Light curtain imaging device and method

Abstract

The invention provides a light curtain imaging device and method. A laser speckle imaging apparatus 100 and method are provided herein. In one embodiment, an apparatus 100 for generating flow information of a biological sample 116 comprises illumination optics 115 operable to generate one or more illumination light curtains for selectively illuminating the biological sample 116 to generate corresponding scattered light; a first image acquisition device 135 operable to acquire respective scattered light of each illuminated layer having the same wavelength as the illumination light curtain; and an image processing device 141 operable to construct three-dimensional flow information of the biological sample 116 from the acquired speckle pattern of the scattered light. Devices and methods for imaging biological samples using specific scanning mirrors and grating elements are also discussed.

Description

Technical Field

[0001] The present invention generally relates to a light-sheet imaging device and method, and in particular but not exclusively, for detecting 3D flow information of a biological sample. Background Art

[0002] Laser speckle imaging (LSI) is one of the established flow imaging methods and is a label-free in vivo flow imaging modality based on the analysis of dynamic fluctuations in laser speckle patterns. Since the imaging method was first introduced in 1980, LSI has been widely used for the visualization of blood flow imaging in living tissues such as the retina, skin, and brain. LSI can be used to monitor the dynamic blood flow response and relative changes in numerical values. However, LSI is a wide-field imaging technique that is typically limited to surface imaging (e.g., surface flow maps) because it does not provide depth selectivity and is essentially a 2D surface imaging technique. In addition, blood flow velocity measurements using traditional non-invasive and non-contact LSI systems and methods are relative or qualitative rather than quantitative and are susceptible to artifacts / noise.

[0003] The object of the present invention is to solve the problems of the prior art and / or to provide useful alternatives to the public. Summary of the Invention

[0004] According to a first aspect of the present invention, there is provided a laser speckle imaging device for generating flow information of a biological sample. The device includes illumination optics operable to generate one or more illumination light-sheets for selectively illuminating the biological sample to produce corresponding scattered light; first image acquisition means operable to acquire the scattered light corresponding to each illuminated layer at the same wavelength as the illumination light-sheet; and image processing means operable to construct three-dimensional flow information of the biological sample from the speckle patterns of the acquired scattered light.

[0005] Illuminating the biological sample with a light-sheet gives the device the ability of optical sectioning to obtain high-quality 3D images of in vivo blood flow and vasculature because the 3D flow in the biological sample can be visualized layer by layer, or batch by batch layer by layer. Therefore, the present device is not limited to the penetration depth limitation and surface detection of traditional LSI systems and broadens the application of this non-invasive imaging method.

[0006] In one embodiment, the illumination optics may include a grating element operable to receive an incident beam and divide the incident beam into at least two illumination light-sheets for simultaneously and selectively illuminating the biological sample. It is conceivable that the grating element may be a transmission grating. The use of the grating element can help improve the scanning speed.

[0007] In one embodiment, the illumination optical device may include a cylindrical lens arranged to generate an incident light beam.

[0008] In one embodiment, the incident light beam may include a plurality of incident sub - light beams, and the illumination optical device may include a cylindrical lens array (or cylindrical microlens array) configured to generate the plurality of incident sub - light beams for a grating element to divide each incident sub - light beam into at least two illumination light curtains. Since a plurality of incident sub - light beams can be generated, the scanning speed can be significantly increased.

[0009] In one embodiment, the illumination optical device may further include a rotatable scanning mirror operable to adjust the angular direction of the incident light beam to generate a reflected light beam, and the incident light beam received by the grating element is the reflected light beam. It is conceivable that the rotatable scanning mirror may include a galvanometer mirror. Since the rotatable scanning mirror can rotate to adjust the angular direction of the incident light beam, this can contribute to significantly increasing the scanning speed.

[0010] It is conceivable that the grating element may be arranged such that the angle between the optical axis of each illumination light curtain and the optical axis of the corresponding scattered light obtained can be between 0 degrees and 90 degrees, or between 30 degrees and 60 degrees. With this configuration of the tilted light curtain, the scattered light obtained will be forward - scattered light, which helps to enhance the detected optical signal. The stronger amplitude of the optical signal can provide greater flexibility in configuring the image acquisition speed and exposure time without worrying about the photon budget.

[0011] In one embodiment, the first image acquisition device may include a diaphragm (iris) having an adjustable aperture for adjusting the scattered light. The adjustable diaphragm can help achieve a relatively uniform image resolution within the field of view defined by the characteristics of the tilted light curtain.

[0012] In one embodiment, the illumination optical device may include a prism operable to transmit one or more illumination light curtains to a biological sample. The use of the prism can help minimize the diffraction and aberration of the illumination light curtain.

[0013] In one embodiment, the first image acquisition device may include an emission filter operable to allow the desired fluorescence to pass through and reach the first image acquisition device. This configuration helps to extend the laser speckle imaging device to applications of fluorescence imaging.

[0014] In one embodiment, a laser speckle imaging device may include a transmission optical device operable to generate a transmission beam for illuminating a biological sample, the transmission beam having a wavelength different from that of an illumination light curtain; and a second image acquisition device operable to acquire corresponding transmission light of the biological sample at the same wavelength as the transmission beam and generate a transmission image, wherein an image processing device is operable to adjust three-dimensional flow information constructed of the biological sample based on the transmission image. Such a configuration can contribute to enhancing the laser speckle imaging device, which can in particular generate complementary blood flow and vasculature information on the one hand and provide cross-validation and calibration methods for laser speckle imaging on the other hand.

[0015] According to a second aspect of the present invention, there is provided a laser speckle imaging method. The method includes generating one or more illumination light curtains to selectively illuminate one or more layers of a biological sample to generate corresponding scattered light; acquiring the corresponding scattered light at the same wavelength as the illumination light curtain; and constructing three-dimensional flow information of the biological sample from the speckle pattern of the acquired scattered light. The method helps to simplify the process of imaging 3D flow information of a biological sample quickly but at low cost and is not limited to imaging the surface of the biological sample. The simplicity of the method can help an operator better monitor the flow state of the biological sample.

[0016] It is conceivable that one illumination light curtain can be used at a time to illuminate the biological sample; and in this case, the biological sample can be illuminated layer by layer to generate scattered light corresponding to each layer. It is also conceivable that two or more illumination light curtains can be used to simultaneously illuminate the biological sample to form a batch of illuminated layers of the biological sample, and in this case, the generated scattered light can correspond to each layer in the batch.

[0017] In one embodiment, the method may include adjusting the positions of two or more illumination light curtains to generate another batch of illuminated layers of the biological sample. It is also conceivable that the position of the illuminated layer can be adjusted using a rotatable scanning mirror, such as a galvanometer mirror. The rotatable scanning mirror can help significantly increase the scanning speed.

[0018] In one embodiment, the method may include generating a transmission beam for irradiating the biological sample to generate transmission light, the transmission beam having a wavelength different from that of the illumination light curtain; acquiring the transmission light at the same wavelength as the transmission beam; and adjusting the three-dimensional flow information constructed of the biological sample based on the acquired transmission light.

[0019] According to a third aspect of the present invention, there is provided a non-transitory computer-readable storage medium for storing a computer program which, when executed by a processor, performs a laser speckle imaging method, the method comprising generating one or more illumination light curtains to selectively illuminate one or more layers of a biological sample, thereby generating corresponding scattered light; acquiring the corresponding scattered light having the same wavelength as the illumination light curtain; and constructing three-dimensional flow information of the biological sample from the speckle pattern of the acquired scattered light.

[0020] According to a fourth aspect of the present invention, there is provided an apparatus for imaging a biological sample. The apparatus may include illumination optics including a rotatable scanning mirror and a grating element, the rotatable scanning mirror being operable to adjust the angular direction of an incident light beam to produce a reflected light beam, the grating element being operable to divide the reflected light beam into at least two illumination light curtains for selectively simultaneously illuminating the biological sample to generate corresponding light; and an image acquisition device operable to acquire the corresponding light of each illuminated layer for imaging the biological sample. By using a combination of a rotatable scanning mirror, a grating element, and light curtain illumination, the speed of scanning the biological sample is significantly increased. This can be applied to systems requiring high scanning speed for 3D scanning, such as LSI systems and fluorescence microscopes.

[0021] In one embodiment, the apparatus may include an image processing device operable to construct three-dimensional imaging information of the biological sample from the contrast pattern of the acquired corresponding light.

[0022] It is contemplated that the grating element may include a transmission grating. It is also contemplated that the rotatable scanning mirror may include a galvanometer mirror.

[0023] In one embodiment, the incident light beam may include a plurality of incident sub-beams, and the illumination optics may include a cylindrical lens array (or cylindrical microlens array) arranged to produce the plurality of incident sub-beams for the grating element to divide each incident sub-beam into at least two illumination light curtains.

[0024] In one embodiment, the illumination optics may include a cylindrical lens arranged to produce the incident light beam.

[0025] In one embodiment, the image acquisition device may include an emission filter operable to filter the acquired corresponding light to allow desired fluorescence to pass through to the image acquisition device.

[0026] It is contemplated that the grating element may be arranged such that the angle between the optical axis of each illumination light curtain and the optical axis of the acquired corresponding light may be between 30 degrees and 60 degrees. More specifically, the angle may be any value between 30 degrees and 60 degrees, and thus, the range may be between 25 degrees and 55 degrees, between 20 degrees and 50 degrees, etc.

[0027] In a fifth aspect, a method for imaging a biological sample is provided, comprising: adjusting an angular direction of an incident light beam by a rotatable scanning mirror to generate a reflected light beam; and splitting the reflected light beam into at least two illumination light curtains by a grating element for simultaneously and selectively illuminating the biological sample to generate corresponding light; and acquiring the corresponding light of each illuminated layer to image the biological sample.

[0028] It will be apparent that features associated with one aspect may be interchangeably applied and used with features associated with other aspects. BRIEF DESCRIPTION OF THE DRAWINGS

[0029] Figure 1 A schematic diagram of a laser speckle imaging device according to a first embodiment of the present invention is depicted;

[0030] Figure 2 Depicts the relationship between anisotropic scattering in a biological sample and the detected signal intensity and the scattering angle, including Figure 1 the angles used therein;

[0031] Figure 3 Depicts a diagram showing the main functional steps of a laser speckle imaging method performed by the device of Figure 1 ;

[0032] Figure 4 Depicts translating a biological sample relative to an illumination light curtain using the device of Figure 1 ;

[0033] Figure 5 Depicts a transmission image of the slice orientation of a 4-dpf zebrafish larva using the device of Figure 1 ;

[0034] Figures 6(a)-(d) depict four representative frames of 300 interpolated frontal images obtained from a scan of Figure 5 ;

[0035] Figure 7 Depicts Figure 5 an angiogram of a single slice, with indications of four test points for analysis;

[0036] Figure 8 Depicts the flow velocity waveforms detected at four test points of Figure 7 ;

[0037] Figure 9 Depicts a cross-sectional view of the trunk structure of a 3-dpf zebrafish larva of an illumination light curtain slice generated by the device of Figure 1 , wherein the flow map is superimposed morphologically;

[0038] Figures 10(a)-(f) depict exemplary original laser speckle images obtained using Figure 1 the device, and the processing results of the exemplary original laser speckle images;

[0039] Figure 11(a) depicts an exemplary light intensity signal picked up from pixels in the DA region of Figure 10(a) using Figure 1 the device; Figure 11(b) depicts the short-time power spectrum obtained by performing time-frequency analysis on Figure 11(a); Figure 11(c) describes the comparison of DA blood flow velocities obtained using LSH-LSI and PIV;

[0040] Figures 12(a)-(f) depict the results of applying PIV analysis to the LSH-LSI scalar velocity map and the transmission image, demonstrating Figure 1 the device's ability in vector velocity mapping;

[0041] Figure 13 depicts a laser speckle imaging device according to a second embodiment of the present invention;

[0042] Figure 14 depicts a schematic diagram of a light curtain imaging device according to a third embodiment of the present invention;

[0043] Figure 15 depicts a schematic diagram of another light curtain imaging device according to a fourth embodiment of the present invention. Detailed Description

[0044] The present disclosure provides a light sheet laser speckle imaging system (LSH-LSI). Specifically, Figure 1FIG. 0 depicts a schematic diagram of a laser speckle imaging device 100 according to a first embodiment of the present invention. The device 100 generally includes an illumination optical device 115 located on an illumination optical path, a sample stage 118, a first image acquisition device 135 located on a detection optical path, and a computing device 141. To generate light, the illumination optical device 115 includes a first light source 102, and in this embodiment, the first light source 102 is a laser diode 102 having a light output of 640 nm in this embodiment. A collimator 104 is configured to receive light from the laser diode 102 and output a light beam. The illumination optical device 115 further includes a beam expander 106, which is configured to receive the light beam and increase the diameter of the light beam collimated by the collimator 104 to a larger collimated output light beam as the incident light beam. The illumination optical device 115 further includes a mirror 108, a first aperture 110, a cylindrical lens 112 (f = 50 mm in this embodiment), and an illumination objective 114. The mirror 108 is configured to adjust the angular direction of the incident light beam to generate a reflected light beam that passes through the first aperture 110. The first aperture 110 with an adjustable aperture is configured to control and adjust the characteristics or parameters of the reflected light beam to effectively control and adjust the parameters of the reflected light beam, such as the thickness and length in the focusing region of the reflected light beam. After passing through the adjustable aperture of the first aperture 110, the reflected light beam controlled or adjusted by the first aperture 110 passes through the cylindrical lens 112 and the illumination objective 114. The cylindrical lens 112 and the illumination objective 114 continuously converge the reflected light beam and output an illumination light curtain to illuminate a biological sample 116 placed on the stage 118 (e.g., a zebrafish larva as shown in Figure 1 ).

[0045] The stage 118 includes a standard glass bottom dish 120 and a glass slide 122. The stage 118 is configured with a central mounting hole for mounting the standard glass bottom dish 120. The glass slide 122 is configured to be mounted below the glass bottom dish 120, with a small air gap between the bottom surface of the glass bottom dish 120 and the top surface of the glass slide 122. The illumination optical device 115 further includes a prism 124. The prism 124 is configured to be attached to the glass slide 122 and located below the stage 118. The prism 124 is configured to minimize the diffraction and aberration of the illumination light curtain received from the illumination objective 114, and reduce the wavefront distortion along the illumination optical path of the optical axis ① perpendicular to the illumination light curtain. The stage 118 further includes an actuator 126, which can be digitally controlled by a data acquisition device or other suitable device. The actuator 126 is configured to drive the movement of the stage 118 to move the stage 118 left and right with a micron-level resolution for depth scanning. By driving the actuator 126, the biological sample 116 in the dish 120 can be shifted or moved in a desired manner (e.g., desired direction / speed) together with the stage 118. After the biological sample 116 is illuminated by the illumination light curtain, the light scattered from the biological sample 116 is collected by the first image acquisition device 135.

[0046] The first image acquisition device 135 includes a collection objective 128, a second aperture 130, a tube lens 132 (f = 100 mm in this embodiment), a dichroic mirror 134, and a first camera (also can be translated as "video camera") 136. The collection objective 128 is configured to collect and direct the light scattered from the biological sample 116 through the second aperture 130. The second aperture 130 with an adjustable aperture is configured to control and adjust the characteristics or parameters of the scattered light collected by the collection objective 128. After passing through the adjustable aperture of the second aperture 130, the scattered light is directed to continuously pass through the tube lens 132 and the dichroic mirror 134 to be collected or received by the first camera 136. The first camera 136 in this embodiment is a high-speed scientific CMOS (sCOMS) camera (e.g., pco.dimax cs1, Excelitas TechnologiesR Corp TM ). The sCMOS camera 136 is capable of capturing raw speckle images at a full-frame (1296×1024 pixels) rate of up to 3086 frames per second (fps), and can reach 10782 fps for a medium image size of 528×528 pixels. At Figure 1 the position shown, the illumination optical path is below the stage 118, and the image is captured by the sCMOS camera 136 above the stage 118.

[0047] From Figure 1As can be seen, the angle between the optical axis ① of the illumination light curtain and the optical axis ④ of the scattered light received by the collection objective lens 128 is greater than 90 degrees (in other words, the angle between the direction of the illumination light curtain and the direction of the collected scattered light is less than 90 degrees). In this case, the scattered light is forward scattered light. Similarly, as Figure 1 shown, the collection objective lens 128 can be arranged at different positions. For example, the optical axis of the scattered light is at ②, and this optical axis is approximately perpendicular to the optical axis ① of the illumination light curtain. In this case, the scattered light is backward scattered light. ③ shows the section direction in which the biological sample 116 can be imaged.

[0048] Figure 2 describes the anisotropic scattering in the biological sample and the relationship between the detected signal intensity and the scattering angle, including Figure 1 the angles used in Figure 2 Specifically, Figure 2 shows the angular distribution of the scattered photon intensity. As Figure 2 shown, when the illumination light curtain is oriented in the direction of the optical axis ①, the density of the collected scattered photons is affected by the angle between the optical axis ① and the optical axis of the scattered light received by the collection objective lens 128. The angular distribution can generally be divided into two regions, the backward scattering region 142 and the forward scattering region 144, as Figure 2 shown. Figure 2 There are multiple concentric circles in Figure 2 Figure 2 The center of the circles represents the biological sample 116. When the biological sample 116 is illuminated by the illumination light curtain from the direction ①, the scattered light can be detected from different directions, as shown by the respective arrows in Figure 2 The length of each arrow indicates the density of the scattered photons in the direction of the arrow. It can be seen that the arrows in the forward scattering region 144 are significantly longer than those in the backward scattering region 142. This indicates that the forward scattered light captured by the sCMOS camera 136 has a stronger amplitude compared to the backward scattered light detected in the confocal setting. Therefore, the image acquisition speed and exposure time can be configured very flexibly without worrying about the photon budget. From Figure 1 it can be seen that the illumination light curtain enters the biological sample 116 at a small angle (about 30°) relative to the horizontal surface of the basic horizontal stage 118 or the bottom surface of the glass bottom dish 120, and the sCMOS camera 136 is vertically oriented. This is also shown in Figure 2 where the angle between the optical axis ① of the illumination light curtain and the optical axis ④ of the scattered light received by the collection objective lens 128 is approximately 120°, and the angle between the optical axis ① and the optical axis ② is approximately 90°.

[0049] As Figure 1As shown, the laser speckle imaging device 100 includes a transmission optical device 137 and a second image acquisition device 139 for simultaneously performing wide-field microscopic imaging. The transmission optical device 137 and the illumination optical device 115 share a component, that is, the transmission optical device 137 includes a prism 124. The transmission optical device 137 further includes a second light source 138, and in this embodiment, the first light source 138 is an LED 138. The laser diode 102 and the LED 138 have different wavelengths, so that the scattered light from the laser diode 102 and the transmitted light from the LED 138 can be separated by the dichroic mirror 134 and detected by the sCMOS camera 136 and the CMOS camera 140 respectively. The second image acquisition device 139 is configured to capture a wide-field transmission microscopic image. As Figure 1 shown, the second image acquisition device 139 and the first image acquisition device 135 share some common components, that is, the second image acquisition device 139 includes a collection objective lens 128, a second aperture 130, a tube lens 132, and a dichroic mirror 134. The difference is that the second image acquisition device 139 does not include the sCMOS camera 136, but includes a second camera 140 (in this embodiment, a general machine vision CMOS camera (e.g., UI-167 3060CP-M-GL Rev.2, IDS)). To obtain a 2D image (or 2D flow map), the stage 118 is adjusted to position the illumination light curtain in an appropriate area inside the biological sample 116. The instantaneous images captured by the sCMOS camera 136 and the CMOS camera 140 are configured to be displayed on a monitor (not shown in the figure) so that the operator can find the area of interest. In this embodiment, a green LED 138 with a central wavelength of 520 nm is used for wide-field illumination, and the transmitted photons are collected by the collection objective lens 128. As Figure 1 shown, the LED 138 is configured to irradiate the biological sample 116 from an angle different from the illumination light curtain output by the illumination objective lens 114. The optical axis of the transmitted light beam generated by the LED 138 is substantially perpendicular to the surface of the stage 118. The collection objective lens 128 collects the light from the biological sample 116, including the scattered photons from the laser diode 102 and the transmitted photons from the LED 138. The dichroic mirror 134 transmits the light for speckle imaging to the sCMOS camera 136, but reflects the light for wide-field microscopic imaging to the CMOS camera 140. The wide-field transmission microscopic image can be used for particle image velocimetry (PIV) analysis, which provides complementary flow and vasculature information on the one hand and a cross-validation and calibration method for laser speckle imaging on the other hand.

[0050] Figure 3 Shows the use of Figure 1 The figure shows the main functional steps of a laser speckle imaging method using the laser speckle imaging device 100 shown. As Figure 3As shown, in order to perform laser speckle imaging using the laser speckle imaging device 100, the first step (S1) is to generate an illumination light curtain to selectively illuminate one or more layers of the biological sample 116, thereby generating corresponding scattered light. To this end, the laser diode 102 is turned on to emit a laser beam, and the collimator 104, beam expander 106, first aperture 110, and cylindrical lens 112 are used to collimate, expand, adjust, and converge the laser beam to generate one or more illumination light curtains. One or more illumination light curtains are arranged to slice through the biological sample 116 in the dish 120 mounted on the stage 118 and selectively illuminate the thin layers of the biological sample 116. After the biological sample 116 is irradiated, the light scattered from the biological sample 116. Therefore, the second step (S2) is to acquire the corresponding scattered light of the same wavelength as the illumination light curtain. To this end, the collection objective lens 128 collects the scattered light and guides the collected scattered light through the second aperture 130 and tube lens 132, and is finally captured by the sCMOS camera 136 for imaging. The sCMOS camera 136 can be configured to acquire an image sequence at an appropriate frame rate and exposure time. The movement of the stage 118 together with the biological sample 116 in the dish 120 mounted on the stage 118 can be driven by the actuator 126, including lateral movement relative to the fixed illumination light curtain. The lateral movement of the stage 118 enables the illumination light curtain to illuminate the biological sample 116 layer by layer and acquire multiple 2D images for 3D visualization. This can also be seen from Figure 4 as follows. Figure 4 depicts the translation of the biological sample 116 relative to the illumination light curtain. As Figure 4 shown, when the illumination light curtain selectively illuminates the biological sample 116 from the direction of arrow A, a part of the biological sample 116 is illuminated (i.e., Figure 4 the region 146 with a length of 453 microns shown in), while the rest of the biological sample 116 remains unaffected. The collection objective lens 128 collects the scattered light from the selectively illuminated region. And the biological sample 116 can then be laterally moved (e.g., in the direction of arrow B) to another region that was previously not affected by irradiation.

[0051] Typically, dozens to hundreds of raw images can be obtained at one location for speckle analysis, and then the biological sample 116 is transferred to another location to repeat the image acquisition process. After capturing sufficient 2D images, the third step (S3) is to construct 3D flow information of the biological sample 116 from the speckle patterns of the collected scattered light. A computing device 141 can be used, such as a personal computer, a server, or any other device suitable for processing 2D images to generate 3D models or information. The computing device 141 can be programmed to receive images from the sCMOS camera 136 to construct 3D information manually or automatically. The images captured by the sCMOS camera 136 and the CMOS camera 140 can be transmitted to the computing device 141 and displayed on a monitor in real time so that the operator can find the area of interest.

[0052] After that, the operator can adjust the frame rate and exposure time of the sCMOS camera 136 to obtain a sequence of raw images. The sequence of raw images is typically processed pixel by pixel with an algorithm to fit the dynamic intensity fluctuations with a theoretical model. Thus, the locally fitted model parameters are converted into flow velocities and assigned to the corresponding pixels. 3D sample scanning can be performed by laterally moving the stage 118 (and thus the biological sample 116) relative to the fixed illumination light curtain using an actuator 126 digitally controlled via a DAQ (data acquisition) device.

[0053] In addition, as Figure 1 shown, the device 100 has a subsystem for parallel wide-field microscopy imaging. The method can include using the LED 138 to generate transmitted light to irradiate the biological sample 116 to produce transmitted light; collecting the transmitted light having the same wavelength as the transmitted light; and processing the transmitted image obtained from the transmitted light, and processing the image as a comparison to adjust (e.g., verify, supplement, and / or correct) the constructed three-dimensional flow information of the biological sample 116.

[0054] To obtain raw data (raw images), the device 100 is configured accordingly. The wavelength of the laser, the thickness of the illumination light curtain, and the acquisition frame rate can be optimally configured for different biological samples 116. The thickness and length of the illumination light curtain can be calculated by equations (1) and (2) respectively:

[0055]

[0056] where w0 is the Gaussian beam waist radius, λ is the laser wavelength, NA IO is the numerical aperture of the illumination objective, Z R is the Rayleigh range, and n is the refractive index of the biological sample. The effective beam thickness is 2w0, and the effective beam length of the uniform illumination is 2Z RA thicker illumination light curtain is associated with a lower axial resolution but a larger available field of view, and vice versa. The thickness of the illumination light curtain also has a great impact on the optical sectioning ability of the system and the accuracy of blood flow velocity quantification. The light curtain characteristics can be adjusted by changing the size of the adjustable aperture of the first aperture 110. The sCMOS camera 136 is used to collect the original laser speckle images, while the CMOS camera 140 is used to collect wide-field transmission images. In this embodiment, both cameras are configured to be triggered by a NIDAQ data acquisition card for synchronous image acquisition, and the frame rate and exposure time can be independently set in the LabVIEW-based software designed for image acquisition and system control. For LSH-LSI imaging, the sCMOS frame rate determines the upper limit of the blood flow velocity that can be measured. Therefore, faster blood flow generally requires a higher frame rate. On the other hand, an excessively high acquisition frame rate can be avoided because it places an unnecessary strain on system resources and slows down the post-processing process.

[0057] In traditional autocorrelation analysis, the intensity autocorrelation function g2(τ) is related to the speckle decorrelation time (τ C ) using Equation (3):

[0058]

[0059] where g2(τ) is the intensity autocorrelation function of the time delay τ, g1(τ) is the corresponding field autocorrelation function, β is a correction factor related to the measurement geometry, and n is another model parameter depending on the type of motion of the light scatterers and the dynamic laser scattering mechanism. Typically, n takes values of 0.5, 1, and 2, corresponding to multiple scattering disordered motion (MU), multiple scattering ordered motion (MO), or single scattering disordered motion (SU) and single scattering ordered motion (SO) states, respectively. Therefore, the fitted decorrelation time is converted to the flow velocity using Equation (4)

[0060]

[0061] where λ is the laser wavelength, and NA DO is the effective numerical aperture of the collection objective 128.

[0062] Due to the inference between scattered photons of the same wavelength as the illumination light curtain, the original images captured by the sCOMS camera 136 contain an inherent speckle pattern. Statistical analysis of the speckle pattern can yield parameters related to the velocity of local microscopic scatterers. In this embodiment, laser speckle contrast analysis, temporal autocorrelation analysis, and time-frequency analysis are used to process the original image sequence to generate the corresponding flow velocity maps.

[0063] In laser speckle contrast imaging analysis, the speckle contrast K is defined using Equation (5):

[0064]

[0065] where σ is the standard deviation, is the average intensity. The standard deviation σ and the average intensity Both can be estimated in space or time.

[0066] In the autocorrelation analysis, the intensity temporal autocorrelation function g2(0,t,τ2) can be derived at any time t using Equation (6):

[0067]

[0068] where I is the speckle intensity, τ is the time delay, and Δ is the width of the time window for averaging.

[0069] Either the speckle contrast K or the autocorrelation function g2(0,t,τ2) can be used to estimate the instantaneous and local motion of the scatterers, and thus the flow information. In the traditional autocorrelation analysis, using Equation (3) as described above, the intensity autocorrelation function g2(τ) is related to the speckle decorrelation time (τ C ).

[0070] To improve Equation (3), a hybrid theoretical model with two independent decay terms and one modulation term is adopted to calculate the intensity autocorrelation g2(τ), as in Equation (7):

[0071]

[0072] where the first exponential term is related to the fitting weight w < and the single-scattering disordered motion (n = 1) with a decorrelation time of τ < while the second exponential term corresponds to another fitting weight w > and another single-scattering ordered motion (n = 2) with another decorrelation time of τ > ; the modulation term Acos(2πf0τ) is related to the frequency shift f0 and the fitting parameter A. In the case where the fitting parameter A is much less than 1, the fitting decorrelation time τ > is thus converted to the flow velocity v using Equation (4), i.e.,

[0073] Otherwise, the frequency shift f0 is converted to the blood cell velocity v through Equation (8):

[0074]

[0075] where n5 is the tissue refractive index and θ is the angle between the illumination light curtain and the flow direction (vessel direction).

[0076] In addition to the temporal autocorrelation analysis (Equation (6)), time-frequency analysis methods can be used to reliably estimate the local flow velocity and its spatial distribution. In this embodiment, the dynamic change of the light intensity is in Matlab TM Processed pixel by pixel using the time - frequency analysis function "pspectrum" in Chinese to calculate the short - time power spectrum estimate. For example, FIG. 11(b) shows the time - dependent power spectrum estimated from the time - series signal plotted in FIG. 11(a). The representative frequency shift f0 of each time window is simply the maximum frequency at which the power spectral density exceeds a threshold empirically estimated based on the system noise level. Convert the local flow velocity from the frequency shift using Equation (8).

[0077] Specifically, to test the laser speckle imaging device 100, a scanning experiment was performed on the head and trunk regions of zebrafish larvae at 4 days post - fertilization (dpf), where the vascular system has a rather complex three - dimensional structure. Figure 5 Depicts a transmission image of the slice orientation of a 4 - dpf zebrafish larva using the device 100 in Chinese and indicates the direction of movement of the stage 118. During the scanning process for laser speckle image acquisition, the head - to - tail central axis of the zebrafish larva was oriented perpendicular to the axis of the actuator 126 (see the arrow 154 in Figure 1 ). The dashed line 155 indicates the position where the illumination light curtain intersects the focal plane of the first camera 136. The stage 118 (and thus the zebrafish larva) was gradually moved from left to right at a translation interval of 25 μm. Thus, the illumination light curtain moves from right to left within the animal (zebrafish) model. The central position of each scanning step is indicated by one of the dashed lines (once) superimposed on the transmission image shown in Figure 9 , i.e., from number 1 to number 15 in the direction of arrow C. For each of the 15 light - curtain positions, the raw laser speckle images were acquired at a speed of 1500 fps for 2 seconds. Overall, the image acquisition time was approximately 8 minutes, including data transfer from the camera memory to the computer (264×476 pixels, 3000 frames per step), stage translation, and stabilization. In this experiment, to focus on obtaining morphological information, simple image - processing methods were used to reconstruct 15 β - maps to depict the vascular system. These reconstructed 2D images were aligned and combined into a 3D image stack, with each slice mapped to a thin layer of the tissue volume being probed. The normal distance between adjacent slices was approximately 25 μm. Figure 5

[0078] ​The raw images of each of the 15 light curtain slices are processed using the time - frequency domain analysis method described above, and a sequence of blood flow images for each slice is generated. The slice - by - slice angiographic image is obtained by averaging the flow velocity over a period of time. The angiographic image is further processed using interpolation to generate a finer three - dimensional stack of 300 frontal images. The depth interval is 1 micron, and the total depth range is 300 μm. Figures 6(a) - (d) depict four representative frames of the 300 interpolated frontal images. Figures 6(a) - (d) depict the images taken at the bottom layer, lower middle layer, upper middle layer, and top layer, respectively. Figure 6(b) shows that the heart 148 is identified, Figure 6(c) shows that the large blood vessel 150 is identified, and Figure 6(d) shows that the small blood vessel 152 is identified.

[0079] Although the original image stack consists of only the Figure 5 15 slices shown, the quality of the 3D rendering seems quite satisfactory. By scanning the same tissue volume with a finer step size and a reduced illumination light curtain thickness, better axial resolution, closer to the lateral resolution, will be obtained.

[0080] Figure 7 An angiographic image of a single slice is depicted. Four test points in Figure 7 are selected to estimate the local blood flow velocity. The first test point is located in the heart region, and the data of the first test point is called Data 1. The second test point is in the dorsal aorta (DA) region, and the data of the second test point is called Data 2. The third and fourth test points are located in the region where the downstream artery is located, and the data of the third and fourth test points are called Data 3 and Data 4, respectively. Figure 8 depicts the flow velocity waveforms detected at the four test points in Figure 7 . Data 1 is represented by a solid line in Figure 8 . Figure 8 shows that the flow velocity waveform of Data 1 has two peaks in each cardiac cycle. The first peak is associated with the inflow of blood cells during diastole, while the second peak is associated with the outflow during systole. Data 2 is shown as a dashed line in Figure 8 , which shows a rapid rising edge of the blood flow in the DA region, which is almost coincident with the cardiac outflow peak of Data 1. However, it reaches a peak velocity value much higher than that of Data 1. Data 3 is represented by a dotted line in Figure 8 , and Data 4 is represented by a dashed line in Figure 8 . Data 3 and Data 4 show a small time delay in the rising edge and a gradual decrease in the peak velocity. This indicates that the present invention is capable of providing a quantitative measurement of the flow velocity. For illustrative comparison, PIV analysis is performed to track blood cells and generate a flow map.

[0081] Using Figure 1 The device 100 shown simultaneously acquires laser speckle images and wide-field transmission images from the trunk region of 3-dpf zebrafish larvae used as biological samples 116. Figure 9 is an imaging geometry that depicts the arrangement of the biological sample 116 with reference to the illumination light curtain. Figure 9 Depicts a cross-sectional view of the trunk structure sliced by the illumination light curtain, with the flow map superimposed on the morphology. As Figure 9 shown, the 3-dpf zebrafish larvae are positioned such that the illumination light curtain intersects the trunk at an angle. The main blood vessels in the trunk, namely the DA and the posterior cardinal vein (PCV), are horizontally oriented to facilitate the acquisition of high-quality wide-field transmission images at 200 fps.

[0082] PIV analysis is performed using PIVlab TM (a GUI (Graphical User Interface)-based Matlab TM program) on the wide-field transmission images, which is designed for particle image velocimetry to plot the flow velocity. The procedure is as follows. First, the stack of transmission images is imported into PIVlab TM . Second, an in-built algorithm is selected for cross-correlation analysis. Generally, the direct Fourier transform correlation with multi-channel and deformed windows (FFT window deformation) is preferred compared to single-channel direct cross-correlation (DCC) and global correlation. Third, the selected analysis is configured and executed, and the results are then calibrated using separately acquired calibration images. Finally, the velocity distribution in the field of view and the instantaneous velocity waveforms in specific regions of interest are generated. Optionally, the vascular network can also be depicted by further processing the velocity map.

[0083] The original laser speckle images were acquired at 3000 fps, which resulted in a stack of 6000 frames in 2 seconds. Fig. 10(a) is an exemplary original laser speckle image (scale bar: 50 μm). The original laser speckle image is a cross-sectional intensity image, where rapid intensity fluctuations can be seen in the large blood vessel region. The high spatio-temporal resolution in the original image stack provides important insights into complex dynamic signal features and is highly beneficial for signal processing and interpretation. To find the time-averaged local flow velocity, the temporal speckle signals were picked up pixel by pixel from the image stack, and the corresponding autocorrelation function was calculated over the entire 2-second time window. Fig. 10(b) is an exemplary optical intensity signal (temporal speckle) picked up from a pixel in the DA region of Fig. 10(a); Fig. 10(c) depicts the autocorrelation function of the signal in Fig. 10(b). The results show that the intensity autocorrelation g2(τ) does not conform to a simple exponential model. The slow fluctuation pattern indicates deterministic correlations on a long time scale, which are caused by the periodic intensity variations of the light scattered by blood cells arriving at the cross-section discretely. Fig. 10(d) depicts a smaller delay time window of the result of the model fitting Fig. 10(c). In the smaller delay time window (Fig. 10(d)), two different exponential decay patterns can be seen in the experimental data (dots), and the curve can be best fitted by applying Equation (7) with the modulation term being zero (or the parameter A approaching zero), and the intensity autocorrelation g2(τ) was calculated as Here, the first exponential term (dashed line in Fig. 10(d)) corresponds to single-scattering disordered motion (n = 1), with a relatively long decorrelation time of 20.6 ms, and the W r value is 0.602. It is related to the slow background random motion of blood cells and the whole animal with an effective velocity of 39.7 μm / s. The second exponential term (solid line in Fig. 10(d)) is associated with single-scattering ordered motion (n = 2), with a much smaller decorrelation time of 3.66 ms, and the W f The value is 0.426, which is converted to a flow velocity of 223 μm / s. Figure 10(e) depicts the velocity map derived from the fitted parameters. The white arrow 156 indicates the DA region with a higher flow velocity, and the white triangle 158 points to the PCV region with a relatively lower flow velocity. To quantify the dynamic blood flow, a much shorter time window was used for the decorrelation time estimation, and the quantified time-correlated velocity is plotted in Figure 10(f). Figure 10(f) depicts the short windows of the flow velocities in the DA and PCV regions analyzed by LSI and PIV analyses. As shown in Figure 10(f), the velocities in the DA and PCV are very consistent between the LSI results and the PIV results. The above analysis is based on the entire 2-second time window and yields time-averaged flow measurements. The arterial flow velocity 160 of the LSI results swings between the extreme values during each cardiac cycle. The peak is approximately 10 times higher than the trough. On the other hand, the venous flow 162 of the LSI results appears to be more continuous. The average (over 2 seconds) flow velocity estimated by LSH-LSI is 135.2 μm / s in the PCV and 259.8 μm / s in the DA. In contrast, the PIV analysis yields an average velocity of 145.6 μm / s for the venous flow and 254.4 μm / s for the arterial flow. The waveforms of the PIV velocities in the DA 164 and the PIV velocities in the PCV 166 are also depicted in Figure 10(f). As can be seen from Figure 10(f), the LSI and PIV results match each other very well. For example, the arterial blood flow has very similar dynamic characteristics in terms of the peak, trough, pulse width, and rising and falling edges. Based on the statistical analysis of five pulses, the peak arterial flow velocity is 636.9 ± 79.6 μm / s (LSI) or 612.4 ± 30.9 μm / s (PIV).

[0084] Figures 11(a)-(c) depict the time-frequency analysis process. Figure 11(a) depicts the same light intensity signal as shown in Figure 10(b). Figure 10(b) shows that the intensity fluctuates in a seemingly random manner. However, the instantaneous oscillation frequency has a periodic pattern. Figure 11(b) depicts the short-time power spectrum of the time-frequency analysis by confirming the periodicity. The frequency shift is caused by the interference between the light waves scattered from the stationary tissue (e.g., blood vessel wall) and the moving red blood cells. They can be easily converted into local and instantaneous flow velocities. The representative frequency shift f0 of each time window is simply the maximum frequency at which the power spectral density is higher than the threshold empirically estimated based on the system noise level. The local flow velocity is converted from the frequency shift through Equation (8). By comparing the results of LSH-LSI and transmissive PIV, the quantitative accuracy of this simple frequency-domain analysis method is verified. Figure 11(c) depicts the comparison between the DA blood flow velocities obtained from LSH-LSI and PIV, which reflects the highly pulsatile arterial (DA) blood flow waveform. As shown in Figure 11(c), similar to that shown in Figure 10(f), the results of LSH-LSI (dashed line) and PIV (solid line) have very similar dynamic characteristics in terms of peak, valley, pulse width, and rising and falling edges. Based on the statistical analysis of five pulses, the peak arterial flow velocity is estimated to be 595.7 ± 23.1 μm / s (LSH-LSI) or 612.4 ± 30.9 μm / s (PIV). Taking the average within a 2-second time window, the flow velocity in the DA is 254.3 μm / s (LSH-LSI) or 254.4 μm / s (PIV). The present invention demonstrates a good ability to quantitatively measure dynamic flow velocities.

[0085] Experimental imaging data were collected from 5-dpf zebrafish larvae, and the blood flow around the heart region was analyzed to demonstrate the ability of the LSH-LSI system in vector velocity mapping. LSH-LSI scalar velocity maps and transmissive images were obtained, including the heart region, the bulbous arteriosus region, and the branchial arch region. PIV analysis was applied to the LSH-LSI scalar velocity maps and transmissive images to generate corresponding vector flow velocity maps in the three regions, which are paired and compared in Figures 12(a)-(f). Figures 12(a), 12(c), and 12(e) depict the laser speckle vector velocity maps, and Figures 12(b), 12(d), and 12(f) depict the transmissive velocity maps, respectively. Figures 12(a) and 12(b) are the heart region, Figures 12(c) and 12(d) are the bulbous arteriosus region, and Figures 12(e) and 12(f) are the branchial arch region. Figures 12(a)-(f) show that the velocity magnitudes and directions derived from the two imaging modalities are basically consistent.

[0086] The various advantages of the present invention can be understood from the foregoing description. Through novel optical design, system optimization, and appropriate flow quantification algorithms, as shown above, the present invention has the capabilities of three-dimensional imaging, dynamic flow velocity quantification, and vector flow mapping, with optimal 3D microcirculation imaging, high spatial and temporal resolution, fast imaging rate, and accurate quantitative flow velocity assessment. This is a simple yet robust quantitative means for 3D visualization of microvascular and blood flow characteristics.

[0087] For transparent and semi-transparent biological samples, device 100 utilizes the inherent optical sectioning ability of selected plane illumination to achieve tomographic, in vivo, and three-dimensional imaging of vascular structures and blood flow velocity distributions with high spatio-temporal resolution. The above-mentioned zebrafish larva imaging experiments conducted with device 100 have revealed complex laser speckle dynamics, and the proposed model above helps to accurately retrieve the decorrelation time related to flow velocity.

[0088] By using one or more illumination light curtains to selectively illuminate the sample area in which flow information will be collected, the selected plane illumination achieves optical sectioning and enables visualization of the 3D flow in the sample layer by layer or, if necessary, slice by slice. With the configuration of light curtain illumination, the device can achieve optical sectioning even without spatial filtering (as in confocal microscopy) or digital post-processing (as in structured illumination microscopy). Different from fluorescence light curtain microscopy, LSH-LSI is based on an inherent contrast mechanism (light scattering) and is a label-free imaging platform. LSH-LSI provides an excellent solution for studying fluid dynamics (such as blood flow) in complex 3D networks.

[0089] The illumination light curtain is disposed obliquely with respect to the normal direction of the surface of the stage 118, or in other words, the angle between the direction of the illumination light curtain and the direction of the scattered light is preferably greater than 0 degrees and less than 90 degrees (more preferably in the range of 30 degrees to 60 degrees), which enhances the detected signal. The forward scattered signal captured in the obliquely oriented LSH-LSI is many orders of magnitude stronger than the backscattered light in a confocal setting. This provides flexibility in configuring the image acquisition speed and exposure time without worrying about the photon budget. In the above imaging experiment, when the illumination light power is less than 10 mW, the camera exposure time can be set as low as a few microseconds (e.g., 5 μs). The high-speed imaging ability is ideal for increasing the dynamic range of flow velocity measurement. However, this is not its only benefit. It is also capable of rapid data acquisition and can reduce the time to determine the local flow velocity to less than 1 millisecond. In contrast, in PIV-based methods, the particles being tracked need to move a significant distance between image frames so that the flow velocity can be accurately estimated. Therefore, transmission images are typically captured at a much lower speed. A short acquisition time for individual slices is highly desirable to increase the total throughput, especially in 3D mapping. In addition, the high imaging speed of LSH-LSI makes it possible to obtain the local flow direction from an instantaneous scalar velocity map that is not time-averaged. As demonstrated above, LSH-LSI is capable of generating a vector velocity map based on additional processing of scalar laser speckle velocimetry results acquired at a high frame rate. The extended function of LSH-LSI to determine the local flow direction is particularly useful when studying the hydrodynamics of the cardiac region, where the movement of blood cells is not confined to one-dimensional small blood vessels.

[0090] The reconstruction of 3D angiography helps to restore morphological information because the initially estimated velocity is relative rather than absolute. As described above, the strong optical sectioning ability of the present invention and the proper separation of single-scattered photons and multiply-scattered photons also enable quantitative measurement of local flow velocity. By using a high-speed camera that captures the scattered light (with the same wavelength as the illumination light) from the sample, the high-speed camera allows for rapid acquisition of the original image sequence from which quantitative flow information can be retrieved. The angle between the illumination axis and the detection axis balances the detection of forward scattered photons, which are much stronger than backscattered photons, and has a depth of focus for wide-field image acquisition. The detection sensitivity and selectivity of the present invention are improved. Therefore, the present invention overcomes the disadvantages of traditional LSI methods, which can only provide qualitative results to indicate relative changes, and precise measurement is very helpful for experimental and computational fluid dynamics analysis, especially as a non-invasive and non-contact method.

[0091] The photons detected by the sCMOS camera 136 have undergone a single scattering event with a relatively small scattering angle, approximately 60 degrees. As described above, for biological tissues, the scattering of visible light is dominated by forward scattering. Therefore, configuring the detection optical path to be vertical on top of the stage 118 (as Figure 1 shown) may help capture more scattered photons, which may improve image quality and image acquisition speed because, in this configuration, the scattered light is forward-scattered light. Thus, most of the singly scattered photons propagate along a direction close to the illumination optical axis, i.e., close to Figure 2 the 0° direction shown. This indicates that light scattering in biological soft tissues is generally dominated by forward scattering that favors small scattering angles. For example, the typical size of red blood cells is about 10 μm, and for an incident beam of 640 nm, the anisotropy factor (average cosine of the scattering angle) is estimated to be about 0.95. However, if the angle between the scattered photons and the 0° direction is too small, the illuminated sample plane may not match the depth of focus of the detection optics. In this embodiment, the optical axis ④ is configured to be approximately 60°, such that both the size of the scattered light collected by the collection objective 128 and the depth of focus of the detection optics are taken into account. This angle can be arranged by considering different factors, such as the optical parameters of the device 100 and the nature of the biological sample. For example, for applications in mouse brain and chicken embryo imaging, the illumination and detection optics can be arranged on the same side of the stage 118, such as the illumination optics with optical axis ① and the detection optics with optical axis ②, as Figure 1 shown. In view of the above, the angle between the optical axis of the illumination light curtain and the optical axis of the collected scattered light preferably varies within the range of 45° to 135°.

[0092] Due to various uncertainties, model fitting in LSI has always been a tricky process. The described embodiments and the above research reveal the complex nature of dynamic scattering signals, which require sophisticated non-traditional theoretical modeling / processing. The LSH-LSI system provided by the present invention is an excellent platform for researchers to further advance the basic theory and instrument design of quantitative laser speckle imaging. The system is enhanced by integrating a transmission imaging subsystem, which provides in vivo independent cross-validation and calibration means.

[0093] The time-frequency analysis method proposed above yields a robust imaging processing algorithm for quantifying local flow velocity. It is particularly suitable for microcirculation imaging because blood cells move near microvessels. Therefore, the Doppler signal becomes strong enough for accurate frequency shift estimation.

[0094] In short, the LSH-LSI described in this embodiment is a label-free, three-dimensional, quantitative, and high spatio-temporal resolution blood flow imaging tool that addresses the limitations of existing imaging methods, which either require fluorescent labeling / tracer particle seeding or are limited to two-dimensional image acquisition.

[0095] Although Figure 1 the laser speckle imaging device 100 is depicted as the first embodiment of the present invention, it is contemplated that the LSH-LSI system of the present invention may be configured and implemented in different ways.

[0096] Figure 13 The laser speckle imaging device 200 according to the second embodiment of the present invention is described. Since most of the devices used in the second embodiment are the same as those in the first embodiment, they are labeled with the same reference numerals. Reference may be made to the above description, where elements, devices, configurations, and / or functions not described in this embodiment are concerned. Along the optical path, the first difference is that, in this embodiment, the illumination optical device 115 does not include Figure 1 the mirror 108 and the first aperture 110, such that the incident light beam output by the beam expander 106 is guided through the cylindrical lens 112. As Figure 13 shown, the beam expander 106, the cylindrical lens 112, and the illumination objective 114 are coaxial. Another difference is that the first image acquisition device 135 does not include Figure 1 the second aperture 130, such that the collection objective 128 guides the scattered light directly through the tube lens 132. Another difference is that the transmission optical device 137 further includes a tube lens 202. The transmitted light from the LED 138 passes through the tube lens 202 before reaching the prism 124. Similar to Figure 1 that, Figure 13 it is also shown that the optical axis of the collection objective 128 is vertical, i.e., perpendicular to the top surface of the substantially horizontal stage 118. In addition, Figure 13 two alternative positions of the first image acquisition device 135 are also shown, which are approximately 45 degrees with respect to the top surface of the stage 118.

[0097] Figure 14 A schematic diagram of a light curtain imaging device 300 according to the third embodiment of the present invention is depicted. Similar to the second embodiment, devices in the third embodiment that are the same as those in the first embodiment are labeled with the same reference numerals. Reference may be made to the above description, where elements, devices, configurations, and / or functions not described in this embodiment are concerned. As Figure 14 As shown, similar to the first embodiment, the illumination optical device 115 includes a cylindrical lens 112 and an illumination objective lens 114 in the illumination optical path, and the illumination optical device 115 is operable to generate a light curtain for illumination. The illumination optical device 115 further includes a rotatable scanning mirror 302 and a grating element 304. The rotatable scanning mirror 302 is a galvanometer mirror 302 in this embodiment, and the grating element 304 is a transmission grating 304 in this embodiment. The incident light beam (converging in the direction perpendicular to the paper) after passing through the cylindrical lens 112 is reflected by the galvanometer mirror 302. When the galvanometer mirror 302 is driven to rotate back and forth, the reflected light beam after the galvanometer scanner is inclined at a time-dependent angle with respect to the detection optical axis of the collection objective lens 128. The illumination objective lens 114 further focuses the reflected light beam in the horizontal direction, and the reflected light beam can be moved left and right to cover the region of interest of the biological sample 116. As Figure 14 shown, the transmission grating 304 is located above the illumination objective lens 114 and below the biological sample 116. The incident light beam is diffracted by the transmission grating 304 to form two inclined illumination light curtains within the biological sample 116. The inclination angle with respect to the detection optical axis of the collection objective lens 128 depends on the light wavelength and the grating groove pitch. The typical range of the inclination angle is 30 to 60 degrees. During the scanning process, the operator can move the two inclined illumination light curtains from left to right (or in the opposite direction) at a high speed provided by the galvanometer mirror 302. Optionally, the first image acquisition device 135 further includes a filter 306, which is an emission filter 306 in this embodiment. The emission filter 306 can be included in the detection optical path after the collection objective lens 128 for fluorescence microscopy imaging. For label-free scattering-based imaging, the emission filter 306 can be removed.

[0098] Compared with the first and second embodiments, the grating-based method in the third embodiment helps to improve the image acquisition speed. For the first and second embodiments, the biological sample 116 is scanned by moving the stage 118 with the actuator 126.

[0099] Figure 15 A schematic diagram of a light curtain imaging device 400 according to a fourth embodiment of the present invention is depicted. The main part of the fourth embodiment is the same as that of the third embodiment, except that the illumination optical device 115 does not include Figure 14 Rather than the cylindrical lens 112, it includes a spherical lens 402 and a lens array 404, which is a cylindrical microlens array 404 in this embodiment. The spherical lens 402 and the cylindrical microlens array 404 can be used to modulate light. The cylindrical microlens array 404 can be used to split an incident light beam into two or more incident sub-beams. The spherical lens 402 receives two or more incident sub-beams and directs them to the rotatable scanning mirror 302. The rotatable scanning mirror 302 reflects one or more incident sub-beams to the illumination objective 114 for the illumination objective 114 to laterally receive two or more incident sub-beams at a distance (Δ). In the case where there are more than two incident sub-beams, the distance (Δ) between any two adjacent incident sub-beams received by the illumination objective 114 is the same. The illumination objective 114 focuses each incident sub-beam in the horizontal direction and can move left and right to cover the region of interest of the biological sample 116.

[0100] Therefore, the scanning range of the galvanometer mirror 302 is reduced to only cover the small distance Δ, which can be one or two orders of magnitude smaller than the length of the entire field of view. Compared with the third embodiment, the implementation of the fourth embodiment is particularly suitable for relatively thin samples and helps to further improve the imaging speed. Compared with the third embodiment, the grating- and lens-array-based method in the fourth embodiment helps to further improve the image acquisition speed. For the third embodiment, the biological sample 116 is scanned layer by layer in batches, with each batch including multiple layers.

[0101] Although in the first embodiment it is described that the illumination optical device 115 includes a laser diode 102, a collimator 104, a beam expander 106, a mirror 108, a first aperture 110, a cylindrical lens 112, an illumination objective 114, and a prism 124, it is conceivable that in different embodiments, the illumination optical device 115 can have different configurations as long as it can generate an illumination light curtain as needed. For example, it may not include the mirror 108 and the first aperture 110 as Figure 13 shown, and it may include different devices for different configurations. For example: in the Figure 14 third embodiment shown, it may include a light source ( Figure 14 not shown in the figure), a cylindrical lens 112, a galvanometer mirror 302, an illumination objective 114, and a grating element 304; in the Figure 15 fourth embodiment shown, it may include a light source ( Figure 15 not shown in the figure), a cylindrical microlens array 404, a spherical lens 402, a galvanometer mirror 302, an illumination objective 114, and a grating element 304.

[0102] Although in the first embodiment it is described that the first image acquisition device 135 includes an objective lens 128, a second aperture 130, a tube lens 132, a dichroic mirror 134, and a first camera 136, it is conceivable that in different embodiments, the first image acquisition device 135 may have different configurations as long as it can collect optical signals and generate images as needed. For example, it may not include the second aperture 130 as shown in Figure 13 ; it may not include the second aperture 130, the tube lens 132, and the dichroic mirror 134 as shown in Figure 14 and Figure 15 , but may further include an emission filter 306.

[0103] Although in the first embodiment it is described that the stage 118 includes a standard glass bottom dish 120, a glass slide 122, and an actuator 126, it is conceivable that in different embodiments, the stage 118 may have different configurations as long as it can support the biological sample 116 as needed. For example, in the third and fourth embodiments shown in Figure 14 and Figure 15 , when the rotatable scanning mirror 302 can be used to move the illumination position for scanning, the actuator 126 may not be required; or alternatively, if needed, both the actuator 126 and the rotatable scanning mirror 302 can be included in one embodiment.

[0104] Although in the first embodiment it is described that the transmission optical device 137 includes a prism 124 and a second light source 138, it is conceivable that in different embodiments, the transmission optical device 137 may have different configurations as long as it can provide transmitted light as needed. For example, it may include a second light source 138, a tube lens 202, and a prism 124, as shown in Figure 13 .

[0105] Although in the first embodiment it is described that the second image acquisition device 139 includes an objective lens 128, a second aperture 130, a tube lens 132, a dichroic mirror 134, and a second camera 140, it is conceivable that in different embodiments, the second image acquisition device 139 may have different configurations as long as it can collect optical signals and generate images as needed. For example, it may not include the second aperture 130 as shown in Figure 13 .

[0106] Although the tilted configuration of the illumination light curtain has advantages, it is conceivable that the LSH-LSI may have different configurations. For example, in the case of sufficient photon budget, a traditional orthogonal detection geometry can be implemented to allow faster depth scanning without using the translation stage 118, that is, the collection optics are tilted (as shown by the dashed arrow 168 in Figure 13 ) to form a right angle with the illumination axis.

[0107] While the light source used in the first embodiment is a laser diode having a light output at a central wavelength of 640 nm, it is contemplated that light sources producing light outputs at other wavelengths may also be used. Similarly, light sources having different central wavelengths from that of the green LED may be used for LED 138.

[0108] While the stage in the first embodiment is used to mount a standard glass-bottom dish, it is contemplated that the stage may be configured to mount other suitable devices adapted to hold a desired biological sample.

[0109] While the apparatuses 100 of the above-described first embodiment and the second embodiment indicate the use of a transmission optical path to capture wide-field images for facilitating sample processing, PIV image analysis, and for providing additional information about the sample, it is contemplated that the transmission optical path may not be included in apparatuses of different embodiments.

[0110] While the lens array 404 is described in the fourth embodiment as including a cylindrical microlens array 404, it is contemplated that the lens array 404 may include a cylindrical lens array or other suitable lens arrays.

[0111] While the biological sample used in the above description is a zebrafish embryo / larva, it is contemplated that any biological material / body capable of scattering light or transmitting light may be used as a biological sample. While LSH-LSI is an excellent platform for zebrafish embryos and larvae, it is contemplated that this platform may be applicable to flow imaging of other small animal models, such as mouse and Drosophila larvae, and even further to medical applications where three-dimensional and quantitative label-free flow imaging is crucial (e.g., in vivo microcirculation imaging applicable to human subjects). Additionally, both the illumination and detection optics may be movable above the sample stage to accommodate other animal models that are less transparent.

[0112] The high spatial resolution and temporal resolution make the present invention a perfect imaging solution for a wide range of applications. It is contemplated that the systems and methods of the present invention may be applied to many fields, such as biomedical research tools for high-quality, non-invasive visualization of microscale to macroscale flows, having high spatial resolution in three dimensions and high temporal resolution for dynamic flow measurements including medical applications.

Claims

1. A laser speckle imaging device for generating flow information of a biological sample, comprising: An illumination optical device that can operate to generate one or more illumination light curtains for selectively illuminating a biological sample to generate corresponding scattered light; A first image acquisition device that can operate to acquire the scattered light corresponding to each illuminated layer at the same wavelength as the illumination light curtain; And An image processing device that can operate to construct three-dimensional flow information of the biological sample from the speckle pattern of the acquired scattered light.

2. The device according to claim 1, wherein The illumination optical device includes a grating element that can operate to receive an incident beam and divide the incident beam into at least two illumination light curtains for simultaneously and selectively illuminating the biological sample.

3. The apparatus according to claim 2, wherein The grating element includes a transmission grating.

4. The device according to claim 2 or 3, wherein, The illumination optical device includes a cylindrical lens for generating the incident beam.

5. The device according to claim 2 or 3, wherein, The incident beam includes a plurality of incident sub-beams, and the illumination optical device includes a cylindrical lens array for generating a plurality of incident sub-beams for the grating element to divide each incident sub-beam into at least two illumination light curtains.

6. The device according to any one of claims 2 to 5, wherein, The illumination optical device further includes a rotatable scanning mirror that can operate to adjust the angular direction of the incident beam.

7. The apparatus according to claim 6, wherein, The rotatable scanning mirror includes a galvanometer mirror.

8. The apparatus according to any one of the preceding claims, wherein, The angle between the optical axis of each illumination light curtain and the optical axis of the acquired corresponding scattered light is between 0 degrees and 90 degrees, or between 30 degrees and 60 degrees.

9. The device according to claim 8, wherein The first image acquisition device further includes a diaphragm with an adjustable aperture for adjusting the scattered light.

10. The device according to claim 1 or 8 or 9, when dependent on claim 1, wherein The illumination optical device further includes a prism that can operate to transmit the one or more illumination light curtains to the biological sample.

11. The device according to any one of the preceding claims, wherein, The first image acquisition device further includes an emission filter that can operate to allow desired fluorescence to pass through and reach the first image acquisition device.

12. The device according to any one of the preceding claims, further comprising: A transmission optical device that can operate to generate a transmission beam for illuminating the biological sample, the transmission beam having a wavelength different from that of the illumination light curtain; And A second image acquisition device that can operate to acquire the corresponding transmission light of the biological sample having the same wavelength as the transmission beam and generate a transmission image, Wherein the image processing device can operate to adjust the three-dimensional flow information of the constructed biological sample based on the transmission image.

13. A laser speckle imaging method, comprising: Generating one or more illumination light curtains to selectively illuminate one or more layers of a biological sample, thereby generating corresponding scattered light; Acquiring the corresponding scattered light having the same wavelength as the illumination light curtain; And Constructing three-dimensional flow information of the biological sample from the speckle pattern of the acquired scattered light.

14. The method according to claim 13, wherein, Illuminating the biological sample with one illumination light curtain at a time, illuminating the biological sample layer by layer, and generating scattered light corresponding to each layer.

15. The method according to claim 14, wherein, Illuminating the biological sample with two or more illumination light curtains simultaneously to form a batch of illuminated layers of the biological sample; wherein the generated scattered light corresponds to each layer in the batch.

16. The method according to claim 15 further comprises adjusting the positions of the two or more illumination light curtains to produce another illuminated layer of the biological sample.

17. The method according to claim 16, wherein, A galvanometer mirror is used to adjust the position of the illuminated layer.

18. The method according to any one of claims 13 - 17 further comprises: generating a transmitted light beam for irradiating the biological sample to produce transmitted light, the transmitted light beam having a wavelength different from that of the illumination light curtain; acquiring transmitted light having the same wavelength as the transmitted light beam; and adjusting the three - dimensional flow information of the constructed biological sample based on the acquired transmitted light.

19. A non - transitory computer - readable storage medium for storing a computer program, which when executed by a processor, executes the laser speckle imaging method according to any one of claims 13 to 18.

20. An apparatus for imaging a biological sample, comprising: An illumination optical device, comprising: a rotatable scanning mirror capable of operating to adjust the angular direction of an incident light beam to produce a reflected light beam; and a grating element capable of operating to divide the reflected light beam into at least two illumination light curtains for simultaneously and selectively illuminating a biological sample to produce corresponding light; and an image acquisition device capable of operating to acquire the corresponding light of each illuminated layer for imaging the biological sample.

21. The apparatus according to claim 20 further comprises an image processing device capable of operating to construct three - dimensional imaging information of the biological sample based on the contrast pattern of the acquired corresponding light.

22. The device according to claim 20 or 21, wherein The grating element comprises a transmission grating.

23. The apparatus according to any one of claims 20 to 22, wherein, The rotatable scanning mirror comprises a galvanometer mirror.

24. The apparatus according to any one of claims 20 to 23, wherein, The incident light beam comprises a plurality of incident sub - beams, and the illumination optical device comprises a cylindrical lens array for generating a plurality of incident sub - beams for the grating element to divide each incident sub - beam into at least two illumination light curtains.

25. The device according to any one of claims 20 to 23, wherein The illumination optical device comprises a cylindrical lens for generating the incident light beam.

26. The apparatus according to any one of claims 20 to 25, wherein The image acquisition device further comprises an emission filter capable of operating to filter the acquired corresponding light to allow desired fluorescence to pass through to the image acquisition device.

27. The device according to any one of claims 20 to 26, wherein, The angle between the optical axis of each illumination light curtain and the optical axis of the acquired corresponding light ray is between 30 degrees and 60 degrees.

28. A method for imaging a biological sample, comprising: adjusting the angular direction of an incident light beam by a rotatable scanning mirror to produce a reflected light beam; and dividing the reflected light beam into at least two illumination light curtains by a grating element for simultaneously and selectively illuminating the biological sample to produce corresponding light; and acquiring the corresponding light of each illuminated layer for imaging the biological sample.

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