A design method for endoscopic optical coherence tomography probe

Abstract

The present invention discloses a method for designing an endoscopic optical coherence tomography (OCT) probe, which relates to the technical field of optical endoscopic probe design. The method includes steps S100: decomposition of a three-dimensional needle-shaped light field; S200: construction of the complex amplitude of a single-layer two-dimensional light field; S300: acquisition of the complex amplitude of a single-layer two-dimensional light field; S400: calculation of the complex amplitude of each single-layer two-dimensional light field; S500: calculation of the phase distribution; S600: geometric distribution conversion; and S700: fabrication of an endoscopic optical coherence tomography (OCT) probe. The present invention improves imaging uniformity and efficiency, resolving the trade-off between probe size, resolution, and DOF. The present invention is closely related to the fields of biomedical imaging and optics, and provides a solution for improving the clinical practicality of endoscopic OCT.

Description

Technical Field

[0001] The present invention relates to the technical field of optical endoscope probe design, and in particular to a design method of an endoscope optical coherence tomography probe. Background Art

[0002] Fiber optics has extended the high-resolution tomographic imaging capabilities of optical coherence tomography (OCT) to the human body, resulting in endoscopic OCT. This technology has been applied in the clinical diagnosis and treatment of coronary arteries and the brain, playing a crucial role in assessing plaque morphology on the vascular wall, evaluating the effectiveness of stent implants, and detecting intracranial artery stenosis. Furthermore, OCT in natural cavities (such as the digestive, respiratory, and reproductive tracts) can serve as an on-site, immediate "optical biopsy" and is increasingly being adopted in clinical practice. In recent years, endoscopic OCT has also begun to monitor treatments such as radiofrequency ablation of the heart and esophagus, and resection of colorectal cancer, enabling the integration of precise diagnosis and treatment.

[0003] The axial resolution of OCT depends on the coherence length of the light source. Using a broadband laser source, an axial resolution of less than 10 μm can be achieved within an imaging depth of several millimeters. In contrast, its lateral resolution is determined by the profile of the focused beam emitted by the sample arm probe. Although focusing the Gaussian beam with a high numerical aperture lens can obtain OCT images with high lateral resolution, it inevitably reduces the depth of focus (DOF). Images with high lateral resolution are only maintained within the depth of focus. In the out-of-focus area, the lateral resolution and signal-to-noise ratio of the image will drop significantly, reducing the accuracy of quantitative analysis of the sample structure.

[0004] To overcome the limitations of OCT's limited focal depth and achieve a balance between high resolution and focal depth, various methods for extending focal depth have been proposed. The main existing technologies are as follows:

[0005] 1. "Miniature all-fiber axicon probe with extended Bessel focus for optical coherence tomography" (Optics express, 2019, 27(2): 358-366, Wang, Wei et al.) proposed to fabricate an aconical lens on the sample arm fiber probe to generate a Bessel beam. First, a gradient refractive index (GRIN) fiber is spliced ​​between a single-mode fiber (SMF) and a coreless fiber (NCF). The NCF is then mechanically polished to produce an aconical lens at the probe tip. The disadvantages of this method are that it requires precise control of the GRIN fiber length and the polishing process of the NCF. In addition, the coherent transfer function of the Bessel beam suffers from severe spatial frequency component loss, which leads to decreased sensitivity and the appearance of sidelobe artifacts in OCT images.

[0006] Second, using the principle of multimode interference, the multimode interference probe consists of a transition fiber and a large core fiber (LCF) (the combination of the two can be regarded as a pupil filter). The transition fiber is used to improve the light transmission efficiency while controlling the excitation mode in the large core fiber. Adjusting the length of the LCF can change the phase difference between the excitation modes, thereby obtaining a controllable light field distribution at the output end of the LCF. Using this principle, "Uniform focusing with an extended depth range and increased working distance for optical coherence tomography by an ultrathinmonolith fiber probe" [Optics Letters, 2020, 45(4): 976-979] (Qiu et al.) proposed that by optimizing the length of the GIF1 fiber (transition fiber) and the LCF, the light will undergo linear polarization mode interference during the propagation of the LCF, forming a multimode interference field at the end face of the LCF-NCF (coreless fiber). After propagation and amplification by the NCF, it is focused by the GIF2 fiber to achieve focal depth extension. The disadvantage of this method is that the working distance is very short.

[0007] 3. Other methods: The depth of focus extension method proposed by Zhao et al. in "Flexible method for generating needle-shaped beams and its application in optical coherence tomography" [Optica, 2022, 9(8): 859-867] requires the use of an objective lens and a diffraction optical element in combination to achieve depth of focus extension. This is obviously not applicable to endoscopic OCT. At the same time, the light field distribution generated by this method has low focusing efficiency and many side lobes, which lead to imaging artifacts. Synthetic aperture technology also requires additional optical element configuration and is difficult to apply in endoscopic OCT. The single-mode optical fiber used in the endoscopic OCT probe has a diameter of 125 microns, and it is extremely difficult to arrange too many elements in such a small space.

[0008] Therefore, those skilled in the art are committed to developing a design method for an endoscopic optical coherence tomography probe. Summary of the Invention

[0009] In view of the above-mentioned defects of the prior art, the technical problem to be solved by the present invention is how to effectively extend the DOF while achieving high resolution and small-sized probes.

[0010] In the field of computational holography, 3D objects are usually constructed flexibly through computational reconstruction. The desired needle-shaped light field (the high-resolution light field distribution after the depth of focus is extended becomes the needle-shaped light field distribution) is used as the object to be reconstructed, and it is divided into several single-layer two-dimensional light fields. Each single-layer two-dimensional light field is preset with an amplitude distribution with the desired high resolution to achieve high resolution. In order to further improve the focusing efficiency and further concentrate the energy in the main lobe area, the energy gradient of the preset amplitude distribution is consistent with the Gaussian; the Gerchberg-Saxton (GS) algorithm is used to optimize the single-layer two-dimensional light field, using the incident light amplitude and random phase as input to optimize the phase that can produce a preset focused light spot in the desired observation plane. The phase is then combined with the desired amplitude to construct a complex amplitude for reverse light field propagation and retain the complex amplitude. , and so on, the phase optimization of each surface is completed and the complex amplitude is retained. Finally, the complex amplitude is superimposed to obtain the phase to obtain the final phase distribution; the phase distribution is then converted into the geometric distribution of the diffraction optical element, and using two-photon printing technology, it is directly printed onto the optical fiber end face to complete the final production of the endoscopic optical coherence tomography probe.

[0011] In one embodiment of the present invention, a method for designing an endoscopic optical coherence tomography probe is provided, comprising:

[0012] S100, three-dimensional needle-shaped light field decomposition, the three-dimensional needle-shaped light field is decomposed into A combination of single-layer two-dimensional light fields, is a positive integer;

[0013] S200, constructing the complex amplitude of a single-layer two-dimensional light field, according to the expected amplitude of each single-layer two-dimensional light field, calculating the phase by the Gerchberg-Saxton (GS) phase retrieval algorithm to construct the complex amplitude of the single-layer two-dimensional light field;

[0014] S300, single-layer two-dimensional light field complex amplitude acquisition, reverse propagation of the single-layer two-dimensional light field complex amplitude, to obtain the complex amplitude of the diffractive optical element output surface ;

[0015] S400, calculating the complex amplitude of each single-layer two-dimensional light field, repeating steps S200 and S400 layer by layer until the calculation of N single-layer two-dimensional light fields is completed, and calculating and saving the complex amplitude of the exit surface of the diffractive optical element corresponding to each single-layer two-dimensional light field;

[0016] S500, phase distribution calculation, accumulating the complex amplitudes of the exit surface of the diffractive optical element obtained in step S400 and taking the phase to obtain a final phase distribution;

[0017] S600, geometric distribution conversion, using the cumulative effect of different phase modulations at different heights of the material to convert the final phase distribution of the diffractive optical element Converted into geometric distribution of diffractive optical elements;

[0018] S700, endoscopic optical coherence tomography probe production, uses two-photon 3D printing technology to print diffraction optical elements onto the end face of the optical fiber to complete the production of the endoscopic optical coherence tomography probe.

[0019] Optionally, in the endoscopic optical coherence tomography probe design method in the above embodiment, the number of single-layer two-dimensional light fields is The selection is made based on the length of the three-dimensional needle-shaped light field.

[0020] Preferably, in the endoscopic optical coherence tomography probe design method in the above embodiment, the number of single-layer two-dimensional light fields is is 80.

[0021] Optionally, in the endoscopic optical coherence tomography probe design method in any of the above embodiments, the single-layer two-dimensional light field complex amplitude includes the desired amplitude and phase , where the expected amplitude is a known condition, namely the light field amplitude generated by the endoscopic optical coherence tomography probe.

[0022] Optionally, in the endoscopic optical coherence tomography probe design method in any of the above embodiments, the Gerchberg-Saxton (GS) phase retrieval algorithm includes:

[0023] S210. Calculate the amplitude of the incident light. The amplitude of the incident light is Gaussian distributed, that is, the amplitude distribution of the single-mode optical fiber after beam expansion is Gaussian distributed. The formula is as follows:

[0024] (1)

[0025] in is the amplitude of the incident light, It is Gaussian light The radius, is the spatial coordinate;

[0026] S220, constructing a complex amplitude. The incident light amplitude and the random phase are used as the starting phase of the phase recovery algorithm to construct the complex amplitude of the output surface of the diffractive optical element. The formula is as follows:

[0027] (2)

[0028] in, is the incident light amplitude, is a random phase;

[0029] S230, calculate the light field distribution of a single-layer two-dimensional light field, assuming that the exit surface of the diffractive optical element is the same as the single-layer two-dimensional light field The distance is , j Represents the index of any layer. A single-layer two-dimensional light field is a three-dimensional needle-shaped light field (formed by the interference of the light field emitted by the exit surface in space) divided into N layers. Therefore, the distance from the exit surface to each layer is different. The light field distribution calculation formula is as follows:

[0030] (3)

[0031] in, represents the Fresnel impulse response, is the wavelength of the incident light, is the wave number;

[0032] S240, performing effective calculation by fast Fourier transform. According to Fourier optics theory, step S230 performs effective calculation by fast Fourier transform, and the formula is as follows:

[0033] (4)

[0034] in and represent forward Fourier transform and inverse Fourier transform respectively, represents the Fresnel transfer function;

[0035] S250, construct light field distribution, in a single layer of two-dimensional light field Producing the desired amplitude ,Will amplitude Replace with , The phase remains unchanged in a single-layer two-dimensional light field Constructing a new light field distribution , by reverse light field propagation, the light field distribution on the exit surface of the diffractive optical element is obtained:

[0036] (5);

[0037] S260, constructing complex amplitude, in order to ensure that the amplitude distribution of the exit surface of the diffractive optical element is consistent with the amplitude of the incident light, Amplitude Replaced by the incident light amplitude , keeping the phase unchanged, constructing the complex amplitude of the output surface of the diffractive optical element;

[0038] S270, iterative calculation, taking the complex amplitude of the exit surface of the diffractive optical element as the value in S230 formula (3) , start the next round of iteration, execute S230-S270 until the iteration meets the preset number of times, and finally obtain the phase .

[0039] Preferably, in the endoscopic optical coherence tomography probe design method in the above embodiment, the preset number of times is 200 times.

[0040] Furthermore, in the endoscopic optical coherence tomography probe design method in the above embodiment, the final phase distribution is:

[0041] (6)

[0042] in, For each single layer of two-dimensional light field The corresponding complex amplitude of the exit surface is, N is the total number of single-layer two-dimensional light fields, represents the phase operation, is the final phase distribution.

[0043] Furthermore, in the endoscopic optical coherence tomography probe design method in the above embodiment, the final phase distribution The conversion to the geometric distribution of the diffractive optical element uses the following formula:

[0044] (7)

[0045] in, is the material refractive index, is the geometric height of the phase element, is the wavelength of incident light.

[0046] Optionally, in the endoscopic optical coherence tomography probe design method in any of the above embodiments, the endoscopic optical coherence tomography probe is manufactured using a femtosecond laser two-photon 3D printer.

[0047] Optionally, in the endoscopic optical coherence tomography probe design method in the above embodiment, the refractive index of the photoresist material used in the femtosecond laser two-photon 3D printer is 1.56.

[0048] This invention utilizes computer-generated holography (CGH) to significantly extend the depth of focus (DOF) of endoscopic optical coherence tomography (OCT). It employs diffraction optics and CGH's multidimensional light field modulation capabilities to precisely control light intensity distribution, eliminating the need for an objective lens. Furthermore, two-photon 3D printing technology can be used to directly fabricate endoscopic OCT probes at the distal end of a single-mode optical fiber. This not only improves imaging uniformity and efficiency but also resolves the trade-off between probe size, resolution, and DOF. The effectiveness of this invention was validated through numerical simulations, beam measurements, and imaging results. This invention has the potential to improve clinical diagnosis and treatment monitoring in various medical applications, including cardiovascular and gastrointestinal assessments. This invention is closely related to the fields of biomedical imaging and optics, providing a solution for enhancing the clinical practicality of endoscopic OCT.

[0049] The concept, specific structure and technical effects of the present invention will be further described below in conjunction with the accompanying drawings to fully understand the purpose, characteristics and effects of the present invention. BRIEF DESCRIPTION OF THE DRAWINGS

[0050] Figure 1 is a flow chart illustrating a method for designing an endoscopic optical coherence tomography probe according to an exemplary embodiment;

[0051] Figure 2 is a flow chart illustrating a Gerchberg-Saxton (GS) phase retrieval algorithm according to an exemplary embodiment;

[0052] Figure 3 is a comparative diagram illustrating effects of generating light beams according to an exemplary embodiment;

[0053] Figure 4 is a comparative diagram illustrating focusing efficiency according to an exemplary embodiment;

[0054] Figure 5 is a schematic structural diagram illustrating an endoscopic optical coherence tomography probe according to an exemplary embodiment;

[0055] Figure 6 is a comparative graph illustrating focusing efficiency of endoscopic optical coherence tomography probes according to an exemplary embodiment. DETAILED DESCRIPTION

[0056] The following describes several preferred embodiments of the present invention with reference to the accompanying drawings to make its technical content clearer and easier to understand. The present invention can be embodied in many different forms of embodiments, and the scope of protection of the present invention is not limited to the embodiments mentioned herein.

[0057] In the drawings, components with identical structures are denoted by the same reference numerals, and components with similar structures or functions are denoted by similar reference numerals. The size and thickness of each component shown in the drawings are arbitrary and are not limited by the present invention. To enhance clarity, the thickness of components in some places in the drawings is schematically exaggerated.

[0058] In one embodiment of the present invention, a method for designing an endoscopic optical coherence tomography probe is provided, such as Figure 1 Shown, including:

[0059] S100, three-dimensional needle-shaped light field decomposition, the three-dimensional needle-shaped light field is decomposed into A combination of single-layer two-dimensional light fields, is a positive integer, and the number of single-layer two-dimensional light fields is selected according to the length of the three-dimensional needle light field. , is 80.

[0060] S200, single-layer two-dimensional light field complex amplitude construction, according to the expected amplitude of each single-layer two-dimensional light field, the phase is calculated by the Gerchberg-Saxton (GS) phase recovery algorithm to construct the single-layer two-dimensional light field complex amplitude, including the expected amplitude and phase , where the expected amplitude is a known condition, namely, the amplitude of the light field generated by the endoscopic optical coherence tomography probe; the Gerchberg-Saxton (GS) phase retrieval algorithm includes:

[0061] S210. Calculate the amplitude of the incident light. The amplitude of the incident light is Gaussian distributed, that is, the amplitude distribution of the single-mode optical fiber after beam expansion is Gaussian distributed. The formula is as follows:

[0062] (1)

[0063] in is the amplitude of the incident light, It is Gaussian light The radius, is the spatial coordinate;

[0064] S220, constructing a complex amplitude. The incident light amplitude and the random phase are used as the starting phase of the phase recovery algorithm to construct the complex amplitude of the output surface of the diffractive optical element. The formula is as follows:

[0065] (2)

[0066] in, is the incident light amplitude, is a random phase;

[0067] S230, calculate the light field distribution of a single-layer two-dimensional light field, assuming that the exit surface of the diffractive optical element is the same as the single-layer two-dimensional light field The distance is , j Represents the index of any layer. A single-layer two-dimensional light field is a three-dimensional needle-shaped light field (formed by the interference of the light field emitted by the exit surface in space) divided into N layers. Therefore, the distance from the exit surface to each layer is different. The light field distribution calculation formula is as follows:

[0068] (3)

[0069] in, represents the Fresnel impulse response, is the wavelength of the incident light, is the wave number;

[0070] S240, performing effective calculation by fast Fourier transform. According to Fourier optics theory, step S230 performs effective calculation by fast Fourier transform, and the formula is as follows:

[0071] (4)

[0072] in and represent forward Fourier transform and inverse Fourier transform respectively, represents the Fresnel transfer function;

[0073] S250, construct light field distribution, in a single layer of two-dimensional light field Producing the desired amplitude ,Will amplitude Replace with , The phase remains unchanged in a single-layer two-dimensional light field Constructing a new light field distribution , by reverse light field propagation, the light field distribution on the exit surface of the diffractive optical element is obtained:

[0074] (5);

[0075] S260, constructing complex amplitude, in order to ensure that the amplitude distribution of the exit surface of the diffractive optical element is consistent with the amplitude of the incident light, Amplitude Replaced by the incident light amplitude , keeping the phase unchanged, constructing the complex amplitude of the output surface of the diffractive optical element;

[0076] S270, iterative calculation, taking the complex amplitude of the exit surface of the diffractive optical element as the value in S230 formula (3) , start the next round of iteration, and execute S230-S270 until the iteration meets the preset number of times, which is 200 times.

[0077] S300, single-layer two-dimensional light field complex amplitude acquisition, reverse propagation of the single-layer two-dimensional light field complex amplitude, to obtain the complex amplitude of the diffractive optical element output surface .

[0078] S400, calculating the complex amplitude of each single-layer two-dimensional light field. Repeat steps S200 and S400 layer by layer until the calculation of N single-layer two-dimensional light fields is completed. The complex amplitude of the exit surface of the diffractive optical element corresponding to each single-layer two-dimensional light field is calculated and saved.

[0079] S500, phase distribution calculation, accumulating the complex amplitude of the exit surface of the diffractive optical element obtained in step S400 and taking the phase to obtain the final phase distribution, which is:

[0080] (6)

[0081] in, A single-layer two-dimensional light field The corresponding complex amplitude of the exit surface is, N is the number of single-layer two-dimensional light fields, represents the phase operation, is the final phase distribution.

[0082] S600, geometric distribution conversion, using the cumulative effect of different phase modulations at different heights of the material, the final phase distribution The formula used to convert to the geometric distribution of diffractive optical elements is as follows:

[0083] (7)

[0084] in, is the material refractive index, is the geometric height of the phase element, is the wavelength of incident light.

[0085] S700, endoscopic optical coherence tomography probe production, uses two-photon 3D printing technology, uses a femtosecond laser two-photon 3D printer, and uses a photoresist material with a refractive index of 1.56. The diffraction optical element is printed onto the end face of the optical fiber to complete the production of the endoscopic optical coherence tomography probe.

[0086] To verify the effectiveness, theoretical simulations and imaging experiments after probe fabrication were conducted. First, numerical simulations were performed, setting the target resolution to 5 μm, or the full width at half maximum (FWHM) of the beam. An aperture of 125 μm, the diameter of the inner cladding of a single-mode fiber, was assumed. Comparative simulations included a Gaussian beam, generated by propagating the beam through a focusing lens with a focal length of 500 μm and using Fresnel diffraction, and a needle-shaped beam generated using the multi-focus method. The Gaussian beam was generated by propagating the beam through a focusing lens with a focal length of 500 μm and using Fresnel diffraction. Using the multi-focus method, nine equidistant focal points were established in the axial (z) direction within the target focal region (500 to 700 μm after the diffractive optical element), and the phase adjuster was set to 0.028π to generate the needle-shaped beam. Using the patented endoscopic optical coherence tomography probe design method, the range from 500 to 700 μm after the phase mask was divided into 80 planes, and the desired value for each plane was set to an ideal point with a 5 μm FWHM.

[0087] like Figure 3 As shown in the figures, (a) and (b) are the results for a Gaussian beam, (c) and (d) are the results for a needle-shaped beam generated using the multifocal method, and (e) and (f) are the results for a needle-shaped beam generated using the present invention's computer-generated hologram (CGH) method. (a), (c), and (e) are cross-sectional views of the beams in the x-z plane (the results in the y-z plane are similar because the generated beams have rotational symmetry about the z-axis). (b), (d), and (f) are the corresponding cross-sectional views in the x-y plane at the dashed lines in (a), (c), and (e). It can be seen that the focal zones of the needle-shaped beams generated using the multifocal method and the present invention are much longer than those of the Gaussian beam. Compared to the multifocal method, the beams generated using the present invention have fewer diffraction sidelobes and more uniform axial and lateral distributions.

[0088] Further comparison of the focusing efficiency of the focused beams generated by different methods, such as Figure 4 As shown in Figure 2. Focusing efficiency is defined as the percentage of the energy within a circle with a diameter of 3 times the FWHM in each x-y plane distributed along the axial z direction to the total incident Gaussian energy. Among them, Gaussian beam has the highest focusing efficiency, about 98%, but its short depth of focus (DOF) limits its imaging performance in endoscopic OCT, as shown in Figure 2. Figure 6(a) and (c) are shown. The multifocal method and the diffractive optical method of the present invention are relatively inefficient. This is because the present invention only uses a diffractive optical element (DOE) with 16 height levels. In the future, as processing accuracy improves, the focusing efficiency can be further improved by using DOEs with more height levels. Even so, the present invention still shows a significant improvement in focusing efficiency compared to the multifocal method, with the maximum focusing efficiency increased from approximately 42% to 67%.

[0089] A commercial femtosecond laser two-photon 3D printer was used to fabricate a diffractive optical element (DOE) using the present invention at the distal end of a single-mode optical fiber. First, a solid cylinder approximately 690 μm long was printed on the end face of the single-mode optical fiber to expand the beam diameter to approximately 100 μm. The designed DOE was then printed on the cylinder. The refractive index of the photoresist material used for 3D printing was 1.56. Figure 5 (a) and (b) are micrographs from the side and top views, respectively; (c) is a photograph of the DOE surface taken using a scanning electron microscope; and (d) is the measured beam profile of the probe. The light source used was an endoscopic OCT system with a central wavelength of 840 nm and a bandwidth of approximately 50 nm. Beam measurements were performed using an imaging system consisting of a 20× objective lens, a 75 mm focal length lens, and a camera. The light emitted from the fiber end face was considered the object, forming a conjugate relationship with the image received by the camera. While the camera was held stationary, the fiber stage was gradually moved to record the lateral light field distribution layer by layer and reconstruct the beam profile. The measured beam had a full-width-at-half-maximum (FWHM) spot diameter of 5 μm and a focal depth of 224 μm. For comparison, a Gaussian beam with the same spot diameter has a focal depth of 67 μm. The probe's insertion loss was also measured using an optical power meter. For an input fiber power of 17 mW and a DOE output power of 12.4 mW, the resulting insertion loss was 1.367 dB.

[0090] like Figure 6 Figures show a comparison of OCT imaging results using an endoscopic probe designed according to the present invention ((a) and (c)) and a free-space Gaussian beam setup with a similar focused spot ((b) and (d)). (a) and (b) show the imaging results of a resolution phantom made of gold nanoparticles (less than 100 nm in size) and polydimethylsiloxane. (c) and (d) show the imaging results of fresh orange pulp. It can be seen that the present invention maintains high resolution across the entire imaging range, while a Gaussian beam with similar resolution exhibits significant loss of resolution and signal intensity due to defocusing.

[0091] The above describes in detail the preferred embodiments of the present invention. It should be understood that those skilled in the art can make numerous modifications and variations based on the concepts of the present invention without inventive effort. Therefore, any technical solutions that can be derived by those skilled in the art through logical analysis, reasoning, or limited experimentation based on the concepts of the present invention and the prior art should be within the scope of protection defined by the claims.

Claims

1. A method for designing an endoscopic optical coherence tomography probe, characterized in that: include: S100, three-dimensional needle-shaped light field decomposition, the three-dimensional needle-shaped light field is decomposed into A combination of single-layer two-dimensional light fields, is a positive integer; S200, constructing the complex amplitude of a single-layer two-dimensional light field, according to the expected amplitude of each single-layer two-dimensional light field, calculating the phase by the Gerchberg-Saxton (GS) phase retrieval algorithm to construct the complex amplitude of the single-layer two-dimensional light field; S300, obtaining the complex amplitude of a single-layer two-dimensional light field, performing reverse propagation on the complex amplitude of the single-layer two-dimensional light field to obtain the complex amplitude of the exit surface of the diffractive optical element ; S400, calculate the complex amplitude of each single layer of two-dimensional light field, and repeat the steps S200 and S400 layer by layer until the calculation is completed. N Calculate the single-layer two-dimensional light field, calculate and save the complex amplitude of the output surface of the diffractive optical element corresponding to each single-layer two-dimensional light field; S500, phase distribution calculation, accumulating the complex amplitudes of the exit surface of the diffractive optical element obtained in step S400 and taking the phase to obtain a final phase distribution; S600, geometric distribution conversion, using the cumulative effect of different phase modulations at different heights of the material to convert the final phase distribution of the diffractive optical element Converted into geometric distribution of diffractive optical elements; S700, endoscopic optical coherence tomography probe production, uses two-photon 3D printing technology to print diffraction optical elements onto the end face of the optical fiber to complete the production of the endoscopic optical coherence tomography probe.

2. The method for designing an endoscopic optical coherence tomography probe according to claim 1, wherein: The number of the single-layer two-dimensional light field The selection is made based on the length of the three-dimensional needle-shaped light field.

3. The method for designing an endoscopic optical coherence tomography probe according to claim 2, wherein: The number of the single-layer two-dimensional light field is 80.

4. The method for designing an endoscopic optical coherence tomography probe according to claim 3, wherein: The complex amplitude of the single-layer two-dimensional light field includes the desired amplitude and phase , the desired amplitude is a known condition, namely the light field amplitude generated by the endoscopic optical coherence tomography probe.

5. The method for designing an endoscopic optical coherence tomography probe according to claim 4, wherein: The Gerchberg-Saxton (GS) phase recovery algorithm includes: S210. Calculate the amplitude of the incident light. The amplitude of the incident light is Gaussian distributed, that is, the amplitude distribution of the single-mode optical fiber after beam expansion is Gaussian distributed. The formula is as follows: ； in is the amplitude of the incident light, It is Gaussian light The radius, is the spatial coordinate; S220, constructing a complex amplitude. The incident light amplitude and the random phase are used as the starting phase of the phase recovery algorithm to construct the complex amplitude of the exit surface of the diffractive optical element. The formula is as follows: ； in, is the incident light amplitude, is a random phase; S230, calculate the light field distribution of a single-layer two-dimensional light field, assuming that the exit surface of the diffractive optical element is the same as the single-layer two-dimensional light field The distance is , the single-layer two-dimensional light field The light field distribution calculation formula is as follows: ； in, represents the Fresnel impulse response, is the wavelength of the incident light, is the wave number; S240, performing effective calculation by fast Fourier transform. According to Fourier optics theory, the effective calculation by fast Fourier transform in step S230 is as follows: in and represent forward Fourier transform and inverse Fourier transform respectively, represents the Fresnel transfer function; S250, constructing a light field distribution, in the single-layer two-dimensional light field Constructing a new light field distribution , by reverse light field propagation, the light field distribution on the exit surface of the diffractive optical element is obtained: ； S260, constructing a complex amplitude, in order to ensure that the amplitude distribution of the exit surface of the diffractive optical element is consistent with the amplitude of the incident light, Amplitude Replaced by the incident light amplitude , keeping the phase unchanged, constructing the complex amplitude of the output surface of the diffractive optical element; S270, iterative calculation, taking the complex amplitude of the exit surface of the diffractive optical element as the value in S230 formula (3) , start the next round of iteration, execute S230-S270 until the iteration meets the preset number of times, and finally obtain the phase .

6. The method for designing an endoscopic optical coherence tomography probe according to claim 5, wherein: The preset number of times is 200 times.

7. The method for designing an endoscopic optical coherence tomography probe according to claim 1, wherein: The final phase distribution is: in, For each single layer of two-dimensional light field The corresponding complex amplitude of the exit surface is, N is the total number of single-layer two-dimensional light fields, represents the phase operation, is the final phase distribution.

8. The method for designing an endoscopic optical coherence tomography probe according to claim 7, wherein: The final phase distribution The conversion to the geometric distribution of the diffractive optical element uses the following formula: in, is the material refractive index, is the geometric height of the phase element, is the wavelength of incident light.

9. The method for designing an endoscopic optical coherence tomography probe according to claim 1, wherein: The endoscopic optical coherence tomography probe is manufactured using a femtosecond laser two-photon 3D printer.

10. The method for designing an endoscopic optical coherence tomography probe according to claim 9, wherein: The refractive index of the photoresist material used in the femtosecond laser two-photon 3D printer is 1.56.

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