Fast re-compensation of flow cytometer data for spillover readjustment

Abstract

To provide a method of performing fast compensation in a flow cytometry experiment.SOLUTION: A method includes: performing calibration of a flow cytometer by using a plurality of single stained compensation controls and generating an initial spillover matrix; running a sample through the flow cytometer; generating a measured sample event vector by measuring fluorescence of a plurality of cells passing through the flow cytometer; generating a compensated sample event vector by using the initial spillover matrix and the measured sample event vector; generating an adjusted spillover matrix by finely adjusting the initial spillover matrix; and calculating a re-compensated event vector by using the adjusted spillover matrix and the measured sample event vector.SELECTED DRAWING: Figure 15

Description

[Technical field]

[0001] [CROSS REFERENCE TO RELATED APPLICATIONS] This non-provisional patent application claims the benefit of U.S. patent application Ser. No. 62 / 674,273, entitled “FAST RECOMPENSATION OF FLOW CYTOMETERY DATA FOR SPILLOVER READJUSTMENTS,” filed May 21, 2018 by inventor Zhenyu Zhang, and is incorporated herein by reference for all intents and purposes.

[0002] This patent application is related to U.S. Patent Application No. 15 / 659,610, entitled "COMPACT DETECTION MODULE FOR FLOW CYTOMETERS," filed July 25, 2017 by inventors Ming Yan et al., which is hereby incorporated by reference for all intents and purposes. This patent application is further related to U.S. Patent Application No. 15 / 498,397, entitled "COMPACT MULTI-COLOR FLOW CYTOMETER," filed April 26, 2017 by David Vrane et al., which describes a flow cytometer with which the present embodiments may be used, and which is hereby incorporated by reference for all intents and purposes. This patent application is further related to U.S. Patent Application No. 15 / 817,277, entitled "FLOW CYTOMETERY SYSTEM WITH STEPPER FLOW CONTROL VALVE," filed Nov. 19, 2017 by David Vrane, which describes a flow cytometer fluidics system and is hereby incorporated by reference for all intents and purposes.

[0003] Embodiments of the present invention generally relate to analyzing flow cytometer data to determine particle counts within a sample. [Background technology]

[0004] A flow cytometer typically has a viewing orifice that is illuminated by one or more lasers. Laser light from the one or more lasers impinges on various fluorochrome-labeled particles that pass through the orifice. The fluorochrome-labeled particles are typically various biological cells within a sample that have been labeled with different fluorochromes (fluorescent dyes) and these biological cells can be analyzed to obtain information about the sample that can be generalized into global information. One or more optical detectors in the flow cytometer are used to detect the fluorescence (fluorescent light) emitted from the fluorochrome-labeled particles that pass through an orifice impinged by the laser light from the one or more lasers.

[0005] One or more different optical filters can be placed before the emitted fluorescence from the fluorochrome-labeled particles reaches each detector. These optical filters are placed in the emitted fluorescence path such that only a particular bandwidth of light associated with the expected fluorescence of the fluorochrome is visible to each detector. That is, the bandwidth of any given filter utilizes a peak in the emission spectrum of a particular fluorescent dye. In this way, for any given particle, the collective signal from the detectors indicates the type or types of fluorochromes attached to the particle. The signals detected by the detectors from the emitted fluorescence allow for rapid and extensive cellular classification of the various particles in the sample.

[0006] However, emission spectra can overlap between dyes. This limits the number of different fluorochromes that can be detected simultaneously on a given particle by a single laser and detector. Because emission bandwidths are typically in the 30 nanometer (nm) to 60 nm wavelength range, conventional flow cytometers can typically detect only four or five fluorochromes per laser line. Increasing the number of lasers provides a convenient, but expensive, way of increasing the number of fluorochromes that can be detected simultaneously.

[0007] Further complicating the detection of emitted fluorescent light is that many dyes used to stain particles are excited over a range of laser wavelengths larger than the usual 30-60 nm bandwidth range, which can result in signal crosstalk between detectors of different lasers.

[0008] Conventional flow cytometry fluorescence detection systems limited divergence by increasing the collimating lens focal length. However, this resulted in a larger diameter of the light beam and limited the number of detectors to six detectors or so depending on the size of the final image required. In these conventional flow cytometry systems, the final image size was constrained by optical imaging of large sizes such as 800 micrometers (μm), i.e. 800 microns, and optical aberrations in collimating broadband light (e.g., wavelengths from 400 nm to 800 nm) into a set of detectors with a diameter of less than 3 millimeters (mm).

[0009] In another flow cytometry system, the incident light is re-imaged using spherical micromirrors for each detector in the detector chain of a set of detectors in the 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 caused by the reflection from the spherical micromirrors. As the image size increases along the detector chain, increasing the number of detector channels along the row in the detector chain requires detectors with large areas, which makes the flow cytometer large, bulky and expensive.

[0010] The resolution of spectrally overlapping emission spectra is also central to increasing the number of detectable fluorochromes. Using a detector array, fluorochromes can be identified based on their collective emission signature across multiple wavelengths to increase the number of detectable fluorochromes. Essentially, the entire fluorescent signal is dispersed in color by either a diffraction grating or a prism into the array of detectors. In this way, the entire emission spectrum is discretized across the detectors. Spectral deconvolution (unmixing) can be used to calculate the contribution of known individual fluorochrome spectra to the total collected signal. However, this approach to increasing the number of detectable fluorochromes has two main limitations.

[0011] The continuous linear nature of the dispersive element / detector array precludes tuning the bandwidth to exploit the true nature of the fluorochrome spectrum. Thus, broader bandwidth discrimination favors longer wavelength fluorochromes, but misses the details of shorter wavelength fluorochromes with compressed spectra. In addition, if there is scattered light from other lasers, this scattered light is inevitably collected by the detector array. This scattered light corrupts the fluorescent signal detected by the detector. Summary of the Invention

[0012] Therefore, there is a need for further improvements in flow cytometers to increase the number of detectable fluorochromes and thus better analyze fluorochrome-labeled particles within a sample.

[0013] Embodiments of the present invention are summarized by the following claims. [Brief description of the drawings]

[0014] [Figure 1] FIG. 2 is a block diagram of a detection module of a flow cytometer system. [Figure 2A] FIG. 1 is a schematic diagram of an image array with successive channel reimaging. [Figure 2B] FIG. 1 is a schematic diagram of a 1f image array with alternating channel reimaging. [Diagram 3] FIG. 1 is a schematic diagram of a compact detection module containing a 16-channel 1f imaging array of a modular flow cytometer system. [Figure 4A] FIG. 2 is a close-up view of fluorescent light exiting from different points across the diameter of an optical fiber. [Figure 4B] FIG. 2 is a close-up view of a detector that converts an optical signal into an electrical signal. [Diagram 5] FIG. 1f is a close-up view of a portion of the image array. [Figure 6] FIG. 1 is a diagram of a detection module in a modular flow cytometer system having a 16-channel 1f imaging array and 16 detector channels. [Figure 7A] 1A-1D are different schematic diagrams of a detection module in a modular flow cytometer system having a pair of 8-channel 1f imaging arrays and a pair of 8 detector channels. [Figure 7B] 1A-1D are different schematic diagrams of a detection module in a modular flow cytometer system having a pair of 8-channel 1f imaging arrays and a pair of 8 detector channels. [Figure 7C]1A-1D are different schematic diagrams of a detection module in a modular flow cytometer system having a pair of 8-channel 1f imaging arrays and a pair of 8 detector channels. [Figure 8] FIG. 1 is a perspective view of a 1f imaging array with spherical micromirrors. [Figure 9A] FIG. 9 is a cross-sectional view of the 1f image array of FIG. [Figure 9B] FIG. 9 is a cross-sectional view of the 1f image array of FIG. [Figure 10] FIG. 1 is a perspective view of a 1f imaging array with concave square micromirrors. [Figure 11] FIG. 11 is a top side view of a portion of the 1f imaging array of FIG. [Figure 12] FIG. 11 is a perspective view of a mounting block adjacent to the imaging block and a low-cost, thin-shell detector used in the detector modules of FIGS. 6 and 7A-7C. [Figure 13] FIG. 2 is a top view of an optical plate assembly in a modular flow cytometry system. [Figure 14] FIG. 1 is a basic conceptual diagram of a flow cytometer system. [Figure 15] FIG. 1 shows the overall method for performing an experiment on a flow cytometer. [Figure 16A] FIG. 1 is a diagram of the calibration process of a flow cytometer using a single stained compensation control to generate an initial spillover or reference matrix with a level of compensation. [Figure 16B] FIG. 1 illustrates a sample being passed through a flow cytometer, resulting in mixed sample event vectors with overlapping spectral profiles due to multi-stained cells or particles. [Figure 16C] FIG. 13 illustrates a process for generating a corrected sample event vector or an unmixed sample event vector using an inverse matrix (derived from an initial spillover matrix and / or an initial reference matrix with refinements) on the event data. DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

[0015] In the following detailed description of the embodiments, numerous specific details are set forth in order to provide a thorough understanding. However, it will be apparent to one of ordinary skill in the art that the embodiments may be practiced without these specific details. In other instances, well-known methods, procedures, components, and circuits have not been described in detail so as not to unnecessarily obscure aspects of the embodiments of the present invention.

[0016] Embodiments of the present invention include methods, devices and systems for a flow cytometer with a compact highly multiplexed detection module.

[0017] Overall Overview A flow cytometer is disclosed having a compact detection module for fluorescence with an increased number of detectors and a minimum image size at the detector compared to conventional flow cytometers. Each detector module is provided with at least one laser. Multiple lasers can be supported in a compact format by multiple detector arrays. The increased number of detectors is made possible by careful control of incremental aberrations in the detector array as light is transmitted through the detection chain. The compact size of the compact detection module is achieved by reducing the distance between the micromirrors and the filters and careful downward scaling, thereby minimizing image degradation along the line or chain of micromirrors and filters in the imaging array.

[0018] Conventional limitations can be overcome by the use of an optical system with multiple individual detectors and adjustable shifts in filter bandwidth. The optical system allows the concentration of spectral collocation points to best resolve the spectra of both the long and short wavelength dyes used to mark the particles analyzed by the flow cytometer. Fluorescence from the excited particles is imaged into a multimode fiber by an objective lens with a high numerical aperture (NA). The broadband fluorescence emerging from the multimode fiber is collimated and then combined into (imaged onto) multiple detectors. Collimation of the broadband fluorescent light emerging from the multimode fiber is a challenge.

[0019] Flow cytometer detection module FIG. 1 shows a functional block diagram of a portion of a detection module 100 of a flow cytometer. An improved flow cytometer can include multiple detection modules. The detection module 100 is a wavelength-separating system. The detection module 100 continuously reflects and re-images the fluorescent light output 101A from the optical fiber 102 at the imaging array 106. The imaging array 106 is a mechanical imaging array that includes multiple reflecting mirrors 112A-112E suspended in air by mechanical mounts and corresponding multiple long-pass dichroic filters 110A-110E. Generally, dichroic filters are precision color filters used to selectively pass light of a range of wavelengths of color light and reflect color light of other wavelengths. Alternatively, the multiple long-pass dichroic filters 110A-110E can be band-pass filters.

[0020] The imaging array 106 is capable of reflecting and re-imaging the fiber spot of light N times (where N is greater than 2) while maintaining the optical quality of the fiber spot at the end of the imaging array. Re-imaging is the function of recreating the original image with certain aberrations, i.e., reimage, at a surface such as the surface of each of the dichroic filters 110A-110E. The detection module 100 further includes a number of detector channels 113A-113E each comprising a number of objective lenses 116A-116E and a number of detectors 118A-118E with which the light is in communication respectively. Optionally, the multiple detector channels 113A-113E can further include a number of bandpass filters 114A-114E that respectively ensure that different desired wavelength ranges are detected by the multiple detectors.

[0021] In the image array 106, light 101A incident on the first mirror 110A is reflected by the mirror into reflected light 103A. The reflected light 103A from the first mirror 110A passes through air and passes onto a first long-pass dichroic filter 112A. The light 103A from the first mirror 110A is split by the long-pass dichroic filter 112A into a continuous light portion 101B and a passed or transmitted light portion 105A. The continuous light portion 101B passes through air and passes onto the next mirror in the series, mirror 110B. The transmitted light portion 105A is coupled into a first band-pass filter 114A of a first detector channel 113A. The transmitted light portion 105A is cleaned up by the band-pass filter 114A and then coupled into an optical detector 118A by an objective lens 116A. This process is repeated for each stage (each detector channel) in image array 106 .

[0022] The imaging array 106 includes five stages with five re-imagings on each longpass dichroic filter. It would still be desirable to include a larger number of detectors. However, after the sixth re-imaging, the beam distortion through the reflecting mirrors can accumulate to the point that it significantly degrades the image quality at the last dichroic filter. In order for the zigzag configuration between the mirrors and dichroic filters in the imaging array 106 to work properly with a larger number of detectors, it is desirable to minimize image degradation along the optical path.

[0023] Minimizing image degradation in the detection module can be achieved using two mechanisms: Image degradation can be reduced if the bending power of each reflecting mirror in the imaging array is reduced (e.g., by a factor of 2, i.e., half the bending power); Image degradation can be further reduced if the number of times the light beam is re-imaged in the imaging array is reduced (e.g., by a factor of 2, i.e., half the number of times).

[0024] By reducing the bending angle in the mirror, all types of aberrations are reduced. Because aberrations increase nonlinearly with bending angle, the improvement obtained by switching from a "fast-focus mirror" to a "slow-focus mirror" is significantly better than 2x, allowing image quality to be preserved through many 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 light beam is re-imaged in the image array. Instead of re-imaging at each dichroic filter in each detector channel, the incident light can be re-imaged on every other dichroic filter and detector channel (e.g., odd-numbered detector channels).

[0025] Now, with reference to Figures 2A and 2B, we will explain the image arrays 106A and 106B, which provide an improvement in image quality by using transparent blocks with micromirrors having different radii of curvature. In general, the image array consists of an array of micromirrors and an opposing array of bandpass and / or dichroic filters for each detection channel. In each case, the thickness L of the transparent block between the serial chain or series of micromirrors M(n), M(n)' on one side and the serial chain or series of dichroic filters D(n) on the opposite side is the same. On the other hand, the focal lengths of the micromirrors M(n) and M(n)' are different in the image array 106A of Figure 2A and the image array 106B of Figure 2B.

[0026] The focal length f of the micromirror M(n)' in FIG. 2B is L, while the focal length f of the micromirror M(n) in FIG. 2A is half of L. The larger the focal length of the micromirrors M(n)' in imaging array 106B, the more the bending angles and aberrations in the imaging along the serial chain of mirrors are reduced. Furthermore, imaging array 106B is a 1f imaging array with a transparent block thickness L and focal lengths of the micromirrors M(n)' such that reimaging is performed onto an odd number of dichroic filters, such as dichroic filters D(3), D(5), D(7) through D(n).

[0027] In Figures 2A and 2B, Spot A(1) through Spot A(n) are the spot sizes (areas) on the dichroic filters D(1) through D(N), respectively. Spot A(0) is the fiber aperture of the multimode fiber where the fluorescent light is input into each array 106A and 106B. In Figures 2A and 2B, the fiber aperture can be considered infinitesimally small to illustrate the image conjugation characteristics of the two designs.

[0028] FIG. 2A shows an image array 106A having multiple micromirrors M(1)-M(N) and multiple longpass dichroic filters D(1)-D(N), where N is an integer value greater than 1 that represents the number of detector channels. Light that is focused to spot A(1) on dichroic filter D(1) through reflection by micromirror M(1) is re-imaged through reflection of micromirror M(1) to focus the light to spot A(2) on dichroic filter D(2). This is repeated by each of the micromirrors M(2)-M(N) along the serial chain or row. Image array 106A is a 2f image array.

[0029] 2B shows a 1f design of imaging array 106B having multiple micromirrors M(1)'-M(N)' and multiple longpass dichroic filters D(1)-D(N). Micromirrors M(1)'-M(N)' have a different radius of curvature than the radius of curvature of micromirrors M(1)-M(N). 1f imaging array 106B provides image quality improvement at a detector (e.g., detectors 118A-118E in FIG. 1) over that of 2f imaging array 106A.

[0030] The chain of micromirrors M(1)' through M(N)' in 1f imaging array 106B is designed to relay images along the chain by properties of the telescope optics. For example, light focused to spot A(1) on dichroic filter D(1) is imaged to spot A(3) on dichroic filter D(3) by the telescope effect of micromirrors M(1)' and M(2)'. Spot A(2), an even spot, is the intermediate spot in collimated space.

[0031] The multiple long-pass dichroic filters D(1)-D(N) can alternatively be passband or bandpass optical filters, or can include both dichroic and bandpass optical filters in combination with each other to ensure that light of a limited wavelength range passes. The dichroic optical filters use the principle of thin film interference and can also be called interference filters. For each channel, the bandpass or passband optical filters are adjusted to pass a different selected range of wavelengths of light (passband) to each detector and reflect the remaining light wavelengths back to the micromirrors in the micromirror array.

[0032] To provide a compact detector module, the image array 106B is formed from a solid transparent material as will be further described with reference to Figures 8 and 10. The solid transparent material of the transparent imaging block used for the 1f image array can be, for example, clear glass or clear plastic with mirrors and dichroic filters formed in or on the transparent material.

[0033] The design of the 1f image array 106B and the design of the 2f image array 106A can be compared under the conditions of the same thickness of the solid transparent material, the same pitch, and the same angle of incidence. The path distance between adjacent micromirrors in the 2f image array 106A and the 1f image array 106B is approximately the same. In this case, the path distance is a physical distance, not a traditional path length where the refractive index of the material is taken into account. However, the focal length of the micromirrors in the 2f image array 106A is shorter (half) than that of the micromirrors in the 1f image array 106B. In other words, the actual focal length of the micromirrors in the 1f image array 106B is twice that of the micromirrors in the 2f image array 106A due to the different radii of curvature of the micromirrors. The longer the focal length, the less the bending force in each reflection, which minimizes the aberrations introduced in the micromirror reflections. Therefore, the aberrations in the 1f image array 106B are improved over those in the 2f image array 106A.

[0034] In the 2f imaging array system 106A, the fiber aperture A(0) is imaged to spot A(1) on dichroic filter D(1). The image at spot A(2) is a re-image of spot A(1) through reflection off micromirror M(1). Continuing through the zigzag optical path in the array 106A, each spot A(n) on 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 twice the focal length of micromirror M(n).

[0035] In the 1f imaging array system 106B, light from fiber aperture A(0) is imaged by the input channel to spot A(1) on dichroic filter D(1). In imaging array system 106B, the light path from micromirror M(1)' to micromirror M(2)' can be considered to form the equivalent of a telescope with a magnification equal to 1. In such a case, spot A(1) is imaged through the telescope of micromirrors M(1)' and M(2)' to spot A(3), with spot A(2) being an intermediate spot in collimated space. Adjacent micromirrors M(3)' and M(4)' form another telescope that reimages spot A(3) to spot A(5). Continuing along the zigzag path in imaging array 106B, the odd spots A(1), A(3), A(5)...A(2n+1) are all conjugates of one another, and the even spots A(2), A(4), A(6),...A(2n) are intermediate spots in collimated space. In this configuration, the path distance between spot A(n) and micromirror M(n)' is the thickness L of the transparent block, which is the focal length of one of the micromirrors M(n)'.

[0036] Each spot on the filter is formed by a bundle of rays whose angular distribution depends on the numerical aperture (NA) of the input multimode fiber 102 and the input channels to the imaging array of the imaging system. The cone angle of light at spot A(1) is proportional to the numerical aperture of the multimode fiber 102. If the image magnification m from fiber aperture A(0) to spot A(1) is in the ratio 1 to m, where m is a number greater than 1, then the cone angle at spot A(1) is m times smaller than the cone angle at fiber aperture A(0).

[0037] In the 2f imaging array system 106A of Figure 2A, the image magnification from any spot A(n) to an adjacent spot A(n+1) is equal to 1 for all values ​​of n. Neglecting aberrations, the cone angles of the light rays at the dichroic filters D(1)-D(N+1) of the imaging array system 106A are the same for all detector channels.

[0038] 2B, the distance between dichroic filter D(1) and micromirror M(1)' is L. If the total number of channels N is an even number in 1f image array 106B, spot A(1) is imaged to spot A(3) through micromirrors M(1)' and M(2)' (considering a 1× telescope), and spot A(2) at dichroic filter D(2) is in collimated space. A portion of the collimated light at spot A(2) is reflected by dichroic filter D(2) towards micromirror M(2)'.

[0039] Thus, 1f image array 106B has odd detector channels at spots A(1), A(3), ... A(2k-1) ... A(N-1) and even detector channels at spots A(2), A(4), ... A(2k), ... A(N), where 1 < k < (N / 2). Spot A(1) is at the front focal point of micromirror M(1)' and the path distance between adjacent micromirrors is twice the focal length, so all odd spots A(1), A(3), A(5) ... A(2n+1) are images of the fiber aperture (as shown in FIG. 2B by the ray convergence to the odd spots) and all even spots A(2), A(4), A(6), ... A(2n) are in collimated space (as shown in FIG. 2B by the parallel rays at the even spots). Thus, as shown in FIG. 5, the cone angle of odd spots on odd detector channels (odd cone angle CAO) differs from the cone angle of even spots on even detector channels (even cone angle CAE).

[0040] In the 1f image array 106B, the central wavelength and passband width of each optical dichroic filter D(1)-D(N) are different from each other. The central wavelength and passband width of each optical dichroic filter are designed to optimize the sampling of the fluorescence spectrum of the dyes to better and more accurately perform the unmixing of a large amount of different dyes. For example, assuming a fluorescence spectrum of light wavelengths from 400 nm to 800 nm and a 16 channel detection module, a light wavelength bandwidth of 25 nm can be analyzed respectively. For example, the first detector channel and the first dichroic filter D(1) can band pass and analyze light wavelengths from 400 nm to 425 nm with a central wavelength of 412.5 nm. Wavelengths outside of 400 nm to 425 nm are substantially filtered out and do not pass to the first detector in the first detector channel. The second detector channel and second dichroic filter D(2) can pass and analyze optical wavelengths from 425 nm to 450 nm with a central wavelength of 437.5 nm, with each succeeding detector channel and dichroic filter passing and analyzing each subsequent bandwidth in 25 nm increments, and the final or sixteenth detector channel and sixteenth dichroic filter D(16) can pass and analyze optical wavelengths from 775 nm to 800 nm with a central wavelength of 787.5 nm.

[0041] The features of the 1f image array 106B allow the initial optical signal to propagate through a larger number of detectors than the 2f image array 106A. The 1f image array 106B reduces off-axis aberrations by reducing the bending force at each mirror reflection. On the other hand, if the odd cone angle CAO is different from the even cone angle CAE for the dichroic filters in the odd and even channels, it is necessary to find the optimal magnification m at the input stage from the fiber aperture A(0) to the spot A(1) on 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 spot A(1) on the dichroic filter D(1) is optimized for both the odd and even detector channels in the 1f image array 106B.

[0042] In terms of spectral resolution, the performance of the dichroic filter decreases as the cone angle of the incident spot increases. In the 1f image array 106B, the cone angle of the spots in the odd channels is different from the cone angle of the spots in the even channels. Essentially, the cone angle of the odd channels is determined by the numerical aperture (NA) of the multimode fiber and the magnification factor m from the fiber aperture A(0) to the spot A(1) in the dichroic filter D(1). In contrast, the cone angle of the even channels is determined by the spot diameter of the spots in the odd channels.

[0043] Assume that the image magnification m is from the aperture A(0) in the fiber to the spot A(1) on the dichroic filter D(1) in the input channel. In the even detector channels (dichroic filters D(2k)), the cone angle is proportional to the magnification m. Meanwhile, in the odd detector channels (dichroic filters D(2k-1)), the cone angle is inversely proportional to the magnification m. The larger the magnification from the fiber aperture A(0) to the spot A(1), the smaller the cone angle in the odd detector channels (dichroic filters D(2k-1)) but the larger the cone angle in the even detector channels (dichroic filters D(2k)). In one example embodiment having a multimode fiber with NA=0.12, aperture diameter of 600 μm, and micromirror (filter) pitch of 5.5 mm, the recommended magnification m was modeled to be about 2. Obviously, other magnifications m can be determined using different inputs, and therefore the embodiments disclosed herein are not limited to 2× magnification. In the example presented for 1f image array 106B, both the numerical aperture (NA) and the number of re-imagings are reduced by a factor of two, allowing at least four times (4x) more detectors along the same length of array than 2f image array 106A.

[0044] Figures 3, 4A and 4B, and 5 show graphs of simulation results for a compact detector module with a 1f image array 106B and example input values ​​for one embodiment with a fiber with NA=0.12, aperture diameter of 600 μm, micromirror (filter) pitch of 5.5 mm, and a recommended magnification of 2×. The different colors of the light rays shown in Figures 3, 4A and 4B, and 5 are only for clarity purposes to show how light at different positions passes through the detector module.

[0045] Figure 3 shows one end of an optical fiber 102 that launches the fluorescent light into a detection system. Near the opposite end of the optical fiber 102 (not shown), a collecting objective lens with a high NA can be used to collect the fluorescent light from the aperture and couple the light into the opposite end of the optical fiber. The optical fiber 102 then collects the light from the objective lens and directs the light to the end shown in Figure 3. Near the end shown in Figure 3, the system can include a fiber numerical aperture converter that lowers the numerical aperture into free space to launch the fluorescent light into a detector array.

[0046] 3, the magnification m in the compact detection module 300 is achieved 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 ratio of the focal lengths of the collimating lens 302 and the focusing lens 304. For example, to set the magnification 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 further include an input portion of a transparent block (e.g., a wedge, which is a thick portion of a block, see FIG. 8) to the image array 106B before reaching the first dichroic filter D(1).

[0047] The collimating lens 302 receives the light emitted from the fiber 102 and collimates the light. The collimated light passes through a blocking filter 303 and enters into a focusing lens 304. The blocking filter 303 is used to clean out the laser light scattered into the collection optics near the opposite end of the optical fiber 102. The fluorescent spectrum of the light associated with the fluorochrome passes through the blocking filter 303 and enters into the focusing lens 304. The focusing lens 304 focuses the fluorescent spectrum of the light onto a first dichroic filter D(1) in the image array 106B to form an image at spot A(1). The image at spot A(1) is enlarged in size (e.g., diameter and area) by a factor m from the size at the fiber aperture A(0). The position of the lenses 302, 304 between the end of the fiber 102 and the image array 106B can be adjusted.

[0048] Compact detection module 300 further includes a 16-channel 1f image array 106B and 16 detector channels 313A-313P in communication with image array 106B (see, e.g., FIG. 6). In an alternative embodiment, a pair of 8-channel 1f image arrays (see, e.g., FIGS. 7A-7C) can be used in parallel to ease the imaging requirements of each compact image array. In a flow cytometer, two or more (e.g., three) of these compact image arrays can be used to increase the number of detector channels to more than 16 (e.g., 3 times 16 for 40 detection channels), as described with the detector module described with respect to FIG. 13.

[0049] The image array 106B is formed from a solid transparent block of material. The 16-channel 1f image array 106B includes 16 dichroic filters D(1)-D(16) on one side of the transparent block and 15 mirrors M(1)-M(15) on the other side. The image array does not require a mirror following the last detector channel 313P. Moreover, the last filter D(16) 314 can be a non-dichroic filter and can instead be a bandpass filter. In the case of a bandpass filter, there is no need to further reflect the incident light to another mirror or filter.

[0050] Each detector channel 313A-313P (collectively detector channel 313) in the detector array includes a focusing lens 316 and a detector 318 (an example is shown in FIG. 3). The detector 318 is packaged in a thin outline (TO) can package 320, and the focusing lens 316 is coupled to or integrated within the TO can package. The focusing lens 316 focuses the fluorescent light passing through the filter onto the small area size of the detector 318.

[0051] Referring now to FIG. 4A, the optical fiber 102 used to carry the captured fluorescent light signal from the image chamber to the detector array is a multimode optical fiber. Light exits the end face of the multimode fiber from various (but not all) locations across the diameter of the fiber, such as locations X1-X5. Lenses 303, 304 in the input channel shown in FIG. 3 focus the light in aperture A(0) to spot A(1) on the first dichroic filter D(1). There is a two-fold (2×) magnification from spot A(0) to spot A(1), so the spot size at aperture A(0) is smaller than the spot size at spot A(1). The different colors of the light beams emanating from different locations X1-X5 in the aperture shown in FIG. 4A are only for clarity purposes to show how light at different positions passes through the detector module. As shown in FIG. 4A, an optical axis 402 extends outward from the circular center of the end of the optical fiber 102. Light is launched out the end of the optical fiber 102 at a launch cone angle (CA) 404 relative to the optical axis 402 .

[0052] 3 shows the results of a simulation of the image array 106B and how light from different locations in the light beam is alternately imaged and collimated through multiple reflections of mirrors and dichroic filters. These results show all of the light reflected, but the dichroic filters D(n) at any particular position are different, allowing transmission of the light signals according to their respective passbands (only those at the last detector channel 313P are shown in FIG. 3).

[0053] 3 and 4B, in each detector channel 313, the desired wavelength range of the optical signal passing through the dichroic filter D(n) can be collected through a lens 316 and detected by a small aperture photosensitive detector 318. An additional bandpass filter 314 can be used instead or in addition in each detector channel. If other wavelengths of light are present in the dichroic filter D(n), those wavelengths are reflected to the next micromirror M(n) along the chain or series of micromirrors. The series or chain of dichroic filters D(n) splits the different ranges of optical wavelengths into the chain of detector channels 313A-313P.

[0054] The close-up of FIG. 5 shows the simulation results of how the optical light beam is alternately imaged and collimated through the reflective surfaces of the micromirrors and dichroic filters. On the odd-numbered dichroic filters (e.g., dichroic filters D(7), D(9), and D(11) shown in FIG. 5), the spots are images of the fiber aperture. That is, the light emitted from the fiber aperture is imaged on each filter surface of the odd-numbered dichroic filters. On the even-numbered dichroic filters (e.g., dichroic filters D(8), D(10), and D(12) shown in FIG. 5), the even-numbered spots (e.g., spots A(8), A(10), and A(12) shown in FIG. 5) are in a collimated space, where the light rays emitted from a certain point in the fiber aperture become a collimated beam. The beam direction in the collimated space of each even-numbered dichroic filter is slightly different at different points from the fiber aperture.

[0055] The flow cytometer may use one or more linear 16-channel compact wavelength detection modules to detect particle-associated light fluorescent signals. Alternatively or in addition, the flow cytometer may use one or more dual 8-channel compact wavelength detection modules to detect particle-associated light fluorescent signals.

[0056] Figures 6 and 7A-7C show embodiments of compact wavelength detection modules having the functionality of the 1f imaging array 106B shown in Figure 2B. Figure 6 shows a linear 16 channel compact wavelength detection module 600. Figures 7A-7C show a dual 8 channel compact wavelength detection module 700.

[0057] 6, a linear 16 channel compact wavelength detection module 600 includes an input stage (head) 601 and a detection module 614 mounted on a base 610. Light is coupled into the input stage (head) 601 by an optical fiber 102. The input stage (head) 601 includes a collimating lens 602, a long pass filter 603, a clean up optical blocker 604, and a focusing lens 605 mounted on an optical bench. The input stage (head) 601 sets up the magnification m of the initial spot size image A(1) on the first dichroic filter.

[0058] From the input stage 601, light is coupled into a detection module 614. The end of the input stage (head) 601 is coupled to a transparent wedge 607 that receives light from a focusing lens 605. The input stage (head) 601 and the detection module 614 are coupled to a flow cytometer chassis or base 610 which maintains their alignment with respect to one another.

[0059] The detection module 614 includes a 1f imaging array 608 and a detector / lens array 611. The imaging array 608 is an embodiment of the imaging array 106B of FIGS. 2B and 5. The imaging array 608 includes a transparent block 680 (see, e.g., block 806 of FIG. 8 and block 1006 of FIG. 10) that includes a wedge 607 and 15 micromirrors 612 on one side. On the other side of the transparent block 680 are 16 dichroic filters 609. The detector / lens array 611, an embodiment of the multiple detectors 313A-313P, includes multiple photodetectors D1-D16 (e.g., detector 318 of FIG. 3), each having a lens (e.g., lens 316 of FIG. 3) that focuses the split light into the photodetector.

[0060] Light coupled into image array 608 by input stage 601 is wavelength separated into detectors D1-D16 of detector / lens array 611. The 16 channel detection module analyzes a range of wavelengths (e.g., wavelengths from 400 nm to 800 nm).

[0061] To provide a different footprint that better fits on a test bench and to provide parallel processing, a dual 8-channel compact wavelength detection module can be implemented in place of the linear 16-channel compact wavelength detection module 600 .

[0062] 7A, a top view of a dual detection module 700 having a pair of 8-channel compact wavelength detection modules 714, 715 is shown. The compact wavelength detection module 700 includes an input stage (head) 701 in communication with a first 8-channel detection module 714 and a second 8-channel detection module 715, all of which are aligned and mounted on a base 710. The first 8-channel detection module 714 splits and analyzes in parallel a first range of wavelengths (e.g., 650 nm to 800 nm-red wavelengths). The second 8-channel detection module 715 splits and analyzes in parallel a second range of wavelengths (e.g., 400 nm to 650 nm-blue wavelengths).

[0063] Light emitted from the optical fiber 102 is coupled into the input stage (head) 701. The light from the fiber 102 passes through a collimating lens 702 into a long-pass dichroic filter 703. The long-pass dichroic filter 703 reflects light at the laser excitation wavelength (e.g., less than 400 nm) at a 45 degree angle to a scatter detector (not shown). Side scatter (SSC) light can be focused onto a small aperture detector with a ball lens similar to that described for the fluorescence light. Fluorescence light in the fluorescence light spectrum (e.g., 400 nm to 800 nm) passes through the long-pass filter 703 into a second clean-up filter 704. The clean-up filter 704 ensures that the excitation laser light does not reach the splitter detection modules 714 and 715.

[0064] After the clean-up filter 704, the fluorescent light is separated into long and short wavelength bands by a long pass filter 705. The long wavelength light (e.g., red - 650 nm to 800 nm) passes through the long pass filter 705 and is focused by a collimating / focusing lens 706 into a first detection module 714. The long wavelength portion of the light that passes through and is focused by the focusing lens 706 is separated by the first detection module 714. The short wavelength light band (e.g., blue - 400 nm to 650 nm) is reflected by the long pass filter 705 at an angle into a collimating / focusing lens 713. The collimating / focusing lens 713 focuses the short wavelength band of light into a second detection module 715. The short wavelength portion reflected by the long pass filter 705 is separated by the second detection module 715. Alternatively, filter 705 can be a short-pass filter, where short wavelength light passes through the filter and is separated by the first detection module 714, while long wavelength light is reflected by the filter and separated by the second detection module 715.

[0065] Referring to the first detection module 714, light from the focusing lens 706 is incident perpendicularly on the 12 degree wedge surface 707 and passes through a transparent block (e.g., block 806 in FIG. 8) of the image array 708 before being imaged onto a first dichroic or bandpass filter 709. The light passes through the bandpass filter 709 and is focused onto a first small area detector D1 in the detector / lens array 711. The light rejected by the bandpass filter 709 is reflected back onto a first micromirror M(1) 712 of the multiple micromirrors M(1)-M(7) in the image array. The first micromirror M(1) 712 collimates the light and reflects it onto a second detection module D2. This operation continues along the serial chain of micromirrors and detection modules of the transparent blocks of the image array 708. The second detection module 715 functions similarly to the first detection module 714.

[0066] The reflected waves travel through the imaging arrays 106B, 708, 708' in the first and second detection modules 714, 715, respectively, as described herein, and the light is alternately focused and collimated through dichroic filters with successively shorter bandpasses of the light to the odd and even detectors 118 in the odd and even detector channels, respectively. Thus, the different wavelengths are separated by the multiple detectors in the first and second detection modules 714, 715, respectively.

[0067] For a given fluorescent event, the signal from each detector (e.g., detector 318 shown in FIG. 4B, lenses / detectors D1-D16 in FIGS. 6 and 7) is amplified, digitized and synchronized by the electronics system to provide a spectral representation of the input fluorescent light signal. Integrating the detection electronics within the optical module assembly allows for compact designs and lower noise by minimizing detector coupling lengths and amplification circuitry. Detector 318 shown in FIG. 4B converts optical signals, such as the input fluorescent light signal, into electrical signals.

[0068] 7B and 7C show right and left perspective views, respectively, of a dual detection module 700 having a pair of eight-channel compact wavelength detection modules 714, 715. Each of the detection modules 714, 715 includes a mounting base 720 and a cover 722 that surrounds a mounting block 1200 (see FIG. 12) to which the lenses / detectors 711 in the detector array or detector chain are attached. The mounting base 720 and cover 722 maintain the elements of the image array 708, 708' within the transparent block 806, 1006 in mutual registration with the detector arrays in each detector module 714, 715. The mounting base 720 of each detection module 714, 715 is coupled to the base 710 by a number of fasteners.

[0069] Input stage 701 includes an optical bench 751 having a number of filter slots to accommodate filters 703-705, a number of lens slots to accommodate lenses 702, 706, 713, and one or more optical channels through which light reflects and propagates through the filters and lenses. Optical bench 751 is coupled to base 710 of detection module 700 to maintain alignment with detection modules 714 and 715.

[0070] 13 shows a top view of an optical plate assembly 1300 in a modular flow cytometry system 100. The optical plate assembly 1300 includes a laser system 1370 having three semiconductor lasers 1370A, 1370B, 1370C that direct excitation into a flow cell assembly 1308 through which sample fluid flows with sample particles. The laser system 1370 attempts to direct multiple (e.g., three) laser beams collinearly toward the flow cell assembly 1308. However, these multiple laser beams may be slightly offset from one another. The laser system 1370 includes semiconductor lasers 1370A, 1370B, 1370C that typically have wavelengths 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 output power.

[0071] An optical system spatially manipulates the optical laser beams 1371A, 1371B, 1371C generated by the semiconductor lasers 1370A, 1370B, 1370C, respectively. The optical system includes lenses, prisms, and steering mirrors that focus the optical laser beams onto the fluid stream carrying the biological cells (bio cells). The focused optical laser beam size is typically focused across the flow stream when it is between 50 microns and 80 microns (μm), and typically focused along the stream flow in the flow cell assembly 1308 when it is between 5 μm and 20 μm. In FIG. 13, the optical system includes beam shapers 1330A-C that receive the laser light 1371A, 1371B, 1371C from the semiconductor lasers 1370A-C, respectively. The laser beams output from the beam shapers 1330A-1330C are coupled into the mirrors 1332A-1332C, respectively, which direct the laser beams 1399A, 1399B, 1399C into the flow cell assembly 1308 to target particles (e.g., biological cells) stained with a fluorescent dye. The laser beams 1399A, 1399B, 1399C are slightly separated from each other but are directed substantially parallel into the flow cell assembly 1308 by the mirrors 1332A-1332C.

[0072] The laser light beams 1399A, 1399B, 1399C reach biological cells (particles) in the flow stream at the flow cell assembly 1308. The laser light beams 1399A, 1399B, 1399C are then scattered by the cells in the flow stream, causing fluorochromes to fluoresce and produce fluorescent light. A forward scatter diode 1314 collects the on-axis scattered light. A collection lens 1313 collects the off-axis scattered light and the fluorescent light and directs them together to a dichromatic mirror 1310. The dichromatic mirror 1310 focuses the off-axis scattered light onto a side scatter diode 1315. The dichromatic mirror 1310 focuses the fluorescent light onto at least one fiber head 1316. At least one fiber assembly 102 sends the fluorescent light towards at least one detector module 600, 700.

[0073] To analyze the biological sample in more detail using different fluorescent dyes and laser wavelengths, multiple fiber heads 1316, multiple fiber assemblies 102, and multiple detector modules 600, 700 can be used. Three fiber heads 1316A, 1316B, 1316C can be arranged in parallel to receive the fluorescent light, and three fiber assemblies 102A, 102B, 102C can be used to direct the fluorescent light to three detector modules 600A, 600B, 600C or 700A, 700B, 700C.

[0074] The three laser light beams 599A, 599B, ​​599C are slightly offset (e.g., not exactly collinear) to enable three fiber heads 1316A, 1316B, 1316C (and three fiber assemblies 102A, 102B, 102C). Thus, the three fiber heads 1316A, 1316B, 1316C can collect light beam data separately from the three laser light beams 599A, 599B, ​​599C having three different wavelengths. The three fiber assemblies 102A, 102B, 102C then direct the light into three different detector modules (e.g., three different detector modules 600A, 600B, 600C or 700A, 700B, 700C).

[0075] Alternatively, the modular flow cytometry system can collect light beam data using one detector module 600, 700. For example, three fiber assemblies 102A, 102B, 102C can direct light into one detector module 600, 700 as opposed to three different detector modules. The separation of the light beam data is then treated as a data processing operation instead of separating the light beam data by using three different detector modules. By using one detector module, the complexity can be reduced in terms of physical devices. On the other hand, the data processing operation can be more complicated because the separation of the light beam data requires more data manipulation (e.g., identifying different wavelengths and separating the light beam data accordingly).

[0076] Cell geometric features can be classified by analysis of forward and side scatter data. Cells in the fluid flow are labeled with dyes of visible wavelengths in the range of 400 nm to 900 nm. The dyes produce fluorescent light when excited by a laser. These lights are collected by the fiber assembly 102 and directed towards the detector modules 600, 700. The modular flow cytometry system maintains a relatively small size for the optical plate assembly due to the compact semiconductor laser, 11.5x magnification collection lens 1313, and compact imaging arrays in the detector modules 600, 700.

[0077] The collection lens 1313 contributes to the design of the detector module 600, 700. The collection lens 1313 has a short focal length for 11.5× magnification. On the side of the objective lens facing the direction of the fluorescence emission, the collection lens 1313 has a high numerical aperture (NA) of about 1.3 to capture more photons in the fluorescence emission over a wide range of incidence angles. On the side facing the collection fiber 102, the collection lens 1313 has a low NA of about 0.12 to launch the fluorescent light into the fiber over a narrow cone angle. The collection lens 1313 thus converts the NA from a high NA on one side to a low NA on the other side to support the magnification m at the input channel of the detector module 600, 700.

[0078] The core diameter of the collection fiber assembly 517 is between about 400 μm and about 800 μm, with a fiber NA of about 0.12 for a core diameter of about 600 μm. The fiber output end can be tapered to a core diameter of between about 100 μm and about 300 μm to control the image size onto the receiving photodiode.

[0079] The input end of the collection fiber 102 may also include a lensed fiber end to increase the collection NA, allowing the use of fiber core diameters less than about 400 μm. Because the fiber 102 has the flexibility to distribute the light anywhere within the flow cytometer system, the use of fiber for collection of the fluorescent light allows for optimization of the location of the receiver assembly and electronics in a compact flow cytometer system.

[0080] To produce a low-cost flow cytometer, lower-cost components can be implemented. The imaging array 106B in each detection module 614, 714, 715 is formed from a solid transparent material to provide a detection module that is reliable, low-cost, and compact. Additionally, the flow cytometer uses low-cost off-the-shelf thin-shell (TO) can detectors.

[0081] 12, there is shown a mounting block 1200 adjacent to the transparent blocks 806, 1006 (see FIGS. 8-11) which are joined together to form a 1f imaging array 708, 708' which is attached to the base 710 of the compact detector module 700 shown in FIGS. 7A-7C by a mounting base 720 and a cover 722. The mounting block 1200 includes a curved aperture 1201 with multiple angles to accommodate multiple TO can lenses / detectors 711. The alignment of the mounting block 1200 with the transparent blocks 806, 1006 of the imaging array 708 and the angle of the curved aperture 1201 are such that light reflected from the micromirror 712E can be bandpass filtered by the dichroic filter 709E and coupled into the lens / detector 711E.

[0082] Each of the plurality of TO can lenses / detectors 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 solid-state photodetector 1213 inside the TO can package. The solid-state photodetector 1213 is electrically coupled to a plurality of electrical pins 1214 that extend outside the TO can package. These pins are electrically coupled to other electronics of the flow cytometer. Similar to detector 318 shown in FIG. 4B, the solid-state photodetector 1213 converts an optical signal, such as an input fluorescent optical signal, into an electrical signal on at least one electrical pin 1214.

[0083] 8, a perspective view of a transparent block 806 formed from a solid transparent material 800 for one embodiment 106B, 608, 708, 708' of a 1f image array is shown. The solid transparent material 800 used for the transparent block 806 can be, for example, transparent glass or transparent plastic. A plurality of micromirrors 810 in a row and a series 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 series chain are formed in or on the other side of the transparent block 806 of the transparent material 800. Each dichroic or bandpass filter 812 is tuned to a different range of light wavelengths to allow detection of a wide range of fluorescent light emitted by the fluorescent dyes. In one embodiment, the plurality of micromirrors 810 are concave spherical mirrors.

[0084] A transparent block 806 formed from solid transparent material 800 further includes a 12 degree wedge surface 820 that receives light from a focusing lens, such as that described with 1f imaging array 708. Light is incident normal to the surface of wedge surface 820 and is directed (bent) toward a first dichroic filter or bandpass filter D(1). The light passes through transparent block 806 and reaches the first dichroic filter or bandpass filter D(1).

[0085] 10, a perspective view of a transparent block 1006 formed from a solid transparent material 800 for another embodiment 106B, 608, 708, 708' of a 1f image array is shown. The solid transparent material 800 used for the transparent block 1006 can be, for example, transparent glass or transparent plastic. A plurality of micromirrors 1010 are concave rectangular mirrors formed within one side of the transparent material. A plurality of dichroic or bandpass filters 1012 are formed within or on the opposite side of the transparent material 800. Each dichroic or bandpass filter 812 is tuned to a different range of light wavelengths to allow detection of a wide range of fluorescent light emitted by the fluorescent dyes.

[0086] The solid transparent material 800 further includes a 12 degree wedge surface 820 that receives light from a focusing lens, such as that described with the imaging array 708. Light is incident normal to the surface of the wedge surface 820 and is directed (bent) toward a first dichroic filter or bandpass filter D(1). The light passes through a block to reach the first dichroic filter or bandpass filter D(1).

[0087] 9A, a cross-sectional view shows the distance (e.g., thickness L) between a spherical micromirror 810 on one side of a 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 shows an axis 815 perpendicular to the transparent block at the center of a dichroic or bandpass filter 812. The axis 815 extends to the opposite side of the transparent material 800 from the transparent block 806. The axes 814 and 815 are parallel to each other.

[0088] A reflective material 811 is formed (e.g., arranged) in the shape of a spherical transparent microlens in solid transparent material 800 to form each spherical micromirror 810 on one side of transparent block 806. A dichroic or bandpass filter 812 is coupled to material 800 on the opposite side of transparent block 806.

[0089] Figure 11 also shows the distance and axis 1014 between the concave square micromirror 1010 on one side of the transparent material 800 forming the transparent block 1006 and the opposite side. Figure 11 further shows a perpendicular axis 1015 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 by the transparent material 800. The optical axes 1014 and 1015 are parallel to each other.

[0090] 11 further shows that the reflective material 1011 is formed (e.g., disposed) in a curved transparent rectangular shape of a solid transparent block 1006 formed from the transparent material 800 to form a square micromirror 1010. A dichroic filter or bandpass filter 1012 is coupled to the opposite side of the solid transparent block 1006.

[0091] Fluorescent dyes used in flow cytometry applications cover the entire visible and near infrared wavelength range. The emission wavelength bandwidth is usually large for long wavelength fluorescent dyes. Each of the dichroic or bandpass filters 812 can have their detector filter passbands and center wavelengths optimized to measure different dyes with the same amount of spectral sampling. Furthermore, individual filter optimization allows for the exclusion of excitation wavelengths from other lasers. In this way, the detectors in each channel can be fully utilized to detect signals of interest. In conjunction with the fluorescence spectrum unmixing algorithm executed by the computer processor, the individual optimized passband detection provides the ultimate detection of multiple fluorescent dyes of interest.

[0092] How to use different detection systems We now describe how to use the various detection systems disclosed herein in a flow cytometer. Fluorescent light of different wavelengths is generated by various fluorescent dyes that mark different particles in the sample in the flow channel excited by the laser light, before emitting the fluorescent light generated by the fluorescent dyes excited by the laser light from the end of the optical fiber 102 shown in the figure. The generated fluorescent light is received by a collection lens near the opposite end of the laser, as can be seen in FIG. 13. To better match the numerical aperture of the optical fiber, a converter is used that converts from a first numerical aperture at the capture side to a second numerical aperture that is smaller than the first numerical aperture. The optical fiber then directs the fluorescent light towards the end of the optical fiber, which flexibly directs the light towards the compact detection module 600, 700.

[0093] The optical fiber 102 couples the fluorescent light into the end of the optical fiber, thereby emitting the fluorescent light outward from the optical fiber. The emitted fluorescent light has different wavelengths produced by different fluorescent dyes attached to different particles in the sample fluid that are excited by the laser light.

[0094] In the input channel, light emitted from the end of the optical fiber is collimated and focused by a lens towards a first dichroic filter of a first plurality of dichroic filters in a first splitting imaging array.

[0095] Further along the input channel, laser light used to excite the different fluorochromes emerging from the optical fiber is prevented from interfering with the detection of the fluorescent light wavelengths by a blocking device.

[0096] Further along the input channel, the image size from the end of the optical fiber is expanded to the spot size of a first dichroic filter in a series chain or series of a first plurality of dichroic filters in a first demultiplexing imaging array.

[0097] In the first demultiplexing imaging array, a first wavelength range of the fluorescent light is alternately reflected between a first plurality of dichroic filters and a serial chain or row of a first plurality of micromirrors, collimating the fluorescent light onto the odd numbered dichroic filters and re-imaging the fluorescent light onto the even numbered dichroic filters. The focal lengths of the first plurality of micromirrors and the separation distance between the first plurality of dichroic filters and the first plurality of micromirrors provide a telescope effect along the chain of micromirrors that collimates the fluorescent light onto the odd numbered dichroic filters and re-images the fluorescent light onto the even numbered dichroic filters.

[0098] In the series chain or array of the first plurality of dichroic filters, different wavelength ranges of the first wavelength range of the fluorescent light are band-passed in each to demultiplex the wavelength spectrum of the first wavelength range of the fluorescent light.

[0099] Adjacent to the series chain or array of a first plurality of dichroic filters is a series chain or array of a plurality of detector channels as shown in Figures 3, 6, and 7A-7C having a first plurality of photodetectors or photodetectors, each detector channel having a lens that focuses a different wavelength range of the fluorescent light into the first plurality of photodetectors or photodetectors.

[0100] A serial chain or array of multiple photodetectors detects fluorescent light in each of the different wavelength ranges of the first wavelength range associated with each fluorescent dye tagged to the particle. The multiple photodetectors convert the fluorescent light received by each of them into an analog electrical signal. Each analog electrical signal can then be digitized / converted by an analog-to-digital (A / D) converter from an analog electrical signal to a digital electrical signal, which can then be analyzed and counted.

[0101] The fluorescent light is converted to an analog electrical signal by the detector and digitized to a digital electrical signal, so that a computer having a processor can then be used to count the number of each of the different particles in the sample fluid, as disclosed in U.S. Patent Application No. 15 / 498,397, entitled "COMPACT MULTI-COLOR FLOW CYTOMETER," filed April 26, 2017 by David Vrane et al., which is incorporated herein by reference.

[0102] A second and / or a third demultiplexing imaging array can be used in parallel with the first demultiplexing imaging array. In this case, the method further includes splitting the fluorescent light into a first wavelength range of fluorescent light for the first demultiplexing imaging array, a second wavelength range of fluorescent light for the second demultiplexing imaging array, and / or a third wavelength range of fluorescent light for the third demultiplexing imaging array. As shown in FIG. 13, a first optical fiber 102A can be used to guide the fluorescent light toward the first demultiplexing imaging array. A second optical fiber 102B can be used to guide the fluorescent light toward the second demultiplexing imaging array. A third optical fiber 102B can be used to guide the fluorescent light toward the third demultiplexing imaging array.

[0103] The steps described herein for the first non-linear imaging array can be performed simultaneously by the second non-linear imaging array and / or the third non-linear imaging array so that different additional wavelength ranges can be analyzed. For the sake of brevity, steps that are repeated will not be repeated but are incorporated herein by reference.

[0104] Computer System Overview 14, a basic conceptual diagram of a flow cytometer system 1400 is shown. Various embodiments of the flow cytometer 1400 may be commercially available. Five major subsystems of the flow cytometer system 1400 include an excitation optics system 1402, a fluidics system 1404, an emission optics system 1406, an acquisition system 1408, and an analysis system 1410. In general, a "system" includes hardware devices, software, and / or combinations thereof.

[0105] The excitation optics system 1402 includes, for example, a laser device 1412, an optical element 1414, an optical element 1416, and an optical element 1418. Examples of optical elements include an optical prism and an optical lens. The excitation optics system 1402 illuminates an optical interrogation region 1420. The fluidics system 1404 conveys a fluid sample 1422 through the optical interrogation region 1420. The emission optics system 1406 includes, for example, an optical element 1430 and optical detectors SSC, FL1, FL2, FL3, FL4, and FL5. The emission optics system 1406 collects photons emitted or scattered from particles passing through it. The emission optics 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. Optical detectors FL1, FL2, FL3, FL4, and FL5 are fluorescence detectors that may include bandpass or longpass filters to detect specific fluorescence wavelengths. Each optical detector converts photons into electrical pulses and transmits these electrical pulses to acquisition system 1408. Acquisition system 1408 processes these signals and prepares them for analysis in analysis system 1410.

[0106] 1-13 described herein include exemplary elements of the excitation optics system 1402, emission optics system 1406, and acquisition system 1408 of a flow cytometer 1400. U.S. patent application Ser. No. 15 / 817,277, entitled "FLOW CYTOMETERY SYSTEM WITH STEPPER FLOW CONTROL VALVE," filed Nov. 19, 2017 by David Vrane, describes exemplary elements of the fluidics system 1404. This U.S. patent application is incorporated herein by reference.

[0107] Analysis system 1410 can store the digital representations of the signals for analysis after acquisition system 1408 has completed acquisition of those digital representations. Analysis system 1410 is a computer that includes a display device, at least one processor, scratchpad memory, and one or more non-volatile storage devices that can store more semi-permanently data (including signal data) and software.

[0108] The one or more storage devices of the analysis system can store analysis software having instructions that can be executed by the processor to obtain test results of a biological sample (or other type of sample, e.g., a chemical sample) being analyzed. The analysis system 1410 and analysis software can further be used to calibrate the flow cytometer with compensation controls when initialized with control particles that are passed through the flow cytometer prior to passing a biological sample through the flow cytometer. The analysis software has instructions executed by the processor to analyze stored signal data acquired by the acquisition system 1408 during calibration with the controls that are passed through the flow cytometer. The analysis software also has instructions executed by the processor to analyze stored signal data acquired by the acquisition system 1408 during analysis of a biological sample that is passed through the flow cytometer after calibration of the flow cytometer.

[0109] Overview of the method 15 and 16A-16C, a flow chart of a method 1500 for performing flow cytometry with high speed correction is described. For example, the flow cytometry system 1400 of FIG 14 can perform the method 1500. Flow cytometry allows for high speed multi-parameter collection and analysis of data on single cells or particles within a sample fluid.

[0110] In step 1501 shown in FIG. 15, the flow cytometry system starts the flow cytometer 1604, 1614 and checks the operation of the flow cytometer. The flow cytometer can be a conventional flow cytometer 1604 or a spectral flow cytometer 1614. The conventional flow cytometer 1604 conventionally uses mirrors, optical filters, and photomultiplier tubes (PMT) to capture or detect signal data. The spectral flow cytometer 1614 is based on similar principles, but more typically uses a multi-channel detector (e.g., a charge-coupled device (CCD)) that has a means to distribute the light spectrum across multiple channels into the detector.

[0111] In step 1505, the system begins calibrating the flow cytometers 1604, 1614 as shown by Figure 16A. The system uses one or more single-stained reference controls or compensation controls 1602A-1602E, collectively referred to as single-stained reference controls or compensation controls 1602, that are run one at a time through the flow cytometers 1604, 1614. The one or more single-stained reference controls 1602A-1602E individually generate spectral profiles or signatures 1606A-1606E from the respective various autoluminescence or luminescence of fluorochromes tagged to the control particles.

[0112] During calibration, in step 1510, the system uses a single-stained compensation control 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 multi-color flow cytometry, the system uses a single-stained sample 1602 passed through the flow cytometer 1604, 1614 to determine a level of compensation as shown in FIG. 16A. A single stain of a particle 1602 can present a spectral profile or signature 1606A-1606E of a fluorescent dye to a fluorescent photodetector of the flow cytometer. This allows the system and / or user to compensate for overlaps (e.g., see overlap 1616 in FIG. 16B) in the spectral profile of fluorescence from the sample 1610 due to multi-stained cells or particles 1612A-1612E contained in the sample, as shown in FIG. 16B.

[0113] The staining of the compensation control 1602 should be as bright as or brighter than the particles 1612A-1612E in the sample 1610. If the fluorescence measured for the control is brighter, antibody capture beads can be substituted for cells, and one fluorophore-conjugated antibody can be substituted for another. The exception to this is tandem dyes, which are not substitutable. Tandem dyes from different vendors or different batches should be treated like individual dyes, and individual single-stained controls should be used for each, since the amount of spillover can be different for each of these dyes. Also, the compensation algorithm should be performed with positive and negative populations. Whether each individual compensation control contains beads, cells used in the experiment, or different cells, the control itself should contain particles with the same level of autofluorescence. The entire set of compensation controls can contain individual samples of either beads or cells, but the individual samples must have the same carrier particles of the fluorophore. Also, the compensation control uses the same fluorophore as the sample. For example, green fluorescent protein (GFP) and fluorescein isothiocyanate (FITC) both emit mostly green photons but have very different emission spectra. Therefore, the system cannot use one of them for the sample and the other for the compensation control. Also, the system must collect enough events (e.g., about 5000 events for both the positive and negative populations) to make a statistically significant determination of spillover.

[0114] During calibration in a conventional flow cytometer 1604, the system obtains an initial spillover matrix from single-stained reference controls 1602 that are passed one by one through the conventional flow cytometer. In the conventional flow cytometer 1604, the fluorescent signals (e.g., colors) are separated into discrete fluorescent bands 1606A-1606E using a series of edge filters and dichroic mirrors. The system detects (e.g., measures) each individual channel using a photomultiplier tube (PMT). Ideally, the fluorescent bands would be perfectly discrete, but that is not reality. During detection of the fluorescent signals in the controls 1602 (and samples 1610), "spillover" can occur between the fluorescent bands, as shown in the binding profile 1616. The system uses the spillover matrix [S] to define the spillover between the fluorescent bands (e.g., the spillover 1618 between the discrete bands of the binding profile 1616 shown in FIG. 16B).

[0115] Alternatively, during calibration in a spectral flow cytometer 1614, the system obtains an initial reference matrix from a single-stain reference control 1602 that is run through the spectral flow cytometer. Spectral flow cytometry is a technique based on traditional flow cytometry, but in spectral flow cytometry, a spectrograph and a multichannel detector (e.g., a charge-coupled device (CCD)) are used in place of the traditional mirrors, optical filters, and photomultiplier tubes (PMTs) of traditional flow cytometer systems. In a spectral flow cytometer, side-scattered light and fluorescent light are collected and coupled directly or through optical fibers into a spectrograph, where the entire light signal is dispersed and displayed as a high-resolution spectrum on a CCD or multichannel detector.

[0116] 15 and 16B, in step 1515, the system passes a sample 1610 through a flow cytometer 1604, 1614. The system generates, acquires, and / or records data (e.g., event data) representing a spectral profile 1616 based on the sample fluid 1610. For example, fluorescing cells are detected within the sample fluid flowing through the flow cytometer. Each detection of a fluorescing cell is an event. The event data can be defined according to a measured sample event vector.

[0117] In step 1520, the system generates a compensated sample event vector (in the case of a conventional flow cytometer 1604) or an unmixed sample event vector (in the case of a spectral flow cytometer 1614) to count the number of different types of cells or particles in the sample 1610 and obtain a measure of concentration. Typically, as shown in FIG. 16C, an inverse matrix 1624 (derived from an initial spillover matrix and / or an initial reference matrix using fine-tuning) is used on the event data representing the spectral profiles 1616 to generate a compensated sample event vector or an unmixed sample event vector representing the individual spectral profiles or signatures 1606A-1606E of the various autoluminescence of the various cells 1612A-1612E in the sample 1610 or the luminescence of the fluorochromes tagged to the cells 1612A-1612E.

[0118] In the case of a conventional flow cytometer 1604, the system calculates a corrected event vector based on the initial spillover matrix and the measured sample event vector. In the case of a spectral flow cytometer 1614, the system calculates an unmixed sample event vector based on the initial reference matrix and the measured sample event vector.

[0119] Unfortunately, the initial spillover matrix and the reference matrix tend not to be accurate enough to obtain reliable results. Thus, in step 1525, the system performs a fast correction. In the case of a conventional flow cytometer 1604, this step involves correcting for inaccuracies in the initial spillover matrix. In the case of a spectral flow cytometer 1614, this step involves correcting for inaccuracies in the initial reference matrix.

[0120] In step 1527, based on the fast correction, the system generates a re-corrected sample event vector in a faster manner than was previously performed. Method 1500 then ends in step 1530, although one or more steps can be repeated for other samples of similar dates / times of calibration / sampling with the flow cytometer.

[0121] Obtaining spillover matrices from single stain controls on a conventional flow cytometer Thus, the system can include a conventional flow cytometer that generates or acquires a spillover matrix from the single stained controls. The step of generating or acquiring a spillover matrix by using a conventional flow cytometer is further discussed.

[0122] 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. Examples of compensation controls include beads 1602 stained or dyed with fluorescent dyes such as fluorescein isothiocyanate (FITC), R-phycoerythrin (PE), peridinin chlorophyll protein complex (PerCP), phycoerythrin and a cyanine dye (PE-Cy7), allophycocyanin (APC), and a tandem fluorescent dye combining APC and a cyanine dye (APC-Cy7).

[0123] Assume that the vector {U} is a measured sample event vector with N values, each of which is from one of the N detectors detecting a compensation control (e.g., FITC, PE, PerCP, PE-Cy7, APC, APC-Cy7).

[0124] Assume that vector {V} is a corrected sample-event vector with N values. The measured sample-event vector {U} is equal to the spillover matrix [S] multiplied by the corrected sample-event vector {V}. This can be expressed using the following matrix relationship with the measured sample-event vector {U}:

number

[0125] Therefore, the inverse spillover matrix [S] -1 Using the above, the corrected sample event vector {V} can be obtained from the following matrix equation:

number

[0126] The initial spillover matrix [S] can be obtained by measuring each single-stained control (e.g., FITC, PE, PerCP, PE-Cy7, APC, APC-Cy7) in each detector to obtain the following matrix:

number

[0127] Each column in the initial spillover matrix [S], which is a separation of variables (SOV) matrix, corresponds to one single stain control (e.g., FITC, PE, PerCP, PE-Cy7, APC, APC-Cy7). For example, column 1 corresponds to the FITC single stain control. As another example, column 2 corresponds to the PE single stain control, and so on for each single stain control run to calibrate the flow cytometer. Each row in the initial spillover matrix [S] corresponds to a given detector number. For example, row 1 corresponds to detector 1. Row 2 corresponds to detector 2, and so on.

[0128] Generally, the initial spillover matrix generated is not accurate enough to accurately separate the spectra and accurately identify cells or particles. Therefore, fine-tuning of the off-diagonal element values ​​of the initial spillover matrix [S] (e.g., fine-tuning to the initial spillover matrix [S] to generate an adjusted spillover matrix [S]' and its associated inverse, an adjusted correction matrix [C]') is required. These fine-tuning can be done based on the experience and judgment of the laboratory technician / operator. Fine-tuning is often done to correct distortions caused by interactions of fluorochromes stained on the same cell or particle, or distortions caused by the system measuring single stained and unstained controls, or both the interaction and the system.

[0129] The adjustment matrix [D] is assumed to be a fine adjustment made (e.g., added) to the off-diagonal element values ​​of the initial spillover matrix [S]. The matrix equation {V R}=[[S]+[D]] -1 The recalculated event vector {V R} can be obtained.

[0130] Acquiring Unmixed Event List Data for Spectral Flow Cytometers Alternatively, the system may include a spectral flow cytometer 1614 that generates or acquires the unmixed event list data. Generating or acquiring the unmixed event list data by using a spectral flow cytometer is further discussed.

[0131] Assume that [R] is the N x M initial reference matrix of the spectral flow cytometer obtained from single-stained reference controls, where N is the number of detectors and M is the number of fluorochromes being measured (e.g., FITC, PE, PerCP, PE-Cy7, APC, APC-Cy7), and M is always smaller than N.

[0132] Assume that {U} is a measured sample event vector with N values, each of the N values ​​being from one of the N detectors.

[0133] Assume {V} is an unmixed sample-event vector with M values ​​(e.g., fluorescence intensity), where each of the M values ​​is the unmixed value of a fluorescent dye (e.g., one of FITC, PE, PerCP, PE-Cy7, APC, APC-Cy7).

[0134] The unmixed sample event vector {V} has the following matrix relationship with the measured sample event vector {U}:

number

[0135] Since the number of variables M in the unmixed sample event vector {V} is less than the number of variables N in the measured sample event vector {U} (e.g., the dimension of the unmixed sample event vector is less than the dimension of the measured sample event vector), in that case the system uses a least squares algorithm to obtain a solution to the above equation.

[0136] In comparison to conventional flow cytometers, the unmixed event vector is equivalent to the corrected event vector. Therefore, the spectral spillover matrix of the unmixed event list data (e.g., the unmixed sample event vector) is the identity matrix [I] as follows:

number

[0137] In general, the unmixed event list data is not accurate enough, so a fine-tuning of the identity spectral spillover matrix (e.g., to produce a tuned spectral spillover matrix) is required. Thus, the equation for the recorrected event vector in a spectral flow cytometer is: {V R}=[[I]+[D]] -1 {V}, where [D] is a fine adjustment δ i,j and n×n delta matrix with zeros where no fine-tuning is needed. For example, the delta matrix is

number

[0138] Fast correction of flow cytometry data Thus, in flow cytometry (e.g., conventional and spectral flow cytometry), the Flow Cytometry Standard (FCS) data collected from the cytometer is linear, unprocessed list data. The list data does not need to be corrected before consumption in plots and use in statistical analysis of cells in a sample. The system performs high-speed corrections to address insufficient precision in the spillover matrix of conventional flow cytometers and / or the unmixed event list data of spectral flow cytometers.

[0139] The correction of the list data is based on an initial spillover matrix that the system obtains from measured single stain correction controls and / or fine-tuning inputs. The obtained initial spillover matrix is ​​generally not accurate enough. Fine-tuning is performed to generate an adjusted spillover matrix by fine-tuning the values ​​in the initial spillover matrix.

[0140] Each time the spillover values ​​are tweaked, the spillover matrix needs to be inverted to obtain a correction matrix, which is then multiplied by each list data event vector to generate the corrected list data (e.g., recorrected event vector).

[0141] For example, take an experiment with N fluorescence parameters. For each event vector correction, N 2 In addition to multiplications, N × (N-1) additions are required. The computational complexity is N 2 order (e.g., O(N 2 )).

[0142] For experiments with a limited number of fluorochrome parameters and a limited number of events, the correction calculations may not be a bottleneck in flow cytometry data analysis. However, when an experiment includes a large number of events (e.g., 2 million events) as well as a large number of fluorescent parameters (e.g., more than 20 fluorescent parameters), the correction calculations may take a very long time. As a result, the response of displaying plots and statistics on the computer interface may be very slow due to the large amount of calculations being processed every time the system changes the spillover value.

[0143] Advantageously, upon receiving or performing fine-tuning of the spillover matrix of the flow cytometry data analysis, the system executes a fast correction algorithm that significantly reduces the amount of calculations without sacrificing any accuracy of the corrected list data. The fast correction algorithm, for example, requires only (3N+1) multiplications / divisions plus (N+1) additions. The complexity of the fast correction algorithm is of the order of N (e.g., O(N)). Thus, the system can significantly improve the responsiveness of plots and display of statistics.

[0144] For example, consider a 20-color experiment with 1 million events. Whenever the system receives or performs a fine adjustment to the spillover values, the normal correction algorithm requires a total of 400 million multiplications plus 399 million additions. In contrast, the fast correction algorithm of the present system requires a total of only 60 million multiplications plus 20 million additions. The total multiplication and addition savings are 566% and 1895%, respectively, compared to the normal correction algorithm.

[0145] Next, the derivation of the present fast correction algorithm is discussed.

[0146] Let us assume that the matrix [C] is a correction matrix. The correction matrix [C] is expressed by the matrix equation [C] = [S] -1 is the inverse of the spillover matrix [S] by. When the correction matrix [C] and the inverse of the spillover matrix [S] are multiplied together, an identity matrix is ​​obtained as in the matrix equation [C][S]=[I]. The corrected event vector {V} can be calculated by multiplying the correction matrix [C], which is represented by the matrix equation {V}=[C][U}, and the uncorrected measured event vector {U}.

[0147] Due to the fine-tuning, the system generates or calculates an adjusted spillover matrix [S]'. The value of one element in the initial spillover matrix [S] is changed, e.g., S i,j '=>S i,j +δ i、jAssuming that the initial spillover matrix [S]' is adjusted to 0, then this adjusted spillover matrix [S]' can be expressed as the sum of the initial spillover matrix [S] plus the adjustments in the delta matrix [D], using the matrix equation [S]'=[S]+[D], where:

number

[0148] [S]+[D]=[S]([I]+[C][D]), [[S]+[D]] -1 =([I]+[C][D]) -1 [C], the recorrected event vector {V R The equation for} can be rewritten as follows:

number

[0149] The recorrection matrix can be simplified as follows:

number

number

[0150] Thus, the use of fast correction algorithms allows cell samples to be analyzed more quickly using a flow cytometer, and results (e.g., cell counts, concentrations, titrations) to be obtained more efficiently. Instead of a researcher or laboratory technician spending a day or more waiting for calculations to be completed to obtain test results for a biological sample (e.g., blood or urine), the test results for the biological sample can be obtained within a few hours by using fast correction algorithms.

[0151] conclusion Although the embodiments have been described in detail, the embodiments should not be construed as being limited by such embodiments, but should be construed according to the scope of the appended claims.

[0152] Although certain specific exemplary embodiments have been described and illustrated in the accompanying drawings, it will be understood that such embodiments are merely illustrative of the broad invention and are not intended to limit the broad invention, and that the embodiments are not limited to the specific constructions and configurations shown and described, as various other modifications may occur to those skilled in the art.

[0153] Some specific functions of a flow cytometer, such as analyzing the electrical signals detected by the detector to count different particles in the sample fluid, can be implemented in software and executed by a computer or processor. The software programs or code segments are used to perform the necessary tasks to perform those functions. These programs or code segments can be stored in a processor-readable medium or transmitted by a computer data signal embodied in a carrier wave over a transmission medium or communication link. The processor-readable medium can include any storage medium capable of storing information. Examples of the processor-readable medium are electronic circuits, semiconductor memory devices, read only memory (ROM), flash memory, erasable programmable read only memory (EPROM), floppy diskettes, CD-ROMs, optical disks, and hard disks. The code segments can be downloaded to the storage medium via a computer network such as the Internet, an intranet, etc.

[0154] Although this specification contains many specifics, these should not be construed as limitations on the scope of the disclosure or on the scope of what may be claimed, but rather as descriptions of features specific to certain embodiments of the disclosure. Some specific features described herein in the context of separate embodiments can also be implemented in combination in a single embodiment. Conversely, various features described in the context of a single embodiment can also be implemented separately or in combination in multiple embodiments. Furthermore, although features may be described above as operating in some specific combinations and may even be initially claimed as such, one or more features from a claimed combination may in some cases be deleted from the combination, and the claimed combination may be directed to a partial combination or a variation of the partial combination. Accordingly, the claimed invention is limited only by the scope of the appended claims.

Claims

1. 1. A method of performing flow cytometry using a conventional flow cytometer, comprising: generating an initial spillover matrix by using multiple single-stain compensation controls; passing a sample through said flow cytometer; generating a measured sample event vector by measuring the fluorescence of a plurality of cells passing through the flow cytometer; generating a corrected sample event vector by using the initial spillover matrix and the measured sample event vector; generating an adjusted spillover matrix by fine-tuning the initial spillover matrix; The method includes:

2. calculating a recorrected event vector by using the adjusted spillover matrix and the measured sample event vector; The method of claim 1 further comprising:

3. The method of claim 1 , wherein the corrected sample event vector is equal to the measured sample event vector linearly multiplied by the inverse of the initial spillover matrix.

4. The method of claim 1 , wherein generating an adjusted spillover matrix comprises adding a delta matrix to the initial spillover matrix.

5. 5. The method of claim 4, wherein the delta matrix includes one or more delta values ​​that fine-tune the initial spillover matrix, the delta matrix having the same dimensions as the initial spillover matrix.

6. 2. The method of claim 1, wherein the initial spillover matrix has dimensions NxN, where N is the number of the single-stain compensation controls.

7. 2. The method of claim 1, wherein each of the single-stain compensation controls comprises one of fluorescein isothiocyanate (FITC), R-phycoerythrin (PE), peridinin chlorophyll protein complex (PerCP), PE-Cy7, allophycocyanin (APC), or APC-Cy7.

8. The method of claim 6 , wherein the measured sample event vector includes N values ​​and the corrected sample event vector has N values.

9. 3. The method of claim 2, wherein the recorrected event vector is equal to the measured sample event vector multiplied by the inverse of the sum of the initial spillover matrix and a delta matrix.

10. 1. A method of performing flow cytometry using a spectral flow cytometer, comprising: generating a reference matrix by using a plurality of single-stained compensation controls; passing a sample through said flow cytometer; generating a measured sample event vector by measuring the fluorescence of a plurality of cells passing through the flow cytometer; generating an unmixed sample event vector by using the reference matrix and the measured sample event vector; generating an adjusted spectral spillover matrix by fine-tuning the spectral spillover matrix; The method includes:

11. calculating a recorrected event vector by using the adjusted spectral spillover matrix and the measured sample event vector; The method of claim 10 further comprising:

12. The method of claim 10 , wherein the unmixed sample event vector is equal to the measured sample event vector linearly multiplied by the reference matrix.

13. The method of claim 12 , wherein the number of variables in the unmixed sample event vector is less than the number of variables in the measured sample event vector.

14. solving for the unmixed sample event vector using a least squares algorithm on the measured sample event vector and the reference matrix; The method of claim 13 further comprising:

15. The method of claim 10 , wherein the initial spectral spillover matrix of the unmixed sample event vectors is an identity matrix.

16. The method of claim 15 , wherein generating an adjusted spillover matrix comprises adding a delta matrix to the initial spectral spillover matrix.

17. 17. The method of claim 16, wherein the delta matrix includes one or more delta values ​​that fine-tune the initial spectral spillover matrix, the delta matrix having the same dimensions as the initial spectral spillover matrix.

18. 16. The method of claim 15, wherein the initial spectral spillover matrix has dimensions NxN, where N is the number of the single-stain compensation controls.

19. The method of claim 10 , wherein the measured sample event vector includes N values ​​and the unmixed sample event vector has N values.

20. 16. The method of claim 15, wherein the recorrected sample event vector is equal to the measured sample event vector multiplied by the inverse of the sum of the initial spectral spillover matrix and a delta matrix.

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