Optical system and method for measuring light

Abstract

The system and method for measuring light collects and measures light emitted from different locations along the length of a flow cell using shared focusing and magnification optics, a light splitting module that splits the magnified light into multiple wavelength ranges, and an array of spherical lenses that directs light of different wavelengths and emission locations to an array of detectors.

Description

[Technical Field]

[0001] The present invention relates to systems and methods for measuring light emissions, and more particularly to optical systems for measuring light emitted from fluorescently labeled biological cells that can be used to analyze biological cell populations. [Background technology]

[0002] Flow cytometry is a technique used to detect and measure the physical and chemical properties of populations of biological cells. Broadly speaking, the process involves labeling cellular biomarkers with fluorescent dyes and flowing the cells through a flow cell. As the cells pass through the flow cell, they are illuminated with a light source tuned to a frequency that causes the fluorescent dyes to fluoresce. This fluorescence is measured to reveal information about the cell population and properties.

[0003] In practice, most flow cytometry involves multicolor flow cytometry, which examines mixed cell populations (e.g., the various cell types found in blood and tissues). These cell types are typically distinguished from one another by labeling antibodies or antibody fragments specific for different biomarkers with different fluorescent dyes that fluoresce at different wavelengths. The complexity of biological systems has greatly increased the number of biomarkers that can be labeled. Similarly, the availability of an increasing number of fluorescent dyes has greatly expanded the ability to analyze more biomarkers in a single experiment.

[0004] In addition, advanced electronics systems are being designed to acquire data from many more fluorescent channels simultaneously and at exceptional speeds. However, even with advanced electronics, traditional flow cytometers remain limited in their capabilities. Because they rely on splitting the collected fluorescent signal into separate channels, each covering a narrow wavelength band for a specific fluorophore, several fluorophores with similar emission spectra cannot be used simultaneously in a single experiment. Furthermore, the ability to simultaneously detect several fluorophores with large spectral overlap is significantly compromised when using traditional flow cytometers, even if possible. Recently, the development of spectral flow cytometers has addressed the limitations of traditional flow cytometers by collecting light across a wide wavelength range, including multiple individual channels (i.e., spectral data) where light from each channel is collected collectively. Deconvolution algorithms are applied to the collected spectral data of samples stained with multiple fluorophores (multiple-stained samples) to calculate the fluorescent signal intensity of each fluorophore used in the sample using the spectral data from each individual fluorophore (single-stained samples) as a reference. However, key challenges to the design of a spectral flow cytometer are (1) how to collect the fluorescence emitted from cells or particles when fluorophores are excited by each laser, and (2) how to disperse or split the emitted fluorescence into a continuous spectrum spanning a wide wavelength range (i.e., 360 nm to 900 nm) and detect the dispersed light in a specific narrow wavelength range (i.e., one fluorescence detection channel).

[0005] In spectral flow cytometers and most flow cytometers, light collection is achieved through the use of optical fibers. This method uses collection optics to collect and image the fluorescence from each of the different lasers, which are typically spatially separated along the flow path, onto different focused spots. For each excitation laser, an optical fiber is required to collect the excited fluorescence. Thus, for N lasers, N optical fibers are used to collect the fluorescence excited by all of the lasers. The collection optics may be a specially designed objective lens or multiple optical elements (e.g., lenses) that collect the emission light from a single laser, focus it onto a single focused spot, and image all the emission light onto different focused spots. An array of optical fibers is positioned at the focusing plane, each corresponding to a focused spot from one laser, and is used to guide the light at each focused spot through a light-splitting component (the split or dispersed light at different continuous wavelength ranges then propagates into an array of separate photodetectors for each laser). To couple light into the optical fiber array, which is located at a focal plane close to the collection optics, the collection optics are typically designed with low magnification (e.g., <10x) and short focal length. For example, lasers are typically spatially separated by 50 μm to 200 μm along the flow path within a flow cytometer. The optical fiber array is typically spaced 500 μm to 1 mm apart. Therefore, the magnification of the collection optics cannot be too high. Otherwise, the light at the focal plane cannot be effectively coupled into the optical fibers. Aligning the optical fiber array to each focused light spot is crucial and requires very high mechanical alignment precision. Furthermore, even with a properly aligned optical fiber array, the focused light spot may shift due to the varying position of each cell or particle passing through the flow cell. This contributes to fluctuations in the optical signal intensity collected by each optical fiber, leading to an increase in the coefficient of variation (CV) of the detected fluorescent signal. CV can become excessively large, especially with increasing sample flow rates. As the sample flow rate increases, the core diameter of the sample stream inside the flow path increases.Cells and particles occupy more space away from the center of the sample core stream. Therefore, the focal spot after the collection optics varies in position more significantly compared to the fixed-position optical fiber. As a result, the intensity of light coupled into the optical fiber at each focal spot varies, leading to an increased CV. Therefore, the use of optical fibers in a collection system results in a large (e.g., worse) signal CV, especially as sample flow rates increase in high-throughput flow cytometry analysis. In addition, the small numerical aperture of optical fibers (see, e.g., Numerical Aperture in Optical Fibers (questtel.com)) reduces the overall light collection efficiency of the system, leading to reduced sensitivity in detecting fluorescent signals. Furthermore, optical fibers also deteriorate over time, requiring frequent maintenance or repair services. Such periodic maintenance of optical fibers adds additional ownership costs. Another problem associated with the optical fiber approach is that one laser requires a set of light-splitting optics. This may include optical components such as light-dispersing elements and / or dichroic mirrors and bandpass filters. As the number of lasers increases, the number of optical components used to disperse or split the light increases, significantly increasing the system's complexity and cost.

[0006] Two common methods are used to split and disperse light in spectral and other flow cytometers: optical filtering and optical dispersion. Optical filtering uses optical components such as dichroic mirrors and bandpass filters to split light into different wavelength ranges. As mentioned above, in a typical flow cytometer, one laser requires one set of such optical components. As the number of lasers increases, the number of such optical components increases, making the system very complex and significantly increasing costs.

[0007] Optical dispersion techniques use optical elements (e.g., dispersive prisms or diffraction gratings) that can disperse light into a continuous spectrum. In this type of design, the dispersion pattern of light is fixed (i.e., from short to long wavelengths) and is usually nonlinear across the spectrum (i.e., not evenly spread in space). Thus, each detection channel (i.e., a specific wavelength range of dispersed light) has a limited narrow wavelength range. In addition, the spread of light from the optical dispersion element is usually very narrow in space, making it very difficult to physically position an array of optical detectors to detect light in each narrow wavelength range. One approach uses multichannel photomultiplier tubes (PMTs) for light detection, but multichannel PMTs are very expensive. As the number of lasers and fluorescence channels increases, the design of such systems becomes even more complex. For example, an array of optical dispersion elements may be required to achieve narrow wavelength ranges for the detection channels, and each dispersion element must be positioned with high mechanical tolerances (i.e., high manufacturing costs). In addition to the system complexity, the array of optical dispersion elements reduces the transmission of the collected light, leading to significant light loss and reduced sensitivity in detecting the fluorescence signal.

[0008] Thus, there remains a need for new systems and methods for use in flow cytometry that increase the number of fluorochromes that can be utilized in a single experiment. Summary of the Invention

[0009] The present invention addresses the aforementioned deficiencies in the art and provides related advantages, comprising an optical system for measuring light, the system comprising: a flow cell having a length and width that allows passage of biological cells; a plurality of different excitation light sources that direct light to different locations along the length of the flow cell; shared collection optics and shared magnification optics configured to collect light emitted from the different locations of the flow cell and magnify and focus the collected light onto a focal plane at specific distances according to different wavelength ranges and to different locations along the focal plane according to each of the different locations of the flow cell from which the light was emitted; and a light splitting module having light splitting optics arranged such that the light splitting optics is shared by the magnified light emitted from each of the different locations of the flow cell, the light splitting optics configured to split the magnified light emitted from each of the different locations of the flow cell into an array of channels, each channel corresponding to a distinct wavelength range. This can be achieved by an optical system comprising: a light splitting module; an array of detection modules, each detection module dedicated to an individual channel of the array of channels, each detection module being an array of spherical lenses arranged at different focal lengths along the propagation direction of collected light for each of different wavelength ranges, each spherical lens in the array being configured to receive magnified light emitted from one of the different positions of the flow cell and direct the received magnified light to a predetermined area on a corresponding photodetector; and an array of photodetectors configured to measure the magnified light directed from each of the spherical lenses, whereby one photodetector measures collected light within one wavelength range from one of the different positions of the flow cell.

[0010] In some embodiments, the multiple excitation light sources include multiple laser light sources, hi further embodiments, the laser light sources are shaped by beam shaping optics into a designed shape and focused at different locations within the flow cell.

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

[0012] In some embodiments, the distance between adjacent locations on the flow cell where light is directed from the excitation light source is between 50 μm and 250 μm. In further embodiments, the distance is optionally between 50 μm and 150 μm.

[0013] In some embodiments, the shared collection optics directs light to a shared magnification optics, which magnifies the vertical separation between light emitted from different locations on the flow cell. In further embodiments, the shared magnification optics includes a set of objective lenses having a magnification of 20x to 250x, optionally 30x to 100x, optionally 35x to 75x, and optionally 40x to 60x. Exemplary vertical separations between adjacent light in an array of spherical lenses include less than 30 mm. In some embodiments, the vertical separation in an array of spherical lenses is 1 mm to 10 mm.

[0014] As a non-limiting example, multiple excitation light sources direct light to adjacent positions 100 μm apart along the flow cell, the shared magnification optics has a magnification of 55X, and adjacent light received at the array of spherical lenses is 5.5 mm apart.

[0015] In some embodiments, the light splitting module propagates the light through a series of dichroic mirrors and bandpass filters to split the expanded light into multiple distinct wavelength ranges, which in some embodiments include the wavelength range 320 nm to 1000 nm.

[0016] In some embodiments, the excitation light source comprises a 349 nm, 405 nm, 488 nm, 561 nm, and 637 nm laser. By way of non-limiting example, 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 locations fluctuates in response to changing fluid properties through the flow path and / or changes in the positioning of the excitation light source. In further embodiments, a spherical lens with each array directs the fluctuating light to the same region 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 specially designed so that the diameter of the light spot on the photodetector matches the active detection area of ​​the detector.

[0020] In some embodiments, the collection optics and spherical mirror are designed so that the detected light spot on the corresponding photodetector is not affected by changes in the position of the collected light focused onto the spherical lens.

[0021] In some embodiments, the spherical lenses are designed such that the detected optical signal on the corresponding photodetector is not affected by lateral positional variations of the biological cells within the channel.

[0022] In some embodiments, a set alignment optics is configured to align the expanded light onto the light splitting module and the array of spherical lenses.

[0023] In some embodiments, any of the optical systems may be integrated into the flow cytometer, thus providing the optical system in the flow cytometer.

[0024] In a related aspect of the invention, there is provided a method for characterizing biological cells, the method comprising the steps of passing biological cells through a flow cell, the biological cells being labeled with a plurality of fluorescent markers that fluoresce at different wavelengths when excited, the markers being attached or absent to portions of the cell that together characterize the biological cell by the presence, absence, or abundance of the marker; directing light to different positions along the length of the flow cell as the biological cell passes the different positions; collecting light emitted from each of the different positions of the flow cell as the biological cell passes; and directing the collected light to different focal lengths according to different wavelength ranges and for each of the different positions of the flow cell from which the collected light was emitted. the method includes the steps of magnifying and focusing the magnified light at different positions within each focal length accordingly, splitting the magnified light into a plurality of different wavelength ranges, wherein different fluorescent wavelengths are split into a plurality of different wavelength ranges, receiving light of the same wavelength range emitted from different positions along the length of the flow cell with the same array of spherical lenses, wherein different spherical lenses in each array receive light emitted from different positions of the flow cell, directing the light from each of the spherical lenses to a different photodetector, and measuring the light at each of the different photodetectors, thereby identifying the presence, absence, or abundance of the marker and therefore identifying a cellular portion that characterizes the biological cell. [Brief explanation of the drawings]

[0025] [Figure 1] FIG. 1 is a diagram of a flow cytometer incorporating an exemplary optical system provided herein. [Figure 2] FIG. 1 illustrates one embodiment for sampling fluorescently labeled cells. [Figure 3]FIG. 1 illustrates an embodiment for splitting light into different wavelengths. [Figure 4] FIG. 1 is a diagram of an exemplary array of spherical lenses. [Figure 5] FIG. 1 shows detection in a single channel. [Figure 6] 4 is a table of specifications for the embodiment shown in FIG. 3. [Figure 7A] FIG. 1 shows the focusing of light at five separate, distinct locations 18a-18e along the length of the flow cell, with various adjacent locations 18a-18e spaced 100 μm apart from each other. [Figure 7B] 7B shows the propagation of fluorescent light emitted from the five different positions shown in FIG. 7A towards the shared focal plane 56. FIG. [Figure 7C] FIG. 7C shows the spacing of fluorescence in the shared focal plane 56 shown in FIG. 7B, where the spacing of light emitted from adjacent locations on the flow cell is increased from 100 μm to 5.5 mm. [Figure 8A] FIG. 10 shows another exemplary propagation of light in the form of fluorescent light from five different positions in a flow cell, where the fluorescent light in a shared focal plane 56 is directed to a photodetector 52 via a spherical lens 54. [Figure 8B] FIG. 10 shows the centering of the fluorescence at the photodetector 52 when the fluorescence strikes the center of the spherical lens 54. [Figure 8C] FIG. 10 shows the centering of the fluorescence at the photodetector 52 when the fluorescence strikes a position that is offset 1.65 mm to the right of the spherical lens 54 but is vertically centered. [Figure 8D] FIG. 10 shows the centering of the fluorescence at the photodetector 52 when the fluorescence strikes the spherical lens 54 at a position offset 2.2 mm to the left and 0.55 mm below. DETAILED DESCRIPTION OF THE INVENTION

[0026] In this disclosure, the optical design of a light collection subsystem for a spectral flow cytometer is optimized to address the following issues: (a) the complex design and potential reduced light collection efficiency and increased cost associated with using an array of optical fibers, each guiding light from one laser excitation to a corresponding subsystem for light splitting and light detection, (b) the large signal coefficient of variation (CV) associated with high sample flow rates for high-throughput cellular analysis in existing systems, especially current spectral flow cytometers with fiber optic collection of fluorescence excited by different lasers, (c) the high cost and complex light collection subsystem design associated with using multiple optical filter sets to split fluorescence from different lasers, and (d) the poor resolution of the spectral wavelength range in individual fluorescence detection channels and the high cost of using an array of light-dispersing optics.

[0027] At least some of the light collection subsystems described herein include a novel light collection optics subsystem and an innovative set of light-splitting (transmission and reflection) filters. To collect light, a customized objective lens is designed with a high numerical aperture (e.g., greater than 1.2) and a high magnification factor (e.g., >30x). A 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 fluorescence emitted from n lasers (n is the number of lasers in the system) into separate light spots on a vertical plane. Each focused spot is separated by a distance long enough to be detected using an array of n separate detectors without any mechanical interference. That is, the magnification factor of the objective lens must be large enough to allow for a large enough distance between the focused spots of the individual lasers. On the other hand, a magnification factor that is too large is undesirable for a compact system because it increases the footprint and complexity of the system.

[0028] To split the light, an optical filter set is designed to split the fluorescence excited by all lasers in the system. The optical filter set's arrangement exploits the differences in focal lengths at different wavelengths across the entire spectral range of the collected fluorescence, causing light of different wavelengths to propagate along optical paths of different lengths. At the end of each optical path for each distinct narrow wavelength range of fluorescence is the corresponding focal plane, where a detection module is located. The detection module includes an array of spherical lenses followed by a corresponding array of SiPM detectors. Each spherical lens is positioned at the focal point of the fluorescence excited by the laser beam in the flow cell, ensuring that the light forms a uniform distribution and covers the largest possible area on the SiPM detector. The lenses in front of the SiPM photodetectors allow the SiPMs to operate at their best performance conditions, significantly reducing the variation in the detected signal, allowing very good signal CVs to be achieved for each fluorescence detection channel, even at high sample flow rates.

[0029] The optical subsystem described herein satisfies one or more of the following objectives: 1) a large numerical aperture of the collection system to maximize light collection; 2) a properly designed collection objective lens that optimizes / balances the desired small footprint of the device (using a smaller magnification factor) with the objective that the images of light from each excitation laser are sufficiently separated (in the millimeter range) at the focal plane so that an array of individual photodetectors can be physically placed to detect light from each excitation laser; 3) a collection objective lens designed with a high magnification factor to allow sufficient length for the light expansion path along which optical filters can be placed to split the fluorescence into a continuous spectrum of different narrow wavelength ranges; and 4) a spherical lens placed at the focused spot of fluorescence collected from each excitation laser after the light passes through a bandpass filter in front of 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 closely matches the effective detection area of ​​the photodetector. For example, if the effective photodetector area is 3 mm x 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.

[0030] While sophisticated design of the focusing objective lens is challenging, the development and placement of the light-splitting filter set requires careful engineering development and innovation. The selection and placement of the individual bandpass filter sets is performed so that the collected light first passes through one bandpass filter positioned at a specific angle relative to the incident light and is detected by a photodetector for one channel, while the remainder of the light is reflected to the next bandpass filter. Thus, the collected side-scattered light and fluorescence light pass through the individual bandpass filter sets in series and are detected by a photodetector array after each bandpass filter. The system innovatively designs the path of light passing through the bandpass filter sets, detecting light for specific wavelength ranges determined by the bandpass filters. The long focal length allows the scattered light and fluorescence light to enter each bandpass filter at a small half-cone angle and be focused behind the bandpass filter. The long focal length allows light to propagate far enough in free space to accommodate a series of bandpass filters.

[0031] This configuration offers several advantages: 1) it is a free-space optical system and does not use optical fibers. The absence of optical fibers results in higher light collection efficiency, greater stability of the collection optics, and improved operability. 2) only one set of optical filters is required to split the light excited by all lasers. Using only one set of filters instead of N sets simplifies the detection system and significantly reduces product costs. Utilizing the same optical filters to split the light excited by different lasers offers the surprising advantage that the fluorescence is split into exactly the same wavelength range, regardless of the laser that triggered the fluorescence. This is particularly important for fluorescent molecules excited by multiple lasers, when reference fluorescence spectra from different lasers are acquired. Splitting the fluorescence from different lasers into exactly the same wavelength range with the same filters can be useful for obtaining consistent or similar reference fluorescence spectra from different lasers and can also be useful for deconvolving measured fluorescence spectra collected for samples stained with multiple fluorophores to calculate the fluorescence signal intensity of each fluorophore used in the sample using the reference spectral data from each individual fluorophore. 3) The filter set splits the light into individual narrow wavelength ranges (i.e., one fluorescence detection channel) across the entire detection spectrum in free space. This significantly reduces light loss and improves overall light collection efficiency. 4) This system allows the use of solid-state photodetectors (i.e., SiPMs and APDs) with good photon detection performance and reduced cost. This design allows spectral flow cytometer systems to be designed with exceptional performance while remaining very simple and significantly reducing cost. 5) Importantly, the spherical lens offers significant advantages in providing optimal CV of the detected optical signal at high sample flow rates for high-throughput cellular analysis. As mentioned above, the focused spot of 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 within the flow channel experiences horizontal fluctuations as the sample fluid flows through the channel.Therefore, the horizontal position of the focused spot of light from each excitation laser changes on the first surface of each corresponding spherical lens. A larger sample flow rate results in a larger fluid fluctuation core diameter, which means a larger horizontal sample distribution width. As an example, if the fluid core diameter is 60 μm and the magnification of the collection optical system is 55X, the focused spot of light collected from each excitation laser has a horizontal spread of 3.3 mm on the first surface of each corresponding spherical lens. If the diameter of the spherical lens is larger than 3.3 mm, light from all samples can be completely collected without any loss. After passing through the spherical lens, the fluorescence from each excitation laser maintains the same position on the photodetector chip surface. Therefore, the detected light signal is hardly affected by fluctuations from the sample. The size of the light on the surface of the photodetector chip can be designed to match the active area of ​​the photodetector chip by appropriately designing the spherical lens. The detected light spot on the photodetector is not affected by the change in the position of the collected light focused on the first surface of the spherical lens. That is, the detected light signal is not affected by fluctuations in the horizontal position of the sample inside the flow channel. The same principle applies to changes in the vertical position of the detected light. Even if the focused laser beam inside the flow cell changes from its original vertical position, the detected light signal can always remain stable as long as the fluorescence from this laser is fully collected by the spherical lens in front of the photodetector. This can be achieved by using a large-diameter spherical lens. In summary, when a spherical lens is placed between the focused spots of each fluorescence from each excitation laser and in front of the photodetector, it has the unique advantage of obtaining a circular light spot with uniform energy distribution and a spot size that 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 with fluctuations in the position of the original light source (i.e., the fluorescence emitted from each cell or particle inside the flow channel).

[0032] In summary, the same set of optical filters is used for all available lasers, significantly reducing system complexity and system cost. The collection optics can be designed with a high numerical aperture (>1.2) to maximize light collection from each laser, resulting in significantly improved fluorescence signal detection sensitivity. In addition, the system has a free-space design, minimizing light loss and reducing the need for frequent servicing, which is typically a problem when using optical fibers. Because the light from each excitation laser is sufficiently spatially separated in the focused spot, arrays of photodetectors, such as silicon photomultipliers (SiPMs) and avalanche photodiodes (APDs), can be used as detectors. These solid-state photodetectors have excellent photodetection performance (i.e., high photon detection efficiency, high gain, low dark count, and low operating voltage), and are also compact, robust, and low-cost.

[0033] The foregoing technical improvements, advantages, and benefits will now be described in a series of non-limiting detailed embodiments with reference to the drawings, where it should be understood that like reference numerals are used to identify like components shown in one or more of the figures.

[0034] Beginning with FIG. 1 , an optical system 10, a flow cytometer 100 incorporating the optical system 10, and a method for measuring light are provided. These are particularly useful for determining the physical and chemical properties of cells labeled with fluorescent markers. The term “cell,” as used herein, refers to a biological cell, which is capable of living independently and is the smallest unit that constitutes all organisms and tissues in the body. A cell has three main parts: a cell membrane, a nucleus, and a cytoplasm. The system 10 and method are particularly useful for profiling cells through the measurement of light emitted from fluorescent labels that target specific cellular biomarkers, such as cell surface antigens or internal molecules. From the light measurements, the system 10 and method can reveal cellular characteristics, such as cell size, cell granularity, DNA content, DNA gene expression, surface receptors, intracellular proteins, and temporal signaling. Furthermore, from these characteristics, cell profiles can be generated according to function or activity, e.g., immune infection-fighting cells (e.g., B cells, T cells), state (e.g., cancerous compared to normal), stage (e.g., mature compared to immature), origin (e.g., transgenic compared to non-transgenic), and other categories known in the art. Additionally, compounds such as potential therapeutic agents can also be characterized or profiled based on measurable changes in treated cell populations using the system 10 and methods described herein.

[0035] As described in more detail in the following paragraphs, characterization and profiling of cells can be achieved by measuring light emitted by the cells themselves (which can be used to determine their size and granularity) and / or by measuring the emission from moieties attached to or inserted into the cells (e.g., fluorescently labeled antibodies, antibody fragments, or other molecules that attach to or bind to cells and fluoresce when excited). By assessing the presence, absence, and / or abundance of these moieties, system 10 and methods can characterize and profile cells and compounds that affect cells on a single-cell basis. Non-limiting examples of such are provided throughout this disclosure.

[0036] The optical measurements described herein are preferably performed on a cell-by-cell basis, so that data is collected and analyzed from individual cells. Furthermore, measurements are very fast, e.g., at rates of approximately 10,000 cells / minute. Turning to FIG. 2 , one embodiment of a flow cytometer 100 incorporating an optical system 10 according to the present technology is shown. This is achieved, in part, by rapidly passing cells through an elongated flow cell 12, whose inner diameter D is slightly larger than the cells themselves. Samples / cells 30 can be passed through the flow cell 12 by incorporating the flow cell 12 into a fluidic system that, on the one hand, delivers cells 30 to the flow cell 12 and, on the other hand, removes cells from the flow cell 12. This can be achieved using any suitable fluid pumping mechanism known in the art, such as those used in flow cytometry systems to pump sheath fluid 14 into and out of the flow cell 12. In some embodiments, the optical system 10 itself and its components constitute part of the flow cytometer 100.

[0037] As cells pass through the flow cell 12, the excitation light source 16 directs light to different locations 18a-18e along the length L of the flow cell 12. This light is provided at a wavelength that causes certain moieties to fluoresce. For example, the excitation light sources 16a-16e direct light of specific wavelengths to the different locations 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 the different locations 18a-18e along the length L of the flow cell 12. As the cells pass through the different locations 18a-18e, fluorescent labels attached to the cells can fluoresce, and this fluorescence is collected, resolved, measured, and analyzed for each of the different locations 18a-18e along the length L of the flow cell 12.

[0038] In a preferred embodiment, as detailed in FIG. 2, a single excitation light source 16 directs light to a single location 18a-18e along the length L of the flow cell 12, thereby selectively pairing each different location 18a-18e along the length L of the flow cell 12 with its own excitation light source 16. For example, a first location 18a along the length L of the flow cell 12 receives light from the light source 16a at a first wavelength (e.g., λ a =349 nm), and the second location 18b can receive light of a second wavelength (e.g., λ = 349 nm) from the light source 16b. b = 405 nm), and a third location 18c can receive light of a third wavelength (e.g., λ c =488 nm), and a fourth location 18d can receive light of a fourth wavelength (e.g., λ d = 561 nm), and a fifth location 18e can receive light of a fifth wavelength (e.g., λ e2, the system 10 and method include five excitation light sources 16a-16e. However, in other embodiments, the system 10 and method include any number n of excitation light sources 16. In some embodiments, the system 10 and method include n=7 or more than seven excitation light sources 16. In some embodiments, n=10 or more than ten excitation light sources 16. In some embodiments, the system 10 and method include n=15 or more than fifteen excitation light sources 16.

[0039] In some embodiments, the excitation light source 16 can direct light to different locations 18a-18e along the length L of the flow cell 12 through a fiber optic cable. In other embodiments, the excitation light source 16 is a laser aimed directly at different locations 18a-18e of the flow cell 12. In still other embodiments, light emitted from the laser is directed to different locations 18a-18e along the length of the flow cell 16, for example, through the use of alignment mirrors. In each of the above configurations, the excitation light source 16 can direct light through beam-shaping optics 20, which change the initial shape of the light beam from circular or near-circular to one or more designed shapes. These designed shapes can take into account the following: On the one hand, the laser beam needs to be shaped to be wide enough to span the width of the flow cell 12, thereby encompassing the entire inner diameter D; and on the other hand, the laser beam needs to be shaped to minimize its height in order to shorten the distance between adjacent locations 18a-18e along the length L of the flow cell 12. In some embodiments, these engineered shapes tend to be elliptical rather than circular, and tend to be aligned such that the major axis of the elliptical shape extends across the inner diameter D of the flow cell 12 rather than along the length L of the flow cell 12. However, in other embodiments, an elliptical beam shape is formed such that the major axis of the ellipse extends along the length L of the flow cell 12.

[0040] The systems and methods resolve fluorescent signals emitted between adjacent locations 18a-18e (that are very close to one another, e.g., in the micron range) along the length L of the flow cell 12. In some embodiments, the systems 10 and methods resolve fluorescent signals between adjacent locations 18a-18e along the length L of the flow cell 12 (where the separation distance 22 is between 50 μm and 300 μm). In some embodiments, fluorescent signals are resolved from adjacent locations 18a-18e along the length L of the flow cell 12 (where the separation distance 22 is between 60 μm and 250 μm). In some embodiments, fluorescent signals are resolved from adjacent locations 18a-18e along the length L of the flow cell 12 (where the separation distance 22 is between 70 μm and 150 μm). In some embodiments, fluorescent signals are resolved from adjacent locations 18a-18e along the length L of the flow cell 12 (where the separation distance 22 is between 75 μm and 100 μm). In some embodiments, the fluorescent signals are measured at adjacent locations 18a-18e along the length L of the flow cell 12 (where the separation distance 22 is between 50 μm and 55 μm, 55 μm and 60 μm, 60 μm and 65 μm, 65 μm and 70 μm, 70 μm and 75 μm, 75 μm and 80 μm, 80 μm and 85 μm, 85 μm and 90 μm, 90 μm and 95 μm, 95 μm and 100 μm, 100 μm and 105 μm, 105 μm and 110 μm). 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 of adjacent positions 18 a-18 e along the length L of flow cell 12 is the same for all adjacent positions 18 a-18 e. That is, the separation distance 22 between adjacent first and second locations 18a, 18b as cells flow along the length L of flow cell 12 is preferably the same as the separation distance 22 between adjacent second and third locations 18b, 18c as cells continue to flow along the length L of flow cell 12, and so on. However, the invention is not so limited, and thus, in other embodiments, adjacent separation distances 22 are not the same.

[0041] As cells 30 flow along the length L of flow cell 12, they are exposed to light from multiple excitation light sources 16, causing them to emit light. That is, the cells and fluorescently labeled molecules paired with the cells emit light when exposed to light of appropriate wavelengths targeted to different locations 18a-18e of flow cell 12. The principles of light emission and fluorescence are well known in the art and are therefore omitted for the sake of brevity.

[0042] Light emitted from different locations 18a-18e along the length L of flow cell 12 is collected and magnified using shared collection optics 24 and shared magnification optics 26. By "shared collection optics" and "shared magnification optics," it is meant that light emitted from two or more, and preferably all, of the different locations 18a-18e along the length L of flow cell 12 being measured is collected using the same collection lens or lenses and magnified using the same magnification lens or lenses. Collection lenses with high numerical apertures (e.g., >1.2) maximize the collection of light from each of the different locations 18a-18e along the length L of flow cell 12, significantly improving detection sensitivity.

[0043] An exemplary configuration is shown in FIG. 2, in which the focusing objective 32 is integrated with both the shared focusing optics 24 and the shared magnification optics 26. Also shown in FIG. 2, the focusing optics 24, embodied as a hemispherical lens 34, is adhered to the surface of the flow cell 12 using an optical adhesive and passes emission light from the flow cell 12 into the focusing objective 32. Preferably, this hemispherical lens 34 is not adhered to the focusing objective 32 so that the focusing objective 32 can move along three different axes to compensate for variations in light emission position due to differences in fluid dynamics and the alignment of the excitation light source 16. Upon entering the focusing objective 32, the emission light passes through a set of lenses, some of which are converging and others diverging, to achieve the desired magnification and focal length. Roughly speaking, converging lenses are thicker in the center and tend to direct light rays passing through the lens toward a focal point, while diverging lenses are thinner in the center and have a negative focal length, so that light rays exiting the lens tend to diverge away from each other. By combining different lenses in series, the magnification of the focusing objective lens 32 can be between 20x and 250x. In some embodiments, the magnification of the focusing objective lens 32 is between 30x and 100x. In some embodiments, the magnification of the focusing objective lens 32 is between 40x and 60x. In some embodiments, the magnification of the focusing objective lens 32 is 40x, 41x, 42x, 43x, 44x, 45x, 46x, 47x, 48x, 49x, 50x, 51x, 52x, 53x, 54x, 55x, 56x, 57x, 58x, 59x, or 60x. In some embodiments, when light enters the focusing objective lens 32, it passes through two concave-convex lenses, followed by a hemispherical lens, and then a biconvex lens, resulting in a magnification of between 40x and 60x.

[0044] In some embodiments, the objective lens 32 can be adjusted in 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.

[0045] While one purpose is to expand all of the light emitted from the flow cell 12 for subsequent splitting and measurement, it is also important to ensure that the light emitted from each of the different locations 18a-18e along the length of the flow cell 12 remains distinguishable from the other locations. This can be achieved by combining a diverging lens (which again spreads the light outward) with one or more collimating lenses (which narrow the cross-section of each beam and redirect the light along the same path). Examples of collimating lenses are known in the art and include various plano-convex lenses.

[0046] Light emitted from different locations 18a-18e along the length L of the flow cell 12 passes through the shared collection optics 24 and the shared magnification optics 26 of the collection objective 32 to further optics (e.g., a reflective mirror 42) and a light-splitting module 38. The light-splitting module 38 includes light-splitting optics configured to split the magnified light into multiple different wavelength ranges (e.g., 40A-40S) before receiving it from the detection module 50. The light-splitting module and the detection module are more clearly shown in FIG. 3 and house different channels 40A-40S (collectively referred to as an "array of channels") for measurement. In the case of the spectral flow cytometer of the present invention, such different fluorescent channels having different wavelength ranges are actually contiguous; i.e., when counting fluorescent channels from a short wavelength range to a long wavelength range, for two adjacent channels, the longer wavelength end of the short wavelength 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 as follows: 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 several wavelength gaps corresponding to the laser wavelength. the gap is 389nm to 417nm for the 405nm laser, 477nm to 500nm for the 488nm laser, 552nm to 573nm for the 562nm laser, and 628nm to 650nm for the 637nm laser. Proper alignment of the focusing objective lens 32 with the light splitting module 38 is ensured by a pair of reflecting mirrors 42 which can be adjusted in different directions.

[0047] In a preferred embodiment, the optical splitting module 38 includes a short-pass or long-pass filter 44, which initially splits the received light into one of two distinct propagation paths 46A and 46B through the optical splitting module 38. It is understood that the filter 44 can be configured based on the number m of desired channels 40. Based on the channel configuration shown in FIG. 3, where M=19 channels (40A-40S), two propagation paths 46A and 46B are implemented. However, it is within the scope of the present invention for all light to travel along the same propagation path or for light to be repeatedly split to pass between three or more propagation paths. Nevertheless, a series of band-pass filters 48 are positioned along each propagation path 46A, 46B. Each band-pass filter 48 is designed to allow the passage of light having a specific wavelength and reject the passage of the remaining wavelengths. These rejected wavelengths are attenuated along the propagation paths 46A, 46B through the optical splitting module 38, thus encountering the next band-pass filter 48, and so on. Therefore, alignment of the bandpass filters 48 with respect to the incident light is critical to ensuring that rejected wavelengths continue to propagate through the light splitting module 38. By way of non-limiting example, it has been found that continuous propagation can be achieved by positioning the bandpass filters 48 at angles of 10° to 20°. However, it is more preferable to position each bandpass filter 48 at an angle of 12° to 16° with respect to the incident light, with angles of 14° or 15° with respect to the incident light being currently most preferred. In this manner, by propagating the expanded light from each of different locations 18a-18e of the flow cell 12, the emission light can be separated into multiple different detection channels 40A-40S (each corresponding to a specific wavelength range). Furthermore, because the collection optics 24 and expansion optics 26 are shared across different locations 18a-18e along the length L of the flow cell 12, the emission from these different locations 18a-18e can be propagated in the same direction, and therefore through the same set of bandpass filters 48.In this way, the bandpass filter 48 can be shared by light emitted from each of the different locations 18a-18e of the flow cell 12, which itself reduces the footprint of the light splitting module 38.

[0048] Light passing through each of the different bandpass filters 48 is received by an array of light detection modules 50. In the embodiment shown in FIG. 3, a dedicated light detection module (50A-50S) is 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 light splitting module 38. As shown in FIGS. 4 and 5, each light detection module 50 includes or is coupled to a dedicated array of spherical lenses 54, which then directs the magnified light from each of the different locations 18a-18e along the length L of the flow cell 12 to a predetermined light detector array 52 for measurement. That is, the array of spherical lenses 54 is assigned to each of the channels 40A-40S defined by the corresponding bandpass filter 48 and is positioned along a propagation path that intersects the bandpass filter 48 at a specific distance determined by the wavelength and optics, and each of the spherical lenses 54a-54e in the same array 54 (shown in FIG. 4) is configured to receive light of the same wavelength range. However, each of the spherical lenses 54a-54e in the array 54 emits light from a different excitation light source 16a-16e (corresponding assigned wavelength (λ a , λ b , λ c λ d λ e ) and positions 18a-18e along the length L of flow cell 12 (see FIG. 2). That is, spherical lens array 54 not only receives light of a certain wavelength range (depending on the channel) and directs it to detector 52 for measurement, but also distinguishes between the light emitted from each of the different positions 18a-18e along the length L of flow cell 12 for measurement.

[0049] The positioning of the spherical lens array 54, and therefore the specific distance, is determined based on the focal length of the light passing through the corresponding bandpass filter 48. That is, each of the spherical lenses 54 in the same array is positioned 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-40S, which is itself determined by the focal lengths of the shared focusing optics 24, the shared magnifying optics 26, and the bandpass filter 48. That is, the array 54 of spherical lenses 54a-54e is positioned at focal points 58a-58e on the plane 56 for each channel 40A-40S. An example is shown in FIG. 6, which is a table summarizing exemplary focal lengths f of the 19 channels (e.g., channels 40A-40S) and the positioning of the 19 arrays 54 of spherical lenses (54a-54e).

[0050] For example, the first listed channel 1 in FIG. 6 corresponds to channel 40A shown in FIG. 3 and has a focal length f A = 692.17 mm. The spherical lens array 54 assigned to the light detection module 50A has a focal length f A = 692.17 mm. M Channels and 50 M The detection module and corresponding M For the system of Figures 2 and 3 with M = 19 channels and n = 5 excitation sources, a total of M x n (19 x 5) = 95 detectors and spherical lenses are used.

[0051] In the example given above, if the focusing optics has a magnification of 55X and a separation distance 22 of 100 μm (FIG. 3), then the vertical separation distance 60 between two adjacent focused light spots (focal points 58a-58e) is 5.5 mm. Thus, spherical lenses 54a-54e are aligned so that they are vertically spaced apart from one another by 5.5 mm (CC) in focal plane 56.

[0052] 1-7C collectively, each of the spherical lenses 54a-54e in a single array 54 of spherical lenses are aligned along the same plane 56 as each other, but the array 54 of spherical lenses also distinguishes between light emitted from different locations 18a-18e along the length L of the flow cell 12 by assigning each spherical lens 54a-54e in the array 50 to a different focal point 58a-58e within the focal plane 56. Each of these focal points 58a-58e within the focal plane 56 is the result of collecting and magnifying light emitted from a different location 18a-18e along the length L of the flow cell 12. As more clearly shown in FIGS. 7A-7C , resolving fluorescence emitted from different locations 18a-18e along the length of the flow cell involves increasing the separation distance between adjacent locations 18a-18e. For example, light may be emitted from adjacent locations 18a-18e that are spaced 100 μm apart (see FIG. 7A ). Resolving such a short distance is difficult. However, by propagating the emission light through an array of different lenses (see FIG. 7B), the separation distance can be significantly increased (FIG. 7C shows 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 and transmit signals from each of the different locations 18a-18e along the length of the flow cell to the photodetector 52. That is, the emission light from each of the different locations 18a-18e of the flow cell 12 is ultimately focused as a spot of light at a different focal point 58a-58e along the focal plane 56, respectively. By positioning the spherical lenses 54a-54e at those points 58a-58e, the emission light can then be directed to a predetermined area of ​​a dedicated photodetector 52a-52e for measurement. Directing the emission light to a specific focal point 58a-58e along the focal plane 56 is achieved by maintaining the identity or source of the light throughout the light splitting module 38. As shown more clearly in FIG. 4, the flow cell 12 is normally arranged in a vertical direction, and therefore the spherical lenses 54a to 54e are also normally arranged in a vertical direction and offset from one another.In some embodiments, the vertical separation distance 60 between adjacent beams (i.e., adjacent focal points 58a-58e) in the array of spherical lenses 50 is less than 30 mm. In some embodiments, the vertical separation distance 60 between adjacent focal points 58a-58e is between 1 mm and 10 mm. In some embodiments, each of the vertical separation distances 60 between adjacent focal points 58a-58e is 5.5 mm.

[0053] Each spherical lens 54a-54e in each spherical lens array 54 directs received light to a different photodetector 52a-52e for measurement, thereby enabling measurements based on the emission wavelength and emission location from the flow cell 12. Among the challenges encountered in delivering light from the flow cell 12 to the detectors 52 is miniaturization, which is due, at least in part, to uneven flow or swaying of cells as they pass through the flow cell 12. Inconsistencies were initially found when cells passed through the flow cell 12 due to variations in fluid properties and the positioning of the excitation light source 16. These inconsistencies were magnified throughout the system, posing significant challenges in manufacturing a flow cytometer 100 with a small footprint. This problem is partially solved by using spherical lenses 54a-54e in the spherical lens array 54, which corrects for inconsistencies in optical alignment to the appropriate photodetectors 52a-52e. Mismatch correction is experimentally demonstrated in Figures 8A-8D. For reference throughout Figures 8A-8D, Figure 8A shows the end portion of fluorescent light propagating through a system where the propagating light is directed to detector 52 by spherical lens 54 located at focal plane 56. Figure 8B shows the results from an ideal alignment where the propagating light strikes spherical lens 54 at the center and is therefore directed to the predetermined area or active area of ​​photodetector 52. Again, due to various factors, in some cases, the light will strike spherical lens 54 off-center. Figure 8C shows the results when the light strikes spherical lens 54 1.65 mm off-center horizontally (to the right). The radius of spherical lens 54 allows the light to be redirected to the predetermined area or active area of ​​photodetector 52. Figure 8D shows the results when the light strikes spherical lens 54 2.2 mm off-center horizontally (to the left) and 0.55 mm off-center vertically (down). In this case, spherical lens 54 corrects the propagation path so that the light again strikes a predetermined or effective area of ​​detector 52. This correction or redirection when light strikes spherical lens 54 off-center is accomplished at least in part by the radius of spherical lens 54.8A-8D, each of the spherical lenses 56 has a radius of 4.1 mm to direct or redirect light to a predetermined or effective area of ​​less than 3 mm x 3 mm. In light of the above, one skilled in the art will appreciate that the radius of the spherical lenses 54 can vary depending, at least in part, on the predetermined or effective area of ​​the detector 52 and the spacing difference of the propagating light. Thus, while the radius is not limited to any particular size, as a general guidance to begin with, the radius can be between 1 mm and 20 mm, e.g., 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 therebetween.

[0054] The photodetectors 52a-52e (collectively referred to as an "array of photodetectors"), which may include detectors such as silicon photomultipliers (SiPMs), avalanche photodiodes (APDs), etc., are collected into M photodetection modules 50 (each containing an array of n photodetectors), with one photodetector 52a-52e in each photodetection module 50A-50S measuring light corresponding to a single wavelength range within a channel 40A-40S from one of different locations 18a-18e along the length L of the flow cell 12. By measuring light from each of the different locations 18a-18e along the length L of the flow cell 12 in each channel 40A-40S, the system 10 and method can identify the presence, absence, and / or abundance of fluorescent markers and thus determine cell moieties used to characterize the cells.

[0055] The foregoing description of various embodiments of the disclosed technology has been presented for purposes of illustration and description. It is not intended to limit the disclosed technology to the precise form disclosed. Many modifications, variations, and improvements will be apparent to those skilled in the art. For example, embodiments of the apparatus and associated optics can be sized and otherwise adapted for various detector and / or flow cell configurations. Moreover, those skilled in the art will recognize, or be able to ascertain using no more than routine experimentation, many equivalents to the specific apparatus and methods described herein. Such equivalents are considered to be within the scope of the disclosed technology and are covered by the following appended claims.

[0056] Elements, features, or acts from one embodiment can be readily combined or substituted with one or more elements, features, or acts from other embodiments to form many additional embodiments within the scope of the technology of the present disclosure. Moreover, elements shown or described as being combined with other elements can, in various embodiments, exist as independent elements. Accordingly, the scope of the technology of the present disclosure is not limited to the details of the described embodiments, but instead is limited only by the scope of the appended claims.

Claims

1. 1. An optical system for measuring light, comprising: a) a flow cell having a length and width that allows the passage of biological cells; b) a plurality of different excitation light sources directing light at different locations along the length of the flow cell; c) shared collection and magnification optics configured to collect light emitted from the different locations of the flow cell, and to magnify and focus the collected light onto a focal plane at a specific distance according to different wavelength ranges and to different locations along the focal plane according to each of the different locations of the flow cell from which light was emitted; d) a light splitting module including a light splitting optics arranged such that the expanded light emitted from each of the different positions of the flow cell shares the light splitting optics, the light splitting optics configured to split the expanded light emitted from each of the different positions of the flow cell into an array of channels, each channel corresponding to a distinct wavelength range; and e) an array of detection modules, each detection module dedicated to an individual channel in said array of channels, each detection module comprising: i) an array of spherical lenses arranged at each of the different focal lengths along a propagation direction of the collected light for each of the different wavelength ranges, each of the spherical lenses in the array configured to receive the magnified light emitted from one of the different positions of the flow cell and direct the received magnified light to a predetermined area on a corresponding photodetector; ii) an array of photodetectors configured to measure the expanded light directed from each of the spherical lenses, whereby one photodetector measures the collected light within one wavelength range from one of the different locations of the flow cell; and an array of detection modules, an optical system comprising:

2. The system of claim 1 , wherein the plurality of excitation light sources comprises a plurality of laser light sources.

3. The system of claim 2 , wherein the laser light source is shaped by beam shaping optics into a designed shape and focused to different positions inside the flow cell.

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

5. 10. The system of claim 1, wherein the distance between adjacent locations on the flow cell to which light is directed from the excitation light source is between 25 μm and 250 μm, optionally between 50 μm and 150 μm.

6. 10. The system of claim 1, wherein the shared collection optics directs light to the shared magnification optics, and the shared magnification optics increases the vertical separation distance between light emitted from the different locations on the flow cell.

7. 7. The system of claim 6, wherein the shared magnification optics comprises a set of objective lenses comprising magnifications between 20x and 250x, optionally between 30x and 100x, optionally between 35x and 75x, optionally between 40x and 60x.

8. The system of claim 6 , wherein the vertical separation distance between adjacent lights in the array of spherical lenses is less than 30 mm.

9. The system of claim 8 , wherein the vertical separation distance for the array of spherical lenses is between 1 mm and 10 mm.

10. 10. The system of claim 9, wherein the multiple excitation light sources direct light to adjacent positions 100 μm apart along the flow cell, the shared magnification optics includes a 55X magnification, and adjacent light received at the array of spherical lenses is 5.5 mm apart.

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

12. The system of claim 1 , wherein the plurality of wavelength ranges includes a wavelength range from 320 nm to 1000 nm.

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

14. 13. The system of 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 of claim 1 , wherein the light emitted from the different locations fluctuates in response to changing fluid properties through the flow path and / or changes in positioning of an excitation light source.

16. 16. The system of claim 15, wherein the spherical lens with each array directs fluctuating light to the same area of ​​the photodetector where the detected light is stable.

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

18. 10. The system of claim 1, wherein each spherical lens is specially designed so that the diameter of the light spot on the photodetector matches the effective detection area of ​​the detector.

19. 10. The system of claim 1, wherein the collection optics and spherical mirror are specified such that the detected light spot on the corresponding photodetector is unaffected by changes in position of the collected light focused onto the spherical lens.

20. The system of claim 1 , wherein the spherical lenses are designed such that the detected optical signals on the corresponding optical detectors are not affected by lateral positional variations of the biological cells within the flow channel.

21. The system of claim 1 , further comprising a set alignment optics configured to align the expanded light onto the light splitting module and the array of spherical lenses.

22. A flow cytometer comprising the system of claim 1.

23. 1. A method for characterizing a biological cell, comprising: a) passing biological cells through a flow cell, the biological cells being labeled with multiple fluorescent markers that fluoresce at different wavelengths when excited, the markers being attached or absent to cellular moieties that together characterize the biological cells by the presence, absence, or abundance of the markers; b) directing light to different locations along the length of the flow cell as the biological cells pass through the different locations; c) collecting light emitted from each of the different locations of the flow cell as the biological cell passes through; d) expanding and focusing the collected light to different focal lengths according to different wavelength ranges and to different positions within each focal length according to each of the different positions of the flow cell from which the collected light was emitted; e) dividing the expanded light into a plurality of different wavelength ranges, wherein the different fluorescent wavelengths are divided into the plurality of different wavelength ranges; f) receiving light of the same wavelength range emitted from the different locations along the length of the flow cell with the same array of spherical lenses, different spherical lenses in each array receiving light emitted from the different locations of the flow cell; g) directing light from each of said spherical lenses to a different photodetector; h) measuring light at each of the different photodetectors, thereby identifying the presence, absence, or abundance of the marker and thus identifying a cellular portion that characterizes the biological cell; A method comprising:

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