Method for performing flow cytometry using conventional or spectral flow cytometry

Abstract

Methods for performing flow cytometry using conventional or spectral flow cytometry. The method comprises: after calibrating the conventional flow cytometer, generating an initial overflow matrix with the conventional flow cytometer by using a plurality of single stain compensation controls; passing the biological sample through a conventional flow cytometer; generating a measured sample event vector by measuring fluorescence of a plurality of cells in a biological sample passing through a conventional flow cytometer; and after passing the biological sample through and generating a measured sample event vector with a conventional flow cytometer, a processor of the computer executes instructions to perform the following functions: generating a compensated sample event vector by using the initial overflow matrix and the measured sample event vector; generating an adjusted overflow matrix by finely adjusting the initial overflow matrix; a recompensated event vector is calculated by precisely separating the spectrum using the adjusted overflow matrix and the measured sample event vector.

Description

[0001] This application is a divisional application of the invention patent application with the original application number 201980034214.1 (International application number: PCT / US2019 / 033419, international application date; May 21, 2019, invention name: Rapid re-compensation of flow cytometry data for overflow readjustment). Technical Field

[0002] Embodiments of the present invention generally relate to analyzing flow cytometer data to determine the count of particles in a sample. Background Art

[0003] Flow cytometers typically have an observation aperture illuminated by one or more lasers. Laser light from the one or more lasers strikes various fluorescein-labeled microparticles that pass through the aperture. Fluorescently labeled microparticles are typically various biological cells in a sample labeled with different fluorochromes, which can be analyzed to obtain information about the sample that can be generalized to the whole. One or more optical detectors in a flow cytometer are used to sense the fluorescence (fluorescence) emitted by the fluorescein-labeled microparticles that pass through the aperture struck by laser light from the one or more lasers.

[0004] One or more different filters can be arranged before the fluorescence emitted from the fluorescein-labeled microparticles reaches each detector. The filters are arranged in the emission fluorescence path so that each detector sees only a specific bandwidth of light associated with the expected fluorescence of the fluorescein. That is, the bandwidth of any given filter utilizes the peak in the emission spectrum of a specific fluorescent dye. In this way, for any given microparticle, the aggregate signal from the detector indicates the type of one or more fluoresceins attached to the microparticle. The signal detected by the detector from the emitted fluorescence allows rapid and comprehensive cell classification of various microparticles in the sample.

[0005] However, emission spectra may overlap between dyes. This limits the number of different fluorophores that can be detected simultaneously on a given particle by a single laser and detector. Because the emission bandwidth is typically in the wavelength range between 30 nanometers (nm) and 60 nm, conventional flow cytometers can typically detect no more than four or five fluorophores per laser line. Increasing the number of lasers provides an advantageous but expensive method to increase the number of fluorophores that can be detected simultaneously.

[0006] Further complicating the detection of emitted fluorescence is the fact that many of the dyes used to color the microparticles are excited in a laser wavelength range that is different from and greater than the typical 30nm to 60nm bandwidth range. This can lead to signal crosstalk between detectors for different lasers.

[0007] Existing flow cytometry fluorescence detection systems limit divergence by increasing the focal length of the collimating lens. However, depending on the required final image size, this results in a larger diameter beam, which limits the number of detectors, such as to six detectors. In these existing flow cytometry systems, the final image size is constrained by optical aberrations when collimating large-scale optical imaging (such as 800 microns (μm)) and broadband light (e.g., wavelengths of 400nm-800nm) into a group of detectors with a diameter of less than 3 millimeters (mm).

[0008] In another flow cytometry system, incident light is re-imaged using a spherical micromirror for each detector in a detector chain of a set of detectors arranged in a row. Re-imaging avoids the divergent collimated light problem of the flow cytometry system described above. However, the number of detectors is limited by the aberrations introduced by reflections from the spherical micromirrors. As the image size increases along the detector chain, in order to increase the number of detector channels arranged in a row in the detector chain, large area detectors are required, resulting in large flow cytometers that are bulky and expensive.

[0009] The resolution of the emission spectrum of spectral overlap is also important for increasing the number of detectable fluoresceins. Detector arrays can be used to identify fluoresceins based on the lumped emission signature across multiple wavelengths, thereby increasing the number of detectable fluoresceins. In essence, the entire fluorescent signal is dispersed into the detector array by a diffraction grating or a prism. In this way, the entire emission spectrum is discretized across multiple detectors. Spectral deconvolution (unmixing) can be used to calculate the contribution of a known single fluorescein spectrum to the total lumped signal. However, this method for increasing the number of detectable fluoresceins has two major limitations.

[0010] The continuous linear nature of the dispersion component / detector array does not allow the bandwidth to be adjusted to exploit the true nature of the fluorescein spectrum. Therefore, the identification of a wider bandwidth favors longer wavelength fluoresceins while ignoring the details of shorter wavelength fluoresceins with compressed spectra. In addition, scattered light from other lasers (if present) is inevitably collected by the detector array. This scattered light detracts from the fluorescence signal that the detector is intended to detect.

[0011] Therefore, further improvements in flow cytometry are needed to increase the amount of detectable fluorescein and to better analyze fluorescein-labeled microparticles in a sample. Summary of the invention

[0012] According to a first aspect of the present invention, there is provided a method for performing flow cytometry using a conventional flow cytometer, the method comprising the following steps:

[0013] An initial spillover matrix was generated by using multiple single-stained compensation controls;

[0014] passing the sample through the flow cytometer;

[0015] generating a measured sample event vector by measuring fluorescence of a plurality of cells passing through the flow cytometer;

[0016] generating a compensated sample event vector by using the initial spillover matrix and the measured sample event vector; and

[0017] generating an adjusted spillover matrix by fine-tuning the initial spillover matrix,

[0018] The step of generating an adjusted overflow matrix comprises: adding an incremental matrix to the initial overflow matrix.

[0019] According to the first aspect of the present invention, the method further comprises the following steps:

[0020] A re-compensated event vector is calculated by using the adjusted spillover matrix and the measured sample event vector.

[0021] According to the first aspect of the present invention, the compensated sample event vector is equal to the measured sample event vector linearly multiplied by the inverse of the initial spillover matrix.

[0022] According to the first aspect of the present invention, the incremental matrix comprises one or more incremental values ​​for fine-tuning the initial spillover matrix, and wherein the incremental matrix has the same dimension as the initial spillover matrix.

[0023] According to the first aspect of the present invention, the initial spillover matrix comprises N×N dimensions, and wherein N is the number of the single-stained compensation controls.

[0024] According to the first aspect of the present invention, each of the single staining compensation controls comprises one of the following: fluorescein isothiocyanate FITC; R-phycoerythrin PE; peridinin chlorophyll protein complex PerCP; PE-Cy7; allophycocyanin APC; or APC-Cy7.

[0025] According to the first aspect of the invention, the measured sample event vector comprises N values, and wherein the compensated sample event vector has N values.

[0026] According to the first aspect of the present invention, the re-compensated event vector is equal to the inverse of the sum of the initial spillover matrix and the increment matrix multiplied by the measured sample event vector.

[0027] According to a second aspect of the present invention, there is provided a method for performing flow cytometry using a spectral flow cytometer, the method comprising the following steps:

[0028] An initial spectral spillover matrix was generated by using multiple single-stained compensation controls;

[0029] passing the sample through the flow cytometer;

[0030] generating a measured sample event vector by measuring fluorescence of a plurality of cells passing through the flow cytometer;

[0031] generating an unmixed sample event vector by using the initial spectral spillover matrix and the measured sample event vector; and

[0032] Generate an adjusted spectral spillover matrix by fine-tuning the spectral spillover matrix,

[0033] The step of generating an adjusted spectral spillover matrix comprises: adding an incremental matrix to the initial spectral spillover matrix.

[0034] According to a second aspect of the present invention, the method further comprises the following steps:

[0035] A re-compensated event vector is calculated by using the adjusted spectral spillover matrix and the measured sample event vector.

[0036] According to a second aspect of the present invention, the unmixed sample event vector is equal to the measured sample event vector linearly multiplied by the initial spectral spillover matrix.

[0037] According to a second aspect of the present invention, the number of variables in the unmixed sample event vector is less than the number of variables in the measured sample event vector.

[0038] According to a second aspect of the present invention, the method further comprises the following steps:

[0039] The step of solving the unmixed sample event vector includes using a least squares algorithm on the measured sample event vector and the initial spectral spillover matrix.

[0040] According to a second aspect of the invention, the initial spectral spillover matrix for the unmixed sample event vector is an identity matrix.

[0041] According to the second aspect of the present invention, the incremental matrix comprises one or more incremental values ​​for fine-tuning the initial spectral spillover matrix, and wherein the incremental matrix has the same dimension as the initial spectral spillover matrix.

[0042] According to the second aspect of the present invention, the initial spectral spillover matrix comprises N×N dimensions, and wherein N is the number of the single-staining compensation controls.

[0043] According to a second aspect of the invention, the measured sample event vector comprises N values, and wherein the unmixed sample event vector has N values.

[0044] According to a second aspect of the present invention, the re-compensated sample event vector is equal to the inverse of the sum of the initial spectral spillover matrix and the increment matrix multiplied by the measured sample event vector. BRIEF DESCRIPTION OF THE DRAWINGS

[0045] Figure 1 is a block diagram of the detection module of the flow cytometer system.

[0046] Figure 2A is a schematic diagram of an image array with sequential channel re-imaging.

[0047] Figure 2B is a schematic diagram of the 1f image array with alternating channel re-imaging.

[0048] Figure 3 is a schematic diagram of a compact detection module comprising a 16-channel 1f imaging array for a modular flow cytometer system.

[0049] Figure 4A is a magnified view of the fluorescence exiting at different points on the fiber diameter.

[0050] Figure 4B A magnified view of a detector that converts light into electrical signals.

[0051] Figure 5 This is a magnified view of a portion of the 1f image array.

[0052] Figure 6 is a view of a detection module in a modular flow cytometer system having a 16-channel 1f image array and 16 detector channels.

[0053] FIG. 7A to FIG. 7C are different schematic views of a detection module in a modular flow cytometer system having a pair of 8-channel 1f image arrays and a pair of 8 detector channels.

[0054] Figure 8 is a stereogram of a 1f image array with spherical micromirrors.

[0055] FIG. 9A to FIG. 9B yes Figure 8 1f image of a cross-sectional view of the array.

[0056] Fig.10is a stereogram of a 1f image array with concave rectangular micromirrors.

[0057] Fig.11 yes Fig.10 1f image of a top-side view of a portion of the array.

[0058] Fig.12 is a mounting block adjacent to the imaging block and for Figure 6 and FIG. 7A to FIG. 7C A perspective view of a low-cost, low-profile detector in a detector module.

[0059] Fig.13 is a top view of the optical plate assembly in the modular flow cytometry system.

[0060] Fig.14 This is a basic conceptual diagram of a flow cytometer system.

[0061] Fig.15 is the overall approach to performing experiments on a flow cytometer.

[0062] Fig.16A A view of the calibration process of a flow cytometer using single-stained compensation controls to generate an initial spillover matrix, or reference matrix, with multiple compensation levels.

[0063] Fig. 16B A view of a sample passing through a flow cytometer resulting in a mixed sample event vector with overlapping spectral profiles due to multi-stained cells or particles.

[0064] Fig. 16C is a view of the processing using an inverse matrix (determined by an initial spill matrix and / or an initial reference matrix with fine tuning) on ​​the event data to generate a compensated sample event vector or an unmixed sample event vector. DETAILED DESCRIPTION

[0065] In the following detailed description of the present embodiment, many specific details are set forth to provide a thorough understanding. However, it will be apparent to those skilled in the art that the present embodiment can be practiced without these specific details. In other cases, well-known methods, processes, components, and circuits are not described in detail to avoid unnecessarily obscuring various aspects of the embodiments of the present invention.

[0066] Embodiments of the present invention include methods, apparatus, and systems for flow cytometry with compact, highly multiplexed detection modules.

[0067] General Overview

[0068] A flow cytometer with a compact detection module for fluorescence having an increased number of detectors and minimal image size at the detectors compared to existing flow cytometers is disclosed. Each detector module is fed by at least one laser. Multiple lasers can be supported by multiple detector arrays in a compact manner. The number of detectors can be increased by carefully controlling the incremental aberrations in the detector array as light is emitted through the detection chain. The compact size of the compact detection module is achieved by a reduced distance between micromirrors and filters and careful down scaling to minimize image degradation along the rows or chains of micromirrors and filters in the imaging array.

[0069] Existing limitations can be overcome by using an optical system with multiple individual detectors and adjustable levels of filter bandwidth. The optical system focuses the spectral configuration point to optimally resolve the spectra of both long wavelength dyes and short wavelength dyes used to label microparticles analyzed by flow cytometry. The fluorescence from the excited microparticles is imaged into a multimode fiber through an objective lens with a high numerical aperture (NA). The broadband fluorescence leaving the multimode fiber is collimated and then coupled (imaged) into multiple detectors. Collimation of broadband fluorescence leaving a multimode fiber is a challenge.

[0070] Detection modules for flow cytometry

[0071] Figure 1 A functional block diagram of a portion of a detection module 100 of a flow cytometer is shown. An advanced flow cytometer may include multiple detection modules. The detection module 100 is a wavelength demultiplexing system. The detection module 100 continuously reflects and re-images the fluorescence output 101A from the optical fiber 102 in the image array 106. The image array 106 is a mechanical image array that includes a plurality of mirrors 112A-112E suspended in the air by mechanical mounting and a corresponding plurality of long-pass dichroic filters 110A-110E. Typically, a dichroic filter is a precision filter that is used to selectively pass light in a range of wavelengths of colored light while reflecting colored light of other wavelengths. Alternatively, the plurality of long-pass dichroic filters 110A-110E may be bandpass filters.

[0072] The image array 106 is capable of reflectively re-imaging the fiber spot N times (where N is greater than 2) while maintaining the optical quality of the fiber spot at the end of the image array. Re-imaging is a function of recreating the original image (re-imaging, reimage) with some aberrations at a surface such as each of the dichroic filters 110A-110E. The detection module 100 also includes a plurality of detector channels 113A-113E, each of which includes a plurality of objective lenses 116A-116E and a plurality of detectors 118A-118E that are optically connected to each other. Optionally, the plurality of detector channels 113A-113E may also include a plurality of bandpass filters 114A-114E, respectively, to ensure that the plurality of detectors detect each desired wavelength range.

[0073] In the image array 106, light 101A incident on the first reflector 110A is reflected by the reflector into reflected light 103A. The reflected light 103A from the first reflector 110A passes through the air and is incident on the first long-pass dichroic filter 112A. The light 103A from the first reflector 110A is separated by the long-pass dichroic filter 112A into a continuous light portion 101B and a passing or transmitted light portion 105A. The continuous light portion 101B passes through the air and is incident on the next reflector in the series of reflectors, i.e., the reflector 110B. The transmitted light portion 105A is coupled to the first bandpass filter 114A of the first detector channel 113A. The transmitted light portion 105A is cleaned up by the bandpass filter 114A and then coupled to the optical detector 118A by the objective lens 116A. This process is repeated for each stage (detector channel) in the image array 106.

[0074] Image array 106 includes five stages that are re-imaged five times at each long pass dichroic filter. It is still desirable to include a greater number of detectors. However, after more than five re-imagings, the distortion of the beams through the mirrors may accumulate to the point where the image quality at the last dichroic filter becomes highly degraded. In order for the zigzag configuration between the mirrors and the dichroic filters in image array 106 to perform properly with a greater number of detectors, it is desirable to minimize image degradation along the optical path.

[0075] Minimizing image degradation in the detection module can be achieved using two mechanisms. Image degradation can be reduced if the curvature of each mirror in the image array is reduced (e.g., reduced by a factor of two - half the curvature). Image degradation can be further reduced if the number of times the beam is re-imaged in the image array is reduced (e.g., reduced by half - half the number of times).

[0076] Reducing the bend angle of the mirrors can reduce aberrations of all types. Since aberrations increase nonlinearly with the bend angle, the improvement obtained by switching from a "fast focusing mirror" to a "slow focusing mirror" is significantly better than 2X, allowing image quality to be maintained through more reflections. The image quality incident on the last detector in the detector chain can be further improved by reducing the number of times the beam is re-imaged in the image array. Instead of re-imaging on each dichroic filter for each detector channel, the incident light can be re-imaged on every other dichroic filter and detector channel (e.g., odd-numbered detector channels).

[0077] Now refer to FIG. 2A to FIG. 2B 106A-106B are described to provide an improvement in image quality by utilizing a transparent block of micromirrors having different radii of curvature. Typically, the image array consists of an array of micromirrors for each detection channel and an opposing array of bandpass and / or dichroic filters. In each case, the thickness L of the transparent block between the serial chain or row of micromirrors M(n), M(n)' on one side and the serial chain or row of dichroic filters D(n) on the opposite side is the same. However, the focal lengths of the micromirrors M(n) and M(n)' are in the range of 106A-106B. FIG. 2A to FIG. 2B The image arrays 106A-106B are different.

[0078] Figure 2B The focal length f of the micromirror M(n)' in is L, and Figure 2A The focal length f of the micromirrors M(n) in the image array 106B is half of L. The larger focal length of the micromirrors M(n)' in the image array 106B reduces the bending angle and aberrations when imaging along the serial chain of mirrors. Moreover, the image array 106B is a 1f image array having a thickness L of a given transparent block and the focal lengths of the micromirrors M(n)' so that re-imaging occurs on odd-numbered dichroic filters, such as dichroic filters D(3), D(5), D(7) through D(n).

[0079] exist FIG. 2A to FIG. 2B In FIG. 1 , the light spots A(1) to A(n) are the spot sizes (areas) on the dichroic filters D(1) to D(N), respectively. The light spot A(0) is the fiber aperture of the multimode fiber, from which the fluorescence is input into each array 106A-106B. Figure 2A and Figure 2B In , the fiber aperture can be considered to be infinitely small to illustrate the image conjugation properties of the two designs.

[0080] Figure 2AAn image array 106A having a plurality of micromirrors M(1) to M(n) and a plurality of long pass dichroic filters D(1) to D(N) is illustrated, where N is an integer value greater than one and represents the number of detector channels. Light focused at a spot A(1) on the dichroic filter D(1) is re-imaged by reflection from the micromirror M(1) to focus the light at a spot A(2) on the dichroic filter D(2). This process is repeated for each of the micromirrors M(2) to M(N) along a serial chain or row. The image array 106A is a 2f image array.

[0081] Figure 2B A 1f design for an image array 106B having a plurality of micromirrors M(1)' to M(N)' and a plurality of long-pass dichroic filters D(1) to D(N) is illustrated. The micromirrors M(1)' to M(N)' have a radius of curvature different from the radius of curvature of the micromirrors M(1) to M(N). The 1f image array 106B has a plurality of micromirrors M(1)' to M(N)' and a plurality of long-pass dichroic filters D(1) to D(N). Figure 1 Provides image quality improvement over 2f image array 106A at detectors 118A-118E in FIG.

[0082] The chain of micromirrors M(1)' to M(N)' in the 1f image array 106B is designed to relay images along the chain through the properties of telescope optics. For example, light focused at spot A(1) on the dichroic filter D(1) is imaged to spot A(3) on the dichroic filter D(3) through the telescopic effect of micromirrors M(1)' and M(2)'. The even-numbered spot (spot A(2)) is the middle spot in the collimated space.

[0083] The plurality of long pass dichroic filters D(1) to D(N) may alternatively be passband or bandpass filters, or may include both dichroic filters and bandpass filters combined together to ensure that a limited range of wavelengths of light pass through. Dichroic filters use the principle of thin film interference and may also be referred to as interference filters. For each channel, the bandpass or passband filters are tuned to transmit light of a different selected wavelength range (passband) to each detector and reflect the remaining wavelengths of light back to the micromirrors in the micromirror array.

[0084] To provide a compact detector module, the image array 106B is formed of a solid transparent material, as shown in FIG. Figure 8 and Fig.10 Further explained, the solid transparent material of the transparent imaging block used for the 1f image array may be transparent glass or transparent plastic, for example, with reflectors and dichroic filters formed in or on the transparent material.

[0085] The design of 1f image array 106B can be compared with the design of 2f image array 106A under the conditions of the same thickness of solid transparent material, the same pitch and the same angle of incidence. The path distances between adjacent micromirrors in 2f image array 106A and 1f image array 106B are similar. In this case, the path distance is a physical distance that takes into account the refractive index of the material, rather than a conventional path length. However, the focal length of the micromirrors in 2f image array 106A is shorter (half) than the focal length of the micromirrors in 1f image array 106B. In other words, due to the different radii of curvature of the micromirrors, the actual focal length of the micromirrors in 1f image array 106B is twice the focal length of the micromirrors in 2f image array 106A. A longer focal length reduces the curvature of each reflection, which minimizes the aberrations introduced in the micromirror reflection. Therefore, the aberrations in 1f image array 106B are improved compared to the aberrations in 2f image array 106A.

[0086] In the 2f image array system 106A, the fiber aperture A(0) is imaged to spot A(1) on the dichroic filter D(1). The image at spot A(2) is a re-image of spot A(1) reflected by micromirror M(1). Continuing through the zigzag optical path in the array 106A, each spot A(n) at each dichroic filter D(n) is imaged by the next micromirror M(n) to the next spot A(n+1) on the next dichroic filter D(n+1). In this configuration, the path distance between spot A(n) and micromirror M(n) is the thickness L of the transparent block, which is 2 times (twice) the focal length of micromirror M(n).

[0087] In the 1f image array system 106B, light from the fiber aperture A(0) is imaged to the spot A(1) on the dichroic filter D(1) through the input channel. In the image array system 106B, the light path from micromirror M(1)' to micromirror M(2)' can be considered to form an equivalent telescope with a magnification of 1. In this case, through the telescopes of micromirrors M(1)' and M(2)', the spot A(1) is imaged to the spot A(3), where the spot A(2) is the middle spot in the collimation space. The adjacent micromirrors M(3)' and M(4)' form another telescope to re-image the spot A(3) to the spot A(5). Continuing the zigzag path in the imaging array 106B, the odd-numbered spots A(1), A(3), A(5), ..., A(2n+1) are all conjugate to each other, while the even-numbered spots A(2), A(4), A(6), ..., A(2n) are intermediate spots in the collimated space. In this configuration, the path distance between the spot A(n) and the micromirror M(n)' is the thickness L of the transparent block, which is one focal length of the micromirror M(n)'.

[0088] Each spot on the filter is formed by a beam of light. The angular distribution of the beam is determined by the numerical aperture (NA) of the input multimode fiber 102 and the input channels of the imaging array for the imaging system. The light cone angle at the spot A(1) is proportional to the numerical aperture of the multimode fiber 102. If the image magnification m from the fiber aperture A(0) to the spot A(1) is a ratio of 1 to m (where m is greater than 1), then the cone angle at the spot A(1) is one-m smaller than the cone angle at the fiber aperture A(0).

[0089] exist Figure 2A In the 2f image array system 106A, for any value of n, the image magnification from any spot A(n) to the adjacent spot A(n+1) is equal to 1. In the absence of visual aberrations, the light cone angles at the dichroic filters D(1) to D(n+1) of the image array system 106A are the same for all detector channels.

[0090] Now refer to Figure 2B , the distance between the dichroic filter D(1) and the micromirror M(1)' is L. If the total number of channels N in the 1f image array 106B is an even number, the light spot A(1) is imaged to the light spot A(3) through the micromirrors M(1)'-M(2)' (considering a 1X telescope), and the light spot A(2) on the dichroic filter D(2) is in the collimated space. A portion of the collimated light at the light spot A(2) is reflected by the dichroic filter D(2) toward the micromirror M(2)'.

[0091] Therefore, the 1f image array 106B has odd-numbered detector channels at spots A(1), A(3), ..., A(2k-1), ... to A(N-1), and even-numbered detector channels at spots A(2), A(4), ..., A(2k), ... to A(N), where 1≤k≤(N / 2). Since spot A(1) is at the front focus of micromirror M(1)' and the path distance between adjacent micromirrors is twice the focal length, all odd-numbered spots A(1), A(3), A(5), ..., A(2n+1) are images of the fiber aperture (e.g., Figure 2B ), while all the even-numbered spots A(2), A(4), A(6), …, A(2n) are located in the collimated space (as shown in Figure 2B , as shown by the parallel rays at the even-numbered spots. Figure 5 As shown, the cone angle of the odd-numbered spots of the odd-numbered detector channels (odd cone angle CAO) is different from the cone angle of the even-numbered spots of the even-numbered detector channels (even cone angle CAE).

[0092] In the 1f image array 106B, the center wavelength and passband width of each optical dichroic filter D(1) to D(N) are different from each other. The center wavelength and passband width of each optical dichroic filter are designed to optimize the sampling of the dye fluorescence spectrum, thereby better accurately unmixing a large number of different dyes. For example, assuming a fluorescence spectrum with a wavelength of light of 400nm to 800nm ​​and sixteen (16) channel detection modules, each of which can analyze a bandwidth of 25nm of light wavelength. For example, the first detector channel and the first dichroic filter D(1) can pass and analyze light wavelengths of 400nm to 425nm with a center wavelength of 412.5nm. Wavelengths outside of 400nm to 425nm are basically filtered out and not transmitted to the first detector in the first detector channel. The second detector channel and the second dichroic filter D(2) can pass and analyze light wavelengths of 425nm to 450nm with a center wavelength of 437.5nm, and so on for each detector channel. The last or sixteenth detector channel and sixteenth dichroic filter D(16) can bandpass and analyze light wavelengths from 775 nm to 800 nm with a center wavelength of 787.5 nm.

[0093] The characteristics of the 1f image array 106B allow the initial light signal to propagate to a greater number of detectors than the 2f image array 106A. The 1f image array 106B reduces off-axis aberrations by reducing the curvature in each mirror reflection. However, in the case where the odd cone angles CAO and the even cone angles CAE at the dichroic filters in the odd and even channels are different, it is necessary to determine the optimal magnification m in the input stage from the fiber aperture A(0) to the light spot A(1) at the dichroic filter D(1). For a given fiber numerical aperture (NA) and aperture diameter, the magnification m from the fiber aperture A(0) to the light spot A(1) at the dichroic filter D(1) is optimized for the odd and even detector channels in the 1f image array 106B.

[0094] From the perspective of spectral resolution, as the cone angle of the incident light spot increases, the performance of the dichroic filter deteriorates. In the 1f image array 106B, the cone angle of the light spot in the odd channel is different from the cone angle of the light spot in the even channel. Basically, the cone angle in the odd channel is determined by the numerical aperture (NA) of the multimode fiber and the magnification factor m of the light spot A(1) from the fiber aperture A(0) to the dichroic filter D(1). In contrast, the cone angle of the even channel is determined by the diameter of the light spot in the odd channel.

[0095] In the input channel, assume that the image magnification from the aperture A(0) at the fiber to the light spot A(1) at the dichroic filter D(1) is m. At the even detector channels (dichroic filter D(2k)), the cone angle is proportional to the magnification m. However, in the odd detector channels (dichroic filter D(2k-1)), the cone angle is inversely proportional to the magnification m. The larger magnification from the fiber aperture A(0) to the light spot A(1) results in a smaller cone angle at the odd detector channels (dichroic filter D(2k-1)), but a larger cone angle at the even detector channels (dichroic filter D(2k)). In an example embodiment of a multimode fiber with NA=0.12, an aperture diameter of 600 μm, and a micromirror (filter) pitch of 5.5 mm, the recommended magnification m is modeled as approximately 2. Obviously, other magnifications m can be determined using different inputs, so the embodiments disclosed herein are not limited to 2X magnification. In the example presented for 1f image array 106B, the numerical aperture (NA) and the number of re-images are both reduced by a factor of two, thereby allowing at least four times (4X) more detectors along a row of the same length compared to 2f image array 106A.

[0096] Figure 3 , FIG. 4A to FIG. 4B as well as Figure 5 Illustrated are simulation results for a compact detection module with a 1f image array 106B and a graph of example input values ​​for an implementation with NA=0.12, 600 μm aperture diameter, 5.5 m micromirror (filter) pitch fiber, and a recommended magnification of 2X. Figure 3 , FIG. 4A to FIG. 4B as well as Figure 5 The different colors of light shown in FIG. 5 are for clarity only to illustrate how light at different positions passes through the detector module.

[0097] Figure 3 One end of the optical fiber 102 is shown that transmits the fluorescence into the detection system. Near the opposite end of the optical fiber 102 (not shown), a light collecting objective with a high NA can be used to collect the fluorescence from the aperture and couple it into the opposite end of the optical fiber. The optical fiber 102 then collects the light from the objective and directs it to the Figure 3 End shown. Close Figure 3 As shown in the end, the system can include a fiber optic numerical aperture converter to reduce the numerical aperture to free space in order to transmit fluorescent light to the detector array.

[0098] Now refer to Figure 3, the magnification m in the compact detection module 300 is implemented by the input stage 301. The input stage 301 includes: a collimating lens 302, a blocking filter 303, and a focusing lens 304. The magnification m is achieved by adjusting the focal length ratio of the collimating lens 302 and the focusing lens 304. For example, in order to set the magnification to be equal to 2 (m=2), the focal length of the focusing lens 304 is twice the focal length of the collimating lens 302. The input channel 301 can be considered to also include the input portion of the transparent block in the image array 106B before reaching the first dichroic filter D(1) (e.g., wedge shape, block thickness, see Figure 8 ).

[0099] Collimating lens 302 receives light emitted from optical fiber 102 and collimates the light. The collimated light passes through blocking filter 303 and is transmitted to focusing lens 304. Blocking filter 303 is used to remove laser light scattered into the collection optical device near the opposite end of optical fiber 102. Light of the fluorescence spectrum associated with fluorescein passes through blocking filter 303 and enters focusing lens 304. Focusing lens 304 focuses the light of the fluorescence spectrum onto the first dichroic filter D(1) in image array 106B to form an image at spot A(1). The size (e.g., diameter and area) of the image at spot A(1) is magnified m times compared to the size at aperture A(0) at the optical fiber. The position of lenses 302, 304 between the end of optical fiber 102 and image array 106B can be adjusted.

[0100] The compact detection module 300 also includes sixteen (16) channels in communication with the image array 106B and the image array 106B and sixteen (16) detector channels 313A-313P (see, e.g., Figure 6 In an alternative embodiment, a pair of eight (8) channel 1f image arrays may be used in parallel (see, for example, FIG. 7A to FIG. 7C ) to relax the imaging requirements of each compact image array. In a flow cytometer, more than one (e.g., three) of these compact image arrays can be used to multiply the number of detector channels to greater than sixteen (e.g., three times sixteen for forty detection channels, as shown in FIG. Fig.13 as explained in the detector module described in ).

[0101] The image array 106B is formed from a solid transparent block of material. The sixteen (16) channel 1f image array 106B includes sixteen (16) dichroic filters D(1) through D(16) on one side of the transparent block and fifteen mirrors M(1) through M(15) on the opposite side. The image array does not require a mirror following the last detector channel 313P. In addition, the last filter D(16) 314 may not be a dichroic filter; instead, a bandpass filter may be used. In the case of a bandpass filter, the incident light does not need to be further reflected to another mirror or filter.

[0102] Each detector channel 313A-313P (collectively referred to as detector channels 313) in the array or detector includes a focusing lens 316 and a detector 318 ( Figure 3 ). The detector 318 is packaged in a thin outline (TO) can package 320, to which the focusing lens 316 is coupled or integrated. The focusing lens 316 focuses the fluorescence passing through the filter onto the small area size of the detector 318.

[0103] Now refer to Figure 4A The optical fiber 102 used to transmit the fluorescence signal captured from the image chamber to the detector array is a multimode optical fiber. Light leaves the end face of the multimode optical fiber from various (if not all) positions on the diameter of the optical fiber (such as positions X1 to X5, for example). Figure 3 Lenses 303, 304 are shown in the input channel focusing light into aperture A(0) to spot A(1) on the first dichroic filter D(1). Because there is a two-fold (2X) magnification from spot A(0) to spot A(1), the spot size at aperture A(0) is smaller than the spot size at spot A(1). Figure 4A The different colors of light emitted from different positions X1 to X5 within the aperture shown in FIG. 1 are only for clarity, to illustrate how light at different positions passes through the detector module. Figure 4A As shown, an optical axis 402 extends outwardly from a circular center of the end of the optical fiber 102. Light is emitted from the end of the optical fiber 102 at a launch cone angle (CA) 404 relative to the optical axis 402.

[0104] Figure 3 The simulation results of the image array 106B are shown and how light from different positions in the beam is alternately imaged and collimated through multiple reflections from mirrors and dichroic filters. Although these results show all light reflections, the dichroic filters D(n) at any particular position are different and allow light signals to be transmitted according to their respective passbands (only at Figure 3 The last detector channel 313P is shown in FIG.

[0105] Now, refer to Figure 3 and Figure 4B In each detector channel 313, the optical signal of the desired wavelength range passing through the dichroic filter D(n) can be collected by a lens 316 and detected by a small aperture photosensitive detector 318. Another bandpass filter 314 can be used alternatively or further in each detector channel. Light of other wavelengths at the dichroic filter D(n), if any, is reflected along the chain or row of micromirrors to the next micromirror M(n). The dichroic filters D(n) of the row or chain demultiplex the different ranges of optical wavelengths into the chain of detector channels 313A-313P.

[0106] Figure 5 The enlarged figure shows the simulation results of how the light beam is alternately imaged and collimated by the reflective surface of the micromirror and the dichroic filter. Figure 5 On the dichroic filters D(7), D(9), and D(11) shown in FIG. 1 , the light spot is an image of the fiber aperture. That is, the light emitted from the fiber aperture is imaged onto each filter surface of the odd-numbered dichroic filters. On the even-numbered dichroic filters (e.g., Figure 5 On the dichroic filters D(8), D(10) and D(12) shown in the figure, the even-numbered spots ( Figure 5 The light spots A(8), A(10) and A(12) are shown in collimated space, where light emitted from a point at the fiber aperture becomes a collimated beam. The beam direction in the collimated space at each even-numbered dichroic filter is slightly different for different points from the fiber aperture.

[0107] In the flow cytometer, one or more linear 16-channel compact wavelength detection modules can be used to detect the fluorescent signal of light associated with the microparticles. Alternatively or in combination, one or more dual 8-channel compact wavelength detection modules can be used in the flow cytometer to detect the fluorescent signal of light associated with the microparticles.

[0108] Figure 6 and FIG. 7A to FIG. 7C An embodiment of a compact wavelength detection module is illustrated, which has Figure 2B The functionality of the 1f image array 106B is shown. Figure 6 A linear 16-channel compact wavelength detection module 600 is illustrated. FIG. 7A to FIG. 7C A dual 8-channel compact wavelength detection module 700 is illustrated.

[0109] Now refer to Figure 6, the linear 16-channel compact wavelength detection module 600 includes an input stage (end) 601 and a detection module 614 mounted to a base 610. Light is coupled into the input stage (end) 601 through an optical fiber 102. The input stage (end) 601 includes a collimating lens 602 mounted to an optical bench, a long pass filter 603, a clean-up optical blocker 604, and a focusing lens 605. The input stage (end) 601 sets the magnification m of the initial spot size image A(1) on the first dichroic filter.

[0110] From the input stage 601, light is coupled into the detection module 614. One end of the input stage (head) 601 is coupled to a transparent wedge 607 to receive light from the focusing lens 605. The input stage (head) 601 and the detection module 614 are coupled to the chassis or base 610 of the flow cytometer to maintain their alignment.

[0111] The detection module 614 includes an image array 608 and a detector / lens array 611. The image array 608 is Figure 2B and Figure 5 The image array 608 includes a transparent block 680 (see, for example, Figure 8 and Fig.10 806, 1006), the transparent block includes an optical wedge 607 and fifteen micromirrors 612 on one side. On the opposite side of the transparent block 680, there are sixteen dichroic filters 609. The detector / lens array 611 (an embodiment of the plurality of detectors 313A-313P) includes a plurality of photodetectors D1 to D16 (e.g., Figure 3 Detectors 318), each photodetector having a lens (e.g., Figure 3 A lens 316 is provided to focus the demultiplexed light into the photodetector.

[0112] Light coupled into the image array 608 by the input stage 601 is wavelength demultiplexed into detectors D1 through D16 of the detector / lens array 611. The 16-channel detection module analyzes a range of wavelengths (eg, 400 nm to 800 nm wavelength).

[0113] To provide a better fit to different footprints of test platforms and to provide parallel processing, the linear 16-channel compact wavelength detection module 600 may instead be implemented as a dual 8-channel compact wavelength detection module.

[0114] Now refer to Fig. 7A, shows a top view of a dual detection module 700 having a pair of 8-channel compact wavelength detection modules 714, 715. The compact wavelength detection module 700 includes an input stage (header) 701 in communication with a first 8-channel detection module 714 and a second 8-channel detection module 715, all of which are mounted to align with a base 710. The first 8-channel detection module 714 demultiplexes and analyzes a first range of wavelengths (e.g., 650nm to 800nm ​​- red wavelengths) in parallel. The second 8-channel detection module 715 demultiplexes and analyzes a second range of wavelengths (e.g., 400nm to 650nm - blue wavelengths) in parallel.

[0115] Light emitted from the optical fiber 102 is coupled into the input stage (end) 701. Light from the optical fiber 102 passes through the collimating lens 702 into the long pass dichroic filter 703. The long pass dichroic filter 703 reflects light at the laser excitation wavelength (e.g., less than 400nm) at a 45 degree angle to a scattering detector (not shown). Side scattered (SSC) light can be focused onto a small aperture detector with a ball lens similar to that described for fluorescence. Fluorescence in the fluorescence spectrum (e.g., 400nm-800nm) passes through the long pass filter 703 and enters the second clean-up filter 704. The clean-up filter 704 ensures that no excitation laser reaches the demultiplexing detection modules 714-715.

[0116] After the purification filter 704, the fluorescence is separated into a long wavelength band and a short wavelength band by a long pass filter 705. The long wavelength light (e.g., red light - 650nm to 800nm) passes through the long pass filter 705 and is focused by the collimating / focusing lens 706 into the first detection module 714. The long wavelength portion of the light that passes through and is focused by the focusing lens 706 is demultiplexed by the first detection module 714. The short wavelength light band (e.g., blue light - 400nm to 650nm) is reflected by the long pass filter 705 at an angle back to the collimating / focusing lens 713. The collimating / focusing lens 713 focuses the light of the short wavelength band into the second detection module 715. The short wavelength portion reflected by the long pass filter 705 is demultiplexed by the second detection module 715. Alternatively, the filter 705 may be a short-pass filter, and short-wavelength light passes through the filter and is demultiplexed by the first detection module 714 , while long-wavelength light is reflected by the filter and demultiplexed by the second detection module 715 .

[0117] Referring to the first detection module 714, light from the focusing lens 706 enters the 12 degree wedge 707 vertically and passes through a transparent block (e.g., Figure 8The light passes through the bandpass filter 709 and is focused onto the first small area detector D1 in the detector / lens array 711. The light rejected by the bandpass filter 709 is reflected back onto the first micromirror M(1) 712 of the plurality of micromirrors M(1) to M(7) in the image array. The first micromirror M(1) 712 collimates the light and reflects it back to the second detection module D2, and so on, to the end of the serial chain of micromirrors and detection modules of the transparent block of the image array 708. The second detection module 715 functions similarly to the first detection module 714.

[0118] The reflections proceed through the image array 106B, where the light is alternately focused and collimated as described herein at 708, 708' in each of the first detection module 714 and the second detection module 715, and successively shorter bandpass light passes through the dichroic filter and enters the odd and even detectors 118 in the odd and even detector channels, respectively. Thus, different wavelengths are demultiplexed by the multiple detectors in each of the first detection module 714 and the second detection module 715.

[0119] For a given fluorescence event, the number of Figure 4B The detector 318 shown, Figures 6 to 7A The signals from the lenses / detectors D1 to D16 in the optical module are amplified, digitized and synchronized by the electronic system to provide a spectral representation of the input fluorescence signal. Integrating the detection electronics into the optical module assembly allows for compact design and lower noise by minimizing the connection length between the detector and the amplification circuit. Figure 4B The detector 318 shown in FIG. 3 converts an optical signal (such as an input fluorescent signal) into an electrical signal.

[0120] Figure 7B and Figure 7C The right and left perspective views of a dual detection module 700 having a pair of 8-channel compact wavelength detection modules 714, 715 are illustrated, respectively. Each detection module 714, 715 includes a mounting base 720 and a cover 722 to enclose a mounting block 1200 (see FIG. 1 ) to which the lens / detector 711 in the detector array or chain is mounted. Fig.12 ). The mounting base 720 and cover 722 align the components of the image array 708, 708' in the transparent block 806, 1006 together with the detector array in each detector module 714, 715. The mounting base 720 of each detection module 714, 715 is coupled to the base 710 by a plurality of fasteners.

[0121] The input stage 701 includes an optical bench 751 having: a plurality of filter slots for accommodating filters 703-705; a plurality of lens slots for accommodating lenses 702, 706, 713; and one or more optical channels along which light is reflected and propagated through the filters and lenses. The optical bench 751 is coupled to the base 710 of the detection module 700 to maintain alignment with the detection modules 714-715.

[0122] Fig.13 A top view of an optical plate assembly 1300 in a modular flow cytometry system 100 is shown. The optical plate assembly 1300 includes a laser system 1370 having three semiconductor lasers 1370A, 1370B, 1370C, which direct excitation into a flow cell assembly 1308, wherein a sample fluid flows with sample particles. The laser system 1370 attempts to direct multiple (e.g., three) laser beams toward the flow cell assembly 1308 in a co-linear manner. However, the multiple laser beams may be slightly offset from each other. The laser system 1370 includes semiconductor lasers 1370A, 1370B, 1370C having wavelengths typically of about 405 nanometers (nm), 488 nm, and 640 nm. The output power of a 405 nm semiconductor laser is typically greater than 30 milliwatts (mW); the output power of a 488 nm semiconductor laser is typically greater than 20 mW; and the output power of a 640 nm semiconductor laser is typically greater than 20 mW. The controller electronics control the semiconductor laser to operate at a constant temperature and a constant output power.

[0123] The optical system spatially manipulates the optical laser beams 1371A, 1371B, 1371C generated by semiconductor lasers 1370A, 1370B, 1370C, respectively. The optical system includes lenses, prisms, and steering mirrors to focus the optical laser beams onto a fluid stream carrying a biological cell (bio cell). The focused optical laser beam size is typically focused across the flow stream to 50 micrometers (μm)-80 μm, and is typically focused along the flow flowing in the flow cell assembly 1308 to 5 μm-20 μm. Fig.13, the optical system includes beam shapers 1330A-1330C, which receive laser light 1371A, 1371B, 1371C from semiconductor lasers 1370A-1370C, respectively. The laser light output from the beam shapers 1330A-1330C is coupled into the mirrors 1332A-1332C, respectively, to direct the laser light 1399A, 1399B, 1399C toward and into the flow cell assembly 1308 to target particles (e.g., biological cells) stained with fluorescein dye. The laser light 1399A, 1399B, 1399C is slightly separated from each other, but directly enters the flow cell assembly 1308 approximately in parallel through the mirrors 1332A-1332C.

[0124] Laser beams 1399A, 1399B, 1399C reach biological cells (particles) in the flow stream in the flow cell assembly 1308. Laser beams 1399A, 1399B, 1399C are then scattered by the cells in the flow stream, causing the fluorescein to fluoresce and produce fluorescence. Forward scatter diode 1314 collects the on-axis scattered light. Collecting lens 1313 collects the off-axis scattered light and the fluorescence and directs them together to beam splitter 1310. Beam splitter 1310 focuses the off-axis scattered light onto side scatter diode 1315. Beam splitter 1310 focuses the fluorescence onto at least one optical fiber tip 1316. At least one optical fiber assembly 102 routes the fluorescence toward at least one detector module 600, 700.

[0125] In order to perform more detailed analysis of biological samples using different fluorescent dyes and laser wavelengths, multiple fiber optic tips 1316, multiple fiber optic assemblies 102, and multiple detector modules 600, 700 can be used. Three fiber optic tips 1316A, 1316B, 1316C can be arranged in parallel to receive fluorescence, and three fiber optic assemblies 102A, 102B, 102C can be used to guide the fluorescence to three detector modules 600A, 600B, 600C or 700A, 700B, 700C.

[0126] The three fiber optic ends 1316A, 1316B, 1316C (and the three fiber optic assemblies 102A, 102B, 102C) are enabled because the three laser beams 599A, 599B, ​​599C are slightly offset (e.g., not exactly collinear). Therefore, the three fiber optic ends 1316A, 1316B, 1316C can collect beam data from the three laser beams 599A, 599B, ​​599C having three different wavelengths, respectively. The three fiber optic assemblies 102A, 102B, 102C then direct the light to three different detector modules (e.g., three different detector modules 600A, 600B, 600C or 700A, 700B, 700C).

[0127] Alternatively, the modular flow cytometry system can use one detector module 600, 700 to collect beam data. For example, three fiber optic assemblies 102A, 102B, 102C can direct light into one detector module 600, 700, rather than three different detector modules. The separation of the beam data is then performed as a data processing operation, rather than using three different detector modules to separate the beam data. From a physical device perspective, using one detector module may be less complex. However, the data processing operation may be more complex because the separation of the beam data requires more data manipulation (e.g., identifying different wavelengths and separating the beam data accordingly).

[0128] Cell geometry can be classified by analysis of forward and side scatter data. Cells in the fluid stream are labeled with dyes having visible wavelengths ranging from 400nm to 900nm. When excited by the laser, the dyes produce fluorescence, which are collected by the fiber optic assembly 102 and routed toward the detector modules 600, 700. The modular flow cytometry system maintains a relatively small size for the optical plate assembly via compact semiconductor lasers in the detector modules 600, 700, a light collecting lens 1313 of 11.5X focal length, and a compact image array.

[0129] The collecting lens 1313 contributes to the design of the detector modules 600, 700. The collecting lens 1313 has a short focal length for a focal length of 11.5X. The collecting lens 1313 (objective lens) has a high numerical aperture (NA) of about 1.3 facing the fluorescent emission to capture more photons in the fluorescent emission over a wide range of incident angles. The collecting lens 1313 has a low NA of about 0.12 facing the collection fiber 102 to emit fluorescence into the fiber over a narrow cone angle. Therefore, the collecting lens 1313 is converted from a high NA on one side to a low NA on the opposite side to support the magnification m in the input channel of the detector modules 600, 700.

[0130] The core diameter of the collection fiber assembly 517 is between about 400 μm and 800 μm, and the fiber NA is about 0.12 for a core diameter of about 600 μm. The fiber output end can be tapered to a core diameter between about 100 μm and 300 μm for controlling the imaging size onto the receiving photodiode.

[0131] The input end of the collection fiber 102 may also include a lensed fiber end to increase the collection NA, thereby allowing the use of a fiber core diameter of less than about 400 μm. Because the optical fiber 102 has the flexibility to deliver light anywhere in the flow cytometer system, using optical fiber for fluorescence collection enables optimization of the location of the receiver assembly and electronics for a compact flow cytometer system.

[0132] In order to manufacture a low-cost flow cytometer, lower-cost components can be introduced. The image array 106B in each detection module 614, 714, 715 is formed of solid transparent material to provide a reliable, low-cost and compact detection module. Moreover, the flow cytometer uses a low-cost, off-the-shelf thin-profile (TO) can detector.

[0133] Now refer to Fig.12 , in transparent blocks 806, 1006 (see Figures 8 to 11 ) shows a mounting block 1200 near the top of the housing 1200, which are connected together to form a mounting base 720 and a cover 722 to be mounted to the FIG. 7A to FIG. 7C The base 710 of the compact detector module 700 is shown with the 1f imaging array 708, 708'. The mounting block 1200 includes a plurality of angled curved openings 1201 to accommodate a plurality of TO-can lens / detectors 711. The alignment of the mounting block 1200 with the transparent blocks 806, 1006 of the imaging array 708 and the angles of the angled curved openings 1201 allow light reflected from the micromirror 712E to be bandpass filtered by the dichroic filter 709E and coupled to the lens / detector 711E.

[0134] Each TO can lens / detector 711 includes a focusing lens 1211 and a low cost TO can detector 1212 coupled together. The TO can detector 1212 includes a window top and a semiconductor photodetector 1213 inside the TO can package. The semiconductor photodetector 1213 is electrically coupled to a plurality of electrical pins 1214 that extend outside the TO can package to which other electronic components of the flow cytometer are electrically coupled. Figure 4B Similar to the illustrated detector 318 , the semiconductor photodetector 1213 converts an optical signal (such as an incoming fluorescent signal) into an electrical signal on at least one electrical pin 1214 .

[0135] Now refer to Figure 8 , a perspective view of a transparent block 806 formed of a solid transparent material 800 for an embodiment of the 1f image array 106B, 608, 708, 708' is shown. The solid transparent material 800 for the transparent block 806 can be, for example, transparent glass or transparent plastic. A plurality of micromirrors 810 in a row and a serial chain are formed in or on one side of the transparent block 806 of the transparent material 800. A plurality of dichroic or bandpass filters 812 in a row and a serial chain are formed in or on the opposite side of the transparent block 806 of the transparent material 800. Each dichroic or bandpass filter 812 is tuned to a different wavelength range of light to allow detection of a wide range of fluorescence emitted by fluorescein. In one embodiment, the plurality of micromirrors 810 are concave spherical mirrors.

[0136] The transparent block 806 formed of solid transparent material 800 also includes a 12 degree wedge surface 820 to receive light from a focusing lens, such as described with reference to the 1f imaging array 708. The light enters perpendicular to the surface of the wedge surface 820 and is directed (turned) toward the first dichroic or bandpass filter D(1). The light passes through the transparent block 806 to the first dichroic or bandpass filter D(1).

[0137] Now refer to Fig.10 , a perspective view of a transparent block 1006 formed of a solid transparent material 800 for another embodiment of the 1f image array 106B, 608, 708, 708' is shown. The solid transparent material 800 for the transparent block 1006 can be, for example, transparent glass or transparent plastic. A plurality of micromirrors 1010 are concave rectangular reflectors formed into one side of the transparent material. A plurality of dichroic or bandpass filters 1012 are formed into or on the opposite side of the transparent material 800. Each dichroic or bandpass filter 812 is tuned to a different wavelength range of light to allow detection of a wide range of fluorescence emitted by fluorescein.

[0138] The solid transparent material 800 also includes a 12 degree wedge surface 820 to receive light from a focusing lens, such as described with reference to the image array 708. The light enters perpendicular to the surface of the wedge surface 820 and is directed (turned) toward the first dichroic or bandpass filter D(1). The light passes through the transparent block to the first dichroic or bandpass filter D(1).

[0139] Now refer to Fig.9A , the cross-sectional view illustrates the distance (eg, thickness L) between the spherical micromirror 810 on one side of the transparent block 806 and the opposite side. An axis 814 perpendicular to the transparent block at the center of the spherical micromirror 810 extends to the opposite side of the transparent block 806. Fig. 9B An axis 815 is illustrated perpendicular to the transparent block at the center of the dichroic or bandpass filter 812. The axis 815 extends to the opposite side of the transparent block 806 of transparent material 800. The axes 814 and 815 are parallel to each other.

[0140] Reflective material 811 is formed (eg, disposed) over the spherical transparent microlens shapes of solid transparent material 800 to form individual spherical micromirrors 810 on one side of transparent block 806. Dichroic or bandpass filter 812 is coupled to material 800 on the opposite side of transparent block 806.

[0141] Fig.11 Similarly shown are the distance and axis 1014 between the concave rectangular micromirror 1010 on one side of the transparent material 800 forming the transparent block 1006 and the opposite side. Fig.11Also illustrated is an axis 1015 that is perpendicular at the center point of the dichroic or bandpass filter 1012. The optical axis 1015 extends to the opposite side of the transparent block 1006 formed of the transparent material 800. The optical axes 1014 and 1015 are parallel to each other.

[0142] Fig.11 Also shown is reflective material 1011 formed (eg, disposed) on the curved transparent rectangular shape of solid transparent block 1006 formed from transparent material 800 to form rectangular micromirrors 1010. Dichroic or bandpass filters 1012 are coupled to opposite sides of solid transparent block 1006.

[0143] Fluorescent dyes used in flow cytometry applications cover the entire visible and near infrared wavelength range. For long wavelength fluoresceins, the emission wavelength bandwidth is generally larger. Each dichroic or bandpass filter 812 can have its detector filter passband and center wavelength optimized to measure different dyes with the same amount of spectral sampling. Moreover, individual filter optimization allows the exclusion of excitation wavelengths from other lasers. In this way, the detector in each channel can be fully utilized to detect the signal of interest. In combination with the fluorescence spectral unmixing algorithm executed by the processor of the computer, individual and optimized passband detection provides ultimate detection of a large number of fluorescent dyes of interest.

[0144] Methods using various detection systems

[0145] The following is a description of methods for using the various detection systems disclosed herein in a flow cytometer. Before the fluorescence generated by the laser-excited fluorescein is emitted from the end of the optical fiber 102 shown in the figure, fluorescence of different wavelengths is generated by the various fluorescein excited by the laser, and these fluorescein mark different particles in the sample in the flow channel. The generated fluorescence is received by a collecting lens near the opposite end of the laser, such as can be Fig.13 A converter is used to convert from a first numerical aperture on the capture side to a second numerical aperture that is smaller than the first numerical aperture to better match the numerical aperture of the optical fiber. The optical fiber then directs the fluorescence toward the end of the optical fiber to flexibly direct it toward the compact detection module 600, 700.

[0146] The optical fiber 102 couples the fluorescence to the end of the optical fiber, thereby emitting the fluorescence from the optical fiber. The emitted fluorescence has different wavelengths generated by different fluoresceins that have been excited by the laser and are attached to different particles in the sample fluid.

[0147] In the input channel, light emitted from the end of the optical fiber is collimated and focused by a lens toward a first dichroic filter of a first plurality of dichroic filters in a first demultiplexing imaging array.

[0148] Further along the input channel, the laser light emitted from the optical fiber used to excite the different fluorophores is blocked by a blocking device so as not to interfere with the wavelengths of the detected fluorescence.

[0149] Further along the input path, the image size from one end of the optical fiber is magnified to a spot size for a first dichroic filter in a serial chain or row of a first plurality of dichroic filters in a first demultiplexed imaging array.

[0150] In the first demultiplexed imaging array, alternatively, fluorescence of the first wavelength range is reflected between the first plurality of dichroic filters and a serial chain or row of the first plurality of micromirrors to collimate the fluorescence on odd-numbered dichroic filters and to re-image the fluorescence on even-numbered dichroic filters. The focal length of the first plurality of micromirrors and the distances between the first plurality of dichroic filters and the first plurality of micromirrors provide a telescope effect along the chain of micromirrors to collimate the fluorescence on odd-numbered dichroic filters and to re-image the fluorescence on even-numbered dichroic filters.

[0151] In the serial chain or row of the first plurality of dichroic filters, a different wavelength range of the fluorescent light of the first wavelength range is band-passed at each dichroic filter to demultiplex the wavelength spectrum of the fluorescent light of the first wavelength range.

[0152] Adjacent to the first plurality of serial chains or rows of dichroic filters is a serial chain or row of a plurality of detector channels having a first plurality of light detectors or photodetectors, such as Figure 3 , Figure 6 as well as FIG. 7A to FIG. 7C Each detector channel has a lens to focus fluorescence of a different wavelength range into a first plurality of light detectors or photodetectors.

[0153] The serial chain or row of multiple photodetectors detects fluorescence in each different wavelength range in the first wavelength range associated with each fluorescein labeled to the microparticle. The multiple photodetectors convert the fluorescence received by each into an analog electrical signal. Each analog electrical signal can then be digitized / converted from an analog electrical signal to a digital electrical signal by an analog-to-digital (A / D) converter and then analyzed and counted.

[0154] As the fluorescence is converted into an analog electrical signal by a detector and digitized into a digital electrical signal, a computer with a processor can then be used to count the number of each different particle in the sample fluid, as disclosed in application No. 15 / 498,397, entitled COMPACT MULTI-COLOR FLOW CYTOMETER, filed by David Vrane et al. on April 26, 2017, which is incorporated herein by reference.

[0155] The second and / or third demultiplexed imaging arrays may be used in parallel with the first demultiplexed imaging array. In this case, the method further comprises separating the fluorescence into fluorescence of the first wavelength range for the first demultiplexed imaging array, fluorescence of the second wavelength range for the second demultiplexed imaging array, and / or fluorescence of the third wavelength range for the third demultiplexed imaging array. Fig.13 As shown, the first optical fiber 102A can be used to guide the fluorescence toward the first demultiplexing imaging array. The second optical fiber 102B can be used to guide the fluorescence toward the second demultiplexing imaging array. The third optical fiber 102B can be used to guide the fluorescence toward the third demultiplexing imaging array.

[0156] The steps described here for the first demultiplexed imaging array may be performed simultaneously by the second and / or third demultiplexed imaging arrays, so that different additional wavelength ranges may be analyzed. For brevity, repeated steps are not repeated, but are incorporated herein by reference.

[0157] Computer System Overview

[0158] Now refer to Fig.14 A basic conceptual diagram of a flow cytometer system 1400 is shown. Various embodiments of the flow cytometer 1400 may be commercially available. The five main subsystems of the flow cytometer system 1400 include: an excitation optical system 1402, a fluidics system 1404, an emission optical system 1406, an acquisition system 1408, and an analysis system 1410. In general, a "system" includes hardware devices, software, and / or a combination thereof.

[0159] The excitation optical system 1402 includes, for example, a laser device 1412, an optical component 1414, an optical component 1416, and an optical component 1418. Example optical components include optical prisms and optical lenses. The excitation optical system 1402 illuminates the optical interrogation area 1420. The jet system 1404 carries the fluid sample 1422 through the optical interrogation area 1420. The emission optical system 1406 includes, for example, optical components 1430 and optical detectors SSC, FL1, FL2, FL3, FL4, and FL5. The emission optical system 1406 collects photons emitted or scattered from the passing particles. The emission optical system 1406 focuses these photons onto the optical detectors SSC, FL1, FL2, FL3, FL4, and FL5. The optical detector SSC is a side scatter channel. The optical detectors FL1, FL2, FL3, FL4, and FL5 are fluorescence detectors, which may include bandpass or longpass filters to detect specific fluorescence wavelengths. Each optical detector converts photons into electrical pulses and sends the electrical pulses to acquisition system 1408. Acquisition system 1408 processes and prepares these signals for analysis in analysis system 1410.

[0160] Described in this article Figures 1 to 13 Exemplary components of the excitation optical system 1402, emission optical system 1406, and acquisition system 1408 of the flow cytometer 1400 are included. U.S. Patent Application No. 15 / 817,277, filed by David Vrane on November 19, 2017, and entitled FLOWCYTOMETERY SYSTEM WITH STEPPER FLOW CONTROL VALVE, describes exemplary components of the fluidics system 1404 and is incorporated herein by reference.

[0161] The analysis system 1410 can store digital representations of the signals for analysis after acquisition of the signals by the acquisition system 1408. The analysis system 1410 is a computer that includes a display device, at least one processor, scratch pad memory, and one or more storage devices that can semi-permanently store more data (including signal data) and software.

[0162] One or more storage devices of the analysis system may store analysis software having instructions that can be executed by the processor to obtain laboratory results for the analyzed biological sample (or other type of sample, such as a chemical). The analysis system 1410 and analysis software may also be used to calibrate the flow cytometer using compensation controls when initializing with control particles passing through the flow cytometer before passing the biological sample through the flow cytometer. The analysis software has instructions executed by the processor that analyze stored signal data acquired by the acquisition system 1408 during calibration with controls passing through the flow cytometer. The analysis software also has instructions executed by the processor that analyze stored signal data acquired by the acquisition system 1408 during analysis of the biological sample passing through the flow cytometer after the flow cytometer has been calibrated.

[0163] Method Overview

[0164] Now refer to Fig.15 as well as FIG. 16A to FIG. 16C , a flow chart depicting a method 1500 for performing flow cytometry using fast compensation. For example, Fig.14 The flow cytometer system 1400 may perform the method 1500. Flow cytometry allows for rapid multi-parameter collection and analysis of data on single cells or particles in a sample fluid.

[0165] exist Fig.15In the illustrated step 1501, the flow cytometer system starts the flow cytometer 1604, 1614 and checks the performance of the flow cytometer. The flow cytometer can be a conventional flow cytometer 1604 or a spectral flow cytometer 1614. Conventional flow cytometers 1604 traditionally use mirrors, filters, and photomultiplier tubes (PMTs) to capture or sense signal data. Spectral flow cytometers 1614 are based on similar principles, but instead, a multi-channel detector (e.g., a charge coupled device (CCD)) is typically used, which has a device that disperses the spectrum across multiple channels into the detector.

[0166] In step 1505, if Fig.16A As shown, the system begins to calibrate the flow cytometer 1604, 1614. The system uses one or more single-stained reference or compensation controls 1602A-1602E (collectively referred to as single-stained reference or compensation controls 1602), which are passed one at a time through the flow cytometer 1604, 1614. The one or more single-stained reference controls 1602A-1602E generate spectral profiles or features 1606A-1606E, respectively, based on the corresponding various autoluminescence or luminescence of the fluorescein labeled to the control microparticles.

[0167] In step 1510, during calibration, the system uses the single-stained compensation controls 1602 to generate an initial spillover matrix for a conventional flow cytometer 1604 or an initial reference matrix for a spectral flow cytometer 1614. When performing multicolor flow cytometry, the system uses the single-stained samples 1602 that pass through the flow cytometers 1604, 1614 to determine compensation levels, such as Fig.16A As shown. The single staining of microparticles 1602 can reveal the spectral curve or characteristics 1606A-1606E of fluorescein to the fluorescence light detector of the flow cytometer. This allows the system and / or user to compensate for the overlap of the spectral curves of fluorescence from the sample 1610 due to multi-stained cells or microparticles 1612A-1612E contained in the sample (e.g., see Fig. 16B The overlap 1616 in Fig. 16B shown.

[0168] The staining of the compensation control 1602 must be as bright or brighter than the microparticles 1612A-1612E in the sample 1610. Antibody capture beads can replace cells, and if the measured fluorescence is brighter for the control, one fluorophore-bound antibody can replace another. The exception is a combination dye, which cannot be replaced. Combination dyes from different suppliers or different batches must be treated like individual dyes, and each dye should use a separate single staining control because the overflow of each of these dyes may be different. Similarly, the compensation algorithm should be performed with positive and negative populations. Regardless of whether each individual compensation control contains beads, cells used in the experiment, or even different cells, the control itself must contain microparticles with the same level of autofluorescence. The entire set of compensation controls can include a single sample of beads or cells, but the individual samples must have the same carrier microparticles for the fluorophore. Moreover, the compensation control uses the same fluorophore as the sample. For example, green fluorescent protein (GFP) and fluorescein isothiocyanate (FITC) both mainly emit green photons, but the emission spectra are very different. Therefore, the system cannot use one for the sample and the other for a compensating control. Furthermore, the system must collect enough events to make a statistically significant determination of spillover (e.g., approximately 5000 events for both the positive and negative populations).

[0169] During calibration of a conventional flow cytometer 1604, the system obtains an initial spillover matrix from single-stained reference controls 1602 that are passed through the conventional flow cytometer one at a time. In the conventional flow cytometer 1604, a series of edge filters and beam splitters are used to separate the fluorescent signals (e.g., colors) into discrete fluorescent bands 1606A-1606E. The system utilizes photomultiplier tubes (PMTs) to detect (e.g., measure) each individual channel. Ideally, the fluorescent bands would be completely discrete, but this is unrealistic. During detection of fluorescent signals in the control 1602 (as well as the sample 1610), "spillover" can occur between fluorescent bands, such as shown in the combined curve 1616. The system utilizes a spillover matrix [S] to define spillover between fluorescent bands (e.g., Fig. 16B Overflow 1618 between discrete bands of combined curve 1616 is shown).

[0170] Alternatively, during calibration of the spectral flow cytometer 1614, the system obtains an initial reference matrix from a single-stained reference control 1602 that passes through the spectral flow cytometer. Spectral flow cytometry is a technique based on conventional flow cytometry, but replaces the traditional mirrors, filters, and photomultiplier tubes (PMTs) of a conventional flow cytometer system with a spectrometer and a multi-channel detector (e.g., a charge-coupled device (CCD)). In a spectral flow cytometer, side scattered light and fluorescence are collected and coupled to a spectrometer directly or through optical fibers, where the entire light signal is dispersed and displayed on the CCD or multi-channel detector as a high-resolution spectrum.

[0171] In step 1515, refer to Fig.15 and Fig. 16B , the system passes a sample 1610 through a flow cytometer 1604, 1614. The system generates, obtains and / or records data (e.g., event data) representing a spectral curve 1616 based on the sample fluid 1610. For example, fluorescent cells are detected in the sample fluid flowing through the flow cytometer. Each detection of a fluorescent cell is an event. Event data can be defined based on the measured sample event vector.

[0172] In step 1520, the system generates a compensated sample event vector (for conventional flow cytometer 1604) or a demixed sample event vector (for spectral flow cytometer 1614) to count the number of various types of cells or particles in sample 1610 to obtain a measure of concentration. Fig. 16C As shown, an inverse matrix 1624 (determined based on an initial spillover matrix and / or an initial reference matrix with fine-tuning) is used on the event data representing the spectral curve 1616 to generate compensated sample event vectors or unmixed sample event vectors representing separate spectral curves or features 1606A-1606E of various autoluminescence or luminescence of fluorescent substances labeled into various cells 1612A-1612E in the sample 1610.

[0173] For conventional flow cytometer 1604, the system calculates compensated event vectors based on the initial spillover matrix and the measured sample event vectors. For spectral flow cytometer 1614, the system calculates unmixed sample event vectors based on the initial reference matrix and the measured sample event vectors.

[0174] Unfortunately, the initial spillover matrix and the reference matrix are often not accurate enough to produce reliable results. Therefore, in step 1525, the system performs a fast compensation. For a conventional flow cytometer 1604, this step includes compensating for inaccuracies in the initial spillover matrix. For a spectral flow cytometer 1614, this step includes compensating for inaccuracies in the initial reference matrix.

[0175] In step 1527, based on the fast compensation, the system generates a re-compensated sample event vector in a faster manner than previously performed. Method 1500 then ends at step 1530, but one or more steps may be repeated for other samples using the flow cytometer for similar date / time calibration / sampling.

[0176] Obtaining spillover matrices from single-stained controls for conventional flow cytometry

[0177] Thus, the system may include a conventional flow cytometer to generate or obtain the spillover matrix from the single staining control. The steps of generating or obtaining the spillover matrix by using a conventional flow cytometer are further discussed.

[0178] Assume that the matrix [S] is an N×N dimensional spillover matrix obtained from single-stained compensation controls, where N is the number of fluorescence detectors. Example compensation controls include beads 1602 stained or colored with fluorescein, such as fluorescein isothiocyanate (FITC), R-phycoerythrin (PE), Peridinin Chlorophyll Protein Complex (PerCP), phycoerythrin and cyanine dye (PE-Cy7), Allophycocyanin (APC), and a combination of fluorescein and APC (APC-Cy7).

[0179] Assume that vector {U} is a measured sample event vector having N values, each value coming from one of the N detectors detecting a compensation control (e.g., FITC, PE, PerCP, PE-Cy7, APC, APC-Cy7).

[0180] Assume that vector {V} is a compensated sample event vector with N values. The measured sample event vector {U} is equal to the overflow matrix [S] multiplied by the compensated sample event vector {V}. This can be represented by the following matrix relationship with the measured sample event vector {U}:

[0181] [S]{V}={U} Equation 1

[0182] Therefore, using the inverse overflow matrix [S] -1 , the compensated sample event vector {V} can be obtained from the matrix equation:

[0183] {V}=[S] -1 {U} Equation 2

[0184] The initial spillover matrix [S] can be obtained by measuring each single-stained control (e.g., FITC, PE, PerCP, PE-Cy7, APC, APC-Cy7) at each detector to obtain the following matrix:

[0185]

[0186] In the subscripts x, y in Equation 3, the x value indicates the detector number. In the subscripts x, y in Equation 3, the y value indicates the column associated with the single-stained control.

[0187] Each column in the initial spillover matrix [S] (separation of variables (SOV) matrix) corresponds to a single stain control (e.g., FITC, PE, PerCP, PE-Cy7, APC, APC-Cy7). For example, the first column corresponds to the FITC single stain control. As another example, the second column corresponds to the PE single stain control; and so on for each stain control that is run to calibrate the flow cytometer. Each row in the initial spillover matrix [S] corresponds to a given detector number. For example, the first row corresponds to detector 1. The second row corresponds to detector 2, and so on.

[0188] Typically, the initial spillover matrix generated is not accurate enough to accurately separate the spectra and identify cells or particles. Therefore, it is necessary to fine-tune the values ​​of the off-diagonal elements of the initial spillover matrix [S] (e.g., fine-tune the initial spillover matrix [S] to generate an adjusted spillover matrix [S]' and its associated inverse matrix, the adjusted compensation matrix [C]'). Fine-tuning can be performed based on the experience and judgment of the laboratory technician / operator. Fine-tuning is typically performed to correct for distortion caused by interactions of fluorescein dyed on the same cells or particles or by the system used to measure single-stained and unstained controls, or both of these interactions and the system.

[0189] Assume that the adjustment matrix [D] is to fine-tune (e.g., add) the non-diagonal element values ​​of the initial spill matrix [S]. It can be calculated according to the matrix equation {V R}=[[S]+[D]] -1 {U} Determine the recompensated event vector for conventional flow cytometers {V R}.

[0190] Obtaining unmixed event list data for spectral flow cytometry

[0191] Alternatively, the system may include a spectral flow cytometer 1614 to generate or obtain the unmixed event list data. The steps of generating or obtaining the unmixed event list data by using a spectral flow cytometer are further discussed.

[0192] Assume that [R] is the N×M initial reference matrix for spectral flow cytometry obtained from single-stained reference controls, where N is the number of detectors and M is the number of fluorophores to be measured (e.g., FITC, PE, PerCP, PE-Cy7, APC, APC-Cy7), where M is always less than N.

[0193] Assume {U} is a measured sample event vector with N values, each value coming from one of the N detectors.

[0194] Assume {V} is an unmixed sample event vector having M values ​​(e.g., fluorescence intensities), where each value is an unmixed value of a fluorescein (e.g., FITC, PE, PerCP, PE-Cy7, APC, APC-Cy7).

[0195] The unmixed sample event vector {V} and the measured sample event vector {U} have the following matrix relationship:

[0196] [R]{V}={U}

[0197] Since the number of variables M in the unmixed sample event vector {V} is smaller than the number of variables N in the measured sample event vector {U} (e.g., the dimension of the unmixed sample event vector is smaller than the dimension of the measured sample uniform vector), the system then uses a least squares algorithm to obtain a solution to the above equation.

[0198] Compared to conventional flow cytometers, the unmixed event vector is equivalent to the compensated event vector. Therefore, the spectral spillover matrix of the unmixed event list data (e.g., the unmixed sample event vector) is the following identity matrix [I]:

[0199]

[0200] Typically, the unmixed event list data is not accurate enough, so that fine adjustment of the unit spectral spillover is required (e.g., fine adjustment is performed to generate an adjusted spectral spillover matrix). Therefore, the equation for the recompensated event vector of the spectral flow cytometer becomes: {V R}=[[I]+[D]] -1 {V}, where [D] is the fine-tuned δ in the i-th row and j-th column respectively. i,j The increment matrix is ​​an n×n increment matrix of , and is zero if fine adjustment is not required. For example, the increment matrix can be

[0201] Fast compensation of flow cytometry data

[0202] Therefore, in flow cytometry (e.g., conventional and spectral), the Flow Cytometry Standard (FCS) data collected from the cytometer is linear raw list data. The list data needs to be compensated before it is used for graphing and for statistical analysis of cells in a sample. The system performs fast compensation to address the lack of accuracy in the overflow matrix of conventional flow cytometers and / or the unmixed event list data of spectral flow cytometers.

[0203] Compensation of the tabulated data is based on an initial spillover matrix obtained by the system from measured single stain compensation controls and / or fine tuning inputs. The initial spillover matrix obtained is usually not accurate enough. Fine tuning can be performed by fine tuning the values ​​in the initial spillover matrix, thereby generating an adjusted spillover matrix.

[0204] Whenever the overflow value is fine-tuned, the overflow matrix needs to be inverted to obtain the compensation matrix. The compensation matrix is ​​then multiplied with each list data event vector to generate a compensated list data (eg, a re-compensated event vector).

[0205] For example, take an experiment with N fluorescence parameters. To compensate each event vector, it requires N 2 The computational complexity is about N times multiplication plus N×(N-1) times addition to generate the compensated event vector. 2 (For example, O(N 2 )).

[0206] For experiments with a limited number of fluorophore parameters and a limited number of events, compensation calculations may not be a bottleneck for flow cytometric data analysis. However, if the experiment contains a large number of fluorescence parameters (e.g., more than 20 fluorescence parameters) and has a large number of events (e.g., 2 million events), then compensation calculations can be extremely time consuming. As a result, the displayed graphs and statistics on the computer interface can be extremely slow to respond due to the large number of calculations that must be processed each time the system changes the overflow value.

[0207] Advantageously, the system implements a fast compensation algorithm that significantly reduces the amount of computation when the system receives or performs fine-tuning of the spillover matrix for flow cytometry data analysis without sacrificing any accuracy of the compensated tabulated data. For example, such a fast compensation algorithm requires only (3N+1) multiplications / divisions plus (N+1) additions. The complexity of such a fast compensation algorithm is approximately N (e.g., O(n)). Thus, the system can significantly improve the responsiveness of displayed charts and statistics.

[0208] For example, consider a 20-color experiment with one million events. The typical compensation algorithm requires a total of 400 million multiplications plus 399 million additions each time the system receives or performs a fine adjustment of the overflow value. In contrast, the fast compensation algorithm of the present system requires only a total of 60 million multiplications plus 20 million additions. The total savings in multiplications and additions are 566% and 1895%, respectively, compared to the typical compensation algorithm.

[0209] The derivation of the current fast compensation algorithm is now discussed.

[0210] Assume that the matrix [C] is the compensation matrix. According to the matrix equation [C] = [S] -1 , the compensation matrix [C] is the inverse of the spillover matrix [S]. If the compensation matrix [C] is multiplied by the inverse of the spillover matrix [S], an identity matrix such as in the matrix equation [C][S] = [I] is obtained. The compensated event vector {V} can be calculated by multiplying the compensation matrix [C] and the uncompensated measured event vector {U} (expressed by the matrix equation {V} = [C] {U}).

[0211] Due to the fine-tuning, the system generates or calculates an adjusted spillover matrix [S]'. Assume that the value of one element in the initial spillover matrix [S] has changed, for example, S i,j '=>S i,j +δ i,j The fine-tuned overflow matrix [S]' can be represented by the sum of the initial overflow matrix [S] and the fine-tuning in the incremental matrix [D] according to the matrix equation [S]' = [S] + [D], where is an increment matrix, where the subscripts i and j denote the i-th row and j-th column, respectively. The recompensated event vector {V R} can be calculated by multiplying the inverse of the fine-tuned spillover matrix [S]' (the fine-tuned compensation matrix [C]') with the uncompensated measured event vector [U], such as the matrix equation {V R}=[[S]+[D]] -1 {U}. The incremental matrix [D] has the same dimensions as the initial spill matrix [S]. The incremental matrix [D] includes the incremental values ​​δ used to fine-tune the initial spill matrix [S]. i,j .

[0212] Since [S] + [D] = [S] ( [I] + [C] [D]), [ [S] + [D]] -1 =([I]+[C][D]) -1 [C], the recompensated event vector {V R The equation can be rewritten as

[0213] {VR}=([I]+[C][D]) -1 [C]{U}=([I]+[C][D]) -1 {V}

[0214] Among them, ([I]+[C][D]) -1 is the recompensation matrix.

[0215] Since the recompensation matrix can be simplified to

[0216]

[0217] Then, the matrix equation for the recompensated event vector can be written as

[0218]

[0219] Each component of the recompensated vector {V} is determined by adding / subtracting and multiplying / dividing the components of the uncompensated measured event vector {U}, thereby significantly reducing the number of calculations. Thus, the recompensated event vector {V} can be calculated more quickly by the processor of the computer using the fast compensation algorithm. R}.

[0220] Thus, using the fast compensation algorithm, cell samples can be analyzed faster using a flow cytometer and results (e.g., cell counts, concentrations, titrations) can be obtained more efficiently. Instead of a researcher or laboratory technician spending a day or more waiting for calculations to be completed to obtain laboratory results for a biological sample (e.g., blood or urine), laboratory results for a biological sample can be obtained within hours by using the fast compensation algorithm.

[0221] in conclusion

[0222] Thus the embodiments have been described.While the embodiments have been described in detail, they should not be construed as being limited to such embodiments but rather construed in accordance with the claims that follow.

[0223] Although specific exemplary embodiments have been described and shown in the drawings, it is to be understood that such embodiments are merely illustrative and not restrictive of the invention, since various other modifications may occur to those skilled in the art, and that these embodiments are not limited to the specific construction and arrangements shown and described.

[0224] Certain functions of the flow cytometer can be implemented in software and performed by a computer or processor, such as analyzing the electrical signals detected by the detector to count different particles in the sample fluid. The program or code segment of the software is used to perform the tasks necessary to perform these functions. The program or code segment can be stored in a processor-readable medium, or sent by a computer data signal specifically implemented in a carrier wave on a transmission medium or a communication link. The processor-readable medium may include any storage medium that can store information. Examples of processor-readable media include: electronic circuits, semiconductor memory devices, read-only memories (ROMs), flash memories, erasable programmable read-only memories (EPROMs), floppy disks, CD-ROMs, optical disks, and hard disks. The code segment can be downloaded to the storage medium via a computer network such as the Internet, an intranet, etc.

[0225] Although this specification includes many specific details, these should not be interpreted as limitations on the scope of the present disclosure or the scope that can be claimed, but should be interpreted as descriptions of features peculiar to the specific implementation of the present disclosure. Some features described in the context of separation implementation in this specification can also be implemented in combination in a single implementation. On the contrary, the various features described in the context of a single implementation can also be implemented in multiple implementations, individually or in a sub-combination. In addition, although features are described as working in some combinations above, even initially equally claimed, one or more features from the claimed combination can be removed from the combination in some cases, and the claimed combination can be directed to a sub-combination or a variation of the sub-combination. Therefore, the claimed invention is limited only by the appended claims.

Claims

1. A method for performing flow cytometry using a conventional flow cytometer, wherein the conventional flow cytometer uses a mirror, a filter, and a photomultiplier tube to capture or sense signal data, the method comprising the following steps: After calibrating (1505) the conventional flow cytometer, Generating (1510) an initial spillover matrix using a conventional flow cytometer using a plurality of single-stained compensation controls; passing (1515) the biological sample through the conventional flow cytometer; generating (1515) a measured sample event vector by measuring fluorescence of a plurality of cells in the biological sample passing through the conventional flow cytometer; as well as After passing the biological sample through and generating the measured sample event vectors using the conventional flow cytometer, the processor of the computer executes instructions to perform the following functions: generating (1520) a compensated sample event vector by using the initial spillover matrix and the measured sample event vector, wherein the compensated sample event vector is equal to the measured sample event vector linearly multiplied by the inverse of the initial spillover matrix; generating (1525) an adjusted spillover matrix by fine-tuning the initial spillover matrix, wherein the fine-tuning includes adjustments required to correct for distortion caused by interactions of different fluorophores stained on the same cell or the same particle, distortion caused by the conventional flow cytometer measuring single-stained controls and unstained controls, or both; and A re-compensated event vector is calculated (1527) by accurately separating the spectra using the adjusted spillover matrix and the measured sample event vector.

2. The method according to claim 1, wherein: forming a recompensation matrix using the initial spillover matrix and the fine adjustment of the initial spillover matrix in the incremental matrix; as well as The step of calculating the re-compensated event vector is accomplished by linearly multiplying the re-compensation matrix and the measured sample event vector, thereby reducing the amount of calculation to quickly obtain the laboratory result of the biological sample.

3. The method according to claim 1, wherein: Generating the adjusted spillover matrix includes: Add the incremental matrix to the initial overflow matrix, wherein the incremental matrix has the same dimensions as the initial spillover matrix, and The incremental matrix includes one or more incremental values ​​for fine-tuning the initial overflow matrix.

4. The method according to claim 1, wherein: The initial spillover matrix comprises N×N dimensions, and wherein N is the number of the single-stained compensation controls.

5. The method according to claim 1, wherein: Each of the single staining compensation controls includes any one of the following: fluorescein isothiocyanate FITC; R-phycoerythrin PE; peridinin chlorophyll protein complex PerCP; PE-Cy7; allophycocyanin APC; and APC-Cy7.

6. The method according to claim 4, wherein: The measured sample event vector includes N values, and wherein the compensated sample event vector has N values.

7. The method according to claim 2, wherein: The re-compensated event vector is equal to the inverse of the sum of the initial spillover matrix and the delta matrix multiplied by the measured sample event vector.

8. A method of performing flow cytometry using a spectral flow cytometer using a multi-channel detector having a means for dispersing a spectrum across multiple channels into a semiconductor photodetector, the method comprising the steps of: After calibrating (1505) the spectral flow cytometer, generating (1510) a reference matrix by using a plurality of single-stained compensation controls; passing (1515) the biological sample through the flow cytometer; generating (1515) a measured sample event vector by measuring fluorescence of a plurality of cells passing through the flow cytometer; After passing the biological sample through and generating the measured sample event vectors using the semiconductor photodetector of the spectral flow cytometer, the processor of the computer executes instructions to perform the following functions: generating (1520) an unmixed sample event vector by using the reference matrix and the measured sample event vector, wherein a number of variables in the unmixed sample event vector is less than a number of variables in the measured sample event vector, and wherein the unmixed sample event vector is equal to the measured sample event vector linearly multiplied by an inverse of the reference matrix; generating (1525) an adjusted spectral spillover matrix by fine-tuning the spectral spillover matrix, wherein the fine-tuning includes adjustments required to correct for distortion caused by interactions of different fluorophores stained on the same cell or the same particle, distortion caused by the spectral flow cytometer measuring single-stained controls and unstained controls, or both; and A re-compensated event vector is calculated (1527) by using the adjusted spectral spillover matrix and the measured sample event vector.

9. The method according to claim 8, wherein forming a recompensation matrix using the spectral spillover matrix and the fine adjustments to the spectral spillover matrix in the delta matrix; and The step of calculating the re-compensated event vector is accomplished by linearly multiplying the re-compensation matrix and the measured sample event vector, thereby reducing the amount of calculation to quickly obtain the laboratory result of the biological sample.

10. The method according to claim 8, further comprising the steps of: The step of solving the unmixed sample event vector includes using a least squares algorithm on the measured sample event vector and the inverse of the reference matrix.

11. The method according to claim 8, wherein: The initial spectral spillover matrix for the unmixed sample event vector is the identity matrix.

12. The method according to claim 11, wherein: Generating an adjusted spectral spillover matrix includes: Add the incremental matrix to the initial spectral spillover matrix, wherein the incremental matrix has the same dimensions as the initial spillover matrix, and The incremental matrix includes one or more incremental values ​​for fine-tuning the initial overflow matrix.

13. The method according to claim 11, wherein: The initial spectral spillover matrix comprises N×N dimensions, and wherein N is the number of the single-stained compensation controls.

14. The method according to claim 8, wherein: The measured sample event vector comprises N values, and wherein the unmixed sample event vector has N values.

15. The method according to claim 11, wherein: The re-compensated sample event vector is equal to the inverse of the sum of the initial spectral spillover matrix and the increment matrix multiplied by the measured sample event vector.

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