Method and biosensor for label-free cell activation profiling using microfluidic impedance cytometry

Abstract

The present invention relates to a microfluidic impedance cytometer including an integrated biosensor and the use of the device for label-free activation profiling of cells in a sample, such as myeloid and / or lymphocytes and neutrophils. The microfluidic impedance cytometer may comprise a spiral shaped flow channel, coplanar electrodes for generating an electric field across the channel, a first sample inlet, and a second sheath fluid inlet, the sample inlet being a stepped sample inlet in that the height of the sample inlet is less than the height of the sheath fluid inlet and the flow channel.

Description

[Technical Field]

[0001] The present invention generally relates to the field of cell sorting and profiling using label-free microfluidic methods based on impedance cytometry. Specifically, the present invention relates to biophysical profiling of neutrophils and lymphocytes in a sample by using microfluidic impedance cytometry having an integrated biosensor specifically designed for the above purpose, as well as to the biosensor. [Background technology]

[0002] Lymphocytes are essential to our adaptive immune system. Upon recognition of specific antigens, they secrete cytokines, undergo proliferation and differentiation, and regulate the overall immune response. These cells are crucial not only in responses to pathogens such as HIV, but also in disorders such as cancer and cardiovascular disease. Therefore, qualitative and quantitative detection of lymphocyte phenotypes is clinically useful for reflecting the host immune state and for disease diagnosis. For example, lymphocyte count and neutrophil-to-lymphocyte ratio are often used as clinical parameters for disease prediction in bacteremia, malaria, and tuberculosis. Activated T lymphocytes, which are critical to cell-mediated immune responses, are also associated with susceptibility to HIV infection and survival in cancer patients.

[0003] Because T lymphocytes are highly diverse, it is important to perform single-cell analysis to identify various lymphocyte subpopulations. Conventional techniques for studying lymphocyte activation include measuring surface marker expression (CD69, CD137, etc.) by flow cytometry (Non-Patent Document 1; Non-Patent Document 2) or measuring secreted cytokines (IFNγ, IL2, etc.) using enzyme-linked immunosorbent assay (ELISA) (Non-Patent Document 3; Non-Patent Document 4). However, these approaches are expensive and cumbersome due to the use of antibodies and are not always scalable for clinical trials.

[0004] Electrical-based microtechnologies (such as dielectrophoresis (DEP), electrorotation, and impedance cytometry) are attractive alternatives that facilitate label-free single-cell measurements with minimal sample requirements and cell perturbation (Non-Patent Documents 5; 6; 7; and 8). Recent studies include lymphocyte profiling using electrorotation (Non-Patent Documents 9 and 10) and mouse lymphocyte activation and lymphocyte differential counting using impedance cytometry (Non-Patent Documents 11, 12, 13, and 14). However, these approaches suffer from low throughput (sample flow rate <20 μL / min) and require additional blood sample preparation steps. While impedance cytometry can improve throughput (Non-Patent Document 15), cells may undergo inertial focusing to have an equilibrium position >1 in the vertical direction (channel height direction). This leads to multiple cell clusters in the electrical profile, which negatively impacts phenotypic determination accuracy (Non-Patent Documents 16, 17, and 18). Several solutions have been proposed, including DEP focusing (Non-Patent Literature 19) and position sensing using multiple electrodes (Spencer et al., cited above; De Ninno et al., cited above). However, these require additional equipment (amplifiers), making the setup complex, and the increased detection area (due to the additional electrodes) may increase the likelihood of simultaneous particle occurrences (the presence of two or more particles in the detection area).

[0005] Neutrophils are the most abundant and important effector cells in innate immunity. Conventional techniques for studying neutrophil activation include measuring surface marker expression by flow cytometry or measuring secreted cytokines / enzymes using enzyme-linked immunosorbent assays (ELISA). However, these approaches are expensive and cumbersome due to the use of antibodies and are not always scalable for clinical trials. Electro-based microtechnologies (such as dielectrophoresis (DEP), electrorotation, and impedance cytometry) are attractive alternatives that facilitate label-free single-cell measurements with minimal sample requirements and cell perturbation (Non-Patent Literature 5 (doi:10.1007 / s10404-010-0580-9 (2010)); Non-Patent Literature 6; Non-Patent Literature 8 (doi:10.1039 / D0LC00840K (2021))). Recent research includes leukocyte profiling using electrorotation (Non-patent Literature 9 (doi:10.1002 / 1522-2683(200207)23:13<2057::AID-ELPS2057>3.0.CO;2-X(2002)); Non-patent Literature 10 (doi:https: / / doi.org / 10.1016 / S0005-2736(98)00253-3(1999)). Examples include )) and differential impedance cytometers (Non-Patent Literature 12 (doi:10.1126 / scitranslmed.3006870 (2013)); Non-Patent Literature 13 (doi:https: / / doi.org / 10.1016 / j.bios.2017.02.047 (2017)); Non-Patent Literature 14 (doi:10.1021 / acsnano.0c03018 (2020))). However, these approaches suffer from low throughput (sample flow rate <20 μL / min) and require additional blood sample preparation steps.

[0006] To mitigate the main obstacles, impedance cytometry has emerged as a label-free single-cell phenotyping technique based on cell-specific electrical properties (Non-Patent Document 7). While this technique is widely used for determining cell types (Non-Patent Document 11 (doi:10.1038 / nprot.2016.038); Non-Patent Document 20; Non-Patent Document 21), the clinical utility of microfabricated impedance cytometers is limited because they require a conventional sample preparation step of purifying target cells before analysis.

[0007] While impedance cytometry can improve throughput, cells may undergo inertial focusing to have an equilibrium position >1 in the vertical direction (channel height direction). This can lead to multiple cell clusters in the electrical profile, negatively impacting phenotypic determination accuracy (Non-Patent Literature 16; Non-Patent Literature 17; Non-Patent Literature 18 (doi:https: / / doi.org / 10.1016 / j.snb.2017.10.113(2018))).

[0008] Several solutions have been proposed, including DEP focusing (Non-Patent Literature 19 (doi:10.1021 / acssensors.6b00286(2016))) and position sensing using multiple electrodes (Spencer et al., cited above; de ​​Ninno et al., cited above). However, these solutions are complex to set up due to the need for additional equipment (amplifiers), and the detection area becomes larger (due to the additional electrodes), which may increase the occurrence of simultaneous particle occurrences (the presence of two or more particles in the detection area).

[0009] An integrated spiral microfluidic device with on-chip impedance detection for leukocyte phenotyping (Dean Flow Fractionation, DFF) has been previously developed (Non-Patent Literature 15 (doi:10.1039 / C9LC00250B(2019))). However, it has been found that even with this device, the use of inertia-based sorting results in particle focusing at two different vertical positions (positions along the channel height) during flow. This negatively impacts the accuracy of impedance detection because the signal magnitude differs depending on the vertical position, even for the same particle / cell. [Prior art documents] [Non-patent literature]

[0010] [Non-Patent Document 1] Wolfl et al., Blood, 2007, 110, 201-210. [Non-Patent Document 2] Cibrian and Sanchez-Madrid, European Journal of Immunology, 2017, 47, 946-953. [Non-Patent Document 3] Versteegen et al., J Immunol Methods, 1988, 111, 25-29. [Non-Patent Document 4] Igietseme and Herrscowitz, Journal of Immunological Methods, 1987, 97, 123-131. [Non-Patent Document 5] Sun & Morgan, Microfluidics and Nanofluidics, 2010, 8, 423-443. [Non-Patent Document 6] Chen et al., International Journal of Molecular Sciences, 2015, 16, 9804. Non-Patent Document 7 Petchakup et al., Micromachines, 2017, 8, 87 Non-Patent Document 8 Honrado et al., Lab on a Chip, 2021, 21, 22-54 Non-Patent Document 9 Pethig et al., Electrophoresis, 2002, 23, 2057-2063 Non-Patent Document 10 Huang et al., Biochimica et Biophysica Acta (BBA)-Biomembranes, 1999, 1417, 51-62 Non-Patent Document 11 Hassan et al., Nat. Protocols, 2016, 11, 714-726 Non-Patent Document 12 Watkins et al., Science Translational Medicine, 2013, 5, 214ra170-214ra170 Non-Patent Document 13 Rollo et al., Biosensors and Bioelectronics, 2017, 94, 193-199 Non-Patent Document 14 Han et al., ACS Nano, 2020, 14, 8646-8657 Non-Patent Document 15 Petchakup et al., Lab on a Chip, 2019, 19, 1736-1746 Non-Patent Document 16 Spencer et al., Lab on a Chip, 2016, DOI:10.1039 / C6LC00339G Non-Patent Document 17 De Ninno et al., Lab on a Chip, 2017, DOI:10.1039 / C6LC01516F Non-Patent Document 18 Caselli et al., Sensors and Actuators B: Chemical, 2018, 256, 580-589. [Non-Patent Document 19] Haandbaek et al., ACS Sensors, 2016, 1, 1020-1027. [Non-Patent Document 20] Simon et al., Lab on a Chip, 16, 2326-2338, doi:10.1039 / C6LC00128A (2016) [Non-Patent Document 21] Song et al., Lab on a Chip, 13, 2300-2310, doi:10.1039 / C3LC41321G (2013) [Overview of the project] [Problems that the invention aims to solve]

[0011] Therefore, further improved methods for cell profiling, particularly for lymphocytes and neutrophils, are still needed in this field. [Means for solving the problem]

[0012] This invention satisfies this need by providing a method for label-free profiling of blood cell (lymphocytes or neutrophils, etc.) activation based on (microfluidic) impedance cytometry.

[0013] In a first embodiment, the present invention relates to a method for label-free cell profiling using impedance cytometry, wherein the method is: (i) A step of providing a fluid sample containing cells; (ii) Perfusing the sample into an impedance cytometer having a first sample inlet and a second sheath fluid inlet and a flow channel, wherein the flow channel comprises coplanar electrodes for generating an electric field across the channel and means for detecting disruptions of the electric field by detecting changes in electrical impedance; (iii) the step of measuring the change in electrical impedance for label-free quantitative profiling of cells in the sample; Includes.

[0014] This method may be suitable for quantitative label-free cell profiling. In various embodiments of such a method, the cells are blood cells, preferably myelocytes and / or lymphocytes. In specific embodiments, the cells are lymphocytes or neutrophils. Thus, this method may be a method for activated lymphocyte profiling or activated neutrophil profiling.

[0015] In various embodiments, impedance measurements are performed at low frequencies in the range of 0.1–1 MHz, preferably 0.2–0.6 MHz, to determine lymphocyte cell size. In various embodiments, the measurements may be supplemented by further performing impedance measurements at high frequencies in the range of >1–20 MHz, preferably 1.2–15.0 or 1.2–12.0 MHz, to determine electrical opacity. Electrical opacity can be calculated as the ratio of the magnitude of the impedance signal at low frequencies (0.1–1 MHz) to the magnitude of the impedance signal at high frequencies (>1–20 MHz), and reflects cell membrane capacitance or cell nuclear characteristics. In various embodiments, measurements may be performed in two higher frequency ranges, one for determining membrane characteristics (>1–10 MHz) and the other for determining nuclear characteristics (>10–20 MHz).

[0016] In various embodiments, the method utilizes microfluidic impedance cytometry. In such embodiments, step (i) may further include providing an impedance cytometer in the form of a microfluidic device comprising a spiral-shaped flow channel having an inner and outer wall with respect to the center of the spiral, and having two inlets and at least two outlets, one of the two inlets being a sample inlet located on the inner wall of the spiral-shaped flow channel, the other of the two inlets being a sheath inlet located on the outer wall of the spiral-shaped flow channel, and at least one of the outlets comprising an impedance detector including a coplanar electrode for direct impedance quantification of lymphocytes. In such embodiments, step (ii) may include introducing the sample from (i) into the sample inlet port at a sample flow rate and introducing the sheath fluid into the sheath inlet at a sheath flow rate, wherein the sample-to-sheath flow rate ratio is 1:15 or less, preferably 1:20 or less, and driving the sample and sheath fluid through the spiral-shaped flow channel for cell sorting and focusing. Step (iii) may then include quantitatively profiling the cells, preferably activated lymphocytes, by directing the selected and focused cells in step (ii) towards an impedance detector and determining the impedance signal readout, by driving the selected and focused lymphocytes into at least one outlet equipped with an impedance detector.

[0017] In these methods, the sample flow rate may be in the range of 10 to 400 μL / min. In various embodiments, the sample-to-sheath flow rate ratio is 1:20 to 1:50. The sample inlet may be a stepped sample inlet, where the height of the sample inlet is preferably at least 2 times, more preferably 3 to 10 times, lower than the height of the sheath fluid inlet and flow channel. This setting is also referred to herein as a “stepped sample inlet.” The sample inlet may be located at the bottom of the flow channel to allow for better cell focusing.

[0018] In various embodiments, the sample inlet has a height of 15 to 50 μm, and / or the sheath fluid inlet and flow channel have a height of 80 to 250 μm. In various embodiments of the method, the opening width of the sheath fluid inlet is in the range of 200 to 700 μm, and / or the opening width of the sample inlet is 50 to 200 μm, preferably 60 to 180, 70 to 160 μm, or 80 to 150 μm. In such embodiments, the aspect ratio (width to height) of the sheath fluid inlet to the sample inlet is in the range of 3 to 10, preferably 4 to 8. In various embodiments, the area ratio of the sheath inlet to the sample inlet is in the range of 10 to 50, preferably 20 to 40.

[0019] The method of the present invention may further include a step of performing magnetically activated cell sorting (MACS) prior to impedance profiling in steps (ii) and (iii), by incubating the sample with magnetic beads that bind to undesirable cells, and separating the magnetic beads to which the undesirable cells are bound from the unbound cells by a magnetic field. Such undesirable cells may include myelocytes such as monocytes, especially when the method is for lymphocyte profiling.

[0020] In another embodiment, the present invention also relates to a biosensor for quantitative label-free cell profiling using impedance cytometry, wherein the biosensor is: (i) A microfluidic device comprising a spiral-shaped flow channel having an inner wall and an outer wall with respect to the center of the spiral, and having two inlets and at least two outlets, wherein one of the two inlets is a sample inlet and is located on the inner wall of the spiral-shaped flow channel, and the other of the two inlets is a sheath inlet and is located on the outer wall of the spiral-shaped flow channel; (ii) A coplanar electrode for direct impedance quantification of cells, the coplanar electrode being integrated with at least one of the outlets of the microfluidic device; It is equipped with.

[0021] All embodiments described above in the context of the method, relating to the microfluidic device and the impedance detector itself, are similarly applicable to the biosensor of the present invention. In yet another embodiment, the present invention also comprises a biosensor of the present invention, further comprising (1) a microchip, (2) an impedance analyzer, and (3) a fluid pump connected to a sample and sheath fluid inlet to enable perfusion control, wherein the microchip is connected to the impedance analyzer and the fluid pump, and the impedance analyzer is connected to a coplanar electrode to enable impedance measurement, characterized in that it is an integrated impedance cytometer. [Brief explanation of the drawing]

[0022] [Figure 1] A schematic diagram of a cell sorting and detection device with improved single-stream particle focusing, utilizing a sample inlet with a stepped height, is shown. The 3D shape of the stepped height inlet is highlighted by a dotted box. [Figure 2] This is a CFD simulation showing 3D sample focusing in the inlet region (cross-sectional view) of a spiral device. Black and gray represent sample streamlines and sheath streamlines, respectively. The particles are equilibrated at the bottom of the channel by inertial focusing and secondary Dean vortices, and remain focused until they reach the detection region (close to the electrodes). [Figure 3]This illustrates the characterization of particle focusing at an inlet with a stepped design. The left image shows a cross-sectional view of a CFD simulation with varying sample-to-sheath flow rates, while the right image shows impedance detection of a 10 μm bead indicating the equilibrium position of a single particle. [Figure 4] This shows the effect of changing the height of the sample inlet, which has a step, on the sample streamlines (black dots). The sheath inlet opening is fixed at 100 μm for all conditions. [Figure 5] This shows the effect of the sample inlet opening width on sample streamlines. The sample inlet height is fixed at 20 μm for all conditions. The gray and black dots represent the streamlines of the sheath and sample, respectively. [Figure 6] This shows the quantification of the percentage of sample streamlines in the lower half of the channel, indicating the efficiency of sample flow confinement. [Figure 7] Numerical impedance simulations are shown to investigate the effect of membrane conductivity on the impedance spectrum (top), differential current response (middle), and opacity (bottom) using the double-shell model described in the section on impedance modeling. Table 1 contains the electrical properties of the cellular components. [Figure 8] Numerical impedance simulations illustrating the effects of varying nuclear size (top) and nuclear membrane conductivity (bottom) are shown. The vertical dotted lines indicate the system's frequency selection. [Figure 9] This paper describes multi-parameter impedance phenotyping of neutrophils. Neutrophils with different phenotypes were tested in an integrated device. By using a multi-frequency measurement scheme (0.3 MHz, 1.72 MHz, and 12 MHz), different phenotypes can be distinguished based on impedance characteristics, namely: 1) cell size (impedance at 0.3 MHz), 2) membrane opacity reflecting plasma membrane characteristics (ratio of impedance at 1.72 MHz to 0.3 MHz), and 3) nuclear opacity representing nuclear characteristics (ratio of impedance at 12 MHz to 0.3 MHz). [Figure 10]This document describes neutrophil phenotyping using an integrated device. Neutrophil phenotypes are: 1) untreated control, 2) death (heat treatment at 75°C), 3) treatment with phorbol 12-myristate 13-acetate (PMA) for 2 hours (NETosis), and 4) spontaneous NETosis at pH 9. The intensity scale indicates the density of single-cell events. [Figure 11] Quantification of NETosis populations. (A) Impedance gating strategy using nuclear opacity for NETosis quantification (top), and the corresponding back-gated population for membrane opacity versus electrical size (bottom). (B) Cell size and (C) Nuclear opacity of healthy and NETosis populations. (D) Spontaneous NETosis quantification of neutrophils in different pH environments. [Figure 12] Impedance profiling of neutrophil bacterial responses to Escherichia coli (EC). Neutrophils were treated for 24 hours with different bacterial suspensions at 1) 10-fold dilutions from bacterial stock (EC-10×, OD625:0.29), 2) 103-fold dilutions (EC-1k×), and 3) 106-fold dilutions (EC-1M×). [Figure 13] This describes a workflow for lymphocyte activation profiling using integrated inertial impedance cytometry. Activated lymphocytes are larger in size and classified into the impedance detection region, while healthy lymphocytes (smaller in size) are classified as waste during the spiral phase. [Figure 14] Lymphocyte impedance profiling using integrated inertial impedance cytometry is shown. (A) Stacked high-speed images showing cell convergence at the sorting stage and the stage before impedance detection for dead lymphocyte populations, control lymphocyte populations, and activated lymphocyte populations. (B) Impedance counts of activated lymphocytes under each condition for 3 minutes. [Figure 15] The scatter plots of the forward and side scattering signals for the pre-sorted (sample) fraction of the lymphocyte sample, the sorted fraction (from the impedance detector outlet), and the discarded fraction (from the spiral sorter's waste outlet) are shown. [Figure 16]The CD137 expression profiles of the lymphocyte activation marker from the original lymphocyte sample (inlet) and selected lymphocytes from the impedance outlet are shown. [Figure 17] The scatter plots show the percentage of activated cells in the pre-sorted fraction versus the pre-sorted fraction. The dotted line represents 5% activated cells in the pre-sorted cells (left). The calibration curve of impedance counts against the percentage of activated cells in the pre-sorted (sample) from two different donors (right). [Figure 18] This document describes the workflow for antigen-specific T lymphocyte response testing using impedance cytometry, including 1) isolation and culture of PBMCs with antigen, 2) MACS isolation of T lymphocytes (CD4 negative selection), and 3) microfluidic impedance profiling. [Figure 19] The impedance profiles of unstimulated control lymphocytes and stimulated lymphocytes (phytohemagglutinin (PHA) and purified tuberculin protein derivatives (PPD)) are shown. n indicates the number of impedance cells detected (normalized to 100k cells as the input sample). [Figure 20] The left image shows the quantification of activated lymphocytes (CD69+CD137+) before and after selection from three different donors. The right image shows representative flow cytometry analysis of pre-selection (inlet) and post-selection (impedance outlet) cells of PPD-treated PBMCs. [Figure 21] Flow cytometry analysis (left) shows the ratio of CD69+CD137+ lymphocytes in stimulated samples (PPD, PHA) normalized to control (unstimulated) in the impedance-sorted fraction (left). Microfluidic differential impedance count ratio (detected cells in stimulated samples (PHA, PPD) normalized to control (unstimulated)) in selected fractions from four different donors (right). [Figure 22]Representative flow cytometry analyses of lymphocyte populations (control, PHA, and PPD treated) in the sample (before sorting), as well as larger (activated) lymphocytes (after sorting) and inactivated lymphocytes (waste) after microfluidic sorting, are shown. [Figure 23] This shows impedance activation profiling using peripheral blood mononuclear cells (PBMCs) stimulated with PHA. (A) Stacked high-speed images showing the experimental workflow and (B) sorting of unstimulated control and PHA-stimulated PBMCs. Scale bar is 150 μm. (C) 3-minute impedance profiles of lymphocytes sorted for control, PHA, and PPD. [Figure 24] This diagram illustrates an impedance system prototyping. It shows a microfluidic cytometer setup including a fluid pump, microfluidic devices, and a PCB chip holder connected to an impedance analyzer. The inset shows a close-up of the PCB and chip holder. [Modes for carrying out the invention]

[0023] The present invention will be better understood by referring to the detailed description when considered in conjunction with non-limiting embodiments and the accompanying drawings. The following detailed description refers, for illustrative purposes, to specific details and embodiments in which the present invention may be carried out. These embodiments are described in sufficient detail to enable those skilled in the art to carry out the present invention. Other embodiments may be utilized, and structural and logical modifications may be made without departing from the scope of the invention. Since several embodiments can be combined with one or more other embodiments to form new embodiments, the various embodiments are not necessarily mutually exclusive.

[0024] Unless otherwise defined, all technical and scientific terms used herein have the same meaning as those commonly understood by those skilled in the art. The singular forms “a,” “an,” and “the” include plural referents unless the context explicitly indicates otherwise. Similarly, the word “or” is intended to include “and” unless the context explicitly indicates otherwise. The term “comprises” means “includes.” In case of any conflict, this specification, including the definitions of terms, shall prevail.

[0025] The inventors have developed a label-free, high-throughput microfluidic lymphocyte phenotyping assay based on impedance signatures. This development was made possible by the initial characterization of different T lymphocyte populations (CD4+), namely, the observation that they show significant biophysical differences in the impedance profiles of healthy lymphocytes, lymphocytes activated (by anti-CD3 / CD28 Dynabeads®), and dead lymphocytes (treated with 10% DMSO). Peripheral blood mononuclear cells (PBMCs) stimulation using mitogenectin phytohemagglutinin (PHA) and purified tuberculin protein derivative (PPD) antigens also revealed an increase in lymphocyte cell size (>8 μm) corresponding to lymphocyte activation (CD69+CD137+). Next, as a proof of concept for clinical use, the inventors applied an integrated cell sorting and impedance profiling platform for label-free quantitative measurement of activated T lymphocytes (>8 μm), hydrodynamically achieving a single-cell focused flow close to the coplanar electrode using a stepped-height sample inlet. This was found to offer several advantages for single-cell impedance measurement: 1. a well-aligned particle focused flow that minimizes cell loss during the initial sorting stage, 2. improved signal detection sensitivity due to the proximity of cells to the electrode, and 3. smaller impedance signal variability due to the more confined particle position / height in the lower half of the channel. These advantages enabled significant enrichment of larger activated T lymphocytes (approximately 11.7-fold) and electrical detection of low levels of lymphocyte activation (<5%). Subsequently, the integrated microchip was coupled with magnetically activated cell sorting (MACS) for impedance profiling of activated T lymphocytes (CD4+) from PBMCs stimulated with PHA and PPD. To identify activated T lymphocytes, we defined differential impedance cell number ratio (stimulated / unstimulated), which showed better sensitivity compared to immunophenotyping by flow cytometry.

[0026] The integrated inertial impedance cytometry platform described herein offers several key advantages, including 1) size-based activated lymphocyte enrichment and 2) hydrodynamically-based single-stream focusing to improve impedance detection sensitivity. When combined with MACS, the platform successfully detected activated CD4 T lymphocyte responses from PHA and PPD-stimulated PBMCs. In summary, the results reported herein demonstrate the potential of the developed platform for label-free antigen-specific lymphocyte profiling. It is anticipated that the developed technology may be further used for the quantitative detection of antigen-specific T lymphocyte responses as a point-of-care test for infectious diseases (e.g., tuberculosis and COVID-19). Label-free cell analysis and sequential collection of eluted cells greatly facilitate downstream bioanalysis to elucidate novel relationships between biophysical cellular properties and antigen-specific responses.

[0027] The inventors further discovered that incorporating a sample inlet with stepped height for single-stream particle focusing enhances the integrated platform for spiral microfluidics and impedance cytometry. Using this novel feature, they were able to successfully demonstrate label-free characterization of different neutrophil phenotypes (healthy, dead, and NETosis) based on nuclear and membrane properties. The developed platform also significantly minimizes sample handling and analysis / testing time with minimal cell perturbation, representing a significant advance for point-of-care applications.

[0028] The method of the present invention enables quantitative label-free cell profiling using impedance cytometry, and involves the following steps: (i) A step of providing a fluid sample containing cells; (ii) Perfusing the sample into an impedance cytometer having a first sample inlet and a second sheath fluid inlet and a flow channel, wherein the flow channel comprises a coplanar electrode for generating an electric field across the channel and means for detecting the breakdown of the electric field by detecting a change in electrical impedance; (iii) the step of measuring the change in electrical impedance for label-free quantitative profiling of cells in the sample; Includes.

[0029] In various embodiments of the present invention, the method of the present invention does not include (a) a centrifugation step; and / or (b) a labeling step in which at least one type of blood cell is labeled with a marker molecule.

[0030] The cells being profiled are typically blood cells, usually leukocytes, i.e., myelocytes and / or lymphocytes, with lymphocytes and neutrophils being used as specific examples to demonstrate the feasibility of the methods described herein. Thus, the methods described and illustrated are primarily methods that enable activated lymphocyte or activated neutrophil profiling.

[0031] The term "blood cells," as used herein, means including platelets, red blood cells, and white blood cells. All white blood cells are nucleated and are distinguished from anucleated red blood cells and platelets. White blood cell types can be classified in a standard manner. The two broadest categories are further divided into five main types: neutrophils, eosinophils, basophils, monocytes, and lymphocytes. These types are distinguished by their physical and functional characteristics. Monocytes and neutrophils are phagocytic. Further subtype classifications are well known in the art.

[0032] The terms “lymphocyte” or “lymphocyte” as used interchangeably herein refer to any one of natural killer (NK) cells, T cells, B cells, or any combination thereof. Peripheral blood mononuclear (PBMC) cells, tumor-infiltrating lymphocytes (TIL) cells, and lymphokine-activated killer (LAK) cells are also considered lymphocytes. Non-limiting examples of lymphocytes include cytotoxic lymphocytes (CTLs, CD8). + or CD4 + ), NK cells (CD2 + ), and T helper cells (CD4 + ) are some examples.

[0033] Therefore, the sample used is typically a blood sample. This sample may be treated before use in the method of the present invention, for example, by various separation, purification, and preservation treatments (all of which are well known to those skilled in the art). For example, red blood cells may be lysed or otherwise separated from the rest of the blood sample to increase the concentration of white blood cells and reduce the red blood cell background.

[0034] In various embodiments, further separation steps are performed to separate subpopulations of leukocytes. For leukocytes that occur in smaller total numbers compared to the dominant neutrophil and lymphocyte populations (e.g., basophils, eosinophils, and monocytes), it may be beneficial to separate lymphocytes and / or neutrophils, as well as erythrocytes and platelets, to enable profiling of these cells. The same applies when profiling neutrophils or lymphocytes. In such embodiments, it is not essential, but still beneficial, to separate subpopulations that are not evaluated from those being profiled before performing the methods described herein. Means for separating different subpopulations of leukocytes from one another are well known in the art, and some are also described below herein.

[0035] The sample may be a finger puncture sample or a sample generated from a vein puncture. "Sample" or "patient sample," as used herein, generally includes biological samples such as tissues and body fluids. "Body fluids" may include, but are not limited to, blood, serum, plasma, saliva, cerebrospinal fluid, pleural fluid, tears, mammary gland fluid, lymph, sputum, urine, amniotic fluid, and semen. Samples may include body fluids that are "cell-free." "Cell-free fluids" contain less than about 1% (w / w) of whole-cellular material. Plasma or serum are examples of cell-free fluids. Samples may include specimens of natural or synthetic origin. "Plasma," as used herein, refers to the cell-free fluid found in blood. "Plasma" can be obtained from blood by removing whole-cellular material from blood by methods known in the art (e.g., centrifugation, filtration, etc.). As used herein, "peripheral blood plasma" refers to plasma obtained from a peripheral blood sample. As used herein, "serum" includes the plasma fraction obtained after plasma or blood has been coagulated and the coagulated fraction has been removed. Serum samples are also cell-free. On the other hand, as used herein, "whole blood" refers to blood itself collected from a subject such as a human, and includes unseparated blood cells. Furthermore, as a whole blood sample, a sample obtained by diluting whole blood with a suitable buffer and / or by adding additives such as anticoagulants or protease inhibitors may be used. A whole blood sample contains blood cells such as platelets, red blood cells, and white blood cells. Since the methods disclosed herein relate to profiling cells, particularly lymphocytes, sample types that typically contain cells are preferred in the sense of the present invention.

[0036] As used herein, a “finger prick sample” is a blood sample consisting of a single drop taken from a fingertip. Preferably, the sample is taken with a small fingertip lancet device that can painlessly puncture the skin and produce a fresh blood flow. Once the blood droplet has fallen onto the finger, a collector can be placed over the droplet to allow the blood to be absorbed by the collector. The volume of a finger prick sample may range from 10 to 500 μL, preferably 50 to 300 μL, and more preferably 80 to 150 μL. Alternatively, the blood droplet may also be taken from another part of the body, such as an earlobe. The characteristics described for a finger prick sample also apply to such blood droplet samples taken from other areas of the body.

[0037] In various embodiments, the method utilizes microfluidic impedance cytometry. In such embodiments, step (i) may further include providing an impedance cytometer in the form of a microfluidic device.

[0038] The microfluidic device comprises a spiral-shaped flow channel having an inner wall and an outer wall relative to the center of the spiral, and having two inlets and at least two outlets. Typically, one of the two inlets is a sample inlet and the other is a sheath inlet. In some embodiments, the sample inlet is located on the inner wall of the spiral-shaped flow channel, and the sheath fluid inlet is located on the outer wall of the spiral-shaped flow channel. The device may have three or more inlets, but in a typical setting, this is not necessary and therefore undesirable. The at least two outlet ports may be at least three or at least four outlet ports, and may be two, three, or four outlets. In one embodiment, the device has three outlet ports.

[0039] The spiral-shaped flow channel may have at least a first end and a second end, the flow channel having two inlet ports at or near the first end, and at least two outlet ports at or near the second end, one of the two inlet ports located on the inner wall of the spiral-shaped flow channel, and the other inlet port located on the outer wall of the spiral-shaped flow channel. As schematically shown in Figures 3, 6 and 10, the first end may be the inner end of the spiral, and the second end may be the outer end.

[0040] In these configurations of microfluidic devices, an impedance detector, including a coplanar electrode required for impedance measurements, particularly for direct impedance quantification of lymphocytes, is located at at least one of the outlets, typically the outlet where the cells to be profiled are sorted by the microfluidic device.

[0041] In our previous research (International Publication No. 2017 / 003380A1), which is incorporated in its entirety herein by reference, particularly in the section relating to microfluidic devices, we found that Dean flow fractionation (DFF) is suitable for separating blood components and / or microparticles (such as bacteria) present in blood, and described how the parameter settings of a microfluidic device need to be adapted for the efficient and distinct purification of different cells and particles. However, this previously reported method focused on the purification of neutrophils. Our method uses DFF for cell sorting and focusing to enable the following impedance measurements:

[0042] Dean Flow Fractionation (DFF) is a microfluidic cell sorting technique known as Dean Flow Fractionation (DFF), which is well-known for separating cells based on their size (Hou et al., Sci. Rep. 3 (2013); Hou et al., Lab Chip, 15, 2297-2307 (2015)). DFF utilizes the dominant inertial force (F) at high flow velocities (Reynolds number, Re approximately 50-100). LThis is a sorting method based on inertial microfluidic technology, involving lateral movement of particles or cells along streamlines to focus at a more precise location due to ). In a DFF system, the fluid flowing through a spiral microchannel is subjected to radially outward centrifugal acceleration, resulting in the formation of two symmetrical, oppositely rotating Dean vortices in the upper and lower halves of the channel. The presence of Dean vortices within the spiral microchannel creates a lateral Dean traction force (F D ) is added, and both forces (F L and F D Because the ) changes nonlinearly depending on the particle size, and their superposition (FL / FD) determines the equilibrium position within the channel cross-section, excellent separation resolution is obtained.

[0043] As used herein, the term “flow channel” means a tubular passage for a liquid. Preferably, the flow channel is spiral in shape. Since the spiral shape represents an asymmetric shape, an inner wall and an outer wall can be defined. As used herein, the expression “microfluidic device” means a physical element that enables the control and manipulation of small, typically sub-millimeter-scale, geometrically constrained fluids. As used herein, a “microfluidic device” refers to a device comprising one or more microfluidic flow channels designed to transport, store, mix, react, and / or analyze liquid samples, typically in volumes of less than 1 milliliter. Typical examples of materials that may be used to fabricate microfluidic devices include, but are not limited to, silicone rubber, glass, plastics, silicon, and metals, and preferably, microfluidic devices are fabricated from polydimethylsiloxane (PDMS).

[0044] As used herein, the term “inlet port” is defined as the end of a flow channel that first comes into contact with the sample and sheath fluid. Conversely, the term “outlet port” refers to the end of a flow channel that releases the fluid material of the sample and the sheath fluid from the microfluidic device, for example, into the container to which it is connected.

[0045] As used herein, the term “connected to a container” refers to the transfer of the fluid recovered within the microfluidic device of the present invention to a second storage unit. “Container” in the sense of the present invention may be a unit whose sole function is to store the recovered fluid, but may also be a device that enables further purification or analysis of the recovered fluid, such as a column or filter unit. Devices used for the analysis of the recovered fluid may include, but are not limited to, FACS devices, mass spectrometers, microscopes, centrifuges, PCR thermocyclers, and a second flow channel containing a coating that may be used to perform functional assays.

[0046] The term “sheath fluid,” as used herein, refers to a variety of fluids, including aqueous or non-aqueous fluids and / or fluids that may contain additional material components, such as suspensions or emulsions of soluble chemical components or at least partially insoluble components. Preferably, the sheath fluid is a buffer, preferably a buffer compatible with blood cells, such as phosphate-buffered saline (PBS) and other well-known buffers. As used herein, the term “buffer” means an aqueous solution of any compound or combination of compounds that controls the pH of the environment in which they are dissolved or dispersed. With respect to pH value, buffers reduce the effect of acids or bases added to the buffer solution. Examples of water-soluble buffers include, but are not limited to, meglumine, sodium bicarbonate, sodium carbonate, sodium citrate, calcium gluconate, disodium hydrogen phosphate, dipotassium hydrogen phosphate, tripotassium phosphate, sodium tartrate, sodium acetate, calcium glycerophosphate, tromethamine, magnesium oxide, or any combination thereof. It is understood that metal cations may be substituted with other suitable metal cations without altering the buffering properties of the buffer. Buffers with low water solubility can also typically be used in combination with other more soluble components. Examples, but not limited to, include magnesium hydroxide, aluminum hydroxide, sodium dihydroxyaluminum carbonate, calcium carbonate, aluminum phosphate, aluminum carbonate, dihydroxyaluminum aminoacetate, magnesium oxide, magnesium trisilicate, magnesium carbonate, and combinations thereof. Buffers can also be supplemented with adjuvants such as salts, surfactants, and BSA (bovine serum albumin). The concentrations that can be used for each of the buffers listed above are well known in the art.

[0047] In the above embodiment of the method using a microfluidic device, step (ii) typically includes introducing a sample into the sample inlet port at a sample flow rate and introducing a sheath fluid into the sheath inlet at a sheath flow rate. The flow rates of the sample and sheath fluid are parameters that affect the cell sorting characteristics of the method.

[0048] In various embodiments, the sample-to-sheath flow rate ratio is 1:15 or less, preferably 1:20 or less, in the range of 1:15 to 1:100, 1:15 to 1:80, 1:15 to 1:60, 1:15 to 1:50, 1:18 to 1:45, or 1:20 to 1:45, for example, 1:20, 1:25, 1:30, 1:35, 1:40, or 1:45. In these methods, the sample flow velocity may be in the range of 10 to 400 μL / min. When the sample-to-sheath flow rate ratio is low, the sample flow rate is also low, and to enable this low ratio, the sheath fluid flow rate must be significantly increased. Typical sample flow rates may be in the range of about 1:15 to about 1:250. Therefore, the sheath fluid flow rate may be in the range of, for example, about 200 to about 5000 μL / min. The flow rates used are typically limited by the dimensions of the flow channel and microfluidic device.

[0049] As used herein in relation to numerical values, "approximately" refers to a range of ±10% or ±5% of the referenced value. The sample can be introduced into the flow channel of a microfluidic device by a syringe or other means. Other means for introducing the sample into the flow channel of a microfluidic device include microdispensers, such as droplet dispensers (e.g., injection nozzles); and in-channel dispensers (e.g., metering dispensers). The means of introducing samples into the flow channels of microfluidic devices have been discussed in more detail by Nguyen et al. and Li et al. (Nguyen et al., "Fundamentals and Applications of Microfluidics," 2nd edition, ARTECH HOUSE, INC. (Norwood, Massachusetts: 2006), pp. 395-417; Li et al., "Injection Schemes for Microchip-based Analysis Systems," Lab on a Chip Technology, Vol. 1: "Fabrication and Microfluidics," edited by Herold et al., Caister Academic Press (Norfolk, UK: 2009), pp. 385-403). Nguyen et al. discuss microdispensers in detail. Li et al. discuss microdispensers in detail.

[0050] The sample inlet may be a stepped sample inlet. As used herein, “stepped sample inlet” means that the height of the sample inlet is preferably at least 2 times, more preferably 3 to 10 times, for example about 4 times, about 5 times, about 6 times, about 7 times, about 8 times, or about 9 times lower than the height of the flow channel. The sample inlet may be located at the bottom of the flow channel to allow for better focusing of cells. In various embodiments, the stepped sample inlet also means that its height is typically lower than the height of the sheath fluid inlet by the same coefficient as in the case of the flow channel. In various embodiments, the sheath fluid inlet and the flow channel have similar or identical heights, and the sample inlet is significantly lower to provide a “step”.

[0051] The inventors have found that a sample with a stepped inlet represents a suitable means for hydrodynamically achieving a single-cell focused flow near the coplanar electrode used for impedance measurements. This brings several advantages to single-cell impedance measurements: namely, 1. a well-aligned particle focused flow that minimizes cell loss during the initial sorting stage, 2. improved signal detection sensitivity due to the proximity of cells to the electrode, and 3. smaller impedance signal fluctuations due to the more confined particle position / height in the lower half of the channel.

[0052] In various embodiments, the sample inlet has a height of 15–50 μm, and / or the sheath fluid inlet and flow channel have a height of 80–250 μm. In one example, the sample fluid inlet has a height of approximately 20 μm, and the flow channel and optionally the sheath fluid inlet also have a height of approximately 110 μm.

[0053] In various embodiments of this method, the opening width of the sheath fluid inlet is in the range of 200 to 700 μm, and / or the opening width of the sample inlet is in the range of 50 to 200 μm, preferably 60 to 180, 70 to 160 μm, or 80 to 150 μm. In some embodiments, the aspect ratio (width to height) of the sheath fluid inlet and the sample inlet is in the range of 3 to 10, preferably 4 to 8. The aspect ratio of the sample inlet and / or the sheath fluid inlet is generally preferred to be greater than 1, typically greater than 1.5, or greater than 2. In various embodiments, the area ratio of the sheath inlet to the sample inlet is in the range of 10 to 50, preferably 20 to 40.

[0054] In various embodiments of the present invention, the at least two outlet ports are at least three or at least four outlet ports. In some embodiments described herein, the device comprises three outlet ports, one connected to an impedance detector and the other two being waste outlet ports. In various embodiments, the sample outlet port connected to the impedance measurement is a middle outlet port, and the flow characteristics are tuned so that target cells are sorted into this middle outlet port.

[0055] In various embodiments, the Reynolds number (Re) of a sample flowing through a spiral-shaped flow channel is 50 to 100. As used herein, the term "Reynolds number" means ρυL / μ, where ρ is the density of the liquid, υ is the velocity of the liquid, L is the characteristic length of the flow channel, and μ is the viscosity of the liquid.

[0056] The spiral-shaped flow channel may contain or consist of polydimethylsiloxane (PDMS). It may have a width of 200-700 μm, for example 300-650 or 400-600 μm, preferably about 500 μm. In various embodiments, it has a height of 110-130 μm and / or a total length of 4-20, 5-15 or 7-13 cm, preferably about 10 cm.

[0057] In various embodiments of the device, at least two outlet ports are two outlet ports. In a more preferred embodiment of the two-outlet port device defined above, the first outlet port covers 5 to 30%, preferably 20%, of the width of the flow channel, the second outlet port covers 70 to 95% of the width of the flow channel, the first outlet port is located on the inner wall, and the second outlet port is located on the outer wall. In various embodiments of the two-outlet port device defined above, the spiral-shaped flow channel (a) contains or consists of polydimethylsiloxane (PDMS); (b) has a total length of 5 to 10 cm, preferably 6.5 cm; (c) has a width of 200 to 400 μm, preferably 300 μm; and / or (d) has a height of 30 to 90 μm, preferably 60 μm. In such embodiments, the width of the first outlet port may be 1 to 50 μm, and the width of the second outlet port may be 51 to 300 μm, measured from the inner wall to the outer wall.

[0058] In other embodiments, the spiral-shaped flow channel (a) contains or consists of polydimethylsiloxane (PDMS), (b) has a width of 300 to 600 μm, preferably 500 μm, (c) has a height of 100 to 130 μm, preferably 110 μm, and / or (d) has a total length of 7 to 13 cm, preferably 10 cm. In such embodiments, the device may have three or four outlet ports, for example, three outlet ports. In various embodiments, the spiral-shaped flow channel may gradually widen to a width of 800 to 1200 μm at the branching of the outlet ports. In further various embodiments of the invention where three to four outlet ports are present, the width of the flow channel may be divided into three or four equal parts to provide three or four outlet ports having essentially the same width. The outlet ports are generally arranged horizontally, with one near the inner wall and another near the outer wall relative to the center of the spiral, and any further outlet ports located between the inner wall outlet port and the outer wall outlet port.

[0059] The dimensions of microfluidic devices and their flow channels, as well as the materials and methods for constructing microfluidic devices, have been described by Nguyen et al., Tabeling, Armani et al., Tsao et al., Carlen et al., Cheung et al., Sun et al., and Waddell (Nguyen et al., "Fundamentals and Applications of Microfluidics," 2nd edition, ARTECH HOUSE, INC. (Norwood, Massachusetts: 2006), pp. 55-115; Tabeling, "Introduction to Microfluidics," Oxford University Press). Published by Caister Academic Press (Oxford, UK: 2005), pp. 244-295; Armani et al., "Fabricating PDMS Microfluidic Channels Using a Vinyl Sign Plotter," Lab on a Chip Technology, Vol. 1: "Fabrication and Microfluidics," edited by Herold et al., Caister Academic Press (Norfolk, UK: 2009), pp. 9-15; Tsao et al., "Bonding Techniques for Thermoplastic Microfluidics," Lab on a Chip Technology, Vol. 1: "Fabrication and Microfluidics," edited by Herold et al., Caister Academic Press Press) (Norfolk, UK: 2009), pp. 45-63;Carlen et al., "Silion and Glass Micromachining," Lab on a Chip Technology, Vol. 1: "Fabrication and Microfluidics," edited by Herold et al., Caister Academic Press (Norfolk, UK: 2009), pp. 83-114; Cheung et al., "Microfluidics-based Lithography for Fabrication of Multi-Component Biocompatible Microstructures," Lab on a Chip Technology, Vol. 1: "Fabrication and Microfluidics," edited by Herold et al., Caister Academic Press Press) (Norfolk, UK: 2009), pp. 115-138; Sun et al., "Laminated Object Manufacturing (LOM) Technology-Based Multi-Channel Lab-on-a-Chip for Enzymatic and Chemical Analysis," Lab on a Chip Technology, Vol. 1: "Fabrication and Microfluidics," edited by Herold et al., Caister Academic Press (Norfolk, UK: 2009), pp. 161-172;Waddell, "Laser Micromachining," Lab on a Chip Technology, Vol. 1: Fabrication and Microfluidics, edited by Herold et al., Caister Academic Press (Norfolk, UK: 2009), pp. 173-184.

[0060] In various embodiments, the spiral-shaped flow channel gradually expands to a width of 800–1200 μm at the branching point of the outlet port, for example. In some embodiments, the expanded width is approximately 1000 μm. As used herein, the term “furcated” means divided into branches. As used herein, the term “branch” means a limited portion of a larger or more complex body, i.e., a smaller flow channel that emerges from or enters or re-enters the primary flow channel. As used herein, the term “primary flow channel” or “first flow channel” means a flow channel through which at least the majority of the sample flows.

[0061] In various embodiments of the present invention in which four outlet ports are used, the first outlet port covers 0 to 100 μm of the width of the flow channel defined from the inner wall to the outer wall, the second outlet port covers 101 to 250 μm of the same width, the third outlet port covers 251 to 650 μm of the same width, and the fourth outlet port covers 651 to 1000 μm of the same width. In an alternative embodiment of the present invention, the first outlet port covers 1 to 20% of the width of the flow channel, the second outlet port covers 10 to 20% of the width of the flow channel, the third outlet port covers 30 to 50% of the width of the flow channel, and the fourth outlet port covers 25 to 45% of the width of the flow channel, where the order of outlet ports 1 to 4 is determined from the inner wall to the outer wall.

[0062] Step (ii), in embodiments employing a microfluidic device, may include driving the sample and sheath fluid through a spiral-shaped flow channel for cell sorting and focusing.

[0063] The sample may be driven by capillary action. Alternatively, the sample may be driven by a pump, by electrical force, or by other means for driving the sample. The same applies to the sheath fluid, which is typically driven by active means such as a pump rather than by capillary action. Suitable pumps include, but are not limited to, rotary (centrifugal) pumps; peristaltic pumps; and ultrasonic pumps. Electrical forces include, but are not limited to, electrohydrodynamic forces; electric field driving forces, e.g., electrophoresis, electroosmosis; and surface tension driving forces, e.g., electrohumidification, electrohumidification on dielectric surfaces. The means for driving the sample have been discussed in more detail by Nguyen et al., Erickson et al., Grover et al., Hunt et al., and Bersano-Begey et al. (Nguyen et al., "Fundamentals and Applications of Microfluidics," 2nd edition, ARTECH HOUSE, INC. (Norwood, Massachusetts: 2006), pp. 255-309; Erickson et al., "Introduction to Electrokinetic Transport in Microfluidic Systems," Lab on a Chip Technology, Vol. 1; "Fabrication and Microfluidics," edited by Herold et al., Caister Academic Press) Press) (Norfolk, UK: 2009), pp. 231-248;Grover et al., "Monolithic Membrane Valves and Pumps," Lab on a Chip Technology, Vol. 1, "Fabrication and Microfluidics," edited by Herold et al., Caister Academic Press (Norfolk, UK: 2009), pp. 285-317; Hunt et al., "Integrated Circuit / Microfluidic Chips for Dielectric Manipulation," Lab on a Chip Technology, Vol. 2, "Biomolecular Separation and Analysis," edited by Herold et al., Caister Academic Press (Press) (Norfolk, UK: 2009), pp. 187-206; Bersano-Begey et al., "Braille microfluidics," Lab on a Chip Technology, Vol. 2: "Biomolecular Separation and Analysis," edited by Herold et al., Caister Academic Press (Norfolk, UK: 2009), pp. 269-285. Nguyen et al. discuss microvalves, micropumps, microflow sensors, microfilters, and microseparators in detail. The chapter by Herold et al. provides detailed explanations of valves, pumps, and separation.

[0064] Next, step (iii) includes driving the selected and focused lymphocytes into an impedance detector, i.e., at least one outlet equipped with a coplanar electrode used for impedance measurement, thereby directing the selected and focused cells from step (ii) towards the impedance detector and determining the impedance signal readout to quantitatively profile the cells, preferably activated lymphocytes.

[0065] In various embodiments, impedance measurements are performed at low frequencies in the range of 0.1–1 MHz, preferably 0.2–0.8 or 0.2–0.6 MHz, to determine cell size. In various embodiments, the measurements may be supplemented by further performing impedance measurements at high frequencies in the range of >1–20 MHz, preferably 1.2–15.0 or 1.2–12.0 MHz, to determine electrical opacity. Electrical opacity can be calculated as the ratio of the magnitude of the impedance signal at high frequencies (>1–20 MHz) to the magnitude of the impedance signal at low frequencies (0.1–1 MHz). Electrical opacity (1.72 MHz / 0.3 MHz) reflects the electrical characteristics of the cell membrane, and electrical opacity (12 MHz / 0.3 MHz) reflects the electrical characteristics of the nucleus. In various embodiments, measurements can be performed in two higher frequency ranges, one for determining membrane properties (>1–10 MHz, preferably 1.2–8 MHz or 1.3–5 MHz or 1.3–3 MHz) and the other for determining nuclear properties (>10–20 MHz, preferably 10.5–15 MHz or 10.5–13 MHz). Thus, the first frequency band of these higher frequency bands is also referred to as the one used to determine membrane opacity, and the second frequency band is used to determine nuclear opacity. The indicated higher frequency range, i.e., >10–20 MHz, can also be used to determine the properties of intracellular structures, for example, to evaluate phagocytosis. It has also been found that this high frequency range makes it possible to profile neutrophils and identify those undergoing NETosis, because this is accompanied by nuclear swelling. In various embodiments, the measurements in the low frequency range described above can be combined with one of two higher frequency ranges to enable the determination of film properties or nuclear properties, or with both higher frequency ranges to determine both.Therefore, in various embodiments, impedance measurements are performed at frequencies of 0.1 to 1 MHz, preferably 0.2 to 0.6 MHz, optionally at frequencies >1 to 10 MHz, preferably 1.3 to 5 MHz, and / or at frequencies >10 to 20 MHz, preferably 10.5 to 13 MHz.

[0066] After step (iii), the cells may be discarded or collected in containers connected to their respective exit ports. This allows for further use and / or analysis of the sorted cells.

[0067] The method of the present invention may further include a step of performing magnetically activated cell sorting (MACS) prior to impedance profiling in steps (ii) and (iii). This can be achieved, for example, by incubating a sample with magnetic beads that bind to undesirable cells and separating the magnetic beads to which the undesirable cells are bound from the unbound cells by a magnetic field. Such undesirable cells may include myelocytes such as monocytes, especially when the method is for lymphocyte profiling. To enable such cell sorting, the magnetic beads are typically coated with a compound or ligand that binds only to a specific type of cell. The compound may be an antibody or antibody-like molecule in various embodiments.

[0068] In various embodiments of the present invention, the methods of the present invention do not include a labeling step in which at least one type of blood cell is labeled with a marker molecule. As used herein, the term “labeling” refers to the attachment or incorporation of one or more detectable markers (or “labels”) to cells, proteins and / or peptides as used herein. As used herein interchangeably, the terms “marker molecule” or “detectable marker” refer to any compound comprising one or more suitable chemicals or enzymes that directly or indirectly produce a detectable compound or signal in a chemical, physical or enzymatic reaction. As used herein, this term should be understood to include both the label itself (i.e., the compound or portion bound to the protein and / or peptide) and the labeling reagent (i.e., the compound or portion before binding to the peptide or protein). The labels used herein may be bound to amino acid residues of proteins and / or peptides via covalent or non-covalent bonds. Typically, the bond is covalent. The labels can be selected from, among others, isotope labels, isobaric labels, enzyme labels, color labels, fluorescent labels, chromogenic labels, luminescent labels, radioactive labels, haptens, biotin, metal complexes, metals, and gold colloids, with isotonic and isobaric labels being particularly preferred. All of these types of labels are well established in the art.

[0069] The present invention also encompasses microfluidic devices disclosed herein. In one embodiment, the present invention also provides a biosensor for quantitative label-free cell profiling using impedance cytometry, wherein the biosensor is: (i) A microfluidic device comprising a spiral-shaped flow channel having an inner wall and an outer wall with respect to the center of the spiral, and having two inlets and at least two outlets, wherein one of the two inlets is a sample inlet and is located on the inner wall of the spiral-shaped flow channel, and the other of the two inlets is a sheath inlet and is located on the outer wall of the spiral-shaped flow channel; (ii) A coplanar electrode for direct impedance quantification of cells, the coplanar electrode being integrated with at least one of the outlets of the microfluidic device; It is equipped with.

[0070] When used in this context, the term "biosensor" refers to a microfluidic device described herein and also defined below, comprising means for performing impedance measurements of selected cells.

[0071] In the context of this method, and with respect to microfluidic devices and impedance detectors, including dimensions and components, all embodiments described above are also applicable to the biosensor of the present invention. Accordingly, in various embodiments, the biosensor includes a sample inlet which is a stepped sample inlet, wherein the height of the sample inlet is preferably at least twice, more preferably three to ten times, lower than the height of the sheath fluid inlet and flow channel, and the sample inlet is optionally located at the bottom of the flow channel.

[0072] In various embodiments of the biosensor, the sample inlet has a height of 15 to 50 μm, and the sheath fluid inlet and flow channel have a height of 80 to 250 μm. The opening width of the sheath fluid inlet may be in the range of 200 to 700 μm, and / or the opening width of the sample inlet may be 50 to 200 μm, preferably both dimensional ranges are satisfied. The aspect ratio (width to height) of the sheath fluid inlet and sample inlet is preferably in the range of 3 to 10, more preferably 4 to 8. In various embodiments, the area ratio of the sheath inlet to the sample inlet is in the range of 10 to 50, preferably 20 to 40. In various embodiments, the spiral-shaped flow channel has a total length of 4 to 20, 5 to 15, or 7 to 13 cm, preferably about 10 cm.

[0073] Such biosensors / microfluidic devices can preferably be used for sorting and profiling both neutrophils and lymphocytes in a setup where other cell populations are removed before the sample is introduced into the device's sample inlet.

[0074] Another embodiment is an integrated impedance cytometer comprising a biosensor of the present invention, further comprising one or more, preferably all, of (1) a microchip, (2) an impedance analyzer, and (3) a fluid pump connected to a sample and sheath fluid inlet to enable perfusion control, wherein the microchip is connected to the impedance analyzer and the fluid pump, and the impedance analyzer is connected to a coplanar electrode to enable impedance measurement. The integrated impedance cytometer may further include software for operating the cytometer.

[0075] The realization of an integrated platform featuring continuous, label-free cell separation and impedance profiling on a single chip is of great commercial interest, with a wide range of clinical applications including immunology research, cancer diagnostics, and point-of-care and immunohealth profiling. This technology provides single-step, label-free cell separation and impedance analysis with minimal human handling. For multi-parameter studies, the developed platform can also be integrated with other cell-based platforms. Using all the features of this integrated platform, it can be developed into products for the immunology research and point-of-care testing markets.

[0076] All embodiments disclosed above in relation to the methods of the present invention are similarly applicable to the biosensors and impedance cytometers of the present invention, and vice versa. The present invention is further illustrated by the following non-limiting examples. However, it should be understood that the present invention is not limited to the exemplary embodiments. [Examples]

[0077] Materials and Methods Device Design, Fabrication and Operation A schematic diagram of the device is shown in Figure 1A. Briefly, it consists of three stages: (1) a spiral cell sorter, (2) an inertial cell focuser, and (3) an impedance detector using coplanar electrodes. This is based on a previous integrated platform for leukocyte (approximately 8~10 μm) research, and details of the impedance cytometer are described in previous documents (Petchakup et al. (2019), supra; Petchakup et al. (2018), supra).

[0078] To further improve the detection sensitivity and consistency of impedance, the present inventors added a stepped sample inlet to hydrodynamically confine the sample flow to the lower half of the channel using sheath fluid (Figure 1B). Under the influence of inertial coupled Dean flow migration in the spiral microchannel, larger cells (hydrodynamic diameter (D h ) to particle size (A p )>0.07) equilibrate near the bottom of the channel (closer to the electrode surface) for neutrophil sorting and impedance detection (Figure 1B).

[0079] To realize the stepped sample inlet in the device, two-layer photolithography for SU-8 master fabrication was used, where the first layer (height of approximately 20 μm) defined the sample inlet and the detection region (width of 30 μm), and the second layer was aligned on top of the first layer to define the rest of the device (height of approximately 110 μm). The area ratio between the sheath inlet (opening width 500 μm) and the stepped-height sample inlet (opening width 100 μm) was approximately 27.5. For the integrated platform, the optimal flow rates of the sample and the sheath fluid were 80 μL / min and 1680 μL / min (1:21), respectively. Polydimethylsiloxane (PDMS) was cast onto the SU-8 master to fabricate the microfluidic device, which was bonded onto a glass substrate with patterned coplanar gold electrodes (width 20 μm, gap 20 μm, and thickness 200 nm).

[0080] Prior to any experiment, the device was coated with 1% bovine serum albumin (BSA, Biowest) in phosphate-buffered saline (PBS, Lonza) for at least one hour to prevent nonspecific adhesion to the device. The sample and sheath (both in 0.1% BSA in PBS) were loaded into the device and perfused using a syringe pump (PHD ULTRA®, Harvard Apparatus).

[0081] Before electrical measurements and high-speed imaging, the sample and sheath were perfused for 1 minute to allow the flow to stabilize. During the experiment, the platform was mounted on an inverted microscope (Nikon Eclipse Ti). High-speed images were acquired using a Phantom V9.1 (AMETEK).

[0082] Measurement setup and data presentation Differential measurement using three electrodes was used for single-cell impedance profiling. Using the central electrode, a 1V excitation signal was applied at three different frequencies (0.3 MHz, 1.72 MHz, and 12 MHz) by a lock-in amplifier (HF2LI, Zurich instrument). The current response from the side electrodes was converted to voltage by a transimpedance amplifier (DHPCA-100, FEMTO®) with a transimpedance gain of 10 kV / A, fed back to the lock-in amplifier, and the magnitude and phase at the applied frequencies were extracted at a sampling rate of 230 k samples / s. The time-series impedance response was recorded for 3 minutes and processed to extract amplitude peaks corresponding to single-particle events. By spiking a reference bead (10 μm bead), the electrical size of the cell could be determined as shown in previous literature (Caselli et al., cited above) (|Z LF |α Electrical Size 3Data calibration was performed to calculate the electrical size and normalized opacity (particle opacity normalized by the opacity of a reference bead, Petchakup et al. (2019), op. cit.). Finally, a 2D scattering plot of electrical size (μm) versus normalized opacity was plotted to show the impedance profile of the cell population. Data processing was performed using a custom-made program written in MATLAB® (MathWorks).

[0083] Sample preparation Blood samples from healthy donors were collected by venipuncture into sodium citrate vacutainer® (BD) for experimental purposes after informed consent. To isolate primary neutrophils, whole blood samples were stratified in a 1:1 ratio onto PolymorphPrep®. After density centrifugation according to the manufacturer's protocol, neutrophils were collected and washed with 0.5× PBS. Subsequently, RBC lysis was performed to further remove RBCs. RBC lysis buffer (eBioscience) was added to the neutrophil pellet (1 mL) for 3 minutes and quenched with 0.5% BSA in PBS. The neutrophils were washed and resuspended in 0.1% BSA in PBS.

[0084] To extract peripheral blood mononuclear cells (PBMCs), whole blood was subjected to Ficoll-Paque® Plus (GE Healthcare) density centrifugation according to the manufacturer's protocol. For experiments requiring purified lymphocytes, a human CD4 T cell negative isolation kit (STEMCELL) was used to remove other residual immune cell populations before or after culturing.

[0085] For comparison across different experiments and devices, 10 μm polystyrene beads (Polysciences) were used as a sample (10 4 ~10 5 Beads ( / mL) were spiked and used as a reference because they exhibited consistent dielectric properties.

[0086] Lymphocyte activation using CD3 / CD28 Dynabeads (registered trademark) Dynabeads® Human T-Activator CD3 / CD28 (Thermofisher) was used for lymphocyte activation. Briefly, lymphocytes were cultured in a well plate with anti-CD3 / CD28 beads at a 1:1 cell-to-bead ratio with 30 U / mL of rIL-2 and maintained at 37°C in a humidified CO2 incubator. After 24 hours, cells were collected and the activator was removed using a magnet. The cells were then resuspended in 0.1% BSA in PBS at the desired concentration. For the untreated control sample, lymphocytes were cultured without the activator and rIL-2.

[0087] PHA and PPD stimulation First, the isolated PBMCs are placed on a 24-well plate in a 2x10⁶ arrangement. 6 Cells were seeded at a concentration of cells / ml and incubated at 37°C for 24 hours in RPMI1640 medium (Hyclone) supplemented with 10% fetal bovine serum (FBS, Gibco) and 1% penicillin-streptomycin (P / S, Life Technologies). For activation experiments, lymphocytes were activated by adding PHA (final concentration 2.5 μg / ml, Sigma) or PPD (final concentration 10 μg / ml, Statens Serum Institute, Denmark) to the medium. After the desired time point, cells were harvested and resuspended in the recommended MACs buffer. Other immune cell populations were removed using CD4 lymphocyte MACs-negative isolation. Subsequently, cells were measured for impedance analysis at approximately 5–50 × 10⁶ cells. 4 The cells were resuspended in 0.1% BSA in PBS at a concentration of cells / mL.

[0088] neutrophil activation To study the effect of different pH levels on the biophysical properties of neutrophils, neutrophils were resuspended in RPMI at different pH levels and incubated at 4°C for 2 hours. For bacterial studies, neutrophils were incubated in an orbital shaker for 24 hours in penicillin-streptomycin-free RPMI at different bacterial concentrations.

[0089] Flow cytometry analysis To assess lymphocyte activation status, different cell fractions were stained with CD3-BUV737, CD4-BV650 (BD Biosciences, USA), CD137-PE, and CD69-FITC (Thermofisher, USA) antibodies on ice for 15 minutes. Cells were washed once to remove unbound antibodies and then analyzed using a BD LSR Fortessa® X-20 flow cytometer (BD Biosciences). Cell sorting was performed using a BD FACSFusion cell sorter (BD Biosciences). Flow cytometry data were analyzed and quantified using FlowJo® (BD Biosciences).

[0090] statistical analysis All statistical analyses were performed using GraphPad Prism® V8.0 (GraphPad software).

[0091] research approval Written informed consent was obtained from all participants during the recruitment process. All protocols were approved by the institutional ethics committees of Nanyang Technical University (IRB-2014-04 / 27 and IRB-2018-10-015-02) and Tan Tock Seng Hospital (2014 / 00416) in accordance with the Human Biomedical Research Act (Ministry of Health, Singapore). For the experiment, blood samples were collected from healthy donors via venipuncture into 4 mL sodium citrate or lithium heparin tubes (BD Vacutainer®).

[0092] Impedance Modeling To understand the effect of membrane conductivity in impedance signals, a double-shell model based on Maxwell's mixing theory (MMT) (Sun et al., IET Nanobiotechnology, 2007, 1, 69-79; Non-patent Literature 5 (Sun & Morgan, Microfluidics and Nanofluidics, 2010, 8, 423-443)) was used to model single cells with different membrane conductivity. In this model, a globular cell is suspended at the channel center between the excitation electrode and the sensing electrode. The impedance of the mixture is given by the following equation.

[0093]

number

[0094] During the ceremony:

[0095]

number

[0096] This represents the complex permittivity given the dielectric constant (ε) and conductivity (σ), and j 2 = -1, ω is the angular frequency, and φ is the volume fraction, which is the ratio of the volume of cells in the detection channel to the volume of the culture medium. The subscript "mix" indicates a mixture. l is the channel width. G is the heterogeneous fringe field. 1 This is the cell constant used to explain the complex dielectric constant of a mixture. The calculation of the complex dielectric constant of a mixture is based on the following formula:

[0097]

number

[0098] Here, φ is the volume fraction, the ratio of the volume of cells to the volume of culture medium in the detection channel. The subscripts "mix," "cell," and "med" indicate the mixture, cells, and culture medium, respectively. To describe the cellular structure, the complex dielectric constant of a cell can be calculated as a function of its internal components (membrane, cytoplasmic contents, and nucleus), as follows:

[0099]

number

[0100] During the ceremony:

[0101]

number

[0102] θ is the effective complex permittivity of two adjacent components (e.g., membrane and cytoplasm), and θ is a volume fraction, which is the ratio of the volume of the inner component to the volume of the outer component. The subscripts "in" and "out" represent the inner and outer components, respectively. In this model, the inventors first calculate the effective complex permittivity of the nucleus using the complex permittivity of the nuclear envelope and nuclear cytoplasm, and then calculate the effective complex permittivity of the cytoplasmic content using the complex permittivity of the cytoplasm and nucleus. Finally, the complex permittivity of the cell can be calculated using the complex permittivity of the cytoplasmic content and membrane.

[0103] In the design of the microfluidic device, a three-electrode design for differential measurement was used to minimize common-mode noise and signal drift. Differential current response

[0104]

number

[0105] The cells are connected to the first and second electrodes.

[0106]

number

[0107] The calculation simulates the current readout when it is between the second and third electrodes, and the second and third electrodes are also calculated to do so.

[0108]

number

[0109] Only a medium exists between them. Electrical opacity is defined as the ratio of the current response at high frequencies to that at low frequencies (0.3 MHz).

[0110] [Table 1]

[0111] computational fluid dynamics Computational fluid dynamics simulations were performed using Ansys FLUENT® (Ansys, Canosburg, Pennsylvania, USA). A partial spiral channel with a truncated channel length after a half-loop from the inlet bifurcation was simulated. It was assumed that the downstream flow after the truncated region, with a channel length equivalent to 76 times the hydraulic diameter (180.32 μm), had a negligible effect on the streamline profile at the inlet bifurcation. A pressure-based solver was used to simulate steady-state laminar flow using water as the medium to mimic the sheath fluid of PBS. Boundary conditions were set to stationary walls with no slip at the channel walls, velocity inlets at both inlets, and a 0 gauge pressure outlet at the outlet. A complete hexahedral mesh with element dimensions of 3.5 μm along the height, 8.5 μm along the width, and 310 μm along the length was generated using ICEM (Ansys, Canosburg, Pennsylvania, USA). Mesh refinement was performed by subdividing the entire mesh by 1.5 times, resulting in an increase in the number of nodes from 152,680 to 432,832. Velocity profiles along the channel height in different cross-sectional regions of the spiral channel were compared between the original mesh and the refined mesh. The difference in peak velocity between meshes was controlled to less than 8%, and the simulation results for the mesh with 152,680 nodes were selected for analysis in this study. Streamlines were initiated at the sample inlet using a discrete-phase model (DPM), and the interaction between particles and the continuous phase was disabled to simulate the dynamics of fluid streamlines. Streamline coordinates were later exported and processed in MATLAB® (Mathworks, USA).

[0112] Results and Discussion Example 1: Features of an entrance with steps To determine the optimal sample pinching effect, we first characterized various sample-to-sheath flow ratios using computational fluid dynamics (CFD) simulations (Figures 2 and 3). At a flow rate ratio of 1:10 (160 μL / min:1600 μL / min), sample streamlines (black) indicating cell / particle flow were equally spread across both the upper and lower halves of the channel. As the flow rate ratio increased (1:21), particle streamlines became hydrodynamically confined to the lower half of the channel. We further validated the simulation results based on impedance characterization of 10 μm beads (similar in size to neutrophils) (Figure 3). At flow rate ratios of 1:10 and 1:14, two distinct clusters were observed within the impedance profile because particles (initially unconfined) were inertially focused to two different equilibrium positions within the spiral channel (upper and lower halves of the channel). In contrast, under higher flow ratio conditions (1:21 and 1:43), a single cluster was present, indicating successful particle focusing in the lower half of the channel. These results highlighted the critical importance of single-stream particle focusing for improving impedance detection, as the particles also remained aligned near the electrodes. Based on these results, a flow ratio of 1:21 (80 μL / min: 1680 μL / min) was used in subsequent experiments.

[0113] The stepped sample inlet design has a height of 20 μm (the same as the detection channel) to simplify the fabrication process for bilayer photolithography. Further CFD simulations were performed to understand how the stepped sample inlet parameters (width and height) affect 3D focusing (Figures 4-6).

[0114] By varying the stepped inlet height from 20 to 60 μm (Figure 4), both 20 μm and 40 μm heights were comparable to confining sample streamlines to the lower half of the channel (>90%) at a flow ratio of 1:21. Lower stepped inlet heights (<20 μm) were not included in this study because they were unsuitable for cell processing and prone to clogging. Regarding opening width, smaller opening widths (50 μm) resulted in the formation of high-velocity flow (Figure 5) and extreme mixing of sample fluid and sheath fluid. This resulted in more dispersed sample streamlines (approximately 50% at 1:21) compared to larger opening widths (>100 μm).

[0115] Example 2: Integrated cell sorting and impedance cytometry for neutrophil activation profiling Three frequency measurement schemes (0.3 MHz, 1.72 MHz, and 12 MHz, 1V for each frequency) were used to capture three electrical characteristics of each cell: 1) the magnitude of the impedance signal at 0.3 MHz, indicating cell size; 2) membrane opacity, defined as the ratio of the magnitude of the impedance signal at 1.72 MHz to the magnitude of the impedance signal at 0.3 MHz, reflecting cell membrane capacitance; and 3) nuclear opacity, defined as the ratio of the magnitude of the impedance signal at 12 MHz to the magnitude of the impedance signal at 0.3 MHz, reflecting nuclear characteristics. Numerical impedance simulations were performed to study the effect of various nuclear parameters (nuclear membrane, nuclear size) on the opacity spectrum (Figure 8). Opacity is defined as the ratio of the impedance response at different frequencies to the impedance response at 0.3 MHz. Since certain biological events, such as apoptosis or necrosis, can lead to nuclear membrane degradation (Lindenboim et al., 2020, Vol. 6, p. 29; Kwon et al., Scientific Reports, 2015, Vol. 5, p. 15623), we investigated the effects of various conductivity levels of the nuclear envelope and observed that opacity increased with increasing conductivity (8). Neutrophils undergoing NETosis exhibit characteristic nuclear changes, such as nuclear swelling (Li et al., EMBO reports, 2020, 21, e48779). To understand whether electrical measurements can capture changes related to nuclear size, we varied nuclear sizes and revealed that opacity decreased with increasing nuclear size (8). Based on these simulations, impedance measurements at 12 MHz were performed to represent changes in the electrical properties of the nucleus. To explain the particle distribution, density scattering plots were used to represent the electrical parameters of the cell population. As a proof of concept for clinical trials (Figure 9), this platform was used to characterize various neutrophil phenotypes (1. control, 2. death (cultured at 75°C), 3. NETosis induced by PMA (phorbol 12-myristate 13-acetate), and 4. spontaneous NETosis in high pH medium (pH 9)).A comparison between dead neutrophils and control neutrophils (Figure 10) reveals that dead cells exhibit smaller electrical size and higher membrane opacity, likely due to high membrane conductivity resulting from a damaged / leaky membrane, which is characteristic of cell death (Figure 7).

[0116] Next, using the platform, we studied neutrophils undergoing NETosis, a newly discovered neutrophil function, where neutrophils release their nuclei along with antimicrobial substances to resolve infection. By treating neutrophils with PMA for 2 hours, NETosis-affected neutrophils showed increased electrical size, decreased membrane opacity (due to increased cell membrane area), and decreased nuclear opacity (due to changes in nuclear morphology (Yipp and Kubes, Blood, 122, 2784-2794, doi:10.1182 / blood-2013-04-457671(2013)). PMA is a potent NETosis inducer (Tatsiy and McDonald, Frontiers in In Immunology, 9, doi:10.3389 / fimmu.2018.02036(2018)), the inventors then determined whether the presence of pH-induced NETosis could be detected using their integrated platform. The inventors promoted spontaneous NETosis by suspending neutrophils in a highly alkaline solution (pH 9) (Khan et al., Frontiers in Medicine, 5, doi:10.3389 / fmed.2018.00019(2018)).

[0117] As a result, the changes in electrical properties of the pH9 sample were smaller than those of the PMA-stimulated sample, and the presence of a tail region with larger cell size and slightly lower membrane opacity (approximately 0.8) was evident. Interestingly, with respect to nuclear opacity, a main population exhibiting impedance characteristics similar to the control sample and a long tail region (ranging from 0.4 to 0.7) indicating the presence of spontaneous NETosis (similar to PMA treatment) were observed. Next, the healthy and NETosis populations were quantified as shown in Figure 11, and it was found that nucleus opacity provided better separation of the NETosis population compared to membrane opacity.

[0118] Finally, we examined the electrical signatures of neutrophils in response to bacterial infection. Isolated neutrophils in RPMI were analyzed at different concentrations (10×, 10⁶). 3 × and 10 6Cells were cultured with E. coli (EC) (× dilution) at 37°C for 24 hours. Figure 12 shows that low EC concentrations (EC-1M×) induced observable changes in the impedance profile, as is evident in increased cell size, decreased membrane opacity, and decreased nuclear opacity. At higher EC concentrations, the membrane opacity profile revealed two distinct populations: 1) a dead cell population with high opacity (approximately 0.9–1), and 2) a high-density activated neutrophil population (membrane opacity ranging from 0.75–0.85). With respect to nuclear opacity, the impedance profile shows a dominant high-density population with lower opacity (approximately 0.4), similar to that observed in PMA treatment. Unlike NETosis studies, neutrophils can exhibit different responses to eradicate bacteria (phagocytosis, degranulation, or NETosis), so relating the neutrophil impedance phenotype to biological function remains unclear. Overall, these results demonstrate the use of multi-frequency impedance cytometry, particularly high-frequency impedance signatures, for studying various neutrophil responses (vitality, NETosis, and bacterial responses). All these capabilities suggest that the platform described herein can be readily adapted for neutrophil response-based clinical diagnosis in various diseases, including cancer, urinary tract infections (UTIs), and sepsis.

[0119] Example 3: Integrated cell sorting and impedance cytometry for lymphocyte activation profiling When cells flow across a coplanar electrode energized with alternating current (AC), they disrupt the electric field, resulting in a change in electrical impedance. Impedance measurements at low frequencies (0.1–1 MHz) correlate with particle size due to the minimum electric field penetration through the particle (beads / cell), while impedance measurements in the high-frequency range (1–10 MHz) reflect membrane properties due to the effect of membrane polarization (Sun & Morgan, op. cit.; Honrado et al., op. cit.). In this specification, a dual-frequency measurement scheme (0.3 MHz and 1.72 MHz) was used to capture two electrical characteristics of each cell: 1. opacity, defined as the ratio of the magnitude of the impedance signal at 1.72 MHz to the magnitude of the impedance signal at 0.3 MHz, which reflects the cell membrane capacitance, and 2. the magnitude of the impedance signal at 0.3 MHz, which indicates the cell size. To prevent damage from electrochemical reactions in a highly conductive medium (0.1% BSA in PBS, approximately 1.6 S / m) (Honrado et al., cited above), the excitation voltage at each frequency was set to 1 V. Density scattering plots were used to represent the electrical parameters of both cell populations in order to explain the particle distribution.

[0120] First impedance profiling was performed on different purified lymphocyte populations: 1. healthy lymphocytes (control), 2. dead lymphocytes (10% dimethyl sulfoxide (dead), 24 hours), and 3. activated lymphocytes (anti-CD3 / CD28 Dynabeads® (activated), 24 hours). The opacity of dead lymphocytes was higher than that of healthy lymphocytes, likely due to lymphocyte apoptosis (de AbreuCosta et al., Molecules, 2017, 22, 1789), resulting in loss of membrane integrity (Zhang et al., Cell Res, 2018, 28, 9-21) and increased membrane conductance (Wang et al., Biochimica et biophysica acta, 2002, 1564, 412-420). These observations were confirmed using a dual-shell impedance model, and the simulation results showed that an increase in membrane conductance occurs at low frequencies (|Z). LF|) reduces the impedance signal and opacity (|Z HF | / |Z LF This demonstrated an increase in |) (Figure 7). For activated lymphocytes, an increase in cell size accompanied by lower opacity was observed, which may be due to the increased membrane surface area as the lymphocytes were activated (Huang et al., Biochimica et Biophysica Acta (BBA)-Biomembranes, 1999, 1417, 51-62; Teague et al., Cytometry, 1993, 14, 772-782). These impedance measurements also coincided with optical images showing a significant increase in cell size for activated lymphocytes.

[0121] Next, we considered using integrated inertial impedance cytometry for quantitative lymphocyte activation profiling (Figure 13). Briefly, the platform consists of three stages: (1) a spiral cell sorter, (2) an inertial cell focuser, and (3) an impedance detector using coplanar electrodes. Since the platform has been previously applied to neutrophil (approximately 8-10 μm) studies (Petchakup et al. (2019), op. cit.), we assumed that the device could enrich and electrically characterize activated lymphocytes (>8 μm). To further improve impedance detection sensitivity and consistency, a stepped sample inlet was added to hydrodynamically confine the sample flow to the lower half of the channel using a sheath fluid. Under the influence of inertial coupled Dean flow movement within the spiral microchannel, larger cells (hydrodynamic diameter (D)) were detected. h Particle size (A) relative to ) pA pulse strength > 0.07 equilibrates near the bottom of the channel (closer to the electrode surface) for activated lymphocyte sorting and impedance detection. To determine the optimal sample pinch effect, various sample-to-sheath flow ratios were first characterized using computational fluid dynamics (CFD) simulations (Figure 2). At a flow rate ratio of 1:10 (160 μL / min:1600 μL / min), sample streamlines (black) indicating cell / particle flow spread equally across both the upper and lower halves of the channel. As the flow rate ratio increased (1:21), particle streamlines were hydrodynamically confined to the lower half of the channel. Simulation results based on impedance characterization of 10 μm beads (similar in size to activated lymphocytes) were further validated (Figure 3). At flow ratios of 1:10 and 1:14, the particles (initially unconfined) were inertially focused to two different equilibrium positions within the spiral channel (upper and lower halves of the channel), resulting in the observation of two distinct clusters within the impedance profile (Hou et al., Scientific Reports, 2016, 6, 29410). In contrast, under higher flow ratio conditions (1:21 and 1:43), a single cluster was present, indicating successful particle focusing in the lower half of the channel. These results highlight the critical importance of single-stream particle focusing for improving impedance detection, as the particles also remained aligned near the electrodes. Based on these results, a flow ratio of 1:21 (80 μL / min: 1680 μL / min) was used in subsequent experiments.

[0122] In terms of device performance, the platform is approximately 1 x 10⁻⁶. 3 It is possible to achieve impedance measurement throughput of particles / second (Petchakup et al. (2019), op. cit.). The passive inertia-based focusing used in this platform has also been established under high flow conditions (Reynolds number, Re>50), which is suitable for high particle concentrations (up to 10 × 10⁻¹⁰). 6 It enables processing of particles / mL. In summary, this tip facilitates the processing of complex biological samples (e.g., blood) for label-free, rapid immunocellular profiling.

[0123] Example 4: Quantitative profiling of activated T lymphocytes using an integrated platform Next, an integrated inertial impedance cytometry platform was applied to the electrical profiling of different T lymphocyte (CD4+) phenotypes (control, dead, and activated). Based on the spiral channel design (height of approximately 110 μm), activated lymphocytes (>8 μm, A) were profiled. p / D h Only cells with an impedance of >0.07 experience a dominant inertial force and equilibrate near the inner wall. These cells are then sorted to the intermediate outlet and sent toward the impedance detector. Smaller, unactivated lymphocytes are located further away from the inner wall region and are removed to the bottom outlet as waste in the spiral sorter. Figure 14A shows stacked time-lapse images of cells focused on the three bifurcation points of the spiral cell sorter and pre-impedance detector for control and activated and dead lymphocyte populations. As expected, very few cells were sorted to the impedance detector for both control and dead lymphocytes. In contrast, larger activated lymphocytes showed a clear shift in flow position as they approached the inner wall, which resulted in a significant increase in cells sorted and detected by the impedance cytometer. This is interpreted as a higher impedance signal readout (higher activated cell count, approximately 15 times higher) compared to healthy controls (Figure 14B).

[0124] Furthermore, using activated CD4T lymphocyte samples (with anti-CD3 / CD28 beads), flow cytometry analysis was performed on various cell fractions, including pre-sorting (sample inlet), sorted cells (from impedance detector outlet), and discarded cells (from spiral sorter discard outlet). Due to the size-based sorting principle, an increase in forward and side scattering (FSC) was observed for cells in the sorted fraction, confirming the presence of larger cells detected by impedance (Figure 15). To determine the sorting sensitivity and impedance detection limit, samples with varying levels (0.3% to 74.6%) of spiked activated (CD137+) lymphocytes (10 5 Samples were processed at a concentration of cells / mL (Figure 16). For samples containing a large number of activated lymphocytes (>5%), approximately 60–75% of CD137+ expressing cells were consistently collected in the sorted fraction (Figure 17). For samples with low lymphocyte activation (<5%), the average post-sorting enrichment was approximately 30% (approximately 11.7 times), which may be due to random noise events. The calibration curve also shows a sensitivity of 278.8 impedance counts / mL and 1.7 × 10⁻⁶. 3 At the detection limit of cells / mL (3 standard deviations of the number of detected cells in 0% activated cells extrapolated from three different experiments), good linearity was observed between impedance counts and the percentage of activated lymphocytes in the sample (Figure 17). It should be noted that these metrics depend on the processing protocol and duration, which may be affected by the amount of initially / spontaneously activated cells present in the sample (Min, Frontiers in Immunology, 2018, 9) and the variability of biophysical properties due to the cell cycle (Huang et al., op. cit.; Kinjyo et al., Nature Communications, 2015, 6, 6301). In summary, the results clearly suggest the feasibility of using the developed integrated platform to quantify low levels of lymphocyte activation (<5%).

[0125] Example 5: Integrated workflow for CD4T lymphocyte antigen activation profiling As a proof-of-concept for quantitative lymphocyte profiling in infectious diseases, PBMCs were stimulated with PHA and PPD for 24 hours to detect lymphocyte activation using the developed platform. High cell counts were observed for control samples, which may be attributable to the presence of (larger) monocytes in the PBMCs (Figure 23). To overcome this problem, CD4 magnetically activated cell sorting (MACs, negative separation) was combined to deplete monocytes and other commensal cells before impedance profiling (Figure 18). Using this strategy, low cell counts for controls (n=552), as well as higher impedance cell counts for PHA and PPD-treated PBMCs (n=31983 and 710, respectively), indicating the presence of activated T lymphocytes, were successfully detected (Figure 19). This was in good agreement with flow cytometry analysis, as activated lymphocytes (CD69+CD137+) were significantly enriched from 0.387% to 7.017% (approximately 18-fold) for PPD treatment, which involved efficient removal of smaller lymphocytes in a spiral sorter (Figure 20). To account for donor variability, the ratio of lymphocyte activation (CD69+CD137+) in sorted cells (from the impedance outlet) of stimulated samples (PHA or PPD) to sorted cells in control samples (unstimulated) was quantified (Figure 21). Similarly, "differential impedance count" was defined as the ratio of the number of cells detected under stimulated conditions (PHA or PPD) to the respective control (unstimulated). As expected, the differential impedance count ratio was higher for PHA (approximately 105-fold) than for PPD (approximately 6-fold), attributable to a stronger mitogen response (Figure 21). Impedance cytometry detection of PPD-specific lymphocyte responses was also more sensitive than flow cytometry (approximately 6 times more sensitive (impedance), approximately 3 times more sensitive (flow cytometry)) and featured differential impedance counts greater than 1 detected in all donors.In summary, these results demonstrate the use of an integrated platform for label-free quantitative lymphocyte activation profiling, which can be further applied as a point-of-care test based on antigen-specific T lymphocyte responses.

[0126] Example 6: System Prototyping for an Integrated Impedance Cytometer To facilitate the commercialization of the integrated impedance cytometer, system prototyping was carried out as shown in Figure 24. First, the feasibility of miniaturizing the system was tested by incorporating 1. a PCB chip holder for the interface of a microchip to the impedance analyzer, and 2. a fluid pump to enable perfusion control of the sample and sheath using integrated software (Figure 24). Next, software with a graphical user interface for end-users / collaborators was developed to provide real-time data visualization and quantification for field result analysis.

[0127] The present invention has been broadly and generally described herein. Each of the narrower species and subgeneral classifications included in the general disclosure also forms part of the present invention. This includes the general description of the present invention having conditions or negative limitations that remove any subject matter from a genus, whether or not the deleted material is specifically described herein. Other embodiments are also within the scope of the following claims.

[0128] Those skilled in the art will readily understand that the present invention is well adaptable to achieve its objectives and obtain the results and benefits mentioned, as well as those unique therein. Furthermore, it will be readily apparent to those skilled in the art that various substitutions and modifications can be made to the present invention disclosed herein without departing from the scope and spirit of the invention. The compositions, methods, procedures, treatments, molecules and specific compounds described herein are current representative examples of preferred embodiments, are illustrative, and are not intended as limitations on the scope of the invention. Those skilled in the art will come up with modifications and other uses that are encompassed in the spirit of the invention as defined by the claims. Any enumeration or discussion of previously published documents herein should not necessarily be construed as an acknowledgment that such documents are part of the latest art or common general knowledge.

[0129] The present invention, as described herein exemplary, can be adequately implemented in the absence of any element(s) or limitation(s) not specifically disclosed herein. Therefore, while the present invention is specifically disclosed by exemplary embodiments and optional features, modifications and variations of the present invention embodied in the embodiments disclosed herein may be used by those skilled in the art, and such modifications and variations are considered to be within the scope of the present invention.

[0130] The contents of all documents and patent documents cited herein are incorporated in their entirety by reference.

Claims

1. A method for quantitative label-free cell profiling using impedance cytometry, wherein the method is: (i) Providing a fluid sample containing cells and an impedance cytometer in the form of a microfluidic device comprising a spiral-shaped flow channel having an inner and outer wall with respect to the center of the spiral, and having two inlets and at least two outlets, wherein one of the two inlets is a sample inlet and is located on the outer wall of the spiral-shaped flow channel, the other of the two inlets is a sheath fluid inlet and is located on the inner wall of the spiral-shaped flow channel, and at least one of the outlets comprises an impedance detector including a coplanar electrode for direct impedance quantification of cells, the sample inlet is a stepped sample inlet whose height is lower than the height of the sheath fluid inlet and the flow channel, and the sample inlet is located at the bottom of the flow channel; (ii) Perfusing the sample into an impedance cytometer having a first sample inlet and a second sheath fluid inlet and a flow channel, wherein the flow channel comprises a coplanar electrode for generating an electric field across the channel and means for detecting the breakdown of the electric field by detecting a change in electrical impedance; (iii) the step of measuring the change in electrical impedance for label-free quantitative profiling of cells in the sample; Methods that include...

2. The method according to claim 1, wherein the fluid sample is a body fluid.

3. A method according to claim 1 or 2, wherein the method is: (a) Activated lymphocyte profiling; and / or (b) Neutrophil profiling; A method intended for that purpose.

4. The method according to any one of claims 1 to 3, wherein the impedance measurement is performed at a low frequency in the range of 0.1 to 1 MHz in order to determine the cell size.

5. The method according to claim 4, wherein the impedance measurement is further performed at high frequencies in the range of >1 to 20 MHz to determine the electrical opacity as the ratio of the magnitude of the impedance signal at low frequencies to the magnitude of the impedance signal at high frequencies.

6. The method according to claim 5, wherein the impedance measurement is performed at a frequency of >1 to 10 MHz to determine film opacity and / or at a frequency of >10 to 20 MHz to determine nuclear opacity.

7. The method according to any one of claims 1 to 6, wherein the method uses microfluidic impedance cytometry.

8. The method according to claim 1, comprising the steps of introducing the sample of (i) into the port of the sample inlet at a sample flow rate and introducing the sheath fluid into the sheath fluid inlet at a sheath flow rate, wherein the sample-to-sheath flow rate ratio is 1:15 or less, and driving the sample and sheath fluid through the spiral-shaped flow channel for cell sorting and focusing.

9. The method according to claim 8, wherein step (iii) is to drive the cells selected and focused in step (ii) into at least one outlet equipped with an impedance detector, thereby directing the selected and focused cells toward the impedance detector and determining the impedance signal readout, thereby quantitatively profiling the cells.

10. The method according to claim 8 or claim 9, wherein the sample flow rate is in the range of 10 to 400 μL / min.

11. The method according to any one of claims 8 to 10, wherein the sample-to-sheath flow rate ratio is 1:20 to 1:

50.

12. The method according to any one of claims 1 to 11, wherein the height of the sample inlet is at least twice lower than the height of the sheath fluid inlet and the flow channel.

13. The method according to any one of claims 1 to 12, wherein the sample inlet has a height of 15 to 50 μm, and the sheath fluid inlet and the flow channel have a height of 80 to 250 μm.

14. The method according to any one of claims 1 to 13, wherein the opening width of the sheath fluid inlet is in the range of 200 to 700 μm, and / or the opening width of the sample inlet is 50 to 200 μm.

15. The method according to any one of claims 1 to 14, wherein the aspect ratio (width relative to height) of the sheath fluid inlet and the sample inlet is in the range of 3 to 10.

16. The method according to any one of claims 1 to 15, wherein the area ratio of the sheath fluid inlet to the sample inlet is in the range of 10 to 50.

17. The method according to any one of claims 1 to 16, further comprising a cell purification step, wherein, prior to impedance profiling in steps (ii) and (iii), the sample is incubated with magnetic beads that bind to undesirable cells, and magnetic cell sorting (MACS) is performed by separating the magnetic beads to which the undesirable cells are bound from the unbound cells using a magnetic field.

18. The method according to claim 17, wherein the undesirable cells include monocytes.

19. A biosensor for quantitative label-free cell profiling using impedance cytometry, wherein the biosensor is: (i) A microfluidic device comprising a spiral-shaped flow channel having an inner wall and an outer wall with respect to the center of the spiral, and having two inlets and at least two outlets, wherein one of the two inlets is a sample inlet and is located on the outer wall of the spiral-shaped flow channel, and the other of the two inlets is a sheath fluid inlet and is located on the inner wall of the spiral-shaped flow channel, the sample inlet is a stepped sample inlet whose height is lower than the height of the sheath fluid inlet and the flow channel, and the sample inlet is located at the bottom of the flow channel; (ii) A coplanar electrode for direct impedance quantification of cells, the coplanar electrode being integrated with at least one of the outlets of the microfluidic device; A biosensor equipped with these features.

20. The biosensor according to claim 19, wherein the height of the sample inlet is at least twice lower than the height of the sheath fluid inlet and the flow channel.

21. The biosensor according to claim 19 or 20, wherein the sample inlet has a height of 15 to 50 μm, and the sheath fluid inlet and the flow channel have a height of 80 to 250 μm.

22. The biosensor according to any one of claims 19 to 21, wherein the opening width of the sheath fluid inlet is in the range of 200 to 700 μm, and / or the opening width of the sample inlet is 50 to 200 μm.

23. The biosensor according to any one of claims 19 to 22, wherein the aspect ratio (width relative to height) of the sheath fluid inlet and the sample inlet is in the range of 3 to 10.

24. The biosensor according to any one of claims 19 to 23, wherein the area ratio of the sheath fluid inlet to the sample inlet is in the range of 10 to 50.

25. An integrated impedance cytometer comprising a biosensor according to any one of claims 19 to 24, further comprising (1) a microchip, (2) an impedance analyzer, and (3) a fluid pump connected to the sample and sheath fluid inlets to enable perfusion control, wherein the microchip is connected to the impedance analyzer and the fluid pump, and the impedance analyzer is connected to the coplanar electrode to enable impedance measurement.

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