Many retinal diseases are related to abnormal eye blood flow, for example, retinopathy caused by diabetes, retinal vein occlusion, and age-related macular degeneration. In the study of glaucoma, insufficient blood supply to the retina is considered to be a possible reason of occurrence and development of glaucoma. Therefore, retinal blood flow measurement is important for the clinical diagnosis, treatment, and research of retinal diseases.
Optical coherence tomography (OCT) technology is a non-invasive detection technique, which is widely used in imaging of living cross-sectional structure of biological tissues. OCT can provide tissue structures with high resolution and high sensitivity by measuring a depth-related scattering light. At the same time, the OCT technology can also be used to detect Doppler shift of the scattering light to obtain motion information of a fluid and sample and therefore, it is suitable for measuring retinal blood flow. Unfortunately, the frequency shift probed by single-beam Doppler OCT only relates to the blood flow rate in the direction of the probe beam, and blood flow information in a direction perpendicular to the direction of the probe beam cannot be directly obtained from Doppler shift, therefore, it is impossible to obtain actual blood flow rate of vessels.
In order to solve the above problem, a series of techniques have been developed to obtain actual blood flow rate of the vessels.
(1) Through a three-dimension scanning of retina, directions in the space of vessels in the retina can be obtained, so as to determine a Doppler angle of a probe light, and then the Doppler angle can be used to calculate the actual blood flow rate. Because the vessels of the retina are almost perpendicular to probe beam, this method is less accurate. In addition, a space vector of a vessel to be measured is determined by continuously scanning two planes or rings and then the Doppler angle can be calculated to obtain the actual blood flow rate. But measurement results of this method can be affected by eye movement, and it can only measure the vessels around the optic disc and cannot measure blood flow in other areas of the retina.
(2) The same point in the sample can be scanned by using multi-beam and multi-angle probe light to obtain the actual blood flow rate of the vessels. The OCT probe light can be split into two beams by a glass plate, and the two light beams can converge in the fluid to form a dual-beam and dual-angle illumination mode. The actual blood flow rate of the vessels can be obtained by analyzing Doppler shift probed by the two light beams. The drawback of this method is that, due to the time delay between the two light beams, it is not applicable for a frequency domain OCT system. Furthermore, retina vessel blood flow and retina vessel blood flow rate can be measured by using a dual-beam OCT system with beams split by a polarized light, or a DOVE prism synchronized with a OCT scanning mechanism can be used to achieve dual-beam circular scanning of the retina. Those dual-beam systems consist of two Michelson interferometers, which are complex in structure and difficult to adjust, and taking into account the safety of probe light, the power of each probe light is much lower than that of a single-beam system, which reduces the sensitivity of the dual-beam OCT system, thereby increasing the phase noise of the system.
In order to measure the blood flow rate and blood flow of a single vessel and all vessels in the optic disc, fundus vessels must be scanned by ophthalmic diagnostic equipment. Please refer to FIG. 3, FIG. 10, and FIG. 11, in which FIG. 3 is a video screenshot of linear scanning of one vessel in the optic disc, and FIGS. 10-11 are schematic diagram illustrating circular scanning of all vessels in the optic disc. As illustrated in FIG. 3, a black bond line indicates a scanning direction that the probe light scans a fundus vessel B, and the scanning direction corresponds to Y-axis direction illustrated in FIG. 9. After scanning by the probe light, an original fundus Doppler image may be obtained as illustrated in FIG. 4. In FIG. 4, there is undesired background Doppler (namely low-frequency background Doppler with alternate dark) and bright and high-frequency background Doppler with obvious vertical lines. Similarly, in the process of circular scanning of all the vessels in the optic disc as illustrated in FIGS. 10-11, an original Doppler image obtained (not illustrated) also has low-frequency background Doppler with alternate dark and bright and high-frequency background Doppler with obvious vertical lines as shown in FIG. 4. After analysis, there are two reasons for generating the background Doppler. First, a central line of a main light of the probe light does not extend through a rotation axis of a scanning unit. When the scanning unit is an X-Y galvanometer, it can be considered that a rotation axis of a Y galvanometer is the rotation axis of the scanning unit. Referring to FIG. 1, when the scanning unit is the X-Y galvanometer, if a central line O1 of the main light of the probe light extends through a rotation axis 902 of an X-Y galvanometer 900, then it is considered that the central line O1 of the main light of the probe light extends through the rotation axis of the scanning unit. With the swing of the X-Y galvanometer 900, an incidence direction e of the probe light is perpendicular to the scanning direction of the probe light on an imaging plane d of a lens 1, in this case, no additional background Doppler is introduced. Referring to FIG. 2, if the central line O1 of the main light of the probe light does not extend through the rotation axis 902 of the X-Y galvanometer 900, with the swing of the X-Y galvanometer, the incidence direction e of the probe light is no longer perpendicular to the scanning direction of the probe light on the imaging plane d of the lens 1. If an incident light deviates from a rotation axis of a galvanometer, scanning angular rate of the galvanometer is denoted by w, and focal length of the lens 1 is denoted by f, and then a frequency shift F obtained can be expressed as: F=2 fwL/λ0√{square root over (L2+f2)}, where λ0 represents a center wavelength of the probe light. High-frequency background Doppler can be removed by removing lines one by one in vertical, and it is easy to get a wrong background, which may directly affect accuracy of measurement. Second, in the process of detecting eyes, eyeballs will be involuntary rotate slightly, and thus the probe light cannot be always incident on a same position of the eyeball at a same angle, which can lead to region-shaped low-frequency background Doppler with alternate dark and bright as illustrated in FIG. 4.
Low-frequency background Doppler can be easily removed by well-known methods for removing Doppler background. For high-frequency background Doppler, because the background between each two adjacent lines is different and irregular, it is necessary to find out background Doppler for each line. This not only makes the process of removing background more tedious, but also different to ensure the accuracy rate of removing background. Wrong removal of the background will directly affect the accuracy of measuring blood flow rate of vessels, and it is necessary to remove high-frequency background Doppler by adjusting the optical path.