WO2023035281A1 - 一种基于点阵激光扫描的流式成像系统 - Google Patents

一种基于点阵激光扫描的流式成像系统 Download PDF

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WO2023035281A1
WO2023035281A1 PCT/CN2021/118291 CN2021118291W WO2023035281A1 WO 2023035281 A1 WO2023035281 A1 WO 2023035281A1 CN 2021118291 W CN2021118291 W CN 2021118291W WO 2023035281 A1 WO2023035281 A1 WO 2023035281A1
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light
objective lens
module
signal
laser
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PCT/CN2021/118291
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English (en)
French (fr)
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尤政
韩勇
赵精晶
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清华大学
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Priority to US18/249,439 priority Critical patent/US11835442B2/en
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    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01N—INVESTIGATING OR ANALYSING MATERIALS BY DETERMINING THEIR CHEMICAL OR PHYSICAL PROPERTIES
    • G01N15/00—Investigating characteristics of particles; Investigating permeability, pore-volume or surface-area of porous materials
    • G01N15/10—Investigating individual particles
    • G01N15/14—Optical investigation techniques, e.g. flow cytometry
    • G01N15/1429—Signal processing
    • G01N15/1433—Signal processing using image recognition
    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01N—INVESTIGATING OR ANALYSING MATERIALS BY DETERMINING THEIR CHEMICAL OR PHYSICAL PROPERTIES
    • G01N15/00—Investigating characteristics of particles; Investigating permeability, pore-volume or surface-area of porous materials
    • G01N15/10—Investigating individual particles
    • G01N15/14—Optical investigation techniques, e.g. flow cytometry
    • G01N15/1434—Optical arrangements
    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01N—INVESTIGATING OR ANALYSING MATERIALS BY DETERMINING THEIR CHEMICAL OR PHYSICAL PROPERTIES
    • G01N15/00—Investigating characteristics of particles; Investigating permeability, pore-volume or surface-area of porous materials
    • G01N15/10—Investigating individual particles
    • G01N15/14—Optical investigation techniques, e.g. flow cytometry
    • G01N15/1429—Signal processing
    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01N—INVESTIGATING OR ANALYSING MATERIALS BY DETERMINING THEIR CHEMICAL OR PHYSICAL PROPERTIES
    • G01N15/00—Investigating characteristics of particles; Investigating permeability, pore-volume or surface-area of porous materials
    • G01N15/10—Investigating individual particles
    • G01N15/14—Optical investigation techniques, e.g. flow cytometry
    • G01N15/1468—Optical investigation techniques, e.g. flow cytometry with spatial resolution of the texture or inner structure of the particle
    • G01N15/147—Optical investigation techniques, e.g. flow cytometry with spatial resolution of the texture or inner structure of the particle the analysis being performed on a sample stream
    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01N—INVESTIGATING OR ANALYSING MATERIALS BY DETERMINING THEIR CHEMICAL OR PHYSICAL PROPERTIES
    • G01N15/00—Investigating characteristics of particles; Investigating permeability, pore-volume or surface-area of porous materials
    • G01N15/10—Investigating individual particles
    • G01N2015/1006—Investigating individual particles for cytology
    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01N—INVESTIGATING OR ANALYSING MATERIALS BY DETERMINING THEIR CHEMICAL OR PHYSICAL PROPERTIES
    • G01N15/00—Investigating characteristics of particles; Investigating permeability, pore-volume or surface-area of porous materials
    • G01N15/10—Investigating individual particles
    • G01N15/14—Optical investigation techniques, e.g. flow cytometry
    • G01N15/1434—Optical arrangements
    • G01N2015/144—Imaging characterised by its optical setup

Definitions

  • the invention belongs to the technical field of flow imaging, in particular to a flow imaging system based on dot matrix laser scanning.
  • Imaging flow cytometry is an imaging technique for imaging each cell in high-speed flowing cells to obtain microscopic images.
  • Imaging flow cytometer is an instrument for flow imaging. It solves the defect that traditional flow cytometry cannot obtain cell images, and can obtain more abundant cell morphological information, which can be widely used in scientific research and clinical testing.
  • imaging flow cytometry can be divided into two types based on array detectors and single-point detectors, and each type of imaging flow cytometry has many different implementation methods.
  • the imaging method based on the array detector adopts CCD or CMOS as the detection device, the most representative one is the ImageStream series instruments of Amnis Company, and the core device is the time-delay integration CCD.
  • imaging methods based on single-point detectors are currently a class of methods suitable for high-throughput imaging flow cytometry.
  • An important imaging method in the single-point detector method is an imaging method based on spectral labeling method and time-stretching technology.
  • the spectral labeling part uses a dispersive device to map different wavelengths of broad-spectrum light to different positions of the imaged sample.
  • the sample Reflected light or scattered light is "time-stretched" through the dispersive fiber, so that the optical signal carrying the sample image information is separated in the time domain, and then detected by a single photomultiplier tube (PMT), and finally the reconstructed image is restored.
  • This imaging method is limited by the limited dispersion coefficient due to the use of time-stretching technology, so it needs to use a higher-speed acquisition card to complete data acquisition; and because the time-stretching technology requires the use of km-level dispersion fibers, The system complexity and stability are low.
  • Another important imaging method based on the single-point detector method is the compressed sensing imaging method excited by random structured light. This method uses a static random illumination laser spot to irradiate the cells.
  • the excited fluorescence signal contains the fluorescence distribution information of the cells, and then the compressed sensing algorithm can be used to solve the inverse problem to obtain the cell image, or use a machine
  • the learning method directly recognizes the original signal.
  • this imaging method can obtain high imaging throughput, its practicality is limited because it cannot directly acquire cell images and requires a lot of time to solve the inverse problem.
  • the purpose of the present invention is to provide a flow imaging system based on fractional laser scanning.
  • the present invention proposes a flow imaging system based on dot matrix laser scanning, including:
  • the laser light source is used to generate continuous laser light, and the laser light is irradiated to the spectroscopic module;
  • a beam splitting module arranged downstream of the laser light source, the beam splitting module is used to divide the received single beam of laser light into multiple beams of laser beams with spatially dispersed angles or offset positions, and the multiple beams of laser light are transmitted to the illumination objective lens;
  • the fluid focusing module is arranged downstream of the illumination objective lens.
  • the fluid focusing module is used to arrange the dispersed cells to be imaged in the sample into a single cell axial flow. When the cells pass through the illumination area of the two-dimensional lattice spot sequentially, the cells excite scattered light or fluorescence as signal light;
  • An acquisition card arranged downstream of the photodetector is used to collect the voltage signal, convert it into a corresponding digital signal through analog-to-digital conversion, and send the digital signal to the computer module;
  • the computer module is arranged downstream of the acquisition card, and the computer module is used to process the received digital signal to obtain the image of the imaged cell in the fluid focusing module.
  • the invention changes the light beam scanning system in the traditional laser scanning microscope into a laser beam splitting module, and changes the moving scanning light beam into a static dot matrix scanning light beam.
  • Level can achieve imaging throughput of more than 10000cell/s, making the imaging speed fast.
  • the two-dimensional lattice light spots scan and excite the moving cells, the distance between the two-dimensional lattice light spots parallel to the direction of cell movement is greater than the length of the cells, and the distance between the two-dimensional lattice light spots perpendicular to the direction of cell movement
  • the position distribution of the array spot covers different positions of the cells.
  • the two-dimensional lattice light spot of the present invention is composed of laser light spots arranged in a specific way.
  • the light spots are projected onto the plane of the imaging object, and the cells move at a uniform speed in the plane of the imaging object.
  • the distance between the two-dimensional lattice light spots parallel to the direction of cell movement is greater than the length of the cells, and the position distribution of the two-dimensional lattice light spots perpendicular to the direction of cell movement covers different positions of the cells, so that when the cells flow through the light spots Ensure that only one light spot illuminates the cells at a time in the area, and realize the effect of using the cell movement to complete the progressive laser scanning of the cells.
  • a back-facing fluorescence detection optical path is formed; when the light-collecting objective lens and the illuminating objective lens are not the same objective lens, a forward-facing fluorescence detection optical path is formed.
  • the optical splitting module is one of a diffractive optical device optical splitting system, a spatial light modulator optical splitting system, a digital micromirror device optical splitting system, a multi-fiber optical splitting system and a multi-beam splitter optical splitting system.
  • the diffractive optical device light-splitting system includes a diffractive optical device, and the surface of the diffractive optical device has a relief structure;
  • the spatial light modulator light-splitting system includes a spatial light modulator, and the spatial light modulator is composed of a liquid crystal array;
  • the digital micromirror device light-splitting system includes The digital micromirror device is a digital micromirror array.
  • the fluid focusing module adopts a traditional flow cytometer flow chamber or a microfluidic chip.
  • the imaging method combined with microfluidic technology has the advantages of fully enclosed and pollution-free detection.
  • the photodetector is one of a single photomultiplier tube, an avalanche phototube, a multi-channel fluorescence detection system or a spectrum detection system.
  • the spectral detection system can obtain images of cells in various spectral bands, effectively solve the effect of fluorescence aliasing, and increase the amount of detection information.
  • the present invention combines a multi-channel fluorescence detection system and a spectral detection system to realize multi-color flow imaging and spectral flow imaging, so that the present invention has good scalability.
  • a confocal structure is added before the photodetector.
  • the invention adds a confocal structure before the photodetector, which improves the imaging resolution and imaging quality.
  • the technical effect of the present invention is: the flow imaging system based on fractional laser scanning of the present invention uses the laser spots of the lattice to scan the moving cells, and finally uses the photodetector to detect the cells excited by the cells. Fluorescence or scattered light is detected, and the measured signal is used to obtain cell images, which has the beneficial effects of simple implementation, high reliability, high imaging quality, easy integration, and high imaging throughput.
  • Fig. 1 is the schematic structural diagram of the stream imaging system based on dot matrix laser scanning of the present invention
  • Fig. 2 is a schematic diagram of the principle of a lattice spot generated by a diffractive optical device spectroscopic system
  • Fig. 3 is a schematic diagram of the principle of generating lattice spots by the spatial light modulator spectroscopic system
  • Fig. 4 is the schematic diagram of the principle of the lattice spot generated by the digital micromirror device spectroscopic system
  • Fig. 5 is a schematic diagram of the principle of a lattice spot generated by a multi-fiber spectroscopic system
  • Fig. 6 is a schematic diagram of the principle of a lattice spot produced by a multi-beam splitter system
  • Fig. 7 is a schematic diagram of the principle of cell scanning by two-dimensional lattice spot
  • Fig. 8 is a schematic diagram of the principle of one-dimensional lattice light spots scanning cells
  • FIG. 9 is a schematic diagram of a traditional flow cytometer flow chamber and microfluidic-based flow focusing
  • FIG. 10 is a schematic structural diagram of the flow imaging system provided in Embodiment 1 of the present invention.
  • FIG. 11 is a schematic structural diagram of a flow imaging system provided in Embodiment 2 of the present invention.
  • FIG. 12 is a schematic structural diagram of the flow imaging system provided in Embodiment 3 of the present invention.
  • FIG. 13 is a schematic structural diagram of a flow imaging system provided in Embodiment 4 of the present invention.
  • Figure 14 is a schematic diagram of the fluorescence signal and the calculated image of the 10 ⁇ m microsphere
  • Fig. 15 is a schematic diagram of the scattered light signal and the calculated image of the 10 ⁇ m microsphere.
  • Laser light source 1 spectroscopic module 2, illumination objective lens 3, fluid focusing module 4, light collection objective lens 5, photodetector 6, acquisition card 7, computer module 8, laser beam 9, diffractive optical device 10, focusing lens 11, projection plane 12.
  • the flow imaging system based on fractional laser scanning includes:
  • Laser light source 1 the laser light source 1 is used to generate continuous laser light, and the laser light is irradiated to the spectroscopic module 2;
  • the fluid focusing module 4 arranged downstream of the illumination objective lens 3, the fluid focusing module 4 is used to arrange the scattered cells to be imaged in the sample into a single cell axial flow, when the cells pass through the illumination area of the two-dimensional lattice spot 30 sequentially, the cells excite and scatter Light or fluorescence as signal light;
  • the computer module 8 arranged downstream of the acquisition card 7 is used to process the received digital signal to obtain images of cells to be imaged in the fluid focusing module 4 .
  • the laser light source 1 is used to generate continuous laser light, and the laser light is irradiated to the spectroscopic module 2 .
  • the laser light source 1 can be a commercial continuous laser light source. Generally, a laser light source commonly used in flow cytometry is used, and the wavelength is one or more of 355nm, 375nm, 405nm, 488nm, 530nm, 561nm and 640nm.
  • the laser power ranges from a few mW to hundreds of mW, and the beam is a Gaussian beam.
  • the light splitting module 2 is used to divide the received single beam of laser light into multiple beams of laser light that are spatially dispersed or shifted in position.
  • the multiple beams of laser light are transmitted to the illumination objective lens 3, and the illumination objective lens 3 is used to focus the multiple beams of laser light.
  • a two-dimensional lattice light spot 30 is generated on the focal plane of the illumination objective lens 3 .
  • the optical splitting module 2 has five implementation methods, including: a diffractive optical device optical splitting system, a spatial light modulator optical splitting system, a digital micromirror device optical splitting system, a multi-fiber optical splitting system and a multi-mirror optical splitting system. As shown in Figure 2-6.
  • the spectroscopic system of diffractive optics is shown in Figure 2.
  • the laser beam 9 is incident on the diffractive optical device 10.
  • the surface of the diffractive optical device 10 has a micro-processed relief structure, which can modulate the phase of the incident light.
  • the modulated light is divided into multiple laser beams with different angles in space from a single beam. .
  • the split multi-beams pass through the focusing lens 11 and are focused onto the projection plane 12 to form a spot array 13 .
  • the diffractive optical device 10 is equivalent to the spectroscopic module 2;
  • the focusing lens 11 is equivalent to the illumination objective lens 3;
  • the projection plane 12 is equivalent to the plane through which cells flow in the fluid focusing module 4;
  • the spot array 13 is equivalent to the lattice spot 30. Due to the design flexibility of the diffractive device, this method can theoretically generate any number of multi-spots.
  • the spatial light modulator splitting system is shown in Figure 3.
  • the laser beam 9 is incident on the surface of the spatial light modulator 14 .
  • the spatial light modulator 14 is composed of a controllable liquid crystal array, in which each individual pixel can be endowed with a specific phase retardation, so as to obtain the same phase modulation effect as the diffractive optical device 10 .
  • the phase of the beam reflected by the spatial light modulator 14 is modulated, and the single beam is divided into multiple laser beams with different angles in space.
  • the split laser beams pass through the focusing lens 11 and are focused onto the projection plane 12 to form a spot array 13 .
  • the spatial light modulator 14 is equivalent to the light splitting module 2; the corresponding relationship of other parts is the same as that of the parts of the diffraction light splitting system.
  • the reflective system of the digital micromirror device is shown in Figure 4.
  • the laser beam 9 is incident on the surface of the digital micromirror device 15 .
  • a digital micromirror device 15 (DMD) is a digital micromirror array, which is a kind of optical switch, and uses a rotating mirror to realize the opening and closing of the optical switch.
  • the laser beam 9 shoots to the reflector of the DMD.
  • the DMD When the DMD is turned on, the light can enter the other end through a symmetrical optical path; when the DMD is turned off, the reflector of the DMD produces a small rotation, and the light cannot be reflected to the At the other end, the effect of turning off the optical switch is achieved.
  • a grating-like diffractive structure can be formed on the surface, thereby realizing phase modulation of the incident light and separating the incident light into beams with different angles.
  • the split laser beams pass through the focusing lens 11 and are focused onto the projection plane 12 to form a spot array 13 .
  • the digital micromirror device 15 is equivalent to the spectroscopic module 2; the corresponding relationship of other parts is the same as that of the diffractive spectroscopic system.
  • the optical fiber splitting system is shown in Figure 5.
  • the laser beam 9 is incident on the first fiber coupler 16, and the optical fiber adopts a multi-stage beam splitting method to split the laser light step by step.
  • the light splitting of each stage is realized by using a fiber optic beam splitter.
  • the position and angle of the first fiber coupler 16 at the end can be adjusted for the split multi-beam laser beams to form multi-beams with specific distribution.
  • the split multi-beams pass through the focusing lens 11 and are focused onto the projection plane 12 to form a spot array 13 .
  • the first optical fiber coupler 16 and the splitting optical fiber 17 are equivalent to the spectroscopic module 2, and the corresponding relationship of other parts is the same as that of the diffractive spectroscopic system.
  • FIG. 6 shows a schematic diagram of 8-splitting of the laser beam in this way, and this figure uses three-level splitting.
  • Laser beam 9 is incident on the first-order beam splitter of beam splitter 18 (the leftmost beam splitter), half of the laser light is directly projected, and half of the laser light is reflected to the second-order beam splitter; the light projected through the first-order beam splitter It is fully reflected by the reflector 19, and also reflected on the second-stage beam splitter; by analogy, after being split and reflected by the third-stage beam splitter, 8 beams of laser light with a total light intensity of 1/8 will be finally generated.
  • n-level beamsplitters are set, 2 n beams of laser light with energy equal to 1/2 n total incident energy can be generated (energy loss is not considered). A slight angle is maintained between the mirror 19 and the plurality of beam splitters, and multiple beams of spatially different angles will be generated after the beams are reflected back and forth.
  • the split multi-beams pass through the focusing lens 11 and are focused onto the projection plane 12 to form a spot array 13 .
  • the spectroscopic mirror 18 and the reflecting mirror 19 are equivalent to the spectroscopic module 2, and the corresponding relationship of other parts is the same as that of the diffractive spectroscopic system.
  • the two-dimensional lattice spot 30 scans and excites the cells in motion.
  • the distance between the two-dimensional lattice spots 30 parallel to the direction of cell movement is greater than the length of the cell, and the distance between the two-dimensional lattice spots 30 perpendicular to the direction of cell movement
  • the position distribution of the light spots 30 covers different positions of the cells.
  • the principle of motion scanning of the two-dimensional lattice spot 30 is shown in FIG. 7 .
  • the spectroscopic module 2 and the illumination lens will generate a two-dimensional lattice light spot 30 on the plane of cell movement. As shown in Figure 7(a), there are n spots on the plane, namely spot 1, spot 2... spot n, assuming that the coordinates of these spots are (x 1 , y 1 ), (x 2 , y 2 ) ...(x n ,y n ).
  • the position distribution of these spots satisfies the following conditions:
  • the interval ⁇ x of the light spots in the x direction is required to be larger than the size of common human cells, generally 20-30 ⁇ m. This setting ensures that when the cells flow through, only one light spot scans the cells at a time.
  • the position of the light spot in the y direction can be distributed arbitrarily, but it should cover all positions in the y direction where the cells flow through with a certain density, that is, the set of positions in the y direction of the light spot ⁇ y i , 1 ⁇ i ⁇ n ⁇ Constitutes a sample of the range of y-positions the cell flows through. This setting ensures that every position expected to be sampled can be scanned by the light spot when the cells flow through.
  • FIG. 7(c) shows a schematic diagram of recovering the cell image according to the fluorescence intensity signal and the scattered light intensity signal.
  • the signal excited by each light spot is extracted separately, and the signal is rearranged according to the order of the position of the light spot to obtain the image of the cell.
  • the one-dimensional linear array light spot is a special example of the two-dimensional lattice light spot 30 , and the distribution of such light spots is equally spaced in both the x and y directions.
  • the principle of linear array light spot scanning for cell movement is shown in Figure 8.
  • the spectroscopic module 2 and the illumination lens will generate linear light spots on the plane of cell movement. As shown in Figure 8(a), these light spots are located on a straight line with equal distances from each other, and there is a small angle between the direction of cell movement and the direction of the line connecting the light spots.
  • the distances between the linear array light spots in the x-direction and y-direction are ⁇ x and ⁇ y, respectively.
  • ⁇ x is required to be greater than the size of the cell. This setting can ensure that only one light spot scans the cell at a time during the cell movement.
  • ⁇ y is generally set to 20-30 ⁇ m.
  • Each pulse contains the fluorescence intensity distribution and the scattering intensity distribution at the corresponding position of the cell.
  • Figure 8(c) shows a schematic diagram of recovering cell images based on fluorescence intensity signals and scattered light intensity signals, the pulse signals at fixed intervals are extracted respectively and rearranged to obtain cell images.
  • the fluid focusing module 4 is used to arrange the dispersed cells to be imaged in the sample into a single cell axial flow, and when the cells pass through the illumination area of the two-dimensional lattice spot 30 sequentially, the cells excite scattered light or fluorescence as signal light.
  • the fluid focusing module 4 adopts a traditional flow cytometer flow chamber or a microfluidic chip.
  • the first focusing method is to use the focusing module of the flow cell of the traditional flow cytometer.
  • Figure 9(a) shows the cross-section of the fluid circuit diagram. This method adopts a hydrodynamic method for focusing, in which the sheath fluid 34 flows around the sample fluid 35 , and forms a single-cell sample axial flow 37 after passing through the conical constraints of the focusing part.
  • the single-cell sample axial flow 37 corresponds to a row of cells in FIG. 1 (the two vertical lines indicate the boundary of the axial flow).
  • the cells contained in the axial flow pass through the detection area 36 one by one after focusing. In this detection area 36, the cells are illuminated by the lattice spot 30, and the excitation light generated by the cells is collected by the light-receiving objective lens 5 and sent to the photodetector 6 for detection. After being collected by the acquisition card 7, it is transmitted to the computer module 8, and the cell image is restored by the computer module 8.
  • the second focusing method is to use a microfluidic chip for focusing, as shown in Figure 9(b).
  • the microfluidic chip uses microchannels for focusing, the sheath liquid 34 is injected into the chip from two microchannels, and the sample solution 35 is injected from the center at the same time. Under the fluid compression of the two sheath fluids 34 , the sample flow is transformed into a constant single-cell sample axial flow 37 .
  • the single-cell sample axial flow 37 corresponds to a column of cells in FIG. 1 (the two vertical lines indicate the axial flow boundaries).
  • the cells are illuminated by the dot matrix light spots 30 , and after the cells also pass through the detection area 36 , they can be scanned by the light spots to detect the images of the cells.
  • the light collecting objective lens 5 is used to receive the signal light, and the signal light is transmitted to the photodetector 6, and the condenser lens 38 focuses the signal light onto the photosensitive surface of the photodetector 6; the photodetector 6 is used to convert the received signal light It is a voltage signal, and sends the voltage signal to the acquisition card 7; the acquisition card 7 is used to collect the voltage signal, converts it into a corresponding digital signal through analog-to-digital conversion, and sends the digital signal to the computer module 8; the computer module 8 uses The received digital signal is processed to obtain the image of the imaged cell in the fluid focusing module 4 . There are no special parameters and model requirements for the illumination objective lens 3 and the light collection objective lens 5.
  • the illuminating objective lens 3 realizes spot focusing, and the light collecting objective lens 5 realizes the collection of fluorescence and scattered light.
  • the light receiving objective lens 5 may be the same objective lens as the illumination objective lens 3, or may not be the same objective lens.
  • a back-facing fluorescence detection optical path is formed; when the light collecting objective lens 5 and the illuminating objective lens 3 are not the same objective lens, a forward fluorescence detection optical path is formed.
  • the photodetector 6 may use a photomultiplier tube (PMT) or an avalanche photodiode (APD) or other detectors, or may be externally connected with multiple optical paths for fluorescence detection and spectrum detection.
  • PMT photomultiplier tube
  • APD avalanche photodiode
  • the multi-channel fluorescence detection system is a standard configuration of a commercial flow cytometer, which consists of a series of dichroic mirrors and filters with different cut-off wavelengths and multiple different PMTs to form multiple fluorescence detection channels. Multiple fluorescence detection channels are responsible for the detection of different fluorochromes.
  • the spectral detection system is a detection system used in a spectral flow cytometer, and is composed of a spectral spectroscopic device and an array photomultiplier tube (PMT).
  • PMT array photomultiplier tube
  • both the acquisition card 7 and the computer module 8 have no special model requirements, and both can be conventional models.
  • This embodiment takes a system of back-facing fluorescence detection light path and forward scattered light detection light path system as an example, as shown in FIG. 10 .
  • the laser light source 1 emits laser light, which is split into the dichroic mirror 20 through the light splitting module 2 .
  • the cut-off wavelength of the dichroic mirror 20 is longer than the wavelength of the laser light source 1 and shorter than the wavelength of the cell-excited fluorescence.
  • the light beam is reflected to the fluorescent objective lens 21 through the dichroic mirror 20, and the light reflected to the fluorescent objective lens 21 should be a light beam that is spatially shifted by a small angle or a small distance.
  • the fluorescent objective lens 21 serves as the illuminating objective lens 3 and simultaneously serves as the light receiving objective lens 5 .
  • a lattice spot 30 is formed at the center of the fluid focusing module 4 .
  • Cells are labeled with fluorescent dyes or fluorescent proteins, and are excited by the laser spot to produce fluorescence.
  • a series of fluorescent signals are excited. Since the fluorescence wavelength is longer than the cut-off wavelength of the dichroic mirror 20, the fluorescence entering the fluorescence objective lens 21 passes through the dichroic mirror 20 and enters the subsequent optical path. Fluorescence is collected onto a fluorescence detector 25 through a condenser lens 22 .
  • the fluorescence detector 25 can be detected by a photomultiplier tube (PMT).
  • a fluorescence filter 27 is placed in front of the fluorescence detector 25.
  • the fluorescence filter 27 allows only the fluorescence signals within the emission wavelength range of the cells to enter the fluorescence detector 25.
  • the signal detected by the fluorescence detector 25 can be used to restore the fluorescence intensity at different positions of the cell.
  • the right part of the optical path of the fluid focusing module 4 is an optical path for detecting scattered light.
  • a light-shielding rod 24 is placed in front of the scattered light objective lens 23, and the light-shielding rod 24 shields most of the original laser light.
  • the scattered light will be collected by the scattered light objective lens 23 and enter the scattered light detector 26 through the condenser lens 38 .
  • the signal detected by the scattered light detector 26 can be used to recover the scattering intensity of each position of the cell.
  • the imaged fluorescence image reflects the distribution of intracellular fluorescent dyes or fluorescent molecules
  • the scattering intensity image reflects the intensity of intracellular scattering media, such as organelles, cell nuclei and other scattering media.
  • This embodiment is a system of back-facing fluorescence detection optical path and forward scattered light detection optical path, as shown in FIG. 11 .
  • the structure on the right side of the fluid focusing module 4 is the same as that in Embodiment 1.
  • the left side of the fluid focusing module 4 is the spot illumination and the back-facing fluorescence detection structure.
  • the laser light source 1 emits laser light, which is split into the dichroic mirror 20 through the light splitting module 2 .
  • the cut-off wavelength of the dichroic mirror 20 is longer than the wavelength of the laser light source 1 and shorter than the wavelength of the cell-excited fluorescence.
  • the light beam is reflected to the fluorescent objective lens 21 through the dichroic mirror 20, and the light reflected to the fluorescent objective lens 21 should be a light beam that is spatially shifted by a small angle or a small distance.
  • the fluorescence objective lens 21 corresponds to the illumination objective lens 3 in FIG. 1 , and is also equivalent to the light collection objective lens 5 in FIG.
  • a lattice spot 30 is formed at the center of the fluid focusing module 4 .
  • Cells are labeled with fluorescent dyes or fluorescent proteins, which can be excited by the laser spot to produce fluorescence.
  • a series of fluorescent signals are excited. Since the fluorescence wavelength is longer than the cut-off wavelength of the dichroic mirror 20 , the fluorescence entering the fluorescence objective lens 21 passes through the dichroic mirror 20 and enters the subsequent optical path.
  • the fluorescence signal can enter the optical fiber 29 after passing through the condenser lens 22 and the second fiber coupler 28, and then transmit to the photodetector 6 through the optical fiber 29.
  • the photodetector 6 is a multi-channel fluorescence detection or spectrum detection system.
  • This embodiment is a system of forward fluorescence detection optical path and forward scattered light detection optical path, which is suitable for the optical path of multicolor laser excitation, as shown in FIG. 12 . Since the fluorescent signal of a short-wavelength laser overlaps with the band of another long-wavelength laser when using a back-facing detection structure, the forward-facing fluorescence detection structure is more suitable for the optical path excited by multicolor lasers.
  • the left side of the fluid focusing module 4 is the illumination light path.
  • the laser light source 1 generates laser light with single or multiple continuous wavelengths, and after passing through the spectroscopic module 2 and the illumination objective lens 3 , a dot matrix spot 30 is formed in the center of the fluid focusing module 4 .
  • the cells in the fluid focusing module 4 flow through the lattice spot 30, they excite fluorescent signals and scattered light signals.
  • the illumination laser is shielded by the light-shielding rod 24, it will not enter the subsequent detection system, but the fluorescence signal and scattered light signal will be collected by the scattered light and fluorescence objective lens 31, which corresponds to the light-receiving objective lens 5 in FIG. 1 .
  • Scattered light and fluorescence signals enter the dichroic mirror 20, wherein the cut-off wavelength of the dichroic mirror 20 is higher than the shortest wavelength of the laser, and shorter than other wavelengths of laser light and fluorescence.
  • the scattered light is reflected by the dichroic mirror 20 to the scattered light detector 26 for scattering imaging of cells.
  • the light passing through the dichroic mirror 20 includes other long-wavelength lasers and fluorescence of various wavelengths excited by the lasers. This part of light enters the optical fiber 29 through the condenser lens 22 and the second fiber coupler 28, wherein, other long-wavelength lasers are unnecessary optical signals, which are filtered by the notch filter 32, and then passed through the notch filter.
  • the light after 32 contains only the desired fluorescent signal.
  • This part of the fluorescence is transmitted to the photodetector 6 through the optical fiber 29, and the photodetector 6 is a multi-channel fluorescence detection or spectrum detection system.
  • This embodiment is suitable for multi-laser excitation imaging, has more imageable wavelength bands, and can provide more imaging information.
  • a confocal structure is added to the imaging system to improve the resolution of the system, as shown in FIG. 13 .
  • This embodiment adds a small hole diaphragm 33 behind the fluorescence filter 27 , and the fluorescence enters the fluorescence detector 25 after passing through the small hole diaphragm 33 .
  • the pinhole diaphragm 33 and the dot matrix spot 30 in the optical path form a confocal structure.
  • the fluorescence excited by the dot matrix spot 30 is collected by the fluorescence objective lens 21 , passes through the dichroic mirror 20 , and then passes through the condenser lens 22 to focus on the plane where the pinhole diaphragm 33 is located.
  • the fluorescence excited by the lattice spot 30 will also form a fluorescent spot on the plane where the pinhole stop 33 is located.
  • the pinhole diaphragm 33 is set to just allow the light at the conjugate image position of the lattice spot 30 to pass through, while shielding the light at other positions, thus forming a confocal structure.
  • the advantage of the confocal structure is that only the fluorescence excited at the position of the light spot of the lattice spot 30 passes through the small hole, while the fluorescence excited at other positions on the focal plane or at positions outside the focal plane is blocked, thereby improving System contrast, resolution and image quality.
  • This embodiment is suitable for improving the imaging resolution and imaging quality, and the imaging system after the pinhole diaphragm 33 can realize the imaging effect of the confocal microscope under the same conditions.
  • Fig. 14 and Fig. 15 are examples of actual system scanning imaging, using the system shown in Embodiment 1, wherein the spectroscopic module 2 adopts a diffractive optical device spectroscopic system.
  • Fig. 14 is the signal and image of fluorescence imaging of 10 ⁇ m microspheres
  • Fig. 15 is the original signal and image of scattered light of 10 ⁇ m microspheres.
  • the 10 ⁇ m microspheres flow in the focusing chip at a speed of 4.8m/s.
  • the diffractive optical device 10 After the laser beam is split by the diffractive optical device 10, it is focused to the center of the flow channel by the objective lens. Fluorescence excited in the center of the flow channel is detected by the back detection optical path, and scattered light is detected by the forward detection optical path. Both back fluorescence and forward scattered light are detected by a photomultiplier tube (PMT), and the signal is sampled by an acquisition card 7 with a maximum sampling rate of 200MHz.
  • PMT photomultiplier tube
  • first and second are used for descriptive purposes only, and cannot be interpreted as indicating or implying relative importance or implicitly specifying the quantity of indicated technical features.
  • the features defined as “first” and “second” may explicitly or implicitly include at least one of these features.
  • “plurality” means at least two, such as two, three, etc., unless otherwise specifically defined.

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Abstract

一种基于点阵激光扫描的流式成像系统，包括激光光源（1）、分光模块（2）、照明物镜（3）、流体聚焦模块（4）、收光物镜（5）、光电探测器（6）、采集卡（7）和计算机模块（8）。分光模块（2）对激光光源（1）产生的连续激光进行调制，并由照明物镜（3）聚焦到焦平面上，形成点阵光斑（30），点阵光斑（30）照射到流体聚焦模块（4）的单细胞轴流（37）上；当单细胞轴流（37）中的细胞经过点阵光斑（30）时，产生荧光信号和散射光信号被收光物镜（5）接收，然后经聚光镜（38）送入到光电探测器（6）；光电探测器（6）将激发的荧光和散射光信号转化为电压信号，采集卡（7）采集该信号后转换为数字信号并发送给计算机模块（8），经由计算机模块（8）恢复得到细胞图像。这一成像系统实现简单、可靠性高、成像质量高、易于集成且成像通量高的优势。

Description

一种基于点阵激光扫描的流式成像系统
本申请要求2021年9月7日提交的中国优先权专利申请202111044232.3的权益，将其全部内容通过引用的方式合并在此。
技术领域
本发明属于流式成像技术领域，尤其涉及一种基于点阵激光扫描的流式成像系统。
背景技术
流式成像是对高速流动的细胞中的每一个细胞进行成像，获取显微图像的一种成像技术。成像流式细胞仪就是进行流式成像的一种仪器，它解决了传统流式细胞术无法获取细胞图像的缺陷，能获取更丰富的细胞形态学信息，可广泛用于科研和临床检测。根据探测器的不同，成像流式技术可分为基于阵列式探测器和基于单点式探测器两类，每一类成像流式技术又有多种不同的实现方式。其中基于阵列式探测器的成像方式，采用CCD或者CMOS作为探测器件，最具代表性的是Amnis公司的ImageStream系列仪器，核心器件为时间延迟积分CCD。但在高通量成像的应用下，细胞高速运动(>1m/s)，阵列式的探测器存在着帧率不足、信噪比较低和数据流量大的问题，难以满足成像需求。针对以上不足，基于单点式探测器的成像方式是目前适用于高通量成像流式的一类方式。单点式探测器方式中重要的一种成像方式是基于光谱标记法和时间拉伸技术的成像方法，光谱标记部分使用色散器件将宽谱光的不同波长映射到被成成像样本不同位置，样本反射光或散射光通过色散光纤进行“时间拉伸”，从而将携带了样本图像信息的光信号在时域上分开，然后使用单个光电倍增管(PMT)进行检测，最后恢复重建图像。这种成像方式由于采用时间拉伸的技术而受限于有限的色散系数，因此需要使用较高速率的采集卡完成数据采集；且由于时间拉伸技术需要采用长达km级的色散光纤，因此系统复杂度和稳定性都较低。基于单点式探测器方式中的另一种重要成像方式是随机结构光激发的压缩感知成像方式。该方法使用静态的随机照明激光光斑照射细胞，当细胞流经照明区域的时候，激发的荧光信号中包含了细胞的荧光分布信息，然后可以使用压缩感知算法求解逆问题获得细胞图像，或使用机器学习方法直接对原始信号进行识别处理。这种成像方法虽然可以获得较高的成像通量，但由于无法直接获取细胞图像，需要耗费大量时间进行逆问题求解，因此实用性有限。
发明内容
针对上述现有技术中流式成像方式存在的耗时长、可靠性低实用性有限的技术问题，本发明的目的在于提供一种基于点阵激光扫描的流式成像系统。
为达到上述目的，本发明提出了一种基于点阵激光扫描的流式成像系统，包括：
激光光源，激光光源用于产生连续激光，激光照射到分光模块；
设置在激光光源下游的分光模块，分光模块用于将接收到的单束激光分为空间上角度分散或位置有所偏移的多束激光，多束激光传送到照明物镜；
设置在分光模块下游的照明物镜，照明物镜用于对多束激光进行聚焦，在照明物镜的焦平面上产生二维点阵光斑；
设置在照明物镜下游的流体聚焦模块，流体聚焦模块用于将样本中分散的待成像细胞排列成单细胞轴流，在依次经过二维点阵光斑的照明区域时，细胞激发散射光或荧光作为信号光；
设置在流体聚焦模块下游的收光物镜，收光物镜用于接收信号光，并将信号光通过聚光镜聚焦后传送到光电探测器；
设置在收光物镜下游的光电探测器，光电探测器用于将接收到的信号光转化为电压信号，并将电压信号发送给采集卡；
设置在光电探测器下游的采集卡，采集卡用于对电压信号进行采集，经过模数转换转化为对应的数字信号，并将数字信号发送给计算机模块；
设置在采集卡下游的计算机模块，计算机模块用于对接收到的数字信号进行处理，得到流体聚焦模块中被成像细胞的图像。
本发明将传统激光扫描显微镜中的光束扫描系统改为激光分光模块，将运动扫描光束变为静止的点阵扫描光束，实现方法简单；另外，流体聚焦系统中，细胞流动速度在m/s量级，可以实现10000cell/s以上的成像通量，使得成像速度快。
进一步地，二维点阵光斑对运动中的细胞进行扫描激发，在平行于细胞运动方向上的二维点阵光斑之间的距离大于细胞的长度，在垂直于细胞运动方向上的二维点阵光斑的位置分布覆盖细胞的不同位置。
本发明的二维点阵光斑由特定排列的激光光斑构成，光斑被投射到成像物平面上，细胞在被成像物平面内匀速运动。在平行于细胞运动方向上的二维点阵光斑之间的距离大于细胞的长度，在垂直于细胞运动方向上的二维点阵光斑的位置分布覆盖细胞的不同位置，使得当细胞流经光斑区域时保证单次只有一个光斑对细胞进行照明，实现利用细胞的运动对细胞完 成逐行激光扫描的效果。
进一步地，收光物镜和照明物镜为同一个物镜时，形成背向荧光检测光路；收光物镜和照明物镜不是同一个物镜时，形成前向荧光检测光路。
进一步地，分光模块为衍射光学器件分光系统、空间光调制器分光系统、数字微镜器件分光系统、多光纤分光系统和多分光镜分光系统中的一种。
进一步地，衍射光学器件分光系统包括衍射光学器件，衍射光学器件表面具有浮雕结构；空间光调制器分光系统包括空间光调制器，空间光调制器由液晶面阵构成；数字微镜器件分光系统包括数字微镜器件，数字微镜器件为数字微镜阵列。
进一步地，流体聚焦模块采用传统流式细胞仪流动室或微流控芯片。其中，结合微流控技术的成像方法具有全封闭无污染检测的优点。
进一步地，光电探测器为单个光电倍增管、雪崩光电管、多路荧光检测系统或光谱检测系统中的一种。其中光谱检测系统可以获得细胞在各个光谱波段范围内的图像，有效地解决荧光混叠的效果，提高检测的信息量。
本发明结合多路荧光检测系统和光谱检测系统，实现多色流式成像和光谱流式成像，使得本发明可扩展性好。
进一步地，在光电探测器之前加设共聚焦结构。
本发明在光电探测器之前加设共聚焦结构，提高了成像的分辨率和成像质量。
相对于现有技术，本发明的技术效果为：本发明涉及的基于点阵激光扫描的流式成像系统利用点阵的激光光斑对运动中的细胞进行扫描，最后使用光电探测器对细胞激发的荧光或散射光进行探测，利用测得的信号得到细胞图像，具有实现简单、可靠性高、成像质量高、易于集成且成像通量高的有益效果。
附图说明
本发明上述的和/或附加的方面和优点从下面结合附图对实施例的描述中将变得明显和容易理解，其中：
图1为本发明基于点阵激光扫描的流式成像系统的结构示意图；
图2为衍射光学器件分光系统产生点阵光斑的原理示意图；
图3为空间光调制器分光系统产生点阵光斑的原理示意图；
图4为数字微镜器件分光系统产生点阵光斑的原理示意图；
图5为多光纤分光系统产生点阵光斑的原理示意图；
图6为多分光镜分光系统产生点阵光斑的原理示意图；
图7为二维点阵光斑对细胞扫描的原理示意图；
图8为一维点阵光斑对细胞扫描的原理示意图；
图9为传统流式细胞仪流动室和基于微流体的流式聚焦的示意图；
图10为本发明实施例1的提供的流式成像系统的结构示意图；
图11为本发明实施例2的提供的流式成像系统的结构示意图；
图12为本发明实施例3的提供的流式成像系统的结构示意图；
图13为本发明实施例4的提供的流式成像系统的结构示意图；
图14为10μm微球的荧光信号和解算出的图像示意图；
图15为10μm微球的散射光信号和解算出的图像示意图。
附图标记：
激光光源1、分光模块2、照明物镜3、流体聚焦模块4、收光物镜5、光电探测器6、采集卡7、计算机模块8、激光光束9、衍射光学器件10、聚焦透镜11、投射平面12、光斑阵列13、空间光调制器14、数字微镜器件15、第一光纤耦合器16、分束光纤17、分光镜18、反射镜19、二向色镜20、荧光物镜21、聚光透镜22、散射光物镜23、遮光棒24、荧光探测器25、散射光探测器26、荧光滤光片27、第二光纤耦合器28、光纤29、点阵光斑30、散射光及荧光物镜31、陷波滤光片32、小孔光阑33、鞘液34、样本液35、检测区36、单细胞样本轴流37、聚光镜38。
具体实施方式
下面详细描述本发明的实施例，所述实施例的示例在附图中示出，其中自始至终相同或类似的标号表示相同或类似的元件或具有相同或类似功能的元件。下面通过参考附图描述的实施例是示例性的，旨在用于解释本发明，而不能理解为对本发明的限制。
下面参照附图描述根据本发明实施例提出的基于点阵激光扫描的流式成像系统。
如图1所示，该基于点阵激光扫描的流式成像系统包括：
激光光源1，激光光源1用于产生连续激光，激光照射到分光模块2；
设置在激光光源1下游的分光模块2，分光模块2用于将接收到的单束激光分为空间上角度分散或位置有所偏移的多束激光，多束激光传送到照明物镜3；
设置在分光模块2下游的照明物镜3，照明物镜3用于对多束激光进行聚焦，在照明物镜3的焦平面上产生二维点阵光斑30；
设置在照明物镜3下游的流体聚焦模块4，流体聚焦模块4用于将样本中分散的待成像细胞排列成单细胞轴流，在依次经过二维点阵光斑30的照明区域时，细胞激发散射光或荧 光作为信号光；
设置在流体聚焦模块4下游的收光物镜5，收光物镜5用于接收信号光，并将信号光传送到光电探测器6；
设置在收光物镜5下游的光电探测器6，光电探测器6用于将接收到的信号光通过聚光镜38聚焦后转化为电压信号，并将电压信号发送给采集卡7；
设置在光电探测器6下游的采集卡7，采集卡7用于对电压信号进行采集，经过模数转换转化为对应的数字信号，并将数字信号发送给计算机模块8；
设置在采集卡7下游的计算机模块8，计算机模块8用于对接收到的数字信号进行处理，得到流体聚焦模块4中被成像细胞的图像。
激光光源1用于产生连续激光，激光照射到分光模块2。激光光源1可采用商用连续激光光源。一般采用流式细胞仪常用的激光光源，波长为355nm、375nm、405nm、488nm、530nm、561nm和640nm中的一种或多种。激光功率在几mW至上百mW范围内，光束为高斯光束。
分光模块2用于将接收到的单束激光分为空间上角度分散或位置有所偏移的多束激光，多束激光传送到照明物镜3，照明物镜3用于对多束激光进行聚焦，在照明物镜3的焦平面上产生二维点阵光斑30。分光模块2有五种实现方式，包括：衍射光学器件分光系统、空间光调制器分光系统、数字微镜器件分光系统、多光纤分光系统和多分光镜分光系统。如附图2-图6所示。
衍射光学器件分光系统如图2所示。激光光束9入射到衍射光学器件10上，衍射光学器件10表面具有微加工的浮雕结构，可以对入射光的相位进行调制，调制过后的光由单束光分为空间上不同角度的多束激光。分束后的多光束经过聚焦透镜11，聚焦到投射平面12上，形成光斑阵列13。其中，衍射光学器件10相当于分光模块2；聚焦透镜11相当于照明物镜3；投射平面12相当于流体聚焦模块4中细胞流动经过的平面；光斑阵列13相当于点阵光斑30。由于衍射器件具有设计上的灵活性，这种方法理论上可以产生数量任意的多光斑。
空间光调制器分光系统如图3所示。激光光束9入射到空间光调制器14的表面上。空间光调制器14由可控液晶面阵构成，其中每个单独的像素都可以被赋予特定的相位延迟，从而获得与衍射光学器件10相同的相位调制效果。经过空间光调制器14反射后的光束相位被调制，由单束光分为空间上不同角度的多数激光。分束后的多束激光经过聚焦透镜11，聚焦到投射平面12上，形成光斑阵列13。其中，空间光调制器14相当于分光模块2；其余各部分的对应关系与衍射分光系统各部分对应关系相同。
数字微镜器件反光系统如图4所示。激光光束9入射到数字微镜器件15的表面上。数字微镜器件15(DMD)是一种数字微镜阵列，是光开关的一种，利用旋转反射镜实现光开关的开合。激光光束9射向DMD的反射镜片，DMD打开的时候，光可经过对称光路进入到另一端；当DMD关闭的时候，即DMD的反射镜产生一个小的旋转，光经过反射后，无法反射到另一端，也就达到了光开关关闭的效果。通过控制DMD阵列的开关模式，可以在表面形成类似光栅的衍射结构，从而实现对入射光的相位调制，使入射光分开为角度上不同的光束。分束后的多束激光经过聚焦透镜11，聚焦到投射平面12上，形成光斑阵列13。其中，数字微镜器件15相当于的分光模块2；其余各部分的对应关系与衍射分光系统各部分对应关系相同。
光纤分光系统如图5所示。激光光束9入射到第一光纤耦合器16上，光纤采用多级分束的方式，对激光进行逐级分光。其中每一级的分光使用光纤分束器实现。分束光纤17第一级分光后，产生能量为1/2总能量的2束激光；第二级分光后，产生能量为1/4的4束激光；以此类推，当进行n级分光之后，将产生能量为1/2 n总入射能量的2 n束激光(不考虑能量损耗)。分束后的多束激光可以调节末端第一光纤耦合器16的位置和角度，形成特定分布的多光束。分束后的多光束经过聚焦透镜11，聚光到投射平面12上，形成光斑阵列13。其中，第一光纤耦合器16和分束光纤17相当于分光模块2，其余各部分的对应关系与衍射分光系统各部分对应关系相同。
多分光镜分光系统如图6所示。图6给出了这种方式对激光束进行8分光的示意图，这张图采用了三级分光。激光光束9入射到分光镜18的第一级分光镜上(最左侧分光镜)，一半的激光直接投射，一半的激光被反射到第二级分光镜；投射过第一级分光镜的光被反射镜19全部反射，同样反射到第二级分光镜上；以此类推，当经过三级分光镜分光和反射后，最终将产生8束光强总光强1/8的激光。如果设置n级分光镜，则可以产生2 n束能量为1/2 n总入射能量的激光(不考虑能量损耗)。其中反光镜19与多个分光镜之间保持一个微小的角度，当光束被来回反射之后将会产生多束空间上不同角度的光束。分束后的多光束经过聚焦透镜11，聚光到投射平面12上，形成光斑阵列13。其中，分光镜18和反射镜19相当于分光模块2，其余各部分的对应关系与衍射分光系统各部分对应关系相同。
二维点阵光斑30对运动中的细胞进行扫描激发，在平行于细胞运动方向上的二维点阵光斑30之间的距离大于细胞的长度，在垂直于细胞运动方向上的二维点阵光斑30的位置分布覆盖细胞的不同位置。二维点阵光斑30运动扫描的原理如图7所示。分光模块2与照明透镜将在细胞运动的平面上产生二维点阵光斑30。如图7(a)所示，平面上共有n 个光斑，分别为光斑1、光斑2……光斑n，假定这些光斑的坐标分别为(x 1,y 1),(x 2,y 2)…(x n,y n)。这些光斑的位置分布满足如下条件：
(1)在细胞流动方向上，即x方向上间隔均匀，均为Δx：Δx＝x i-x i-1,(1<i≤n)。
(2)光斑在x方向上的间隔Δx要求大于常见人体细胞的大小，一般取20～30μm。这样设置保证细胞流经时，每次只有一个光斑扫描细胞。
(3)光斑在y方向上的位置可以任意分布，但应该以一定的密度覆盖细胞流经的y方向的所有位置，也即光斑y方向所在位置的集合{y i,1≤i≤n}构成对细胞流经的y位置范围的一个采样。这样设置保证细胞流经时，每个期望采样的位置都能被光斑扫描到。
当细胞流动时，分别经过光斑1、光斑2……光斑n。光斑所激发的荧光或者散射光信号如图7(b)所示，由于光斑在x方向上的间隔距离大于细胞的大小，因此该信号在时域上表现为分立的脉冲。信号的第一个脉冲为光斑1所激发，表示细胞在位置y 1所在直线上的荧光和散射强度分布；以此类推，信号中第i个脉冲为光斑i激发，表示细胞在位置y i所在直线上的荧光和散射强度分布。图7(c)显示了根据荧光强度信号和散射光强度信号恢复出细胞图像的示意图。将每个光斑激发的信号分别提取出来，按照光斑所在的位置顺序将信号重新排列即可获得细胞的图像。
一维线阵光斑作为二维点阵光斑30的一个特殊样例，这种光斑分布在x和y方向上都是等间隔的分布。线阵光斑对细胞运动扫描的原理如图8所示。分光模块2与照明透镜将在细胞运动的平面上产生线阵光斑。如图8(a)所示，这些光斑位于一条直线上且相互之间距离相等，细胞运动的方向与光斑连线的方向之间呈一个小角度。以细胞运动的方向为x轴，垂直于细胞运动的方向为y轴，那么线阵光斑之间在x方向和y方向上的距离分别为Δx和Δy。其中Δx要求大于细胞的大小，这样设置可以保证细胞运动过程中，每次仅有一个光斑对细胞进行扫描。根据常见人体细胞的大小，Δy一般设置为20～30μm。当细胞依次经过光斑1、光斑2……光斑n时，分别激发出荧光信号和散射光信号，这些光信号在时间上表现为一系列脉冲，如图8(b)所示。其中每个脉冲包含了位于细胞对应位置处的荧光强度分布和散射强度分布。图8(c)显示了根据荧光强度信号和散射光强度信号恢复出细胞图像的示意图，将固定间隔的脉冲信号分别提取出来，重新排列即可获得细胞的图像。
流体聚焦模块4用于将样本中分散的待成像细胞排列成单细胞轴流，在依次经过二维点阵光斑30的照明区域时，细胞激发散射光或荧光作为信号光。流体聚焦模块4采用传统流式细胞仪流动室或微流控芯片。第一种聚焦方式为使用传统流式细胞仪流动室的聚焦模块，图9(a)给出了液路图的横截面。这种方式采用流体动力学的方式进行聚焦，其中，鞘液34 包裹样本液35流动，当经过聚焦部分的锥形约束之后，形成单细胞样本轴流37。单细胞样本轴流37对应于图1中一列细胞(两条竖线表示轴流边界)。轴流中含有细胞，聚焦后逐个经过检测区36，在这个检测区36内，细胞被点阵光斑30照明，细胞产生的激发光由收光物镜5收集并送至光电探测器6检测，并经过采集卡7采集后传输至计算机模块8，由计算机模块8恢复出细胞图像。第二种聚焦方式为使用微流控芯片进行聚焦，如图9(b)所示。微流控芯片采用微流道进行聚焦，鞘液34从两路微流道中注入芯片，样本液35同时从中心注入。在两路鞘液34的流体压缩下，实现将样本流转化为恒定的单细胞样本轴流37。单细胞样本轴流37对应于图1中的一列细胞(两条竖线表示轴流边界)。在检测区36内，细胞被点阵光斑30照明，细胞同样经过检测区36之后，即可被光斑扫描而检测到细胞的图像。收光物镜5用于接收信号光，并将信号光传送到光电探测器6，聚光镜38将信号光聚焦到光电探测器6的感光面上；光电探测器6用于将接收到的信号光转化为电压信号，并将电压信号发送给采集卡7；采集卡7用于对电压信号进行采集，经过模数转换转化为对应的数字信号，并将数字信号发送给计算机模块8；计算机模块8用于对接收到的数字信号进行处理，得到流体聚焦模块4中被成像细胞的图像。照明物镜3和收光物镜5无特殊参数和型号要求。照明物镜3实现光斑聚焦，收光物镜5实现对荧光和散射光的收集。其中收光物镜5可以和照明物镜3为同一个物镜，也可以不是同一个物镜。收光物镜5和照明物镜3为同一个物镜时，形成背向荧光检测光路；收光物镜5和照明物镜3不是同一个物镜时，形成前向荧光检测光路。光电探测器6可以使用光电倍增管(PMT)或者雪崩光电管(APD)等探测器，也可以外接多路荧光检测和光谱检测光路。其中，多路荧光检测系统为商用流式细胞仪的标准配置，由一系列不同截止波长的二向色镜和滤光片以及多个不同的PMT构成多个荧光检测通道。多个荧光检测通道负责检测不同的荧光染料的。光谱检测系统为光谱流式细胞仪所用检测系统，由光谱分光装置以及阵列光电倍增管(PMT)构成。另外，采集卡7和计算机模块8均无特殊型号要求，均采用常规型号即可。
实施例1
本实施例以一种背向荧光检测光路、前向散射光检测光路系统为例，如图10所示。其中激光光源1发射出激光，经过分光模块2分光进入二向色镜20。二向色镜20的截止波长长于激光光源1的波长，短于细胞激发荧光的波长。光束经过二向色镜20反射到荧光物镜21，反射到荧光物镜21的光应该为在空间上以小角度或微小距离偏移的光束。荧光物镜21作为照明物镜3，同时作为收光物镜5。光束经过荧光物镜21聚焦之后，在流体聚焦模块4的中心形成点阵光斑30。细胞经过荧光染料或荧光蛋白标记，被激光光斑激发产生荧光。当细胞经过点阵光斑30之后，激发出一系列荧光信号。由于 荧光波长比二向色镜20的截止波长要长，因此进入荧光物镜21的荧光透过二向色镜20进入后续光路。荧光经过聚光透镜22汇聚到荧光探测器25上。荧光探测器25可以使用光电倍增管(PMT)进行探测，同时荧光探测器25前放置荧光滤光片27，荧光滤光片27使得仅让细胞发射波长范围内的荧光信号进入荧光探测器25。利用荧光探测器25检测到的信号即可恢复细胞不同位置的荧光强度。
流体聚焦模块4右侧部分光路为对散射光探测的光路。散射光物镜23前放置遮光棒24，遮光棒24将大部分原有激光遮蔽。当细胞流经点阵光斑30时，散射光会由散射光物镜23收集经聚光镜38进入散射光探测器26。利用散射光探测器26检测到的信号即可以恢复出细胞各个位置的散射强度。
本实施例中，成像的荧光图像反映了细胞内荧光染料或荧光分子的分布，散射强度图像反映了细胞内散射介质，如细胞器、细胞核等散射介质的强度。
实施例2
本实施例为背向荧光检测光路、前向散射光检测光路系统，如图11所示。其中流体聚焦模块4右侧的结构与实施例1一致。
流体聚焦模块4左侧为光斑照明和背向荧光检测结构。其中激光光源1发射出激光，经过分光模块2分光进入二向色镜20。二向色镜20的截止波长长于激光光源1的波长，短于细胞激发荧光的波长。光束经过二向色镜20反射到荧光物镜21，反射到荧光物镜21的光应该为在空间上以小角度或微小距离偏移的光束。荧光物镜21对应于图1中的照明物镜3，同时也相当于图1中的收光物镜5，在这里同时起到了照明和收集荧光的功能。光束经过荧光物镜21聚焦之后，在流体聚焦模块4的中心形成点阵光斑30。细胞经过荧光染料或荧光蛋白标记，可以被激光光斑激发产生荧光。当细胞经过点阵光斑30之后，激发出一系列荧光信号。由于荧光波长比二向色镜20的截止波长要长，因此进入荧光物镜21的荧光透过二向色镜20进入后续光路。这里的荧光信号经过聚光透镜22和第二光纤耦合器28后可进入光纤29，经过光纤29传输到光电检测器6，光电检测器6为多路荧光检测或光谱检测系统。
实施例3
本实施例为前向荧光检测光路、前向散射光检测光路系统，适用于多色激光激发的光路，如图12所示。由于使用背向检测结构时，会出现某个短波长激光的荧光信号和另一个长波长激光的波段重叠的问题，因此前向荧光检测结构更适用于有多色激光激发的光路。
流体聚焦模块4左侧为照明光路。其中激光光源1产生单个或多个连续波长的激光，经过分光模块2和照明物镜3之后，在流体聚焦模块4中心形成点阵光斑30。流体聚焦 模块4中的细胞流经点阵光斑30时，激发出荧光信号和散射光信号。照明激光经过遮光棒24遮蔽之后，不会进入后续检测系统，而荧光信号和散射光信号会被散射光及荧光物镜31收集，散射光及荧光物镜31对应于图1中的收光物镜5。散射光及荧光信号进入到二向色镜20，其中二向色镜20的截止波长高于激光器的最短波长，短于其它波长激光及荧光波长。散射光被二向色镜20反射到散射光探测器26上，用于细胞的散射成像。透过二向色镜20的光包含有其它长波长的激光，以及激光激发出的各个波段的荧光。这部分光经过聚光透镜22和第二光纤耦合器28进入光纤29，其中，其它长波长的激光是不需要的光信号，通过陷波滤光片32予以滤除，经过陷波滤光片32之后的光仅包含需要的荧光信号。这部分荧光经过光纤29传输到光电检测器6，光电检测器6为多路荧光检测或光谱检测系统。本实施例适用于多激光激发成像，具有更多的可成像波段，可提供更多的成像信息。
实施例4
本实施例在成像系统中增加共聚焦结构，以提升系统分辨率，如图13所示。本实施例与实施例1的不同是：本实施例在荧光滤光片27之后增加了小孔光阑33，荧光经过小孔光阑33后再进入荧光探测器25。小孔光阑33和光路中点阵光斑30形成共聚焦结构。具体为：点阵光斑30激发的荧光经过荧光物镜21收集之后，透过二向色镜20，再经过聚光透镜22之后，聚焦到小孔光阑33所在的平面上。由于点阵光斑30所在平面和小孔光阑33所在平面为共轭平面，因此点阵光斑30激发的荧光也会在小孔光阑33所在平面上形成荧光光斑。小孔光阑33设置为刚好使得点阵光斑30共轭像位置处透光，而遮蔽其它位置的光，这样形成共聚焦的结构。共聚焦结构的优点是，仅仅让位于点阵光斑30的光斑所在位置处激发的荧光通过小孔，而焦平面上其它位置或者位于焦平面之外的位置激发的荧光则被遮蔽，从而提高系统对比度、分辨率与成像质量。
本实施例适用于提升成像分辨率和成像质量，经过小孔光阑33之后的成像系统，可以实现同等条件下共聚焦显微镜的成像效果。
图14和图15为实际系统扫描成像的样例，采用实施例1所示的系统，其中分光模块2采用衍射光学器件分光系统。图14为对10μm微球进行荧光成像的信号和图像，图15为对10μm微球的散射光原始信号和图像。10μm微球以4.8m/s的速度在聚焦芯片中流动，激光经过衍射光学器件10分光之后，由物镜聚焦到流道中心。流道中心激发的荧光被背向检测光路探测到，散射光使用前向检测光路进行检测。背向荧光和前向散射光均由光电倍增管(PMT)进行探测，信号由一个最大采样率200MHz的采集卡7进行采样。
在本说明书的描述中，参考术语“一个实施例”、“一些实施例”、“示例”、“具体示例”、或“一些示例”等的描述意指结合该实施例或示例描述的具体特征、结构、材料或者特点包含于本发明的至少一个实施例或示例中。在本说明书中，对上述术语的示意性表述不必须针对的是相同的实施例或示例。而且，描述的具体特征、结构、材料或者特点可以在任一个或多个实施例或示例中以合适的方式结合。此外，在不相互矛盾的情况下，本领域的技术人员可以将本说明书中描述的不同实施例或示例以及不同实施例或示例的特征进行结合和组合。
此外，术语“第一”、“第二”仅用于描述目的，而不能理解为指示或暗示相对重要性或者隐含指明所指示的技术特征的数量。由此，限定有“第一”、“第二”的特征可以明示或者隐含地包括至少一个该特征。在本发明的描述中，“多个”的含义是至少两个，例如两个，三个等，除非另有明确具体的限定。
尽管已经示出和描述了本发明的实施例，本领域的普通技术人员可以理解：在不脱离本发明的原理和宗旨的情况下可以对这些实施例进行多种变化、修改、替换和变型，本发明的范围由权利要求及其等同物限定。

Claims (8)

  1. 一种基于点阵激光扫描的流式成像系统，其特征在于，包括：
    激光光源，所述激光光源用于产生连续激光，所述激光照射到分光模块；
    设置在所述激光光源下游的所述分光模块，所述分光模块用于将接收到的单束所述激光分为空间上角度分散或位置有所偏移的多束所述激光，多束所述激光传送到照明物镜；
    设置在所述分光模块下游的所述照明物镜，所述照明物镜用于对多束所述激光进行聚焦，在所述照明物镜的焦平面上产生二维点阵光斑；
    设置在所述照明物镜下游的流体聚焦模块，所述流体聚焦模块用于将样本中分散的待成像细胞排列成单细胞轴流，在依次经过所述二维点阵光斑的照明区域时，所述细胞激发散射光或荧光作为信号光；
    设置在所述流体聚焦模块下游的收光物镜，所述收光物镜用于接收所述信号光，并将所述信号光通过聚光镜聚焦后传送到光电探测器；
    设置在所述收光物镜下游的所述光电探测器，所述光电探测器用于将接收到的所述信号光转化为电压信号，并将所述电压信号发送给采集卡；
    设置在所述光电探测器下游的所述采集卡，所述采集卡用于对所述电压信号进行采集，经过模数转换转化为对应的数字信号，并将所述数字信号发送给计算机模块；
    设置在所述采集卡下游的所述计算机模块，所述计算机模块用于对接收到的所述数字信号进行处理，得到所述流体聚焦模块中被成像所述细胞的图像。
  2. 如权利要求1所述的基于点阵激光扫描的流式成像系统，其特征在于，所述二维点阵光斑对运动中的所述细胞进行扫描激发，在平行于所述细胞运动方向上的所述二维点阵光斑之间的距离大于所述细胞的长度，在垂直于所述细胞运动方向上的所述二维点阵光斑的位置分布覆盖所述细胞的不同位置。
  3. 如权利要求1或2所述的基于点阵激光扫描的流式成像系统，其特征在于，所述收光物镜和所述照明物镜为同一个物镜时，形成背向荧光检测光路；所述收光物镜和所述照明物镜不是同一个物镜时，形成前向荧光检测光路。
  4. 如权利要求3所述的基于点阵激光扫描的流式成像系统，其特征在于，所述分光模块为衍射光学器件分光系统、空间光调制器分光系统、数字微镜器件分 光系统、多光纤分光系统和多分光镜分光系统中的一种。
  5. 如权利要求4所述的基于点阵激光扫描的流式成像系统，其特征在于，所述衍射光学器件分光系统包括衍射光学器件，所述衍射光学器件表面具有浮雕结构；所述空间光调制器分光系统包括空间光调制器，所述空间光调制器由液晶面阵构成；所述数字微镜器件分光系统包括数字微镜器件，所述数字微镜器件为数字微镜阵列。
  6. 如权利要求5所述的基于点阵激光扫描的流式成像系统，其特征在于，所述流体聚焦模块采用传统流式细胞仪流动室或微流控芯片。
  7. 如权利要求6所述的基于点阵激光扫描的流式成像系统，其特征在于，所述光电探测器为单个光电倍增管、雪崩光电管、多路荧光检测系统或光谱检测系统中的一种。
  8. 如权利要求7所述的基于点阵激光扫描的流式成像系统，其特征在于，在所述光电探测器之前加设共聚焦结构。
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