WO2016112079A1

WO2016112079A1 - Microfluidic aptasensor including a graphene nanosensor

Info

Publication number: WO2016112079A1
Application number: PCT/US2016/012297
Authority: WO
Inventors: Qiao Lin; Xuejun Wang; Jaeyoung Yang; Cheng Wang
Original assignee: Columbia University in the City of New York
Current assignee: Columbia University in the City of New York
Priority date: 2015-01-06
Filing date: 2016-01-06
Publication date: 2016-07-14
Anticipated expiration: 2017-07-06

Abstract

Graphene nanosensors for monitoring a target analyte utilizing anti-target analyte aptamers can include a single conductance sensor on a substrate platform, where the graphene sensor can be functionalized with aptamers for binding the target analyte, or alternatively, a nonosensor can include microbeads functionalized with aptamers which can allow for selective enrichment and isocratic elution of the target analyte, where the concentration of the enriched target analyte can be measured on a graphene surface functionalized with a target analyte of interest.

Description

MICROFLUIDIC APTASENSOR INCLUDING A GRAPHENE

NANOSENSOR

CROSS REFERENCE TO RELATED APPLICATIONS

This application claims priority from United States Provisional Application No. 62/100,412, filed January 6, 2015, which is incorporated by reference herein in its entirety.

STATEMENT REGARDING FEDERALLY FUNDED RESEARCH

This invention was made with government support under ECCS- 1509760 awarded by National Science Foundation. The government has certain rights in this invention.

BACKGROUND

Graphene is a single atom thick two-dimensional nanomaterial with honeycomb lattice of carbon. It exhibits desirable physical and chemical properties, such as high carrier mobility, good biocompatibility, and stability, for bio-sensing. Rapid and precise detection of biomarkers are important for disease diagnostics; aptameric graphene nanosensors are attractive because of the specificity afforded by the aptamer (i.e., target analyte-binding single-strand oligonucleotides) and the high sensitivity by graphene. Graphene nanosensors configured as field-effect transistors ("FET") can provide a sensitive, label-free biomarker detection platform because of the ability of surface-based binding to cause changes in the bulk properties of graphene.

However, certain existing aptameric graphene nanosensors have been limited to well-conditioned buffers, while analyte detection in physiological samples such as human serum has remained a challenge. One main difficulty involves the hydrophobic nature of pristine graphene, as background molecules in serum tend to nonspecifically adsorb to the graphene surface, thus interfering with the detection of the target analyte. Another difficulty is their application to the detection of small biomolecules is still limited, as the binding of low charged small molecules does not directly induce detectable changes in graphene conductance, hence suffering low sensitivity.

There is a need to develop graphene nanosensors that offer improved specificity and sensitivity, including for diagnostic testing.

SUMMARY

The disclosed subject matter provides systems and methods for detecting small biomolecules using a microfluidic nanosensors including a graphene nanosensor. The disclosed subject matter also provides a microdevice and techniques for monitoring a target analyte. In one aspect, the disclosed subject matter provides a microdevice and techniques for monitoring a target analyte in a sample using an aptamer capable of binding to the target analyte.

In certain embodiments, a microdevice includes a graphene nanosensor, on a substrate platform, wherein the graphene nanosensor can be functionalized with an aptamer that binds the target analyte. Target analyte binding of the aptamer changes the charge density on the graphene nanosensor surface, inducing changes in the carrier concentration of the sensor. In certain embodiments, the aptamers bind reversibly with essentially all analytes. In certain embodiments, the aptamers bind reversibly with the target analyte.

The nanosensor can have a FET configuration. In accordance with an exemplary embodiment, the nanosensor can include a substrate. The substrate can be, for example, a SiC^-coated silicon wafer. Source and drain electrodes can be patterned on the substrate. A graphene sheet can be coupled to the source and drain electrodes. The graphene surface can be functionalized with aptamers that binds the target analyte. A polymeric material (e.g., PDMS sheet) can define the well for testing the sample.

The disclosed subject matter also provides a microfluidic aptasensor including an enrichment chamber and a sensing chamber. The enrichment chamber can include one or more microbeads. Aptamers can be immobilized on the surface of the microbeads.

The sensing chamber can include a graphene nanosensor. The nanosensor can have a FET configuration. In accordance with an exemplary embodiment, the nanosensor can include a substrate. The substrate can be, for example, a SiC^-coated silicon wafer. Gate electrodes can be patterned on the substrate. A single-layer graphene sheet can connect the source and drain electrodes. The graphene surface can be functionalized with a control target analyte. A polymeric material (e.g., PDMS sheet) can define microchambers and channels. The enrichment chamber and the sensing chamber can be coupled via a serpentine channel. In accordance with embodiments of the disclosed subject matter, the aptasensor can further include a temperature control unit, an inlet coupled to the enrichment chamber, a waste outlet coupled to the serpentine channel, and a sample outlet coupled to the sensing chamber.

In accordance with another embodiment, the disclosed subject matter provides methods for detecting small biomolecules using a microfluidic graphene nanosensor. A sample can be introduced into the polymeric well. Target analytes (e.g., IgE) can be selectively captured by aptamers immobilized on the graphene nanosensor. Non-target molecules, including impurities, can then be removed through, e.g., buffer washing. The electrical conductance through the change in conductance caused by the target analyte binding to the aptamer immobilized on the graphene can be measured by measuring the drain current at a fixed drain voltage.

In accordance with yet another embodiment, the disclosed subject matter provides methods for detecting small biomolecules using a microfluidic aptasensor including an enrichment chamber and a sensing chamber. A sample can be introduced into the enrichment chamber. Target analytes (e.g., AVP) can be selectively captured by aptamers immobilized on microbead surfaces. Non-target molecules, including impurities, can then be removed through, e.g., buffer washing. The enriched target molecules can then be released at an elevated temperature and eluted with free aptamer. This mixture can travel to the sensing chamber via a serpentine channel. The eluate can be incubated with graphene pre-functionalized with a control (e.g., reagent) target analyte (e.g., with reagent AVP). The graphene- bound control (e.g., reagent) target analyte can compete with the sample target molecules to bind to and capture some of the free aptamer molecules, which can cause an increase in charge on the graphene. The electrical conductance through the graphene can be measured by measuring the drain current at a fixed drain voltage.

In one aspect, the aptamer can be a single-stranded oligonucleotide or a peptide. In certain embodiments, the aptamer can be a single-stranded oligonucleotide, which can be DNA, RNA, XNA, or a combination thereof. In certain embodiments, the aptamer can be functionalized to the surface of graphene via a linker. The linker can be a pyrene-terminated agent (e.g., 1-pyrenebutanoic acid succinimidyl ester). In certain embodiments, the graphene can be serially functionalized with a polymer (e.g₃ polyethlene glycol) and aptamers. In accordance with certain embodiments, the polymer can be coupled to the aptamer and the opposite end of the polymer can be coupled to the linker.

In certain embodiments, the nanosensor as disclosed herein can enable reliable monitoring of a target analyte in a sample. In certain embodiments, the sample can be a bodily fluid, a non-bodily fluid liquid, or a laboratory sample. In certain embodiments, the nanosensor can be used to measure the amount or change in the amount of a target analyte in a sample. In certain embodiments, the bodily fluid can be tears, blood, saliva, mucus, interstitial fluid, spinal fluid, intestinal fluid, amniotic fluid, lymphatic fluid, pericardial fluid, peritoneal fluid, pleural fluid, semen, vaginal secretions, sweat, or synovial fluid of a subject.

BRIEF DESCRIPTIONS OF THE DRAWINGS

Figure 1 : (a) Schematic of an aptameric graphene nanosensor (100) for analyte detection according to some embodiments of the disclosed subject matter; (b) Principle of aptameric graphene nanosensing according to some embodiments of the disclosed subject matter: (i) PEG and aptamer-functionalized graphene; (ii) Sample incubation; and (iii) target analyte (e.g., IgE) capture and purification on the graphene surface by the aptamer.

Figure 2: Image of a fabricated aptameric graphene nanosensor.

Figure 3 : Fabrication of an aptameric graphene nanosensor according to some embodiments of the disclosed subject matter.

Figure 4: Functionalization of the graphene surface according to some embodiments of the disclosed subject matter.

Figure 5 : (a) Schematic of a microfluidic graphene nanosensor that integrates aptamer-based selective analyte enrichment and detection (500) according to some embodiments of the disclosed subject matter; (b) Images of a fabricated the graphene nanosensor.

Figure 6: Fabrication of a microfluidic graphene nanosensor according to some embodiments of the disclosed subject matter.

Figure 7: Principle of aptameric graphene nanosensing that integrates aptamer-based selective analyte enrichment according to some embodiments of the disclosed subject matter: The sample target (e.g., AVP) was enriched on microbead surfaces by aptamer binding (a); After the buffer washing (b); the temperature was raised to 55 °C, which disrupted aptamer-sample target (e.g., AVP) complexes and released sample target (e.g., AVP) into a free aptamer solution (c); The mixture of the free aptamer and released sample target (e.g., AVP) was incubated with graphene functionalized with standard target (e.g., AVP) (d), inducing the binding of the free aptamer to the standard target (e.g., AVP) on graphene via competitive binding (e), thus changing the graphene conductance.

Figure 8: A plot showing the Raman spectrum of the graphene used in the aptameric graphene nanosensor.

Figure 9: A plot showing the fluorescence of single-strand of DNA complementary to an IgE-specific aptamer introduced to graphene with different surface modifications.

Figure 10: A plot showing the transfer characteristics measured before (dashed line) and after (solid line) serum incubation on different surfaces: (a) Serum on bare graphene surface and (b) Serum on PEG-grafted graphene surface.

Figure 1 1 : A plot showing the transfer characteristics in different serum samples with various IgE concentration. Inset: reduction in drain-source current with respect to increased IgE concentration at Vg=0.05 V.

Figure 12: A plot showing the dependence of the Dirac point voltage on IgE concentration.

Figure 13 : A plot showing the control testing on the PEG-aptamer functionalized graphene.

Figure 14: A method for CVD graphene synthesis and transfer procedure according to some embodiments of the disclosed subject matter: (a) CVD graphene synthesis in quartz tubing furnace and (b) CVD graphene transfer onto the substrate.

Figure 15 : Characterization of a graphene sheet according to some embodiments of the disclosed subject matter: (a) AFM micrograph; (b) Height profile; and (c) Raman spectra (532 nm laser excitation).

Figure 16: A plot showing the characterization of the graphene nanosensor by measuring changes in graphene conductance with varying free aptamer concentrations (0-1000 nM).

Figure 17: A plot showing the measurements with standard mixtures of free aptamer at 25 nM and sample AVP at varying concentrations (1-100 nM).

Figure 18: Fluorescence based time resolved measurements of aptamer-based specific enrichment during continuous introduction of a 100 pM TAMRA-labeled AVP (TVP) solution at 37 °C, followed by infusing a free aptamer solution at 55 °C for thermally-activated release of AVP.

Figure 19: A plot showing the bead-control testing. Integrated devices with and without aptamers on microbeads were tested to verify aptamer-based specific enrichment.

Figure 20: A plot showing the normalized signals obtained from testing with varying sample AVP concentrations (1-500 pM). Three independent testings were conducted at each concentration.

DETAILED DESCRIPTION

The disclosed subject matter provides for devices and techniques to monitor target analytes. More specifically, the disclosed subject matter provides for field-effect transistor ("FET")-based sensors and systems that can be used with aptamers that bind to targets with high specificity and affinity.

As used here, the term "coupled" refers to the connection of a system component to another system component by any suitable means known in the art. The type of coupling used to connect two or more system components can depend on the scale and operability of the system. For example, and not by way of limitation, coupling of two or more components of a system can include connecting the nanosensor to the substrate platform or coupling the enrichment chamber to the sensing chamber via the channel.

As used herein, the term "analyte" is a broad term and is used in its ordinary sense and includes, without limitation, any target the presence or concentration of which is sought in a sample by the sensors and systems disclosed herein. For example, the analyte(s) include, but not are limited to, small organics, peptides, proteins, cells, and tissues.

As used herein, the term "functionalized" means to have a capability of being reactive to an analyte. For example, functionalized refers to a substrate that has a substance attached, conjugated, or grafted to it, wherein the substance has a functional group that is capable of reacting with an analyte. For example, the substance can be covalently attached, conjugated, or grafted to the surface of the functionalized substrate.

As used herein, the term "aptamer" means a stable single-stranded oligonucleotide (e.g., DNA, RNA, Xeno Nucleic Acid (XNA)), or peptide that binds to a specific target analyte. In certain embodiments, the aptamer binds with high affinity and specificity to the target analyte.

The disclosed subject matter provides a graphene nanosensor for monitoring a target analyte. In certain embodiments, the graphene nanosensor utilizes a single conductance sensor on a substrate platform, wherein the graphene sensor can be functionalized with aptamers for binding the target analyte. In certain embodiments, the graphene nanosensor utilizes microbeads functionalized with aptamers which can allow for selective enrichment and isocratic elution of the target analyte. In certain embodiments, the concentration of the enriched target analyte can be measured on a graphene surface functionalized with a target analyte of interest.

In certain embodiments, the sensor can be made of graphene. Graphene is a flat monolayer of carbon atoms tightly packed into a two-dimensional honeycomb lattice. In certain embodiments, the FET sensing element with graphene as the conducting channel has an electric resistance of about 0.1 kQ - about 3 kQ. In certain embodiments, graphene as the conductance channel has an electric resistance of about 0.1 kQ - about 3 kQ, about 0.25 kQ - about 2.75 kQ, about 0.5 kQ - about 2.5 kQ, about 0.75 kQ - about 2.25 kQ, about 1 kQ - about 2 kQ, about 1.25 kQ - about 1.75 kQ, about 1.5 kQ - about 2 kQ, or about 2 kQ - about 3 kQ. In certain embodiments, graphene as the conductance channel has an electric resistance of at least about 0.1 kQ, at least about 0.2 kQ, at least about 0.3 kQ, at least about 0.4 kQ, at least about 0.5 kQ, at least about 0.6 kQ, at least about 0.7 kQ, at least about 0.8 kQ, at least about 0.9 kQ, at least about 1 kQ, at least about 1.2 kQ, at least about 1.4 kQ, at least about 1.6 kQ, at least about 1.8 kQ, at least about 2 kQ, at least about 2.2 kQ, at least about 2.4 kQ, at least about 2.6 kQ, at least about 2.8 kQ, at least about 3 kQ, at least about 4 kQ, at least about 5 kQ, at least about 6 kQ, at least about 7 kQ, at least about 8 kQ, at least about 9 kQ,or at least about 10 kQ. In certain embodiments, graphene as the conductance channel has an electric resistance of no more than about 0.1 kQ, no more than about 0.2 kQ, no more than about 0.3 kQ, no more than about 0.4 kQ, no more than about 0.5 kQ, no more than about 0.6 kQ, no more than about 0.7 kQ, no more than about 0.8 kQ, no more than about 0.9 kQ, no more than about 1 kQ, no more than about 1.2 kQ, no more than about 1.4 kQ, no more than about 1.6 kQ, no more than about 1.8 kQ, no more than about 2 kQ, no more than about 2.2 kQ, no more than about 2.4 kQ, no more than about 2.6 kQ, no more than about 2.8 kQ, no more than about 3 kQ, no more than about 4 kQ, no more than about 5 kQ, no more than about 6 kQ, no more than about 7 kQ, no more than about 8 kQ, no more than about 9 kQ, or no more than about 10 kQ. In certain embodiments, graphene as the conductance channel has an electric resistance of no more than about 3 kQ.

The graphene monoatomic sheet can be thin (0.34 nm), possess strong mechanical strength (Young's modulus: 1 TPa), adhere to underlying substrates, and be flexible, optically transparent, and chemically stable. Such properties can enable new, transformative methods for detection of biological analytes. The monoatomic structure as well as electric conductivity (-1738 S/m) and charge carrier mobility (2xl0⁵ cm2/Vs) of graphene can be exploited to enable sensitive analyte detection.

In certain embodiments, the graphene sensor includes a single layer sheet. In certain embodiments, the graphene sensor includes a multilayered sheet. In certain embodiments, the graphene sensor includes at least one layer of graphene. In certain embodiments, the graphene sensor includes at least two layers of graphene, at least three layers of graphene, or at least four layers of graphene. In certain embodiments, the graphene sheet can be formed by mechanical exfoliation, chemical exfoliation, chemical vapor deposition, or silicon carbide. In certain embodiments, the graphene sheet can be formed by chemical vapor deposition ("CVD"). In certain embodiments, the graphene sheet can be formed by mechanical exfoliation, which can include the removal of a layer of graphene from a block of graphite using tape or other sticky substance. Exemplary techniques for fabrication of the graphene sheet is illustrated in Figure 14 and discussed in further detail in Example 2.

In certain embodiments, the nanosensor can be constructed on a substrate platform. In using graphene as a functional material on a substrate, the nanosensors disclosed herein can be miniaturized and mechanically flexible for placement on a substrate platform. In certain embodiments, the substrate platform can be rigid. As used herein, "rigid" means self-supporting over its length when held at an edge. In certain embodiments, the substrate platform can be flexible. As used herein, "flexible" refers to a material that is not rigid and that conforms to the surface of whatever object the material contacts.

A graphene nanosensor as disclosed herein can be used in several ways. For example, binding of the target analyte with the aptamer immobilized on graphene can cause a change in surface charge density, which can penetrate the atomically thin graphene to significantly change the graphene' s conductance, leading to a detectable signal even at low analyte concentrations.

In certain embodiments, the nanosensor as disclosed herein can enable reliable monitoring of a target analyte in a sample. In certain embodiments, the sample can be a bodily fluid, a non-bodily fluid liquid, or a laboratory sample. In certain embodiments, the nanosensor can be used to measure the amount or change in the amount of a target analyte in a sample. For example, the nanosensor can be part of a standalone device that monitors the target analyte in a sample added to the device (e.g., a piece of lab equipment or home monitor). In certain embodiments, the bodily fluid can be tears, blood, serum, saliva, mucus, interstitial fluid, spinal fluid, intestinal fluid, amniotic fluid, lymphatic fluid, pericardial fluid, peritoneal fluid, pleural fluid, semen, vaginal secretions, sweat, or synovial fluid of the subject.

The disclosed method can be applied to the diagnosis testing. For example, the development of nanosensing systems can be used to detect analytes, such as the presence of organics, peptides, proteins, cells, tissues, or the like in a sample. For example, but not by way of limitation, the target analyte can be immunoglobulin molecules (e.g., IgA, IgD, IgE, IgG, IgM), hormones (e.g., AVP, ADH, glucagon, amylin, GIP, GLP, epinephrine, Cortisol, and growth hormone, estrogens (e.g., estradiol, strone, estriol), calcitriol, progesterone, UPL, LH, FSH, PRL, oxytocin, androgens, DHEA, DHT, orexin, EPI, MT, T₃, T₄, prostalandins, leukotrienes, prostacylin, thromboxane, Insulin, leptin, inhibin, PTH, somatostatin, thrombopoietin, TSH, TRH, VIP, sectetin, relaxin, renin, GnRH, GHRH, HAMP, or the like), complement factors (e.g., Complement factor C3, complement factor C4A (Anaphylatoxin) complement factor C9, complement factor D (CFAD), or the like), enzymes (e.g., lysozyme, carboxypeptidase N, pancreatic trypsin inhibitor, enzyme inhibitors: cy statin C, plasminogen, a-2-antiplasmin, inter-a-trypsin inhibitor complex component II (ITI2), a- 1 -antitrypsin, hexokinase type II, ribonuclease, or the like), transport proteins (e.g., transthyretin, serotransferrin, retinol binding protein (RJ P), transforming growth factor-binding protein (TGF-BP), insulinlike growth factor- binding protein (IBP3), or the like), or plasma proteins (e.g., albumin, fibrinogen A (RGD Peptides), fibrinogen B, a- 1 -microglobulin, b-2-microglobulin, zinc-a-2- glycoprotein (ZAG), a-2-HS-glycoprotein (fetuin), serum amyloid A protein (SAA), haptoglobin, profilin, vitronectin, desmocollin, thymosin b4, apolipoprotein C-III, uteroglobin, ubiquitin, gelsolin, somatomedin B, hemopexin, or the like), oligo nucleotides, glucose. The disclosed method can be applied to the diagnosis of disease, for example, to detect analytes, such as human viruses and bacteria, as many of those microorganisms carry glycoproteins on the exterior surface that can be targeted by the aptamers. Additionally, the disclosed techniques can be applied to a noninvasive method for monitoring cancer treatment. For example, aptamers can be designed to bind to target analytes associated with the cancer cell itself. In certain embodiments, the aptamers can be specific for apoptosis of the cancer cells. In certain embodiments, the microdevice can be used to track overall cell apoptosis to monitor when levels of cell death before they become too toxic. In certain embodiments, the microdevice is able to detect the level of cancer cells in the bodily fluid.

Furthermore, the methods disclosed herein can be used for determining the drug distribution of a treatment regimen in order to maximize the therapeutic effects of the drug. In such methods, the aptamers of the device can be designed to bind to the drug found in the bodily fluid. The design parameters can be varied without departing from the scope of the disclosed subject matter. Such design parameters include, for example, the number and dimensions of atomic layers and substrate platform layers, shape and dimensions of the graphene, and molecular weight of the aptamers, linkers, and polymers.

Aptameric Graphene Nanosensor

In certain embodiments, a graphene nanosensor utilizes a single conductance sensor on a substrate platform, wherein the graphene sensor can be functionalized with aptamers for binding the target analyte. The graphene nanosensor can be functionalized with the aptamer via a linker. In certain embodiments, the nanosensor can include a nanolayer of polymer. In certain embodiments, the polymer can be conjugated to a linker attached to the graphene and also conjugated to an aptamer.

In certain embodiments, the aptamer can bind specifically to the target analyte. The aptamer can bind reversibly or irreversibly to the target analyte. The aptamer can bind to more than one target analyte. The graphene can be functionalized with at least one type of aptamer. The graphene can be functionalized with two or more types of aptamers specific for either the same or different analyte. In certain embodiments, target analyte binding with the aptamer changes the charge density on the nanosensor surface, inducing changes in the carrier concentration of the nanosensor. With reference to Fig. 1, the nanosensor can be configured as a solution-gated graphene-based FET ("GFET") in that the graphene layer (101) is the conducting channel, which can be formed between two electrodes (source (102) and drain (103)) on an substrate platform (104). The substrate platform can be one or multi -layered. For example, the substrate platform can include two layers (see 104) such as, but not limited to a silicon wafer based device. An example of a silicon wafer based device, the lower layer can be silicon as the substrate while the upper layer can be silicone oxide which can serve as an insulating layer. In certain embodiments, the substrate platform can be a single layer. In certain embodiments the single later substrate platform can be a polymer substrate.

In certain embodiments, the graphene layer can be formed between the source and drain electrodes. In certain embodiments, the graphene layer lies over at least a portion of the source and drain electrodes.

In certain embodiments, the graphene surface (101a) can be functionalized with aptamers (105a) via π-π stacking interactions between the linker (105b) attached to the aptamer (105a) and the graphene surface (101a) (Fig. 1). The graphene (101) can be modified with a polymer layer (105c) that can be serially functionalized between the linker (105b) and aptamer (105a) layers. In certain embodiments, the polymer layer can be a nanolayer. The linkers, polymers, and aptamers are further discussed below.

In certain embodiments, the analyte sample solution (106) can be added above the aptamer or polymer/aptamer modified graphene in a polymeric well (107). In certain embodiments, a wire electrode can be inserted to the solution as a gate electrode (108). In this configuration, an electrical double layer ("EDL") can be formed at the interface of the graphene and solution serving as the gate dielectric layer.

During operation and with reference to Fig. lb, the modified graphene nanosensor (Fig. lb(i)) can be incubated with the sample (106) containing the target analyte (109) (Fig. lb(ii)). After removal of the sample followed with buffer washes (Fig. lb(iii)), the target analyte can be purified from the sample by the aptamer (105a) that binds the target analyte. In certain embodiments, the aptamer can be functionalized along with a polymer (105c), which can reduce the nonspecific adsorption of background molecules (110) to the graphene surface (101a).

The transfer characteristics of graphene can be measured in buffer by sweeping the gate voltage. A bias drain-source voltage ( ds) can be applied between the drain (103) and source (102) electrodes which can generate a drain-source current (Ids) through the graphene channel, and the ids can be controlled by the applied gate voltage (Vg). This can yield the transfer characteristics of the GFET (i.e., the functional dependence of I^ on V_g). When the target analyte (109) in a sample (106) can be captured by the aptamer (105a), the charged target analyte can alter the carrier concentration of the graphene (101), and thus induce a shift of the Dirac point (Vuimc), the voltage at which the ids reaches its minimum.

In certain embodiments, the electrode wire/gate electrode can be, for example, but not limited to ITO, Ti/Pd/Pt, Ag/AgCl, Ag, Pt, Cu, Cr, or combinations thereof.

In certain embodiments, the source and drain electrodes, are separately made from material such as, but not limited to ITO, Ti/Pd/Pt, Ag/AgCl, Ag, Pt, Cu, Cr, or combinations thereof.

In certain embodiments, the polymeric well can be made of a silicone substance, such as a polymeric organosilicon compound. In certain embodiments, the polymeric organosilicon compound can be polydimethylsiloxane ("PDMS").

In certain embodiments, the substrate support can be made from material such as, but not limited to polyethylene terephthalate ("PET"), polycarbonate polystyrene, polymethyl methacrylate ("PMMA"), polymacon, silicones, Si0₂, Si0₂- coated silicon wafer, fluoropolymers, silicone acrylate, fluoro-silicone/acrylate, poly hydroxyethyl methacrylate, Hf0_2; parylene, Si₃N₄ or combinations thereof.

In certain embodiments, the dimensions of the graphene conducting channel can be a length of about 10 μιη to about 20 μιη by a width of about 10 μιη to about 20 μιη by a thickness of the graphene layer. In certain embodiments, the graphene conducting channel can have a length of about 5 μιη to about 15 μιτι, about 7.5 μιη to about 22.5 μιτι, about 10 μιη to about 20 μιτι, or about 12.5 μιη to about 17.5 μιη. In certain embodiments, the graphene conducting channel can have a width of about 5 μιη to about 15 μιτι, about 7.5 μιη to about 22.5 μιτι, about 10 μιη to about 20 μιη, or about 12.5 μιη to about 17.5 μιη. In certain embodiments, the graphene channel in the sensor can be about 20 x 10 μιη² with a thickness of about 0.34 nm

Exemplary techniques for fabrication of the devices is illustrated in, for example, Figures 3 and 4 and certain embodiments will be discussed in further detail in Examples 1 and 2.

With reference to Fig. 3, in certain embodiments, the device can be fabricated using micro and nanofabrication techniques on a substrate platform (300), such as an oxidized silicon wafer. After cleaning the substrate platform, a layer of the material for the source and drain electrodes can be deposited on the substrate (301). The electrodes can be deposited by any technique known to one skilled in the art. For example, the material can be placed using thermal evaporation. Next, a layer of photoresist can be added (e.g. spin-coated) on top of the electrode material (302) and then baked (303). For example, the material can be baked at 1 15 °C for 1 min. The source and drain electrodes can be patterned (304). The electrodes can be patterned by any technique known to one skilled in the art, such as but not limited to photolithography. The wafer can be developed (305) and then etched in gold (306) and chrome etchant (307) sequentially. Graphene can be synthesized via methods known to those of skill in the art. For example, the graphene layer can be synthesized using chemical vapor deposition ("CVD") substrate (see Example 2 for an example protocol) and transferred on top of the source and drain electrodes (see Figs, la, lb, and 3) (308). A polymeric well (e.g., 107) can be placed above the graphene sensing region (309) that could be filled with a sample. A wire can be inserted into the solution to apply the gate voltage (e.g., 108).

With reference to Fig. 4, in certain embodiments, to achieve functionalization of the graphene surface (400), the GFET can be immersed with a solution of the linker (401). For example, the GFET can be immersed with a solution of the linker for at least 1, at least 2, at least 3, at least 4, or at least 5 hours at room temperature. After rinsing thoroughly with ethanol and then with phosphate buffered saline ("PBS") buffer (402), the GFET can optionally be immersed in a solution of polymer (403). For example, the GFET can be immersed with a solution of polymer for at least 2 hours, at least 4 hours, at least 6 hours, at least 8 hours, at least 10 hours, at least 12 hours, at least 14 hours, at least 16 hours, at least 18 hours, or at least 20 hours at room temperature. In certain embodiments, the GEFT can then be immersed with a mixture of l-ethyl-3-(3-dimethylaminepropyl) carbodiimide hydrochloride (EDC-HC1) and N-hydroxysulfosuccinimide ("NHS") (e.g., at a pH=6 in PBS solution) to activate the free end of the polymer (404). In other embodiments, the polymer activated using solutions such as, but not limited to, DIC/DIPC, DCC, HATU, HBTU, HCTU, BOP, PyOxim, or T3P. For example, the GFET can be immersed with a solution of EDC^»HC1 and NHS for at least 1, at least 2, at least 3, at least 4, or at least 5 hours at room temperature. Finally, the GFET can be immersed in a solution of aptamer (405). For example, the GFET can be immersed with a solution of aptamer for at least 2 hours, at least 4 hours, at least 6 hours, at least 8 hours, at least 10 hours, at least 12 hours, at least 14 hours, at least 16 hours, at least 18 hours, or at least 20 hours at room temperature.

In certain embodiments, the aptamer is attached directly to the linker.

For example, the aptamer can be an amino group-terminated aptamer which is directly attached to the linker by forming an amid bond with the carboxylic group on the linker.

Graphene Nanosensor with Integrated Selective Enrichment

In certain embodiments, the graphene nanosensor utilizes microbeads functionalized with aptamers. In certain embodiments, aptamers which can allow for selective enrichment and isocratic elution of the target analyte. In certain embodiments, the aptamer can bind specifically to the target analyte. In certain embodiments, the aptamer can bind reversibly to the target analyte. In certain embodiments, the concentration of the enriched target analyte can be measured on a graphene surface functionalized with a target analyte of interest.

In certain embodiments, the microbeads are functionalized with at least one type of aptamer. In certain embodiments, the graphene nanosensors use at least one set of microbeads functionalized with one type of aptamer. In certain embodiments, the graphene nanosensors use at least two sets of microbeads, where each set of microbeads can be functionalized with a different set of aptamers specific for either the same or different analyte.

In certain embodiments, the nanosensor can be configured as a FET in that the graphene (501) can be the conducting channel, which is formed between two electrodes (source (502) and drain (503)) on an graphene nanosensor substrate platform (504) (Fig. 15). A sweeping voltage of about 0-1 V can be applied to the gate electrode (515) forming a gate voltage (V_gs) through a sample solution. In certain embodiments, the electrical conductance through graphene (501), which depends on the charge on the graphene surface, can be measured from the drain current (I_ds) at a fixed drain voltage (!¾. Iⁿ certain embodiments, the graphene surface can be fictionalized with target analyte. In certain embodiments, the graphene layer can be formed between the source and drain electrodes. In certain embodiments, the graphene layer lies over at least a portion of the source and drain electrodes.

In certain embodiments, the integrated microdevice includes at least one enrichment microchamber (505), at least one graphene-based sensing microchamber (506) combined with at least one graphene nanosensor (501), at least three temperature control units (507), and at least one flow gate (508) connected to at least one serpentine channel (509) for sample (510) transfer to the sensing microchamber (506). As used herein, a serpentine channel can include a channel with several segments that alternate in direction. In certain embodiments, the channels may a have at least 4, at least 6, at least 8, or at least 10 alternating sections. In certain embodiments, the channel is a straight channel. In certain embodiments, a weir structure (51 1) in the at least one enrichment microchamber (505) retains microbeads (512) within the enrichment microchamber (505) during the selective enrichment process. A magnet can hold the microbeads in the enrichment microchamber (505).

In certain embodiments, the microdevice can include at least three temperature control units (507) which can include a resistive heater (513) and a temperature sensor (514) separately control temperatures inside the at least one enrichment (505 (heater 513a; sensor 514a)) and sensing (506 (heater 513b; sensor 514b)) microchambers and the at least one serpentine channel (509 (heater 513c; sensor 514c)) for forming or disrupting aptamer-target analyzte complexes. The microchambers (505 and 506) and the channels (e.g., 509) can be formed by a polymeric material (516). In certain embodiments, the polymeric material also includes at least one buffer/sample inlet (518), at least one waste outlet (519), and at least one sample outlet (520). In certain embodiments, the microdevice can be formed on a support platform (517). In certain embodiments, an inlet (518) can be coupled to the enrichment chamber, a waste outlet (519) can be coupled to the serpentine channel, and a sample outlet (520) can be coupled to the sensing chamber.

In certain embodiments, the microdevice can have at least 1, at least 2, at least 3, at least 4, at least 5, at least 6, at least 7, at least 8, at least 9, or at least 10 enrichment chambers. In certain embodiments, each enrichment chamber can have its own separate temperature control unit.

In certain embodiments, the microdevice can have at least 1, at least 2, at least 3, at least 4, at least 5, at least 6, at least 7, at least 8, at least 9, or at least 10 sensing chambers. In certain embodiments, there can be a sensing chamber for each enrichment chamber. In certain embodiments, there can be two sensing chambers for each enrichment chamber. In certain embodiments, each sensing chamber can have its own separate temperature control unit. In certain embodiments, valves (e.g., pneumatic) control the movement of the sample and/or buffer to the different chambers.

In certain embodiments, the dimensions of the enrichment chamber can be a length of about 1 mm to about 5 mm by a width of about 1 mm to about 5 mm by a height of about 100 μιη to about 1 mm. In certain embodiments, the enrichment chamber can be a length of about 2 mm by a width of about 2 mm by a height of about 200 μιη. In certain embodiments, the length can be from about 1.25 mm to about 4.75 mm, about 1.5 mm to about 4.5 mm, about 1.75 mm to about 4.25 mm, about 2 mm to about 4 mm, or about 2.25 to about 3.75 mm. In certain embodiments, the width can be from about 1.25 mm to about 4.75 mm, about 1.5 mm to about 4.5 mm, about 1.75 mm to about 4.25 mm, about 2 mm to about 4 mm, or about 2.25 to about 3.75 mm. In certain embodiments, the height can be from about 200 μιη to about 900 μιη, about 300 μιη to about 800 μιτι, about 400 μιη to about 700 μιτι, or about 500 μιη to about 600 μιη.

In certain embodiments, the enrichment chamber can have a volume of about 0.1 μΐ to about 25 μΐ. In certain embodiments, the enrichment chamber can have a volume of about 0.5 μΐ to about 20 μΐ, about 1 μΐ to about 15 μΐ, or about 5 μΐ to about 10 μΐ. In certain embodiments, the enrichment chamber can have a volume of about 0.1 μΐ to about 1 μΐ, about 0.2 μΐ to about 0.9 μΐ, about 0.3 μΐ to about 0.8 μΐ, about 0.4 μΐ to about 0.7 μΐ, or about 0.5 μΐ to about 0.6 μΐ.

In certain embodiments, the dimensions of the sensing chamber can be a length of about 2 mm to about 5 mm by a width of about 2 mm to about 5 mm by a height of about 1 mm to about 3 mm. In certain embodiments, the dimensions of the sensing chamber can be a length of about 2.25 mm to about 4.75 mm, about 2.5 mm to about 4.5 mm, about 2.75 to about 4.25, about 3 mm to about 4 mm, or about 3.25 mm to about 3.75 mm. In certain embodiments, the dimensions of the sensing chamber can be a width of about 2.25 mm to about 4.75 mm, about 2.5 mm to about 4.5 mm, about 2.75 to about 4.25, about 3 mm to about 4 mm, or about 3.25 mm to about 3.75 mm. In certain embodiments, the dimensions of the sensing chamber can be a height of about 1.25 mm to about 2.75 mm, about 1.5 mm to about 2.5 mm, or about 1.75 mm to about 2.25 mm. In certain embodiments, the sensing chamber can be a length of about 2 mm by a width of about 2 mm by a height of about 1 mm.

In certain embodiments, the sensing chamber can have a volume of about 4 μΐ to about 7 μΐ. In certain embodiments, the sensing chamber can have a volume of about 4.25 μΐ to about 6.75 μΐ, about 4.5 μΐ to about 6.5 μΐ, about 4.75 μΐ to about 6.25 μΐ, about 5 μΐ to about 6 μΐ, or about 5.25 μΐ to about 5.75 μΐ.

In certain embodiments, the dimensions of the serpentine channel can be a width of about 500 μιη to about 1 mm by a height of about 20 μιη to about 50 μιη. In certain embodiments, the serpentine channel can be a width of about 500 μιη by a height of about 20 μιη. In certain embodiments, the serpentine channel can have a length of about 4 cm to about 10 cm. In certain embodiments, the dimensions of the serpentine channel can be a width of about 550 μιη to about 950 μιτι, about 600 μιη to about 900 μιη, about 650 μιη to about 850 μιτι, or about 700 μιη to about 800 μιη. In certain embodiments, the dimensions of the serpentine channel can be a height of about 22.5 μιη to about 47.5 μιτι, about 25 μιη to about 45 μιτι, about 27.5 μιη to about 42.5 μιη, about 30 μιη to about 40 μιτι, or about 32.5 μιη to about 37.5 μιη. In certain embodiments, the serpentine channel can have a length of about 4.5 cm to about 9.5 cm, about 5 cm to about 9 cm, about 5.5 cm to about 8.5 cm, about 6 cm to about 8 cm, or about 6.5 cm to about 7.5 cm.

In certain embodiments, the weir structure can be from about 1/5 to about 1/20 the height of the enrichment chamber. In certain embodiments, the weir structure can be from about 1/6 to about 1/19, from about 1/7 to about 1/18, from about 1/8 to about 1/17, from about 1/9 to about 1/16, from about 1/10 to about 1/15, or from about 1/11 to about 1/14 the height of the enrichment chamber. In certain embodiments, the dimensions of the weir structure can be slightly smaller than the microbeads. In certain embodiments, the weir structure can be from about 1/2 to about 1/4 the size of the microbeads. In certain embodiments, the weir structure can be from about 10 μιη to about 70 μιτι, about 20 μιη to about 60 μιτι, or about 30 μιη to about 50 μιη. In certain embodiments, the weir structure can be about 20 μιη.

In certain embodiments, the microbeads can have a diameter of about 40 μιη to about 90 μιτι, about 50 μιη to about 80 μιτι, or about 60 μιη to about 70 μιη.

In certain embodiments, the gate electrode can be made from material such as, but not limited to ITO, Ti/Pd/Pt, Ag/AgCl, Ag, Pt, Cu, Cr, or combinations thereof.

In certain embodiments, the substrate and/or support platform can be made from material such as, but not limited to polyethylene terephthalate ("PET"), polycarbonate polystyrene, polymethyl methacrylate ("PMMA"), silicones, silicon wafers, Si0₂, Si0₂-coated silicon wafer, fluoropolymers, silicone acrylate, fluoro- silicone/acrylate, poly hydroxyethyl methacrylate, Hf0_2; parylene, Si₃N₄ or combinations thereof.

In certain embodiments, the polymeric material can be made of a silicone substance. In certain embodiments, the silicone substance can be a polymeric organosilicon compound. In certain embodiments, the polymeric material can be polydimethylsiloxane ("PDMS").

In certain embodiments, the temperature control unit can be a Cr/Au or

Cr/Pt resistive microheater and temperature sensor.

Exemplary techniques for fabrication of the devices are illustrated in, for example, Fig. 6 and certain embodiments will be discussed in further detail in Examples 2 and 3.

With reference to Fig. 6, in certain embodiments, the device can be fabricated using micro and nanofabrication techniques on a substrate platform (600), such as an oxidized silicon wafer. After cleaning the substrate platform, a layer of the material for the source and drain electrodes can be deposited on the substrate platform (601). The electrodes can be deposited by any technique known to one skilled in the art. For example, the material can be placed using thermal evaporation, graphene can be transferred to connect the source and drain electrodes (602). In certain embodiments, one layer of graphene is transferred to connect the source and drain electrodes. = Graphene can be synthesized via methods known to those of skill in the art. For example, the graphene layer can be synthesized using chemical vapor deposition ("CVD") substrate (see Example 2 for an example protocol). The graphene surface can be functionalized with control (e.g., reagent) target analyte (603). The graphene can be functionalized by any technique known to one skilled in the art. For example, the graphene can be immersed with a solution of the linker, for at least 1, at least 2, at least 3, at least 4, or at least 5 hours at room temperature. After rinsing the graphene can be immersed with a solution of target analyte for least 1, at least 2, at least 3, at least 4, or at least 5, at least 6 hours, at least 7, at least 8 hours, at least 9, at least 10 hours, at least 12 hours, at least 14 hours, at least 16 hours, at least 18 hours, or at least 20 hours at room temperature.

The microchambers and channels can be fabricated in polymeric material (604). The microchambers and channels can be patterned by any technique known to one skilled in the art, such as but not limited to, soft photolithography. Holes can be punched into the polymeric material to create inlets and outlets for buffer and sample (605). The sheet can then be bonded onto a support platform on which a resistive microheater and temperature sensor was placed (e.g., patterned via photolithography) (606). The graphene sensor chip can be stacked and bonded on the polymeric layer, while the functionalized graphene region faces the open sensing microchamber (607).

Fig. 7 describes a method of target analyte measurement (700) according to some embodiments. In certain embodiments, the microbeads (702) can be functionalized with at least one type of aptamer (704). The surface of the graphene (701) can be functionalized with control (e.g., reagent) sample (708) of the target analyte. The graphene surface can be functionalized with the control sample (708) of the target analyte via a linker (709). In certain embodiments, the linker can be 1-pyrenebutanoic acid succinimidyl ester. The linkers and aptamers are discussed in further detail below.

In certain embodiments, the device can initially be rinsed with buffer (e.g., about 10 μΐ/min), and the sensing microchamber (e.g., 506) can be filled with a buffer solution to measure a reference signal (7ds,ref)- In certain embodiments, a sample solution containing the target analyte (703) can be introduced into the enrichment microchamber (e.g., 505). In certain embodiments, the sample solution can be introduced into the enrichment microchamber using a syringe, syringe pump, or peristaltic pump. In certain embodiments, the enrichment microchamber can be maintained between about 25 °C to about 60 °C using a temperature control unit (e.g., 507a) an integrated microheater and a temperature controlled sensor) integrated on the support platform (e.g., 517). In certain embodiments, the enrichment microchamber can be maintained at about 37 °C. During this stage, the target analyte (703) can be selectively captured by an aptamer (704) immobilized on a microbead (702) surface, and can be enriched by continuous sample infusion into the device.

Following the enrichment, the chamber can be rinsed with a buffer solution (e.g., about 10 μΐ/ min) to remove the impurities (706) at the same temperature the enrichment microchamber is held at following the introduction of the sample. In certain embodiments, the buffer enters the buffer/sample inlet (e.g., 518) and passes out the waste outlet (e.g., 519). In certain embodiments, the sample outlet is closed so that the sate goes through the waste outlet. For example, the sample outlet tube can be clamped or closed by other mechanical means.

After the removal of nonspecifically adsorbed target analyte, molecules, and impurities (706), the enriched target analytes can be released at an elevated temperature and eluted with free aptamers (707), causing isocratic elution of target analytes (703) with the free aptamer (707). The waste outlet can be closed, a plug (i.e., a small volume of solution ,e.g., 4 - 75 μΐ). of the free aptamer (707) solution (i.e., an aptamer solution unattached to any support) can be flowed through the enrichment microchamber at an elevated temperature. The aptamer solution can be at a concentration of about 5 nM to about 50 nM, about 10 nM to about 45 nM, about, about 15 nM to about 40 nM, about 20 nM to about 35 nM, or about 25 nM to about 30 nM.

In certain embodiments, the aptamer solution can be at a concentration of about 25 nM. In certain embodiments, the amount of aptamer solution can be about 5 μΐ to about 15 μΐ or about 7.5 μιη to about 12.5 μΐ. The flow rate of the aptamer solution can be about 10 μΐ. The flow rate of the aptamer solution can be about 5 μΐ/min to about 15 μΐ/min, or about 7.5 μιη/ιηίη to about 12.5 μΐ/min. The flow rate of the aptamer solution can be about 10 μΐ/min. The enrichment microchamber can be maintained at a temperature of about 25 °C to about 60 °C while the aptamer solution is introduced. The enrichment microchamber can be maintained at a temperature of about 55 °C while the aptamer solution is introduced.

The mixture of eluted sample target analytes (703) and the free aptamer (707) can then be transferred to the sensing microchamber (e.g., 506). The sensing microchamber can be kept at the same temperature as the enrichment chamber when the sample was first introduced. The sensing microchamber can be maintained between about 25 °C to about 60 °C using a temperature control unit (e.g., 507b) an integrated microheater and a temperature controlled sensor) integrated on the support platform (e.g., 517). The sensing microchamber can be maintained at about 37 °C.

In certain embodiments, the sample can be continuously infused. The sample can be continuously infused for at least about 20 minuets, at least about 30 minutes, at least about 40 minutes, at least about 50 minutes, at least about 60 minutes, at least about 70 minuets, at least about 80 minutes, at least about 90 minutes, at least about 100 minutes, at least about 110 minutes, at least about 120 minutes, at least about 130 minuets, at least about 140 minutes, at least about 150 minuets, or at least about 160 minuets. The sample can be continuously infused for at least 120 minutes. In alternative embodiments, the sample can be continuously infused for about 30 minutes to about 160 minuets, about 40 minuets to about 150 minutes, about 50 minutes to about 140 minutes, about 60 minutes to about 130 minutes, about 70 minuets to about 120 minutes, about 80 minutes to about 1 10 minuets, or about 90 minutes to about 100 minutes.

In certain embodiments, the binding of the free aptamers to the control target analyte immobilized on the graphene surface changes the charge density on the nanosensor surface, inducing changes in the carrier concentration of the nanosensor. In certain embodiments, the measurement circuit can use a DC power supply (e.g., E3631A, Agilent) to provide the drain voltage ds, a function generator (e.g., 33220A, Agilent) to supply the gate voltage V_gs, and a digit multimeter (e.g., 3441 OA, Agilent) to measure the drain current (/ds)- During the electrical measurement, Ids values (at fixed V_gs and ds) can be automatically collected once per second for a determined period.

Aptamers

In certain embodiments of the disclosed subject matter, the nanosensor can be used to determine the level of a target analyte in the body, for example small organics, peptides, proteins, cells, tissues, or the like. In certain embodiments, the aptamers that identify the target analyte can be a stable single-stranded oligonucleotide (e.g., DNA, RNA, Xeno Nucleic Acid (XNA)) or peptide that binds to a specific target analyte.

Oligonucleotide aptamers are nucleic acid species that have been engineered through repeated rounds of in vitro selection or SELEX (systematic evolution of ligands by exponential enrichment) to bind to various molecular targets.

Peptide aptamers include a short variable peptide domain. In certain embodiments, the variable loop length can be typically composed of ten to twenty amino acids, and the scaffold may be any protein which has good solubility and stability properties. The peptide domain can be attached at both ends to a protein scaffold. The double structural constraint can increase the binding affinity of the peptide aptamer. The double structural constraint can increase the binding affinity of the peptide aptamer to nM range. Peptide aptamer selection can be made using different systems. For example, peptide aptamer selection can occur by, but not limited to, yeast two-hybrid system, selection and combinatorial peptide libraries constructed by phage display and other surface display technologies such as mRNA display, ribosome display, bacterial display and yeast display. Peptide aptamers can also include affimers, which are small, stable proteins engineered to display peptide loops which provides a high affinity binding surface for a specific target analytes. In certain embodiments, the affimers can be low molecular weight (e.g., 12-14 kDa).

The binding between the aptamer and the analyte of interest can be irreversible or reversible. In certain embodiments, the aptamer binds reversibly to the target analyte. For example, the binding and dissociation between the target analyte and the aptamer can be an equilibrium phenomenon driven by the concentration of the analyte in the sample and/or conducting channel. The amount of the analyte bound with the aptamer depends on the concentration of the analyte in the sample and/or conducting channel.

In certain embodiments, the aptamer can be conjugated to a substance that can immobilize the aptamers on the graphene sensor and/or microbead surface (i.e., functionalized surface (e.g., graphene sensor and/or microbead).

In certain embodiments, the aptamers can be irreversibly attached to graphene with strength comparable to a covalent attachment. The linker can be a pyrene-terminated agent. The pyrene-terminated agent can be 1-pyrenebutanoic acid succinimidyl ester ("PASE"). A pyrene-terminated agent can be synthesized that can irreversibly attached to graphene using a sticky point for π-π stacking interactions without disrupting the graphene' s conjugation or altering its electronic properties. The amino group-terminated aptamer can be directly attached to PASE by forming amide bond.

In certain embodiments, the aptamers can be irreversibly attached to microbead. For example, the attachment can occur using interactions between avidin- biotin or NHS ester-amine.

In certain embodiments, the nanosensor can be serially functionalized with a polymer and aptamer. The polymer can minimize the non-specific adsorption of background molecules onto the graphene. In certain embodiments, the nanosensor can have a nanolayer of polymer, and, the polymer can be polyethlene glycol ("PEG"). The PEG molecule can have a molecular mass below about 2000. In certain embodiments, the PEG molecule can have a molecular mass between about 1000 to about 2000. In certain embodiments, the PEG molecule can have a molecular mass between about 1000 to about 1500, about 1100 to about 1400, or about 1200 to about 1300. In certain embodiments, the PEG molecule can have a molecular mass between about 1500 to about 2000, about 1600 to about 1900, or about 1700 to about 1800. In certain embodiments, the PEG is linear. In certain embodiments the PEG is branched. In certain embodiments, the polymer can be conjugated to a linker on one end of the polymer and the aptamer to the other end of the polymer. The linkage for both can be formed by amide bonds.

In certain embodiments, in order to screen out effects not caused by the target analyte, for example, environmental factors such as temperature or other analytes, the reference chamber can include a graphene sensor that is functionalized with the linker and/or linker and polymer without the aptamer (the reference linker). The reference linker does not bind with the target analyte. In certain embodiments, the reference linker is the same linker and/or linker and polymer used to functionalize the graphene support. Also, the reference linker should not bind or react specifically with any other substance in the sample. The reference linker and polymer should, however, respond similarly to non-target analytes and environmental conditions.

In certain embodiments, in order to screen out effects not caused by the target analyte, control testing can be carried out with a control molecule (i.e., a molecule that is not the target analyte). The graphene nanosensor can be exposed to samples spiked with various concentrations of the control molecule. The transfer characteristics can be examined to see if there is a shift introduced by non-specific binding of control molecules to the graphene surface.

Additional aspects and embodiments of the disclosed subject matter are illustrated in the following examples, which are provided for better understanding of the disclosed subject matter and not limitation.

Example 1. An aptameric graphene nanosensor for analyte detection

Illustrated herein is an aptameric graphene nanosensor that can allow accurate detection of a target analyte (e.g., 109). Presented is an aptameric graphene field-effect transistor ("GFET") nanosensor for sensitive and label-free detection of biomarkers in serum. The graphene nanosensor was serially functionalized with a polyethylene glycol ("PEG") nanolayer and an aptamer for specific detection of a target analyte and effective rejection of background molecules.

This examples presents a graphene nanosensor for the detection of immunoglobulin E ("IgE") as a representative analyte. IgE is an antibody that plays a role in mammalian immune defense eliciting protective response including local inflammation, itching, and mucus production. The PEG reduced nonspecific adsorption of serum background molecules onto graphene. The binding of IgE with the aptamer on the surface induced changes in the carrier density and mobility in the bulk of the graphene, generating a detectable signal. Experimental results demonstrated the specific IgE detection in serum in a physiologically relevant range of 50 pM to 35 nM. Hence this nanosensor can be utilized in clinical diagnostics.

An example graphene nanosensor (e.g., 100) was configured as an electrolyte-gated FET where graphene (e.g., 101) was the conducting channel, formed between drain (e.g., 103) and source electrodes (e.g., 102) on the Si0₂ substrate platform (e.g., 104) (Fig. 1). Pyrene group terminated 1-pyrenebutanoic acid succinimidyl ester ("PASE") (e.g., 105b)) was attached to graphene (e.g., 101a) via π- π stacking. Then, PEG (e.g., 105c) was conjugated to PASE (e.g., 105b), and the aptamer (e.g., 105a) to the free end of PEG (e.g., 105c), both by forming amide bonds. The analyte sample (e.g., 106) was held above the PEG (e.g., 105c) and aptamer (e.g., 105a) modified graphene (e.g., 101a) in a poly dimethyl siloxane ("PDMS") well (e.g., 107), and an Ag/AgCl wire was inserted to the solution as a gate electrode (e.g., 108). In this embodiment, an electrical double layer ("EDL") was formed at the interface of the graphene and solution serving as the gate dielectric layer.

During operation, the modified graphene nanosensor (Fig. lb) was incubated in serum (e.g., 106) containing IgE (e.g., 109) (Fig. lb(ii)). After removal of the serum followed with buffer washes (Fig. lb(iii)), the IgE was purified from serum by the anti-IgE aptamer (e.g., 105a) while the functionalized PEG (e.g., 105c) drastically reduces the nonspecific adsorption of background molecules to the graphene surface (e.g., 101a). A bias drain-source voltage (¥&) applied between the drain (e.g., 103) and source (e.g., 102) s generated a drain-source current (I_ds) through the graphene channel, and the Ids was controlled by the applied gate voltage (F_g). This yielded the transfer characteristics of the GFET (z^'.e.,the functional dependence of s on _g). When the IgE (e.g., 109) in a serum sample (e.g., 106) was captured by the anti-IgE aptamer (e.g., 105a), the charged IgE altered the carrier concentration of the graphene (e.g., 101), and thus induced a shift of the Dirac point (Vuimc), the voltage at which the Ids reaches its minimum.

An example device was fabricated using micro and nanofabrication techniques on an oxidized silicon wafer. After cleaning by piranha, a layer of 5/45 nm Cr/Au was deposited using thermal evaporation (e.g., 301). A layer of photoresist was then spin-coated on top of Au layer (e.g., 302) and baked (e.g., 303) at 1 15 °C for 1 min. Photolithography was used to pattern the source and drain electrodes (e.g., 304), and the wafer was then developed (e.g., 305) and etched in gold (e.g., 306) and chrome etchant (e.g., 307) sequentially. Graphene was synthesized via chemical vapor deposition ("CVD") substrate (see Example 2 for an example protocol) and was transferred onto the to connect the source and drain electrodes (Figs, la, lb, and 3) (e.g., 308).

To achieve noncovalent functionalization of the graphene surface (e.g., 400), the GFET was immersed for 2 hours at room temperature in a dimethylformamide ("DMF") solution of 5 mM 1-pyrenebutanoic acid succinimidyl ester ("PASE"), which serves as a linker (e.g., 401). After rinsing thoroughly with ethanol followed by phosphate buffered saline ("PBS") buffer (e.g., 402), the graphene sensor was immersed in a 1 mg/mL solution of PEG (MW: 2,000) in PBS solution for 12 hours at room temperature (e.g., 404). The mixture of l-ethyl-3-(3- dimethylaminepropyl) carbodiimide hydrochloride (EDC^»HC1) and N- hydroxysulfosuccinimide ("NHS") at pH=6 in PBS solution was used to activate carboxylic group at the free end of the PEG for 1 hour (e.g., 405). Finally, a solution of IgE specific aptamer in PBS (concentration: 1 μΜ) was added to the channel of GFET for 12 hours at room temperature (e.g., 406). A rinse step was used between each step.

A PDMS well (-10 L in volume) was placed above the graphene sensing region (e.g., 107) that could be filled with a serum sample. An Ag/AgCl wire was inserted into the solution to apply the gate voltage (e.g., 108) .

In testing, serum samples with different IgE concentrations were introduced into the PDMS well (e.g., 107) for 5 minutes and then washed with PBS buffer, and the transfer characteristics of graphene then were measured in buffer by sweeping the gate voltage.

The graphene was verified to be a single layer via Raman spectroscope (Fig. 8), where a G band at -1580 cm^"1 and 2D band at 2685 cm^"1 in Raman spectrum are characteristics of single-layer graphene.

To characterize the serial functionalization of the graphene surface with PEG and the IgE-specific aptamer, the modified graphene was exposed to a fluorescently labeled single-strand DNA probe that was complementary to the aptamer. Negligible fluorescence was observed on the PEG functionalized graphene while a strong fluorescence signal was observed on the aptamer and PEG modified graphene (Fig. 9), indicating that the PEG and aptamer were successfully immobilized on graphene surface.

The nonspecific adsorption of background molecules to graphene was observed by measuring the transfer characteristics of the bare graphene in buffer before and after the serum incubation. After incubation of the bare graphene in serum, the Dirac point V_Oi_mc shifted to negative gate voltage by 50 mV, which may be caused by the n-doping of the background molecule adsorption to graphene surface. In contrast, the PEG modified graphene did not exhibit appreciable shift of Vuimc after serum incubation. This supports that the nonspecific adsorption was inhibited by the PEG nanolayer (Fig. 10), which was attributed to the improved hydrophilicity and water solubility of graphene by the grafted PEG.

To investigate the response of the graphene nanosensor at varying IgE concentrations, the nanosenor was incubated in serum samples with various IgE concentrations, and the transfer characteristics were measured. The V_{O mc} was observed to consistently shift toward the negative gate voltage from 101 mV to 85 mV as the IgE concentration increased from 50 pM to 35 nM (Fig. 11 and Fig. 12). The IgE is positively charged based on the IgE concentration dependent transfer characteristics. The IgE concentration dependent transfer characteristics can be illustrated by the p-type behavior of the graphene in the FET structure. Graphene is known as a p-type semiconductor with holes as the majority of carrier, such that the charged IgE protein captured by the IgE specific aptamer screens the graphene surface and lowers the carrier concentration in the bulk of graphene and hence reduce the conductance, suggesting n-type doing was induced in graphene. The transconductance was found to be approximately constant regardless of IgE concentrations implying the constant carrier mobility. Therefore, electrostatic interaction was believed to be the dominating effect, otherwise the carrier mobility would be significantly altered. A higher IgE concentration leads to a greater decrease of the carrier concentration and conductance of the graphene (inset in Fig. 11), as described by the transfer characteristics equation (1). ι_Λ = M(ir / L cv y₉ -v_^ ( l ) where C is the top-gate capacitance per unit area, is the carrier mobility, W and L are respectively the width and length of the graphene channel.

To exclude the possibility of nonspecific binding and false signals, control testing was carried out with a control molecule, immunoglobulin G (IgG). The graphene nanosensor was exposed to samples spiked with various IgG concentrations. The transfer characteristic curves showed negligible shifts (Fig. 13), which demonstrated the specific capture of the IgE in human serum and the capability of preventing nonspecific adsorption of the grafted PEG nanolayer. Example 2. Construction of CVD Graphene

CVD graphene (1401) was synthesized by heating annealed Cu foil (1402) in a quartz tubing furnace (Fig. 14). The Cu foil (1402) was first sharply heated to 1000 °C in Argon (Ar) environment (200 mTorr), and annealed in hydrogen (H2) environment (10 mTorr). The mixture of methane (CH₄) and H₂ were then introduced and allowed to react for 18 min (CH₄: 170 mTorr, H₂: 10 mTorr), after which the sample was cooled down to room temperature in Ar flow at200 mTorr and then retrieved from the tube. After graphene growth, 500nm PMMA (1403) was spin coated on top for protection and a PDMS stamp (1405) and glass slide (1404) was attached by pressing. Cu (1402) was removed by wet etching.

Graphene (1401) was transferred onto the substrate (1406) at 170 °C

(Fig. 14b) to realize the graphene conducting channel stretching over drain/source electrodes. After the protective PMMA layer (1403) on graphene surface was dissolved by acetone, AFM (XE-100, Park System) and Raman spectroscopy (Renishaw, 532 nm laser) were used to verify the single-layer graphene sheet (Fig. 15).

Example 3. A Microfluidic Aptasensor Graphene Nanosensor with Integrated Specific Enrichment

Illustrated herein is an aptameric graphene nanosensor that can allow accurate detection of a target analyte (e.g., 500).

This example presents an integrated microfluidic aptasensor which combined aptamer-based specific enrichment and conductance-based graphene nanosensing, allowing detection of a low charged and small biomolecule with high sensitivity. An aptamer specific to a target analyte was immobilized on microbeads for solid phase-based selective enrichment and isocratic elution of the analyte. A FET configuration was used to measure changes in graphene conductance, which were used to determine the analyte concentration. Microfluidic integration of enrichment and sensing on a single chip eliminates the need for off-chip sample handling, thus minimizing potential sample contamination or loss.

For demonstration, arginine vasopressin (A VP), a small peptide and a clinically important biomarker for septic and hemorrhagic shocks, was chosen as a target analyte. Results discussed below (e.g. Fig. 20) show that the integrated aptasensor is capable of detecting label-free AVP at clinically relevant low concentrations (1-500 pM). Hence this nanosensor can be utilized in clinical diagnostics.

An example graphene nanosensor of a FET configuration is shown in Fig. 5. The integrated microdevice consisted of an enrichment microchamber (height: 200 μπι, volume: 2.5 μΐ), a graphene-based sensing microchamber (height: 1 mm, volume: ~5 μΐ) combined with a graphene nanosensor, three temperature control units (Cr/Au resistive microheater and temperature sensor), and a flow gate connected to a serpentine channel (width: 500 μπι, height: 20 μπι) for sample transfer to the sensing microchamber. A weir structure (height: 20 μπι) in the enrichment microchamber retained microbeads (diameter: 50 to 80 μπι) within the chamber during the selective enrichment process. Three sets of a resistive heater and a temperature sensor separately controlled temperatures inside two microchambers and the serpentine channel for forming or disrupting aptamer-AVP complexes.

The source and drain electrodes were connected by graphene which served as a conducting channel. A voltage was applied to the gate electrode forming a gate voltage (V_gs) through a sample solution. In this FET configuration, the electrical conductance through graphene, which depends on the charge on the graphene surface, was measured from the drain current (Ids) at a fixed drain voltage (ϊ¾.

The microfluidic device was fabricated as follows. Graphene was synthesized via chemical vapor deposition (see Example 2), and transferred onto gold electrodes that were fabricated on a silicon dioxide-coated silicon substrate. The graphene surface was then functionalized with AVP by incubating in 1- pyrenebutanoic acid succinimidyl ester (PASE linker, 2 mM) solution for 1 hr and AVP solution (1 μΜ) for 4.5 hrs, sequentially. A PDMS sheet defining microchambers and channels was fabricated via soft lithography, followed by punching through-holes which were used as inlet/outlet ports and an open chamber. The sheet was then bonded onto a glass substrate on which a resistive microheater and temperature sensor was patterned via photolithography. The graphene sensor chip was finally stacked and bonded on the PDMS layer, while the AVP-functionalized graphene region was included within the open sensing microchamber.

The principle of AVP detection in the device is schematically shown in Fig. 7. AVP molecules in a sample ("sample AVP") were selectively captured by an aptamer immobilized on microbead surfaces at 37 °C (Fig. 7a), and were enriched by continuous sample infusion into the device. After removal of nonspecifically adsorbed AVP molecules and impurities through buffer washing (Fig. 7b), the enriched AVP molecules were released at an elevated temperature (55 °C) and eluted with solution-borne aptamer molecules ("free aptamer"), causing isocratic elution of AVP with the free aptamer (Fig. 7c). The mixture of the free aptamer and released sample AVP was incubated with graphene functionalized with standard AVP (Fig. 7d), inducing the binding of the free aptamer to the standard AVP on graphene via competitive binding (Fig. 7e), thus changing the graphene conductance.

The device was initially rinsed with buffer (10 μΐ/min), and the sensing microchamber was filled with a buffer solution to measure a reference signal ( _ef)- sample solution (10 μΐ/min for 2 hrs) prepared in buffer with varying AVP concentrations was continuously introduced into the enrichment microchamber using a syringe pump (KD210P, KD Scientific Inc.) while the chamber was maintained at 37 °C using a temperature control unit (an integrated microheater and a temperature sensor controlled by Lab VIEW) integrated on the bottom glass substrate. Following the enrichment, the chamber was thoroughly rinsed with a buffer solution (10 μΐ/ min) at 37 °C. After the waste outlet was closed, a plug of the free aptamer solution at 25 nM was flowed through (10 μΐ at 10 μΐ/ min) the enrichment microchamber at an elevated temperature, 55 °C. The mixture of eluted sample AVP and the free aptamer was then transferred to the sensing microchamber whose temperature was kept at 37 °C.

The measurement circuit used a DC power supply (E3631A, Agilent) to provide the drain voltage _ds, a function generator (33220A, Agilent) to supply the gate voltage V_gs, and a digit multimeter (3441 OA, Agilent) to measure the drain current (7_ds). During the electrical measurement, I_ds values (at fixed V_gs and _ds) were automatically collected once per second for a determined period through a PC-based Lab VIEW.

The aptamer binding-induced change in graphene conductance was characterized using the graphene nanosensor. A free aptamer solution at different concentrations in the range of 0-1000 nM was incubated with the graphene nanosensor until ids reached a saturated level (Fig. 16). The magnitude of the conductance difference, δ ids =| Us - ^ds,ref|, was found to increase with the free aptamer concentration and saturate above 100 nM, indicating that the increase of surface charge arises from the affinity binding of charged nucleic acid aptamer molecules to standard AVP on the graphene surface. Least square fitting of a monovalent binding model to the experimental data yielded a dissociation constant ¾) of 16.7 nM for aptamer- AVP binding, approximately 10 times higher than the value previously obtained in solution (1.7 nM ( Vroc. Natl. Acad. Sci, vol. 103, pp. 5173-5178)), which can be attributed to the restriction of configurational freedom of surface-immobilized AVP, which substantially reduces their binding kinetics.

The nanosensor was then tested with standard mixtures of the free aptamer at a fixed concentration (25 nM) and sample AVP at a varying concentration (cAVP, 1-100 nM). Results showed that 6Us decreased with increasing cAVP in the range of 1-100 nM, which is attributed to a higher AVP concentration in a standard mixture causes a higher aptamer occupancy, which impeded the aptamer binding to surface-immobilized standard AVP (Fig. 17). However, no appreciable variation in δ/ds with varying cAVP was observed below 1 nM (data not shown), necessitating the enrichment of sample AVP prior to graphene nanosensing to achieve the AVP detection in the low picomolar range, required in clinical settings.

The aptamer-based selective enrichment of AVP was characterized on microbead surfaces via fluorescence measurements. A TAMRA (as a fluorophore)- labeled AVP (TVP) solution at 100 pM was continuously infused into the enrichment chamber (10 μΐ/min) while the chamber temperature was maintained at 37 °C, and the fluorescence intensity of the bead surface was measured every 15 min. It was observed that mean fluorescence intensities increased over time and saturated after 1.5 hr (Fig. 18), suggesting that continuous sample infusion for more than 1.5 hr was needed to attain sufficient analyte enrichment. The thermally activated release of captured AVP molecules was also tested by flushing beads with a free aptamer solution at 55 °C. The fluorescence intensity was quickly (< 1 min) dropped down to the baseline, indicating that enriched sample AVP molecules were released from aptamer at 55 °C which disrupts aptamer-AVP complexes. This temperature- dependent behavior of aptamer-AVP binding enables isocratic elution of the analyte, allowing subsequent graphene nanosensing.

To verify the specificity of aptamer-based enrichment, integrated devices were tested with and without aptamer on microbead surfaces for AVP detection. Devices with bare beads produced higher δ/ds values than those with aptamer-functionalized beads (Fig. 19) at 100 pM AVP. This indicates that the sample AVP was significantly enriched by aptamer-based specific recognition on the bead surface, thus resulting in the higher sample AVP concentration in a mixture.

Detection of AVP was tested at physiologically relevant picomolar concentrations (1-500 pM) using the integrated device. To account for experimental variability caused by device-to-device variation of the graphene nanosensor, 5Ids was normalized with respect to δ/ds max, the maximum conductance change of an individual graphene nanosensor. Experimental results showed that δ/ds decreased with increasing cAVP (Fig. 20), implying that a higher cAVP results in a higher aptamer occupancy in the mixture, thus less free aptamer molecules can bind to standard AVP on graphene. The results also confirmed the effectiveness of aptamer- based selective enrichment by demonstrating the picomolar detection that was difficult to attain without the enrichment process. It can thus be concluded that this approach is capable of quantitatively detecting AVP at picomolar levels, confirming the potential of the device as a quantitative AVP assay in clinical diagnostics. It was speculate that the large error bars could arise from a variation of microbead packing density in the enrichment chamber between devices.

The foregoing merely illustrates the principles of the disclosed subject matter. Various modifications and alterations to the described embodiments will be apparent to those skilled in the art in view of the inventors' teachings herein. Features of existing methods can be integrated into the methods of the exemplary embodiments of the disclosed subject matter or a similar method. It will thus be appreciated that those skilled in the art will be able to devise numerous methods which, although not explicitly shown or described herein, embody the principles of the disclosed subject matter and are thus within its spirit and scope.

Claims

1. A microdevice for monitoring a target analyte using at least one aptamer that binds to a target analyte, the microdevice comprising: a substrate platform and a nanosensor, wherein the nanosensor is coupled to the substrate platform and comprises a conductance element functionalized with an aptamer for detecting the target analyte.

2. The microdevice of claim 1, wherein the conductive element comprises an element having a surface adapted for a change in charge density thereon upon the binding of the aptamer with the target analyte.

3. The microdevice of claim 1, wherein the conductive element comprises graphene.

4. The microdevice of claim 3, wherein the graphene comprises a single layer sheet.

5. The microdevice of claim 1, wherein the nanosensor further comprises a gate electrode formed from a nanolayer including at least one of ITO, Ti/Pd/Pt, gold, chrome or copper.

6. The microdevice of claim 3, wherein the nanosensor further comprises a source electrode and a drain electrode, and wherein the graphene sheet makes contact to both the source electrode and the drain electrode.

7. The microdevice of claim 6, wherein the source and drain electrodes comprise a nanolayer including at least one of ITO, Ti/Pd/Pt, chromium, gold, or chrome.

8. The microdevice of claim 1, wherein the platform substrate comprises at least one of polyethylene terephthalate (PET), polycarbonate polystyrene, polymethyl methacrylate (PMMA), polymacon, silicones, fluoropolymers, silicone acrylate, fluoro-silicone/acrylate, or poly hydroxyethyl methacrylate.

9. The microdevice of claim 1, wherein the target analyte comprises a small organic, peptide, protein, cell, or tissue.

10. The microdevice of claim 1, wherein the aptamer is a single-stranded oligonucleotide or a peptide.

11. The microdevice of claim 10, wherein the aptamer is a single-stranded oligonucleotide comprising DNA, RNA, XNA, or a combination thereof.

12. The microdevice of claim 3, wherein the aptamer is functionalized to the surface of graphene via a linker.

13. The microdevice of claim 12, wherein the linker is a pyrene-terminated agent.

14. The microdevice of claim 13, wherein the the pyrene-terminated agent is 1- pyrenebutanoic acid succinimidyl ester.

15. The microdevice of claim 12, wherein the graphene is serially functionalized with a polymer and aptamers.

16. The microdevice of claim 15, wherein one end of the polymer is coupled to the aptamer and the other end of the polymer is coupled to the linker.

17. The microdevice of claim 15, wherein the polymer is polyethlene glycol.

18. A microdevice for monitoring a target analyte comprising: a support platform, at least one enrichment chamber, wherein the enrichment chamber comprises microbeads functionalized with an aptamer that binds target analyte; at least one sensing chamber, wherein the sensing chamber comprises a graphene nanosensor functionalized with control target analyte; and at least one serpentine channel connecting the at least one enrichment chamber and the at least one sensing chamber.

19. The microdevice of claim 18, wherein the graphene comprises a single layer sheet.

20. The microdevice of claim 18, wherein the graphene nanosensor further comprises a gate electrode formed from a nanolayer including at least one of ITO, Ti/Pd/Pt, gold, chrome or copper.

21. The microdevice of claim 20, wherein the graphene nanosensor further comprises a source electrode and a drain electrode, and wherein the graphene sheet makes contact to both the source electrode and the drain electrode.

22. The microdevice of claim 21, wherein the source and drain electrodes comprise a nanolayer including at least one of ITO, Ti/Pd/Pt, chromium, gold, or chrome.

23. The microdevice of claim 18, wherein the support platform substrate comprises at least one of polyethylene terephthalate (PET), polycarbonate polystyrene, polymethyl methacrylate (PMMA), polymacon, silicones, fluoropolymers, silicone acrylate, fluoro-silicone/acrylate, or poly hydroxy ethyl methacrylate.

24. The microdevice of claim 18, wherein the target analyte comprises a small organic, peptide, protein, cell, or tissue.

25. The microdevice of claim 18, wherein the aptamer is a single-stranded oligonucleotide or a peptide.

26. The microdevice of claim 25, wherein the aptamer is a single-stranded oligonucleotide comprising DNA, RNA, XNA, or a combination thereof.

27. The microdevice of claim 18, wherein the aptamer is functionalized to the surface of microbead via a linker.

28. The microdevice of claim 27, wherein the linker is a pyrene-terminated agent.

29. The microdevice of claim 28, wherein the the pyrene-terminated agent is 1 - pyrenebutanoic acid succinimidyl ester.

30. The microdevice of claim 18, further comprising a temperature control unit for each enrichment chamber, each sensing chamber, and each serpentine channel.

31. A method for monitoring a target analyte using an aptamer capable of binding to the target analyte, comprising: placing a nanosensor in contact with a sample containing a target analyte, wherein the nanosensor comprises a conductance element functionalized with an aptamer for detecting the target analyte; detecting a change, if any, in the conductance of the conductance element; and electrodebased on the detected change, determining a presence and/or concentration of the target analyte in the sample.

32. The method of claim 31, wherein the binding of the aptamer with the target analyte causes a change in the charge density on the conductive element surface.

33. The method of claim 31, wherein the aptamer reversibly binds with the target analyte.

34. The method of claim 31, wherein the nanodevice detects low concentrations of the target analyte.

35. A method for monitoring a target analyte using an aptamer capable of binding to the target analyte, comprising: placing a device in contact with a sample containing a target analyte, wherein the device comprises at least one enrichment chamber; at least one sensing chamber comprising a conductance element; and at least one serpentine channel connecting the at least one enrichment chamber and the at least one sensing chamber; enriching the target analyte in the enriching chamber; releasing the enriched target analyte from the enriching chamber; moving the enriched target analyte to the sensing chamber; detecting a change, if any, in the conductance of the conductance element in the sensing chamber; and based on the detected change, determining a presence and/or concentration of the target analyte in the sample.

36. The method of claim 35, further comprising removing the nonspecifically adsorbed target analyte and impurities from the enriching chamber prior to releasing the enriched target analyte.

37. The method of claim 35, further comprising adding a solution of free aptamer to the enriching chamber following release of the target analyte.

38. The method of claim 35, wherein the binding of the aptamer with the control target analyte causes a change in the charge density on the conductive element surface.