transform to the full range spectrum of wavelengths (λ) to produce a confocal function corrected high resolution OCT image (22); 10. An OCT imaging system, comprising:

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