Optical system and method for measuring light

Abstract

Systems and methods for measuring light that collect and measure light emanating from different locations along the length of a flow cell using a shared collection and magnification optical element; the light splitting module is used for splitting the amplified light into a plurality of wavelength ranges; and the spherical lens array guides the light with different wavelengths and emission positions to the detector array.

Description

Technical Field

[0001] The present disclosure relates to systems and methods for measuring light emission, and more particularly, to an optical system for measuring light emitted from fluorescently labeled biological cells, which can be used to analyze a population of biological cells. Background Art

[0002] Flow cytometry is a technique for detecting and measuring the physical and chemical properties of a population of biological cells. Broadly speaking, the process involves labeling cell biomarkers with a fluorescent dye and flowing the cells through a flow cell. As the cells pass through the flow cell, they are illuminated by a light source tuned to a frequency at which the fluorescent dye fluoresces. The fluorescence is measured to reveal information related to the cell population and properties.

[0003] In fact, most flow cytometry involves multi-color flow cytometry, which examines mixed cell populations, such as the different cell types found in blood and tissue. These cell types are typically distinguished from one another by labeling antibodies or antibody fragments specific to different biomarkers with different fluorescent dyes that fluoresce at different wavelengths. Due to the complexity of biological systems, the number of biomarkers used for labeling has increased significantly. Similarly, an increasing number of fluorescent dyes have become available, which has greatly expanded the ability to analyze more biomarkers in one experiment.

[0004] In addition, advanced electronic systems have been designed to acquire data from more fluorescence channels simultaneously and at extremely high speeds. However, even with advanced electronics, the capabilities of traditional flow cytometers remain limited because they are based on splitting the collected fluorescence signals into discrete channels and using a single channel that covers a narrow wavelength band to detect a specific fluorescent dye, and thus some fluorescent dyes with similar emission spectra cannot be used simultaneously in one experiment. Additionally, using traditional flow cytometers (if possible) significantly sacrifices the ability to simultaneously detect some fluorescent dyes with significant spectral overlap. In recent years, the development of spectral flow cytometers has addressed the limitations of traditional flow cytometers in the following way: collecting light over a wide wavelength range, containing multiple discrete channels, where the light from each channel is collected jointly (i.e., spectral data). A deconvolution algorithm is applied to the collected spectral data of samples stained with multiple fluorescent dyes (multi-stained samples) to calculate the fluorescence signal intensity of each fluorescent dye used in the sample, using the spectral data from each individual fluorescent dye (single-stained samples) as a reference. However, the key challenges in designing spectral flow cytometers are: (1) how to collect the emitted fluorescence from cells or particles when the fluorescent dyes are excited by each laser; (2) how to disperse or split the emitted fluorescence into a continuous spectrum within a larger wavelength range (i.e., 360 - 900 nm) and detect the dispersed light within a specific narrow wavelength range (i.e., one fluorescence detection channel).

[0005] For spectral flow cytometers and most flow cytometers, light collection is achieved by using optical fibers. This method uses light collection optics to collect fluorescence from each of the different lasers, which are typically spatially separated along the flow channel, and images the fluorescence to different focal points. For each excitation laser, one optical fiber is required to collect the fluorescence excited by it. Therefore, for N lasers, N optical fibers are needed to collect the fluorescence excited by all lasers. The light collection optics can be a specially designed objective, or multiple optical elements (e.g., lenses), which collect the emitted light from a single laser and focus the emitted light to a single focal point, thereby imaging all the emitted light to different focal points. An optical fiber array is placed at the focal plane, where each optical fiber matches one focal point from one laser and is used to guide the light at each focal point through a spectral splitting component (and then the split or dispersed light at different and continuous wavelength ranges will propagate to discrete photodetector arrays for each laser). To couple the light to the optical fiber array placed at the focal plane close to the light collection optics, the light collection optics system is typically designed to have a low magnification (e.g., <10x) and a short focal distance. For example, the lasers are typically spatially separated by 50 μm to 200 μm along the flow channel in a flow cytometer. The optical fiber array is typically spaced 500 μm to 1 mm. Therefore, the magnification of the light collection optics cannot be too high. Otherwise, the light at the focal plane cannot be effectively coupled to the optical fiber. The alignment of the optical fiber array with each focal spot is very important and requires very high mechanical alignment accuracy. However, even with a well-aligned optical fiber array, due to the different positions of each cell or particle passing through the flow cell, the focal spots will change. This results in a change in the intensity of the light signal collected by each optical fiber, leading to a large coefficient of variation (CV) of the detected fluorescence signal. Especially at higher sample flow rates, the CV may become unreasonably large. The higher the sample flow rate, the larger the core diameter of the sample flow in the flow channel. Cells and particles have a larger space away from the center of the sample core flow. Therefore, relative to the optically fixed fibers, the focal spots after the light collection optics change more in position. And as a result, the intensity of the light coupled to the optical fiber at each focal spot changes, resulting in a large CV. Therefore, using optical fibers in the light collection system results in a large (e.g., poor) signal CV, especially when the sample flow rate is high in high-throughput flow cytometry analysis. In addition, due to the small numerical aperture of the optical fiber (see, for example, the numerical aperture in optical fibers (questtel.com)), the overall light collection efficiency of the system is reduced, which leads to a decrease in the detection sensitivity of the fluorescence signal. In addition, the optical fiber also degrades over time and requires frequent maintenance or repair services. Such continuous maintenance of the optical fiber increases the additional cost of ownership.Another problem associated with fiber optic methods is that a single laser requires a set of beam splitting optics, which may involve beam splitting elements and / or optical components such as dichroic mirrors and bandpass filters. As the number of lasers increases, the number of optical components for splitting or dispersing light also increases, the system becomes very complex, and the cost increases significantly.

[0006] Two general methods (optical filtering method and optical dispersion method) have been used to split and disperse light in spectral and other flow cytometers. The optical filtering method uses optical components such as dichroic mirrors and bandpass filters to split light into different wavelength ranges. As described above, in a typical flow cytometer, a single laser requires a set of such optical components. As the number of lasers increases, the number of such optical components also increases, the system becomes very complex, and the cost increases significantly.

[0007] The optical dispersion method uses an optical element capable of dispersing light into a continuous spectrum, for example, a dispersive prism or a diffraction grating. In this type of design, the dispersion pattern of light is fixed (i.e., from short wavelength to long wavelength) and is typically non-linear across the entire spectrum (i.e., not evenly distributed in space). Thus, each detection channel (i.e., a specific wavelength range of dispersed light) will be limited in terms of wavelength narrowness. Additionally, since the light emitted from the optical light dispersion element typically propagates very narrowly in space, it is very difficult to physically place an optical detector array to detect light within each narrow wavelength range. In one method, a multi-channel photomultiplier (PMT) is used for light detection, but multi-channel PMTs are very expensive. As the number of lasers and fluorescence channels increases, the design of such systems becomes more complex. For example, an array of light dispersion elements may be required to achieve the narrow wavelength ranges of the detection channels, where each dispersion element needs to be arranged with high mechanical tolerance (i.e., high manufacturing cost). In addition to system complexity, the array of optical dispersion elements also reduces the transmission of the collected light, which results in significant light loss and reduced fluorescence signal detection sensitivity.

[0008] Accordingly, there is still a need for new systems and methods for flow cytometry that increase the number of fluorescent dyes available in a single experiment. SUMMARY OF THE INVENTION

[0009] The present disclosure addresses the above deficiencies in the art and provides related benefits. In one embodiment, this is achieved by an optical system for measuring light, the system comprising: a flow cell, the length and width of the flow cell allowing biological cells to pass through; a plurality of different excitation light sources that direct light to different positions along the length of the flow cell; a shared light collection optical element and a shared amplification optical element configured to collect light emitted from different positions of the flow cell, amplify and focus the collected light onto a focal plane at a specified distance according to different wavelength ranges, and amplify and focus the collected light onto different positions along the focal plane according to each different position of the flow cell from which the light is emitted; a spectral splitting module having spectral splitting optical elements arranged in such a way that the spectral splitting optical elements are shared by the amplified light emitted from each different position of the flow cell, and the spectral splitting optical elements are configured to split the amplified light emitted from each different position of the flow cell into a channel array, each channel corresponding to a separate wavelength range; an array of detection modules, wherein each detection module is dedicated to a single channel in the channel array, and wherein each detection module comprises: an array of spherical lenses located at each different focal distance for each different wavelength range along the propagation direction of the collected light, each spherical lens in the array being configured to receive the amplified light emitted from one of the different positions of the flow cell and direct the received amplified light to a predetermined area on a corresponding photodetector; an array of source photodetectors configured to measure the amplified light directed out of each spherical lens such that one photodetector measures the collected light from one of the different positions of the flow cell within one wavelength range.

[0010] In some embodiments, the plurality of excitation light sources include a plurality of laser light sources. In a further embodiment, the laser light sources are shaped into a designed geometry by beam shaping optical elements and focused to different positions within the flow cell.

[0011] In some embodiments, the plurality of different excitation light sources include more than 5 excitation light sources. In a further embodiment, there are more than 10 excitation light sources. In a further embodiment, optionally there are more than 15 excitation light sources.

[0012] In some embodiments, the distance between adjacent positions on the flow cell to which light is directed from the excitation light sources is from 50 μm to 250 μm. In a further embodiment, the distance is from 50 μm to 150 μm.

[0013] In some embodiments, a shared light collection optical element guides light to the shared amplification optical element, which amplifies the vertical separation distance between light emitted from different positions of the flow cell. In further embodiments, the shared amplification optical element includes a set of objective lenses with a magnification of 20 times to 250 times, optionally 30 times to 100 times, optionally 35 times to 75 times, optionally 40 times to 60 times. An exemplary vertical separation distance between adjacent lights is less than 30 mm at the spherical lens array. In some embodiments, the vertical separation distance is 1 mm to 10 mm at the spherical lens array.

[0014] As a non-limiting example, multiple excitation light sources guide light to adjacent positions spaced 100 μm apart along the flow cell, the shared amplification optical element has a magnification of 55 times, and the adjacent lights received at the spherical lens array are separated by 5.5 mm.

[0015] In some embodiments, the beam splitting module propagates light through a series of dichroic mirrors and bandpass filters to split the amplified light into multiple different wavelength ranges. In some embodiments, the multiple wavelength ranges include wavelength ranges from 320 nm to 1000 nm.

[0016] In some embodiments, the excitation light sources include lasers at 349 nm, 405 nm, 488 nm, 561 nm, and 637 nm. As a non-limiting example, the different wavelength ranges include 372 nm to 389 nm, 417 nm to 431 nm, 431 nm to 449 nm, 449 nm to 461 nm, 461 nm to 477 nm, 500 nm to 517 nm, 517 nm to 534 nm, 534 nm to 552 nm, 573 nm to 592 nm, 592 nm to 608 nm, 608 nm to 628 nm, 650 nm to 672 nm, 672 nm to 691 nm, 691 nm to 709 nm, 709 nm to 726 nm, 726 nm to 744 nm, 744 nm to 770 nm, 770 nm to 798 nm, and 798 nm to 832 nm.

[0017] In some embodiments, the light emitted from different positions sways in response to changes in different fluid properties of the flow channel and / or the positioning of the excitation light source. In further embodiments, the spherical lenses of each array guide the swaying light to the same area of the photodetector, where the detected light is stable.

[0018] In some embodiments, the photodetector system includes a silicon photomultiplier (SiPM) detector.

[0019] In some embodiments, each spherical lens is specifically designed such that the diameter of the light spot on the photodetector corresponds to the effective detection area of the detector.

[0020] In some embodiments, the collection optical element and the spherical mirror are specified such that the detected light spot on the corresponding photodetector is not affected by the position change of the collected light focused on the spherical lens.

[0021] In some embodiments, the spherical lens is designed such that the detected optical signal on the corresponding photodetector is insensitive to the horizontal position fluctuations of the biological cells in the flow channel.

[0022] In some embodiments, a set of alignment optical elements is configured to align the amplified light to the spectroscopic module and the spherical lens array.

[0023] In some embodiments, any optical system is integrated into a flow cytometer, and thus the flow cytometer can be provided with the optical system.

[0024] In a related aspect of the present invention, there is provided a method for characterizing biological cells, the method comprising: passing biological cells through a flow cell, the biological cells being labeled with a plurality of fluorescent markers that emit fluorescence at different wavelengths when excited, and tags attached to or absent from cell moieties co-characterize the biological cells by the presence, absence, or abundance of the markers; guiding light to different positions as the biological cells pass through different positions along the length of the flow cell; collecting the light emitted from each different position of the flow cell as the biological cells pass through; amplifying and focusing the collected light to different focal distances according to different wavelength ranges, and amplifying and focusing the collected light to different positions within each focal distance according to each different position of the flow cell from which the collected light is emitted; splitting the amplified light into a plurality of different wavelength ranges, wherein different fluorescence wavelengths are divided among the plurality of different wavelength ranges; receiving light of the same wavelength range emitted from different positions along the length of the flow cell by the same array of spherical lenses, wherein different spherical lenses within each array receive light emitted from different positions of the flow cell; guiding the light from each spherical lens to different photodetectors; and measuring the light at each different photodetector to identify the presence, absence, or abundance of the markers and to identify cell moieties to characterize the biological cells. BRIEF DESCRIPTION OF THE DRAWINGS

[0025] Figure 1 is a diagram of a flow cytometer including an exemplary optical system provided herein.

[0026] Figure 2It is a diagram depicting an embodiment for sampling fluorescently labeled cells.

[0027] Figure 3 It is a diagram depicting an embodiment for splitting light into different wavelengths.

[0028] Figure 4 It is a diagram of an exemplary spherical lens array.

[0029] Figure 5 It is a diagram showing detection within a single channel.

[0030] Figure 6 is Figure 3 a specification table of the illustrated embodiment.

[0031] Figure 7A It depicts the focusing of light at five separate and different positions 18a - e along the length of the flow cell, where the different adjacent positions 18a - e are spaced 100 μm apart from each other; Figure 7B It depicts the Figure 7A propagation of fluorescence emitted from the five different positions shown in Figure 7C towards a shared focal plane 56; and Figure 7B It depicts the

[0032] Figure 8A spacing of fluorescence at the shared focal plane 56 shown in Figure 8B where the spacing of light emitted from adjacent positions of the flow cell increases from 100 μm to 5.5 mm. Figure 8C It shows that when the fluorescence hits the middle of the spherical lens 54, the fluorescence is centered at the photodetector 52. Figure 8D It shows that when the fluorescence hits a position 1.65 mm to the right but vertically centered of the spherical lens 54, the fluorescence is centered at the photodetector 52. Detailed Description

[0033] In the present disclosure, the optical design of the light collection subsystem of a spectral flow cytometer is optimized to address the following issues: (a) the complex design associated with using a fiber optic array and the potentially lower light collection efficiency and higher cost, where each fiber in the fiber optic array guides light from a single laser excitation to a corresponding subsystem for light splitting and light detection; (b) the relatively large coefficient of variation (CV) of the signal associated with existing systems, particularly with respect to high sample flow rates for high-throughput cell analysis in current spectral flow cytometers that use fibers to collect fluorescence excited by different lasers; (c) the higher cost and complexity of the light collection subsystem design associated with using multiple sets of optical filters to split fluorescence from different lasers; (d) the poor spectral wavelength range resolution of individual fluorescence detection channels and the higher cost of using an array of light dispersive optical components.

[0034] For at least some of the light collection subsystems described herein, a novel light collection optical element subsystem and an innovative beam splitting (transmission and reflection) filter bank are included. For light collection, a custom objective lens is designed to have a high numerical aperture (e.g., greater than 1.2) and a large magnification factor (e.g., >30x). The high numerical aperture allows for the collection of more fluorescence emitted from cells or particles as they pass through the excitation laser beam. The objective lens with a large optical magnification focuses the emitted fluorescence from n lasers (n being the number of lasers in the system) onto separate spots in a vertical plane, where each focused spot is separated by a sufficiently large distance to be detected using an array of n individual detectors without any mechanical intervention. That is, the magnification factor of the objective lens needs to be large enough to allow for a sufficiently large distance between the focused spots due to the individual lasers. On the other hand, an overly large magnification factor may result in a larger footprint and a more complex system, which is undesirable for compact systems.

[0035] For beam splitting, the optical filter bank is designed to split the fluorescence excited by all the lasers in the system. The placement of the optical filter bank utilizes the focal length differences at different wavelengths across the entire spectral range of the collected fluorescence, such that light with different wavelengths travels in optical paths of different lengths. The end of the optical path for each discrete narrow wavelength range of fluorescence is the corresponding focal plane where the detection module is placed. The detection module includes an array of spherical lenses followed by an array of corresponding siPM detectors. Each spherical lens is placed at the focus of the fluorescence excited by the laser beam in the flow cell, thereby allowing the light to form a uniform distribution and cover as large an area as possible on the siPM detector. The lens located in front of the siPM photodetector enables the siPM to operate under its optimal performance conditions and can significantly reduce the variability of the detection signal, such that even at high sample flow rates, each fluorescence detection channel can achieve a very good signal CV.

[0036] The optical subsystem described in this article meets one or more of the following objectives: 1) A large numerical aperture for the optical collection system to maximize light collection. 2) A properly designed light collection objective lens that is used to optimize / balance the desired small footprint of the instrument (using a smaller magnification factor) with the objective of having the light from each excitation laser well separated at the focal plane (in the millimeter range), where an array of discrete photodetectors can be physically placed to detect the light from each excitation laser. 3) Design a light collection objective lens with a large magnification factor that is used to allow the light amplification path to have a sufficient length along which optical filters can be placed to split the fluorescence into a continuous spectrum of different narrow wavelength ranges. 4) A spherical lens is placed at the focal point of the collected fluorescence from each excitation laser (after the light has passed through a bandpass filter located before the SiPM photodetector). The focal length of the spherical lens is carefully designed so that after the light passes through the spherical lens, the diameter of the light spot on the detection chip behind the spherical lens matches well with the effective detection area of the photodetector. For example, if the effective photodetector area is 3 mm × 3 mm, the diameter of the light spot on the detection surface is designed to be close to but less than 3 mm so as to be effectively detected by the photodetector.

[0037] Despite the challenge of the complex design of the light collection objective lens, the development and placement of the spectral filter bank also require careful engineering development and innovation. A set of discrete bandpass filters is selected and arranged in the following way: The collected light first passes through one bandpass filter placed at a specific angle relative to the incident light and is detected by a photodetector for one channel, and the remaining light is reflected to the next bandpass filter. Thus, the collected side-scattered light and fluorescence sequentially pass through a set of discrete bandpass filters and are detected by an array of photodetectors located behind each bandpass filter. The system innovatively designs the path for the light to pass through a set of bandpass filters and detects the light for each specific wavelength range determined by the bandpass filters. In the case of a long focal distance, the scattered light and fluorescence enter each bandpass filter with a small half-cone angle and are focused behind the bandpass filter. The long focal distance allows the light to propagate in free space for a long enough time so that a series of bandpass filters can be placed.

[0038] This configuration has several advantages: 1) It is a free-space optical element and does not use optical fibers. Not using optical fibers provides higher light collection efficiency, higher stability of the collection optical elements, and more operational convenience. 2) Only one set of optical filters is required to split the light excited by all lasers. Using only one set of optical filters (instead of N sets of optical filters) simplifies the detection system and significantly reduces the product cost. There are surprising benefits to using the same optical filters to split the fluorescence excited by different lasers. Regardless of how the laser inducing the fluorescence changes, the fluorescence will be precisely split into the same wavelength range. This is particularly important for those fluorescent molecules excited by multiple lasers (when obtaining reference fluorescence spectra due to different lasers). Using the same filter to split the fluorescence generated by different lasers into exactly the same wavelength range helps to obtain consistent or similar reference fluorescence spectra due to different lasers and may help to deconvolve the measured fluorescence spectra collected for samples stained with multiple fluorescent dyes using the reference spectral data of each individual fluorescent dye, thereby calculating the fluorescence signal intensity of each fluorescent dye used in the sample. 3) The filter bank splits the light into discrete narrow wavelength ranges (i.e., one fluorescence detection channel) across the entire detection spectrum in free space, which significantly reduces light loss and improves the overall light collection efficiency. 4) The system allows the use of solid-state photodetectors (e.g., SiPM and APD), which have good photon detection performance and low cost. With this design, the spectral flow cytometry system can be designed to be simple in structure, with significantly reduced cost but excellent performance. 5) Importantly, the spherical lens has great advantages in providing the best CV of the detected optical signal at high sample flow rates for high-throughput cell analysis. As described above, the focal point of the fluorescence from each excitation laser is located on the first surface of the spherical lens. Due to inevitable fluid fluctuations, each individual cell or particle in the flow channel will fluctuate in the horizontal position as the sample fluid flows through the flow channel. Therefore, the horizontal position of the light focal point from each excitation laser on the first surface of each corresponding spherical lens surface will change. A higher sample flow rate will result in a larger fluid fluctuation core diameter, which means a larger sample distribution width in the horizontal direction. For example, if the fluid core diameter is 60 μm, considering the 55-fold magnification of the collection optical system, the focal points of the light collected from each excitation laser on the first surface of each corresponding spherical lens will have a diffusion width of 3.3 mm in the horizontal direction. When the diameter of the spherical lens is greater than 3.3 mm, the light from all samples can be fully collected without any loss. After passing through the spherical lens, the fluorescence from each excitation laser remains in the same position on the surface of the photodetection chip. Therefore, the detected optical signal is largely unaffected by sample variations.By reasonably designing the spherical lens, the size of the light on the surface of the photodetector chip can be designed to match the effective area of the photodetector chip. The detected light spot on the photodetector is not affected by the position change of the collected light focused on the first surface of the spherical lens, that is, the detected optical signal is insensitive to the sample horizontal position fluctuation in the flow channel. The same principle also applies to the vertical position change of the detected light. If the focused laser beam in the flow cell changes from its original vertical position, as long as the fluorescence from this laser is collected by the spherical lens in front of the photodetector, the detected optical signal can always remain stable, which can be achieved by using a spherical lens with a large diameter. In summary, the spherical lens located at the focal point of each fluorescence from each excitation laser and in front of the photodetector has unique advantages. It can provide a circular light spot with a uniform energy distribution, and the size of the light spot matches the effective detection area of the photodetector. The size and position of such a light spot on the surface of the photodetector do not change due to the position change of the original light source (i.e., the fluorescence emitted from each cell or particle in the flow channel).

[0039] In summary, all available lasers use the same set of optical filters, which significantly reduces the complexity of the system and reduces the cost of the system. The light collection optical element can be designed with a high numerical aperture (>1.2) to maximize the collection of light from each laser, which greatly improves the fluorescence signal detection sensitivity. In addition, the system is a free-space design, which minimizes light loss and reduces the need for potential frequent services that usually occur when using optical fibers. Since at the focal point, the light from each excitation laser is well separated spatially, an array of photodetectors (e.g., silicon photomultiplier (SiPM) and avalanche photodiode (APD)) can be used as detectors. These solid-state photodetectors have good light detection performance (i.e., high photon detection efficiency, high gain, low dark count, and low operating voltage), and are also compact, robust, and low-cost.

[0040] The above technical improvements, advantages, and benefits will be described in a series of non-limiting detailed embodiments in conjunction with the accompanying drawings. It should be understood that the same reference numerals are used to identify the same elements shown in one or more of the drawings.

[0041] From Figure 1Beginning, there is provided an optical system 10, a flow cytometer 100 including the optical system 10, and a method for measuring light, which are particularly applicable for determining the physical and chemical characteristics of cells labeled with fluorescent markers. The term "cell" as used herein refers to a biological cell, which is the smallest unit capable of independent survival and constituting all organisms and body tissues. A cell has three main parts, namely, a cell membrane, a nucleus, and cytoplasm. The system 10 and the method are particularly applicable for profiling cells by measuring light emitted from a fluorescent label that targets a specific cell biomarker, such as a cell surface antigen or an internal molecule. Based on the light measurement, the system 10 and the method can reveal the characterizing features of the cells, including cell size, cell granularity, DNA content, DNA gene expression, surface receptors, intracellular proteins, and transient signals. In addition, based on these characteristics, a cell profile can be generated based on function or activity, such as cells that resist immune infection (e.g., B cells, T cells), status (e.g., cancerous, compared to normal), stage (e.g., mature, compared to immature), origin (e.g., transgenic, compared to non-transgenic), and other known categories in the field to which the present disclosure pertains. In addition, compounds (e.g., potential therapeutic drugs) can also be characterized or profiled using the system 10 and the method described herein based on the measurable changes in a population of treated cells.

[0042] As will be described in more detail in the following paragraphs, cell characterization and profiling can be achieved by the following means: measuring the light emitted by the cells themselves, which can be used to determine the size and granularity of the cells; and / or measuring the emitted light from components attached to or inserted into the cells, such as fluorescently labeled antibodies, antibody fragments, or other molecules attached to or associated with the cells and capable of emitting fluorescence when excited. By evaluating the presence, absence, and / or abundance of these components, the system 10 and the method are capable of characterizing and profiling cells as well as compounds that affect the cells on a single-cell basis. Non-limiting examples of this type are provided in the present disclosure.

[0043] The light measurement described herein is preferably performed on a cell-by-cell basis such that data is collected and analyzed from individual cells. In addition, the measurement is very fast, for example, at a rate of about 10,000 cells per minute. Moving on to Figure 2 , Figure 2An embodiment of a flow cytometer 100 including an optical system 10 according to the technology of this specification is shown, which is achieved in part by rapidly passing cells through an elongated flow cell 12 having an inner diameter D slightly larger than the cells themselves. The sample / cells 30 can be passed through the flow cell 12 by integrating the flow cell 12 into a fluid system that, on the one hand, transports the cells 30 to the flow cell 12 and, on the other hand, removes the cells from the flow cell 12. This can be achieved using suitable fluid pumping mechanisms known in the art to which this disclosure pertains, such as those used in flow cytometer systems that transport sheath fluid 14 to and remove it from the flow cell 12. In some embodiments, the optical system 10 itself and its components form part of the flow cytometer 100.

[0044] As the cells pass through the flow cell 12, the excitation light source 16 directs light to different positions 18a - 18e along the length L of the flow cell 12. The wavelength of this light is capable of causing certain components to fluoresce. For example, the excitation light sources 16a - e direct light of specific wavelengths to different positions 18a - 18e along the length L of the flow cell 12. In some embodiments, the excitation light source 16 directs light having wavelengths of 349 nm, 405 nm, 488 nm, 561 nm, and 637 nm to different positions 18a - 18e along the length L of the flow cell 12. As the cells pass through the different positions 18a - 18e, the fluorescent tags attached to the cells can fluoresce, and this fluorescence is collected, resolved, and measured for each different position 18a - 18e along the length L of the flow cell 12 for analysis.

[0045] In Figure 2 the preferred embodiment described in detail, a single excitation light source 16 directs light to a single position 18a - 18e along the length L of the flow cell 12, thereby selectively pairing each different position 18a - 18e along the length L of the flow cell 12 with its own excitation light source 16. For example, the first position 18a along the length L of the flow cell 12 can receive light of a first wavelength (e.g., λ a = 349 nm) from the light source 16a, the second position 18b can receive light of a second wavelength (e.g., λ b = 405 nm) from the light source 16b, the third position 18c can receive light of a third wavelength (e.g., λ c = 488 nm) from the light source 16c, the fourth position 18d can receive light of a fourth wavelength (e.g., λ d = 561 nm) from the light source 16d, and the fifth position 18e can receive light of a fifth wavelength (e.g., λ e = 637 nm) from the light source 16e. Thus, in Figure 2In the illustrated embodiment, system 10 and method include five excitation light sources 16a - 16e; however, in other embodiments, system 10 and method include any number n of excitation light sources 16. In some embodiments, system 10 and method include n = seven or more than seven excitation light sources 16. In some embodiments, n = ten or more than ten excitation light sources 16. In some embodiments, system 10 and method include n = fifteen or more than fifteen excitation light sources 16.

[0046] In some embodiments, excitation light source 16 can direct light through an optical fiber cable to different positions 18a - 18e along the length L of flow cell 12. In other embodiments, excitation light source 16 is a laser directly aimed at different positions 18a - 18e of flow cell 12. In yet other embodiments, light emitted from a laser is directed to different positions 18a - 18e along the length of flow cell 16, for example, by using alignment mirrors. In each of the above configurations, excitation light source 16 can direct light through beam shaping optical element 20, which changes the initial beam geometry of the light from circular or substantially circular to one or more designed geometries. These designed geometries can take into account the following: on the one hand, the laser beam should be shaped wide enough to span the width of flow cell 12, thereby allowing the entire inner diameter D to be covered; on the other hand, the laser beam should be shaped to minimize the height to reduce the distance between adjacent positions 18a - 18e along the length L of flow cell 12. In some embodiments, these designed geometries are more preferably elliptical rather than circular and are aligned such that the major axis of the quasi - elliptical geometry extends across the inner diameter D of flow cell 12 rather than along the length L of flow cell 12. However, in other embodiments, the elliptical beam geometry is formed such that the major axis of the ellipse extends along the length L of flow cell 12.

[0047] The system and method analyze fluorescence signals emitted between adjacent positions 18a - 18e along the length L of the flow cell 12, where these adjacent positions are very close to each other, e.g., in the micron range. In some embodiments, the system 10 and method analyze fluorescence signals between adjacent positions 18a - 18e along the length L of the flow cell 12 that are separated by a distance 22 of 50 μm to 300 μm. In some embodiments, the fluorescence signals are analyzed from adjacent positions 18a - 18e along the length L of the flow cell 12 that are separated by a distance 22 of 60 μm to 250 μm. In some embodiments, the fluorescence signals are analyzed from adjacent positions 18a - 18e along the length L of the flow cell 12 that are separated by a distance 22 of 70 μm to 150 μm. In some embodiments, the fluorescence signals are analyzed from adjacent positions 18a - 18e along the length L of the flow cell 12 that are separated by a distance 22 of 75 μm to 100 μm. In some embodiments, the fluorescence signals are analyzed from adjacent positions 18a - 18e along the length L of the flow cell 12 that are separated by a distance 22 of the following distances: 50 μm to 55 μm, 55 μm to 60 μm, 60 μm to 65 μm, 65 μm to 70 μm, 70 μm to 75 μm, 75 μm to 80 μm, 80 μm to 85 μm, 85 μm to 90 μm, 90 μm to 95 μm, 95 μm to 100 μm, 100 μm to 105 μm, 105 μm to 110 μm, 110 μm to 115 μm, 115 μm to 120 μm, 120 μm to 125 μm, 125 μm to 130 μm, 130 μm to 135 μm, 135 μm to 140 μm, 140 μm to 145 μm, or 145 μm to 150 μm, 150 μm to 160 μm, 160 μm to 170 μm, 170 μm to 180 μm, 180 μm to 190 μm, and 190 μm to 200 μm. Preferably, the separation distance 22 between each pair of adjacent positions 18a - 18e along the length L of the flow cell 12 is the same among all adjacent positions 18a - 18e. That is, the separation distance 22 between adjacent first position 18a and second position 18b as the cell flows along the length L of the flow cell 12 is preferably the same as the separation distance 22 between adjacent second position 18b and third position 18c as the cell continues to flow along the length L of the flow cell 12, and so on. However, the present disclosure is not limited thereto, and thus in other embodiments, the adjacent separation distances 22 are not the same.

[0048] As the cells 30 flow along the length L of the flow cell 12, the cells are exposed to light from multiple excitation light sources 16, which causes light emission. That is, the cells and the fluorescently labeled molecules paired with the cells emit light when exposed to light of an appropriate wavelength, which is targeted at different positions 18a - 18e of the flow cell 12. The principles of light emission and fluorescence are well known in the art to which this disclosure pertains and are thus omitted for the sake of brevity.

[0049] A shared light collection optical element 24 and a shared amplification optical element 26 are used to collect and amplify the light emitted from different positions 18a - 18e along the length L of the flow cell 12. "Shared collection optical element" and "shared amplification optical element" mean that the light (intended for measurement) emitted from two or more (preferably all) of the different positions 18a - 18e along the length L of the flow cell 12 is collected using the same one or more collection lenses and amplified using the same one or more amplification lenses. The light collection lenses with a high numerical aperture (e.g., > 1.2) maximize the collection of light from each different position 18a - 18e along the length L of the flow cell 12, which significantly improves the detection sensitivity.

[0050] Figure 2 An exemplary configuration is shown in which the light collection objective lens 32 is integrated with both the shared light collection optical element 24 and the shared amplification optical element 26. Also as Figure 2As shown, the light collection optical element 24 is implemented as a hemispherical lens 34, which is adhesively attached to the surface of the flow cell 12 using an optical adhesive such that the emitted light from the flow cell 12 enters the light collection objective lens 32. Preferably, the hemispherical lens 34 is not attached to the light collection objective lens 32, such that the light collection objective lens 32 can be moved along three different axes to compensate for variations in the light emission position due to differences in hydrodynamic and excitation light source 16 alignment. After entering the light collection objective lens 32, the emitted light passes through a set of lenses, some of which may be converging lenses and some of which are diverging lenses, to achieve the desired magnification and focal length. Generally speaking, a converging lens tends to be thicker in the middle and directs the light passing through the lens towards the focal point; while a diverging lens tends to be thinner in the middle and has a negative focal length such that the light rays leaving the lens are separated from each other. By combining different lenses in series, the magnification of the light collection objective lens 32 can be from 20 times to 250 times. In some embodiments, the magnification of the light collection objective lens 32 is from 30 times to 100 times. In some embodiments, the magnification of the light collection objective lens 32 is from 40 times to 60 times. In some embodiments, the magnification of the light collection objective lens 32 is 40 times, 41 times, 42 times, 43 times, 44 times, 45 times, 46 times, 47 times, 48 times, 49 times, 50 times, 51 times, 52 times, 53 times, 54 times, 55 times, 56 times, 57 times, 58 times, 59 times or 60 times. In some embodiments, the light entering the light collection objective lens 32 passes through two plano-concave lenses, then a hemispherical lens, then a biconvex lens, and finally achieves a magnification of 40 times to 60 times.

[0051] In some embodiments, the objective lens 32 can be adjusted on three axes by mounting it on an X, Y, and Z stage 28. The three-axis alignment of the objective lens 32 can be used to compensate for tolerances in the light emission position.

[0052] Although one objective lens is used to magnify all the light emitted from the flow cell 12 (which is subsequently used for splitting and measurement), it is also important to ensure that the light emitted from each different position 18a - 18e along the length of the flow cell 12 remains distinguishable from each other. This can be achieved by combining a diverging lens (which again diffuses the light outwards) and one or more collimating lenses (which reduce the cross-section of each light beam and redirect the light along the same path). Examples of collimating lenses are well known in the art and include various plano-convex lenses.

[0053] Light emitted from different positions 18a - 18e along the length L of the flow cell 12 is transmitted from the shared light collection optical element 24 and the shared amplification optical element 26 of the light collection objective 32 to additional optical elements (e.g., mirror 42) and the spectral splitting module 38, which includes spectral splitting optical elements configured to split the amplified light into a plurality of different wavelength ranges (e.g., 40A - 40S) and then received by the detection module 50. Figure 3 The spectral splitting module and the detection module are shown more clearly, which house different channels 40A - S (collectively referred to as the "channel array") for measurement. For the spectral flow cytometer of the present disclosure, such different fluorescence channels or different wavelength ranges of the array are actually continuous, that is, the fluorescence channels are counted from the short wavelength range to the long wavelength range. For two adjacent channels, the longer wavelength end of the short range is the same as the shorter wavelength end of the 'next long wavelength range channel'. In one exemplary embodiment, the wavelength ranges of the channels are: 371nm - 389nm, 417nm - 431nm, 431nm - 449nm, 449nm - 461nm, 461nm - 477nm, 500nm - 517nm, 517nm - 534nm, 534nm - 552nm, 573nm - 592nm, 592nm - 608nm, 608nm - 628nm, 650nm - 672nm, 672nm - 691nm, 691nm - 709nm, 709nm - 726nm, 726nm - 744nm, 744nm - 770nm, 770nm - 798nm, 798nm - 920nm. Obviously, there are some wavelength gaps corresponding to the laser wavelengths, for example, the 389nm - 417nm gap for the 405nm laser, the 477nm - 500nm gap for the 488nm laser, the 552nm - 573nm gap for the 562nm laser, and the 628nm - 650nm gap for the 637nm laser. The correct alignment between the light collection objective 32 and the spectral splitting module 38 is ensured by a pair of mirrors 42 that can be adjusted in different directions.

[0054] In a preferred embodiment, the spectral splitting module 38 includes a short - pass or long - pass filter 44 that initially splits the received light into one of two different propagation paths 46A and 46B passing through the spectral splitting module 38. It can be understood that the filter 44 can be configured based on the desired number m of channels 40. Based on Figure 3The configuration of M = 19 channels (40A - 40S) shown enables two propagation paths 46A and 46B. However, the following also falls within the scope of the present disclosure: all light travels along the same propagation path, or the light is split multiple times to enable the light to be transferred between three or more propagation paths. Nevertheless, along each propagation path 46A, 46B, a series of bandpass filters 48 are aligned. Each bandpass filter 48 is designed to allow light of a specific wavelength to pass through while rejecting light of the remaining wavelengths. These rejected wavelengths are attenuated gradually along the propagation paths 46A, 46B passing through the beam splitting module 38, and thus encounter the next bandpass filter 48, and so on. Therefore, the alignment of the bandpass filters 48 with respect to the incident light is crucial for ensuring that the excluded wavelengths continue to propagate through the beam splitting module 38. As a non-limiting example, it has been found that arranging the bandpass filters 48 at an angle between 10° and 20° can achieve continuous propagation. However, it is more preferable to tilt each bandpass filter 48 at an angle of 12° to 16° with respect to the incident light, and an angle of 14° or 15° with respect to the incident light is currently most preferred. Thus, by propagating the amplified light from each different position 18a - 18e of the flow cell 12, the emitted light can be split into multiple different detection channels 40A - S, each channel corresponding to a specific wavelength range. Additionally, since the collection optical element 24 and the amplification optical element 26 are shared across different positions 18a - 18e along the length L of the flow cell 12, the emitted light from these different positions 18a - 18e can propagate in the same direction and thus pass through the same set of bandpass filters 48. Therefore, the bandpass filters 48 can be shared by the light emitted from each different position 18a - 18e of the flow cell 12, which itself reduces the footprint of the beam splitting module 38.

[0055] The light passing through each different bandpass filter 48 is received by the array of light detection modules 50. In Figure 3 the embodiment shown, dedicated light detection modules (50A - 50S) are coupled to each of the m channels (40A - 40S). Thus, the number of light detection modules in the array 50 corresponds to or matches the number of channels m generated by the beam splitting module 38. As Figure 4 and Figure 5 shown, each light detection module 50 will include or be coupled to a dedicated spherical lens array 54, which will then direct the amplified light from each different position 18a - 18e along the length L of the flow cell 12 to a predetermined photodetector array 52 for measurement. That is, the spherical lens array 54 is assigned to each channel 40A - 40S defined by the corresponding bandpass filter 48 and is positioned at a specific distance along the propagation path passing through the bandpass filter 48, determined according to the wavelength and the optical elements, such that each spherical lens 54a - 54e within the same array 54 (as Figure 4as shown) is configured to receive light within the same wavelength range. However, each spherical lens 54a - 54e within the array 54 receives light emitted from different excitation light sources 16a - 16e (having corresponding assigned wavelengths (λ a , λ b , λ c , λ d , λ e )) and positions 18a - 18e along the length L of the flow cell 12 (see Figure 2 ). That is, the spherical lens array 54 not only receives and directs light within a specific wavelength range (according to its channel) to the detector 52 for measurement, but also differentiates between light emitted from each different position 18a - 18e along the length L of the flow cell 12 for measurement.

[0056] The positioning and assigned distances of the spherical lens array 54 are determined based on the focal length of the light passing through the corresponding band - pass filter 48. That is, each spherical lens 54 within the same array is arranged on the same plane 56 (also referred to herein as the "focal plane"). This same plane 56 corresponds to the focal length f of the light assigned to channels 40A - S, which is itself determined by the focal lengths of the shared collection optical element 24, the shared amplification optical element 26, and the band - pass filter 48. That is, the array 54 of spherical lenses 54a - 54e is positioned at the foci 58a - 58e on the plane 56 for each channel 40A - S. Figure 6 An example is shown, which Figure 6 is a table summarizing the exemplary focal lengths f of nineteen (19) channels (e.g., channels 40A - 40S) and the positioning of nineteen arrays 54 of spherical lenses (54a - 54e).

[0057] For example, Figure 6 the first - listed channel 1 in Figure 3 corresponds to channel 40A shown in A which has a focal length f A = 692.17 mm for a wavelength of 380 nm (the mid - value of the 371 nm to 389 nm range for channel 40A), where the spherical lens array 54 assigned to the light detection module 50A is positioned in the focal plane 56 at the focal length f A = 692.17 mm. Thus, there are f M focal lengths corresponding to 40 M channels and 50 M detection modules. For a system with Figure 2 and Figure 3 having M = 19 channels and n = 5 excitation light sources, a total of M x n (19 x 5)=95 detectors and spherical lenses are used.

[0058] In the above example, a collection optical element with a magnification of 55 times and a separation distance 22 of 100 μm are used ( Figure 3 ), and the vertical separation 60 between two adjacent light focus points (foci 58a - 58e) is 5.5 mm. Thus, the spherical lenses 54a - 54e are aligned on the focal plane 56 and are vertically spaced 5.5 mm apart from each other (C - C).

[0059] Now referring comprehensively to Figures 1 - 7C , although each of the spherical lenses 54a - 54e in the single spherical lens array 54 is aligned with each other along the same plane 56, the spherical lens array 54 also differentiates between the light emitted from each different position 18a - 18e along the length L of the flow cell 12 by assigning each spherical lens 54a - 54e within the array 50 to a different focus 58a - 58e within the focal plane 56. Each of these foci 58a - 58e within the focal plane 56 is the result of collecting and magnifying the light emitted from different positions 18a - 18e along the length L of the flow cell 12. As shown more clearly in Figures 7A - 7C , resolving the fluorescence emitted from different positions 18a - 18e along the length of the flow cell includes increasing the separation distance between adjacent positions 18a - 18e. For example, light can be emitted from adjacent positions 18a - 18e that are only 100 μm apart (see Figure 7A ). Such small distances are difficult to resolve. However, by passing the emitted light through an array of different lenses (see Figure 7B ), the separation distance can be significantly increased ( Figure 7C depicts a separation distance of 5.5 mm at the focal plane). This increased separation distance at the focal plane 56 allows the spherical lenses 56a - 56e to resolve the signals from each different position 18a - 18e along the length of the flow cell for transmission to the photodetector 52. That is, the light emitted from each different position 18a - 18e of the flow cell 12 is finally focused into light spots, each light spot located at a different focus 58a - 58e along the focal plane 56, and by positioning the spherical lenses 54a - 54e at these points 58a - 58e, the emitted light can then be directed to a predetermined area of dedicated photodetectors 52a - 52e for measurement. By maintaining the identity of the light source throughout the spectroscopic module 38, the emitted light is directed to specific foci 58a - 58e along the focal plane 56. Since the flow cell 12 is typically arranged vertically, as shown in Figure 4As shown more clearly therein, the spherical lenses 54a-54e are also generally arranged to be vertically offset from each other. In some embodiments, the vertical separation distance 60 between adjacent lights (i.e., adjacent foci 58a-58e) is less than 30 mm at the spherical lens array 50. In some embodiments, the vertical separation distance 60 between adjacent foci 58a-58e is from 1 mm to 10 mm. In some embodiments, each vertical separation distance 60 between adjacent foci 58a-58e is 5.5 mm.

[0060] Each spherical lens 54a-54e within each spherical lens array 54 directs the received light to a different photodetector 52a-52e for measurement, thereby allowing measurement based on the emission wavelength and emission position from the flow cell 12. One of the challenges in transmitting light from the flow cell 12 to the detector 52 is miniaturization, at least in part due to the non-uniform flow or wobbling of the cells as they pass through the flow cell 12. That is, due to changes in the fluid properties and the positioning of the excitation light source 16, inconsistencies were initially found as the cells passed through the flow cell 12. These inconsistencies are amplified throughout the system, which poses a significant challenge in manufacturing the flow cytometer 100 with a small footprint. This problem is partially solved by using the spherical lenses 54a-54e in the spherical lens array 54, which correct the inconsistencies in the optical alignment to the appropriate photodetectors 52a-52e. The correction of the inconsistencies was experimentally demonstrated in Figure 8A -D. For reference throughout Figure 8A -D, Figure 8A depicts the terminal portion of the propagated fluorescence through the system, wherein the propagated light is directed to the detector 52 by the spherical lens 54 located at the focal plane 56. Figure 8B depicts the result of an ideal alignment, wherein the propagated light hits the center of the spherical lens 54 and is thus directed to a predetermined or effective area of the photodetector 52. However, due to various factors, in some cases, the light hits an off-center position of the spherical lens 54. Figure 8C shows the result where the light hits a position that is horizontally off-center (to the right) by 1.65 mm of the spherical lens 54. Due to the radius of the spherical lens 54, the light can be redirected to a predetermined or effective area of the photodetector 52. Figure 8D shows the result where the light hits a position that is horizontally off-center (to the left) by 2.2 mm and vertically off-center (downward) by 0.55 mm of the spherical lens 54. In this case, the spherical lens 54 corrects the propagation path such that the light hits the predetermined or effective area of the detector 52 again. In instances where the light hits an off-center position of the spherical lens 54, this correction or redirection is at least partially achieved by the radius of the spherical lens 54. In Figure 8AIn the demonstration depicted by -D, the spherical lenses 56 each have a radius of 4.1 mm and are used to direct or redirect light to a predefined or effective area smaller than 3 mm × 3 mm. Given the above, those skilled in the art will recognize that the radius of the spherical lens 54 can vary at least in part depending on the predefined or effective area of the detector 52 and the spacing difference of the propagating light. Thus, the radius is not limited to any specific size, but as a general guide to start with, the radius can be from 1 mm to 20 mm, for example, 1 mm, 2 mm, 3 mm, 4 mm, 5 mm, 6 mm, 7 mm, 8 mm, 9 mm, 10 mm, 11 mm, 12 mm, 13 mm, 14 mm, 15 mm, 16 mm, 17 mm, 18 mm, 19 mm, 20 mm or any size in between.

[0061] The photodetectors 52a - 52e (collectively referred to as the "photodetector array") (which may include silicon photomultipliers (SiPMs), avalanche photodiodes (APDs), or similar detectors) are grouped into M light detection modules 50, each light detection module 50 including an array having n photodetectors such that one photodetector 52a - 52e of each light detection module 50A - 50S measures light that corresponds to a single wavelength range within channels 40A - 40S and comes from one of the different positions 18a - 18e along the length L of the flow cell 12. By measuring the light located at each of the channels 40A - 40S and coming from each of the different positions 18a - 18e along the length L of the flow cell 12, the system 10 and method are able to identify the presence, absence, and / or abundance of a fluorescent marker and thus determine the cellular components for characterizing the cell.

[0062] The above description of various embodiments of the technology of the present disclosure is presented for purposes of illustration and description. It is not intended to limit the technology of the present disclosure to the specific forms disclosed. Many modifications, variations, and improvements will be apparent to those skilled in the art. For example, embodiments of the device and its optical elements can be sized and otherwise adapted for various detector and / or flow cell configurations. Additionally, those skilled in the art will recognize or be able to ascertain using routine experimentation numerous equivalents to the specific devices and methods described herein. These equivalents are considered to fall within the scope of the technology of the present disclosure and are covered by the appended claims.

[0063] Elements, features, or acts in one embodiment can be readily recombined with one or more elements, features, or acts in other embodiments or replaced with one or more elements, features, or acts in other embodiments, thereby forming numerous other embodiments within the scope of the technology of the present disclosure. Additionally, elements shown or described as combined with other elements can exist as independent elements in various embodiments. Accordingly, the scope of the technology of the present disclosure is not limited to the specific details of the described embodiments but is only limited by the appended claims.

Claims

1. An optical system for measuring light, the system comprising: a) A flow cell, the length and width of the flow cell allowing biological cells to pass through; b) A plurality of different excitation light sources, the plurality of different excitation light sources guiding light to different positions along the length of the flow cell; c) Shared light collection optical elements and shared amplification optical elements, the shared light collection optical elements and shared amplification optical elements being configured to collect light emitted from different positions of the flow cell, amplify and focus the collected light onto a focal plane at a specified distance according to different wavelength ranges, and amplify and focus the collected light onto different positions along the focal plane according to each different position of the flow cell from which the light is emitted; d) A spectroscopic module, the spectroscopic module including spectroscopic optical elements arranged in such a way that the spectroscopic optical elements are shared by the amplified light emitted from each different position of the flow cell, and the spectroscopic optical elements being configured to split the amplified light emitted from each different position of the flow cell into a channel array, each channel corresponding to a separate wavelength range; e) An array of detection modules, wherein each detection module is dedicated to a single channel in the channel array, each detection module comprising: i) A spherical lens array, the spherical lens array being located at each different focal distance for each different wavelength range along the propagation direction of the collected light, each spherical lens in the array being configured to receive the amplified light emitted from one of the different positions of the flow cell and guide the received amplified light to a predetermined area on a corresponding photodetector; and ii) A photodetector array, the photodetector array being configured to measure the amplified light guided out from each spherical lens, such that one photodetector measures the collected light from one of the different positions of the flow cell within one wavelength range.

2. The system according to claim 1, wherein, The plurality of excitation light sources includes a plurality of laser light sources.

3. The system according to claim 2, wherein The laser light sources are shaped into a designed geometric shape by beam shaping optical elements and focused to different positions within the flow cell.

4. The system according to claim 1, wherein, The plurality of different excitation light sources includes more than 5 excitation light sources, optionally more than 10 excitation light sources, optionally more than 15 excitation light sources.

5. The system according to claim 1, wherein, The distance between adjacent positions on the flow cell to which light is guided from the excitation light sources is 25 μm to 250 μm, optionally 50 μm to 150 μm.

6. The system according to claim 1, wherein, The shared light collection optical elements guide light to the shared amplification optical elements, and the shared amplification optical elements amplify the vertical separation distance between the lights emitted from different positions of the flow cell.

7. The system according to claim 6, wherein, The shared amplification optical elements include a set of objective lenses, the magnification of the set of objective lenses being 20 times to 250 times, optionally 30 times to 100 times, optionally 35 times to 75 times, optionally 40 times to 60 times.

8. The system according to claim 6, wherein The vertical separation distance between adjacent lights is less than 30 mm at the spherical lens array.

9. The system according to claim 8, wherein The vertical separation distance is 1 mm to 10 mm at the spherical lens array.

10. The system according to claim 9, wherein, The plurality of excitation light sources direct light to adjacent positions spaced 100 μm apart along the flow cell, the shared amplification optical element has a magnification of 55 times, and the adjacent light received at the spherical lens array is separated by 5.5 mm.

11. The system according to claim 1, wherein, The spectroscopic module propagates light through a series of dichroic mirrors and bandpass filters to split the amplified light into a plurality of different wavelength ranges.

12. The system according to claim 1, wherein The plurality of wavelength ranges include wavelength ranges from 320 nm to 1000 nm.

13. The system according to claim 12, wherein, The excitation light sources include lasers at 349 nm, 405 nm, 488 nm, 561 nm, and 637 nm.

14. The system according to claim 12, wherein, The different wavelength ranges include 372 nm to 389 nm, 417 nm to 431 nm, 431 nm to 449 nm, 449 nm to 461 nm, 461 nm to 477 nm, 500 nm to 517 nm, 517 nm to 534 nm, 534 nm to 552 nm, 573 nm to 592 nm, 592 nm to 608 nm, 608 nm to 628 nm, 650 nm to 672 nm, 672 nm to 691 nm, 691 nm to 709 nm, 709 nm to 726 nm, 726 nm to 744 nm, 744 nm to 770 nm, 770 nm to 798 nm, and 798 nm to 832 nm.

15. The system according to claim 1, wherein, The light emitted from different positions sways in response to changes in the fluid properties of the flow channel and / or changes in the positioning of the excitation light sources.

16. The system according to claim 15, wherein, The spherical lenses of each array direct the swaying light to the same area of the photodetector, where the detected light is stable.

17. The system according to claim 1, wherein, The photodetector system includes a silicon photomultiplier (SiPM) detector.

18. The system according to claim 1, wherein, Each spherical lens is specifically designed such that the diameter of the light spot on the photodetector corresponds to the effective detection area of the detector.

19. The system according to claim 1, wherein The collection optical element and the spherical mirror are specified such that the detected light spot on the corresponding photodetector is not affected by changes in the position of the collected light focused on the spherical lens.

20. The system according to claim 1, wherein The spherical lenses are designed such that the detected light signals on the corresponding photodetectors are insensitive to horizontal position fluctuations of the biological cells in the flow channel.

21. The system according to claim 1, further comprising: A set of alignment optical elements, configured to align the amplified light to the spectroscopic module and the spherical lens array.

22. A flow cytometer, comprising the system according to claim 1.

23. A method for characterizing biological cells, the method comprising: a) passing biological cells through a flow cell, the biological cells being labeled with a plurality of fluorescent markers that emit fluorescence at different wavelengths when excited, and tags attached to or absent from cell components co-characterize the biological cells by the presence, absence, or abundance of the markers; b) directing light to different positions along the length of the flow cell as the biological cells pass through; c) collecting light emitted from each different position of the flow cell as the biological cells pass through; d) Amplify and focus the collected light to different focal distances according to different wavelength ranges, and amplify and focus the collected light to different positions within each focal distance according to each different position of the flow cell from which the collected light is emitted; e) Split the amplified light into multiple different wavelength ranges, wherein different fluorescence wavelengths are divided among the multiple different wavelength ranges; f) Receive light of the same wavelength range emitted from different positions along the length of the flow cell by the same array of spherical lenses, wherein different spherical lenses within each array receive light emitted from different positions of the flow cell; g) Guide the light from each of the spherical lenses to different photodetectors; and h) Measure the light at each of the different photodetectors to thereby identify the presence, absence or abundance of the marker and identify cell components to characterize the biological cell.

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