CROSS-REFERENCE TO RELATED APPLICATIONS

[0001] The present application claims priority to U.S. Provisional Application No. 63/089,905, filed October 9, 2020, U.S. Provisional Application No. 63/134,690, filed January 7, 2021, and U.S. Provisional Application No. 63/155,963, filed March 3, 2021, the entire contents of each of which are incorporated herein by reference.

TECHNICAL FIELD

[0001] The present disclosure pertains to devices and methods for detecting infection by a pathogen in a mammalian subject

BACKGROUND

[0002] SARS-CoV-2, the virus that causes Covid- 19, has killed over 1.45 million people worldwide. Despite the urgency of the current pandemic, available diagnostic methods for Covid- 19 use RT-PCR tests that detect nucleic acid sequence specific to SARS-CoV-2. These tests are limited by their requirement of large laboratory space, high reagent costs, multistep sample preparations including an RNA extraction step, the potential for crosscontamination, and the need to train highly skilled workers. Moreover, results usually take hours to days to become available, and the need for transport of specimens to a central laboratory leads to further delays. Pneumonia caused by secondary bacterial infections has been found to contribute to the lethality of individuals with COVID-19.

[0003] Therefore, there is an urgent need to develop approaches to detect and diagnose both viral and bacterial infections.

SUMMARY

[0004] Provided herein are devices for assessing the presence of a pathogen, such as, SARS-CoV-2, in a biological sample. The devices can comprise a substrate comprising a top surface and a back surface; and, an electrode on the top surface of the substrate, wherein the electrode is functionalized with a detection moiety, such as one that binds SARS-CoV-2 spike protein.

[0005] Also provided are wearable articles comprising a device as described herein. [0006] The present disclosure also pertains to methods for assessing the presence of SARS-CoV-2 in a biological sample comprising contacting a device according to the present disclosure with the biological sample; exposing the device to an electrical current in order to generate a signal from the device; and, assessing the signal that is generated by the device electrochemical impedance spectroscopy (EIS) in order to determine the absence or presence of SARS-CoV-2 in the biological sample.

BRIEF DESCRIPTION OF THE DRAWINGS

[0007] The file of this patent or application contains at least one drawing/photograph executed in color. Copies of this patent or patent application publication with color drawings/photographs will be provided by the Office upon request and payment of the necessary fee.

[0008] FIG. 1 illustrates the detection capabilities of a device according to the present disclosure that is configured as a bandage.

[0009] FIG. 2 depicts the in vitro detection of infectious agents by a device according to the present disclosure.

[0010] FIG. 3 illustrates molecular dynamic simulations of the region of the SARS CoV-2 viral spike protein that binds to the human ACE2 protein.

[0011] FIG. 4 illustrates a process by which the inventive devices may be used.

[0012] FIG. 5 depicts the concept under which the inventive devices are used for rapid SARS-CoV-2 detection.

[0013] FIG. 6 provides the results of a real time assessment for diagnosing COVID- 19, in which detection = pg-ng of virus.

[0014] FIG. 7 depicts elements of point-of-care detection of SARS-CoV-2 using the DETECT 1.0 system.

[0015] FIG. 8 illustrates the characterization and calibration of an inventive system.

[0016] FIG. 9 depicts the use of miniaturized and portable device according to the present invention for rapid point-of-care diagnosis of a pathogen, such as COVID-19.

[0017] FIG. 10 provides Nyquist plots showing the response of the modified eChip to different concentrations of angiotensin II, the natural substrate of ACE2, ranging from 1 pg mL' ¹ to 10 pg mL' ¹

[0018] FIG. 11 illustrates the results of an investigation concerning Nafion concentration optimization for a permselective membrane on the present devices. [0019] FIG. 12 shows the results of a study of the effect of sample pretreatment steps on the detection of free SARS-CoV-2 SP

[0020] FIG. 13 depicts the results of a kinetic study of the interaction between SARS-CoV-2 SP and DETECT 1.0.

[0021] FIGS. 14A and 14B provide calibration curves for free SP in PBS solution (FIG. 14 A) and in VTM medium (FIG. 14B).

[0022] FIG. 15 depicts an equivalent circuit used for the extraction of the RCT values used in all EIS measurements. Rs = electrolyte resistance, RCT = charge transfer resistance, CPE = constant phase element, and W= Warburg component (diffusion-limited mass transport).

[0023] FIG. 16 illustrates the relative RCT response extracted from the Nyquist plots for 21 successive EIS measurements of PBS medium using the same biosensor (eChip). The relative standard deviation (RSD) of the RCT values obtained for 21 consecutive measurements was 5.3%, demonstrating an adequate stability for a long operation time (1.5 hours).

[0024] FIG. 17 illustrates the results of recording open circuit potential for 60 minutes from an inventive biosensor. During the initial 30 min, the sensor was exposed to a PBS solution, after which it was subjected to 1 ng mL' ¹ SP for the remaining 30 minutes of the experiment. The biosensor exhibited high stability with an RSD of 0.76% in the potential over the 30 minutes of exposure to SP.

[0025] FIG. 18 illustrates the results of a reproducibility test in which normalized sensitivity for 10 different biosensors (10 electrodes from different fabrication batches) was assessed. An analytical curve using free SP in the concentration range of 1 pg mL' ¹ to 1 ng mL' ¹ was constructed for each eChip. The relative standard deviation (RSD) value obtained was 6.8%, which represents an adequate reproducibility of the method considering that the functionalization step was not automated.

[0026] FIG. 19 depicts the results of an assessment of the stability (shelf-life) of DETECT in different conditions of storage (25 °C-black square, 8 °C - red circles, and -20 °C - blue triangles) over 10 days.

[0027] FIG. 20 provides the results of a test involving measurement of samples of SARS-CoV-2 subjected to heat inactivation.

[0028] FIG. 21 illustrates how the inventive system was used for detection of SARS-CoV-2 in a prospective cohort study. [0029] FIGS. 22 A and 22B provide information concerning a clinical study that was performed in the context of the COVID-19 pandemic in Philadelphia, PA.

DETAILED DESCRIPTION OF ILLUSTRATIVE EMBODIMENTS

[0030] The presently disclosed inventive subject matter may be understood more readily by reference to the following detailed description taken in connection with the accompanying examples, which form a part of this disclosure. It is to be understood that these inventions are not limited to the specific formulations, methods, articles, or parameters described and/or shown herein, and that the terminology used herein is for the purpose of describing particular embodiments by way of example only and is not intended to be limiting of the claimed inventions.

[0031] The entire disclosures of each patent, patent application, and publication cited or described in this document are hereby incorporated herein by reference.

[0032] As employed above and throughout the disclosure, the following terms and abbreviations, unless otherwise indicated, shall be understood to have the following meanings.

[0033] In the present disclosure the singular forms “a,” “an,” and “the” include the plural reference, and reference to a particular numerical value includes at least that particular value, unless the context clearly indicates otherwise. Thus, for example, a reference to “a detection moiety” is a reference to one or more of such moieties and equivalents thereof known to those skilled in the art, and so forth. Furthermore, when indicating that a certain element “may be” X, Y, or Z, it is not intended by such usage to exclude in all instances other choices for the element.

[0034] When values are expressed as approximations, by use of the antecedent “about,” it will be understood that the particular value forms another embodiment. As used herein, “about X” (where X is a numerical value) preferably refers to ±10% of the recited value, inclusive. For example, the phrase “about 8” can refer to a value of 7.2 to 8.8, inclusive. This value may include “exactly 8”. Where present, all ranges are inclusive and combinable. For example, when a range of “1 to 5” is recited, the recited range should be construed as optionally including ranges “1 to 4”, “1 to 3”, “1-2”, “1-2 & 4-5”, “1-3 & 5”, and the like. In addition, when a list of alternatives is positively provided, such a listing can also include embodiments where any of the alternatives may be excluded. For example, when a range of “1 to 5” is described, such a description can support situations whereby any of 1, 2, 3, 4, or 5 are excluded; thus, a recitation of “1 to 5” may support “1 and 3-5, but not 2”, or simply “wherein 2 is not included.”

[0035] In the present disclosure, relevant publications are cited in abbreviated format, except in the section, infra, following the heading “References”, in which the full citations of such references are provided.

[0036] As noted above, there is an urgent need to develop approaches to detect and diagnose both viral and bacterial infections. The present inventors have developed devices that may be cheaply produced and sold, and are capable of diagnosing microbial infections in 10 seconds, representing a vastly cheaper and faster alternative to current state-of-the-art methods used in hospitals (>$100 and diagnosis time of 24 hours) (FIGS. 1 and 2). For example, the devices may be purposed to rapidly detect the virus SARS-CoV-2. The instant technology provides the transformative ability of detecting dangerous infections through its simple design, speed, disposability and ease of operation. The presently disclosed portable electrochemical paper-based devices can use minimal sample volumes (10 pL), costs less than $1 to produce and can detect pathogens such as SARS-CoV-2 within 10 minutes, and are vastly cheaper and faster than current state-of-the-art diagnostics. The portable and easily operable test devie disclosed herein will enable widespread deployment, large-scale testing, and population-level surveillance. Thus, the presently disclosed invention has the potential to transform the way we diagnose pathogenic infections, including those that are currently untreatable, thus improving treatment outcome, potentially extending patient survival, and minimizing healthcare costs.

[0037] Accordingly, provided herein are devices for assessing the presence of a pathogen, such as SARS-CoV-2, in a biological sample, the devices comprising a substrate comprising a top surface and a back surface; and, an electrode on the top surface of the substrate, wherein the electrode is functionalized with a detection moiety, such as one that binds SARS-CoV-2 spike protein.

[0038] The substrate may comprise any material that does not interfere with the ability of the electrode to function as intended. For example, the substrate may comprise paper, cardboard, plastic (e.g., polymer), or textile. When the substrate is intended for use as a wearable, it may be of the same material as a traditional bandage, such as plastic or flexible fabric.

[0039] The electrode may be adhered to the substrate according to any suitable approach, and those of ordinary skill in the art can readily identify numerous approaches for applying an electrode material (e.g., a conductive paste) to a substrate in order to form an electrode. In some embodiments, the electrode is screen-printed onto the top surface of the substrate. In some embodiments, the electrode is wax-printed onto the top surface of the substrate.

[0040] The surfaces of the electrode on the substrate may be modified in order to enable binding to the detection moiety. For example, the electrode may be surface- functionalized with thiol groups. Functionalization with thiol groups can be used to form a disulfide bond with a detection moiety. In some embodiments, a disulfide bond occurs between the surface-functionalized electrode and an N-terminal cysteine residue that is engineered onto a detection moiety. For example, the detection moiety that binds SARS- CoV-2 spike protein is human Angiotensin Converting Enzyme 2 (ACE2), the amino acid sequence IEEQAKTFLDKFNHEAEDLFYQS (SEQ ID NO: 1), or an antibody. Any of these detection moieties may be engineered to include an N-terminal cysteine residue that can form a disulfide bond with thiol groups on the electrode in order to securely attach the detection moiety to the electrode. In some embodiments, a detection moiety may be immobilized on the surface of the electrode by crosslinking the detection moiety, such as by using a chemical cross-linker. For example, the detection moiety may be immobilized on the surface of the electrode by crosslinking the detection moiety using the bifunctional chemical cross-linker glutaraldehyde (GA). In certain embodiments, ACE2 or SEQ ID NO: 1 is immobilized on the electrode via an amide bond between the glutaraldehyde and the N-terminus of ACE2 or SEQ ID NO: 1. Full-length ACE2 and the 23-mer peptide of SEQ ID NO: 1 can be recombinantly generated in E. coli using previously established methods (Chan et al., 2020). The peptide of SEQ ID NO: 1 can alternatively be synthesized chemically. In some embodiments, the detection moiety is ACE2, and the ACE2 is applied onto the electrode such that the resulting amount of ACE2 on the electrode is 2.68 pg.

[0041] The present inventors have developed an electrochemical analytical device for detecting infections in real time. FIG. 1 A depicts a down-side photograph of a device coupled to an adhesive wearable for detecting Pyo through an electrochemical redox process, as provided in FIG. IB. FIG. 1C shows the effect of pH on the electrochemical behavior of Pyo. FIG. ID provides an EP vs. pH plot, and FIG. IE provides square wave voltammograms for successive additions of Pyo with concentrations ranging from 50 to 1000 nmol/L. The inset provides an analytical curve constructed with the peak current for both electrochemical processes. [0042] The potentially wearable device detects, through cyclic voltammetry, redoxactive metabolites uniquely produced by pathogenic infectious agents. In FIG. 2A, redox bacterial biomarkers (left) and ACE2 protein (right) were detectable by the device. On the right, the SARS-CoV-2-ACE2 structure is depicted. ACE2 and peptides derived from its structure are detectable by the present devices. FIG. 3 depicts the results of molecular dynamics simulations performed by the inventors of the region of the SARS-CoV-2 viral spike protein (blue) that binds to the human ACE2 protein (red and yellow). FIG. 2B provides pseudomonas aeruginosa CFU/mL counts of overnight culture dilution compared to current measured at pH2. FIG. 2C shows bacterial growth over time in LB medium determined by the device in relation to CFU/mL counts.

[0043] Impedimetric measurements by electrochemical impedance spectroscopy (EIS) provide qualitative and quantitative data for diagnosing COVID-19 directly from biological samples, such as human blood serum or saliva, through the precise detection of changes in charge transfer resistance due to the detection moiety -virus interaction.

[0044] Thus, electrochemical impedance spectroscopy measurements can be used to detect the selective binding of SARS-CoV-2 with the detection moiety, such as ACE2, which interacts specifically with the spike protein of SARS-COV-2, or a SEQ ID NO: 1, which represents a 23-mer peptide that interacts directly with SARS-CoV-2 (FIG. 4). As provided in FIG. 4, electrochemical impedance spectroscopy readings indicate differences in resistance after application of a steady potential and a range of frequency. The specificity of the interactions between ACE2 or peptides and the viral spike protein allow detection of the SARS-CoV-2 in a sample. In some embodiments, portable screen-printed carbon electrodes are chemically functionalized by anchoring the detection moiety to the electrode surface. As described above, functionalization can be achieved through chemical deposition and formation of disulfide bonds between an N-terminal cysteine residue that will be engineered into both ACE2 and the 23-mer peptide, and the thiol-functionalized electrode surface. The present inventors have previously engineered numerous peptides with an added cysteine for functionalization purposes.

[0045] Blocking agents, such as ethanolamine and bovine serum albumin, may be used to cover the remaining exposed surface of the electrode to avoid unspecific interactions and biofouling of the transductor surface, providing sensitive and selective SARS-COV-2 recognition. Thus, the present devices may comprise a blocking layer over the electrode. [0046] The surface of the electrode can also or alternatively be functionalized by forming a membrane that is protective, permselective, or both in order to enhance the robustness of the analytical device. The phrase “on the electrode” with reference to the membrane can refer to a condition in which the membrane is in direct contact with the electrode, or to a condition in which there are intervening structures between the membrane and the electrode. For example, there may be a blocking layer between the membrane and the electrode, and in such a situation, the membrane may still be referred to as being “on the electrode”, albeit in an indirect fashion. The membrane may be formed from a polymeric material. In some embodiments, the protective membrane can be formed by applying a solution that contains Nafion to the surface of the electrode. The Nafion solution can contain, for example, about 0.1% to about 5.0% v/v Nafion. In some embodiments, the Nafion solution contains about 0.5% to about 3% v/v Nafion. In some embodiments, the solution contains 0.1, 0.2, 0.3, 0.4, 0.5, 0.6, 0.7, 0.8, 0.9, 1.0, 1.1, 1.2, 1.3, 1.4, 1.5, 1.6, 1.7, 1.8, 1.9, 2.0, 2.1, 2.2., 2.3, 2.4, 2.5, 2.6, 2.7, 2.8, 2.9, 3.0, 3.1, 3.2, 3.3., 3.4, 3.5, 3.6, 3.7, 3.8, 3.9, 4.0, 4.1, 4.2, 4.3, 4.4, 4.5, 4.6, 4.7, 4.8, 4.9, or 5% v/v Nafion.

[0047] The EIS may be recorded using the Squidstat Plus (Admiral Instruments) analyzer at open circuit potential and a frequency range from 10 ⁵ to 10' ² Hz using an alternated current signal of 10 mV amplitude. The changes in resistance to charge transfer (RCT), before and after exposure of the biosensor to contaminated biofluids (e.g., human blood serum and saliva samples), can used to provide qualitative and quantitative results for COVID-19 diagnosis. The RCT response will increase due to the binding between ACE2- SARS-CoV-2 or peptide-SARS-CoV-2 and this response can used to calibrate the doseresponse between the virus and the detection moiety (FIG. 4).

[0048] Accordingly, the presently disclosed devices may be configured to generate a signal that can be assessed via electrochemical impedance spectroscopy (EIS) when a current is run through the electrode. The device may be configured to generate a signal when the detection moiety is bound to SARS-CoV-2 spike protein that is different from the signal that the device generates when the detection moiety is not bound to SARS-CoV-2 spike protein.

[0049] In some embodiments, the device is configured to accept a current that is generated by a potentiostat, and to generate a signal from the current that can be detected by the potentiostat. The potentiostat may be an external component, such as of the conventionally used device. However, in some embodiments, the present devices include a miniaturized potentiostat that can perform at least the essential functions of a traditional, external potentiostat, including generating and delivering a current to the electrode, and detecting the signal produced by the device when a current is run through the electrode.

[0050] In some embodiments, the present devices can be used to detect SARS-CoV- 2 on cell phones through the use of an app and a miniaturized potentiostat.

[0051] The device may be wearable, and as such may include an adhesive on the back face of the substrate that is compatible with a subject’s skin.

[0052] The present devices retain a favorable degree of stability following storage. For example, the devices may retain about 50% of their original sensitivity following storage at 8°C for 48 hours. In some embodiments, the devices may retain more than 50% of their original sensitivity following storage at -20°C up to about 10 days. The devices may also retain about 50% of their original sensitivity following storage at -20°C for about 10 days.

[0053] The devices according to the present disclosure are extremely sensitive relative to prior devices for the detection of pathogens. In some embodiments, the limit of detection of the present devices is about 3-10 PFU of pathogen per mL of a biological sample containing the pathogen. For example, the limit of detection of the present devices may be about 10, 9, 8, 7, 6, 5, 4, or 3 PFU of pathogen per mL of a biological sample containing the pathogen. In some embodiments, the limit of detection of SARS-CoV-2 of the present devices is about 3-10 fg of SARS-CoV-2 spike protein per mL of a biological sample containing the pathogen. For example, the limit of detection of the present devices may be about 10, 9, 8, 7, 6, 5, 4, or 3 fg of SARS-CoV-2 spike protein per mL of a biological sample containing SARS-CoV-2. In one embodiment, the limit of detection of SARS-CoV-2 of the present devices is about 2.8 fg of SARS-CoV-2 spike protein per mL of a biological sample containing SARS-CoV-2.

[0054] Also provided are wearable articles comprising a device according to any of the embodiments described herein. The article is may be, for example, a self-adhesive bandage, a band for wrapping around an appendage of a subject (including an upper or lower arm, a calf, or a forearm, for example), a glove, or a mask. When in the form of a mask, the article may incorporate a device according to the present disclosure at a location that will contact droplets that are expelled from a subject’s mouth or nose during breathing, sneezing, or coughing. The article may include a colorimetric functionality that displays a certain color or that changes color when the device detects the presence of SARS-CoV-2. [0055] The present disclosure also pertains to methods for assessing the presence of a pathogen, such as SARS-CoV-2, in a biological sample comprising contacting a device according to the present disclosure with the biological sample; exposing the device to an electrical current in order to generate a signal from the device; and, assessing the signal that is generated by the device electrochemical impedance spectroscopy (EIS) in order to determine the absence or presence of the pathogen in the biological sample. In certain embodiments, the electrical current is an alternating current (AC). The alternating current may have an amplitude of about 5 to about 15 mV. For example, the alternating current may have an amplitude of about 5, 6, 7, 8, 9, 10, 11, 12, 13, 14, or 15 mV. In a specific embodiment, the alternating current has an amplitude of about 10 mV.

[0056] Hereinafter, the present disclosure will be described in more detail through Examples, which are intended to be illustrative to the present disclosure, although present disclosure is not limited to the Examples.

Example 1: Detection of SARS CoV-2

[0057] An electrode is screen-printed onto a paper substrate. The electrode is functionalized with thiol groups. An ACE2 protein that further includes an N-terminus cysteine group is bonded to the thiol-functionalized electrode via disulfide bonds. Bovine serum albumin is used to block the remaining exposed surfaces of the electrode.

[0058] The device comprising the electrode and the substrate is contacted with blood serum from a subject suspected of being infected with SARS CoV-2. A potentiostat is used to deliver a current to the electrode, and the resulting EIS signal is recorded using a Squidstat Plus analyzer at open circuit potential and a frequency range from 10 ⁵ to 10' ² Hz using an alternated current signal of 10 mV amplitude. The changes in resistance to charge transfer (RCT), before and after exposure of the electrode to the blood serum is used to provide qualitative and quantitative results that enable COVID-19 diagnosis. FIGS. 6A and 6B provide the results of the assessment.

Example 2 — Device with Portable Potentiostat

[0059] Inventors next developed a simple, inexpensive, and rapid test for detection of SARS-CoV-2, dubbed “DETECT 1.0” (DETECT 1.0 (Detection through Electrochemical Technology for Enhanced COVID-19 Testing prototype 1.0) (FIG. 7) The device transforms biochemical information from a specific molecular binding event between the SARS-CoV-2 spike protein (SP) and ACE2 into an electrical signal that can easily be detected.

As illustrated in FIG. 7A-7C, DETECT 1.0 enables diagnosing neat saliva and NP/OP swab samples infected with SARS-CoV-2 (FIG. 7A). FIG. 7B provides a schematic for the preparation of the electrodes. Briefly, the screen-printed electrodes in a three-electrode configuration cell (counter electrode - CE, working electrode - WE, and reference electrode - RE) were printed in phenolic paper circuit board or filter paper with conductive carbon and Ag/AgCl inks. The WE was functionalized with glutaraldehyde to enable anchoring of ACE2, which was stabilized by the addition of bovine serum albumin. Detection was improved by adding a Nafion permeable membrane enabling chemical preconcentration of cation species and protecting the electrode’s surface against biofouling with proteins, lipids, and other macromolecules present in biological samples. FIG. 7C provides a cost and detection time comparison matrix between DETECT 1.0 and existing FDA-approved antigen, serological and molecular tests (Government, A.C. (2020). Information of Coronavirus (COVID-19) Testing; Service, R. (2020); Administration, U.S.F.& D. (2020). In Vitro Diagnostics EUAs).

[0060] DETECT 1.0 (also referred to herein as DETECT) uses electrochemical impedance spectroscopy (EIS), an electrochemical technique extensively utilized for the characterization of functionalized electrode surfaces and the transduction of biosensors. In our test, the EIS transducer signal reported the selective interaction/binding between the biological receptor immobilized on the electrode surface (i.e., ACE2) and its binding element (i.e., spike protein). The binding between these two molecules causes a change in interfacial electron transfer kinetics between the redox probe, ferricyanide/ferrocyanide in solution and the conducting electrode sites. This electrochemical change is then detectable by monitoring the charge-transfer resistance (RCT), the diameter of the semi -arc on the Nyquist plot, which correlates with the number of targets bound to the receptive surface. The selectivity of an EIS biosensor mostly relies on the specificity between the target and the recognizing bioelement immobilized on the electrode surface and its robustness through the designed architecture surfaces to minimize non-specific binding of the analyte or adsorption of other biomolecules in solution.

[0061] The electrochemical device was designed to explore the remarkable binding affinity of SARS-CoV-2 spike protein (SP) to ACE2, its receptor in the human body. FIGS. 8A-8E provide information concerning the characterization and calibration of the DETECT 1.0 device. FIG. 8A is a schematic representation of the DETECT diagnostic process. FIG. 8B provides a cyclic voltammetry plot, and FIG. 8C provides a Nyquist plot (inset shows the zoomed region of the curve with the semi -arc) of all functionalization steps showing progressive increased resistivity between the bare electrode (in black) and the four modification steps: addition of glutaraldehyde (in red), functionalization of ACE2 (in blue), addition of the blocking agent bovine serum albumin (in green), and addition of the Nafion permselective membrane (in purple). FIG. 8D provides Nyquist plots for different SP concentrations ranging from 100 fg mL-1 to 100 ng mL-1 with 10-fold increments in neat saliva from a healthy donor (negative result by RT-qPCR). The inset shows the linearized correlation between normalized RCT values and the concentration of SP exposed to the electrode. FIG. 8E provides Nyquist plots for tittered inactivated virus solutions at concentrations ranging from 101 to 106 PFU mL-1 with 10-fold increments. The upper left inset shows the linearized correlation between the normalized RCT values and concentration of inactivated virus in solution. The lower right inset shows a zoomed region of the curve with the Nyquist plots’ semi-arc (RCT). The analytical curves presented in FIGS. 8D and 8C were based on triplicate measurements. All data were recorded using the eCHIP version of DETECT.

[0062] We designed the electrochemical device to explore the remarkable binding affinity of SARS-CoV-2 spike protein (SP) to ACE2, its receptor in the human body (Andersen et al., 2020; Yang et al., 2020) (FIG. 8A). The working electrode (WE), where the (electro)chemical reaction/interaction takes place and is converted to a detectable analytical signal, was functionalized by a drop-casting method. Enzyme immobilization was achieved by cross-linking ACE2 using the bifunctional chemical cross-linker glutaraldehyde (GA) (Barbosa et al., 2014). This dialdehyde reacts mainly with the primary amino groups of proteins, for example, the s-amino group of lysine residues or the N-terminal group of the protein chain (Pereira et al., 2018). We used bovine serum albumin (BSA) to block the electrode’s surface after immobilization of ACE2. BSA is a functionally inert protein with a high density of superficial lysine residues that is commonly used for biosensor development (Pereira et al., 2018).

[0063] Using these well-established protocols for bioelectrode development, we first added GA for 1 hour at 37 °C to fully cover the carbon electrode surface generating a cross-linked polymer that enables the covalent anchoring of ACE2 at 37 °C for 1.5 hours (FIG. 7B). Next, BSA was added to the surface of the electrode for 30 minutes at 37 °C to block possible remaining active sites (i.e., working electrode’s surface areas that were not functionalized with ACE2) thus preventing nonspecific adsorption to the GA layer by other proteins. We also incorporated an additional functionalization step using a 1.0% Nafion solution (FIG. 11) to create a protective polymeric membrane enhancing the robustness of the biosensor (Mauritz and Moore, 2004). Interestingly, Nafion increased up to 2-fold the sensitivity of the biosensor, particularly when used at a concentration ranging between 1.0% and 1.5% (FIG. 11). Given these results, we selected 1.0% Nafion (wt%) for subsequent optimization steps because of its optimal analytical response to low reagent usage ratio (Mauritz and Moore, 2004). This anionic membrane enables small positively charged species to cross and preconcentrate close to the biosensing surface. The Nafion layer also enhanced the robustness of DETECT by protecting against biofouling of the electrodic surface when exposed to the sample’s complex matrix (e.g., proteins, lipids, and other macromolecules present in biological samples) that may interfere with the detection (e Silva et al., 2020; Mauritz and Moore, 2004)

[0064] The optimized protocol generated the best analytical signal for the detection of SARS-CoV-2 in human biofluid samples (FIG. 7A). It consists of the following 4-steps: 1) modifying the working electrode with the immobilizing agent (GA); 2) covalent attachment of the recognition agent ACE2; 3) addition of the stabilization and active site blocking agent BSA; and 4) incorporating the permselective membrane (Nafion). A detailed protocol describing biosensor preparation, including the production of the screen-printed devices and functionalization, is provided in the Examples, infra.

[0065] Our test can be performed at room temperature with minimal equipment and reagents, and costs $4.67 to produce [$0.07 to produce the bare electrode, $4.50 to functionalize the electrode with the recognition agent ACE2, and $0.10 to coat the electrode with GA, BSA, and Nafion used (Figure 1c)]. The overall cost of DETECT may be further reduced through recombinant production of ACE2 and ACE2 variants (Chan et al., 2020). Our technology is also highly scalable, as the electrodes can be rapidly mass-produced by using commercially available screen-printers. One laboratory-sized unit is able to produce 35,000 electrodes daily (1.05 M electrodes/month) and this could scaled-up to 10.5 billion electrodes monthly with only 10,000 screen-printers (Table 1). These estimates take into account both the time needed to print the electrodes and all functionalization steps (i.e., 1 hour for GA functionalization, 1.5 hours to incorporate ACE2, 0.5 hours for BSA, and 1 hour for Nafion; total of 4 hours). However, it must be noted that these steps can be fully automated into a production line for industrial purposes, drastically reducing time requirements. [0066] Table 1. DETECT 1.0: a scalable technology. Scalability of the production of electrodes over a one-year period with laboratory screen-printers and industrial screenprinters. The numbers shown reflect both the number of printed electrodes over time considering the printing rate of the screen printer and all functionalization steps (addition of the anchoring agent, anchoring the recognition agent, addition of the blocking agent, and generation of the perm-selective membrane, the latter of which may take 4 additional hours after electrodes are printed.

Number of electrodes produced

Time -

1 Screen- 100 Screen- 10,000 Screen- 1 Industrial 100 Industrial 10,000 Industrial

(days) printer printers printers Screen-printer Screen-printers Screen-printers

1 35,000 3,500,000 350,000,000 150,000 15,000,000 1,500,000,000

7 245,000 24,500,000 2,450,000,000 1,050,000 105,000,000 10,500,000,000

15 525,000 52,500,000 5,250,000,000 2,250,000 225,000,000 22,500,000,000

30 1,050,000 105,000,000 10,500,000,000 4,500,000 450,000,000 45,000,000,000

365 12,775,000 1,277,500,000 127,750,000,000 54,750,000 5,475,000,000 547,500,000,000

The key steps required for the electrode’s functionalization were optimized and characterized (FIGS. 8B-C). Additionally, we evaluated the incubation time (i.e. time of exposure of the sample to the biosensor to enable sensitive detection), and whether a centrifugation/dilution step was needed to detect SARS-CoV-2 in complex biological samples such as saliva ⁸ . These optimization steps revealed that an additional centrifugation step was not needed (FIG. 12) since the use of neat saliva yielded similar results to those obtained using centrifuged samples. FIG. 12 provides calibration curves of the SP ranging from 500 fg mL' ¹ to 100 ng mL' ¹ , where the saliva samples were incubated using three different setups: (i) direct use, i.e., without any pretreatment; (ii) neat saliva after 2 min of centrifugation at 10,000 rpm; and (iii) after simple 1 : 1 dilution in PBS. We can observe that the use of neat saliva allows the same detection efficacy and greater linear behavior when compared to the other pretreatment conditions. All measurements were recorded in triplicate using eChips.

[0067] These results demonstrated that our approach is robust and can directly use human samples (NP/OP or saliva) without a prior pretreatment step, thus allowing the application of DETECT for streamlined and rapid point-of-care diagnosis. We selected 2 minutes as the optimal incubation period of the sample on the working electrode’s surface for sensitive SARS-CoV-2 detection in samples considering the detectability and analytical frequency of the tests (FIG. 13). Our very minimal incubation time requirement (2 minutes) confirms the favorable configuration of the modified electrode that allows rapid interaction kinetics between the SP and immobilized ACE2 [kinetics constant rate of 10 ⁴ M ^-1 s ^-1 in its natural environment (Yang et al., 2020)]. Overall, DETECT provides a result in 4 minutes (2 minutes of sample incubation + 2 minutes to perform the EIS analysis), which is vastly faster than methods currently available for diagnosing COVID-19 (FIG. 7C). It is important to note that the total time required to run each blank is an additional 4 minutes. However, we did not take this into account in our testing time calculations because the blanking step can be done before analyzing clinical samples, and we can use the RCT values obtained for the blanks (PBS or VTM) to compare with the patient sample values.

[0068] Taking into account the optimal analytical conditions evaluated (Table 1), we built calibration curves for free SP (FIGS. 8D, 14 A, 14B) and heat-inactivated virus using the normalized RCT response, defined by the following equation:

Z - Z _o normalized R _CT = — - — ZQ where Z is the RCT of the sample and Zo is the RCT of the respective blank solution: phosphate buffer saline (PBS), virus transportation medium (VTM), or healthy saliva. The normalization process of RCT aims to correct eventual fluctuations in the sensor operation, such as the temperature at the testing point or variations due to analyst operation.

Table 2. Analytical parameters of DETECT 1.0.

Parameter Value

Linear concentration range (SP in PBS) 10 fg mL' ¹ - 100 ng mL' ¹

Linear concentration range (SP in VTM) 10 fg mL' ¹ - 1 ng mL' ¹

Linear concentration range (SP in saliva) 100 fg mL' ¹ - 100 ng mL' ¹

Limit of detection (SP in PBS) 2.18 fg mL' ¹

Limit of detection (SP in VTM) 6.29 fg mL' ¹

Limit of detection (SP in saliva) 1.39 pg mL' ¹

Limit of quantification (SP in PBS) 7.26 fg mL' ¹

Limit of quantification (SP in VTM) 20.96 fg mL' ¹

Limit of quantification (SP in saliva) 4.63 pg mL' ¹

Working concentration range (IV in VTM) 10 ¹ - 10 ⁶ PFU mL' ¹ Limit of detection (IV in VTM) 1.16 PFU mL' ¹

Limit of quantification (IV in VTM) 3.87 PFU mL' ¹

[0069] The dose-response curve for the free SP in PBS solution ranged from 1 fg mL' ¹ to 10 pg mL' ¹ (FIG. 14 A). A linear concentration range from 10 fg mL' ¹ to 100 ng mL' ¹ was obtained (R ² = 0.993) and limits of detection (LOD) and quantification (LOQ) were calculated as 2.18 fg mL' ¹ and 7.26 fg mL' ¹ SP based on signal to noise ratios (S/N=3) and (S/N=10), respectively. We built a dose-response for the free SP in VTM medium at a concentration range from 10 fg mL' ¹ to 100 pg mL' ¹ (FIG. 14B). A linear concentration range from 10 fg mL' ¹ to 1 ng mL' ¹ was obtained (R ² = 0.995) and limits of detection (LOD) and quantification (LOQ) were calculated as 6.29 fg mL' ¹ and 20.96 fg mL' ¹ SP based on the signal to noise ratio (S/N=3) and (S/N=10), respectively. When performed in neat saliva, the calibration curve was built at a concentration ranging from 100 fg mL' ¹ to 100 ng mL' ¹ (FIG. 8D). The calculated LOD and LOQ were 1.39 pg mL' ¹ and 4.63 pg mL' ¹ , respectively. The higher LODs obtained in saliva and VTM are consistent with the increased sample complexity compared to PBS solution.

[0070] The RCT values of Nyquist plots were extracted by the application of an equivalent circuit (FIG. 15). The equivalent circuit comprises two semi-arc regions observed in the Nyquist plots, where the first is a non-defined semi-arc at a high-frequency range due to inhomogeneity or defects in the electrode modification step (during drop-casting functionalization) and considerably small (RCT -10 Q) (Bertok et al., 2019; Uygun and Ertugrul Uygun, 2014). The second parallel component of the equivalent circuit comprises an RCT, whose signal intensity was proportional to the logarithm of the SP/virus concentration, and also presented a Warburg element to describe the mass transport (diffusional control).

[0071] The concentration range of SP detected by our device was 10-1,000 times lower than that reported in previous studies (Rashed et al., 2021; Seo et al., 2020), thus underscoring the sensitivity of our approach. To assess the diagnostic capability of DETECT, we calibrated our biosensor using tittered solutions of inactivated SARS-CoV-2 ranging from 10 ¹ to 10 ⁶ PFU mL' ¹ (Figure 2e). DETECT exhibited high sensitivity presenting a limit of detection (LOD) of 1.16 PFU mL' ¹ , which corresponds to the order of 10° RNA copies pL' ¹ (Rao et al., 2020; Uhteg et al., 2020), a viral load that correlates with the initial stages of COVID-19 (i.e., 2 to 3 days after onset of symptoms)(Zou et al., 2020). Thus, DETECT’s LOD and LOQ values are comparable to those of gold-standard approaches such as Real Star®

SARS-CoV-2, CDC COVID-19, and e-Plex® SARS-CoV-2 (Uhteg et al., 2020) with the advantage of detecting symptomatic and asymptomatic individuals at the earliest stages of the infection allowing for rapid decision-making and the subsequent use of more appropriate and effective countermeasures. To ensure the repeatability, stability, and reproducibility of the results, we carried out three different experiments. First, 21 successive EIS measurements of the medium (PBS) were performed using the same device to verify the drift of the EIS response, yielding an RSD value of 5.3% (FIG. 16). These results demonstrated that the device exhibits a repeatable and stable response. Next, a measurement of open circuit potential before and after the addition of 1.0 ng mL' ¹ of SP in PBS was recorded for 60 minutes (FIG. 17) and a small change in the potential (RSD value of 0.76%) was observed during the 30 minutes of exposure to 1.0 ng mL' ¹ of SP solution. Finally, the reproducibility of DETECT was evaluated by analytical curves in the range of 1 pg mL' ¹ to 1 ng mL' ¹ of SP and the analytical sensitivity of 10 electrodes from different batches was assessed (FIG. 18). An RSD value of 6.8% was obtained, indicating that the electrochemical device fabrication and functionalization protocols display high reproducibility.

[0072] Next, we evaluated the stability of DETECT at different temperature storage conditions (25 °C, 8 °C, and -20 °C) over 10 days (FIG. 19). Analytical curves were generated with SP at a concentration ranging from 1 pg mL' ¹ to 1 ng mL' ¹ and the sensitivity was normalized by the mean value of the three different biosensors used immediately after the functionalization steps. The biosensors stored at room temperature did not detect the SP after 24 hours due to loss of enzymatic activity (FIG. 19). The sensors stored at 8 °C were stable after 24 hours, but after 48 hours presented decreased sensitivity (around 50% of the initial response) keeping this low sensitivity for 7 days (FIG. 19). Biosensors stored at -20 °C exhibited the most promising results since they were as sensitive as those used right after functionalization even after 96 hours and retained 50% of their sensitivity after 10 days of storage (FIG. 19).

[0073] Next, the performance of DETECT was assessed using both SARS-CoV-2- positive and negative clinical samples from the Hospital of the University of Pennsylvania (HUP) (Tables 3 and 5, below), including a highly contagious SARS-CoV-2 UK B.1.1.7 variant (Tables 3 and 4, below). All samples were heat-inactivated at 56 °C for 1 hour. The effect of heat inactivation of SARS-CoV-2 samples on the analytical response of our biosensor was evaluated through measurements taken before and after sample inactivation at 56 °C for 1 hour (FIG. 20). The results indicated that thermal inactivation affected the ability of SP to bind to ACE2, since a decrease of up to 60% was detected in the analytical response for sample 2 after heat inactivation (FIG. 20). These results indicate that heat-inactivated clinical samples with very low viral titers may fall below our current limit of detection. Rath and Kumar (Rath and Kumar, 2020) demonstrated using molecular dynamics simulations that temperatures >50°C trigger the closing of the spike receptor binding motif (RBM), which buries the receptor binding residues preventing contacts between the SP and the ACE2 receptor. These insights may help explain the results obtained upon thermal inactivation of our biosensor (FIG. 20). However, despite this decrease in SP binding to ACE2 upon heat inactivation, the sensitivity of our method still enabled accurate viral detection in clinical samples containing a range of viral titers (FIG. 8E).

[0074] We also observed that centrifuging the samples did not lead to increased impedimetric detection of the SP (FIG. 12). Therefore, the NP/OP and saliva samples were used in VTM and PBS, respectively, following the Food and Drug Administration (FDA) recommendation for regulatory applications. Of note, the detectability of impedimetric measurements after 2 minutes of incubation of the sample on the working electrode’s surface was as high as longer incubation times of 5 and 10 minutes (FIG. 13), thus demonstrating DETECT’s fast interaction kinetics between the SP and functionalized WE, as discussed above. Thus, we selected 2 minutes of incubation and set as a cut-off value a 10% change in the RCT when compared to the blank solution. Such a cut-off threshold takes into account the LOQ obtained for inactivated virus analysis (FIG. 8E), thus allowing discrimination between SARS-CoV-2 negative and SARS-CoV-2 positive samples (Tables 3 and 5, below).

[0075] In blinded tests, we analyzed 139 NP/OP swab samples (in VTM) obtained from patients after heat-inactivation, 109 of which were COVID-19 positive and 30 CO VID- 19 negative as determined by RT-qPCR and clinical assessment (Table 3, below). DETECT demonstrated high sensitivity, specificity and accuracy for NP/OP (83.5%, 100% and 87.1%, respectively; Table 4) and saliva (100%, 86.5% and 90.0%, respectively; Table 4) samples. DETECT missed a single sample, which presented a viral count lower than its LOD (10' ¹ RNA copies pL' ¹ ). It is worth noting that although the heat inactivation protocol decreased the response of our biosensor due to the inactivation of SP (FIG. 20), the outstanding sensitivity of DETECT (Table 2) enabled high detectability (Table 2) and hit rate (Table 4, below). Out of the 12 negative NP/OP swab samples present in our sample set, 100% were confirmed as SARS-CoV-2 negative by DETECT (data not shown). In addition, the highly contagious SARS-CoV-2 UK variant B.1.1.7 was obtained from a government testing site in Philadelphia (Tables 4 and 5, below). DETECT successfully identified this sample as positive with a normalized RCT value of 1.10 (Table 3), thus underscoring its ability to detect emerging mutant variants of SARS-CoV-2.

Table 3. Diagnosis of NP/OP samples from patients of the Hospital of the University of Pennsylvania (HUP) with COVID-19 symptoms using DETECT 1.0.

COVID-19 DETECT RCT

NP/OP Sample ID RT-qPCR

Status 1.0

257 + + . o

312 + + + 0.177

357 . . . 0.024

255 . . . 0.013

307 + + + 0.161

Mock-1 . . . o

312 + + + 0.159

356 . . . 0.021

263 + + + 0.149

290 + + + 0.123

314 + + + 0.263

256 + + + 0.210

334 + + + 0.254

257 + + + 0.155

251 + + + 0.171

309 + + + 0.136

332 . . . 0.029

336 + + + 0.16

290 + + + 0.145 Mock-2 0

353 0

Mock-3 0

262 + + + 0.118

360 0

348 + + + 0.261

358 0

363 0.080

346 + + + 0.128

348 + + + 0.168

361 0.0722

UPHS COVID 1 + + + 0.275

UPHS COVID 4 + + 0

UPHS COVID 5 + + + 0.255

UPHS COVID 6 + + + 0.200

UPHS COVID 7 + + + 0.106

UPHS COVID 8 + + 0.098

UPHS COVID 9 + + + 0.109

UPHS CO VID 10 + + + 0.272

UPHS COVID 11 + + 0

UPHS CO VID 12 + + 0

UPHS CO VID 13 + + 0

UPHS CO VID 14 + + + 0.167

UPHS CO VID 15 + + + 0.114

UPHS CO VID 16 + + + 0.137

UPHS CO VID 17 + + + 0.143 UPHS CO VID 18 + + + 0.241

UPHS CO VID 19 + + + 0.241

UPHS COVID 20 + + + 1.011

UPHS CO VID 21 + + + 1.082

UPHS COVID 22 + + + 0.183

UPHS COVID 23 + + + 0.725

UPHS COVID 24 + + + 0.107

UPHS COVID 25 + + + 0.175

UPHS COVID 26 + + + 0.104

UPHS COVID 27 + + + 0.110

UPHS COVID 28 + + + 0.171

UPHS COVID 32 + + + 0.129

UPHS CO VID 35 + + + 0.191

UPHS CO VID 36 + + + 0.745

UPHS CO VID 37 + + + 0.110

UPHS CO VID 39 + + + 0.130

UPHS COVID 40 + + + 0.261

UPHS COVID 42 + + + 0.108

UPHS COVID 44 + + + 0.179

UPHS COVID 45 + + 0.024

UPHS COVID 46 + + + 0.114

UPHS COVID 47 + + + 0.103

UPHS COVID 48 + + + 0.146

UPHS COVID 49 + + + 0.101

UPHS COVID 50 + + + 0.126

UPHS COVID 51 + + + 0.132 UPHS COVID 52 + + + 0.131

UPHS COVID 53 + + + 0.191

UPHS COVID 54 + + + 0.107

UPHS CO VID 55 + + + 0.225

UPHS COVID 56 + + + 0.104

UPHS COVID 57 + + + 0.188

UPHS CO VID 58 + + + 0.194

UPHS COVID 59 + + 0

UPHS COVID 60 + + 0

UPHS CO VID 61 + + 0

UPHS COVID 62 + + + 0.139

UPHS COVID 63 + + + 0.229

UPHS COVID 64 + + + 0.481

UPHS COVID 65 + + 0

UPHS COVID 66 + + + 0.152

UPHS COVID 67 + + + 0.197

UPHS COVID 68 + + + 0.105

UPHS COVID 69 + + + 0.101

UPHS COVID 70 + + + 0.129

UPHS CO VID 71 + + + 0.101

UPHS COVID 72 + + 0

UPHS COVID 73 + + + 0.159

UPHS COVID 74 + + 0.003

UPHS COVID 75 + + + 0.222

UPHS COVID 76 + + + 0.101

UPHS COVID 77 + + + 0.236 UPHS COVID 78 + + + 0.342

UPHS COVID 79 + + 0.012

UPHS COVID 80 + + + 0.102

UPHS CO VID 81 + + 0.031

UPHS COVID 82 + + + 0.302

UPHS COVID 83 + + + 0.127

UPHS COVID 84 + + + 0.165

UPHS COVID 85 + + + 0.130

UPHS COVID 86 + + + 0.102

UPHS COVID 87 + + + 0.221

UPHS COVID 88 + + + 0.196

UPHS COVID 89 + + 0.035

UPHS COVID 90 + + + 0.137

UPHS CO VID 91 + + 0

UPHS COVID 92 + + + 0.13

UPHS COVID 93 + + + 0.184

UPHS COVID 94 + + 0

UPHS COVID 95 + + + 0.115

UPHS COVID 96 + + + 0.101

UPHS COVID 97 + + + 0.236

UPHS COVID 98 + + + 0.630

UPHS COVID 99 + + + 0.582

UPHS CO VID 100 + + + 0.102

700067571 0.05

601112 0

466721776 0 633400 - - - 0.06

349368993 - - - 0.02

442134375 . . . 0

468444690 - - - 0.07

346496821 - - - 0.073

357098938 . . . 0

440956795 . . . 0

363618695 . . . 0

042 . . . o

044 . . . 0.028

046 . . . 0.078

047 . . . 0.044

049 . . . o.O9

053 . . . 0.077

054 . . . 0.077

100667644* + + + 1.098 individual infected with a highly contagious SARS-CoV-2 UK variant B.1.1.7.

[0076] DETECT demonstrated high sensitivity, specificity and accuracy (96.2%, 100% and 97.4%, respectively; Table 4).

Table 4. Positive and negative values obtained by RT-qPCR, and sensitivity, specificity, and accuracy of DETECT 1.0 using NP/OP and saliva samples.

RT-qPCR

(NP/OP Positive Negative Total Sensitivity Specificity Accuracy

(N=109*) (N=30) (N=139)

91/109

Positive 91 0 91 (83.5%) 121/139

(87.1%)

30/30

Negative 18 30 48 (100%) RT-qPCR

DETECT _o . tSalival . Sensitivity Specificity Accuracy

' Positive Negative Total

(N=13) (N=37) (N=50)

*Clinical sample set includes a highly contagious SARS-CoV-2 UK variant B.l.1.7 from a patient.

[0077] To evaluate DETECT’s diagnostic efficacy in a more complex biological environment, we tested saliva samples from 50 patients (Table 5) under the same conditions used for the NP/OP swab samples.

Table 5. Diagnosis of saliva samples from patients of the Hospital of the University of Pennsylvania (HUP) with COVID-19 symptoms using DETECT 1.0.

ED Saliva _CO VID-19 Status RT-qPCR DETECT 1.0 RCT

Sample ID

1 - - - 0

2 - - - 0

3 + + + 0.261

4 . . . 0.099

6 + + + 0.573

9 - - - 0

14 - + + 0.252

21 - + + 0.121

24 - - - 0.050

33 + + + 0.303

41 - - - 0.069

42 + - + 0.751

43 - + + 0.154

44 - - - 0.076

45 - + + 0.176

46 - - - 0.096 51 - - - 0

52 - - - 0

53 - - - 0

54 - - - 0

55 - - - 0.035

56 - - + 0.232

58 + + + 0.223

69 - - - 0.081

70 + + + 1.103

72 - - - 0.083

77 + + + 0.181

79 - - - 0.012

82 + + + 0.302

90 + + + 0.137

91 + + + 0.132

700067571 . . . o.O8

453299679 - - - 0

468349915 - - - 0.03

633400 - - - 0

349368993 - - - 0

464333574 + + + 0.134

468444690 - - - 0.08

335835294 + + + 0.102

346496821 - - - 0.072

357098938 - - - 0

440956795 . ₀

363618695 . ₀

041 - - - 0

042 - - - 0

043 - - - 0

044 - - - 0

046 - - - 0

047 - - - 0

096 + + + 0.293 [0078] The greater complexity of saliva, compared to swab samples, is known to hinder the accurate detection of infectious agents (Jamal et al., 2020; Zou et al., 2020). Saliva is a biofluid that is susceptible to large variations in composition depending on different factors such as the ingestion of food and drinks prior (30-60 minutes) to sample collection, which can lead to the dilution of the saliva matrix, and the insertion of exogenous molecular species that may interfere with accurate detection. Even using highly heterogenous saliva samples, the sensitivity of DETECT remained high (100%), however false positives led to decreased specificity (86.5%), and an accuracy of 90.0% (Table 4). The latter results may be explained by potential interactions between ACE2, which is a carboxypeptidase and amino acid transporter, and other biomolecules that can be found in neat biofluids, such as regulatory peptides and peptide hormones (e.g., angiotensin, bradykinin, ghrelin, apelin, neurotensin, and dynorphin) (Turner, 2015). Thus, we believe the performance of DETECT will improve when using fresh saliva samples at the point-of-care. It is worth noting that among the SARS-CoV-2-positive saliva samples, our test identified as positive two samples that had been previously erroneously detected as negative by RT-qPCR, therefore indicating that DETECT may help correctly diagnose COVID-19 in samples previously misdiagnosed by other methods.

[0079] Several key analytical features were used to compare the performance of DETECT with respect to other electrochemical methods reported in the literature (Table 6).

Table 6. Comparison of methods reported to CO VID-19 diagnosis.

Lowest Number of

Biological Price Time

Sensor Technique Concentration Clinical Reference Target (USS) (min) Detected Samples

SARS-

DETECT CoV-2 „ .

_{1 A} EIS .. 2.8 fg mL 151 4.67 4 This work

1.0 spike ° protein /i _ra 1

SARS- .. (Torrente-

CoV-2 AC FK an igen 500 pg mL' ¹ 16 - 10 Rodriguez et

RapidFlex ^0CP - ^EIS ™deocap _S l * d protein

SARS- , IgM and (Torrente-

CoV-2 OCP EIS IgG 250 ng mL' ¹ 16 - 10 Rodriguez et

RapidFlex antibodies al., 2020)

SARS- .• (Torrente- DPV and C-reactive _T , , ₁ „ ,

CoV-2 „ _TC , . 50 ng mL 16 - 10 Rodriguez et

D JT71 OCP-EIS protein

RapidFlex ^{1 &} al., 2020) SARS-

231 RNA (Alafeef et al.,

SCC CoV-2 48 copies pL' ³ 2020)

RNA ¹

SARS-

200 RNA (Zhao et al.,

DPV CoV-2 33 copies pL' ¹ ° 2021)

RNA SARS- CoV-2 (Rashed et al.,

EIS 0.1 mg mL' ¹ 4 spike ³ 2021) protein IgM and

(Yakoh et al.,

SWV IgG 1 pg mL' ¹ 17 2021) antibodies

CRISPR E gene and 10 RNA copies (Broughton et

DETECTR technology N gene pL ¹ al., 2020)

Colorimetric (Moitra et al.,

N gene 0.18 ng pL assay ^{1 3U} 2020)

Localized

2.26 x 10 ⁴ surface (Qiu et al.,

RdRp plasmon RNA copies 2020) resonance pL ¹

DNA Synthetic nanoscaffoldRNA (Jiao et al.,

0.96 pmol L' based hybrid ¹ conserved 2020) chain reaction region

20-200 RNA (Yan et al.,

RT-LAMP orflab copies pL ¹ 2020)

100 RNA (Baek et al.,

RT-LAMP N gene copies pL ¹ 2020)

EIS - Electrochemical impedance spectroscopy; DPV - Differential pulse voltammetry; OCP-EIS - Open-circuit potential -electrochemical impedance spectroscopy; SCC - Signal conditioning circuit; SWV - Square-wave voltammetry; RT-LAMP - Reverse transcription loop-mediated isothermal amplification.

Our method provides the highest sensitivity (LOD of 2.8 fg mL' ¹ ) for the detection of SARS- CoV-2 spike protein with excellent time of detection and overall cost (Table 6). Additionally, the robustness of DETECT was evaluated in a large clinical sample set (Tables 3 and 5), and all results were compared with those obtained by RT-qPCR (Table 4), thus highlighting the reliability of our method. All experiments described thus far (e.g., detection of SARS-CoV-2 spike protein and clinical samples) were performed using the eChip version of the electrode (e.g., FIG. 8, Table 4, FIGS. 11 and 12). After successfully applying the eChip (composed of printed circuit board) to clinical samples (Tables 3 and 5) and obtaining robust and sensitive results (Table 2), we sought to construct an optimized electrode composed of a material that was more accessible and inexpensive to enable scale-up production of DETECT. We selected filter paper as the main component of electrochemical paper-based analytical device (ePAD) as it is easy to handle (maleable), accessible, and inexpensive [paper filter costs $0.50 per 1 m ² whereas printed circuit board (PCB) costs $40.00 per 1 m ² ] (Ataide et al., 2020; Ozer et al., 2020). We adapted and demonstrated the applicability of ePAD in a portable potentiostat connected to a smart device (FIG. 9 A). We used the screen-printing method to fabricate the electrodes and combined wax -printing technology to pattern the electrochemical cell onto the paper filter. Thus, the ePAD is composed of more accessible and low-cost material, enabling scalable manufacturing and on-demand testing at the point-of-care (Ataide et al., 2020; Ozer et al., 2020).

[0080] To demonstrate the portability of DETECT and its potential as a point-of- care diagnostic test, we adapted and demonstrated its applicability in a portable potentiostat connected to a smart device. FIG. 9A and FIG. 9B illustrate the of the miniaturized and portable DETECT 1.0 for rapid point-of-care diagnosis of COVID-19. FIG. 9 represents an image of the mobile device-compatible handheld DETECT 1.0 during real-time sample analysis. FIG. 9B provides Nyquist plots obtained using ePAD coupled to a smart-device for different concentrations of SP ranging from 1 pg mL-1 to 100 ng mL-1. The inset shows the calibration curve for the normalized RCT values of the different concentrations of SP.

[0081] In this case, a paper-based electrode (ePAD) was used, as this is a more accessible and low-cost material for onsite analysis. However, the cellulosic structure of the paper presents higher wettability compared to the phenolic circuit boards (eChip), causing the absorption of the sample by the electrode’s paper surface. Therefore, in order to enhance the detectability (/.< ., the LOD) of DETECT, we added 2.5-fold increased volumes of the modifiers (GA, ACE2, BSA, and Nafion) on the surface of the WE during the fabrication process. This approach allowed higher sensitivity towards the detection of SP, which was used to generate a calibration curve (FIG. 9B). We attribute the enhanced detection (7-fold increase) of the paper-based version of DETECT compared to the phenolic-based electrode (eChip) to the higher amount of recognition element (ACE2) used on the working electrode’s surface. However, it is worth noting that the eChip version already demonstrated excellent performance at detecting SARS-CoV-2 (Tables 2 and 4) and, although its sensitivity can be further increased by using a higher concentration of ACE2 (FIG. 9B), this would increase the cost of the test since recombinant ACE2 accounts for 95% of the final cost of DETECT 1.0 (FIG. 7C).

[0082] DETECT diagnoses COVID-19 at its early stages compared to serological tests, which take 5-7 days to ensure reliable detection of IgG and IgM antibodies ¹⁷ . Our device presented higher accuracy, specificity, and selectivity than most existing methods available for SARS-CoV-2 detection ⁿ . Our biosensor is inexpensive and portable, enabling decentralized diagnosis at the point-of-care. The time of detection of our approach (4 minutes) is significantly lower than existing diagnostic tests ¹⁰ ’ ^11,18 , _anc | could potentially be lowered even more by using engineered versions of human ACE2 with enhanced selective binding towards SARS-CoV-2 SP ⁷ . The use of such ACE2 variants would also help reduce the rate of false positives in complex biofluids such as saliva ^{7 19} > ²⁰

[0083] DETECT presented higher accuracy, specificity, and selectivity than most existing electrochemical methods available for SARS-CoV-2 detection (Table 6) (Uhteg et al., 2020). We also assessed DETECT’s specificity in cross-reactivity assays by exposing our sensor to the following seven different viruses: three coronaviruses (MHV - murine hepatitis virus, HCoV-OC43 - human coronavirus OC43, and human coronavirus 229E; Table S4) and four non-coronavirus viral strains (H1N1 - A/California/2009, H3N2 - A/Nicaragua, Influenza B - B/Colorado, HSV2 - herpes simplex virus-2; Table 7).

Table 7. Cross-reactivity analysis of DETECT 1.0 when exposed to other coronaviruses and non-coronavirus strains.

ED Saliva Sample ID DETECT 1.0 RCT

MHV - 0

HCoV-OC43 - 0

229E - 0.06

H1N1 - 0

H3N2 - 0.04

Influenza B - 0

HSV2 - 0

We did not detect cross-reactivity events against any of the viruses tested (RCT < 10%) (Table 7) thus further highlighting the translatability of our diagnostic test. Our biosensor is inexpensive and portable, enabling decentralized diagnosis at the point-of-care. The time of detection of our approach (4 minutes) is significantly lower than existing diagnostic tests (Kaushik et al., 2020; Rashed et al., 2021; Uhteg et al., 2020), and could potentially be lowered even more by using engineered versions of human ACE2 with enhanced selective binding towards SARS-CoV-2 SP (Chan et al., 2020). The use of such ACE2 variants would also help reduce the rate of false positives in complex biofluids such as saliva (Chan et al., 2020; Glasgow et al., 2020; Sorokina et al., 2020).

[0084] DETECT can also be multiplexed to allow detection of other emerging biological threats such as bacteria, fungi, and other viruses. Thus, our technology serves as a platform for the rapid diagnosis of COVID-19 and future endemic/pandemic outbreaks at the point-of-care. Its low cost, speed of detection, scalability, and implementation using smart devices and telemedicine platforms may facilitate much needed population-wide deployment. Additional Information Concerning Materials and Methods

[0085] The electrochemical sensors were screen-printed in a three-electrode configuration cell on two accessible substrates (i) a qualitative filter paper and (ii) phenolic paper circuit board material. Electrically conductive carbon and Ag/AgCl inks were used for the screen-printing process to fabricate the working/auxiliary electrodes and reference electrodes, respectively. The working electrode’s carbon surface was modified using the drop casting method. First, 4 pL of 25% glutaraldehyde (GA) solution was added for 1 hour at 37 °C for the formation of a cross-linked polymer, which enabled the anchoring of ACE2 (4 pL at 0.32 mg mL' ¹ ), then incubated at 37°C for 1.5 hours. Next, 4 pL of bovine serum albumin (BSA) at 1 mg mL' ¹ was added and the WE was allowed to dry for 0.5 hours at 37 °C to stabilize the enzyme through the co-reticulation and allow blockage of potential remaining active sites of the carbon electrode to avoid any nonspecific adsorption by other proteins to the glutaraldehyde layer and ensure stabilization of the ACE2 tertiary structure. Both concentrations of GA and BSA solutions were used in excess to ensure the complete functionalization and blocking of the WE’s surface.

[0086] To test ACE2 conformational integrity after the addition of BSA to the functionalized electrode, the response of the electrode to angiotensin II, ACE2’s natural substrate (FIG. 10) was analyzed. FIG. 10 provides Nyquist plots showing the response of the modified eChip to different concentrations of angiotensin II, the natural substrate of ACE2, ranging from 1 pg mL' ¹ to 10 pg mL' ¹ . A sensitive linear response was observed in the range of 1 pg mL' ¹ to 10 pg mL' ¹ of angiotensin II, demonstrating that our anchoring and stabilization strategies maintained the functionality of ACE2’s active sites and revealing that the biosensor architecture did not obstruct ACE2. The calibration curve was built based on triplicate measurements. The results showed that the anchoring and stabilization steps were effective on the WE’s surface and there was no loss of ACE2’s conformation integrity since it was able to interact with its natural substrate.

[0087] Since the objective was to simplify detection of SARS-CoV-2 in complex biological samples, such as neat saliva and NP/OP swabs, we added a 1% Nafion solution as an extra protective layer. Nafion solution, an anionic and permselective membrane, is commonly used to enhance the sensitivity and robustness of electrochemical sensors. In our study, the membrane formed by 1% Nafion solution enhanced the sensitivity of DETECT 1.0 (FIG. 11), by enabling chemical preconcentration of cation species and protecting the electrode’s surface against biofouling of biomolecules present in biological samples, such as protein, lipids, and other macromolecules f FIG. 11 depicts calibration curves for the SP (ranging from 1 pg mL' ¹ to 100 ng mL' ¹ ) that were built with different Nafion concentrations (0, 1%, 3% and 5%; wt%) to test the effect of the permselective membrane on the analytical signal of DETECT 1.0. The optimal concentration of Nafion found was found to be 1% (wt%). It is worth noting that Nafion at 5% created a thicker film on the working electrode’s surface that did not present high adherence to the surface and detached during the impedimetric measurements. Therefore, it was not possible to measure different concentrations of SP in solution.

[0088] The effect of each modifier layer on the electrochemical response of our modified electrode was characterized, recording cyclic voltammetry (CV) and EIS measurements in the presence of 5 mmol L' ¹ potassium ferricyanide/ferrocyanide (the redox probe), Figures 2B and 2C. These results demonstrated that the peak current signal of the redox probe decreased when using CV and the resistance to charge transfer increased after each functionalization step. The decrease in the peak current signal occurs due to the addition of nonconductive materials (e.g., proteins) that block the active sites of the electrodic surface, hindering the kinetics of charge transfer of the redox probe.

[0089] We next evaluated the stability of the biosensor by measuring 6 successive EIS measurements in undiluted healthy human saliva (negative result for COVID-19 by RT- qPCR) and the same sample spiked with 1 pg mL' ¹ free SP. Relative standard deviations of 3.58% and 5.21% were obtained, respectively. These results demonstrate that the developed biosensor presents a very stable architecture and provide effective robustness for the detection of SP in complex sample. [0090] We proceeded to analyze patient samples obtained from symptomatic patients at the Hospital of the University of Pennsylvania. We tested 35 NP/OP swabs (Table 3) and 31 saliva samples (Table 5) that were complementary confirmed as SARS-CoV-2 positive or SARS-CoV-2 negative by RT-qPCR.

[0091] Chemicals. All chemicals were of analytical grade and used without additional purification. Solutions were obtained by dissolving or diluting the reagents in appropriate electrolytes prepared in deionized water. Human angiotensin converting enzyme 2 (ACE2) was purchased from GenScript (USA), sulfuric acid, potassium chloride (KC1), potassium ferricyanide K3[Fe(CN)e], potassium ferrocyanide K4[Fe(CN)e], bovine serum albumin (BSA), Nafion (5%) and glutaraldehyde (25%) were obtained from Sigma Aldrich (USA), and phosphate buffer saline (PBS) solution was purchased from VWR (USA). Viral transport medium (VTM) was obtained from Thermo Fisher. Conductive carbon and Ag/AgCl inks were acquired from Creative Materials, USA. SARS-CoV-2 spike protein was kindly donated by Scott Hensley (University of Pennsylvania) and the inactivated samples were donated by Sara Cherry, Michael Feldman and Ronald Collman (University of Pennsylvania).

[0092] Fabrication of electrochemical devices. The electrochemical sensors were screen-printed in a three-electrode configuration cell (dimensions: 1.8 x 1.2 cm) on two accessible substrates (i) a qualitative filter paper and (ii) phenolic paper circuit board material. First, specific patterns were wax printed on A4 size filter paper using a commercial Xerox ColorQube 8570 printer (Xerox, Brazil). The patterns consist of small white rectangles (1.1 x 1.7 cm) to delimit the electrochemical cell on paper substrates. In a single A4 size paper, 80 patterns were printed, thus affording 80 disposable ePADs. Following, the screenprinting process was performed in the previously patterned paper using electrically conductive carbon and Ag/AgCl inks (Creative Materials, USA) to fabricate the working/auxiliary electrodes and reference electrodes, respectively. The printed filter paper sheets were then placed in a thermal oven for 30 minutes at 100 °C. The heating process induces the curing step of the conductive tracks and melts the deposited wax layer that then penetrated in the cellulosic structure, forming a 3D hydrophobic barrier around the hydrophilic patterns (electrochemical cell). Finally, the electrochemical paper-based analytical devices (ePADs) were cut with scissors and the backside of the devices was covered with a transparent tape to prevent solution leakage through the device and to add structural integrity. The phenolic paper is a material largely used as a printed circuit board substrate. The boards were washed thoroughly with deionized water and isopropyl alcohol. The screen-printing process on the paper phenolic resin was performed using the same design and dimension reported for the filter paper platform. The electrochemical circuit board-based devices (eChip) present a rigid substrate and low wettability that dispenses the use of a hydrophobic barrier. After the curing step of printed electrodes, they were cut into small pieces (2 x 2 cm) and a non-conductive layer was applied to delimit the electrode area.

[0093] Modification of the eChips and ePADs. The electrodes were washed with deionized water and cleaned/activated electrochemically by cyclic voltammetry (CV) recorded in sulfuric acid solution (0.1 mol L' ¹ ) in the potential range from -1.3 to 1.5 V at the scan rate of 100 mV s' ¹ for 5 cycles. The eChips were dried at room temperature and 4 pL of GA solution (25% in water) was added on the surface of the working electrode using the drop-casting method. After 1 hour, 4 pL of ACE2 solution (0.32 pg mL' ¹ ) prepared in PBS medium was added on top of the working electrode and left to dry at room temperature for 1.5 hours. Subsequently, 4 pL of BSA solution (1 mg mL' ¹ ) was added on the surface of the working electrode to stabilize the protein and block unspecific sites of the electrode. After 30 minutes, 4 pL of Nafion solution (1.0% in PBS) was added to the working electrode’s surface and left for 1 hour before the final washing with deionized water. The ePADs were modified using the same protocol but applying 2.5-fold higher volume of the modifying agent solutions.

[0094] Electrochemical measurements. SquidStat Plus (Admiral Instruments) and Sensit Smart (PalmSens) potentiostats controlled by a laptop running the software SquidStat and a smartphone running the software PSTouch, respectively, were used to record all electrochemical data. The electrodes were characterized by CV technique using a mixture of 5 mmol L' ¹ potassium ferricyanide/ferrocyanide in the medium of 0.1 mol L' ¹ KC1 solution prior and after electrode modification using a potential range of 0.7 to -0.3 V at the scan rate of 50 mV s' ¹ . Electrochemical impedance spectroscopy (EIS) was used to characterize the biosensor and for SARS-CoV-2 detection. The EIS measurements were performed using 200 pL of a mixture of 5 mmol L' ¹ ferricyanide/ferrocyanide prepared in 0.1 mol L' ¹ KC1 solution added after the sample incubation on the electrode (10 pL of OP/NP or saliva samples) and the gentle washing process using PBS solution to remove the unbound SP/SARS-CoV-2. A sinusoidal signal was applying in the frequency range between 10 ⁵ and 10' ¹ so using a typical open circuit potential of 0.15 V and an amplitude of 10 mV at room temperature. [0095] Optimization tests. We evaluated the main experimental parameters and processes that affect the efficiency of the developed biosensor. For modification steps, both GA and BSA were used at high concentration levels to ensure the complete recovery of the electrode surface providing the best condition to covalently attachment of ACE2 and its stabilization. The formation of permselective membrane was evaluated by using different Nafion concentrations in the range of 0.5 to 3.0 wt%. After the biosensor preparation, we evaluate its response to different concentrations (1 pg ml; ¹ - 10 pg ml; ¹ ) of angiotensin II (Angll), the natural substrate of ACE2, to verify if the anchoring and stabilization strategies maintain the biological activity of ACE2. To assess the kinetics of interaction between SP and the architecture of the modified electrode, we carried out calibration curves ranging from 1 pg mL' ¹ to 1 ng mL' ¹ SP using different times of incubation (from 1-10 minutes) to obtain the best analytical response to DETECT 1.0. Finally, the need for sample pretreatment of saliva samples was evaluated using 3 different approaches: (i) direct use of raw saliva, (ii) 2 minutes of centrifugation at 10,000 rpm, and (iii) simple dilution of sample 1 : 1 (v/v) with PBS. We performed this study with saliva samples because it presents greater matrix complexity (high viscosity and content of proteins, lipids, and other biomolecules that can cause biofouling of the electrodic surface) when compared to NP/OP swab samples.

[0096] Cross-reactivity experiments. Cross-reactivity assays were carried out by exposing the sensor to three coronaviruses (MHV - murine hepatitis virus at 10 ⁸ PFU mL' ¹ , HCoV-OC43 - human coronavirus OC43 at 10 ⁴ PFU mL' ¹ , and human coronavirus 229E at 10 ⁷ PFU mL' ¹ ), and four non-coronavirus viral strains (H1N1 - A/California/2009, H3N2 - A/Nicaragua, Influenza B - B/Colorado, HSV2 - herpes simplex virus-2, all at 10 ⁵ PFU mL' ¹ ) were used to assess the specificity of our biosensor. The conditions used were the same as those used for all SARS-CoV-2 samples: incubation time of 5 minutes, 10 L of virus sample, and EIS measurements as specified above (Electrochemical Measurements section).

[0097] Quantification and statistical analysis. Cyclic voltammetry and electrochemical impedimetric spectroscopy measurements are presented as an average of 3 or 7 different replicates for each condition and it is described in each figure caption. Graphs were created and statistical tests conducted in GraphPad Prism 8.

Example 3 — Cohort Study

[0098] To assess the clinical performance of the instant diagnostic platform, an accuracy study was conducted for detecting SARS-CoV-2 in anterior nare samples and compared the results obtained to those from RT-PCR. [0099] Clinical enrollment was performed over the period of 10 weeks between

January and March 2021, following the period with the most COVID-19 cases in Philadelphia (from November to December 2020), where an average of 40,000 tests were performed with around 500 daily COVID-19 cases confirmed (prevalence of -1.25% from November to December) (FIG. 22A). All samples collected for the study were aliquoted and frozen at -80°C promptly after collection. The anterior nare samples were immersed in VTM following the Food and Drug Administration (FDA) recommendation for regulatory applications. A total of 321 nare swab samples were analyzed from incoming patients that agreed to donate their samples.

[00100] Clinical samples were incubated for 2 minutes onto the surface of the electrode, as this was the optimal amount of time needed to ensure viral detection using the inventive RAPID system (Torres MDT, et al. (2021) Low-cost biosensor for rapid detection of SARS-CoV-2 at the point-of-care. Matter 4: 1-14). The configuration of the modified electrode favors rapid interaction kinetics between the SARS-CoV-2 spike protein and immobilized ACE2 (kinetics constant rate of 10 ⁴ M ^-1 s ^-1 (Yang J, et al. (2020) Molecular interaction and inhibition of SARS-CoV-2 binding to the ACE2 receptor. Nat Commun 11 (1):4541). The RAPID system provides a result within 4 minutes (2 minutes of sample incubation + 2 minutes to perform the EIS analysis), which is faster than currently available methods for diagnosing COVID-19 (Bhalla N, et al. (2020) Opportunities and Challenges for Biosensors and Nanoscale Analytical Tools for Pandemics: COVID-19. ACS Nano 14(7):7783-7807). An additional 4 minutes was needed to run each blank, however we did not consider this when calculating our testing time because the blanking step is performed prior to clinical sample analysis. Before starting our clinical study, we calibrated our biosensor using tittered solutions of inactivated SARS-CoV-2 ranging from 10 ¹ to 10 ⁶ PFU mL' ¹ . FIG. 22A shows the number of tests, number of cases, and prevalence of COVID-19 in Philadelphia as per official records (CO VID data for Pennsylvania (2021) Commonw Pennsylvania). FIG. 22B shows the number of tests, number of cases, and prevalence in the present retrospective cohort study. Complete clinical data paired with the gold-standard method (RT-PCR) were used to confirm the CO VID-19 status of each of the 321 samples (FIG. 22B). A total of 31 positive and 290 negative COVID-19 samples were obtained. As provided in Table 8, below, RAPID demonstrated high sensitivity (80.7%), specificity (89.0%), and accuracy (88.2%). Table 8. Clinical assessment of RAPID detection of SARS-CoV-2 Positive and negative values obtained by RT-qPCR, and sensitivity, specificity, and accuracy of RAPID 1.0 using nare samples.

RT-qPCR

RAPID Sensitivity Speciricily Prevalence Accuracy

Positive Negative Total (N = 31) (N = 290) (N = 321) i P->ositive 2 n5c o 32n 5 c-77 25/31

M

Negative 6 258 264

[00101] The presence or absence of symptoms and other medical conditions did not interfere with the results obtained with RAPID, and no correlation was found between other medical conditions, race, gender or age with the false positives and negative data obtained. Compared to other electrochemical methods, molecular tests, colorimetric assays, and diagnostic tests reported in the literature, RAPID presents the highest sensitivity reported to date (LOD of 2.8 fg ml; ¹ SARS-CoV-2 spike protein). In addition, RAPID displays a rapid detection time for SARS-CoV-2 (4 minutes) and is low cost (<US$5.00) (Parihar A, et al. (2020) Point-of-Care Biosensor-Based Diagnosis of CO VID-19 Holds Promise to Combat Current and Future Pandemics. ACS Appl Bio Mater 3(11):7326-7343).

[00102] Currently available diagnostic tests (prior to the present disclosure) do not provide an accurate, rapid, and affordable diagnosis of COVID-19. For instance, commercial SARS-CoV-2 antigen tests only detect virus concentrations characteristic of later stages of the disease at which patients are already highly infectious (Corman VM, et al. (2021) The Lancet Microbe, doi: 10.1016/S2666-5247(21)00056-2), thus not accurately controlling viral spread. RT-PCR, the current gold standard for testing, presents optimal accuracy 3-5 days after the onset of symptoms (Bourn Y, et al. (2021). Lancet Infect Dis. doi: 10.1016/S1473- 3099(21)00132-8). The affordability aspect is also particularly important in order to ensure health equity and increased access to valuable tools, such as diagnostic tests, for preventing viral spread in disadvantaged communities.

[00103] In the present cohort study, the performance of RAPID was assessed using 321 anterior nare swab samples from a diversified pool of subjects with age ranging from 18 to 78 years old, different races, genders, CO VID-19 related symptoms and other medical conditions (Table 9, below).

Table 9. Demographic information of the subjects tested. Total Cohort Positive Subjects (n Negative Subjects

Median Age 37 (13) 36 (14) 37 (13)

Gender

Male 91 (28%) 9 (29%) 82 (28%)

Female 230 (72%) 22 (71 %) 208 (72%)

Race

Caucasian 133 (41 %) 13 (42%) 120 (41 %)

African American 147 (46%) 16 (52%) 131 (45%)

Hispanic 13 (4%) 2 (6%) 11 (4%)

Other 29 (9%) 0 29 (10%)

Medical Problems

Asthma 66 (21 %) 7 (23%) 59 (20%)

Hypertension 61 (19%) 8 (26%) 53 (18%)

History of Smoking 41 (13%) 3 (10%) 38 (13%)

Diabetes 28 (9%) 5 (16%) 23 (8%)

No Medical History 176 (55%) 16 (52%) 160 (55%)

Symptoms

Cough 93 (29%) 14 (45%) 79 (27%)

Headache 68 (21 %) 11 (35%) 57 (20%)

Fever/Chills 67 (21 %) 11 (35%) 56 (19%)

Shortness of Breath 33 (10%) 3 (10%) 30 (10%)

No Symptoms 127 (40%) 6 (19%) 121 (42%)

The clinical prevalence of positive COVID-19 cases in the set of samples analyzed was 9.7%, which is higher than the mean observed for the same period in Philadelphia (1-2%; FIG. 22). We did not find statistical correlations between the erroneously diagnosed samples by RAPID and the clinical status or any relevant information obtained from the participants (Table 1 and Data SI). False-positive results may be due to the use of angiotensin-converting enzyme inhibitors or angiotensin receptor blockers that may interact with RAPID’s ACE2 -modified electrode. However, the lack of information about the medication usage of participants limited our ability to draw such a correlation. Another important source of potential errors is the self-collection of swabs that took place during testing, as this may lead to samples with no (or very few) viral counts even though the patient was COVID-19 positive and had a medium-to-high viral load.

[00104] Additional details concerning the performance of the present cohort study are as follows.

RAPID biosensor preparation.

[00105] The testing platform comprised two components: the electrochemical sensor and a potentiostat. The electrochemical sensors were prepared following established protocols (Torres MDT, et al. (2021) Low-cost biosensor for rapid detection of SARS-CoV-2 at the point-of-care. Matter 4:1-14). Briefly, the portable devices were screen-printed in a three-electrode configuration cell on phenolic circuit board material (2x2 cm). Electrically conductive carbon and Ag/AgCl inks were used for the screen-printing process to fabricate the working/auxiliary electrodes and reference electrodes, respectively. The working electrode’s carbon surface was modified using the drop-casting method. First, 4 pL of 25% glutaraldehyde (GA) solution was added for 1 hour at 37 °C to allow the formation of a crosslinked polymer, which enabled subsequent anchoring of ACE2 (4 pL at 0.32 mg mL' ¹ ). ACE2 was then incubated at 37 °C for 1.5 hours. Next, 4 pL of bovine serum albumin (BSA) were added at 1 mg mL' ¹ and allowed the working electrode (WE) to dry for 0.5 hours at 37 °C to stabilize the enzyme and block potential active sites present within the carbon electrode, in order to avoid nonspecific adsorption of other proteins to the glutaraldehyde layer and ensure stabilization of the ACE2 tertiary structure. Since the goal was to simplify the detection of SARS-CoV-2 in complex biological samples, such as anterior nare swabs, a 1 wt. % Nafion solution was added as an additional protective layer. Nafion, an anionic and selective membrane that allows the permeation of cationic species, is commonly used to enhance the sensitivity and robustness of electrochemical sensors (Mauritz KA, Moore RB (2004) State of Understanding of Nafion. Chem Rev 104(10):4535-4586). In the present study, the membrane formed by 1 wt. % Nafion solution enhanced the sensitivity of RAPID 1.0, by enabling chemical preconcentration of cation species and protecting the electrode’s surface against biofouling by macromolecules present in biological samples, such as proteins and lipids (e Silva RF, et al. (2020) Simple and inexpensive electrochemical paper-based analytical device for sensitive detection of Pseudomonas aeruginosa. Sensors Actuators B Chem 308: 127669).

Anterior nare sample collection and processing.

[00106] The collection of the anterior nare samples was performed by the subjects tested under supervision by clinical research staff at the Penn Presbyterian Medical Center (PPMC). All the demographic information, as well as the presence or absence of symptoms of the individuals tested, are shown in Table 9, above. The samples were stabilized and stored in viral transport medium (VTM) following CDC guidelines (CDC SOP#: DSR-052-05). The anterior nare samples were maintained on ice during the collection period, separated into identical aliquots and subsequently stored at -80 °C until tested. Care was taken to ensure samples were thawed only once before testing. RAPID test for SARS-CoV-2 diagnosis.

[00107] SquidStat Plus (Admiral Instruments) and Multi Autolab Ml 01 (NOVA 2.1) potentiostats controlled by a laptop running the software SquidStat and a smartphone running the software PSTouch, respectively, were used to record all electrochemical data. The electrodes were characterized by Cyclic Voltammetry (CV) and EIS techniques using a mixture of 5 mmol L' ¹ potassium ferricyanide/ferrocyanide in 0.1 mol L' ¹ KC1 solution before and after electrode modification with glutaraldehyde, ACE2, BSA, and Nafion. CVs and EIS were recorded using a potential ranging from 0.7 to -0.3 V at the scan rate of 50 mV s' ¹ and a frequency ranging from 10 ⁵ to 10' ¹ Hz using a sinusoidal signal with 10 mV of amplitude at room temperature, respectively.

[00108] RAPID reports the selective binding between ACE2, the biological receptor immobilized on the electrode surface, and SARS-CoV-2 spike protein, its binding element. The interaction between these two molecules causes a change in interfacial electron transfer kinetics between the redox probe, ferricyanide/ferrocyanide in solution and the conducting electrode sites. This electrochemical change is then detectable by monitoring the chargetransfer resistance (RCT) and the diameter of the semi -arc on the Nyquist plot, which correlates with the number of spike protein molecules bound to the electrode’s surface (5). The selectivity of an EIS biosensor mostly relies on the specificity between the target and the recognizing bioelement immobilized on the electrode surface, and the robustness of the latter to minimize non-specific binding or adsorption of other biomolecules present in biofluids. The EIS measurements were performed using 200 pL of a mixture of 5 mmol L' ¹ ferricyanide/ferrocyanide prepared in a 0.1 mol L' ¹ KC1 solution added after incubating the clinical sample (10 pL of anterior nare sample) for 2 minutes on electrode surface. A gentle washing step using PBS was performed to remove the sample and any unbound SARS-CoV- 2. For the EIS measurement, a sinusoidal signal was applied at room temperature in the frequency range between 10 ⁵ and 10' ¹ so using a typical open circuit potential of 0.15 V and an amplitude of 10 mV.

[00109] RAPID enables viral detection of SARS-CoV-2 in anterior nare samples stored in VTM within 4 minutes (2 minutes of incubation and 2 minutes of measurement time). Each test was performed at room temperature requiring only a potentiostat, PBS, and a redox probe solution (/.< ., mixture of 5 mmol L' ¹ ferricyanide/ferrocyanide prepared in 0.1 mol L' ¹ KC1 solution). Each RAPID test cost $4.67 to produce ($0.07 to produce the bare electrode, $4.50 to functionalize the electrode with the recognition agent ACE2, and $0.10 to coat the electrode with GA, BSA, and Nafion). RAPID display high sensitivity (1.16 PFU ml; ¹ ) comparable to that of RT-PCR assays (1-10 PFU ml; ¹ ).

RT-PCR analysis.

[00110] For the RT-PCR assays, RNA was extracted and purified using the QIAmp DSP Viral RNA Mini Kit (Qiagen) from a 140 pL aliquot. The first step of this process chemically inactivated the virus from the anterior nare samples under highly denaturing conditions (guanidine thiocyanate) and was performed in a biosafety cabinet under BSL-2 enhanced protocols. The remainder of the process was performed at the lab bench under standard conditions using the vacuum protocol as per manufacturer’s instructions. Next, RNA present in the samples was analyzed in duplicate using the TaqPath™ 1-Step RT-qPCR reagent (Life Technologies) on the Quantstudio 7 Flex Genetic Analyzer (AB I). The oligonucleotide primers and probes for detection of 2019-nCoV were selected from regions of the virus nucleocapsid (N) gene. The panel was designed for specific detection of the 2019-nCoV viral RNA (two primer/probe sets, N1 and N2). An additional primer/probe set to detect the human RNase P gene (RP) in control samples and clinical specimens was also included in the panel (2019-nCoVEUA-01). RNaseP is a single copy human-specific gene and can indicate the number of human cells collected.

Prospective cohort study design and participants.

[00111] The performance of RAPID was assessed using both SARS-CoV-2-positive and negative samples from an ambulatory COVID-19 testing site for the general public, led by staff at the Penn Presbyterian Medical Center (PPMC). All participants underwent anterior nare testing for SARS-CoV-2 using CLIA-approved RT-PCR by PPMC staff for testing, and subsequent to this testing underwent study procedures. Adult (age > 17 y) subjects were eligible if they (1) underwent PPMC staff-led testing immediately prior to study enrollment, (2) were deemed competent for written consent, (3) were English fluent, and (4) did not have any contraindications to anterior nare samples collection procedures, such as recent facial surgery or active head and neck cancer. Subjects completed standard written consent, and then completed a short survey including demographic information and recent infectious symptoms, if any. Subjects then underwent anterior nasal swabbing supervised by trained clinical research coordinators. This work was approved by the University of Pennsylvania Institutional Review Board (IRB 844145). Diagnosis and statistical analysis.

[00112] The RCT values of Nyquist plots obtained using Squidstat Plus (Admiral Instruments) and Multi Autolab Ml 01 (Metrohm) were extracted by the application of an equivalent circuit using the softwares Zahner Analysis and Nova 2.1, respectively. The equivalent circuit comprises two semi-arc regions observed in the Nyquist plots, where the first is a non-defined semi-arc at a high-frequency range due to inhomogeneity or defects in the electrode modification step (during drop-casting functionalization) and considerably small (RCT -10 Q) (Uygun ZO, Ertugrul Uygun HD (2014) A short footnote: Circuit design for faradaic impedimetric sensors and biosensors. Sensors Actuators B Chem 202:448-453; Bertok T, et al. (2019) Electrochemical Impedance Spectroscopy Based Biosensors: Mechanistic Principles, Analytical Examples and Challenges towards Commercialization for Assays of Protein Cancer Biomarkers. ChemElectroChem 6(4):989-1003). The second parallel component of the equivalent circuit comprises an RCT, whose signal intensity was proportional to the logarithm of the concentration of SARS-CoV-2 and presented a Warburg element to describe the mass transport (diffusional control).

[00113] To diagnose a given sample, the normalized RCT, defined by the following equation, was used: where Z is the RCT of the sample and Zo is the RCT of the blank solution (VTM).

[00114] A cut-off value was set as a 10% change in the RCT when compared to the blank solution. Such a cut-off threshold considers the LOQ value previously obtained for inactivated virus, thus allowing discrimination between SARS-CoV-2 negative and SARS- CoV-2 positive samples.

[00115] The presently disclosed RAPID system is an inexpensive and portable alternative to existing COVID-19 tests, allowing for decentralized diagnosis at the point-of- care. The fast detection (4 min) enabled by the present approach is significantly lower than commercially available tests, and could potentially be lowered even more by using alternative recognition agents, such as engineered versions of human ACE2 with enhanced selective binding towards SARS-CoV-2, or engineered receptors to the SARS-CoV-2 spike protein, such as antibodies (Chan KK, et al. (2020). Science (80- ) 369(6508): 1261-1265).

[00116] Finally, RAPID can be multiplexed to allow detection of emerging biological threats such as bacteria, fungi, and other viruses, simply by adding other recognition agents and modifying the electrodes disposition (array configuration). Its ability to detect minimal viral particles within a sample allows diagnosing COVID-19 at the onset of the infection. Collectively, its low-cost, rapid detection time, and high analytical sensitivity make RAPID an exciting alternative tool for high-frequency COVID-19 testing and effective population surveillance (Mina MJ, et al. (2020) N Engl J Med 383(22):el20).

[00117] References. The following references may be relevant to the present disclosure:

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