OA17023A - Counting particles using an electrical differential counter. - Google Patents

Abstract

This disclosure relates to methods and devices to count particles of interest, such as cells. The methods include obtaining a fluid sample that may contain particles of interest; counting all types of particles in a portion of the sample using a first electrical differential counter to generate a first total; removing any particles of interest from the portion of the fluid sample; counting any particles remaining in the portion of the fluid sample using a second electrical differential counter after the particles of interest are removed to generate a second total; and calculating a number of particles of interest originally in the fluid sample by subtracting the second total from the first total, wherein the difference is the number of particles of interest in the sample. These methods and related devices can be used, for example, to produce a robust, inexpensive diagnostic kit for CD4+ T cell counting in whole blood samples.

Description

This invention relates to counting particles such as cells, and more particularly to counting particles using electrical differential counters.

BACKGROUND

Counting of particles, such as cells, is of significant use in medicine and public health. One widely used cytometry System involves optical devices, such as flow cytometers, and tags cells of interest with optical labels (such as fluorescent markers) and interrogates them with light sources such as lasers.

The Coulter principle of impédance cytometry, based on resistive-pulse sensing, is well-established for counting cells non-optically. In its original format, Coulter counting allowed for différentiation of cells by size, to enable counting of individual subsets of a mixed population, such as a whîte blood cell diflerential. A second génération of impédance spectroscopy methods builds on the original Coulter principle and interrogates cells across a sweep of altemating current (AC) frequencies.

Microfluidic Systems hâve shown unique promise for studying cell function, cell and tissue engineering, disease diagnosis, blood sample préparation, and drug discovery. Very recently, the use of microfluidics to isolate pure populations of leukocyte subsets from whole blood has attracted significant interest for point-of-care diagnostics. While the principle behind a cell isolation approach can be easily adapted to a wide spectrum of clinical applications, detecting these isolated cells remains a technical challenge to be addressed.

SUMMARY

This disclosure describes Systems and methods for counting particles of interest in a mixed population of particles using a simple, low-cost electrical method. Using a differential counting method with an electrical différentiel counter, these Systems and methods can be used to count a subset of particles, e.g., white blood cells, from a starting sample, e.g., of whole blood, beyond the capability of current Coulter type Systems and methods. For example, Systems with two electrical impédance sensors can be used to obtain an absolute CD4+ T cell count from a blood sample.

In one aspect, the disclosure includes methods of counting particles of interest, such as cells, e.g., white blood cells, e.g., CD4+ T cells, in a sample, e.g., whole blood, that includes two or more different types of particles. These methods include obtaining a fluid sample that may contain particles of interest; counting ail types of particles in a portion of the sample using a first electrical differential counter to generate a first total; removing any particles of interest from the portion of the fluid sample; counting any particles remaining in the portion of the fluid sample using a second electrical differential counter after the particles of interest are removed to generate a second total; and calculating a number of particles of interest originally in the fluid sample by subtracting the second total from the first total, wherein the différence is the number of particles of interest in the sample.

In these methods, the first and second electrical differential counters can be the same or a different electrical differential counter. In some implémentations, these methods can further include reversing a flow direction of the fluid sample after removing the particles of interest from the portion of the fluid sample. In other implémentations, the methods can further include maintaining a flow direction of the fluid sample while counting ali types of particles in the portion of the sample; removing particles of interest; and counting any particles remaining in the portion of the fluid sample. In these methods, the particles, e.g., cells, of interest are removed from the portion of the fluid sample using one or more binding agents or moieties, such as antibodies, e.g., that specifically binds to a spécifie surface marker on the particle of interest, such as a white blood cell, such as a CD4+ T cell, or a particulate type of white blood cell, or a plateïet, or other spécifie cell in the sample, such as a tumor cell, e.g., a circulating tumor cell (CTC).

In other implémentations, the methods can further include depleting selected particles from the portion of the sample before counting ail types of particles in the portion of the sample. For example, if the sample is whole blood, the method can include depleting érythrocytes in the blood using a lysis technique. In other implémentations, for example, the fiuid sample can include whole blood and the method can include depleting érythrocytes, monocytes, neutrophils, CD8+ lymphocytes, or other cellular components of blood by immuno-depletion.

In certain implémentations, the particles of interest are CD4+ T cells, and removing the particles of interest includes capturing CD4+ T cells in a capture chamber functionalized with anti-CD4 antibodies. The methods can further include removing nonspecifically adsorbed leukocytes by purging the capture chamber with phosphate buffered saline. The methods can also further include determining a cell flow direction based on a polarity of an impulse signal generated by the first electrical differential counter.

In another aspect, the disclosure includes devices that include a microfluidic chip defining a channel including an inlet and an outlet; a capture chamber arranged along the channel between the inlet and the outlet, wherein the chamber is configured to capture particles of interest from fluid flowing through the channel; a first electrical differential counter arranged to count ail types of particles in a fluid flowing into the capture chamber; a second electrical differential counter arranged to count ail types of particles remaining in the fluid flowing out of the capture chamber; and a computing mechanism in electronîc communication with the first and second electrical differential counters, wherein the computing mechanism calculâtes a number of particles of interest based on signais from the first and second electrical differential counters.

In different implémentations of these devices, the first and second electrical differential counters can be the same or different electrical differential counters. The devices can further include a pump System in fluid communication with the channel, wherein the pump System is opérable in a first mode to cause fluid to flow in a first direction in the channel past the first electrical differential counter and opérable in a second mode to cause fluid to flow in a second direction in the channel opposite the first direction and back to the first electrical différent! al counter.

In certain implémentations, a portion of the channel can define a flow path that extends in a loop from the first electrical différentiel counter through the capture chamber and back to the first electrical différentiel counter.

In various implémentations, the capture chamber includes surfaces functionalized with bindîng agents, such as anti-CD4 antibodies.

In another aspect, the disclosure includes kits that include a device as described herein; a solution that includes a binding agent or moiety, such as an antibody, e.g., that specifically binds to a spécifie surface marker on a white blood cell, such as aCD4+ T cell, with an affinity for the particles of interest; and a solution comprising a lysing agent effective to lyse selected particles without lysing the particles of interest. In the devices in these kits, the first and second electrical differential counters can be the same or different electrical differential counters, and the devices can further include a pump System in fluid communication with the channel, wherein the pump system is opérable in a first mode to cause fluid to flow in a first direction in the charme! past the first electrical differential counter and opérable in a second mode to cause fluid to flow in a second direction in the channel opposite the first direction back to the first electrical differential counter.

In some implémentations, a portion of the channel defines a flow path that extends in a loop from the first electrical differential counter through the capture chamber and back to the first electrical differential counter.

In another aspect, the disclosure describes micro fluidic chips that include a plurality of capture chambers, wherein the capture chambers are configured to capture particles of interest from fluid flowing through the chambers; an electrical differential counter opérable to count particles in a mixed population of particles in fluid flowing into the capture chambers and to count particles remaining in fluid flowing out of the capture chamber; and a computing mechanism in electronîc communication with the electrical differential counter, the computing mechanism opérable to calculate a number of particles of interest based on signais from the electrical differential counter.

These microfluidic devices can further include a fluidic channel coupled to the plurality of chambers, wherein the fluidic channel includes a first channel région and a second channel région, wherein the first channel région is configured to receive a lysing solution and a sample fluid, and mix the sample fluid with the lysing solution, and wherein the second channel région is configured to receive a quenching solution and a lysed solution from the first channel région, and mix the quenching solution with the lysed solution.

In any of the forgoîng aspects and implémentations, the binding agents or moieties can be selected from antibodies, antibody fragments, oligo- or polypeptides, nucleic acids, cellular receptors, ligands, aptamers, MHC-peptide monomers or oligomers, biotîn, avidin, oligonucleotides, coordination complexes, synthetic polymers, and carbohydrates.

Also in any of the forgoing aspects, the sample can be a blood sample, the binding moiety can bind to CD66, CDI4, CD4, CDS, EpCAM, E-Selectin, or P-Selectîn, and the desired cell can be selected from neutrophils, monocytes, lymphocytes, circulating tumor cells (CTCs), HIV infected CD8 lymphocytes, circulating endothélial cells, and platelets. In some implémentations, the desired cells of interest are CD4+ lymphocytes. In this implémentation, the sample may be obtained from a patient at risk of developing AIDS.

By a patient is meant a living multicellular organism. The term patient is meant to include humans, mice, dogs, cats, cows, sheep, horses, non-human primates, and fish.

By binding moieties or “binding agents’’ is meant a molécule that specifically bînds to an analyte (e.g., a cell). Binding moieties include, for example, antibodies, aptamers, receptors, ligands, antigens, biotin/avîdin, métal ions, chelating agents, nucleic acids, MHC-peptide monomers, tetramers, pentamers, or other oligomers.

By cell surface marker is meant a molécule bound to a cell that is exposed to the extracellular environment. The cell surface marker can be a protein, lipid, carbohydrate, or some combination of the three. The term cell surface marker includes naturally occurring molécules, molécules that are aberrantly présent as the resuit of some disease condition, or a molécule that is attached to the surface of the cell.

By lysis is meant disruption of the cellular membrane. For the purposes of this invention, the term lysis” is meant to inciude complété disruption of the cellular membrane (complété lysis), partial disruption of the cellular membrane (partial lysis), and permeabilization of the cellular membrane.

By binding moîety is meant a chemical species to which a cell binds. A binding moiety may be a compound coupled to a surface or the material making up the surface. Exemplary binding moieties inciude antibodïes, antibody fragments (e.g., Fe fragments), oligo- or polypeptides, nucleic acids, cellular receptors, ligands, aptamers, MHC-peptide monomers or oligomers, biotin, avidin, oligonucleotides, coordination complexes, synthetic polymers, and carbohydrates.

The term chamber Îs meant to inciude any designated portion of a micro fluidic channel, e.g., where the cross-sectional area is greater, less than, or the same as channels entering and exiting the chamber.

The methods and devices described herein provide several benefits and advantages. In particular, the approaches described herein can be used to provide novel devices for cell analysis that are smaller, less expensive, and simpler to use than presently existing large, expensive, and complex flow cytometers, Coulter counters and impédance spectroscopes. The devices described herein can be used to discriminate a wider number of cell types and subtypes than currently known Coulter counters and impédance spectroscopes. The smaller, less expensive, microfabricated devices described herein can require much smaller volumes of blood or plasma and expensive reagents. They can be less expensive to operate and maintain. These devices represent mobile platforms that can be used at the point of care, independent of health care infrastructure. As closed, one-time use, disposable devices for the handiing of blood and other biohazardous fluids, these devices reduce System risks and costs. Thus, the new methods and devices can be used to diagnose various diseases such as H1V/A1DS and cancers such as leukemia, and can be used to monitor a patient’s progress with médication, e.g., to détermine the overall efficacy of a particular treatment regimen used for a given patient.

Compared to optical cytometry methods, the simplicity of the electrical interrogation methods as described herein, and extended to multî-frequency impédance methods can be used to create a more streamlined, cost-effective, and mechanicaliy robust solution for portable cellular analysis. The devices described herein are simpler and less expensive, in part, because they do not require a stable light path and the associated lensing, filtering, and focusing mechanisms that can add cost and complexity to optical détection methods. Moreover, the devices described herein can hâve higher throughput, than optical détection devices, which tend to hâve low throughput because of the small détection area available at a single time.

The microfabricated cell counters described herein are unlike Coulter counters, in that they can be used to count complex subsets of cells in a simple, handheld System without the need for extemal cell surface labels and other reagents, which add complexity and cost to the assay. Moreover, unlike cell counting strategies like flow cytometry and impédance measurement, the microfabricated cell counters described herein can be used with cells attached to surfaces even to count small numbers of cells on large surface areas in a relatively large volume.

Détection and énumération of cells are essential for medical diagnostics, especially for AIDS and cancer diagnosis, and pathogen détection. While most existing methods to detect cells are optical (i.e., microscopy), electrical détection is significantly simpler, cheaper, and more amenable to point-of-care devices. To date, electrical détection and énumération of intact cells based on impédance spectroscopy (i.e., détection of changes in electrical impédance caused by the presence of cells) hâve proven to be extremely practical and inexpensive, but limited to large cell populations or homogenous cell types (e.g., Coulter counting of red blood cells or total lymphocytes).

The combination of sélective particle déplétion in a microfluidic device using controlled shear flow, with double counting provides the new particle counting Systems based on a subtraction assay concept. Both the microfluidic particle capture methods and the résistive puise particle count methods are extremely robust and simple, and can thus be used to produce a robust, inexpensive diagnostic kit, e.g., for CD4 cell counting. For example, referring to FIG. 1, a droplet of whole blood provided by a finger stick can be applied to the inlet of a chip incorporating the cell counting techniques described herein. Red blood cell lysis and absolute CD4+ T cell counting, as well as on-chip sample préparation for a subséquent viral load test, can be performed on the chip.

Unless otherwise defined, ail technical and scientific terms used herein hâve the same meaning as commonly understood by one of ordinary ski 11 in the art to which this invention belongs. Although methods and materiaîs similar or équivalent to those described herein can be used in the practice or testing of the présent invention, suitable methods and materiaîs are described below. Ail publications, patent applications, patents, and other référencés mentioned herein are incorporated by référencé in their entirety. In case of conflict, the présent spécification, inciuding définitions, wili control. In addition, the materiaîs, methods, and examples are illustrative only and not intended to be limitîng.

Other features and advantages of the invention wili be apparent from the following detailed description, and from the daims.

DESCRIPTION OFDRAWINGS

FIG 1 is a schematic of a cell counting method and device, inciuding test cartridges that include the microfluidic chips described herein.

FIGS. 2A-2E are schematics of use a microfluidic circuit in a cell counting device.

FIG 3 is a schematic of a chip incorporating a cell counting device.

FIG 4 is a schematic illustrating fabrication of a cell counting device.

FIG 5 is a schematic of a differential cell counter experimental setup.

FIG 6 is a circuit schematic of a self-referencing electrical sensor using three électrodes connected in a Wheatstone bridge configuration.

FIG 7 is graph comparing estimated inlet concentrations with measured chip concentrations.

FIG 8 is a graph presenting entrance and exît counts for a passivated capture chamber experiment.

FIG 9 is a graph illustrating the relationshîp between white blood cell concentration and the dîscrepancy between the entrance and exit counts.

FIGs. lOAto 10C are a sériés of graphs illustrating the effect of inciuding a shearing module in a cell counter.

FIG 11 is a schematic of a difierential cell counter experimental setup based on a reverse flow concept using of a single pair of électrodes for a difierential CD4+T cell count.

FIG 12 is a graph comparing the error in counts found using the reverse-flow difierential counter protocol with the total number of cells counted.

FIG 13 is a graph présenting entrance and exit counts for a passivated capture chamber experiment using the reverse-flow difierential counter protocol.

FIG 14 is a sériés of merged images of an entire difierential counter chip with magnification of régions near the entrance (1), mid-section (2), and exit (3).

FIG 15 is an area histogram of circular objects on a chip as observed using optical counting.

FIG 16 is a graph presenting forward and reverse flow counting of CD4+ T cells.

FIG. 17 is a schematic of a device using a single electrode set for counting cells flowing into and out ofthe capture chamber. The device includes a counting device in which a portion of the channel defines a flow path that extends in a loop from the first electrical difierential counter through the capture chamber and back to the first electrical difierential counter as shown in FIG. 17.

FIGs. 18 is a schematic of a particle counting device and a graph of impédance signal as a fonction of time showing the signais caused by partîcles flowing in opposite directions.

FIG. 19A is a graph comparing % error found using the reverse-flow difierential counter protocol with the total number of cells counted.

FIG. I9B is a graph that illustrâtes the cumulative forward and reverse counts for cells using the reverse-flow difierential counter protocol.

FIG. 20A is a graph comparing electrical and optical counts. FIG. 20A shows results from 14 CD4+ T cell counting experiments using white blood cells purified from human whole blood samples and the close corrélation (y = 0.994x, R2 = 0.997) between the electrical difierential method and the optical control.

FIG. 20B is a graph depicting a Bland-Altman analysis of the data in FIG. 20A.

FIG. 21 is a graph comparing % errer with CD4+T cell counts. FIG. 21 illustrâtes how the percent errer (absolute différence in optical and electrical counts, normalîzed by the CD4+ T cell count) relates to the total number of CD4+ T cells counted.

FIG. 22 is a graph that illustrâtes the génération of discrète impédance signal trigger threshold levels.

FIGs. 23A to 23C show the results of the dynamic threshold analysis procedure. FIG. 23A shows differential counts vs. trigger level and shows stabilîty between 8x and 12x trigger levels. Slope FIG. 23B and curvature FIG. 23C analysis identifies 12x as the optimal trigger level because it is part of the most stable régime in the curve.

FIG. 24 is a graph that illustrâtes the cumulative forward and reverse counts for cells using the 12x trigger threshold level.

FIG. 25 is a schematic of a differential cell counter.

FIG. 26 is a plot of percent errer of differential cell counts for whole blood samples.

Like reference symbols in the various drawings indicate like éléments.

DETAILED DESCRIPTION

The new Systems and methods are based on a simple and low-cost electrical counting method and can be used to count particles of interest in a mîxed population of particles in a sample, such as a fluid sample, or a particulate sample dispersed in a fluid. Using differential counting methods with an electrical differential counter, these Systems and methods can be used to count a subset of white blood cells from a starting sample of whole blood. For example, Systems with two electrical impédance sensors can be used to obtain an absolute CD4+ T cell count from a blood sample.

The new micro-scale devices operate using a novel subtraction impédance interrogation technique. In the described methods, a complex mixture of particles in a starting sample is passed through an electrode configuration for resistîve-pulse or impédance sensing, and a total count of particles in the collective starting sample can be obtained. Next, particles of interest can be selectively retained in a microchannel through the use of a spécifie, immobilized capture reagent under controlled shear flow. Finally, the remaining population of particles in suspension can be passed through a second electrode configuration for resistive-pulse sensing, and a second count of the total population, depleted ofthe particles ofinterest, can be obtained. The différence between the two counts represents the count of the captured particles, and thus, the particle count of interest.

This approach can be used, for example, in a CD4+ T cell micro-cytometer, which is a micro-scale device for CD4+ T cell counting and which can be used as part of a kit for use in a point-of-care System for monitoring CD4+ T cell counts. In this implémentation, whole blood is passed through an electrode sensing région, and the total particle count is obtained for the collective starting sample. The CD4+ T cells in the sample are selectivcly depleted through the use of anti-CD4 antibodies, immobilized in a microfluidic chamber or channel under controlled shear flow. The remaining population of particles in the CD4+ T cell depleted whole blood is passed then through a second electrode sensing région, and a second count of the total population depleted of the particles of interest is obtained. The différence between the two counts represents the count of the captured CD4+ T cells. This kit, device, and method can be used for counting CD4+ T cells from a finger stick of blood at the point of care.

As shown in FIG. 1, the device can be fully realized as (1) a one-time use, disposable cartridge 10 that contains all the micro fluidics and sensing éléments described herein, and (2) a hand-held cartridge reader 20, which provides the electrical sensing, stimuli, and fluidic controls (e.g., pumping mechanisms). The top of FIG. 1 also shows a flow diagram of the path of a droplet of whole blood, e.g., provîded by a finger stick, from application to the inlet of a device (e.g., a sample cartridge and reading unit) incorporating the cell counting techniques described herein. The blood passes through a red blood cell lysis station and an absolute CD4+ T cell counting station, as well as an on-chip sample préparation station and a subséquent viral load test station.

As shown in FIG. 1, the drop of blood 30 (-10 to 20 μί volume) would be dropped onto the cartridge's receiving port after (or before) the cartridge 10 is inserted into the reading unit 20. The reading unit 20 would control the infusion ofthe blood and other fluids through the cartridge in addition to applying the electrical signal to the cartridge's sensing région and reading the change in the electrical signal caused by the passage ofcells through the cartridge 10. The reading unit 20 would then analyze the electrical signais and calculate the concentration of the target cells, which would be displayed to the operator. As discussed in more detail below, different cartridges can be designed to sense for different diseases simply by changing the binding agent, e.g., antibody type, in the chip's capture région.

A Cell Counting Device in Operation

Use of an exemplary cell counting device 100 is illustrated in FIGS. 2A- 2E. Cell counting device 100 indudes two impédance sensors 110 and 111. The cell counting device 100 defines a microfluidic circuit or channel 112, which extends from a sample inlet 114 through a sélective particle déplétion chamber or capture chamber 116 to a sample outlet 118. The sample inlet 114 receives an unprocessed or a processed sample to be analyzed. The following discussion describes the use of cell counting device 100 to count CD4+ T cells in a sample of whole blood. Cell counting devices as described herein can also be used to analyze other samples including, for example, plasma, urine, sputum, or other biological or other fluids, e.g., industrial fluids, that contain two or more different types of particles.

Cell counting device 100 indudes an optional reagent inlet 120, where one or more sample processing reagents can be introduced and mixed with the sample. In some instances, reagents introduced through this manner can be red blood cell lysing reagents, sample stabilization reagents, particle surface labels, or other reagents of interest.

Channel 112 can include an optional sample processing area 122, where the starting sample can be further processed or purified to make particle counting faster, more accurate, or more efficient. In cell counting device 100, the sample processing area 122 is a red blood cell lysis area and a monocyte déplétion area. For example, the sample processing area 122 can include surfaces coated with a monocyte capture reagent such as an anti-CD 14 antibody. In general, the capture chambers are functionalized or coated with binding agents or binding moieties as described herein. These binding moieties are selected to specifically bind to the particles, e.g., to surface markers on cells, and not to other particles that may be présent in the sample. The sample processing area 122 can be a red blood cell lysis area, or a monocyte déplétion area, or both.

Impédance sensors 110 and 111 are located in channel 112 on each side of capture chamber 116. Impédance sensors 110 and 111 are electrode configurations for the countïng of particles in fluid flowin g through the channel 112. The impédance sensors 110 and 111 can be two-electrode or three-electrode résistive puise sensors of the Coulter type, for the countïng of btood cells. The current implémentation uses a coplanar electrode configuration, meaning ail électrodes are on the same surface, and an AC signal is being passed between the électrodes. In other implémentations, the impédance sensors 110 and 111 may be configured where each electrode and its mate are parallel to each other (still perpendicutar to fluid flow direction), but one electrode is on the floor of the chamber while the other is on the ceiling of the chamber. The électrodes could also be placed parallel to each other, but at the sides of the channel (still perpendicular to the flow of cells). Another implémentation is a fluidic electrode, where an electrical signal is passed through a small channel with a conductive solution that flows perpendicularly to the cell flow direction. The electrical teads in this case could be microfabricated or métal wires placed in each end of the fluidic electrode channel.

In addition, an AC (altemating current) or DC (direct current) signal can be used to sense cell passage. For a DC signal, Ag-AgCl (silver/silver chloride) électrodes could be used, as they provide excellent redox reaction efficiency even under high electrical current. In other implémentations, the impédance sensors 110 and 111 can be, for example, capacitive sensors, résistive sensors, or other sensor modalities that measure the intrinsic optical or magnetic properties of the cells in a label free manner, or sensor modalities that measure labels associated with the cells.

Capture chamber 116 is a sélective particle déplétion or capture chamber, where particles of interest are selectively captured onto a surface or surfaces of the chamber using binding moieties such as analyte capture or binding agents and controlled shear, substantially as described in US 2009/0298067 Al, “Devices and Methods for Detecting Cells and Other Analytes” (which is incorporated herein in its entirety). In some implémentations, capture chamber 116 is functionalized with anti-CD4 antibodies and serves as a sélective CD4+ T cell déplétion chamber. Of course, capture chamber 116 can be functionalized with any other binding agents, e.g., antibodies, aptamers, and binding pairs, which selectively bind to the spécifie particle or particles of interest. Such binding agents are known, or can be easily determined, for a given particle, e.g., cell, of interest.

In some implémentations, the cell countîng device 100 includes an optional fluidic entry channel 124 for sendïng reagents into the capture chamber 116 and an optional fluidic exit channel 126 for removing reagents sent into the capture chamber 116. The optional fluidic entry channel 124 and the optional fluidic exit channel 126 can be used, for example, to selectively functionalize the chamber with binding moieties.

The sample outlet 118 collects flow-through sample and sends it downstream, for example to a self-contained waste area. In some instances, the sample outlet 118 collects flow-through sample and sends it downstream to a downstream assay, or a further processing area on the microfluidic chip.

In some implémentations, the cell counting device 100 also includes an optional sélective sample processing area 128, where the sample is processed prior to mixing with reagents introduced through the reagent entry inlet 120. For exempte, the sélective sample processing area 128 can be a sélective filtration area where unwanted particles are filtered mechanically or chemically.

Before use, the cell counting device 100 is prepared by using the fluidic entry channel 124 and the fluidic exit channel 126 to selectively functionalize the capture chamber 116 with a binding agent, e.g., an antibody spécifie to the CD4 antigen that résides on the surface of the helper T cells and monocytes (though containing an order of magnitude less than the helper T cells.

In use, the sample, e.g., whole blood is introduced into the cell counting device 100 through the sample inlet 114 and a chemical to lyse the red blood cells is introduced into the cell counting device 100 through the reagent inlet 120 (see FIG. 2B). Flowing through the sample processing area 122, red blood cells are lysed as the whole blood mixes with the red blood cell lysing agent and monocytes are captured on surfaces coated with a monocyte capture reagent such as an anti-CD 14 antibody (see FIG. 2C). Ail white blood cells are counted as they pass the entrance impédance sensor 110. The enumerated cells enter a large capture chamber 116 that is functionalized with an antibody spécifie to the CD4 antigen. The capture chamber 116 retains CD4 T cells and monocytes while the remainder of the white blood cells exit the capture chamber 116 and are enumerated by the exit counter 111 (see FIG. 2D). PBS can then be introduced through reagent inlet 120 to wash away non-specifically bound cells in the capture chamber 116. A first electronic processor 130 is linked to the first electrode configuration 110, and records individual particle signais as résistive puises or other electrical measurements. A second electronic processor 132 is linked to the second electrode configuration 111, and records individual particle signais as résistive puises or other electrical measurements. With a known sample volume, the concentration of helper T cells can be obtained by finding the différence between the entrance and exit counts (see FIG. 2E).

This method can be adapted to count other cell types simply by choosing different antibodies for the particular cell surface antigen. The red blood cell lysis région can increase throughput, as érythrocytes' hâve a concentration of 5 x 10⁹/mL in whole blood, 10 which would prove quite difficult to count in a timely manner necessary for a global health diagnostics application. In addition, the sensitivity and accuracy in finding helper T cell counts would be severely diminished by the presence ofthe red blood cells. For example, if 10 pL of blood sample is analyzed, approximately 5x10 red blood cells, 1 x IO⁵ white blood cells, and 1x10* helper T cells (in a healthy adult) would be counted.

Only 0.02 percent of the counted cells would be helper T cells, which could easily be masked by the non-ideal situation of red blood cells being counted at the entrance, but not at the exit (one reason being that some red blood cells non-specifïcally adsorb to the capture chamber). Removal ofthe red blood cells would increase the percentage of helper T cells to 10% out of the total cells counted, greatly increasing the chip's accuracy and précision in providing cell counts.

Design and Fabrication

FIG. 3 shows a differential counter device 100 without the red blood cell lysis région 122. The fluidics layer (a) contains inlet and outlet ports for cell sample flow and two ports used to functionalize the 50 pm-high (6.6 pL) capture région with antibodies.

The two impédance sensing régions are made with 15 pm-wide and 15 pm-high channels that funnel the cells over three 10 pm-wîde platinum électrodes, spaced 10 pm apart (b).

The height of the capture région was chosen to increase the volume of sample and ensure the proper shear stresses at the wall-fluid interface. According to Cheng et al. (“Cell détection counting through cell lysate impédance spectroscopy in microfluidic devices, 30 Lab on a Chip, vol. 7, pp. 746- 755,2007), a shear stress of >3 dyn’cm'² resulted in less effective CD4 T cell capture. The équation

can be used to estimate the shear stress at the walls of a rectangular microfluidic channel 5 of a constant width, ω>, where μ is the dynamic viscosity of the fluid, Q is the volumétrie flow rate, and h is the height of the channel (Usami et a!., “Design and construction of a linear shear stress flow chamber,” Annals of Biomédical Engineering, vol. 21, no. 1, pp.

77-83, January 1993). This shows the sensitive, inverse-squared relationship between the channel height and the shear stress at the chamber's ceiling and floor. A 15 pm capture channel would give a shear stress of 10 dyn*cm‘², well above the aforementioned maximum shear stress Hmit. This shear stress would create a force of -155 pN on a 10 pm cell's membrane, which is the same order of magnitude as the dissociation force of antibody-antigen interactions (see, e.g., Hinterdorfer et al., “Détection and localization of individual antibody-antigen récognition events by atomic force microscopy, Proceedings of the National Academy of Sciences of the United States of America, vol. 93, no. 8, pp.

3477-3481, 1996; Dammer et al., “Spécifie antigen/antibody interactions measured by force microscopy,” Bîophysical Journal, vol. 70, pp. 2437-2441, May 1996; and Harada et a!., “Spécifie and quantized antigen-antibody interaction measured by atomic force microscopy,” Langmuir, vol. 16, no. 2, pp. 708-715, November 2000).

A 50 pm capture channel height greatly reduces the average shear stress to 0.45 dyn’cm'², resulting in a force of -14 pN on the cell and greatly increasing the cell's surface antigen interactions with the immobilized Ab to facilitate cell capture. The 34 mm capture channel length ensures sufficient interaction time (about 80 seconds at sample flow rate of 5 jL^min'¹).

Three-dimensional hydrodynamic focusîng was desired, but would hâve efiectively increased the entrance flow rate 0125 pL/minute for a 5 pL/minute cell sample flow rate) and corresponding shear stress of 11.1 dyn*cm*², which is well beyond the maximum to facilitate CD4+ T cell capture. In addition, the cell passage time through the 15 pm x 15 pm counter pore at this flow rate would resuit in transition times faster than 90 ns, which is well below the minimum transition time of ~2 ps that can be resolved using the lock-in amplifier described in the experimental section.

The fluidics and electrical sensing layers are then aligned and bonded to form the completed differential counter (c).

Fabrication of the differential counter is illustrated for one counter région in FIG.

4. The electrical sensing layer can be fabricated using the standard métal lift-off process.

A4 glass wafer (Pyrex® 7740) is first spin-coated with LOR2A liftoff resîst, soft-baked at 183 °C for 5 minutes and is coated with S-1805. After another soft-bake at 110 °C for 90 seconds, the wafer is aligned to the électrodes mask on a Quintel Q7000 IR backside mask aligner and exposed for a total dose of 2.8 mJ*cm*². The wafer is then placed on a 110 °C hotplate for a 60 seconds post-exposure bake before being îmmersed into

Microposit MF CD-26 developer for 80 seconds and rinsed with DI water for 2 minutes (FIG. 4(a)). The wafer is then de-scummed in an O₂plasma system for 20 seconds before being placed in a CHA Evaporator for the déposition of 25 nm of Ti seed layer, followed by a 75 nm Pt conduction layer (FIG. 4(b)). The undesired métal is lifted off by placing the wafer in a 70° C bath of Microchem Remover PG for 15 minutes, creatîng the necessary conduction paths for the referenced counters (FIG. 4(c)).

The multi-height fluidics layer is created by fabricating a négative image of the desired channels using Microchcm SU-8 25 photoresist. SU-8 25 is spun on a 4 Si wafer to a height of 15 pm, and is pre-baked in two steps for 2 minutes at 65 °C and then 95 °C for 5 minutes. The wafer is aligned and exposed to a mask defming ail of the fluidic channels, including the capture région, counters, samplc inlet and outlet, and Ab functionalization ports (FIG. 4(d)). A second layer of SU-8 is spun on to obtain a total thickness of 50 pm for the entire wafer, and is pre-baked at 65 °C for 5 minutes and then 95 °C for 15 minutes. The wafer is then exposed to a second mask only defining the capture chamber, allowing it to hâve a height of 50 pm, compared to the other fluidic régions of 15 pm in height. The wafer is devcloped in Microchem SU-8 developer for 2 minutes at room température, rinsed with isopropyl alcohol, and hard-baked at 125 65 °C for 15 minutes (FIG. 4(e)). Polydimethylsiloxane (PDMS), l:10::curing agenf.base, is poured over the négative mold and allowed to cure ovemight at 65 °C (FIG. 4(f)). The polymerîzed mold is peeled off, and ports are punched for ail inlets and outlets using a blunt syringe needle.

The sealed fluidic chip is completed by aligning and bonding the electrode sensing layer to the fluidies layer after oxygen plasma activation in a barrel etcher (FIG. 4(g)). Teflon microbore tubing is used to make fluidic connections between the chip and syringe pumps. The lysis région can be completed using the techniques described in Sethu et al. hâve shown that it is feasible to croate a microfluidic red blood cell lysis device using diffusive mixing (see, e.g., Sethu et al., “Continuons flow microfluidic device for rapid érythrocyte lysis,” Analytical Chemistry, vol. 76, pp. 6247-6253,2004 and Sethu et al., “Microfluidic isolation of leukocytes from whole blood for phenotype and gene expression analysis,” Analytical Chemistry, vol. 78, pp. 5453-5461,2006). Differential Counter Setup

FIG. 5 illustrâtes an example of a setup that can be used to differentially count CD4+ T cells. Initially, a pump, such as a Harvard Apparatus PicoPlus syringe pump, is used to flow a known volume of sample, e.g., white blood cells (from whole blood samples with lysed red blood cells), into the chip inlet and through the entrance counter, capture chamber, e.g., a CD4 Ab-functionalized capture chamber, and exit counter at a steady flow rate, e.g. 2,3,4, 5,6, 7, 8,9, or 10 pL/minute. After sample flow, PBS is pumped into the chip at a higher flow rate to remove any non-specifically bound cells from the capture chamber. An amplifier, e.g., a Zurich Instruments HF2LI dual lock-in amplifier, is used to inject an AC signal, e.g., a 5 V (rms) 1.1 MHz AC signal, into the exit and entrance sensors. Relative impédance is measured using a two-electrode arrangement that is self-referencing in a Wheatstone bridge configuration balanced with resistors and capacitors, e.g., 10 kQ resistors (R) and a 68 pF capacitor.

FIG. 6 provides a doser look at the balancing bridge configuration. When no partie le is passing through the sensing région, the current on both branches is approximately the same, because both electrode impédances are similar and R is equal for both branches. Therefore, VI ~ V2 and V_out is -Ο V. When a cell passes through the sensor région (going from left to right), it will temporarily increase the impédance between the first and middle électrodes, reducing the current in the left branch and decreasing the voltage drop across VI, creatîng a négative puise for V. The cell then passes between the middle and third electrode and conversely causes a positive puise at

V. As a resuit, each cell passage créâtes a down-up (or up-down, dépending on the définition of V_oul) puise pair. This bridge balancing provides several benefits, including providing a baseline signal that varies little with changes to fluid conductivity or flow rate and providing a more sensitive détection method creatîng a largcr impédance puise signal-to-noisc ratio. In addition, one can accurately déterminé whether cells are flowing past the sensor in a forward or reverse direction to ascertain total forward and reverse counts, rcspcctivcly. Puise polarity will reverse when direction reverses. Each cell passage créâtes either an up-down (or down-up) puise signature in the forward flow direction, while in the reverse flow direction, ail cells creatc down-up (or up-down) puise signatures, respectivcly, enabling a straight forward method to differentiate between cells entering and exiting the chip.

The bridge potcntial différence signais for the entrance (Voulut i) and exit (V^^) are input into the amplifier, and the impédance magnitude and phase angle (R and Θ, respectively) are output to a computer for real-time observation and recording of data, e.g., at a 115.2 kHz sampling rate using, for example, Lab-VIEW® software. The data is imported into and analyzed with Clampfit software. Impédance puises can be counted using various threshold levels, and entrance and exit counts are compared. Another computer connected to a digital caméra on a microscope, such as a Nikon Eclipse E600FN microscope (Nikon Instruments, Inc., Melville, NY), can be used to observe cell passage through the channels as well as cellular interactions with the capture région.

Reverse-Flow Differential Counter

Although the shearing unit hclps improve the operation of the differential counter device, another major problem arises in that it has proven difficult to objectively choose the correct trigger level for each counter to provide accurate counts. Ideally, both sensors should hâve the same electrical characteristics and require the use of the samc trigger threshold levels. However, it seems that different threshold levels should be used, but several systematic methods to objectively choose the levels hâve failed (e.g., using triggers based on each electrode’s baseline noise and calculating one counter's trigger level based on the weighted average of the other counter's puise amplitude distribution). This may arise from the possibility that the electrical characteristics of each sensing région are different enough to cause an error in cell énumération. Although micro fabrication may provide entrance and exit counters with almost identical electrode 19 geometries, other factors may cause each sensor to hâve different electrochemical properties. The métal lift-offprocedure may leave nanoscale imperfections that vary from sensor to sensor, creating different field edge effects that may affect a counter's response to cell passage. Non-homogenous métal layer thicknesses from uneven 5 évaporation (sometimes observed by a gradient in color of the métal layer through the entire die) would change the conductivity of the métal leads and the sensing région itself, especially between two counters on a single die that are separated by 34 mm. The connecting micromanipulator probing tips and extemal circuitry may also hâve different electrical characteristics between each branch. Some symptoms from these possible to sources are (1 ) a counter's signal-to-noise ratio does not necessarily scale with its baseline's standard déviation, (2) différences between VO,t_l and Vo„t_2 for two sensors on the same chip, which should be the same, and (3) sometimes slowly changing Vo„t_l or Va,„t2 values over time may point to electrochemical reactions occurring at the electrode-electrolyte interface.

To solve thîs threshold ambiguity problem, a single sensor can be used. FIG. 11 illustrâtes the concept of flowing white blood cells through the entrance of the chip and reversing the flow to push the cells back out the entrance. Cells are injected into the entrance port and flow into the functional ized capture chamber to capture helper T cells. When puises are observed at the exit counter, the fluidic valves are switched to allow

PBS to flow through the chip via the exit port, forcing al! unattached cells to be counted again through the entrance counter. Washîng continues until ali unattached cells are washed from the chip. Because this method only uses the entrance counter to enumerate white blood cells, the problem of finding an objective threshold level is significantly reduced. The threshold can simply be chosen as the minimum level in which baseline noise is not counted as cellular events. The exit counter is used only qualitatively to see when cells hâve filled the capture région volume and to begin the reverse washing process.

The self-referencing sensor aliows for easy discrimination between cells entering and exiting the entrance counter port. For example, depending on the extemal electrical 30 configuration, a cell entering the entrance counter may create an up-down impédance puise pair in time, while the same configuration will create a down-up signature for cells exiting under reverse ilow past the entrance counter port (see, e.g., FIG. 18).

The improved accuracy of using a single electrode set for counting cells flowing into and oui of the capture chamber described above with respect to the reverse flow implémentation can also be provided by a counting de vice in which a portion of the channel defines a flow path that extends in a loop from the first electrical differential counter through the capture chamber and back to the first electrical differential counter as shown in FIG. 17. As discussed above with respect to FIG. 18, the puise shape can be used to détermine when cells are entering the chip and when cells are exiting the chip.

Obtaining Pure Leukocyte Samples from Whole Blood

Red blood cells can be lysed before flowing the cells through the differential counter chip. A lysis solution, e.g., of 0.12 % (v/v) formic acid and 0.05% (w/v) saponin in DI, is used for érythrocyte lysis. A large excess of the lysis solution, e.g., 12 mL of lysis solution, is added to 1 mL of whole blood (drawn the same day and kept on a rotator at room température and încubated for 6 seconds with agitation). Lysis is immediately stopped by the addition of quenching solution (such as 5.3 mL of 0.6% (w/v) sodium carbonate and 3% (w/v) sodium chloride in DI) (see, e.g., D. Holmes, D. Pettigrew, C. Reccius, J. Gwycr, C. van Berkel, J. Holloway, D. Davies, and H. Morgan, Leukocyte analysis and différentiation using high speed microfluidic single cell impédance cytometry, Lab on a Chip, vol. 9, pp. 2881-2889,2009). The solution is centrifuged for 5 min. at 200 x gravity at room température, supematant is aspirated, and pellet resuspended in 5 mL PBS + 1% (w/v) bovine sérum albumin (BSA). The quenching solution is centrifuged for 5 minute at 200 x gravity at room température, supematant is aspirated, and pellet resuspended in 5 mL PBS + 1% (w/v) bovine sérum albumin (BSA). The suspension is centrifuged again and resuspended in 1 mL PBS + 1% BSA, giving the physiological concentration of white blood cells.

In a point of care implémentation of the cell counting device 100, the red blood cell lysis could be performed on chip as described with reference to FIGS. 2A-2E.

Dynamic Threshold Analysis for Objective Enumération of Cells

The impédance signal threshold level is the single most important variable in the electrical énumération of cells in electrical differential counting; fmdîng an objective method to choose the threshold is equally important. By définition, this threshold level détermines whether impédance puises are the entities of interest (cells, beads, etc.), or simply débris, electrical noise, or other entities that should be ignored during analysis. Generally, the threshold level can be based on intégral multiples of the standard déviation of the baseline electrical signal when no cells are passing through the sensor région. In this way, most false positives from electrical noise are excluded when the threshold level is set at or above four to six times the standard déviation of the baseline signal level. However, choosing the threshold level based on electrical signal’s standard déviation alone remains to be a subjective analysis method.

Even a small change in the threshold level can resuit in a large change in cell counts, especially at lower threshold levels. Listed below are some additîonal issues that can render this threshold scaling method impractical, because of large counting errors when performing differential counts; whether using the forward flow method with two counting électrodes (FIG. 5), the reverse-flow method with one counting electrode (FIG. 9), or other implémentations (e.g., FIG. 17).

(1) A cell may not produce the same impédance puise amplitude when passing through the second sensor in a forward flow, two-counter design or when passing back through the entrance counter in a reverse-flow, one-counter design. This introduces counting error because a cell may be counted entering the capture chamber, but not counted when leaving the capture chamber.

(2) The electrical noise level may vary enough during or between analyses to possibly trigger false positive counts if only a static threshold level was chosen.

(3) Débris or small entities (e.g., fragments of dead cells, platelets, etc.) may create impédance puises with amplitudes that exceed the threshold, creating false positives.

(4) The optimal threshold levels may change from chip to chip because of the possible physical and/or electrical différences among fabricated chips. A static threshold levcl for al! chips could resuit in inconsistent measurements that would seriously undermine the advantage of the micro fabricated technology.

The présent solution for the task of objectively choosing a cell counting triggcr threshold is to dynamîcally choose the proper thrcshold level by analyzing the impédance signal(s) with a range of discrète threshold lcvels. During or immediately after blood analysis, differential counts (i.e., entrance count - exit count, or forward count — reverse count) are plotted against their corresponding threshold trigger levels, and the optimal threshold level is chosen based on curve stabïlity (i.e., “flatness”). This method has shown to hâve a low inhérent counting error of ~9 cells- pL'¹ (FIG. 19A, Table 2).

FIGS. 22,23, and 24 illustrate this concept using data from an actual differentia! counting experiment.

First, discrète threshold levels are obtained. One method to create these levels is to obtain the standard déviation of the baseline impédance signal (before cell flow commences) to create a multiplicative standard (i.e., “!x” is the standard déviation). Trigger levels can either be calculated linearly (e.g., multiplication of the !x standard), or through more complex, nonlinear methods. FIG. 22 shows a range of trigger threshold lcvels gcncrated using the linear method, and how the 6x trigger (six times the standard déviation of the noise) encounters the baseline noise signal, which would rcsult in false positive cell counts. The range (e.g., 6x to 20x) and multiplicative values (e.g., 6x, 6.5x, 7x, etc.) can be modified to ensure optimal analysis with proper dynamic range and resolution, respect ively.

Second, the impédance signal(s) are analyzed with the generated range of trigger levels, and différentiel counts are plotted against their respective trigger levels. FIG. 23A shows the variation ofdifferential CD4+T cell counts for a range of threshold trigger levels (6x to 20x). In the idéal situation where each entity’s puise amplitude is identical for entering and exiting the capture chamber, the plot should be a horizontal line, showing that the differential counts are constant for ail trigger levels. However, the smaller threshold levels encounter the signal’s baseline noise level, creating many false positives that statistically conceal the true positives. The sudden increase in differential counts from 6x to 8x illustrâtes this non-ideality, as the 6x threshold level is too low in that it is falsely counting noise peaks as “cells.” The differential count levels off at 8x and romains fiat until 12x, where the counts gradually decreased. This plateau contains the optimal threshold trigger level and corresponding difierential count because it best resembles the idéal horizontal line. Another déviation from the idéal plot is shown by the graduai decrease in the difierential counts at larger trigger levels. This can possibly be explained that the average puise height for the exiting entities is lower than the average puise height for the entering entities (e.g., complication #1, listed above).

Third, the variation of counts between contiguous trigger levels is plotted to further investigate the most stable région of the count vs. trigger level curve. This is analogous to finding the slope of the plot in FIG. 23A, and is shown in FIG. 23B.

Specifically, the slope values (sJ are calculated from Equation 1 :

where c_x is the difierential count and t_x is the trigger level at index x. In this case, x is limited to indices 2 to n, where n is defîned as the total number of trigger levels used for analysis. Index 1 is excluded because, by définition, no slope can be calculated for index

1. Noteworthy: 5, gives the slope immediately before the trigger value at index x. Fourth, the variation in slope values between trigger levels is plotted to make the final stability assessment of the count vs. trigger level curve. This is analogous to finding 20 the curvature of the plot in FIG. 23a (or equally the slope of the plot in FIG. 23b) and is shown in FIG. 23(c). Specifically, the variation in slope values (v_x) between two contiguous triggers is calculated from Equation 2:

(G-'.-.)

In this case, x is limited to indices 3 to n. This is because no slope values exist to calculate the slope variation for indices 1 and 2. Noteworthy: s_x gives the curvature immediately before the trigger value at index x.

Fiflh, average curvature values are obtained for adjacent trigger levels to find the threshold level that is within the most stable régime of the counting analysis curve. The smallest average curvature corresponds to the optimal trigger level. Specifically, the average curvature (¾) for two adjacent curvature values for a trigger level at index x is calculated using Equation 3:

In this case, x is limited to indices 3 to n - 1, as curvature values are not available for indices 1,2, and n.

The aforementîoned methodology to identify the proper trigger threshold level can be succinctly described in the following steps:

1. Generate a range of discrète threshold values (FIG. 22).

2. Obtain differential counts for a range of threshold values, c, (FIG. 23A).

3. Find the count variation vs. trigger level, (FIG. 23B and Equation I).

4. Obtain curvature vs. trigger level, v, (FIG. 23C and Equation 2).

5. Calculate averages of contiguous v, values (Equation 3).

6. Search for the minimum value and note its index, which belongs to the optimal threshold trigger level. The count for this index is chosen to be the actual differential count for diagnostic results.

Table 1 provides the data displayed in FIG. 22A-C and is used to illustrate the dynamic threshold optimization process described above. The average curvature value at index 4 (a₄) corresponds to a 12x trigger level, resulting in a differential count of 1,804 CD4+ cells (sélection highlighted in FIG. 23C). FIG. 24 shows the cumulative forward and reverse counts found using a 12x trigger level for the duration of the experiment.

Table 1

Index ω	Trigger Level (x Ix standard)	Differential Count (ex)	S tope (s,)	Curvature O’x)	Avg. Curv. (¾)
1	6	-7	n/a	n/a	n/a

2	8	1810	908.5	n/a	n/a
3	10	1759	-25.5	-467.0	245.5
4	12	1804	22.5	24.0	31.5
5	14	1693	-55.5	-39.0	35.4
6	16	1455	-119.0	-31.8	28.9
7	18	1113	-171.0	-26.0	19.4
8 (H)	20	720	-196.5	-12.8	n/a

This dynamic threshold analysis method has been shown to provide counts which correlate closely (y = 0.994x, R² = 0.997) with an optical énumération method (FIG. 20A). This shows it to be a feasible method for the automatic énumération of particles and cells using an electrical differential counting technique. FIG. 15 is an area histogram of circular objects on a chip as observed using the optical counting method. FIG. 20B shows Bland-Altman comparison analysis between the electrical differential and optical counting methods. A bias ofonly about 9 cells confirms the accuracy of the electrical differential counting method for the entire range of enumerated CD4+ T cells.

The aforementioned methods do not limit the scope of the dynamic threshold analysis method, but serve as an example to prove its feasibility and efficacy. The following are additional notes regarding other implémentations of the dynamic threshold analysis method. First, integer multiples were used to generate discrète threshold values, but fractions of whole numbers can be used as well (e.g., 4.25x). Second, plotting the different data (e* s_x, v_Xt a_x) is not necessary, but was used for illustrative purposes. The operating device’s microcontroller or microprocessor would only need the raw differential counting data (c_x) to calculate the average curvature values (a_x). Third, analysis is not limited to Equations 1 - 3, as other implémentations may be used to find the optimal thresholds more efficiently and/or effectively. Fourth, nonlincar methods can be used to generate threshold levels in addition to the linear method used in the above example. Fifth, threshold analysis is not limited to puise amplitude (or height), but can be used on other variables, such as puise width, puise area, or other implémentations. Sixth, threshold analysis is not limited to puises with positive polarity, but can also be used for negative-going puises. Seventh, the number of and spacing between threshold levels can be adjusted to provide a more accurate rendering of the threshold level vs. differential count plot to locate the optimal threshold level with higher précision.

Cell Counting Devices with Lvsis and Quenching Régions

In some implémentations, an on-chip lysis région, e.g., a red blood cell lysis région, can be included in the counting device, e.g., a CD4+ T cell counting device. The addition of the lysis région can eliminate requirements for additional laboratory equipment and personnel that are needed to lyse the red blood cells off-chip, enhancing the portability of the device. For example, FIG. 25 is a schematic that illustrâtes a CD4+ T cell counting device 2500 that incorporâtes a cell lysing région 2502 (e.g., for lysing red blood cells). During operation ofthe device 2500, whole blood flows into the chip and is surrounded by a lysis solution, which mixes in the serpentine mixing channels 2504 and rapidly ruptures the red blood cells within about 6 to about 10 seconds. Different conditions can be used to lyse other types of particles, e.g., cells. To ensure lysis during a desired time period, the volume of the lysis région channels and the flow rate of the lysis and sample, e.g., blood, solutions can be controlled. For example, the lysis region's channel width can range from about 50 pm to about 1 mm and height can be from about 10 pm to about 400 pm with lysis and blood solutions combined flow rates ranging from about l pL/minute to about 100 pL/minute.

Lysis is rapidly stopped to preserve the remaining cells, such as white blood cells, by the addition of a quenching solution and quench duration is extended via serpentine mixing channels 2506 to ensure quenching of the lysis process, which should hâve a duration of greater than about 10 seconds. The quenching channel dimensions and the combined flow rates of the lysing, blood, and quenching solutions can be controlled to ensure quenching duration is above this minimum. For example, the quenching channel dimensions can be formed to be similar to the lysis région channels and the combined flow rates of the lysis, blood, and quenching solutions can range from about 1 pL/minute to about 1000 pL/minute. The quenched solution then flows through a filter 2508 comprised of pores to prevent possible clogging of the counting pore having the same dimensions as the filter pores. The filter and counting pores can range in size from a height and width each of about 0.5 pm to about 50 pm.

The sensing électrodes of the counter 2510 can be made of a conduction layer of either platinum or gold or other high conductivity métal with an adhesion layer (optional) of chromium or titanium. The sensing électrodes can hâve widths and gaps ranging from less than about 1 pm to about 1 mm. The Coulter principle can be employed to electrically count cells individually by observing the temporal impédance changes (i.e., electrical puises). White blood cells then pass through an identical filter before being dis tribu ted among eight identical capture chambers 2516, which can be from 10 pm to 100 pm high and 0.5 mm to 10 mm wide. The number of capture chambers 2516 can vary from 1 to over 32. Capture chamber height can be tailored to control the shear stresses at the fluid/chamber wall interface for optimal capture of CD4+ T cells or other cells/particles of interest.

The devices can be made with a glass substrate (with micro-pattemed platinum or gold électrodes) bonded to PDMS (polydimethylsiloxane) fluidics via oxygen plasma treatment. Another method uses plastics for the substrate and fluidics (e.g., injection molding) with the sensing électrodes defîned by laser ablation or similar processes.

Cell Counting Devices That Distinguish Between Different Types of Cells

In some implémentations, the cell counting devices can differentiate between different types of white blood cells, red blood cells, and platelets based solely on using multiple interrogation frequencies. This technique enables counts of red blood cells, platelets, and white blood cell subtypes (monocytes, neutrophils, lymphocytes, etc.) in addition to the spécifie énumération of CD4+ T cells using the antibody-coated capture chamber, as already described. For example, referring to FIG. 2, multiple signais of different frequencies can be applied simultaneously to one or more of the impédance sensors 110, providing a discrète impédance spectrum for any particular cell type.

Cells can be differentiated based on their different impédance spectra. For example, Holmes et al. used a 503 kHz frequency to obtain the volume of each cell, but also used a higher frequency (1.7 MHz) to simultaneously inspect a cell's membrane capacitance. They were able to differentiate among some of the different white blood cell subsets (monocytes, neutrophils, and T-lymphocytes) via observing the opacity of a cell (high frequency impédance divided by the low frequency impédance) with the assistance of a red blood cell lysis solution (see Holmes et a!., “Leukocyte analysis and différentiation using high speed microfluidic single cell impédance cytometry,” Lab on a

Chip, 2009, 9,2881-2889; see also Ledis et al., “Lysing reagent System for isolation, identification and/or analysis of leukocytes from whole blood samples, U.S. Patent No. 5,155,044, October 1992). In addition, Cheung et al. used a 6 MHz frequency to differentiate between red blood cells and white blood cells (see, Chreung et al. “Microfluidic Impedance-Based Flow Cytometry, Cytometry: Part A, 2010,77A, 648666).

Accordîngly, a low frequency (e.g., from about 1 kHz up to about I MHz) can be applied to the impédance sensors 110 to obtain a cell's volume and additional higher frequencies (e.g., from about 1 MHz to over 100 MHz) can be applied to the impédance sensors 110 to provide a discrète impédance spectrum for differentiating among several cell types. The more discrète frequencies used, the higher the resolution to differentiate between different cell types that can be indistinguishable at a smaller number of interrogation frequencies used. In particular, platelets can be discriminated among other cell types based simply on their size, as they are approximately I to 2 pm in size—much smaller than other cell types. As a result a low frequency measurement alone can differentiate platelets from other cell types. Red blood cells can be distinguished from white blood cells using a low frequency (500 kHz) and a hîgh frequency (6 MHz), as red blood cells hâve a similar volume to the smaller white blood cells. In some implémentations, different white blood cell types may require one or more frequencies in addition to the low frequency (500 kHz) for différentiation among the white blood cell subtypes.

EXAMPLES

The following examples are illustrative and not limiting.

Testing Maximum Puise Densitv Limits

It is desired that the differential counter can enumerate the physiological concentration of white blood cells flowing at the desired range of 5-10 pL/min to provide a rapid helper T cell count. As the concentration of cells increases with a constant flow rate, the amount of average volume (and time) decreases between events (i.e., puises caused by cell passage through the sensing région). Eventually, the concentration becomes high enough where two cells will be in the same sensing région, creating coïncident events that reduce the accuracy of the counter. In addition, for a finite sampling frequency, even if the cells are not coïncident in the sensing région, a high enough velocity wili eventually cause overlap of the puises from two subséquent cell passages.

Diluted whole blood was used to test the puise density limits of the differential counter, because it contains an abundance of flexible particles, as opposed to polystyrène and latex beads, which hâve been prone to clog the counting channel. A constant flow rate of 5 pL/min was used to inject varying dilutions ( 1:1000 to 1:100) of whole blood into the chip. Puises were only analyzed for the entrance counter. Puise density was calculated by énumération of puises in known duration Windows at random times throughout the raw data.

FIG. 7 illustrâtes the results as a comparison between the cell concentration found using the microfluidic chip (calculated by the number of puises for a known volume flown) compared to the calculated concentration of each dilution (assuming a whole blood concentration of 5 x IO⁹ cells/mL). At a 5 pL /minute flow rate, the microfluidic chip could handle the 1:200 dilution of whole blood (-2.5 x 10⁷ cells/mL), but failed to count every puise for the 1:100 dilution (-5 x 10⁷ cells/mL). The maximum puise density the chip could handle was 2,236 cells/s, équivalent to a concentration of 2.68 x 10⁷ cells/mL at a flow rate of 5 pL/minute. This is well above the upper limit of leukocyte concentration in healthy adults, ensuring no coïncident events, even at a flow rate of 10 pL /minute.

Testing Capture Chamber Sensitivitv and Accuracy

The next expert ments were done to verify that the entrance count is the same as the exit count for a passivated capture chamber. A 10 pL sample of healthy adult blood (with lysed érythrocytes) can hâve over 100,000 leukocytes, in which 10,000, or 10%, are helper T cells. A patient with AIDS can hâve helper T cell counts less than 200 cells/ pL, which results in only 2,000 cells per 10 pL, or 2% of total leukocytes. Any errors in counting can negatively affect the sensitivity and accuracy of this method.

Before cells were flowed into the microfluidic chip, the capture chamber was passivated by flowing in PBS + 1% BSA and incubatîng for 30 minutes at room température to prevent the non-specific adsorption of cells to the glass and PDMS surfaces. BSA is a well-known protein for surface passivation, and readily binds to the hydrophilic glass substrate at pH 7.4 (see e.g., Sweryda-Krawiec et al., A new interprétation of sérum albumin surface passivation, Langmuir, vol. 20, pp. 2054-2056, September 2004). In this particular experiment, three dilutions of white blood cells were flown into the chip at 5 pL /minute, followed by a 10 pL /min PBS + 1 % BSA wash to ensure ail cells exit through the exit counter. Impédance data for each counter is recorded during the entire experiment.

FIG. 8 illustrâtes a typical resuit for the négative control experiment. Ideally, the entrance and exit sums should be équivalent at the end of the experiment, but hâve a différence of over -3,500. It was interesting to note that the exit count was higher than the entrance count, which is true for the 1:1 dilution of white blood cells, but not as dominant in the lower dilutions. FIG. 9 shows the relationshîp between the white blood cell concentration and the différence between the exit and entrance counts for various trigger levels. A trigger level is the voltage threshold that détermines whether an impédance puise is a cell, and is set manually in Clampfit. It is a common convention to base the trigger level on the standard déviation of the baseline signal's noise (with no cells présent).

In this experiment, a trigger level of ten times the standard déviation (SD) of the noise was the minimum threshold that could be used to ensure baseline noise puises were not counted as cellular events. The threshold level for the entrance and exit counters was identical. A noticeable trend is the less diluted samples intersect the X-axis (Entrance Exit = 0) at higher threshold values (67 x SD for 1:1 ; 40 x SD for 1:2; 20 x SD for 1:5) in the direction of increasing trigger level value (left to right). This, combined with the fact that the exit count is higher than the entrance count, can explain the large discrepancy in the entrance and exit counts. Cell aggregates form more frequently as the concentration of the purified leukocytes increases, because there is more interaction between cell surfaces. These aggregates pass through the entrance counter port and its relatively high shear stresses (1,320 dyne/cm) separate the aggregates back into individual cells, which are then counted by the exit counter. An aggregate is counted as a single entity by the entrance counter, but can become three or more entities by the time it reaches the exit counter. The entrance and exit counts only become equal when the threshold level is large enough to not count smaller entitîes such as single cells, and only counts larger objects that remain physically intact after passing through the entrance counter.

The aggregation of leukocytes prevents a true évaluation of the differential counter and can be remedied by larger dilutions. However, diluting has several drawbacks, most importantly, analyzing only a fraction of the cells needed to provide a more robust helper T cell test and requiring a much larger chip volume. Therefore, it is désirable to hâve physîological concentration of white blood cells enter the chip, and can possibly still be allowed using a micro fabricated 10 pm x 13 pm PDMS/glass pore, or shearer, to separate cell aggregates before the chip entrance. FIG. 10A shows the resultsafter repeating the passivated experiment for kl diluted leukocytes. The shearer proves to decrease the number of aggregates before entering the differential counter chip (X-intercept at 9 x SD vs. 12,5 x SD for cell samples injected directly into the chip without the shearer).

FIG. 10B shows the différence in cell size (puise amplitude) and cell passage duration when using the shearer. The population undergoing shear before making it to the entrance counter is a tighter distribution at lower puise duration with similar puise height amplitude as the un-sheared population because the larger aggregates block the impédance sensing région longer. The amplitude does not change much because even the single cells are large enough to block most of the electrical current passing between the sensing électrodes.

FIG. I0C illustrâtes the size and passage duration sîmilarities of cells that hâve been sheared prior to and counted at the entrance sensor and cells that did not undergo pre-chip shearing, but pass through the entrance counter pore and are counted at the exit counter. This shows that the entrance counter ïndeed îs shearing aggregates into smaller entitîes, performing the same job as the pre-chip shearer. It is therefore necessary to hâve the shearing unit placed before the chip to ensure most aggregates are separated into single cells.

Testing a Reverse-Flow Differential Counter

The passivated capture chamber experiments noted above were repeated using the reverse-flow protocol, and the results of fourteen different experiments are shown in FIG. 12 and FIG. I9A. The forward count is équivalent to the number of leukocytes that entered the capture chamber during forward flow; the absolute error count is the différence from the idéal di fferential count of zéro; the percent error is the absolute error count normalized by the forward count. FIG. 12 shows how the absolute error count remained roughly constant for the entire forward counting range. This resulted in a decreased percent error for larger forward counts (FIG. 19A), which is désirable. FIG. 13 is a graph presenting entrance and exit counts for a passivated capture chamber experiment using the reverse-flow differential counter protocol.

FIG. 19B illustrâtes the accumulated forward and reverse counts during the experiment highlighted in FIG. 19A. This demonstrates how the reverse count eventually leveled off and became close to the forward count. As Table 2 below shows, forward counts greater than 2,000 resulted in a much smaller error. This ensures that larger leukocyte numbers—found in clinical situations—will resuit in the lowest error. The decreasîng % error for increasing total forward cell counts can be explained by the fact that the counting errors do not scale with the total number of cells flown, and remain relatively constant.

Table 2

Data Range	Error (%) (Fig. 12 inset)	Abs. Counting Error (cells) (Fig. 12)	Est Sensitivity (cells-pL*1)
X	SD	X	SD	X	SD
AI1WBC	2.91	3.93	44.2	31.3	8.84	6.26
WBC < 2000	7.25	5.37	38.8	25	7.76	5
WB0 2000	1.18	1.02	46.4	34.5	9.28	6.9

Table 2 summarizes the data from FIG. 12 and FIG. 19A for different ranges of total white blood cells counted. The estimated sensitivity can be obtained by assumîng approximately 5 pL of sample was flown into the chip (approximate because current metering methods are in need of improvement). As a resuit, base sensitivity is -9 cells/pL for the more realistic range of greater than 2,000 white blood cells counted, which is similar to the best sensitivity in electrical CD4+ T cell counts in the literature (Cheng et al., “Cell détection counting through cell lysate impédance spectroscopy in microfluidic devîces, Lab on a Chip, vol. 7, pp. 746-755,2007). The main source of counting errors was caused by non-specific adsorption of cells onto the chamber surface, despite passivation with BSA. A more successful passivation using more incubation time and/or PBS with a pH doser to BSA’s isoelectric point of 5 would substantially decrease this error and illustrate that the difierential counting method would provide the most sensitive énumération technique (Freeman et al., “Real time, high resolution studies of proteîn adsorption and structure at the solid-liquid interface using dual polarization interfèremetry,” Journal of Physics: Condensed Matter, vol. 16, pp. S2493-S2496,2004). Another possible source of error may be dead/dying cells rupturing under the high shear rates found in the counter channel after forward counting.

Enumération of CD4+ T Cells Using the Reverse-Flow Technique

The reverse-flow technique was used to electrically enumerate the number of CD4+ T cells captured on a microfluidic chip. The capture région was first coated with an anti-CD4 antibody (Ab)(l:10 in PBS) by adsorption for 30 minutes, followed by several itérations of flowing in more Ab and waiting 10 minutes between each itération. Unbound Ab was removed by rinsing the chamber with PBS + 1% BSA, which also passivates any surface which does not hâve Ab adsorbed to it. White blood cells were flown into the chip at 5 pL /minute until cells were electrically detected at the exit counter. PBS + 1% BSA was then infused through the exit counter port initially at 5 pL /minute to increase the interaction time between the helper T cells and the CD4 Ab. The washing flow rate was încreased to 10 pL /minute after most cells had exited the chip to wash away any non-specifically bound cells.

After electrical counting, an optical control was obtained by imaging the captured cells for subséquent énumération using image processîng software. Phase contrast images of the entîre capture région were taken using an Olympus 1X81 inverted microscope at 40x total magnification. The 42 images were aligned and merged using Adobe Photoshop image processing software, and cells were counted using ImageJ software. FIG. 14 shows the merged images and résultant image of the entire capture and counter régions. It was found that the highest density ofcaptured cells was found before the midpoint of the capture chamber's length (inset 2). A smaller density of cells were found near the inlet (inset 1), which is expected since the cells hâve not had enough time to interact with the Ab on the chamber surface. The lowest density is found near the exit 34 of the chamber, where very few cells are attached (inset 3). Most likely the washing process began before the higher concentration of cells made it to the exit, but could also be because the majority of the helper T cells had ample time to bind to the immobilized CD4 Ab.

It was also noted that the cell path does not span the entire width of the capture channel. This results because the relatively narrow counter channel acts as a highlyfocused nozzle which causes most of the cells to travel within ± 850 pm of the centerline of the channel's length. This can be resolved by placing the entrance and exit counters diagonally opposite of each other (in opposite corners of the capture chamber), which would force the cells to travel the diagonal length of the capture chamber. Another solution may simply be found by curving or fanning the counter outlets so that the cells will not be as focused once entering the capture chamber.

FIG.l5 shows the automated counting of circular objects of various internai areas. The dotted lines dénoté the range of areas assumed for the helper T cells and gives a helper T cell count of 926. This range encompasses cell diameters from 10 to 12.5 pm, which is somewhat larger than the diameter of lymphocytes reported in the literature, but these cells are not in optimal physiological conditions and may hâve initiated apoptosis. Also, the phase contrast imaging créâtes a halo around the cell diameter, which could cause an apparently larger cell, especially when taken at a low magnification, where the size of the pixels are relatively larger and may not create an accurate représentation of the cell's perimeter.

FIG. 16 shows the results of the reverse-flow differential counting of captured helper T cells. The obtained count of 931 cells closely matches the count found by image processing, and shows that the differential counter method is viable method of enumerating helper T cells in a microfluidic chip.

FIG. 20 shows results from 14 CD4+ T cell counting experiments using white blood cells purified from human whole blood samples and the close corrélation (y = 0.994x, R² = 0.997) between the electrical differential method and the optical control. FIG. 21 illustrâtes how the percent error (absolute différence in optical and electrical counts, normalized by the CD4+ T cell count) relates to the total number of CD4+ T cells counted. For less than 1,000 cells captured on the chip, the average error is 4.5% (n = 3).

Assuming a 5 pL sample volume, this would be for CD4+ T cell concentrations less than 200 cells/pL, the concentration lîmit which defines AIDS. This shows to be highly accurate, as a patient with an actual CD4+ T cell concentration of 100 cells· pL'¹ would hâve a counting error of only+/- 4.5 cells/pL. For counts above 1,000 cells captured in the entire chip, the average error is 2.1% (η = 11). The 25+/-10% (n - 14) ratio of captured cells to total cells counted agréés with the literature conceming the 25-33% of leukocytes being CD4+ T cells (Daniels et al., “Functional histology: A text and colour Atlas,” Churchill Livingstone, 1979).

Cell Counting Using Device With Lysine and Ouenching Régions

Experiments were set up to evaluate the reverse electrical differential counting method with the additional red blood cell lysing and quenching régions to ensure its feasibility in diagnostics testing using the device 2500 described above. The chip’s capture régions and exit holding coil were passivated from cellular interactions using a 1% BSA (bovine sérum albumin) solution in PBS (pH 4.5) for three hours. The holding coil was used to ensure cells exiting the chip during forward flow direction would not be lost to waste before flow reversai. Various sample sizes of whole blood (0.5 to 10 pL) were injected into the chip at a flow rate of 1.5 pL/min. The lysing solution (0.12% (v/v) formic acid and 0.05% (w/v) saponin) and quenching solution (2x PBS and 0.6% sodium carbonate) were infused at 17.5 pL/min and 8.5 pL/min, respectively, using an HPLC pump. Flow was reversed once the desired blood volume was injected and the experiment duration ended when cells were completely washed from the chip and holding coil.

FIG. 26 illustrâtes the percent error of twenty-three differential cell counts for whole blood s amples. The percent error is calculated as the absolute différence between the forward and reverse counts, normalized by the forward count, and multîplied by 100. Ideally, the forward and reverse counts would be identical, resulting in a percent error of 0%. The average percent error for ail twenty-three experiments was about 3.3%, which is similar to the percent error of about 2.9% in the previous implémentation that did not hâve a red blood cell Iysis and quenching module (Table 2). This shows that the differential counting chip with the addition of the red blood cell lysis and quenching modules results in a feasible device that can analyze unprocessed whole blood samples with low inhérent error-making it practica! for the use as a portable diagnostic device.

OTHER IMPLEMENTATIONS

It is to be understood that while the invention has been described in conjunction with the detailed description thereof, the foregoing description is intended to illustrate and not limit the scope of the invention, which is defîned by the scope of the appended daims. Other aspects, advantages, and modifications are within the scope of the following daims.

Claims (23)

1. A method of counting particles of interest in a sample that comprises two or more different types of particles, the method comprising:

obtaining a fluid sample that may contain particles of interest;

counting ail types of particles in a portion of the sample using a first electrical differential counter to generate a first total;

removing any particles of interest from the portion of the fluid sample;

counting any particles remaining in the portion of the fluid sample using a second electrical differential counter after the particles of interest are removed to generate a second total; and calculating a number of particles of interest originatly in the fluid sample by subtracting the second total from the first total, wherein the différence is the number of particles of interest in the sample.

2. The method ofclaim 1, wherein the first and second electrical differential counters are the same electrical differential counter.

3. The method of claim 2, further comprising reversing a flow direction of the fluid sample after removing the particles of interest from the portion of the fluid sample.

4. The method of claim 2, further comprising maintaining a flow direction of the fluid sample while counting ail types of particles in the portion of the sample; removing particles of interest; and counting any particles remaining in the portion of the fluid sample.

5. The method of claim 1, wherein the particles of interest are cells.

6. The method of claim 1, further comprising depleting selected particles from the portion ofthe sample before counting ail types ofparticles in the portion of the sample.

7. The method of claim 6, wherein the fluid sample comprises whole blood and the method comprises depleting érythrocytes by lysis.

8. The method of claim 6, wherein the fluid sample comprises whole blood and the method comprises depleting érythrocytes, monocytes, neutrophils, CD8+ lymphocytes, or other cellular components of blood by immunodepletion.

9. The method of claim 1, wherein the particles of interest are CD4+ T cells, and removing the particles of interest comprises capturing CD4+ T cells in a capture chamber functionalized with anti-CD4 antibodies.

10. The method of claim 9, further comprising removing non-specifically adsorbed leukocytes by purging the capture chamber with phosphate bufïered saline.

11. A device comprising:

a microfluidic chip defining a channel including an inlet and an outlet;

a capture chamber arranged along the channel between the inlet and the outlet, wherein the chamber is configured to capture particles of interest from fluid flowing through the channel;

a first electrical differential counter arranged to count ail types of particles in a fluid flowing into the capture chamber;

a second electrical differential counter arranged to count ail types of particles remaining in the fluid flowing out of the capture chamber; and a computing mechanism in electronic communication with the first and second electrical differential counters, wherein the computing mechanism calculâtes a number of particles of interest based on signais from the first and second electrical differential counters.

12. The device of claim 11, wherein the first and second electrical differential counters are the same electrical difTerential counter.

13. The device of claim 12, further comprising a pump System in fluid communication with the channel, wherein the pump System is opérable in a first mode to cause fluid to flow in a first direction in the channel past the first electrical differential counter and opérable in a second mode to cause fluid to flow in a second direction in the channel opposite the first direction and back to the first electrical differential counter.

14. The device of claim 12, wherein a portion of the channel defines a flow path that extends in a loop from the first electrical difierential counter through the capture chamber and back to the first electrical difierential counter.

15. The device of claim 1, wherein the capture chamber includes surfaces functionalized with anti-CD4 antibodies.

16. A kit comprising:

a device of claim 11;

a solution comprising a binding moiety with an affinity for the particles of interest; and a solution comprising a lysing agent effective to lyse selected particles without lysing the particles of interest.

17. The kit ofclaim 16, wherein the first and second electrical difierential counters are the same electrical difierential counter.

18. The kit of claim 17, comprising a pump System in fluid communication with the channel, wherein the pump System is opérable in a first mode to cause fluid to flow in a first direction in the channel past the first electrical difierential counter and opérable in a second mode to cause fluid to flow in a second direction in the channel opposite the first direction back to the first electrical difierential counter.

19. The kit of claim 18, wherein a portion of the channel defines a flow path that extends in a loop from the first electrical difierential counter through the capture chamber and back to the first electrical difierential counter.

20. The kit of claim 16, wherein the binding moiety comprises anti-CD4 antibodies.

21. A microfluidic chip comprising:

a plurality of capture chambers, the capture chambers being configured to capture particles of interest from fluid flowing through the chambers;

an electrical difierential counter opérable to count particles in a mixed population of particles in fluid flowing into the capture chamber and to count particles remaining in fluid flowing out of the capture chamber; and a computing mechanism in electronic communication with the electrical differential counter, the computing mechanism opérable to calculate a number of particles 5 of interest based on signais from the electrical differential counter.

22. The microfluidic device of claim 21, further comprising:

a fluidic channel coupled to the plurality of capture chambers, wherein the fluidic channel includes a first channel région and a second channel région, wherein the first channel région is configured to receive a lysing solution and a sample fluid, and mix the 10 sample fluid with the lysing solution, and the second channel région is configured to receive a quenching solution and a lysed solution from the first channel région, and mix the quenching solution with the lysed solution.

23. The method of claim l, furthercomprisingdetermining a cell flow direction based on a polarity of an impulse signal generated by the first electrical differential

15 counter.