HK1238717A1 - Noninvasive body fluid stress sensing - Google Patents

Abstract

Electrochemical impedance-based label-free and rapid biosensor for select bodily fluid biomolecule levels. Monoclonal antibodies to of biomolecule such as Cortisol were covalently attached to a 16-mercaptohexadecanoic acid functionalized gold working electrode using zero-length crosslinkers N-(3-dimethylaminopropyl)-N- ethylcarbodiimide and 10mM N-hydroxysulfosuccinimide. Cortisol was detected in phosphate buffered saline (simulated tear fluid) using a simple ferrocyanide reagent with a lower limit of detection of 18.73pM and less than 10% relative standard deviation.

Description

Non-invasive body fluid pressure sensing

Cross Reference to Related Applications

This application claims the benefit of U.S. provisional patent application No. 62/037,006 filed on 8/13/2014 and U.S. provisional patent application No. 62/101,143 filed on 8/2015, which are incorporated herein by reference in their entireties.

Technical Field

The present invention relates to the field of electrochemical sensing.

Background

Chronic stress (stress) has had an increasing implication over the last 40 years with a wide and growing variety of the most deadly and life-changing diseases of humans. Such diseases include serious conditions such as diabetes, alzheimer's disease, heart attack, depression, osteoporosis and immunosuppression, as well as non-fatal, but still unfortunate problems such as the common cold, back pain, and even erectile dysfunction. Indeed, the scientific literature shows that stress has an impact on life expectancy in developed countries that exceeds genetic and behavioral factors such as smoking.

In view of the enormous impact of stress on human life and health worldwide, there is great potential for large-scale detection and treatment of stress in the population. Although stress is often described as a subjective emotional state, it has medically important biochemical and physiological effects. These effects can be quantified, for example, by the increased levels of a particular group of hormones, including glucocorticoids and catecholamines. However, even at elevated levels, the physiological concentrations of these hormones in tears, saliva and serum are often very low (38.9 ± 15.5, 46.3 ± 16.0 and 489.7 ± 177.4nM, respectively), making accurate measurements a continuing technical challenge.

Disclosure of Invention

Improved electrochemical sensors using microfluidic tear capture systems have been prepared to detect biomolecules associated with stress and/or trauma, such as cortisol. In addition, other body fluids such as saliva or blood may be utilized.

In one embodiment, the monoclonal antibody is covalently attached to a 16-mercaptohexadecanoic acid functionalized gold working electrode using a zero length crosslinker, N- (3-dimethylaminopropyl) -N-ethylcarbodiimide and 10mM N-hydroxysulfosuccinimide. Cortisol was detected in phosphate buffered saline (simulated tears) using a simple ferrocyanide reagent at a lower detection limit of 18.73pM and a relative standard deviation of less than 10%. The cortisol assay shown herein maintains high reproducibility and extremely low detection levels in a label-free and fast-responding configuration, with sensitivity higher than that sufficient for tear cortisol detection.

These and other aspects of the embodiments disclosed herein will become more apparent by reference to the following detailed description and accompanying drawings.

Drawings

Fig. 1.a. basic scheme of an embodiment of an apparatus with a three-electrode system comprising (a) an Ag/AgCl reference electrode, (b) a sensing well, (c) a sample, (d) an Au working electrode, (e) GDE, and (f) a Pt counter electrode. Note that all materials are exemplary and may be replaced by other suitable materials. In addition, a Multiplexed Electrochemical Impedance Spectroscopy (MEIS) system is schematically depicted in operative connection with the apparatus. B. A sample with target (a) cortisol was placed in a sensing well on the surface of a covalently immobilized monoclonal antibody (MAb) on the surface of a gold working electrode, where MAb (b) was covalently immobilized on the gold surface (d) using 16-mhda (c) and EDC/NHS. The cortisol (a) target in the sample binds to the MAb.

Fig. 2 Nyquist plots for nine different MAb-immobilized electrodes run in cortisol target solutions: (a)0pg/ml, (b)1pg/ml, (c)5pg/ml, (d)10pg/ml, (e)50pg/ml, (f)100pg/ml, (g)500pg/ml, (h)1000pg/ml, (i)5000pg/ml, and (j)10000pg/ml in PBS buffer with 100mM potassium ferrocyanide redox probe.

Fig. 3A depicts the calculation of (a) slope and (b) R-square (fitted seal) to determine the optimal frequency of detection, derived and plotted against frequency from the concentration gradient.

Fig. 3b uses impedance at 1.184Hz and plots against cortisol concentration in PBS over and beyond the physiological range, showing the dynamic range of the sensor (n-3). A slope of 31.672 ohms/pg/ml was observed at the highest concentration variance, R2 was 0.9532 and RSD was 10%.

Fig. 4.a. basic scheme of an embodiment of an apparatus with a three-electrode system comprising (a) an Ag/AgCl reference electrode, (b) a sensing well, (c) a sample, (d) an Au working electrode, (e) GDE, and (f) a Pt counter electrode. Note that all materials are exemplary and may be replaced by other suitable materials. In addition, a Multiplexed Electrochemical Impedance Spectroscopy (MEIS) system is schematically depicted in operative connection with the apparatus. B. A sample with target (a) cortisol was placed in a sensing well on the surface of a covalently immobilized monoclonal antibody (MAb) on the surface of a gold working electrode, where MAb (b) was covalently immobilized on the gold surface (d) using 16-mhda (c) and EDC/NHS. The cortisol (a) target in the sample binds to the MAb.

FIGS. 5A-5I detection of biomarkers in simulated tears. Calibration curves for use in cortisol devices to relate measured impedance to cortisol concentration, and graphs showing the detection of many different biomolecules by the device of the invention.

FIGS. 6A-6F show schematic diagrams of biomarker detection data in blood.

Fig. 7 depicts a summary of biomarker data.

Figure 8. results of cortisol interferent tests. The signal-to-noise ratio obtained from ELISA assays using IgG anti-cortisol antibodies, typically cortisol gradients, and test interferents at 200pg/mL for each analyte are depicted relative to the standard provided.

Fig. 9 depicts a summary of pressure biomarker data using cyclic voltammetry.

FIG. 10 depicts pressure biomarker data using a amperometric technique.

Figure 11 depicts pressure biomarker data using SWV (square wave voltammetry) techniques.

Detailed Description

Although blood has historically been the standard diagnostic test fluid, in recent years tears have gained attention as a powerful sensing medium for three main reasons. First, the tear film includes a number of biomarkers. Second, it is relatively easy to obtain tears as compared to obtaining blood from a patient, which makes tears an ideal replacement for blood in diagnostic tests. Finally, tears, like saliva, are simpler in composition than blood and include fewer proteins that may interfere with electrochemical sensing.

While there are some disadvantages in using tears (e.g., the achievable volumes and target concentrations are much smaller than blood), these disadvantages are outweighed by the advantages of better sensor performance due to their easier and less invasive sampling and less background interference from non-target substances, demonstrating that the tear film is an ideal diagnostic liquid for pressure sensors while still including detectable levels of cortisol.

Thus, in one aspect disclosed herein, screen printed electrodes (one embodiment of which is shown in fig. 1A-B), tear fluid samples are captured by a novel microfluidic capture system that brings the sample to a reagent, and one or more molecular recognition units for cortisol (or other pressure markers found in tears) encapsulated in the mesoporous carbon ink of the sensor itself have been developed using rapid, label-free and Multiplexed Electrochemical Impedance Spectroscopy (MEIS), which can be used at the point of care/injury. The molecular recognition unit may comprise one or more of antibodies, aptamers, peptides, synthetic antibodies (synbodies), nucleic acids, tentacle probes, proteins, and the like. Also, mesoporous carbon inks have been found to block interferents, leading to better experimental results.

Although stress is often described as a subjective emotional state, it has been shown to have important biochemical and physiological effects, with dramatic effects on human health. Therefore, monitoring stress levels by sensing biochemical markers has the potential to have a dramatic impact on stress management. Electrochemical Impedance Spectroscopy (EIS) is one such sensing method that has been successfully used in label-free detection of a variety of very low concentration targets, including whole cells, protein biomarkers, and small molecule targets. Advantages of EIS over other electrochemical methods include speed (90 seconds per test), simplicity (no labeling is required as with "sandwich" assays), and sensitivity (detection of picomolar-concentration targets below the detection limit of many other methods). This label-free sensing ability and extremely low detection limit make EIS an ideal sensing mechanism for cortisol in tears.

Example 1

A standard three-electrode system was used for impedance spectroscopy detection. The system comprised an Ag/AgCl reference electrode (CH instruments, austin, texas), gold disk working electrode (GDE) (CH instruments, austin, texas), and a platinum counter electrode (CH instruments, austin, texas), an anti-cortisol antibody (Sigma-Aldrich, st louis, missouri) covalently attached to the working electrode surface to detect cortisol in the sample solution. A 1000 μ L pipetting head (vwrinternationary, rad, pa) with a razor-tipped tip was mounted close to the GDE to form a plastic "well" capable of holding approximately 0.2ml of sample fluid. A schematic of this system is shown in fig. 1.

Phosphate Buffered Saline (PBS) at pH 7.4 (EMD Biosciences, rahoya, ca) was used to prepare all solutions unless otherwise indicated. To immobilize the anti-cortisol antibody onto the surface of Gold Disk Electrode (GDE), GDE was first subjected to 120-pass wet-milling on 3 μm alumina grit (CH instruments, austin, tx) and washed with distilled water. Then 120 repetitions of figure 8 sanding were performed with 1 μm and then 0.05 μm grit (CH instruments, austin, texas) after which the GDE was sonicated in distilled water for 20 minutes. Next, 100. mu.L of a reagent grade ethanol solution of 1mM 16-mercaptohexadecanoic acid (16-MHDA) (Sigma-Aldrich, St. Louis, Mo.) was placed into the sensing well and sealed with parafilm for 1 hour at room temperature. Next, the surface and sides of the GDE and the sensing well were carefully washed with distilled water. Control EIS assays were performed on 16-MHDA-functionalized GDEs using a "redox probe" of 100mM potassium ferrocyanide (Sigma-Aldrich, St. Louis, Mo.) in PBS buffer to ensure that sufficient and similar amounts of MHDA were immobilized to each GDE. This is determined by analyzing the impedance response of each individual GDE to obtain mutual similarity.

Next, 100 μ L of PBS solution containing 40mM N- (3-dimethylaminopropyl) -N-Ethylcarbodiimide (EDC) (Pierce Biotechnology, Inc.) and 10mM N-hydroxysulfosuccinimide (sulfo-NHS) (VWR International, Inc.) was placed into the sensing wells. After incubation at room temperature for 1 hour, the electrodes were washed with PBS buffer. Next, 100. mu.L droplets of a 10. mu.g/ml solution of anti-cortisol IgG (Aldrich) in PBS buffer were placed on the electrodes and left at room temperature for 1 hour, followed by rinsing with PBS buffer. Finally, 100 μ L of a distilled aqueous solution of 1mM ethanolamine (Sigma-Aldrich, St. Louis, Mo.) was added to the sensing wells and incubated at room temperature for 30 minutes to block all unreacted carboxyl groups of 16-MHDA and EDC/NHS. The electrodes were then carefully washed with PBS buffer and stored in PBS at 4 ℃ until use.

Electrochemical impedance measurements were performed using a CHI660C electrochemical workstation (CH instruments, houston, tx). Samples of cortisol (Sigma-Aldrich, St. Louis, Mo.) were prepared at concentrations from 0 to 10,000pg/mL (0 to 27.59nM) in the redox probe solution and stored at 4 ℃ until use. Each cortisol concentration is then detected at each antibody-immobilized electrode.

For each assay, 100 μ Ι _ of cortisol and redox probe solution was placed within the sensing well of the antibody-immobilized GDE. The AC voltage applied to the sample had an amplitude of 5mV and the formula potential (DC offset) was 150mV, as determined by running the CV on a bare (pre-fixation) electrode with a redox probe. The AC voltage was applied over a 90 second sweep in a frequency range from 1 to 100,000Hz and the impedance magnitude and phase were recorded at each frequency for this sample. Real (real) and imaginary (imaginary) impedances are calculated and plotted in the nyquist plot for each sample. After each detection, GDE and sensing wells were washed thoroughly with PBS before the next sample was added.

For each electrode tested at each AC frequency, the magnitude of the impedance at each cortisol concentration was correlated to the log (concentration) and the slope and R were calculated². The impedance slope and R²Values are plotted against frequency separately, looking for results in high slope and R²The frequency of the optimum balance. Impedance detected at the "optimum" frequencyThe values are then used to generate the final concentration gradient, thereby enabling cortisol concentration to be derived from the impedance estimate.

AC scanning of the bare, antibody-immobilized, biomarker (cortisol) -bound electrode produced nyquist plots. Nyquist plots of 9 different concentrations of cortisol bound to one representative electrode are shown in figure 2. As the sample cortisol concentration increases, more cortisol binds to the antibodies on the electrode surface, thereby increasing the impedance magnitude at all frequencies and pulling the nyquist plot further away from the origin.

As expected, as the concentration of the biomarker ex vivo and thus bound to the antibody increases, the impedance (and thus the resulting signal) detected by the system also increases.

Slope and R of the correlation plot of impedance versus target concentration²As a function of the frequency of the AC voltage. One representative electrode is shown in fig. 3 a. A larger slope is desirable because it corresponds to a larger signal magnitude (larger difference in impedance values between different cortisol concentrations), which is easier to detect with low relative error. Greater R²Is desirable because it represents a higher accuracy of the estimate of cortisol concentration provided in the assay. In FIG. 3a, R can be seen²It is quite high for frequency ranges below 100Hz, but the slope is especially best (maximum) at very low frequencies, and the slope decreases rapidly with increasing frequency. The optimum frequency for maximum slope was found to be 1.18Hz for all electrodes tested. Therefore, this is the best frequency to evaluate at which cortisol-antibody interactions are most effectively detected by EIS.

At this frequency (1.18Hz), the impedance data for the multiple sensors was compiled and plotted against the response concentration to generate an impedance gradient, as shown in fig. 3b. This gradient shows the non-questionable accuracy of this method in detecting very low concentrations of cortisol in tears. From the standard assay definition of the Lower Limit of Detection (LLD), i.e., 3.3 × slope divided by standard deviation, an LLD of 6.79pg/mL (18.73pM) was determined at a detection time of 90 seconds per sample. This clearly identifies the cortisol levels of the tear being measured with high accuracy and the difference between sensors at each concentration is < 10%.

The LLD of 18.73pM is three integer orders of magnitude lower than the typical cortisol concentration range of about 40nM in tears, which at first glance appears to be a very high level of sensitivity for tear sensing applications. However, in practice, such ultra-sensitive detection is needed to allow the cortisol assay to be converted from a laboratory to a physical real-world sensor device. Reproducible and reliable sensors require small differences not only in electrochemical measurements but also in physical device implementation. The tear sample size is unlikely to exceed a 10 microliter volume because no much tear fluid can be collected and some volume will inevitably be lost by adhering to the walls of the sampling system or other fluidic device required to carry the sample from the ocular surface to the sensing electrodes.

As a result, ensuring a consistent reproducible volume of liquid in the electrochemical sensing region (functionalized electrode) is extremely difficult without increasing the volume-diluting the target concentration by a known factor (which can then be considered to calculate the original sample concentration). In addition, tears do not include high concentrations of redox mediators, such as ferrocyanide, which are required for electrochemical work. Therefore, the following is difficult for an actual sensor device: dilution of cortisol around 40nM with additional reagents is avoided in order to increase the total liquid volume to a workable amount that ensures a consistent volume of functionalized electrode area can be reached each time and to provide sufficient mediator concentration for the sensor electrochemistry. For this reason, commercially viable tear cortisol sensors must provide reproducible measurements not in the 1-100nM range, but in the range below 10nM (e.g. when samples with abnormally low cortisol levels are diluted even further by the device during processing).

This is what is allowed by the EIS-based assay presented herein. With an LLD below 0.02nM, a 10 μ L tear sample with 40nM cortisol can be diluted 100X and still be well within the linear range of the EIS-based cortisol sensor. The cortisol assay presented here is therefore more able to meet the technical challenges of discriminating low cortisol concentrations in tears.

In this work, the detection of very low concentrations of cortisol was demonstrated to be reproducible and highly sensitive by using simple and label-free EIS-based biosensors. The replicated sensor components had the best binding at 1.184Hz, were reproducible, and had the highest variability at 10% relative standard deviation. The degree of match was detected as 0.9532, with a response rate of 31.672 ohms/pg/mL and a lower limit of detection of 18.73 pM. This work shows that small changes in cortisol levels can be measured accurately and quickly, even if as little as is technically feasible as those typically found in human tears; it is feasible even after considering the practicality of physical sensor design that may require further dilution of already low concentrations of target for reproducible performance.

In another aspect disclosed herein, screen printed electrodes (one embodiment of which is shown in fig. 4), bodily fluid samples are captured by a novel microfluidic capture system that brings the sample to a reagent, and one or more molecular recognition units for cortisol (or other pressure markers found in the liquid) encapsulated in the sensor's own mesoporous carbon ink have been developed using rapid, label-free and Multiplexed Electrochemical Impedance Spectroscopy (MEIS), which can be used at the point of care/injury. Although tears are used in this embodiment, blood may also be used, as shown in fig. 6.

The molecular recognition unit may comprise one or more of an antibody, an aptamer, a peptide, a synthetic antibody, a nucleic acid, a tentacle probe, a protein, and the like. Also, mesoporous carbon inks have been found to block interferents, leading to better experimental results.

Example 2

Although the following examples are for the detection of cortisol, similar protocols are used for the detection of other target biomolecules. Tears or blood are used in this embodiment, but other body fluids may also be used.

A standard three-electrode system was used for impedance spectroscopy detection. The system comprised an Ag/AgCl reference electrode (CH instruments, austin, texas), gold disk working electrode (GDE) (CH instruments, austin, texas), and a platinum counter electrode (CH instruments, austin, texas), an anti-cortisol antibody (Sigma-Aldrich, st louis, missouri) covalently attached to the working electrode surface to detect cortisol in the sample solution. A 1000 μ L pipetting head (vwrinternationary, rad, pa) with a razor-tipped tip was mounted close to the GDE to form a plastic "well" capable of holding approximately 0.2ml of sample fluid. A schematic of this system is shown in fig. 4.

Phosphate Buffered Saline (PBS) at pH 7.4 (EMD Biosciences, rahoya, ca) was used to prepare all solutions unless otherwise indicated. To immobilize the anti-cortisol antibody onto the surface of Gold Disk Electrode (GDE), GDE was first subjected to 120-pass wet-milling on 3 μm alumina grit (CH instruments, austin, tx) and washed with distilled water. Then 120 repetitions of figure 8 sanding were performed with 1 μm and then 0.05 μm grit (CH instruments, austin, texas) after which the GDE was sonicated in distilled water for 20 minutes. Next, 100. mu.L of a reagent grade ethanol solution of 1mM 16-mercaptohexadecanoic acid (16-MHDA) (Sigma-Aldrich, St. Louis, Mo.) was placed into the sensing well and sealed with parafilm for 1 hour at room temperature. Next, the surface and sides of the GDE and the sensing well were carefully washed with distilled water. Control EIS assays were performed on 16-MHDA-functionalized GDEs using a "redox probe" of 100mM potassium ferrocyanide (Sigma-Aldrich, St. Louis, Mo.) in PBS buffer to ensure that a sufficient and similar amount of MHDA was immobilized to each GDE. This is determined by analyzing the impedance response of each individual GDE to obtain mutual similarity.

Next, 100 μ L of PBS solution containing 40mM N- (3-dimethylaminopropyl) -N-Ethylcarbodiimide (EDC) (Pierce Biotechnology, Inc.) and 10mM N-hydroxysulfosuccinimide (sulfo-NHS) (VWR International, Inc.) was placed into the sensing wells. After incubation at room temperature for 1 hour, the electrodes were washed with PBS buffer. Next, 100. mu.L droplets of a 10. mu.g/ml solution of anti-cortisol IgG (Aldrich) in PBS buffer were placed on the electrodes and left at room temperature for 1 hour, followed by rinsing with PBS buffer. Finally, 100 μ L of a distilled aqueous solution of 1mM ethanolamine (Sigma-Aldrich, St. Louis, Mo.) was added to the sensing wells and incubated at room temperature for 30 minutes to block all unreacted carboxyl groups of 16-MHDA and EDC/NHS. The electrodes were then carefully washed with PBS buffer and stored in PBS at 4 ℃ until use.

Electrochemical impedance measurements were performed using a CHI660C electrochemical workstation (CH instruments, houston, tx). Samples of cortisol (Sigma-Aldrich, St. Louis, Mo.) were prepared at concentrations from 0 to 10,000pg/mL (0 to 27.59nM) in the redox probe solution and stored at 4 ℃ until use. Each cortisol concentration is then detected at each antibody-immobilized electrode.

For each assay, 100 μ Ι _ of cortisol and redox probe solution was placed within the sensing well of the antibody-immobilized GDE. The AC voltage applied to the sample had an amplitude of 5mV and the formula potential (DC offset) was 150mV, as determined by running the CV on a bare (pre-fixation) electrode with a redox probe. The AC voltage was applied over a 90 second sweep in a frequency range from 1 to 100,000Hz and the impedance magnitude and phase were recorded at each frequency for this sample. Real and imaginary impedances are calculated and plotted in the nyquist plot for each sample. After each detection, GDE and sensing wells were washed thoroughly with PBS before the next sample was added.

The impedance at each cortisol concentration will be large for each electrode tested at each AC frequencySmall is associated with log (concentration) and the slope and R are calculated². The impedance slope and R²Values are plotted against frequency separately, looking for results in high slope and R²The frequency of the optimum balance. The impedance values detected at this "optimal" frequency are then used to generate the final concentration gradient, thereby enabling the cortisol concentration to be estimated from the impedance.

As expected, as the concentration of the biomarker ex vivo and thus bound to the antibody increases, the impedance (and thus the signal) detected by the system also increases. See fig. 5A-5I through fig. 8.

In this work, the detection of very low concentrations of cortisol was demonstrated to be reproducible and highly sensitive by using simple and label-free EIS-based biosensors. The replicated sensor components had the best binding at 1.184Hz, were reproducible, and had the highest variability at 10% relative standard deviation. The degree of match was detected as 0.9532, with a response rate of 31.672 ohms/pg/mL and a lower limit of detection of 18.73 pM. This work shows that even for reproducible performance, it is technically feasible to accurately and quickly measure small changes in cortisol levels, even those typically found in human tears, after considering the practicality of physical sensor design that may require further dilution of already low concentrations of target.

Also, as summarized in fig. 7, a number of target biomolecules can be detected, such as cortisol, glucose, lactic acid, lactoferrin, IgE, catecholamines, S-100 β, neuron-specific enolase, glial fibrin, and tumor necrosis factor- α.

Turning to fig. 9-11, a summary of pressure biomarker data is depicted. Cyclic Voltammetry (CV) is an electrochemical technique that detects the current generated in an electrochemical cell under conditions where the voltage exceeds the voltage predicted by the nernst equation. CV is performed by cycling the voltage of the working electrode and detecting the resulting current. FIG. 9 shows the CV superposition of A EP, B NE, C DA and D Cort (see structure in the figure). DA. The concentrations of EP, Cort and NE were 0.04M, 0.04M, 0.04M and 0.1M, respectively. The minimum superposition of the signals is denoted by E, where F denotes the large superposition of the signal peaks.

Fig. 10 depicts pressure biomarker data from a measure current technique. Amperometry in chemistry and biochemistry is based on the detection of ions in solution based on current or changes in current. Amp-it of (embedded) DA, where the voltage applied at the oxidation peak of CV is 0.52V, at A2 seconds, B12 seconds, C20 seconds during AMP-it. The outer graph is a calibration curve in which the current is plotted against the DA concentration at times (a), (b) and (c) during AMP-it. The logarithmic fit of the calibration curve at different times A, B and C had R of 0.9566, 0.9547 and 0.9540, respectively²。

FIG. 11 depicts the SWV (square wave voltammetry) technique used to determine the current at EP concentration versus oxidation peak (0.23V) at 30 Hz. The SWV technique was used to determine the current at 20Hz of DA concentration versus the oxidation peak (0.22V). The SWV technique was used to determine the current at 20Hz at the NE concentration versus the oxidation peak (0.23V). The SWV technique was used to determine the current at the concentration of Cort versus oxidation peak (0.18V) at 15 Hz.

The above-described embodiments are not intended to be limiting.

Claims (10)

1. An electrochemical sensor, the sensor comprising:

a reference electrode and a counter electrode;

a sensing well disposed between the reference electrode and the counter electrode; and

a functionalized working electrode disposed within the sensing well;

wherein the content of the first and second substances,

one or more molecular recognition units for a pressure marker are coupled to the functionalized working electrode.

2. The sensor of claim 1, wherein the one or more molecular recognition units comprise a monoclonal antibody covalently attached to a 16-mercaptohexadecanoic acid functionalized working electrode.

3. The sensor of claim 2, wherein the one or more monoclonal antibodies are attached to the functionalized working electrode using a zero length crosslinker N- (3-dimethylaminopropyl) -N-ethylcarbodiimide and 10mM N-hydroxysulfosuccinimide.

4. The sensor of claim 3, wherein the reference electrode comprises an Ag/AgCl electrode.

5. The sensor of claim 4, wherein the counter electrode comprises a Pt electrode.

6. The sensor of claim 1, further comprising a multiplexed electrochemical impedance spectroscopy system in operable arrangement therewith.

7. A method of sensing biomolecules in a bodily fluid, the method comprising:

after contacting a sensor having a molecular recognition unit for the biomolecule coupled to a working electrode with a sample of bodily fluid, detecting the amount of the biomolecule bound to the molecular recognition unit of the sensor using a multiplexed electrochemical impedance spectroscopy.

8. The method of claim 7, wherein the one or more molecular recognition units comprise a monoclonal antibody covalently attached to a 16-mercaptohexadecanoic acid functionalized working electrode.

9. The method of claim 7, wherein the bodily fluid comprises tears or blood.

10. The method of claim 7, wherein the biomolecule is one or more selected from the group consisting of cortisol, glucose, lactic acid, lactoferrin, IgE, catecholamines, S-100 β, neuron-specific enolase, glial fibrillary protein, and tumor necrosis factor- α.