Single photon avalanche diode implementation method and system for flow cytometry

Abstract

The invention discloses a single photon avalanche diode implementation method and system for flow cytometry. A computer-implemented method: receiving one or more optical signals at a point in time, wherein each optical signal is captured by a respective pixel of an SPAD array; and measuring the number of activated pixels of the SPAD array at the time point, and adjusting the count of the optical signals captured by the SPAD detector at the time point according to the number of the pixels.

Description

Technical Field

[0001] The present invention mainly describes the fluorescence detection method, with an emphasis on single photon avalanche diode (SPAD). Background Art

[0002] Flow cytometry systems typically require optical detectors with a wide dynamic range, spanning at least five orders of magnitude, and capable of detecting light signal intensities as low as a few photons. These light pulses may last as little as 0.5μs, requiring high-bandwidth devices to accurately measure the pulse characteristics. Historically, such applications have typically required photomultiplier tubes (PMTs), which are well-suited for most instrument platforms. However, PMTs capable of detecting photon wavelengths exceeding 750nm are expensive.

[0003] Although several new technologies have been invented to replace PMTs, their performance remains unsatisfactory. Multi-pixel photon counters (MPPCs), also known as silicon photomultipliers (SiPMs), offer excellent gain and bandwidth performance, but their signal-to-noise ratio is much lower than that of PMTs and they are subject to linearity limitations. Avalanche photodiodes (APDs) offer good photon detection efficiency, but their gain is extremely low, resulting in a decrease in bandwidth and similarly limited linearity. SPAD direct photon counting technology can address these issues, but the response speed of the best SPADs lags behind by several orders of magnitude, making it impossible to achieve accurate counting across the dynamic range within the required bandwidth.

[0004] It is worth emphasizing that new flow cytometers are equipped with advanced detectors to meet the needs of high-throughput multiplexed experiments. The design of such system architectures continues to evolve to meet two major trends: (a) fluorescence detectors that receive information from a single laser source; and (b) laser sources that interact with samples at spatially separated locations. If fluorescence detectors that receive signals from a single laser source are used, the space occupied by the detectors increases as their number increases. The wide-spectrum fluorescence signal from a single laser source can be effectively distributed through a linear array multi-detector configuration, and this design has gradually become a mainstream instrument design method. However, this design method also has its shortcomings. (1) If a diffraction grating or prism is used as a spectroscopic element, the spacing between these detectors is crucial to determining the wavelength center and bandwidth of each channel. (2) Independent detectors can set gain and peak wavelength sensitivity, while detector arrays are not good at setting gain and material composition, making it impossible to optimize the signals of all channels at the same time. The space occupied by the detectors can be reduced by having a single detector receive light signals from multiple detection points at the same time. This technique assigns signals to corresponding laser sources based on their arrival times, but has the following drawbacks: parameter settings must balance the parameters of the signals from all laser sources, which can lead to improper gain settings (saturation or insufficient gain). The sample concentration must be low to avoid signal overlap, which occurs when different cells are simultaneously exposed to different laser sources, resulting in overlapping signals received by a single detector. Summary of the Invention

[0005] This invention describes a method and system for implementing single-photon avalanche diodes (SPADs) for flow cytometry. A computer-implemented method comprises receiving one or more light signals at a given point in time, each captured by a corresponding pixel of a SPAD array; measuring the number of activated pixels of the SPAD array at that point in time; and adjusting the light signal count captured by the SPAD detector at that point in time based on this number of pixels. Multiple SPADs can be integrated into a single device, with the spacing between the detectors determined by the distribution and density of detection points in the fluid channel. BRIEF DESCRIPTION OF THE DRAWINGS

[0006] This patent or patent application publication document is provided in color. A color copy of this patent or patent application publication document shall be provided upon payment of the prescribed fee and approval by the Office.

[0007] For the purpose of illustrating the invention, the drawings show the best mode of carrying out the invention, but it is to be emphasized that the invention is not limited to the precise arrangements shown.

[0008] Figure 1 A flow cytometry system according to the present invention is shown.

[0009] Figure 2 A SPAD array according to the present invention is shown.

[0010] Figure 3 FIG. 2 shows a detector equipped with the SPAD according to the present invention.

[0011] Figure 4 A curve diagram of the simulated detection count of the SPAD array described in the present invention is shown.

[0012] Figure 5 A curve diagram of the simulated detection count of the SPAD array described in the present invention is shown.

[0013] Figure 6 A curve diagram of the simulated detection count of the SPAD array described in the present invention is shown.

[0014] Figure 7 A curve diagram of the simulated detection count of the SPAD array described in the present invention is shown.

[0015] Figure 8 A graph showing the photon detection efficiency (PDE) of the detector within different wavelength ranges according to the present invention is shown.

[0016] Figure 9 The graph shows the fluorescence band curve of the flow cytometer of the present invention.

[0017] Figure 10 The graph shows the fluorescence band curve of the flow cytometer of the present invention.

[0018] Figure 11A -D shows the fluorescence band curve diagram of the flow cytometer of the present invention. Figure 11A Detection results in the range of 300-600 nm are shown. Figure 11B Detection results in the range of 700-900 nm are shown. Figure 11C and Figure 11D An embodiment equipped with 21 detector arrays is shown.

[0019] Figure 12 A standard configuration of a flow cytometry system is shown.

[0020] Figure 13 A model of a flow cytometry system is shown, in which the fluorescence detection module is placed in the imaging and amplification optical path of the collection lens. A primary filter is located behind the lens to separate individual laser emissions from the main optical path and direct them to a separate physical location for scattered emission detection.

[0021] Figure 14 A flow cytometry system equipped with an expandable optical architecture is shown, which directs the focused light beam into different optical paths through a beam splitter element and achieves synchronous focusing of multiple detector arrays through optical segmentation.

[0022] Figure 15FIG. 4 shows an efficiency curve diagram of the diffraction grating according to the present invention.

[0023] Figure 16 A SPAD array according to the present invention is shown. DETAILED DESCRIPTION

[0024] The present invention may be understood with reference to the following detailed description, accompanying drawings, and examples, all of which are within the scope of the present invention. It is emphasized that the present invention is not limited to the specific apparatus, methods, applications, conditions, or parameters described and / or illustrated herein, and that the terminology used herein is illustrative only and does not constitute a limitation on the scope of the patent claims. It is emphasized that throughout the present invention and claims, the singular terms "a," "an," and "the" are to be understood to include the plural; a reference to a specific value includes at least that value unless the context clearly excludes such reference. The term "plurality" as used herein refers to a quantity greater than one. When a range of values ​​is expressed, another embodiment includes the range from the specific value to the next specific value. Similarly, when a value is prefixed with "about" to indicate an approximation, the specific value itself constitutes another embodiment. The scope definitions herein are inclusive and combinable, and the order of experimental steps may be interchanged. All documents cited herein for any purpose are incorporated herein by reference in their entirety.

[0025] It is worth emphasizing that, for the sake of clarity, the technical features of the present invention that are described in different embodiments may actually be integrated into the same embodiment and presented. Conversely, the multiple technical features of the present invention that are collectively included in the same embodiment for the sake of concise description may be presented individually or in the form of any sub-combination. It is worth emphasizing that the numerical values ​​stated in the range include any numerical value within the range. In addition, the term "provided with" should be understood to have its standard open meaning, but also covers the closed definition of "consisting only of..." For example, a device provided with component A and component B may include other components in addition to component A and component B, or may be composed only of component A and component B.

[0026] The present invention describes a method and system for implementing a single-photon avalanche diode for flow cytometry. Flow cytometry system detectors need to have the characteristics of high bandwidth and wide dynamic range, so SPADs are generally not used as detection elements in such systems. The bandwidth of the SPAD may be limited by the pixel reset time. By counting the number of pixels in the reset state when the SPAD is sensitive to light in real time, invalid pixels and activated pixels can be corrected, and the count value of the light signal can be adjusted accordingly. Therefore, even if some SPAD pixels cannot be sensitive to light because they are in the reset state, the corrected count value can still more accurately reflect the intensity of the light signal received by the SPAD at a given time. Therefore, the method described in this article can improve the bandwidth of the SPAD, allowing it to be applied to fields with higher bandwidths.

[0027] Figure 1 Flow cytometry system 100 according to the present invention is shown. A sample is ejected under pressure through nozzle 105, forming fluid stream 110 at the outlet of nozzle 105. Fluid stream 110 separates into a series of droplets 115. Laser 120 can illuminate cells in the sample at detection site 125. The optical response characteristics of the cells produced by illumination with laser 120 depend on the cell type. The optical response can be forward scattered light or forward fluorescent emission 130. The optical response can also be side scattered light or side fluorescent emission. The optical signal is transmitted to photodetector 140 via filter 135.

[0028] The photodetector 140 can generate an output signal whose waveform and intensity reflect the intensity of cellular light captured by the filter 135 at a specific frequency. In certain embodiments, the controller 145 can generate an event data signal based on the output signal transmitted by the photodetector 140. For example, the event data signal can include a timestamp to identify data about a specific cell. In certain embodiments, the event data signal can include sorting information. For example, the controller 145 can execute sorting logic and make sorting decisions based on the biological response of the cell type to be sorted. The probability of an event belonging to a population can be statistically calculated and applied to sorting decisions. The event data signal can be used to charge the liquid stream 110 before the droplets are separated from the liquid stream 110, enabling the deflection plates 150 to control the deflection of the charged droplets (e.g., for sorting). In certain embodiments, the output signal of the photodetector 140 can include a photon count value. As described below, in certain embodiments, the controller 145 can adjust the photon count value.

[0029] The content of the present invention is not limited to Figure 1 The specific flow cytometry system 100 shown is for illustrative purposes only. For example, the methods of the present invention can be implemented using a flow cytometry system equipped with the following configurations: multiple lasers of different wavelengths, various filters for filtering or directing emitted light, and (if necessary) various deflection plates or sorting instruments for sample sorting. Furthermore, the present invention can also be implemented using a flow cytometer and scanner equipped with cell sorting capabilities.

[0030] Figure 2 A SPAD 200 that can be implemented according to the technical solution of the present invention is shown. Figure 2 The top perspective view of the SPAD and its pixel unit 205 is shown. The SPAD 200 includes a plurality of pixel units 205, which can be arranged in a variety of configurations. For example, the SPAD 200 is shown as a nearly square structure (such as a 32×32 pixel array); however, it is worth emphasizing that the actual shape of the SPAD device is not limited to Figure 2In some embodiments, each pixel unit 205 may have a uniform geometric shape and form a square top surface profile; however, it should be understood that the specific shape and size of the pixel are not limited to the following. Figure 2 In addition, each SPAD 200 may be composed of a SPAD array according to the present invention. For example, in some embodiments, the SPAD array may include multiple SPADs connected to the underlying electronic components (such as Figure 3 In some embodiments, multiple SPAD arrays can be integrated to construct a specific SPAD element. For example, multiple SPAD arrays can be arranged in an edge-aligned manner (such as Figure 16 This configuration can extend the overall size of the SPAD array along a single dimension or multiple dimensions.

[0031] Each corresponding pixel unit 205 is activated after capturing a light signal (e.g., a photon) at a given time. The pixel unit 205 can transmit the electronic signal generated by the captured light signal to the underlying electronic components. Therefore, at a given time, the SPAD 200 can be used as a photoelectric counter, where each pixel unit 205 can capture a corresponding light signal, and the total count of the light signals captured at a given time can be used as a representative indicator of the intensity of the light received.

[0032] Each pixel unit 205 has a reset time, that is, after the pixel unit 205 is activated to capture a light signal, it will enter an interval during which it cannot respond to new light signals. For example, the reset time may be 10ns, during which the pixel unit cannot capture light signals. In some embodiments, the SPAD pixel unit may have a passive reset (e.g., passive quenching), which is achieved by a passive circuit (such as a passive quenching circuit). In some embodiments, the SPAD pixel unit may have an active reset (e.g., active quenching), which is achieved by an active quenching circuit (e.g., triggered by a digital output pulse to reduce the bias voltage across the active quenching circuit). Therefore, at any given point in time, the SPAD may have pixel units in the following three states at the same time: an inactive pixel that can capture light signals, an active pixel that needs to be reset before it can detect photons again, and a reset pixel that is unable to capture light signals during the reset period (e.g., the reset time window). Therefore, the photon count value of the SPAD 200 at a given time may depend on the number of pixels that capture light signals in that given time. Furthermore, in some embodiments, the reset time can be set individually per pixel (pixel-level reset) or uniformly across all pixels (global reset). For example, a specific reset time can be set to 10ns, which causes that specific pixel to remain in reset for 10ns or longer (enabling operation independent of other pixels in the array). In some embodiments, a global reset time can be set to 10ns, which causes all pixels in the SPAD array that enter the reset phase to reset simultaneously.

[0033] Figure 3 FIG3 shows a detector SPAD 300 equipped with the SPAD of the present invention. The detector may include a SPAD 200 and its supporting underlying electronic circuit 305, which is dedicated to receiving an electrical signal converted from an optical signal captured by a SPAD pixel unit. In some embodiments, the underlying electronic circuit 305 may include one or more application-specific integrated circuits (ASICs) for generating an electrical signal representing the photon count value of a specific time period based on the number of optical signals captured by the SPAD 200 in the specific time period. In some embodiments, the underlying electronic circuit 305 may also transmit the photon count signal to other components of the flow cytometry system (e.g., Figure 1 In some embodiments, the underlying electronic circuitry 305 may also process the photon count signal, such as by adjusting the photon count value, as will be described in more detail below.

[0034] In certain embodiments, the detector interface may include a power input 310, a ground input 325, a clock signal input 315, a control command input 320, and a data output 330. The clock signal input 315 can serve as a system-level reference clock, enabling full sampling synchronization across multiple detectors. The control command input 320, which can utilize a bus protocol such as I2C, provides functions such as setting control parameters, reading back status registers, disabling noisy pixels, shutting down unused pixels in the optical system, and performing other control functions. Extended functions, such as selecting an imaging mode, can encode the pixel value grid back to provide more information.

[0035] The data output 330 can transmit the total number of photon detection values ​​within the sampling period, and can use low-voltage differential signaling (LVDS) to output in a high-speed, low-noise manner. To facilitate data transmission, a prefix (such as a 0100 bit pattern) can be added before the data to achieve optimal transmission. If the normal output data is set to a 12-bit value, the additional 0100 bit pattern will generate a 16-bit output value after adding the prefix. The advantage of adding extra bits is that it can force the data stream to produce a level jump, allowing processing devices such as FPGAs to maintain synchronization even when transmitting all-0 or all-1 signals, while providing a clear data alignment identifier. The ground input 325 can provide a reference data to the detector, which can be used for voltage or current measurement of the detector.

[0036] When used in a system, the detector can serve as a replacement for a photomultiplier tube (PMT) and is suitable for flow cytometers, laser scanning fluorescence microscopes, and other instruments. Optical detection involves capturing the target with a collection objective, then splitting the beam through optical fibers (e.g., using a single fiber for each detection point), and directing and filtering the target beam using dichroic mirrors and filters. Furthermore, focusing and homogenizing optics can be added to evenly distribute the light onto the detector.

[0037] Figure 4 Figure 3 shows simulated detection counts of a SPAD for a Gaussian light pulse, plotting the number of detected photons as the number of photons increases. An ideal detection system would have a detection ratio of 1.0, representing a horizontal line. This means the system perfectly detects 1.0 for every incident photon. This also simulates quantum efficiency. Therefore, for a simulated system with a quantum efficiency of 40%, perfect detection would be a constant 0.4, representing a horizontal line. Figure 5 Shown Figure 4 The upper graph shown is a partial enlarged view of the graph (after the photon detection ratio axis is magnified). It can be seen that dark noise will significantly increase the detection ratio under extremely low photon number conditions. Under extremely low photon number conditions, the detection ratio is increased due to the influence of dark noise because the dark noise activation signal of the pixel is indistinguishable from the real photon signal. The detection noise of the flow cytometer is an inherent characteristic, and there are significant distribution differences in the weak biological signals themselves, so the detection result is within the allowable error range. However, the problem worthy of attention is the lack of detection linearity. Zooming in and observing the details of the linearity shows that when the input is on the order of 10,000 photons, its linear deviation is about 1.5% of the total number of photons, which is equivalent to 3.75% of the number of detected photons.

[0038] Figure 4 and Figure 5 The simulation results shown here are for a SPAD chip setup with 4096 pixels, a 30MHz readout frequency, and a sampling period of 33.333ns. The noise counts per pixel average 100 times per second, and the pixel reset time is 10ns. This reset time applies to previously triggered pixels that need to be quenched, and is performed at the beginning of the sampling period. During this reset time, the pixel cannot detect photons.

[0039] This nonlinearity stems from the pixel saturation effect and is the fundamental reason why the dynamic range of multi-pixel photon counters (MPPCs) is limited at high light intensities. MPPCs enable the pixel unit circuit to quickly reset to improve linearity. This improves linearity to some extent, but still fails to meet the expected requirements. However, by using a photon counting detection method that periodically resets the pixels, SPADs can cleverly take advantage of the characteristic that "pixel detection efficiency decreases as the photon count increases." When 50% of the pixels are activated, the probability of detecting a new photon decays to 50% of the baseline value. At this time, the statistical significance of a new photon detected is equivalent to 2 photons. A photon detection level correction lookup table (LUT) can be constructed to map the actual number of detected photons to the theoretical number of incident photons corresponding to the pixel fill state.

[0040] Setting the pixel reset time can effectively suppress photon detection efficiency at high light fluxes, thereby optimizing the detection dynamic range. This problem can also be corrected by counting the number of reset pixels within a specific time period and using the ratio of the reset time to the sampling period.

[0041] Integrating these correction factors corrects linearity and improves the dynamic range by approximately two orders of magnitude.

[0042] Figure 6 Figure 2 shows the simulated detection count curve results for a SPAD where the detection counts have been corrected for the number of active pixels at the time of detection (e.g., linearity correction is enabled). Figure 5 similar, Figure 7 Shown Figure 6 A zoomed-in view of the upper graph (with the photon detection magnified relative to the coordinate axis) is shown. As shown, the system linearity is significantly improved after correction, now maintaining a linear response over a range of over 1,000,000 incident photons. The simulator is configured by default to randomly distribute photons across the entire surface of the chip, but this random distribution is not required; other photon distributions can also be implemented. In some embodiments, the linearity correction value can be adjusted based on the type of photon distribution.

[0043] Figure 8 The figure shows the detector detection efficiency (PDE) for photons of different wavelengths. The blue curve represents typical SPAD performance parameters, while the other colors represent the performance parameters of different photomultiplier tube (PMT) models. The light green and dark green curves represent standard tube types, respectively. The orange and purple curves represent improved PMTs, respectively.

[0044] The photon detection efficiency of this SPAD has significant advantages in the band above 500nm, especially in the red and near-infrared bands, where its performance far exceeds that of PMT. Correction factor calculation

[0045] For a pixel array 4096, the input value for a single detection cycle ranges from 0 to 4096 to represent the number of pixels activated. A correction factor is calculated for each input value (pixel count) to obtain the expected number of photons obtained by the activated pixel count.

[0046] Where: "Pixel count" represents the pixel count currently being counted, which is marked as (1) and (2) in chronological order, and so on. "Number of pixels" represents the total number of pixels. Taking a pixel array of 4096 as an example, the initial count result is: Pixel Count Correction value 1 1.00000 2 2.00024 3 3.00073 4. 4.00147 … 4093 28925.01223 4094 30290.34556 4095 32338.34556 4096 36434.34556

[0047] In some embodiments, these correction factors may be calculated in advance and archived in a photon detection level correction lookup table (LUT) for real-time correction. Active reset system reset correction factor calculation

[0048] The reset correction factor can be calculated once based on the number of pixels activated in the previous cycle, or the correction factor can be applied to the correction value. The correction factor is calculated as follows: (number of pixels activated in the previous cycle / total number of pixels) × (Reset time / Cycle time)

[0049] Where: "Previous cycle pixel count" represents the number of activated pixels in the previous cycle; "Total pixels" represents the total number of pixels received by the detector; "Pixel reset time" refers to the time required for a pixel to reset from the activated state to the enabled state; and "Detection cycle" represents the duration of a complete detection cycle (i.e., 1 / frequency).

[0050] Taking a 4096-pixel array as an example: if 1000 pixels were activated in the previous clock cycle, the pixel reset time is 10ns, and the detection period is 33.33333ns, the correction value can be calculated to be 1146.31, and the adjusted photon count is 1236.90. Calculation of reset correction factor for passive reset system

[0051] The previous discussion focused on active pixel reset, also known as synchronous quenching, in which the pixel reset occurs at the start of the clock cycle. However, ST Micro's current standard pixels feature passive quenching, meaning the pixel resets as soon as possible after being set, independent of the clock cycle. The current setting is to reset approximately 10ns after a photon triggers the pixel. This feature allows a pixel to be activated multiple times within a single clock cycle.

[0052] For passive reset systems, the correction value only requires a single calculation. The algorithm is: number of detected photons divided by the system's effective integration time (i.e., the cumulative time that the pixel is not in the reset state within the total cycle). Because a single pixel can reset and detect multiple photons within a clock cycle, the calculated value range of the correction table needs to be expanded to detect more photons. To cover all possible cases, the number of iterations is calculated as follows:

[0053] Among them, "Total Pixels" represents the total number of pixels received by the detector; "Pixel Reset Time" represents the time required for the pixel to recover from the enabled state to the active state; and "Detection Cycle" represents the length of time for a complete detection cycle (i.e., 1 / Frequency).

[0054] A photon detection level correction lookup table (LUT) is created through iterative calculation, and the corresponding theoretical incident photon value is output using the number of detected photons as an index.

[0055] Where: "Photon count" represents the number of activated pixels in the previous cycle; "Total pixels" represents the total number of pixels received by the detector; "Pixel reset time" represents the time required for the pixel to recover from the enabled state to the activated state; and "Detection cycle" represents the length of time for a complete detection cycle (i.e., 1 / frequency).

[0056] For example, if the photon count for the previous clock cycle of pixel array 4096 was 1000, the pixel reset time is 10ns, and the detection cycle is 40ns, the calculated correction value is 1065.00260. This calculation can be done in advance and can usually be converted into a multiplication correction factor, so no division operation is required.

[0057] In some embodiments, a rectangular array structure can be created to achieve the following dual detection optimization: spatial dimension: improving the detection efficiency of single-wavelength photons by distributing multiple pixels (as described above); and spectral dimension: using pixels in a second dimension to distinguish light signals of different wavelengths. This may require the use of prisms, gratings, or other spectroscopic techniques to expand the light by wavelength.

[0058] Another optical processing solution can expand the detection range to encompass multiple detection sites on a flow cytometer. Since cells move linearly in the fluid flow and sequentially pass through multiple laser detection sites, an objective lens can be used to simultaneously focus the emitted light from all detection sites onto the detector surface, spatially separating the optical signals from different detection sites and distributing them to specific regions of the detector. This optical system utilizes a combination of dichroic mirrors and filters, with simultaneous light splitting behind the lens (in the free-space optical path), achieving the following functions: wavelength routing: allocating the light beam to different paths based on wavelength; and confocal design: precisely focusing each independent light path at the corresponding multiple detection sites on the detector surface. For the detector, a detector array can be used to receive optical signals from lasers in different regions, thereby achieving multi-laser detection. Alternatively, multiple independent SPAD detectors can be integrated into a detector array at a specific spacing to simultaneously collect optical signals from multiple detection sites. The present invention utilizes a suitable arrangement of such a detector array. For example, a detector measuring 1 mm x 10 mm can be configured as an array of 32 x 320 pixels or 64 x 640 pixels. Light signals from different laser collection points can be received by different portions of the detector array. Therefore, the laser can be configured with a detector array to select its corresponding dedicated pixel area. For example, a spectral cell sorter should be equipped with 7 laser acquisition points, so the detector can be configured as 7 independent working areas to receive different laser signals. This function can be achieved through software configuration, or the detector can be designed or pre-configured to define different pixel groups as independent detection units. Taking a 1×10mm array as an example, the detector array can be configured as 10 independent 1×1mm detector units. To meet the use requirements of a spectral cell sorter, the detector array can be configured as 7 independent 1×1 detector units, and an unactivated spacing area should be provided between each detector unit.

[0059] Such detectors can share the same chip input interface, but multiple output interfaces should be configured to ensure that each output channel has a reasonable bandwidth. For example, a chip designed for a spectral cell sorter could be configured with seven independent LVDS output channels, one for each detector interface. Alternatively, to reduce wiring complexity, each output port could use an interleaved transmission mode to process data from one or more output detectors. At a 30MHz operating frequency, if 16 bits of data are output according to the aforementioned scheme, the data transmission rate per channel can reach 480Mbps. Two channels can use a single output interface to interleave data transmission, reducing the data link by half, allowing a data transmission rate of 960Mb / s per output channel. Standard LVDS signals at 960Mbps can be stably received by the FPGA's LVDS input interface. For a spectral cell sorter, this solution allows the existing electronic system with 60 LVDS input signals to process twice the number of channels.

[0060] To identify the data in the output channel, the prefix bits in the data stream can be changed, prefixing the first channel data with 0100 and the second channel data with 0101.

[0061] This design makes the system structure reasonable, improves efficiency, reduces product costs, and reduces the number of detectors used. Optical configuration

[0062] The flow cytometer described in the present invention may include a single detector output mode or a multi-detector partitioning mode, in which a single detection area receives the same fluorescent channel signal from multiple laser receiving points, and different areas of the same detector array independently receive the same fluorescent channel signal from different laser points. It is characterized in that the detector array located in the focal plane of the lens realizes multi-detection point signal acquisition through a shared filter set, and the signal of each detection point is processed as an independent signal on the detector array. The total number of processed signals can be equal to (or less than) the total number of laser detection points. Since each detection point has a unique imaging area and data processing channel, data processing will not be affected by the signal overlap problem, and each channel can be optimized to achieve optimal performance.

[0063] Rather than using the optical system of a flow cytometer to create a completely separate optical detection library, the optical configuration described in this invention reduces the size of the detection architecture of multiple integrated areas, maximizes shared filters in the free-space optical path, and controls lens and filter costs. The instrument adopts a compact structural design.

[0064] This configuration also facilitates detection of side-scatter signals from laser detection points. Since fluorescence signal detection across multiple detection points requires coverage of a wide wavelength range and high-quality collection lenses, this configuration is well-suited for signal acquisition from the scatter channel. Since the intensity of the scatter channel is several orders of magnitude higher than the fluorescence signal, adding a custom filter (scatter sensor) at the front end of the optical system to filter out the wavelength band surrounding the laser detection point is advantageous. This allows the fluorescence signal to enter the fluorescence filter region without laser interference. This scattered light splitting optical element directs the scattered light from the laser excitation at all detection points to independent detector arrays (similar to the fluorescence detector array structure described later), enabling more precise signal separation and parallel processing. In this scenario, the scattered light signals are spatially separated, but their wavelengths are the same as the laser wavelength at the corresponding detection points. Depending on the required scatter signal quality, spatial filtering or bandpass filtering of the scattered light at different wavelengths may be required before detection. It is worth noting that the advantage of this architecture is that high-quality scattered light spot signals can be directly acquired in the free-space optical path, allowing secondary signal processing of one, two, or all scattered light signals before detection.

[0065] For example, the optical configuration of a flow cytometry system creates a set of common fluorescence channels for multiple laser detection points. The preferred embodiment of the present invention directly utilizes the system's own lasers to establish a common spectral breakpoint reference between multiple detection points, thereby achieving synchronous calibration and signal correlation for the entire system. Figure 9 Shown is a schematic diagram of the filter configuration for a flow cytometer.

[0066] For example, an exemplary configuration of a flow cytometer includes lasers 6 and output channels 51, where filters in each channel can be used to consistently detect similar fluorescence emitted by different lasers. Using the optical configuration described herein, the same filters can be used across multiple lasers, resulting in significant improvements. Figure 10 A method of expanding the filter set is shown to maximize filter utilization for each laser.

[0067] For example, the total number of fluorescence channels can be increased from 51 to 78, but by adding detector arrays at multiple detection points, the total number of bandpass filters can be reduced from 51 to 21. Figure 11 shows that the detector array can efficiently collect light signals. Figure 11A The total number of fluorescence channels shown in Figure 1 is 78, and the total number of detector arrays is 21. The optical configuration described in this article (e.g. Figure 13 and Figure 14 As shown) can realize the precise separation of light by wavelength (or wavelength band). The optical signal can originate from multiple detection points (for example, 6 detection points), and by at least partially sharing common optical elements (for example) in the emission light path of multiple detection points, the optical hardware configuration is simplified. In some embodiments, a SPAD detector or a SPAD detector array can be used in the system described herein. Other detectors or detector arrays can also be used instead of the SPAD detector or the SPAD detector array, or in combination with the SPAD detector or the SPAD detector array. For example, other detectors that can be used in the system may include an avalanche photodiode (APD), a multi-pixel photon counter (MPPC), a photomultiplier tube (PMT), and the like.

[0068] The optical configuration can further reduce the number of dichroic filters required by reusing the same filters for all lasers. The detection paths can be centrally managed to produce high-quality optical imaging, generating multiple light paths to the detector array at all detection points.

[0069] like Figure 12As shown, a single light-collecting objective lens typically utilizes a high-resolution design, achieving separation of each detection point with a high magnification factor on the imaging plane. At the imaging plane, the scattered light and fluorescence data from each detection point are typically separated from other data using filters, objectives, or prisms. In a typical optical path, the light-collecting objective lens must maintain a specific working distance to properly amplify and image the scattered light signal. After maintaining this specific working distance, the light-collecting objective lens uses additional optical hardware to separate the fluorescence channel. For example, when six detection points are involved, the second section of the optical path enables separate detection of scattered light and fluorescence from each laser.

[0070] The optical configuration described in this article can use the optical system to amplify the imaging module to achieve fluorescence detection at all detection points. Figure 13 As shown, the dichroic filters 1320 in system model 1300 are all short-pass filters, with their center wavelengths matching the wavelengths of each laser within the system. This generates discrete optical signal "blocks" that do not span any laser wavelengths and exist only within a predetermined laser wavelength range. This optical design directs each discrete optical signal "block" along a separate physical path. While the spectral bandwidth of each discrete optical signal "block" is narrow, it still preserves data from all detection sites. The optical system design must determine how to perform secondary processing on this "block" optical signal, filtering it into one or more wavelength bands before focusing it onto the detector array 1335 and associated circuitry described above.

[0071] System model 1300 includes an objective subsystem 1305, which can be equipped with a high-quality objective lens. The objective lens can be adapted to different lens modules for different systems (e.g., flow analyzers, flow sorters, spectral confocal microscopes, etc.). The light beam passes through the objective subsystem 1305 and optically interacts with a scattering filter 1315 (e.g., a multi-notch scattering filter). Certain portions of the light beam can be directed to an optical signal detector, such as a side scattered light detector 1310, which can receive scattered light generated by some or all of the excitation lasers used in system 1300. When interacting with the scattering filter 1315, a portion of the light beam can be transmitted to the dichroic filter 1320, as described above. Certain light beams that interact with the dichroic filter 1320 can be directed to a spectrometer module 1325, which can emit the light beam to a diffraction grating to achieve continuous dispersion on the detector. Some of the light beams that interact with dichroic filter 1320 can be directed to a branch detection module 1330, which transmits the light beams to a cascaded dichroic mirror and bandpass filter for detection of the corresponding wavelength bands by a detector. Some of the light beams that interact with the dichroic filter can be directed to a detector 1335, which is equipped with an array detection surface 1345. For example, the array detection surface is divided into six functional areas from top to bottom: the first area: collects the fluorescence signal of the first laser; the second area: collects the fluorescence signal of the second laser; the third area: collects the fluorescence signal of the third laser; the fourth area: collects the fluorescence signal of the fourth laser; the fifth area: collects the fluorescence signal of the fifth laser; and the sixth area: collects the fluorescence signal of the sixth laser. The light signals reaching detector 1335 can be presented as multiple detection points 1340. These detection points are imaged by the objective lens and focused on the confocal plane of detector 1335.

[0072] Modular optical configurations are available. Flow cytometry systems use a variety of configurations in which the number of excitation lasers can be increased or decreased. When a laser is removed from this configuration, its corresponding dichroic beamsplitter and all associated filter assemblies must be removed. The principles for adding additional lasers are different: either a dichroic mirror can be added to capture the fluorescence signal from the additional laser, or the existing optical architecture can remain unchanged. In both cases, the addition of spatially separated laser sources requires the addition of detectors to the detector array. If additional lasers are added, special coatings for the scatter beamsplitter or bandpass filters must be added to block scattered light from the new laser from reaching the fluorescence detector.

[0073] like Figure 13As shown, the light extracted from the main imaging section can be further processed to separate the fluorescence into its final optical channel. However, due to the imaging characteristics, the target color light remains continuously focused during the deflection path and ultimately forms a common focal plane at the same total distance from the lens. Furthermore, dichroic mirrors and bandpass filters or other dispersive techniques such as prisms and diffraction gratings can be used to separate the beams. The signal processing of each optical path segment can be customized, even supporting user-level parameter adjustment. Figure 14 The general design and new technologies that are compatible with the proposed architecture are shown. For example, Figure 14 The objective lens system 1405 shown (can be used with Figure 13 1305 corresponding to), side scattered light detector 1410 (can be with Figure 13 corresponding to the side scattered light detector 1310), scattering filter 1415 (which can be Figure 13 scattering filter 1315), dichroic filter 1420 (which can be Figure 13 The dichroic filter 1320 corresponds to the dichroic filter 1320) and the branch detection module 1425 (which can be Figure 13 Such components include the branch detection module 1330 (corresponding to the branch detection module 1330). The portion of the light beam that interacts with the dichroic filter 1420 can be directed to the spectrometer module, which transmits the beam to a diffraction grating, achieving continuous dispersion at the detector. For example, the optical signal can be spectrally separated, spatially separated, or both.

[0074] In some embodiments, an alignment configuration of a flow cytometer can be performed. For example, a SPAD array can be configured as a terminal detector of the main optical path (e.g., a 32×320 pixel array). The flow cytometer can be configured to operate in pixel imaging mode (outputting a raw pixel distribution image) rather than in integral accumulation mode. In some embodiments, the SPAD array can be instructed to operate in this specific mode, for example, via I2C or other communication buses, instructing the SPAD array to operate in this mode for one sampling cycle. The SPAD array can then output data for inactive pixels (e.g., "0") and active pixels (e.g., "1"), which, after analysis, can generate a black and white image of the binary distribution of photon density.

[0075] In some embodiments, the alignment process / mode can include simultaneously observing the aforementioned images and adjusting optical components to optimize beam focus and positioning to ensure optimal light signal at the detector surface. Furthermore, the alignment process helps determine the availability of detection area on the SPAD array, which can then be allocated for detection. Exemplary embodiments

[0076] The following embodiments are merely illustrative and do not limit the scope of the present invention and the appended claims. It should be understood that any part of any one or more embodiments may be combined with any part of any other one or more embodiments. Example 1

[0077] A computer-implemented method comprising: receiving one or more light signals at a point in time, wherein each light signal is captured by a corresponding pixel of a SPAD array; measuring the number of activated pixels of the SPAD array at that point in time, and adjusting the light signal count captured by the SPAD detector at that point in time based on the number of pixels. Example 2

[0078] The computer-implemented method of embodiment 1 further comprising storing the adjusted count in a memory. Example 3

[0079] The computer-implemented method according to any one of embodiments 1-2, wherein the SPAD array consists of SPAD detectors. Example 4

[0080] The computer-implemented method according to any one of embodiments 1 to 3, wherein the SPAD array consists of a 32×32 pixel array. Example 5

[0081] The computer-implemented method of any one of embodiments 1 to 4, for receiving a plurality of light signals, wherein each light signal is captured by a corresponding pixel in a different one of a plurality of SPAD arrays. Example 6

[0082] The computer-implemented method according to any one of embodiments 1 to 5, wherein the reset time is set to 10 ns. Example 7

[0083] The computer-implemented method according to any one of embodiments 1 to 6, wherein the reset time of each corresponding pixel is unique. Example 8

[0084] The computer-implemented method according to any one of embodiments 1 to 7 further includes: determining a correction factor of the received one or more light signals based on the number of activated pixels at the time point, and adjusting the light signal count based on the correction factor. Example 9

[0085] According to the computer-implemented method of any one of embodiments 1 to 8, the correction factor may be determined by retrieving the correction factor from a photon detection level correction lookup table (LUT). Example 10

[0086] The computer-implemented method of any one of embodiments 1 to 9, wherein the correction factor is determined according to the following formula:

[0087] Where "correction value" is the correction factor, "pixel count" is the number of activated pixels, and "number of pixels" is the total number of pixels in the SPAD array. Example 11

[0088] The computer-implemented method of any one of embodiments 1 to 10, wherein the light signal count is adjusted based on:

[0089] Where "correction value" is the correction factor, "total number of pixels" is the total number of pixels in the SPAD array, "pixel count in the previous cycle" is the number of activated pixels in the SPAD array in the previous cycle, "pixel reset time" is the corresponding pixel reset time of the SPAD array, and "detection cycle" is the corresponding pixel detection cycle of the SPAD array. Example 12

[0090] The computer implemented method as described in any one of embodiments 1 to 12, wherein the optical signal comprises photons. Example 13

[0091] The computer-implemented method of claim 1 , further comprising correlating the light signal counts with properties of particles emitting the light signals, the particles being cells. Example 14

[0092] A flow cytometric system comprises: a flow cytometer configured to illuminate particles at a detection site; and a SPAD detector configured to capture emission signals generated by the illuminated particles at the detection site. Example 15

[0093] The flow cytometry system according to embodiment 14 is further provided with a controller configured to perform the embodiments according to any one of claims 1 to 13. Example 16

[0094] A non-transitory computer-readable medium comprising a processor, a memory, and

[0095] The computer executable instructions stored in the memory can cause the processor to perform the computer-implemented method described in any one of embodiments 1 to 13 when executing the instructions. Example 17

[0096] A flow cytometry system is provided with: a plurality of lasers configured to illuminate particles at corresponding detection sites; A first filter, which: (i) is configured in a main optical path of the emission light generated after the multiple lasers excite the particles, and (ii) is configured to receive laser excitation light from at least one laser among the multiple lasers and redirect a first portion of the received laser excitation light along a first auxiliary optical path; and at least one SPAD array, which is configured to receive the emission light of the first auxiliary optical path. Example 18

[0097] According to the flow cytometry system of embodiment 17, the first portion of the laser excitation light is located within a specific wavelength spectrum range, and the at least one SPAD array includes at least one detection area configured to receive the partial area of ​​the wavelength spectrum. Example 19

[0098] The flow cytometry system according to any one of embodiments 17 to 18, wherein the partial region of the wavelength spectrum comprises a fluorescence channel. Example 20

[0099] The flow cytometry system of any one of embodiments 17 to 19, wherein the detection area comprises a pixel of at least one SPAD array. Example 21

[0100] The flow cytometry system according to any one of embodiments 17 to 20 is further provided with an optical detector positioned at the end of the main light path and configured to collect the laser excitation light from the main light path. Example 22

[0101] The flow cytometry system according to embodiment 21, wherein the optical detector is configured to detect the alignment status of each element of the main optical path. Example 23

[0102] The flow cytometry system of any one of embodiments 17 to 22 further comprises a second filter: (i) positioned along the auxiliary optical path, and (ii) configured to redirect a second portion of the excitation light from the plurality of lasers along the second auxiliary optical path. Example 24

[0103] The flow cytometry system according to embodiment 23 is further provided with a SPAD array, which is positioned along the second auxiliary optical path and configured to receive the laser excitation light of the second auxiliary path. Example 25

[0104] The flow cytometry system according to claim 23, wherein the optical signals of the first auxiliary optical path and the second auxiliary optical path are received by the at least one SPAD array. Example 26

[0105] The flow cytometer system according to any one of embodiments 17 to 25, wherein the filter is provided with a dichroic mirror. Example 27

[0106] The flow cytometry system of embodiment 26, wherein the dichroic mirror is a 45 degree dichroic mirror. Example 28

[0107] The flow cytometry system according to any one of Examples 17 to 26 is further provided with a scattering filter, which is positioned along the main light path and located between multiple lasers and at least one filter, wherein the scattering filter is configured to deflect the side scattering signal of the laser excitation light away from the main light path. Example 29

[0108] Embodiment 28: The flow cytometry system of any one of embodiments 17 to 28, wherein each laser of the plurality of lasers is configured to emit light at a different wavelength than the other lasers. Example 30

[0109] The flow cytometry system according to any one of embodiments 17 to 29 is further provided with an objective lens configured to receive laser excitation light from the plurality of lasers and guide it into a main optical path. Example 31

[0110] A flow cytometry system according to any one of embodiments 17 to 30, further comprising a controller configured to perform the computer-implemented method according to any one of claims 1 to 13.

Claims

1. A method comprising: One or more optical signals are received at a certain point in time, wherein each optical signal is captured by a corresponding pixel of at least one SPAD array, wherein the pixel remains in an activated state for a reset time after capturing the optical signal. determining a number of activated pixels of at least one SPAD array during the time point; and The light signal count captured by the at least one SPAD array at the time point is adjusted according to the number of activated pixels of the at least one SPAD array at the time point. The light signal counts can be correlated to the nature of the particle emitting the light signal, which can be a cell.

2. The method according to claim 1, further comprising: The adjusted count is stored in memory.

3. The method according to any one of claims 1 and 2, wherein the at least one SPAD array is arranged in at least one SPAD detector.

4. The method of any one of claims 1 to 3, wherein the SPAD array comprises a 32x32 pixel array.

5. The method according to any one of claims 1 to 4, comprising receiving a plurality of optical signals, wherein each optical signal is captured by a corresponding pixel of a different one of the plurality of SPAD arrays. The method according to claim 1 , wherein the reset time is approximately 10 ns. The method of claim 1 , wherein the reset time of each corresponding pixel is unique.

8. The method according to claim 1, further comprising: determining a correction factor for the received one or more light signals based on the number of activated pixels at that point in time; and Adjust the light signal counts according to the correction factor.

9. The method of claim 8, wherein the correction factor is determined by retrieving the correction factor from a photon detection level correction lookup table.

10. The method of claim 8, wherein the correction factor is determined by: Where "Correction Value" is the correction factor, "Photon Count" is the number of pixels activated in the previous detection cycle, "Pixel Reset Time" is the time used to reset the pixel to the active state, "Detection Cycle" is the time period of the detection cycle, and "Total Pixels" is the total number of pixels in the SPAD array.

11. The method according to claim 8, wherein the light signal count is used as a final correction value, and its formula is as follows: Where "Correction value" is the correction factor, "Total number of pixels" is the total number of pixels in the SPAD array, "Pixel count in previous cycle" is the number of activated pixels in the SPAD array in the previous cycle, "Pixel reset time" is the corresponding pixel reset time of the SPAD array, and "Detection cycle" is the corresponding pixel detection cycle of the SPAD array.

12. The method of claim 1, wherein the optical signal comprises photons.

13. The method of claim 1, further comprising correlating the light signal counts with properties of particles emitting the light signals, the particles being cells.

14. A flow cytometry system comprising a controller configured to perform the method of claim 1.

15. A non-transitory computer-readable medium comprising: processor; Memory; as well as The computer executable instructions stored in the memory can cause the processor to perform the computer-implemented method of claim 1 when executed.

16. A flow cytometry system comprising: a plurality of lasers configured to illuminate the particles at corresponding detection sites; A first filter which: (i) disposed in a main optical path of laser excitation light generated after the plurality of lasers excite the particles, and (ii) configured to receive laser excitation light from at least one laser among the plurality of lasers and redirect a first portion of the received laser excitation light along a first auxiliary optical path; and At least one detector array is positioned and configured to receive the laser excitation light of the first auxiliary optical path.

17. The flow cytometry system of claim 16, wherein the first portion of the laser excitation light is within a specific wavelength spectrum range, and the at least one detector array comprises at least one detection region configured to receive the portion of the wavelength spectrum.

18. The flow cytometry system of claim 17, wherein the portion of the wavelength spectrum comprises a fluorescence channel.

19. The flow cytometry system of claim 17, wherein the detection area comprises pixels of at least one detector array.

20. The flow cytometry system according to claim 16, further comprising an optical detector positioned at an end of the main light path and configured to collect the laser excitation light from the main light path.

21. The flow cytometry system of claim 20, wherein the optical detector is configured to detect the alignment status of each element of the main optical path.

22. The flow cytometry system according to claim 16, further comprising a second filter, which: (i) positioned along an auxiliary optical path, and (ii) configured to redirect a second portion of the laser excitation light of the plurality of lasers along a second auxiliary optical path.

23. The flow cytometry system according to claim 22, further comprising a detector array positioned along the second auxiliary optical path and configured to receive the laser excitation light of the second auxiliary path.

24. The flow cytometry system of claim 22, wherein the optical signals of the first auxiliary optical path and the second auxiliary optical path are received by the at least one detector array.

25. The flow cytometry system of claim 16, wherein the filter is provided with a dichroic mirror.

26. The flow cytometry system of claim 25, wherein the dichroic mirror is a 45-degree dichroic mirror.

27. The flow cytometry system of claim 16, further comprising a scatter filter positioned along a main optical path and between the plurality of lasers and the at least one filter, wherein the scatter filter is configured to deflect side scatter signals of the laser excitation light away from the main optical path.

28. The flow cytometry system of claim 16, wherein each laser of the plurality of lasers is configured to emit light at a different wavelength than the other lasers.

29. The flow cytometry system of claim 16, further comprising an objective lens configured to receive laser excitation light from the plurality of lasers and direct it into the main optical path.

30. The flow cytometry system of claim 16, further comprising a controller configured to perform the computer-implemented method of claim 1.

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