Microfluidic particle-labeled impedance sensor array for enhancing bioassay sensitivity

Abstract

Described herein are devices and methods that provide improvements in the detection and quantification of biomolecules. The devices and methods may include a microfluidic biosensing platform for measuring electrical impedance changes in the presence of target analytes from a sinusoidal input voltage to improve the sensitivity of the immunoassay. In one embodiment, the described devices and methods may provide a quantitative diagnostic point-of-care immunoassay platform that uses signal-enhancing microparticles to improve sensitivity and detection limits.

Description

[Technical Field]

[0001] Cross-reference of related applications This application claims priority to U.S. Provisional Patent Application No. 63 / 160,594, filed on 12 March 2021, which is incorporated herein by reference as is. Research funded by the federal government This invention was made with government support under authorization number NSF-EECS / 1509746 granted by the National Science Foundation. The government has certain rights in this invention.

[0002] This specification describes devices and methods that have been improved in the detection and quantification of biomolecules. The devices and methods may include a microfluidic biosensing platform for measuring changes in electrical impedance in the presence of a target analyte from a sinusoidal input voltage to improve the sensitivity of the immunoassay. In one embodiment, the described devices and methods may provide a point-of-care immunoassay platform for quantitative diagnostics that uses signal-enhanced microparticles to improve sensitivity and detection limits. [Background technology]

[0003] Impedance-measuring biosensors are a class of electrochemical biosensors that measure changes in electrical impedance caused by the presence of biomolecules, cells, or labeled biomaterials on the working electrode by applying a sinusoidal voltage. Due to their low cost, ease of miniaturization, multiplexing capability, and label-free operation, these types of sensors have proven promising for digitized point-of-care (POC) diagnostics. To date, considerable effort has been made, in conjunction with the development of miniaturized platforms with reduced complexity. Previous research has led to the development of an integrated cell-counting assay system for malaria diagnosis, consisting of a microfluidic chip and a miniature impedance circuit board. Another integrated diagnostic platform, consisting of a printed gold electrode chip and a microfluidic flow cell, has been proposed for the detection of the transgenic protein Cry1Ab. For a miniaturized portable impedance-measuring biosensor platform, an impedance-measuring reader based on the AD5933 chip has been designed using microfluidic channels on an IDE array for the diagnosis of deep vein thrombosis and pulmonary embolism. Despite considerable effort in developing miniaturized impedance biosensors, few integrated platforms that are easy to use and possess sufficient sensitivity for actual immunoassays have been developed.

[0004] Capillary microfluidic technology has been used to perform different types of immunoassays on different platforms. In one of the earliest studies, the potential of on-chip immunoassays using the capillary filling phenomenon was investigated. To obtain an autonomous fluid pathway, a capillary holding valve was introduced to allow sequential delivery and multiple cycles. By controlling the geometric shape of the channel, and therefore the hydrodynamic resistance, a one-step capillary-driven microfluidic system was developed that allowed for flow velocity control and variation of incubation time in the reaction chamber. This compact device demonstrated a quantitative immunoassay for C-reactive protein (CRP), with the lowest detectable fluorescence signal corresponding to CRP of 10 pg / mL. While such platforms paved the way for autonomous microfluidic chips to perform highly sensitive immunoassays, their low sensitivity and the need for large, bulky downstream sensors to read the output signal hindered their development into commercially viable products. [Overview of the project] [Problems that the invention aims to solve]

[0005] Therefore, novel, integrated, user-friendly methods and devices are needed to improve the detection and quantification of various disease-related biomarkers. Such methods and devices would be useful in a variety of point-of-care (POC) diagnostic applications. [Means for solving the problem]

[0006] One embodiment described herein is a microfluidic device for detecting and quantifying biomolecules, comprising: (a) a multilayer microfluidic network including a microfluidic channel, a buffer inlet, a sample inlet, and a waste outlet, wherein the buffer inlet, the sample inlet, and the waste outlet are in fluid communication with each other via the microfluidic channel, and the microfluidic network is configured to accept a buffer solution and a sample; and a microfluidic chip including a substrate layer on which a first antibody is covalently bound, which is adapted so that the first antibody specifically binds to a target analyte, and the first antibody is mounted between a first electrode and a second electrode, wherein the substrate layer is in fluid communication with a microfluidic channel, and the first antibody is located within the microfluidic channel; and (b) a detector for detecting changes in electrical impedance.

[0007] In one embodiment, the microfluidic network is composed of a polymer material selected from one of the following: polycarbonate (PC), polymethyl methacrylate (PMMA), cyclic olefin copolymer (COC), polyimide, polydimethylsiloxane (PDMS), or a combination thereof. In another embodiment, the microfluidic network is composed of PDMS.

[0008] In another embodiment, the microfluidic network has geometric dimensions for autonomous capillary flow of buffer solution and sample. In another embodiment, the microfluidic network comprises a chamber layer, a capillary valve, and a bridging hole located above the microfluidic channel, the chamber layer and the microfluidic channel being fluidly connected to each other via the capillary valve and the bridging hole, and the sample inlet being located within the chamber layer. In another embodiment, the chamber layer contains a second antibody, which is adapted to specifically bind to a target analyte, and the second antibody is conjugated to a microparticle. In another embodiment, the second antibody contains a biotin moiety, the microparticle contains a streptavidin coating, and the second antibody is conjugated to the microparticle through the binding of the biotin moiety to the streptavidin coating. In another embodiment, the microparticles include magnetic beads, polystyrene beads, silica beads, or a combination thereof. In another embodiment, the microparticles have a size in the range of approximately 1 μm to approximately 5 μm in diameter. In another embodiment, the microparticles include magnetic beads having a size of approximately 2.8 μm in diameter. In another embodiment, the chamber layer further includes a porous polycarbonate (PC) membrane connected to the sample inlet, and the microparticles conjugated to the second antibody are immobilized on the porous PC membrane.

[0009] In another embodiment, the capillary valve comprises an orifice having a diameter in the range of approximately 100 μm to approximately 300 μm. In another embodiment, the capillary valve comprises an orifice having a diameter of approximately 250 μm. In another embodiment, the bridging hole comprises an orifice having a diameter in the range of approximately 0.5 mm to approximately 2.5 mm. In yet another embodiment, the bridging hole comprises an orifice having a diameter of approximately 1 mm.

[0010] In another embodiment, the buffer solution contains phosphate-buffered saline (PBS) at a concentration ranging from about 0.001 mM to about 1 mM. In yet another embodiment, the buffer solution contains PBS at a concentration of about 0.01 mM. In yet another embodiment, the buffer solution further contains bovine serum albumin (BSA) at about 1 wt%.

[0011] In another embodiment, the microfluidic network further includes one or more absorbent pads. In another embodiment, the sample inlet further includes a serum separation membrane. In another embodiment, the substrate layer includes a glass substrate or a plastic substrate. In another embodiment, the first and second electrodes are coated with a conductive metal. In another embodiment, the conductive metal is selected from gold (Au), titanium (Ti), or a combination thereof. In another embodiment, the distance between the first and second electrodes is approximately 1 μm to approximately 10 μm. In another embodiment, the distance between the first and second electrodes is approximately 10 μm. In another embodiment, the first and second electrodes operate at a frequency in the range of approximately 1 kHz to approximately 100 kHz. In another embodiment, the first and second electrodes operate at a frequency of approximately 10 kHz. In another embodiment, the first and second electrodes are part of a plurality of electrodes, which are combined with each other.

[0012] Another embodiment described herein is a method for detecting and measuring the presence of a target analyte in a sample using any one of the embodiments or aspects of the disclosed device, comprising: (a) loading a buffer solution into a buffer inlet; (b) flowing the buffer solution onto a substrate layer; (c) loading a sample into a sample inlet; (d) mixing the buffer solution with the sample; (e) sequentially flowing the mixture of the buffer solution and the sample onto the substrate layer, wherein the target analyte binds to a first antibody; (f) continuously flowing the buffer solution onto the substrate layer to remove any unbound target analyte; and (g) detecting a change in electrical impedance to quantify the concentration of the target analyte in the sample.

[0013] In one aspect, the flowing step is autonomous by capillary action. In another aspect, the sample is incubated with microparticles conjugated to a second antibody, then mixed with a buffer solution, and the target analyte binds to the second antibody conjugated to the microparticles. In another aspect, the microparticles conjugated to the second antibody are immobilized on a porous PC membrane. After loading the sample into the sample inlet, when the sample is dropped onto the porous PC membrane, the microparticles are released from the membrane, and the sample is incubated with the released microparticles conjugated to the second antibody. In another aspect, the sample is loaded into the sample inlet after the buffer solution reaches the capillary valve. In another aspect, the capillary valve opens when the sample contacts the buffer solution.

[0014] In another aspect, the sample includes a blood sample or other biological liquid sample derived from a subject. In another aspect, the method includes a total assay time in the range of about 5 minutes to about 10 minutes.

[0015] The patent or application file includes at least one drawing in color. Copies of this patent or patent application publication, including the color drawing(s), will be provided by the Patent Office or the like upon request and payment of the necessary fees.

Brief Description of the Drawings

[0016] [Figure 1] It is a diagram showing a schematic view of an exemplary microfluidic impedance sensor array for a microparticle-labeled immunoassay. It shows eight mutually combined electrodes (IDEs) integrated with a special specification impedance analyzer and a microfluidic channel (left). It shows an illustrative diagram of the IDEs for a negative control (upper) and a positive control (lower), and represents the formation of a complete immune complex on the IDE (right). [Figure 2A]FIG. 0 shows an exemplary system configuration consisting of a gold (Au) IDE array chip, two impedance analysis circuits, and a data acquisition (DAQ) board for signal transmission related to a LabVIEW software program. A microfluidic channel network can be placed on top of the IDE array to deliver analytes and buffer solutions and to control fluid forces for washing purposes. [Figure 2B] FIG. 3 shows an overview of an exemplary IDE chip fabrication process. [Figure 2C] FIG. 6 shows an illustration of an IDE and an equivalent circuit, and an equation showing how the IDE measures impedance variations from an applied sinusoidal voltage and a measured alternating current. [Figure 3] FIG. 9 is a graph showing gain coefficients calculated for different frequencies based on measurements made by an impedance analyzer and an LCR meter in PBS buffers of different concentrations. [Figure 4] FIG. 12 shows an equivalent circuit of an impedance biosensor for a particulate-labeled immunoassay. A similar circuit consists of the capacitance effect of the electrodes, the double-layer capacitance of the electrode and particle surfaces, and the resistance of the solution. [Figure 5] FIG. 15 is a graph showing a plot of buffer conductivity against buffers of different concentrations when measured by a conductivity meter while using the circuit shown in FIG. 4. [Figure 6] FIG. 18 shows a simulated equivalent circuit using Simulink, MATLAB®. [Figure 7A] FIG. 21 shows a surface functionalization process for a human tumor necrosis factor-α (TNF-α) immunoassay using improved detection with detector antibody-conjugated microparticles. Formation of hydroxyl groups by oxygen plasma cleaning on a glass substrate is shown. [Figure 7B]This figure shows a surface functionalization process for a human tumor necrosis factor-α (TNF-α) immunoassay using improved detection with detection antibody conjugate microparticles. It shows a 30-minute incubation of 3% APTES after placing a removable PDMS masking film. [Figure 7C] This figure shows a surface functionalization process for a human tumor necrosis factor-α (TNF-α) immunoassay using improved detection with detection antibody conjugate microparticles. It shows covalent immobilization of activated capture antibodies by carbodiimide coupling. [Figure 7D] This figure shows a surface functionalization process for a human tumor necrosis factor-α (TNF-α) immunoassay using improved detection with detection antibody-conjugated microparticles. It illustrates incubation with the target analyte, TNF-α, and binding with the immobilized capture antibody. [Figure 7E] This figure shows a surface functionalization process for a human tumor necrosis factor-α (TNF-α) immunoassay using improved detection with detection antibody-conjugated microparticles. It illustrates mask removal and incorporation of a PDMS microfluidic channel. A phosphate-buffered saline (PBS) (1% PBSB) solution containing 1% w / v bovine serum albumin (BSA) was initially injected through the microchannel and incubated for 30 minutes for BSA surface immobilization. [Figure 7F] This figure shows a surface functionalization process for a human tumor necrosis factor-α (TNF-α) immunoassay using improved detection with antibody-conjugated microparticles as detectors. The flow and hydrodynamic washing of the detection antibody conjugated with magnetic microparticles are shown. In Figure 7F, a sandwich immunocomplex is formed between the immobilized capture antibody, target analyte, and detection antibody-conjugated microparticles. [Figure 8]This graph shows the PBS concentrations in deionized water at different concentrations, measured on an IDE using a commercially available LCR meter and impedance analyzer, and calculated from a simulation model. Buffer concentrations above 0.1 mM show saturation, making it impossible to detect impedance fluctuations due to immobilized biomolecules. At 0.01 mM PBS, the observed saturation effect was lower, and considering the pH buffering capacity, 0.01 mM PBS was selected as the buffer solution during the impedance signal measurement period. [Figure 9] This graph shows three different types of microparticles, including selected and characterized magnetic, polystyrene, and silica beads. All microparticles were the same size (2.8 μm) as the original microparticles and had the same surface coating consisting of streptavidin. When the microparticles were thermally denatured, the particles were boiled on a hot plate, and these denatured microparticles were then used for impedance measurements. All data were collected for a 5% IDE surface coverage. At each frequency tested, the samples were ordered from left to right as follows: magnetic beads, denatured magnetic beads, polystyrene beads, denatured polystyrene beads, silica beads, and denatured silica beads. [Figure 10] This graph shows the percentage change in impedance after normalization during the surface functionalization period. The percentage change in impedance signal increases with APTES and capture antibody on the IDE sensor, respectively. These changes decrease slightly as the frequency increases from 11 to 91 kHz, due to the effect of the resulting double-layer capacitance. [Figure 11A] This figure shows the impedance variation and IDE surface coverage for various target analytes with and without magnetic particle conjugate detection antibodies. It also shows the magnetic particle density on the IDE for various target concentrations. [Figure 11B]This graph shows the impedance changes and IDE surface coverage for various target analytes with and without magnetic microparticle conjugate detection antibodies. It also shows the impedance fluctuations after standardization and compares the capture antibodies when coexisting with various target concentrations of TNF-α at 11 kHz for both unlabeled (without magnetic microparticles) and sandwich microparticle-labeled (with magnetic microparticles) samples. By coupling the impedance sensor with a microparticle-labeled immunoassay format, TNF-α could be detected even at a low concentration of 83.46 pg / mL. This combined sensing technology can improve sensitivity by an order of magnitude. [Figure 11C] This graph shows the impedance change and IDE surface coverage for various target analytes with and without magnetic microparticle conjugate detection antibody concentrations. The impedance fluctuation graph compares the capture antibody (cAb) when coexisting with various target concentrations of TNF-α at 11 kHz for both unlabeled (cAb vs. TNF-α) and sandwich microparticle labeled (cAb vs. detector) samples, with the right axis showing surface coverage in percent. [Figure 12] This graph shows the magnitude of the impedance difference at 11 kHz for 2.8 μm magnetic nanoparticles when the surface coverage of the IDE is different. [Figure 13] This figure shows a schematic diagram of an exemplary microfluidic chip comprising a two-stage microfluidic network with a capillary valve and bridging holes. [Figure 14] This diagram illustrates the flow of the operating principle of a capillary valve in a two-stage microfluidic channel network, using different dyes: buffer solution - blue dye; and sample solution containing the target analyte - red dye. [Figure 15A] This figure shows an exemplary design of an integrated portable POC device including a microfluidic chip, impedance analysis circuit, power supply, data acquisition board, and processing unit. [Figure 15B]This figure shows another exemplary design of an integrated portable POC device with an integrated IDE and microfluidic chip. Using the design in Figure 15B, all enclosures and circuits are fabricated using a 3D printer. [Figure 16] This graph shows the buffer elevation time within the bridging hole based on the microfluidic design shown in Figure 13, for different bridging hole diameters. [Figure 17] This figure illustrates exemplary autonomous and sequential flow processes in an integrated IDE system with a capillary-driven microfluidic chip, using different dyes: buffer solution - blue dye; sample solution containing target analyte - red dye; and detection antibody conjugate microparticles - yellow dye. [Figure 18A] This figure shows the impedance fluctuation at 11 kHz obtained from a sandwich immunoassay on an integrated IDE sensing platform with a capillary-driven microfluidic network. Capture antibodies (cAb, mouse monoclonal anti-cardiac troponin I antibody) were immobilized on the IDE array, and various concentrations of human cardiac troponin I target analytes (cTnI) were tested using detection antibody conjugated microparticles (cTnI vs. detector) and without detection antibody conjugated microparticles (cAb vs. cTnI). The detection antibody was also a mouse monoclonal anti-cardiac troponin I antibody and was conjugated to 2.8 μm M-270 magnetic microparticles. The cTnI analytes were derived from human heart tissue and vacuum freeze-dried. [Figure 18B]This graph shows the impedance fluctuation at 11 kHz obtained from a sandwich immunoassay on an integrated IDE sensing platform with a capillary-driven microfluidic network. A capture antibody (cAb, mouse monoclonal anti-cardiac troponin I antibody) was immobilized on the IDE array, and various concentrations of human cardiac troponin I target analyte (cTnI) were tested using detection antibody conjugated microparticles (cTnI vs. detector) and without detection antibody conjugated microparticles (cAb vs. cTnI). The detection antibody was also a mouse monoclonal anti-cardiac troponin I antibody, conjugated to 2.8 μm M-270 magnetic microparticles. The cTnI analyte was derived from human heart tissue and vacuum freeze-dried. [Modes for carrying out the invention]

[0017] Unless otherwise defined, all technical and scientific terms used herein have the same meaning as commonly understood by those skilled in the art. For example, all nomenclature used in and in relation to cell and tissue culture, molecular biology, immunology, microbiology, genetics, and protein and nucleic acid chemistry, and hybridization described herein is well known and commonly used in the art. In case of any conflict, this disclosure, including definitions, shall control. Exemplary methods and materials are described below, although similar or equivalent methods and materials may be used when practicing or testing the embodiments and aspects described herein.

[0018] As used herein, the terms “amino acid,” “nucleotide,” “polynucleotide,” “vector,” “polypeptide,” and “protein” have their general meanings as understood by a biochemist skilled in the art. Standard single-letter nucleotide (A, C, G, T, U) and standard single-letter amino acid (A, C, D, E, F, G, H, I, K, L, M, N, P, Q, R, S, T, V, W, or Y) notation are used herein.

[0019] As used herein, terms such as “include,” “including,” “contain,” “containing,” and “having” mean “comprising.” This disclosure also considers embodiments or elements presented herein, whether expressly or not, that “comprising,” “consisting of,” and “consisting essentially of.”

[0020] Where used herein, “a,” “an,” “the,” and similar terms (in particular, in the context of the claims) are construed to encompass both singular and plural forms unless otherwise indicated herein or the context clearly contradicts them. In addition, unless otherwise specified, “a,” “an,” or “the” means “one or more.”

[0021] As used herein, the term "or" may be conjunctive or non-conjunctive. As used herein, the term “substantially” means to a very large or extremely large degree, but not entirely.

[0022] As used herein, the terms “about” or “approximately” mean, when applied to one or more numerical values ​​of interest, a numerical value similar to the reference value stated, or, as determined by those skilled in the art, a numerical value within an acceptable range of error for a particular numerical value (partially depending on how the numerical value is measured or determined, e.g., the limits of the measurement system). In one embodiment, the term “about” means any numerical value including both integers and fractional parts that fall within a maximum variation of ±10% of the numerical value modified by the term “about.” Alternatively, “about” may mean three or more standard deviations, according to the practices of the art. Alternatively, for example with respect to biological systems or processes, the term “about” may mean within a certain multiple of a numerical value, such as within five times in some embodiments and within two times in other embodiments. As used herein, the symbol “~” means “about” or “approximately.”

[0023] All ranges disclosed herein include both endpoints as individual numerical values, as well as all integers and fractions defined within that range. For example, the range 0.1 to 2.0 includes 0.1, 0.2, 0.3, 0.4, ... 2.0. Where an endpoint is modified by the term “approximately”, the defined range is extended by a variation of up to ±10% of any numerical value within the range, including the endpoint, or within three standard deviations or more.

[0024] As used herein, the terms “control” and “reference” are interchangeable. A “reference” or “control” level may be a predetermined numerical value or range adopted as the baseline or benchmark to which measured results are taken when evaluating the measured results. “Control” also refers to a control experiment or control cells.

[0025] As used herein, the term “dose” refers to any form of an active ingredient formulation or composition containing cells that contains an amount sufficient to initiate or produce a therapeutic effect with at least one or more doses. “Formulation” and “Composition” are used interchangeably herein.

[0026] As used herein, the term “prevention” means preventing or reducing the progression of a disorder to a statistically significant extent or to an extent detectable by a person skilled in the art. As used herein, the terms “effective dose” or “therapeutic dose” mean a substantially non-toxic but sufficient amount of an action, drug, composition, or cell(s) that, when administered to a subject, prevents, treats, or improves, to some extent, one or more of the symptoms of a disease or condition that the subject experiences or is susceptible to. The result may be a reduction or mitigation of the signs, symptoms, or causes of the disease, or any other desirable change in the biological system. The effective dose may be based on factors specific to each subject, including, but not limited to, the age, size, type or severity of the disease, stage of the disease, route of administration, type or degree of supplemental therapy used, ongoing disease process, and type of desired treatment.

[0027] As used herein, the term “subject” refers to an animal. Generally, the subject is a mammal. The subject may also refer to primates (e.g., humans, male or female; children, minors or adults), non-human primates, rats, mice, rabbits, pigs, cattle, sheep, goats, horses, dogs, cats, fish, birds, etc. In one embodiment, the subject is a primate. In one embodiment, the subject is a human.

[0028] As used herein, an object “requires treatment” if it derives a biological, medical, or quality-of-life benefit from such treatment. An object requiring treatment does not necessarily exhibit symptoms, especially in the case of preventive or prophylaxis treatments.

[0029] As used herein, the terms “inhibit,” “inhibit,” or “inhibiting” mean the reduction or suppression of a given biological process, condition, symptom, disorder, or disease, or a significant decrease in the baseline activity of a biological activity or process.

[0030] As used herein, “treatment” or “treating” means the prevention, prevention, inhibition, suppression, reversal, mitigation, improvement, or inhibition of the progression of a biological process, including a disorder or disease, or the complete elimination of the disease. Treatment may be carried out in either an acute or chronic manner. The term “treatment” also means reducing the severity of a disease or symptoms associated with such a disease before the disease becomes distressing. “Suppression” or “improvement” of a disease, disorder, or its symptoms relates to the administration of the cells, compositions, or compounds described herein to a subject after such a disease, disorder, or its symptoms have clinically manifested. “Prevention” or “prevention” of a disease, disorder, or its symptoms relates to the administration of the cells, compositions, or compounds described herein to a subject before the manifestation of a disease, disorder, or its symptoms. “Suppression” of a disease or disorder relates to the administration of the cells, compositions, or compounds described herein to a subject after the disease or disorder has been induced, but before it has clinically manifested or its symptoms have become apparent.

[0031] As used herein, “sample” or “target sample” means any sample in which the presence and / or level of a target analyte or target biomarker is to be detected or determined. Samples may include liquids, solutions, emulsions, or suspensions. Samples may include medical samples. Samples may include any biological fluid or tissue, such as blood, whole blood, blood fractions, such as plasma and serum, muscle, interstitial fluid, sweat, saliva, urine, tears, synovial fluid, bone marrow, cerebrospinal fluid, nasal secretions, sputum, amniotic fluid, bronchoalveolar lavage fluid, gastric lavage fluid, vomit, excrement, lung tissue, peripheral blood mononuclear cells, total leukocytes, lymph node cells, spleen cells, tonsil cells, cancer cells, tumor cells, bile, digestive fluids, skin, or combinations thereof. In some embodiments, a sample contains a certain amount. In other embodiments, a sample contains biological fluid or body fluid. Samples can be obtained by any means known in the art. If the sample is obtained from a patient, it may be used directly, or it may be pre-treated in some way to modify the properties of the sample, as discussed herein or as known in the art, such as by filtration, distillation, extraction, concentration, centrifugation, inactivation of interfering components, or addition of reagents. In one embodiment, the sample includes a blood sample or other biological liquid sample derived from the subject.

[0032] As used herein, “target analyte” or “target biomarker” refers to a substance associated with a biological state or process, such as a disease state, or a diagnostic or prognostic indicator of a disease or disorder (e.g., an indicator that identifies the presence or likelihood of delayed onset of a disease or disorder). The presence or absence of a biomarker, or an increase or decrease in biomarker concentration, may be associated with and / or suggest a particular state or process. Biomarkers may include, but are not limited to, cells or cellular components (e.g., viral cells, bacterial cells, fungal cells, cancer cells, etc.), small molecules, lipids, carbohydrates, nucleic acids, peptides, proteins, enzymes, antigens, and antibodies. Biomarkers may originate from infectious pathogens, such as bacteria, fungi, or viruses, or they may be endogenous molecules found in greater or less abundant amounts in subjects with a disease or disorder compared to healthy individuals (e.g., increased or decreased gene expression or gene product). In one embodiment, the target analyte is derived from a blood or serum sample obtained from a subject.

[0033] As used herein, “microfluidic” refers to the behavior, precise control, and manipulation of fluids geometrically and dimensionally reduced to a small scale (generally submillimeter) such that surface forces outweigh volume forces. “Microfluidic device” refers to a device or circuit comprising at least one microchannel having a cross-sectional dimension of less than 1 millimeter. In some embodiments of the disclosed invention, a microfluidic device may include a multilayered microfluidic chip.

[0034] As used herein, “impedance” means electrical impedance or electronic impedance. As used herein, “detecting a change in electrical impedance” or “detecting an impedance fluctuation” means that the impedance between one or more electrodes has a significant change that can be detected by an impedance analyzer or impedance measuring circuit when a molecular bonding reaction occurs between the electrodes. An impedance change or impedance fluctuation refers to the difference in impedance values ​​between one or more electrodes of a device when a molecular bonding reaction occurs and when no molecular bonding reaction occurs. The impedance between electrodes is generally a function of the frequency of the electric field or sinusoidal input voltage applied for detection and measurement.

[0035] As used herein, “Limit of Detection” or “LOD” means the lowest signal observable with sufficient confidence or statistical significance, or the corresponding lowest quantity determined from the signal. In certain embodiments of the disclosed invention, the lowest concentration of the target analyte that can be reliably detected is [LOD = [Average ブランク +1.645 × (SD ブランク )] + 1.645 × (SD 低濃度試料 It is reported as follows:

[0036] As used herein, “combined” refers to a structure in which projections originating in one direction are combined with projections originating in a different direction, such as fingers of a clasped hand. In one aspect of the disclosed invention, the first and second electrodes are part of a plurality of electrodes, which are combined with each other. The combined electrode (IDE) elements of this application preferably do not come into contact with each other. In one aspect of the disclosed invention, the distance between the first and second electrodes is about 1 μm to about 10 μm. In another aspect, the distance between the first and second electrodes is about 10 μm. In one embodiment, the disclosed device includes an IDE array.

[0037] As used herein, “conductive metal” means any metal having the ability to conduct electric current. In one embodiment, the conductive metal may be selected from gold (Au), titanium (Ti), aluminum (Al), copper (Cu), silver (Ag), zinc (Zn), iron (Fe), nickel (Ni), cobalt (Co), or any combination thereof. In some embodiments, the conductive metal may be deposited and coated on one or more electrodes of the disclosed device. In one embodiment, the first and second electrodes are coated with the conductive metal. In another embodiment, the conductive metal is selected from gold (Au), titanium (Ti), or any combination thereof.

[0038] As used herein, “bioassay” refers to a biochemical test for detecting the presence of a target analyte or target biomolecule in a solution and / or measuring its concentration, for example, through the use of one or more biomolecules, including antibodies or single-stranded DNA (ssDNA). In one embodiment, the bioassay includes an immunoassay comprising a first “capture antibody” or “cAb” that is covalently bound to or immobilized on a substrate layer and binds to and captures a target analyte in a solution. In one aspect of this embodiment, the immunoassay further comprises a second “detection antibody” or “dAb” used to create a sandwich immunoassay by binding to the target analyte captured by the first capture antibody.

[0039] As used herein, "specifically binds to" generally means that an antibody is adapted to bind to a target analyte more readily than it would to bind to a random, unrelated biomolecule or analyte.

[0040] In some embodiments of the disclosed invention, an impedance-based immunoassay biosensor device is described as measuring electrical impedance fluctuations in the presence of biomolecules from a sinusoidal input voltage. Several design parameters have been devised to improve the sensitivity of impedance-based biosensors. Generally, IDEs are used in impedance-based biosensors because they have advantages such as a high signal-to-noise ratio, low resistance drop, and rapid attainment of a steady state. The most important parameters of IDE design are the sensing area, electrode gap, and frequency range. The design of the IDE array itself can increase the overall sensitivity of the IDE by considering the overall sensing area. The electrode gap is the most important parameter in the sensitivity of an IDE biosensor because the electric field between the two electrodes can change effectively when biomolecular binding events occur. Sensitivity can increase further when the distance between the two electrodes decreases from a microgap to a nanogap. As the inter-electrode distance decreases, the chance of short-circuiting of the electrodes by the sample and label increases, which may lead to unsuccessful measurements. In the case of nanogap IDEs, it should be noted that buffer solutions with very low ionic strength are required to maintain a proper baseline, which may affect antigen-antibody interactions. Furthermore, nanogap IDEs require complex assembly procedures, leading to increased manufacturing costs, reduced yields, and decreased practicality. Therefore, a typical gap between electrode fingers for biosensing platforms is proposed to be 1–10 μm.

[0041] The selection of a suitable frequency is another important parameter related to assay sensitivity. At lower frequencies (<1 kHz), impedance is dominated by the leakage resistance of the IDE (which is highly sensitive to the electrode material). At higher frequencies above 100 kHz, solution resistance contributes to the net impedance, and measurement errors increase due to parasitic capacitance and inductance. At intermediate frequencies (1 kHz to 100 kHz), the measurement signal relies on the electrode surface capacitance, enabling the detection of affinity binding on the IDE. Therefore, most impedance-measuring biosensors employ frequencies from 1 kHz to 100 kHz, generally around 10 kHz, in which case the signal is relatively stable, and the impedance response is dominated by interfacial changes. In one embodiment, the first and second electrodes operate at frequencies in the range of approximately 1 kHz to approximately 100 kHz. In another embodiment, the first and second electrodes operate at frequencies of approximately 10 kHz.

[0042] Any biosensor intended for the detection of biomolecules should include a signal interpretation element for recognizing interfacial events within the reaction chamber. This element is typically integrated with a microfluidic biochip in a POC device and can be miniaturized in actual portable and practical devices. Impedance-based biosensors are ideal candidates due to their high sensitivity, improved signal-to-noise ratio, and suitability for integration, enabling signal digitization and smooth information transmission. IDEs used in well-established bioelectric sensors fall into the category of microelectrodes and have been widely used in POC platforms due to their inherent advantages such as real-time recognition, low resistance drop, and rapid establishment of a steady state. IDEs typically consist of a series of electrode fingers in the micrometer size range and detect interfacial bioreactions by applying sinusoidal currents.

[0043] Two different approaches exist to improve the overall immunoassay performance using this type of impedance-based biosensor. Electrochemical coating of gold nanoparticles onto the IDE significantly increases the overall surface area of ​​the electrode, thereby enhancing the signal-to-noise ratio for detecting the target analyte. While this approach provides a reasonable improvement in sensitivity, the cumbersome surface preparation of the IDE, as well as the biofunctionalization and further assembly steps, limit the realization of a more promising POC platform. Another approach is to improve overall sensitivity using micro / nanoparticles. The overall sensitivity of impedance-based biosensors can be improved by using microparticles or nanoparticles as labels and amplifying the impedance signal by inducing interfacial changes. Although particle labeling in electrochemical biosensors is widely used, it should be noted that it is better to record data under PBS buffer solution, as the electrochemical buffers used in Faradic measurements typically act as weak oxidizing agents that can denature some proteins. Furthermore, microparticle labeling in non-Faraday (or volumetric) biosensors is inherently simpler and more suitable for point-of-concept (POC) testing, which has the ability to perform measurements related to changes in interfacial volume during affinity binding without requiring the complete redox reaction that is necessary in Faraday biosensors.

[0044] Several studies have demonstrated the effects of particle size and material on detection sensitivity. Theoretical analysis suggests that labeled particles can efficiently block leakage electromagnetic fields, causing an increase in impedance with respect to the width and gap of the IDE. Using this principle, a sandwich immunoassay for quantifying fetal carcinogenic antigen was performed using gold nanoparticles (GNP), achieving a detection limit of 1 ng / mL. Similarly, an impedance-based real-time immunosensor in which galectin-1 antibody was conjugated to alumina nanoparticles was used to improve sensitivity and immobilization efficiency in the quantification of galectin-1 protein. While labeling with such particles has been utilized in various assays to improve sensitivity, overall results have been limited due to a lack of understanding of the effects of these particles, a lack of comparisons of signal enhancement by labeling, and a lack of implementation of appropriate immunoassay procedures.

[0045] Typical microparticle-based assays require microfluidic channels to induce controllable hydrodynamic forces. Hydrodynamic forces of 0.1–10 pN have been shown to break nonspecific binding, while those of 6–250 pN maintain specific binding. Therefore, controlling and utilizing the hydrodynamic cleaning force induced by flow velocity within the microfluidic channels is crucial for knocking down nonspecifically bound or unbound target analytes or microparticles from the IDE array and improving the overall signal-to-noise ratio.

[0046] By employing such a concept that utilizes signal enhancement and microfluidic systems based on microparticles, it is possible to develop a miniaturized impedance biosensor that incorporates all the components necessary for immunoassays within a single platform. In one embodiment, this device platform includes a gold IDE array chip for multiplexed assays, a small, specially designed 8-channel impedance analyzer connected to semi-real-time data acquisition software, and a microfluidic channel network for sample delivery and hydrodynamic cleaning.

[0047] In some embodiments, the microfluidic network is fabricated by one or more of the following methods: laser cutting, injection molding, die-cutting, grinding, press cutting, layer-by-layer fabrication, 3D printing, photolithography, or a combination thereof, to create microfluidic channels, capillary valves, bridging holes, reservoirs, reaction areas, inlets, and outlets. In one embodiment, the microfluidic network may have geometric dimensions for autonomous capillary flow of immunoassay reagents, such as buffer solutions and samples.

[0048] Capillary-driven microfluidic technology utilizes surface tension to drive the liquid within a channel, and this can be controlled by adjusting the channel shape. This technology employs a passive method of capillary-driven microfluidic technology to flow reagents. This passive flow handling method is independent of all external peripherals and therefore demonstrates developmental capabilities for miniaturized POC platforms. The original capillary-driven microfluidic platform is described as an autonomous fluid circuit typically combined with different capillary elements, such as a capillary pump, capillary stop valve (CSV), and trigger valve.

[0049] In some embodiments, a CSV designed to stop liquid in a microchannel based on surface properties with abrupt geometric changes is described. As the cross-section of the microchannel abruptly expands, the driving force induced by the capillary effect decreases, stopping the liquid when its leading edge reaches that point. Such CSVs are easy to fabricate and are reliable passive elements in capillary-driven microfluidic technology. However, the operating period, surface contact angle, and geometry should be carefully investigated to prevent corner flow, bubble inclusion, and malfunction. To improve the reliability of the stop valve, a two-level stop valve with a hydrophilic silicon microchannel covered with hydrophobic PDMS was developed. The burst pressure of such a stop valve was studied numerically and experimentally, taking into account the microchannel dimensions and liquid contact angle. The stop valve can be easily modified into a capillary trigger valve at the intersection of two perpendicular channels. In this configuration, liquid moving from one channel stops at the intersection until other liquid moving from the orthogonal channel reaches that point. When the intersection is wetted by the fluid in the linear channel, capillary force drives the flow into the channel, and the two fluids mix in a common channel located downstream. To enhance the functionality of the stop valve and reduce fabrication constraints, two-layer trigger valves have been introduced to achieve greater reliability and robustness, and have been tested to stop the fluid for up to 30 minutes. While these advances to date in the design of autonomous capillary-driven microfluidic technology have been brilliant, further investigation is needed to better control the affected parameters. Furthermore, only a few studies have been conducted to date to test these fluid control elements in actual immunoassays with reliable applications for proof-of-concept.

[0050] In one embodiment, a capillary-driven microfluidic technology platform integrated with a portable impedance analyzer is described for use in microparticle-labeled immunoassays. To effectively manipulate the fluid, a vertical CSV (Cellular Circulation System) linking two-stage or two-layer microfluidic technology can be introduced. This CSV allows reagents to be injected into the first stage once the fluid reaches a certain point after being prepared in the upper microfluidic stage or chamber. The concept of two-stage microfluidic technology not only enables the design of miniaturized microfluidic devices but also brings about the addition of a convenient pre-treatment incubation step. By characterizing the fluid parameters, the overall time for a detection assay (e.g., human troponin I) can be completed in less than 6 minutes using this integrated platform.

[0051] One embodiment described herein is a microfluidic device for detecting and quantifying biomolecules, comprising: (a) a multilayer microfluidic network including microfluidic channels, a buffer inlet, a sample inlet, and a waste outlet, wherein the buffer inlet, the sample inlet, and the waste outlet are in fluid communication with each other via the microfluidic channels, and the microfluidic network is configured to accept a buffer solution and a sample; and a microfluidic chip comprising a substrate layer including a surface to which a first antibody or capture antibody is covalently bound, wherein the first antibody is adapted to specifically bind to a target analyte, and the first antibody is mounted between a first electrode and a second electrode, wherein the substrate layer is in fluid communication with microfluidic channels, and the first antibody is positioned within the microfluidic channels; and (b) a detector for detecting changes in electrical impedance.

[0052] In some embodiments, microfluidic networks are described that include hydrophilic polymer materials comprising one or more of the following: polyacrylic acid, polymethyl methacrylate (PMMA), polycarbonate (PC), cyclic olefin copolymer (COC), polyimide, polydimethylsiloxane (PDMS), polyester, nylon, polyvinyl chloride, polyethylene, polypropylene, polyethylene terephthalate glycol, polybutylene adipate terephthalate, ethylene tetrafluoroethylene, fluorinated ethylene propylene, perfluoroalkoxyalkane, polylactic acid, polycaprolactone, polyoxymethylene, cellulose, its copolymers, or combinations thereof. In one embodiment, the microfluidic network is composed of polymer materials selected from any one of PC, PMMA, COC, polyimide, PDMS, or combinations thereof. In another embodiment, the microfluidic network is composed of PDMS.

[0053] Polymer materials, such as PDMS, are generally hydrophobic. In certain embodiments of the disclosed invention, surface modification techniques, including layer-by-layer (LBL) deposition, deposition of polyvinyl alcohol (PVA) after oxygen plasma treatment, or production of poly(ethylene glycol) coatings, can be implemented on the polymer material surface to precisely control the hydrophilicity of the polymer material and control the overall flow velocity of the microfluidic network. In one embodiment, the modification may involve deposition of BSA on the PDMS or substrate surface. In another embodiment, the modification may involve oxygen plasma treatment of the PDMS or substrate surface. Plasma oxidation treatment can be used to alter the surface chemistry of the PDMS or substrate by adding silanol (SiOH) groups to the surface. Atmospheric and argon plasmas are effective for this plasma treatment application. Plasma treatment makes the PDMS surface more hydrophilic, allowing for the maintenance of surface moisture by aqueous solutions. The oxidized surface can be further functionalized by reactions using trichlorosilane. Alternatively, for applications requiring long-term hydrophilicity, technologies such as hydrophilic polymer grafting, surface nanostructuring, and dynamic surface modification with embedded surfactants can also be used.

[0054] In one embodiment, the microfluidic network may comprise a chamber layer, a capillary valve, and a bridging hole located above the microfluidic channel, in which case the chamber layer and the microfluidic channel are fluidly connected to each other via the capillary valve and the bridging hole, and the sample inlet is located within the chamber layer. In one embodiment, the capillary valve comprises an orifice having a size in the range of approximately 100 μm to approximately 300 μm in diameter. In another embodiment, the capillary valve comprises an orifice having a size in the range of approximately 250 μm in diameter. In another embodiment, the bridging hole comprises an orifice having a size in the range of approximately 0.5 mm to approximately 2.5 mm in diameter. In yet another embodiment, the bridging hole comprises an orifice having a size in the range of approximately 1 mm in diameter.

[0055] In another embodiment, the chamber layer comprises a second antibody or detection antibody, the second antibody being adapted to specifically bind to a target analyte, and the second antibody is conjugated to microparticles. In one embodiment, the second antibody comprises a biotin moiety, the microparticles comprises a streptavidin coating, and the second antibody is conjugated to the microparticles through the binding of the biotin moiety to the streptavidin coating. In another embodiment, the microparticles comprise magnetic beads, polystyrene beads, silica beads, or a combination thereof. In another embodiment, the microparticles have a size in the range of approximately 1 μm to approximately 5 μm in diameter. In another embodiment, the microparticles comprise magnetic beads having a size of approximately 2.8 μm in diameter. In another embodiment, the chamber layer further comprises a porous polycarbonate (PC) membrane connected to the sample inlet, and the microparticles conjugated to the second antibody are immobilized on the porous PC membrane.

[0056] Another embodiment described herein is a method for detecting and measuring the presence of a target analyte in a sample using any one of the embodiments or aspects of the disclosed device, comprising: (a) loading a buffer solution into a buffer inlet; (b) flowing the buffer solution onto a substrate layer; (c) loading a sample into a sample inlet; (d) mixing the buffer solution with the sample; (e) sequentially flowing the mixture of the buffer solution and the sample onto the substrate layer, wherein the target analyte binds to a first antibody or capture antibody; (f) continuously flowing the buffer solution onto the substrate layer to remove any unbound target analyte; and (g) detecting a change in electrical impedance to quantify the concentration of the target analyte in the sample. In one embodiment, the flowing step is autonomous by capillary action. In another embodiment, the sample is incubated with microparticles conjugated to a second antibody or detection antibody, then mixed with a buffer solution, wherein the target analyte binds to the second antibody conjugated to the microparticles. In another embodiment, microparticles conjugated with a second antibody are immobilized on a porous PC membrane, and after the sample is loaded into the sample inlet, when the sample is dropped onto the porous PC membrane, the microparticles are released from the membrane, and the sample is incubated with the released microparticles conjugated with the second antibody. In another embodiment, the sample is loaded into the sample inlet after the buffer solution has reached the capillary valve. In another embodiment, the capillary valve is opened when the sample comes into contact with the buffer solution. In another embodiment, the method includes a total assay time ranging from approximately 5 minutes to approximately 10 minutes.

[0057] In one embodiment, a portable impedance-measuring biosensor device is described, used to measure changes in electrical impedance in the presence of biomolecules from a sinusoidal input voltage. The portable impedance-based biosensor platform can improve the sensitivity and LOD of immunoassays by using microparticles as labels. In one embodiment, a 2×4 IDE array having 10 / 10 μm electrodes / gap and a miniaturized impedance analyzer is described. By integrating microfluidic channels, immunoassays using microparticles are feasible to evaluate signal enhancement. As described herein, three different types of microparticles were tested at fixed sizes to clarify the effects of material properties and surface charge on reading sensitivity. Magnetic microparticles, silica microparticles, and polystyrene microparticles were tested to understand the material dependence of microparticles on the sensor array. Of these microparticles, magnetic microparticles showed a high degree of enhancement of the signal generated from the sensor array, with associated stability. While using magnetic microparticles, the detection of human TNF-α is demonstrated for a series of immunoassays, and the signal enhancement levels are compared by measuring the LOD. Based on preliminary testing, magnetic microparticles demonstrate optimal performance for labeling purposes in TNF-α immunoassays. Anti-human TNF-α antibodies can be covalently bound to the IDE surface via EDC / s-NHS-mediated bioconjugation. Anti-TNF-α antibodies conjugated with different concentrations of target analytes and magnetic microparticles are introduced as detectors. By obtaining the immunosensor response (normalized impedance variation) from the surface coverage of microparticles immobilized on the IDE, the limit of detection (LOD) can be improved tenfold in this immunoassay from 0.99 ng / mL for unlabeled detection to 83 pg / mL for microparticle-labeled detection.

[0058] It will be apparent to those skilled in the art that suitable modifications and adaptations to the compositions, formulations, methods, processes, and applications described herein can be made without departing from the scope of any embodiment or aspect thereof. The compositions and methods presented are illustrative and are not intended to limit the scope of any of the specified embodiments. Any of the various embodiments, aspects, and options disclosed herein can be modified or repeated and combined as desired. The scope of the compositions, formulations, methods, and processes described herein includes all actual or potential combinations of the embodiments, aspects, options, examples, and preferences described herein. In the exemplary compositions and formulations described herein, any component may be omitted, any component disclosed herein may be replaced, or any component disclosed herein may be included. The ratio of the mass of any component of any composition or formulation disclosed herein to the mass of any other component in the formulation, or to the total mass of the other components in the formulation, is disclosed herein as if they were expressly disclosed. If the meaning of any term found in any patent or publication incorporated by reference conflicts with the meaning of a term used in this disclosure, the meaning of the term or idiom in this disclosure shall prevail. Furthermore, the above discussion discloses and describes only exemplary embodiments. All patents and publications cited herein are incorporated by reference for the purpose of providing specific teachings.

[0059] The various embodiments and aspects of the present invention described herein are summarized in the following sections. Item 1. A microfluidic device for detecting and quantifying biomolecules, (a) A multilayer microfluidic network comprising microfluidic channels, a buffer inlet, a sample inlet, and a waste outlet, wherein the buffer inlet, sample inlet, and waste outlet are in fluid communication with each other via the microfluidic channels, and the microfluidic network is configured to accept a buffer solution and a sample, and A substrate layer comprising a surface to which the first antibody is covalently bound, adapted so that the first antibody specifically binds to the target analyte, and the first antibody is mounted between the first electrode and the second electrode. The substrate layer is in fluid communication with a microfluidic channel, and the first antibody is located within the microfluidic channel. Microfluidic chips and (b) A detector for detecting changes in electrical impedance and A device that includes this.

[0060] Item 2. The device according to Item 1, wherein the microfluidic network is composed of a polymer material selected from one of the following: polycarbonate (PC), polymethyl methacrylate (PMMA), cyclic olefin copolymer (COC), polyimide, polydimethylsiloxane (PDMS), or a combination thereof. Item 3. The device described in Item 2, in which the microfluidic network is composed of PDMS. Item 4. A device according to any one of items 1-3, wherein the microfluidic network is the geometric dimensions of the autonomous capillary flow of the buffer solution and sample. Item 5. The device according to Item 4, wherein the microfluidic network comprises a chamber layer, a capillary valve, and a bridging hole located above the microfluidic channel, the chamber layer and the microfluidic channel being in fluid communication with each other via the capillary valve and the bridging hole, and the sample inlet being located within the chamber layer. Item 6. The device according to Item 5, wherein the chamber layer contains a second antibody, which is fitted to specifically bind to a target analyte, and the second antibody is conjugated to a microparticle. Item 7. The device according to Item 6, wherein the second antibody comprises a biotin moiety, the microparticle comprises a streptavidin coating, and the second antibody is conjugated to the microparticle through the binding of the biotin moiety to the streptavidin coating. Item 8. A device according to item 6 or 7, wherein the particulate matter includes magnetic beads, polystyrene beads, silica beads, or a combination thereof. Item 9. The device described in Item 8, wherein the fine particles have a size in the range of approximately 1 μm to approximately 5 μm in diameter. Item 10. The device described in Item 9, wherein the microparticles include magnetic beads having a size of approximately 2.8 μm in diameter. Item 11. The device according to any one of items 6 to 10, wherein the chamber layer further comprises a porous polycarbonate (PC) membrane connected to the sample inlet, and microparticles conjugated to a second antibody are immobilized on the porous PC membrane. Item 12. A device according to any one of items 5 to 11, wherein the capillary valve comprises an orifice having a size in the range of approximately 100 μm to approximately 300 μm in diameter. Item 13. The device described in Item 12, wherein the capillary valve comprises an orifice having a diameter of approximately 250 μm. Item 14. A device according to any one of items 5 to 13, wherein the bridging hole comprises an orifice having a size in the range of approximately 0.5 mm to approximately 2.5 mm in diameter. Item 15. The device described in Item 14, wherein the bridging hole comprises an orifice having a diameter of approximately 1 mm. Item 16. A device according to any one of items 1 to 15, wherein the buffer solution contains phosphate-buffered saline (PBS) at a concentration in the range of approximately 0.001 mM to approximately 1 mM. Item 17. The device described in Item 16, wherein the buffer solution contains PBS at a concentration of approximately 0.01 mM. Item 18. The device according to item 16 or 17, wherein the buffer solution further comprises approximately 1 wt% bovine serum albumin (BSA). Item 19. A device according to any one of items 1 to 18, wherein the microfluidic network further comprises one or more absorbent pads. Item 20. A device according to any one of items 1 to 19, wherein the sample inlet further comprises a serum separation membrane. Item 21. A device according to any one of items 1 to 20, wherein the substrate layer includes a glass substrate or a plastic substrate. Item 22. The device described in any one of items 1 to 21, wherein the first and second electrodes are coated with a conductive metal. Item 23. The device described in Item 22, wherein the conductive metal is selected from gold (Au), titanium (Ti), or a combination thereof. Item 24. A device according to any one of items 1 to 23, wherein the distance between the first and second electrodes is approximately 1 μm to approximately 10 μm. Item 25. The device described in Item 24, wherein the distance between the first and second electrodes is approximately 10 μm. Item 26. A device described in any one of items 1 to 25, wherein the first and second electrodes operate at frequencies in the range of approximately 1 kHz to approximately 100 kHz. Item 27. The device described in Item 26, wherein the first and second electrodes operate at a frequency of approximately 10 kHz. Item 28. A device according to any one of items 1 to 27, wherein the first and second electrodes are part of a plurality of electrodes, and the plurality of electrodes are combined with each other. Item 29. A method for detecting and measuring the presence of a target analyte in a sample using a device described in any one of items 1 to 28, (a) A step of introducing the buffer solution into the buffer inlet, (b) The step of pouring the buffer solution onto the substrate layer, (c) The step of loading the sample into the sample inlet, (d) The step of mixing the buffer solution with the sample, (e) A step of sequentially flowing a mixture of buffer solution and sample onto a substrate layer, wherein the target analyte binds to the first antibody. (f) A step of continuously flowing the buffer solution over the substrate layer to remove any unbound target analytes, (g) A step of detecting a change in electrical impedance and quantifying the concentration of the target analyte in the sample. A method that includes this. Item 30. The method described in Item 29, wherein the flowing step is autonomous due to capillary action. Item 31. The method according to Item 29 or 30, wherein the sample is incubated with microparticles conjugated to a second antibody, then mixed with a buffer solution, and the target analyte binds to the microparticle-conjugated second antibody. Item 32. The method according to Item 31, wherein microparticles conjugated with a second antibody are immobilized on a porous PC membrane, and after the sample is loaded into the sample inlet, when the sample is dropped onto the porous PC membrane, the microparticles are released from the membrane, and the sample is incubated together with the released microparticles conjugated with the second antibody. Item 33. The method according to any one of items 29-32, wherein the sample is loaded into the sample inlet after the buffer solution has reached the capillary valve. Item 34. The method according to Item 33, wherein the capillary valve opens when the sample comes into contact with the buffer solution. Item 35. The method described in any one of items 29-34, wherein the sample includes a blood sample or other biological liquid sample derived from the subject. Item 36. The method described in any one of items 29-35, requiring a total assay time ranging from approximately 5 minutes to approximately 10 minutes. [Examples]

[0061] [Examples]

[0062] Design of an Impedance Sensing Platform Impedance-based immunosensors utilize the formation of immune complexes (e.g., antibodies as bioreceptors and specific antigens as their corresponding analytes) within a thin-layer configuration on the electrode surface. This complex formation alters the interfacial capacitance and resistance at the electrode / electrolyte interface. The electrical impedance signal is expressed as the ratio of the voltage phase to the current phase. A difference arises between these two phases when the electric field on the sensing electrode is disrupted and / or altered due to the presence of biomolecules at the electrode interface. Here, we designed a custom microfluidic impedance measurement system, as shown in Figure 1, consisting of a gold (Au) IDE array chip, impedance analysis circuitry, LabVIEW software and associated data acquisition (DAQ) board, and a microfluidic channel.

[0063] Based on conventional photolithography processes and Ti / Au deposition methods, an Au IDE array chip with a width of approximately 10 μm (approximately the width of a finger and the distance between electrodes) was fabricated on a glass wafer, and eight sets of IDEs were generated on a rectangular glass chip approximately 30 × 30 mm in size. The electrical impedance in the analyzer circuit is measured by a 12-bit impedance converter chip AD5933 (Analog Devices Inc.). A detailed diagram of the apparatus is illustrated in Figure 2A. A schematic of the exemplary IDE chip fabrication process is shown in Figure 2B. The glass substrate is first coated with a photoresist (PR) material. The PR-coated glass substrate is then exposed to UV light to fully develop the PR. E-beam deposition is then performed to coat with a conductive metal, such as titanium or gold (3 / 50 nm). The PR is then stripped to produce the IDE chip. Figure 2C illustrates the IDE, the equivalent circuit, and the formula that shows how the IDE measures impedance variation from a sinusoidal applied voltage and the measured AC current. Sinusoidal excitation signal (V PP Apply (=200mV) to each IDE pair, read the current obtained from the circuit board, and calculate the Discrete Fourier Transform (DFT). Since the DFT measures the frequency-dependent energy of the signal, the magnitude and phase of the impedance Z at a certain frequency are given by the following equation.

[0064]

Number

[0065] can be obtained more. However, in the formula, v i , i o , V i , I o , φ i , φ o , R fb , R A , V o are the sinusoidal input voltage, output current, input voltage amplitude, output current, phase of the input signal, phase of the output signal, feedback resistance, internal amplifier gain, and output voltage amplitude, respectively. The AD5933 holds the real (R) and imaginary (I) parts of the DFT in two 16-bit registers. The two data registers are accessible by a DAQ board using the I2C protocol and are stored in LabVIEW software with special specifications after data processing. The gain coefficient and system phase offset were first calibrated by measuring a resistor with known impedance using an LCR-meter (VSP / VMP3, Bio-Logic Science Instrument) and sweeping the frequency within a preferred range of 11 kHz to 91 kHz (Figure 3).

Example

[0066] Equivalent circuit of the particulate amplification type impedance sensor When the IDE biosensor uses particulate amplification, it is considered a non-Faradaic type. Therefore, in order to gain a deeper understanding of its operating principle, an investigation can be conducted on an equivalent circuit model consisting of a two-electrode system. A schematic diagram of this equivalent circuit is shown in Figure 4. C g is the geometric capacitance of the electrode determined by the dimensions of the electrode (thickness, gap, etc.) and the dielectric properties of the surrounding solution. C dl represents the double-layer capacitance of the two electrodes and appears at the interface between the conductive electrode and the adjacent buffer solution. The resistance of the solution between the two electrodes is

[0067]

number

[0068] It is given by the above formula, where n, l, w sp L, and L are the number of electrodes, the length of the fringe, the solution conductivity, the distance between two electrodes, and the distance between the centers of two adjacent electrodes, respectively. K(m) is the exact elliptic integral with coefficients of the first kind,

[0069]

number

[0070] It is defined as follows. The buffer conductivity (κ) was measured using a portable conductivity meter (Oakton CON 6+, Cole-Parmer, USA) and plotted for different concentrations of PBS buffer in Figure 5.

[0071] At this boundary, when a voltage is applied, two ion layers with opposite polarities are formed on the electrode surface and the buffer. These two ion layers are separated by a single layer of solvent molecules, which adhere to the electrode surface and act as a dielectric in a conventional capacitor. s R is the resistance of the buffer solution present between the two electrodes, and it changes depending on the buffer solution, but is not affected by interfacial affinity. sThis depends on the space filled by the buffer, and therefore on the length of the electrodes and the size of the gap between them. To verify each of these components, a simplified Randles equivalent circuit was simulated using Simulink and MATLAB (Figure 6). As shown in Figure 6, the two double-layer capacitances, C(DL1) and C(DL2), are considered to be similar due to the symmetry of the electrodes and are fixed at 30 nF. The values ​​for R (dielectric) and Cg (dielectric) are estimated to be 20 kΩ and 0.6 nF, respectively. The magnitude of the impedance of this solution was also measured using a specially designed impedance analyzer, and the same trend was observed. Note that the values ​​obtained from the impedance analyzer are the reciprocals of the absolute values ​​measured by the LCR meter. From these measurements, the specially designed impedance analyzer is given by the equation (G(f) = 1 / (Z LCR (f) × Z アナライザー (f))) is calibrated and plotted against the desired frequency range.

[0072] Considering the antibody immobilization procedure, oxygen plasma treatment generates hydroxyl groups on the glass surface, making it highly likely that the antibody will be immobilized between the electrodes. Therefore, by performing a microparticle amplification IDE immunoassay, the microparticles are primarily located between the electrodes. In this case, the microparticles affect the impedance between the electrodes (by introducing their resistance) and the bilayer capacitance to the gap. The portion highlighted as "particle effect" in Figure 4 is the equivalent component induced by a single microparticle. At high target concentrations, multiple microparticles are located between the electrodes, and the corresponding equivalent components are successively repeated, causing changes in the impedance measurement. [Examples]

[0073] Incorporation of microfluidics To incorporate a simple microfluidic, a PDMS-based microfluidic channel was fabricated by soft lithography. The channel height and width were 110 μm and 2.5 mm, respectively, which was sufficient to cover the entire area of ​​the IDE. A mini vacuum pump (12 / 02EB, Thomas Pump) was placed at the outlet to apply the sample solution into the inlet of the microfluidic channel using a pipette, drive all the solutions for the immunoassay, and perform hydrodynamic washing to remove unbound or nonspecific particles. [Examples]

[0074] Characterization of microparticles To evaluate the amplification effect of various material microparticles on impedance signals, streptavidin-coated magnetic microparticles (Dynabeads M-280, Thermofisher Scientific, USA), polystyrene microparticles (08-19-303, Micromod, Germany), and silica microparticles (43-19-303, Micromod, Germany) were selected. First, to estimate the effect of the materials on impedance signals, these three different microparticles, each 2.8 μm in diameter, were investigated with a 5% IDE surface coverage using impedance spectroscopy. To control the surface coverage (%) by the microparticles, the microparticle suspension was diluted, and after it settled on the surface (approximately 30 seconds), the surface coverage was measured under a microscope. Next, to eliminate the effect of surface charge due to material variations, these microparticles were boiled at 120°C for 120 minutes to denature all proteins on the surface. From this measurement, the optimal microparticles were determined for a series of immunoassays. In addition, the resolution of the IDE sensing platform was validated by measuring impedance with respect to the surface coverage of the magnetic microparticles. After preparing various particulate coverages (0.01–10%) on the IDE array, the change in impedance magnitude was measured to estimate the overall LOD of this platform. [Examples]

[0075] Surface functionalization and on-chip human TNF-α immunoassay To demonstrate an immunoassay using an impedance sensor, the IDE surface was functionalized with a capture antibody (anti-TNF-α, Abcam, UK) as a receptor for detecting a TNF-α target analyte. In this study, carbodiimide-induced crosslinking was used to immobilize the capture antibody in the glass surface gap between the Au electrodes. Figures 7A–7F illustrate the entire process of surface functionalization and immunoconjugation. The IDE chips were first cleaned by treating them with a freshly prepared piranha solution (H2SO4 / H2O23:1V / V) for 30 seconds, followed by thorough washing with deionized (DI) water and drying with nitrogen. Next, a patterned PDMS film (HT6240 Rogers Corp., USA) with an opening window on the IDE array was placed on the cleaned IDE chips so that the biochemical reaction would occur only in the sensing area. After treatment with oxygen plasma, the hydroxylated sensor surface was functionalized with 3% (3-aminopropyl)triethoxysilane (APTES) to obtain an amine-functionalized surface. To adapt the carbodiimide coupling method, the capture antibody was activated with 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide (EDC) and sulfo-N-hydroxysuccinimide (sulfo-NHS) (Thermo Scientific, USA), and then coupled to the APTES-modified surface. After incubating the capture antibody with the crosslinking reagent for 40 minutes, the IDE tip was rinsed with DI water. Next, human TNF-α (ab9642, Abcam, USA) target analytes at seven different concentrations (10 μg / mL to 100 pg / mL after 10-fold dilution) were further incubated on the patterned area for 30 minutes. Finally, the PDMS microfluidic device was coupled to a glass slide via oxygen plasma treatment. A diluted phosphate-buffered saline (PBS) (1% PBSB) solution containing 1% w / v bovine serum albumin (BSA) was first injected through the microchannel and incubated for 30 minutes to immobilize the surface. Next, a biotinylated anti-human TNF-α antibody (R&D Systems, USA) was conjugated with streptavidin-coated microparticles, and the microparticles were introduced into the microchannel using an external minipump.The impedance signal was measured at each step in the formation of the immune complex layer. Before each measurement, careful washing was performed using an appropriate buffer solution to thoroughly remove unbound molecules and ionic residues. To eliminate the influence of baseline variability, the impedance response was standardized during the measurement period (ΔZ) using the initial impedance (Z0) for each IDE pair, which is called the standardized impedance variation (ΔZ / Z0). [Examples]

[0076] Effect of buffer concentration on impedance sensor performance Since impedance sensors are highly sensitive to specific buffer solutions, a suitable buffer should be selected for optimal particulate label impedance sensing. For various PBS concentrations, impedance values ​​were first measured using an impedance analyzer (a simulated Randles model) and then validated using an LCR meter. Figure 8 shows all measured and simulation results for PBS concentrations from 1 mM to 0.001 mM at the investigated frequencies (11 kHz to 91 kHz), demonstrating similar trends and consistent results for all obtained values. A detailed comparative study using simulation results at 11 kHz and 0.001 mM PBS revealed differences of 11% and 13.5% from the LCR meter and portable impedance analyzer, respectively. Since impedance signals are more sensitive to electrodes at lower frequencies, these differences at low buffer concentrations may be due to manufacturing variations in the IDE electrodes. In this work, standardized impedance variations are primarily used for immunoassays, and therefore the differences may have minimal impact on the measurements. As shown in Figure 8, PBS buffers with concentrations of 0.1 mM or higher exhibit similar impedance magnitudes in the range of less than 5 kΩ (the signal is dominated by the buffer solution). Further dilution of PBS to 0.01 mM increases the impedance magnitude to over 10 kΩ. Since impedance fluctuations originating from the immunoassay are indistinguishable under highly conductive saturated buffer solutions, 0.01 mM PBS or lower PBS concentrations are suitable for impedance-based immunoassays. However, excessively low PBS concentrations reduce buffering capacity. Considering background signaling and buffering capacity, 0.01 mM PBS was selected as the sample buffer solution for all subsequent experiments. Table 1 below shows the following formula:

[0077]

number

[0078] This is the solution resistance to IDE calculated using [the specified method].

[0079] [Table 1] [Examples]

[0080] Fine particle material composition for impedance sensor performance Since the material and surface properties of microparticles have a significant impact on the impedance signal, the effects of various microparticles on impedance measurements were investigated. Generally, microparticles affect capacitance by forming a dielectric layer and resistance by blocking electric field fringes, as can be predicted from the equivalent circuit (Figure 4). Figure 9 shows the normalized impedance variation induced by various microparticles in the frequency range of 11 kHz to 91 kHz, along with the resistance and capacitance values.

[0081]

number

[0082] Both values ​​are represented. The impedance values ​​were standardized for each IDE with respect to the baseline signal obtained from the buffer solution. To separate the material and surface effects of the microparticles, streptavidin-coated microparticles and modified microparticles were prepared for such characterization experiments. For polystyrene and silica microparticles with streptavidin, the impedance fluctuation rate (%) varied from 1.10% to 3.04% and 1.11% to 2.02%, respectively, depending on the frequency. After modifying these microparticles, the impedance fluctuation increased to 3.98–4.53% for polystyrene and to 4.08–4.85% for silica. However, for magnetic particles, the impedance fluctuation was observed to decrease to 2.16–3.36% for modified microparticles. These results can be explained by the interfacial interaction and enhancement of the microparticle surface charge when streptavidin is present on the microparticles. In the case of silica and polystyrene microparticles, the surface charge bridges the electric field and contributes to the reduction in impedance fluctuation compared to the modified state. On the other hand, in electrically transparent superparamagnetic nanoparticles, if no surface charge is present, the dielectric constant increases, and the overall impedance fluctuation decreases slightly. Furthermore, for electrically impermeable materials, a decrease in the normalized impedance fluctuation is observed by gradually increasing the measurement frequency, clearly indicating that the capacitance effect is more dominant than the resistance effect under both streptavidin coating and modification conditions. However, in the case of magnetic nanoparticles, the impedance fluctuation is less dependent on the applied frequency, demonstrating that the resistance formed by the magnetic nanoparticles is more dominant than the bilayer capacitance induced by nonconductive particles (Figure 4). In the case of magnetic nanoparticles, the polarization effect also contributes to the magnitude of the overall impedance. When magnetic nanoparticles are located between I and D, the external electric field polarizes the magnetic nanoparticles, thus changing the uniformly distributed electric field near the nanoparticles and resulting in higher resistance.Furthermore, by calculating the relative standard deviation (RSD%), it was demonstrated that the normalized impedance variation in the magnetic nanoparticles exhibited better accuracy (11%) compared to polystyrene (21%) and silica (19%) in the studied frequency range of 11kHz to 91kHz. Based on the above considerations, magnetic nanoparticles were selected for the final measurements due to their higher output signal and better accuracy. [Examples]

[0083] Effect of surface functionalization on impedance fluctuations Figure 10 shows the normalized impedance variation after surface functionalization. APTES is the first layer located at the top of the bare electrode, and the immobilized capture antibody is the second layer. Although the thickness of the single layer of APTES is assumed to be less than 10 nm, its coverage as the first capacitance layer causes a significant increase in impedance. This indicates the importance of the first layer, suggesting that minimizing its capacitance as much as possible improves sensitivity during the immunoassay. After the addition of the capture antibody, the normalized impedance variation increased due to the antibody layer. From these thickness variations, 11 kHz corresponds to the most responsive frequency due to electrical double-layer capacitance compared to other frequencies, and therefore the frequency of 11 kHz was selected for further measurements. [Examples]

[0084] Hydrodynamic washing for effective immunoassays To ensure the high sensitivity of the microparticle-labeled immunoassay, the substrate surface should be washed with sufficient fluid force to remove any nonspecific and unbound microparticles. The washing step was performed in a microfluidic channel after introducing the detection antibody conjugated with microparticles. A constant buffer volume flow rate of 20 μL / min was maintained by applying negative pressure within the microchannel outlet using a small minipump. This volume flow rate corresponds to an average channel velocity of 1.3 mm / s, generating an overall force of 40 pN against the microparticles. This overall force is lower than the adhesion strength within the immunocomplex; however, it is sufficient to remove nonspecific binding from the surface. [Examples]

[0085] On-chip human TNF-α immunoassay Human TNF-α immunoassays were performed on impedance-based biosensors. TNF-α possesses important pro-inflammatory properties and plays a crucial role in innate and adaptive immunity, cell proliferation, and apoptosis processes. Increased TNF-α concentrations have been observed in acute and chronic inflammatory states (e.g., trauma, sepsis, infection, arthritis). Human TNF-α contains 157 amino acids, has a molecular weight of 17.4 kDa, and an isoelectric point (pI) of 5.8.

[0086] To perform the immunoassay, seven different concentrations of human TNF-α (10 μg / mL to 100 pg / mL) were initially incubated on a capture antibody-modified IDE for label-free detection. After measuring the impedance signal, anti-TNF-α antibody-conjugated magnetic microparticles were added to investigate signal enhancement by performing a sandwich immunoassay. Figure 11A shows the distribution of microparticles associated with different concentrations of TNF-α. From the surface coverage image shown in Figure 11A, it can be observed that the microparticle coverage is proportional to the concentration of the target analyte biomolecule. As shown in Figure 12, by obtaining a series of impedance fluctuations adjusted for the background signal, it was found that the magnitude of these differences increased due to the increase in the number of microparticles and the surface coverage. From LOD calculations at 11 kHz, it was found that this sensor can detect such small changes, even if the surface coverage of the magnetic microparticles changes by about 0.1%. Further measurements were performed using an impedance analyzer. Figures 11B to 11C show standard curves representing the relationship between the relative impedance signal and analyte concentration in the label-free and microparticle-labeled immunoassays. Statistical analysis showed that LODs of 0.9 ng / mL and 83.46 pg / mL were achieved for the unlabeled and microparticle-labeled immunoassays, respectively. Microparticles enable an order of magnitude improvement in LOD. In the unlabeled assay, TNF-α molecules bind to the capture antibody and contribute to an additional capacitance layer by forming a thin TNF-α molecular layer. Since the pI of TNF-α is 5.8, it has a partially negative charge in PBS buffer at pH 7.4. As the concentration of TNF-α antigen increases, more negative charge accumulates between the electrodes, causing greater electric field dispersion and impedance fluctuations. When labeled with magnetic microparticles, greater impedance fluctuations are observed compared to the unlabeled assay due to the blocking of the electric field and the material properties of the microparticles. At higher target concentrations, the magnetic microparticles will be closer to each other and may reside on the electrodes, although they are expected to be fixed between the electrodes on the functionalized area of ​​the glass. In this case, polarized particles between I / D points placed under an electric field generate an additional magnetic field, especially when the distance is less than the Debye length.This results in a high impedance fluctuation signal when the target concentration is higher (>100 ng / mL). However, in such cases, the standard deviation also increases due to variations in the relative distances of the microparticles and random polarization. Furthermore, the coefficient of variation (CV) was investigated to quantify the accuracy of the assay method. In this work, a CV of less than 20%, which is appropriate for the dynamic range of the target biomolecule concentration, was achieved for both the unlabeled and microparticle-labeled assay methods.

[0087] Table 2 below focuses on non-Faraday IDE-based immunosensors and interfaces for various targets based on previous research. Generally, label-free detection methods exhibit lower sensitivity compared to detection methods utilizing signal enhancement techniques. While the sensitivity of some label-free immunosensors has been improved by optimizing the detection protocol and using finger-shaped IDEs, signal enhancement techniques show more promising and cost-effective results for improving the overall LOD. In this study, given that the overall LOD is improved by more than an order of magnitude by using particulate labeling and controlled hydrodynamic cleaning force via microfluidic channels, the impedance signal can be further enhanced by supplementing IDE spacing characteristics optimized for signal enhancement techniques.

[0088] [Table 2]

[0089] Much attention has been paid to enhancing the sensitivity of these platforms by developing various biosensors for detecting biomolecules. These studies demonstrate an impedance-measuring biosensor consisting of a compact impedance analyzer, a data acquisition board, and an IDE array integrated with a successfully developed microfluidic channel. The impedance signal obtained from this platform was initially validated by simulations in MATLAB Simulink and an LCR meter. The platform for detecting monolayers of biomolecules and the resolution of microparticle detection were investigated, and human TNF-α was quantified using actual assays. After testing three different types of microparticles, the detection antibody was labeled with microparticles that produced the maximum and most consistent impedance signal. Using magnetic microparticles, the LOD (Level of Disturbance) was improved by an order of magnitude compared to the unlabeled bioassay. This novel, highly sensitive impedance biosensor, using microparticle labeling to enhance the signal, leverages microfluidic technology to obtain controlled hydrodynamic cleaning power, promising for POC (Point of Computing) applications.

[0090] Among the various digitized biosensors, impedance-based biosensors have proven promising for point-of-concept (POC) applications due to their ease of integration, miniaturization, rapid response, cost-effective assay methods, and convenient communication with smartphones. The platform described herein can be integrated with capillary-driven microfluidic technology for autonomous and sequential delivery of analytes. When used in microparticle-labeled immunoassays, this microfluidic approach should be precisely designed to have controlled hydrodynamic cleaning power by controlling the flow rate within a specific range. Further characterization and optimization may be performed to improve functionality and LOD. As demonstrated previously, micro or nano-sized particles do not significantly affect the impedance signal. However, optimal biosensing performance can be ensured by considering the ratio of the electrode gap to the micro / nanoparticle size. Furthermore, by incorporating a serum separation membrane on the microfluidic chip and including a communication chip that interacts with a smartphone or laptop, this device would become an ideal standalone POC platform with high sensitivity and the ability to perform diverse diagnostics. [Examples]

[0091] Capillary-driven microfluidic technology This paper describes an integrated biosensor based on capillary microfluidic technology and impedance, equipped with all the necessary components for performing a signal-amplifying sandwich immunoassay using microparticle labeling. Controlled capillary-driven force is utilized for fluid manipulation, and magnetic microparticles are adapted to enhance the impedance signal for improved LOD (Liquid Output Distance). To perform the bioassay, a small volume of sample (in the μL range) is introduced into a microfluidic chip that sequentially delivers all necessary samples and reagents into the sensing area. This integrated chip (or cartridge) is inserted into a portable impedance-measuring biosensor for signal reading. In a real immunoassay, using capillary microfluidic technology, reagents are sequentially delivered onto this integrated platform within approximately 6 minutes, minimizing user intervention. For demonstration purposes, human troponin I can be tested to check the overall functionality of this integrated platform. The microfluidic technology and IDE integrated platform enables a sample-in / answer-out immunoassay platform for rapid molecular diagnostics.

[0092] Capillary microfluidic design As shown in Figure 13, a two-stage capillary-driven microfluidic system was developed to perform a precisely controlled sandwich immunoassay. Both stages were designed in AutoCAD, and the corresponding masks were obtained from CAD / Art Services. A channel width of 500 μm was adopted for the buffer storage and waste sections, while the channel width from the branch to the sensing area was 200 μm. After developing the desired pattern as a master mold on a silicon wafer, a conventional photolithography process was followed. The channel heights of both stages were fixed at 100 μm. PDMS was mixed with a curing reagent in a 10:1 w / w ratio and cured overnight in a convection oven. After curing, the PDMS was peeled from the master mold and punched using a biopsy punch. On the first stage, buffer inlet, bridging hole, and outlet were punched with diameters of 2.5 mm, 1 mm, and 0.5 mm, respectively. The intermediate layer was cut from a PDMS film (HT6240, Rogers Corporation, USA), and a 250 μm pore was created on the intermediate layer to act as a capillary stop valve (CSV) between two PDMS stages. The top layer, designed to deliver the target analyte into the sensor, has two pores, 1.5 mm and 0.5 mm in size, for target inlet and degassing (venting), respectively.

[0093] Operating principle of capillary-driven microfluidic technology A buffer solution (blue dye) of 1% bovine serum albumin (1%PBSB) in 10 mM PBS is initially added to the buffer inlet, while the outlet vent is covered (see Figure 14). Capillary action drives the buffer through the sensing area towards the outlet. Since the ends of the channel are closed (if the outlet vent is covered), a total volume of 0.2 μL advances into the sensing area within each branch and stops 5 mm behind the bridging hole. Once the buffer reaches the bridging hole, it begins to rise in the pore until it is obstructed by the CSV located in the center of the two stages. The device is then ready to operate by introducing a sample (red dye) containing the target analyte into the sample inlet hole on the upper microfluidic stage. The sample first collides with detection antibody (dAb) conjugate particles located in the center of the channel and, upon forming an immunocomplex, is rapidly captured by capture antibody (cAb) on the sensing spot. By bringing this sample solution into contact with the bridging hole, the sample solution mixes with the buffer coming from the lower stage and flows into the sensing area. At that moment, by removing the cover from the outlet vent, all the analytes begin to flow over the sensing area towards the outlet and form immunocomplexes at the top of the IDE.

[0094] Design and assembly of an integrated IDE sensing platform An integrated, portable device containing all the necessary components for performing a capillary-driven microparticle-labeled sandwich immunoassay was designed, along with an impedance-measuring biosensor. The microfluidic chip at the top of the IDE chip is positioned on a slider that can be inserted into the device's box (Figures 15A-15B). The electrical impedance in the analyzer circuit, using a 12-bit impedance converter chip AD5933 (Analog Devices Inc.), was designed to be contacted using electrode pads and their spring pins on a microchip cartridge. As shown in Figure 15A, the device connects to a laptop using a USB port provided by the DAQ board. In the alternative design shown in Figure 15B, all enclosures and circuits were fabricated using a 3D printer.

[0095] Flow control in capillary microfluidic channels In capillary-driven microfluidic systems, controlling the flow velocity is a crucial parameter that should be carefully considered when designing the channel, as for certain materials, the flow velocity depends entirely on the geometric shape of the channel. By simplifying the Navier-Stokes equations for stratified steady flow that does not involve any other physical forces (e.g., gravity), the volumetric flow rate is:

[0096]

number

[0097] It is possible to estimate more precisely, where h, w, μ, and L(t) are the channel height, width, liquid viscosity, and time-dependent liquid length within the channel, respectively. This estimation has an error of 13% when h=w and decreases to 0.2% when h=w / 2. By substituting the Young-Laplace equation for ΔP, the flow velocity for a capillary-driven system can be calculated.

[0098] In the case of CSV, the pressure barrier is given by the following formula:

[0099]

number

[0100] According to the formula, it also depends on the geometric parameters and the liquid contact angle, where γ and h are the interfacial tension and channel height, respectively. It was reported that after 1 minute of oxygen plasma treatment, the water contact angle reached approximately 45° after 1 hour, and this was used as θ to calculate the burst pressure in the CSV. Assuming γ = 20 mN / m and h = 250 μm (capillary valve size), the burst pressure in each CSV is 55 Pa (equivalent to 5.5 mmH2O). The volumetric capacity of the fluid channel for h = 110 μm is 55 mm 3 (=μL), and the breakdown is 13mm from the entrance to the branching point. 3 and 10.5mm at each branch 3 This consists of the following: When the buffer reaches the bridging hole, 35 μL is the total volume required to complete the assay. Given the buffer inlet size of 3 mm and considering the PDMS layer thickness of 2 mm, the water head generated by this 35 μL buffer volume is less than the burst pressure of the CSV. This ensures the performance of the CSV designed for fluid networks.

[0101] To estimate the overall assay time, we investigated the filling time for different sections. It takes approximately 3 minutes for the buffer to completely fill the priming section and reach the sensing area. As plotted in Figure 16, we investigated the effect of the bridging hole diameter on the time it takes for the buffer to rise to the CSV within the bridging hole. It was found that a bridging hole diameter of less than 1 mm did not significantly change the rise time, and therefore a 1 mm bridging hole diameter was selected. With this 1 mm hole size, it takes 2 minutes for the capillary-driven buffer to fill. It takes approximately 1 minute to introduce the target analyte (approximately 1 minute), wait for the analyte to diffuse (approximately 1 minute), and then approximately 1 minute to remove the cover tape from the exit, fill the waste container, and complete the assay. Taking all times into account, it takes a total of approximately 6 minutes to complete the entire assay using this design.

[0102] Demonstration of a bioassay using human troponin I analyte As shown in Figure 17, a detailed fluid sequence was demonstrated using different dyes. According to the capture frame in Figure 17, the assay first begins with the buffer loading step from the buffer inlet. The buffer (blue dye) flows through the capillary channel until the Young-Laplace pressure reaches a steady state. Next, the sample solution (red dye) and detection antibody conjugation particles (yellow dye) are loaded into each well sequentially to allow the reaction to occur. After removing the cover tape above the outlet, pressure builds up, and force is applied to all the liquids, causing them to move through the capillary channel, and the antigen-antibody reaction occurs during the flow period.

[0103] Following these demonstrations, a troponin I assay was performed to further validate the integrated IDE platform. Figures 18A and 18B show preliminary assay results confirming all fluid sequences. By testing three different concentrations of troponin I, a linear increase in the IDE signal was observed as the concentration of human troponin I increased. This assay demonstration validated the functionality of the impedance analyzer and the integrated IDE chip, including the IDE and capillary-driven microfluidic chips.

[0104] We designed this capillary-driven microfluidic device and tested its autonomous and sequential dispensing of reagents necessary for performing microparticle-labeled immunoassays. To further miniaturize and add sample preparation, we designed a two-stage microfluidic system. First, we designed a vertical CSV and characterized it to connect the two-stage microfluidic technology. By characterizing the flow within the channel, the entire assay using this disposable biochip takes approximately 6 minutes. The chip is designed to be inserted into a portable, compact case containing all the necessary components for impedance-based biomolecular measurements.

[0105] An essential step in performing a one-step immunoassay using capillary microfluidic technology is the isolation of human serum from a whole blood sample. In microfluidic-based point-of-care (POC) applications, this isolation is typically achieved by means of a serum separation membrane. The two-layer microfluidic chip described herein allows for the easy introduction of a serum separation membrane into a disposable cartridge through a sample inlet.

[0106] To enable long-term use of this platform and maintain the surface wetting properties, the flow velocity within the microfluidic chip can be controlled by two main modifications. Firstly, different polymer materials, such as polycarbonate (PC), polymethyl methacrylate (PMMA), cyclic olefin copolymer (COC), and polyimide, are available to overcome the inherent shortcomings of PDMS. Secondly, surface modification techniques, including layer-by-layer deposition, deposition of polyvinyl alcohol after plasma treatment, or the creation of a poly(ethylene glycol) coating, can be applied to the PDMS surface to precisely control the hydrophilicity of PDMS and the overall flow velocity of the microfluidic network. Furthermore, by adding one or more absorbent pads within the waste area of ​​the microfluidic network or within the upper microfluidic stage, the flow velocity of buffers and sample solutions can be easily controlled, thus enabling the acquisition of a wider range of velocities. [Examples]

[0107] DNA microarray applications The disclosed devices and sensor arrays can also be used in DNA hybridization chips (i.e., DNA microarrays) by functionalizing the sensor surface with immobilized “captured” single-stranded DNA (ssDNA) molecules instead of capture antibodies. To introduce this concept, the same surface chemistry and functionalization processes can be used to immobilize various ssDNA molecules. Thus, target analyte DNA amplified from polymerase chain reaction (PCR) can be analyzed on the sensor array. Instead of microarray readers that commonly utilize fluorescence signals, differential impedance signals can be measured to determine the concentration levels of numerous genes derived from a DNA or RNA sample without requiring any additional equipment. Therefore, the disclosed devices and methods can be used in a wide range of bioassay applications, including, for example, immunoassays and DNA microarrays.

Claims

1. A microfluidic device for detecting and quantifying biomolecules, (a) A multilayer microfluidic network comprising a microfluidic channel, a buffer inlet, a sample inlet, and a waste outlet, wherein the buffer inlet, sample inlet, and waste outlet are fluidly connected to each other via the microfluidic channel, and the microfluidic network is configured to accept a buffer solution and a sample. A substrate layer comprising a surface to which a first antibody is covalently bound, adapted so that the first antibody specifically binds to a target analyte, and the first antibody is mounted between a first electrode and a second electrode, wherein the first electrode and the second electrode are part of a plurality of electrodes, the plurality of electrodes are combined with each other, and the distance between the first electrode and the second electrode is 1 μm to 10 μm. The substrate layer is in fluid communication with a microfluidic channel, and the first antibody is located within the microfluidic channel. Microfluidic chips and (b) A detector for detecting changes in electrical impedance, A device that includes this.

2. The device according to claim 1, wherein the microfluidic network is composed of a polymer material selected from one of the following: polycarbonate (PC), polymethyl methacrylate (PMMA), cyclic olefin copolymer (COC), polyimide, polydimethylsiloxane (PDMS), or a combination thereof.

3. The device according to claim 2, wherein the microfluidic network is composed of PDMS.

4. The device according to claim 1, wherein the microfluidic network has geometric dimensions for autonomous capillary flow of the buffer solution and the sample.

5. The device according to claim 4, wherein the microfluidic network comprises a chamber layer, a capillary valve, and a bridging hole located above the microfluidic channel, the chamber layer and the microfluidic channel are in fluid communication with each other via the capillary valve and the bridging hole, and a sample inlet is located within the chamber layer.

6. The device according to claim 5, wherein the chamber layer contains a second antibody, which is adapted to specifically bind to a target analyte, and the second antibody is conjugated to microparticles.

7. The device according to claim 6, wherein the second antibody comprises a biotin moiety, the microparticle comprises a streptavidin coating, and the second antibody is conjugated to the microparticle through the binding of the biotin moiety to the streptavidin coating.

8. The device according to claim 6, wherein the fine particles include magnetic beads, polystyrene beads, silica beads, or a combination thereof.

9. The device according to claim 8, wherein the fine particles have a size in the range of 1 μm to 5 μm in diameter.

10. The device according to claim 9, wherein the fine particles include magnetic beads having a diameter of 2.8 μm.

11. The device according to claim 6, wherein the chamber layer further comprises a porous polycarbonate (PC) membrane connected to the sample inlet, and microparticles conjugated to the second antibody are immobilized on the porous PC membrane.

12. The device according to claim 5, wherein the capillary valve comprises an orifice having a diameter in the range of 100 μm to 300 μm.

13. The device according to claim 12, wherein the capillary valve comprises an orifice having a diameter of 250 μm.

14. The device according to claim 5, wherein the bridging hole comprises an orifice having a diameter in the range of 0.5 mm to 2.5 mm.

15. The device according to claim 14, wherein the bridging hole comprises an orifice having a diameter of 1 mm.

16. The device according to claim 1, wherein the buffer solution contains phosphate-buffered saline (PBS) at a concentration in the range of 0.001 mM to 1 mM.

17. The device according to claim 16, wherein the buffer solution contains PBS at a concentration of 0.01 mM.

18. The device according to claim 16, wherein the buffer solution further comprises 1 wt% bovine serum albumin (BSA).

19. The device according to claim 1, wherein the microfluidic network further comprises one or more absorbent pads.

20. The device according to claim 1, wherein the sample inlet further comprises a serum separation membrane.

21. The device according to claim 1, wherein the substrate layer includes a glass substrate or a plastic substrate.

22. The device according to claim 1, wherein the first electrode and the second electrode are coated with a conductive metal.

23. The device according to claim 22, wherein the conductive metal is selected from gold (Au), titanium (Ti), or a combination thereof.

24. The device according to claim 1, wherein the distance between the first electrode and the second electrode is 10 μm.

25. The device according to claim 1, wherein the first electrode and the second electrode operate at a frequency in the range of 1 kHz to 100 kHz.

26. The device according to claim 25, wherein the first electrode and the second electrode operate at a frequency of 10 kHz.

27. A method for detecting and measuring the presence of a target analyte in a sample using a device according to any one of claims 1 to 26, (a) The step of introducing the buffer solution into the buffer inlet, (b) The step of pouring the buffer solution onto the substrate layer, (c) A step of loading the sample into the sample inlet, (d) The step of mixing the buffer solution with the sample, (e) A step of sequentially flowing a mixture of buffer solution and sample onto a substrate layer, wherein the target analyte binds to the first antibody. (f) A step of continuously flowing the buffer solution over the substrate layer to remove any unbound target analytes, (g) A step of detecting changes in electrical impedance and quantifying the concentration of the target analyte in the sample, A method that includes this.

28. The method according to claim 27, wherein the flowing step is autonomous due to capillary action.

29. The method according to claim 27, wherein the sample is incubated with microparticles conjugated with a second antibody, then mixed with a buffer solution, and the target analyte binds to the second antibody conjugated with the microparticles.

30. The method according to claim 29, wherein microparticles conjugated with a second antibody are immobilized on a porous PC membrane, and after a sample is loaded into the sample inlet, when the sample is dropped onto the porous PC membrane, the microparticles are released from the membrane, and the sample is incubated together with the released microparticles conjugated with the second antibody.

31. The method according to claim 27, wherein the sample is introduced into the sample inlet after the buffer solution has reached the capillary valve.

32. The method according to claim 31, wherein a capillary valve opens when the sample comes into contact with the buffer solution.

33. The method according to claim 27, wherein the sample comprises a blood sample or other biological liquid sample derived from the subject.

34. The method according to claim 27, comprising a total assay time in the range of 5 to 10 minutes.

Images