The present disclosure pertains to devices and methods for detecting infection by a pathogen in a mammalian subject
COVID-19, the severe acute respiratory illness caused by SARS-CoV-2, has led to over 6.57 million deaths worldwide and continues affecting millions of people, primarily in low-income countries, low-resource settings, and communities with low vaccination coverage. Emerging SARS-CoV-2 variants may have harmful interactions with host immunity, as well as increased infectivity, disease severity, and mortality. Low-cost and rapid response technologies that enable accurate, frequent testing of SARS-CoV-2 variants are crucial for outbreak prevention and infectious disease control. Biosensor technologies represent an alternative low-cost approach for detecting infectious diseases including COVID-19. The most widely used substrate for manufacturing electrical circuits and consequently electrodes is the printed circuit board (PCB). PCBs contain Cu, Al, and Sn, consisting of nearly 28% metal. These metals in PCBs are more than 10 times purer than the metals in rich-content minerals. Because PCBs are used extensively and discarded afterward, the recycling of PCBs is not trivial. Moreover, the high percentage of nonmetals in PCBs is around 70%, consisting mostly of thermoset resins and reinforcing materials; these materials pose a particularly challenging recycling problem. The network structure of thermoset resins hinders them from being remelted or reformed. Due to inorganic fillers like glass fiber, with considerably lower fuel efficiency, incineration is not appropriate for treating nonmetals. Nonmetal components of PCBs are mostly disposed of in landfills, which can waste resources and produce significant secondary contamination.
Therefore, there is an urgent need to develop approaches to detect and diagnose both viral and bacterial infections that are also compatible with environmental considerations.
Both types of herpes simplex virus (HSV), HSV-1 and HSV-2, are prevalent in humans, and both cause neonatal infections. Furthermore, these viruses can establish lifelong latency in the sensory neuronal ganglia. Subsequent reactivation of latent virus may cause significant health problems and result in viral transmission to healthy individuals. HSV-1, also known as oral herpes, infects the lips, mouth, eyes, and brain, while HSV-2, also known as genital herpes, is associated mainly with genital infections. Herpes simplex virus type 2 (HSV-2) infection is almost exclusively sexually transmitted.
The World Health Organization (WHO) recently estimated the global prevalence of HSV-1 in individuals aged 0-49 years to be 66.6%, or more than 3.7 billion people who have been infected by HSV-1. Additionally, the WHO estimates the global prevalence of HSV-2, which is transmitted almost exclusively through sexual contact, to include 13.2% of the world's population, or 491.6 million people aged 15-49 years. The attachment of the virus to the cell surface initially involves two glycoproteins on the HSV envelope, glycoprotein C (gC), and to a lesser extent, glycoprotein B (gB). Glycoprotein D (gD), found within the viral envelope, then binds to host cell receptors, initiating a sequence of events that allow HSV to fuse with the host's cell plasma membrane. Studies of the binding of gD to cell surface receptors have led to an understanding of the interaction between human cell receptors and HSV.
Despite the prevalence of HSV-2 infections, there are currently no rapid tests available to detect this infectious agent. Historically, viral culturing has been the main test used for HSV detection in the clinic. However, recently, molecular methods such as polymerase chain reaction (PCR) have been widely used in clinical practice due to their increased sensitivity and selectivity compared to viral culture. Currently, there are very few FDA-cleared molecular tests available for HSV detection. Examples include PCR-based MultiCode-RTx kit, ProbeTec HSV Qx test, and IsoAmp HSV assay with sensitivity and selectivity ranging from 92.4% to 98.4% and 83.7% to 97.0%, respectively. Other commercial serological methods such as immunoblot (IB), ELISA, Western blotting, and chemiluminescence immunoassay (CLIA) have also been used to detect HSV. However, immunoassays rely on the availability of HSV antibodies, and thus, the sensitivity of these tests is influenced by the amount of time since the infection. Indeed, immunoassays display the highest sensitivity when performed at least 21 days after the initial infection and may improve if performed more than 40 days after the primary infection, thus clearly hindering early HSV diagnosis. In addition, these diagnostic methods are time-consuming, costly, and laborious, requiring highly trained staff and sophisticated laboratory infrastructure.
Rapid and accessible diagnostic technologies could improve the management of HSV infections, particularly in low-resource settings and in labor and delivery wards. In fact, several portable devices have been reported as alternative methods for the diagnosis of HSV, and electrochemical detection methods are attractive for developing such devices. Electrochemical detection has adequate sensitivity and selectivity and can be associated with accessible and portable instrumentation. Generally, these portable diagnostic devices are DNA-based biosensors aiming to detect the viral genetic material. Detecting viral DNA or RNA present in biofluids can lead to base-pairing mismatches and hybridization problems that compromise the selectivity of the tests. Moreover, these methods commonly require preconcentration or amplification protocols to achieve the desired sensitivity, decreasing the ability to conduct rapid, frequent, and inexpensive tests.
Rapid and accessible diagnostic technologies constitute promising approaches to help manage HSV-2 infections.
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 bacterial cellulose and the substrate can include 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, and a chemical cross linker comprising polyethylene glycol (PEG) that enables immobilization of the detection moiety that binds SARS-CoV-2 spike protein on the electrode.
Also provided are wearable articles comprising a device as described herein for assessing the presence of a pathogen, such as, SARS-CoV-2.
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.
Also disclosed herein are devices for assessing the presence of herpes simplex virus (HSV) in a biological sample comprising a substrate that includes 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 that binds HSV glycoprotein gD2, such as nectin-1. A chemical cross-linker may be present in order to enable immobilization of the detection moiety on the electrode.
Also provided are wearable articles comprising a device as described herein for assessing the presence of HSV.
The present disclosure also pertains to methods for assessing the presence of HSV 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 HSV in the biological sample.
Also 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.
Also disclosed are wearable articles comprising a device as described herein.
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.
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.
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.
The entire disclosures of each patent, patent application, and publication cited or described in this document are hereby incorporated herein by reference.
As employed above and throughout the disclosure, the following terms and abbreviations, unless otherwise indicated, shall be understood to have the following meanings.
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.
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”. In addition, when the term “about” precedes a range, it is intended to modify both the recited lower end and the recited upper end of the range. For example, the phrase “about 1 to 5” means “about 1 to about 5”. 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.”
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.
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) (
Accordingly, provided herein are devices comprising a substrate that includes 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, and a chemical cross linker comprising polyethylene glycol (PEG) that enables immobilization of the detection moiety that binds SARS-CoV-2 spike protein on the electrode.
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. In order to address the previously described concerns associated with the use of PCB substrates, the present inventors have developed substrate materials that comprise bacterial cellulose (BC). BC is an extracellular polymer synthesized by species of bacteria belonging to several genera: Agrobacterium, Gluconacetobacter, and Sarcina. As a material, BC is nontoxic and low cost and also exhibits several advantages over commercial paper, such as reduced fiber diameter, no use of chemical methods or processes in its manufacture, and high purity. Accordingly, the substrate may comprise bacterial cellulose. In some embodiments, the substrate comprises bacterial nanocellulose.
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.
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), an amino acid sequence representing a fragment of ACE2, 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 polyethylene glycol (PEG). The PEG be conjugated with graphene oxide, and thereby be used as G-PEG. In certain embodiments, ACE2 or a fragment thereof is immobilized on the electrode via an amide bond between the G-PEG and the N-terminus of ACE2 or the fragment thereof. Full-length ACE2 can be recombinantly generated in E. coli using previously established methods (Chan et al., 2020). A peptide of representing a fragment of ACE2 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 μg.
The present inventors have developed an electrochemical analytical device for detecting infections in real time. 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. For the presently disclosed devices, 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 peptide representing a fragment of ACE2 that interacts directly with SARS-CoV-2 (
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.
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% m/v Nafion. In some embodiments, the Nafion solution contains about 0.5% to about 3% m/v Nafion. In certain embodiments, the Nafion solution contains about 0.5% to about 2% m/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% m/v Nafion.
The EIS may be recorded using the Squidstat Plus (Admiral Instruments) analyzer at open circuit potential and a frequency range from 105 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 dose-response between the virus and the detection moiety.
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.
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.
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.
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.
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.
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 4-10×10−18 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, 9.5, 9, 8.5, 8, 7.5, 7, 6.5, 6, 5.5, 5, 4.5, or 4×10−18 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 4-10×10−18 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, 9.5, 9, 8.5, 8, 7.5, 7, 6.5, 6, 5.5, 5, 4.5, or 4×10−18 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 4.3×10−18 of SARS-CoV-2 spike protein per mL of a biological sample containing SARS-CoV-2.
Also provided are wearable articles comprising a device according to any of the embodiments described herein. The article 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.
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.
As noted above, despite the prevalence of HSV-2 infections, there have been no rapid tests available to detect this infectious agent. The present inventors have developed impedimetric biosensors for the rapid, ultrasensitive detection of HSV-2 (
Accordingly, provided herein are devices comprising a substrate that includes 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 that binds HSV glycoprotein gD2.
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.
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.
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 HSV glycoprotein gD2 is nectin-1 or an antibody. Any of the 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 polyethylenimine (PEI). In certain embodiments, nectin-1 is immobilized on the electrode via an amide bond between the PEI and a carboxyl group on nectin-1. For example, carboxyl groups on nectin-1, when exposed to EDC-NHS, may be activated to form a stable ester, which undergoes a nucleophilic addition with amino groups on the PEI-modified electrode, such that a stable amide bond is formed between the PEI-modified carbon electrode and nectin-1. Human herpes virus entry mediator (HveC), also called human nectin-1 (residues 31-346), can be recombinantly produced, for example, by baculoviruses. Their purification from infected insect cells was described previously.
The present inventors have developed an electrochemical analytical device for detecting infections by HSV in real time. Impedimetric measurements by electrochemical impedance spectroscopy (EIS) provide qualitative and quantitative data for diagnosing HSV 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. For the presently disclosed devices, electrochemical impedance spectroscopy measurements can be used to detect the selective binding of HSV with the detection moiety, such as nectin-1, which interacts specifically with the glycoprotein gD2 of HSV. As disclosed herein, 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 nectin-1 and the glycoprotein gD2 allow detection of the HSV 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, and the functionalized electrode surface.
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 HSV recognition. Thus, the present devices may comprise a blocking layer over the electrode.
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. In other embodiments, the protective membrane can be formed by applying a solution that contains chitosan to the surface of the electrode. The solution can contain, for example, about 0.05% to about 5.0% m/v of the membrane material, e.g., of chitosan. In some embodiments, the solution contains about 0.075% to about 3% m/v, about 0.1% to about 2% m/v, about 0.1% to about 1% m/v, or about 0.25% to about 0.75% m/v of the membrane material, e.g., of chitosan. In some embodiments, the solution contains 0.05, 0.06, 0.07, 0.08, 0.09, 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% m/v of the membrane material, e.g., of chitosan.
The EIS may be recorded using the Squidstat Plus (Admiral Instruments) analyzer at open circuit potential and a frequency range from 105 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 HSV diagnosis. The RCT response will increase due to the binding between the detection moiety (e.g., nectin-1) and HSV glycoprotein gD2 and this response can used to calibrate the dose-response between the virus and the detection moiety.
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 glycoprotein gD2 that is different from the signal that the device generates when the detection moiety is not bound to glycoprotein gD2.
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.
In some embodiments, the present devices can be used to detect HSV on cell phones through the use of an app and a miniaturized potentiostat.
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.
The present devices retain a favorable degree of stability following storage. For example, the devices may retain at least 60, 70, 80, or 90% of their original sensitivity following storage at 4° C. for 48 hours. In some embodiments, the devices may retain at least 60, 70, 80% of their original sensitivity following storage at 4° C. for 120 hours. In some embodiments, the devices may retain more than 50% of their original sensitivity following storage at −20° C. up to about 5 days. The devices may also retain about 50% of their original sensitivity following storage at −20° C. for about 7 days.
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 0.055-0.210 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 0.055, 0.06, 0.065, 0.07, 0.075, 0.08, 0.085, 0.09, 0.095, 0.1, 0.11, 0.12, 0.13, 0.14 0.15, 0.16, 0.17, 0.18, 0.19, 0.20, or 0.21 of pathogen per mL of a biological sample containing the pathogen. In some embodiments, the limit of detection of HSV of the present devices is about 0.015-0.09 fg of glycoprotein gD2 per mL of a biological sample containing the HSV. For example, the limit of detection of the present devices may be about 0.015, 0.016, 0.017, 0.018, 0.019, 0.02, 0.022, 0.024, 0.026, 0.028 0.03, 0.032, 0.034, 0.036, 0.038, 0.04, 0.042, 0.044, 0.046, 0.048, 0.05, 0.052, 0.054, 0.056, 0.058, 0.06, 0.062, 0.064, 0.066, 0.068, 0.07, 0.072, 0.074, 0.076, 0.078, 0.08, 0.082, 0.084, 0.086, 0.088, or 0.09 fg of glycoprotein gD2 per mL of a biological sample containing HSV.
Also provided are wearable articles comprising a device according to any of the embodiments described herein. The article 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 HSV.
The present disclosure also pertains to methods for assessing the presence of a pathogen, such as HSV, 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.
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) (
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.
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.
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.
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 μg.
The present inventors have developed an electrochemical analytical device for detecting infections in real time.
The potentially wearable device detects, through cyclic voltammetry, redox-active metabolites uniquely produced by pathogenic infectious agents. In
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.
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 (
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.
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.
The EIS may be recorded using the Squidstat Plus (Admiral Instruments) analyzer at open circuit potential and a frequency range from 105 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 dose-response between the virus and the detection moiety (
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.
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.
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.
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.
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.
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.
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.
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.
The present invention is further defined in the following Examples. It should be understood that these examples, while indicating preferred embodiments of the invention, are given by way of illustration only, and should not be construed as limiting the appended claims. From the above discussion and these examples, one skilled in the art can ascertain the essential characteristics of this invention, and without departing from the spirit and scope thereof, can make various changes and modifications of the invention to adapt it to various usages and conditions.
In this study, for the sensitive diagnosis of SARS-CoV-2 directly at the point of need, the focus was on an approach that does not require sophisticated instrumentation, using instead a low-cost and portable potentiometer that detects the difference in electrical potential between a stable reference electrode (RE) and the functional working electrode (WE) fabricated on a flexible bacterial cellulose (BC) substrate. The WE is selective for the target analyte, which causes a charge change at its surface upon target species recognition, eliminating the need for a redox probe for analysis.
BC is typically a pure mat of nanosized cellulose fibers. Briefly, for BC production, Gluconacetobacter hansenii was incubated in Hestrin-Schramm (HS) medium with 20 g L−1 glucose. After 27 days, a BC material was collected and treated with 5 mmol L−1 NaOH at 80° C., which was subsequently washed with deionized water abundantly and, after drying, resulted in a clear sheet. The BC substrate was used as a platform for the screen-printing of the electrochemical systems, which were cut to 2.5×2.0 cm dimensions (
To obtain a selective and sensitive biosensor, an evaluation was performed regarding the best of two approaches to anchoring the ACE2 receptor on the carbon screen-printed electrodes. First, the WE was modified with amine-functionalized G-PEG. Second, the WE was modified with the conducting polymer PEI, which also contains NH2-functional groups (25, 26). G-PEG provided significant discrimination of the analytical signal at the low concentrations of SP analyzed (10−14-10−11 g mL−1) (
The fabrication, modification, and functionalization steps were then optimized to obtain a more robust and sensitive biosensor for SARS-CoV-2 SP detection. The WE was modified with G-PEG using the drop-casting method and incubating for 60 min at 37° C. to dry. This procedure introduces amine groups on the WE surface for bioconjugation. Next, the ACE2 receptor containing EDAC (1-ethyl-3-(−3-dimethylaminopropyl) carbodiimide)+NHS was dropped on the WE modified with G-PEG and kept for 30 min at 37° C. When the carboxyl groups of ACE2 were exposed to EDAC-NHS, they were activated to form a stable ester, which undergoes a nucleophilic addition with the amino groups present on the WE, resulting in a stable amide bond between the carbon WE/G-PEG and ACE2 (25, 28). The remaining unmodified sites of the WE surface were then blocked using a 1.0% (m/v) BSA solution.
Polymeric membranes can protect the electrode surface against biofouling when this surface is exposed to the sample's complex matrix, and can also provide superficial preconcentration of chemical species. For this study, analytical curves were made at concentrations ranging from 1×10−14 to 1×10−11 g mL−1 SP in 0.1 mol L−1 phosphate buffer solution (PBS) (pH=7.4) to compare 3 strategies of biosensor modification: (1) using 0.5% Nafion®; (2) using 0.5% chitosan; and (3) without any permeable membrane. The Nafion® layer resulted in the highest sensitivity of the biosensor (
Given the results presented in
An evaluation was performed concerning time of incubation of SP with the surface of the modified biosensor would yield the best analytical performance for SARS-CoV-2 detection. The experiment was carried out in triplicate with an interval concentration ranging from 10−14 to 10−11 g mL−1 of SARS-CoV-2 SP (
Materials. All reagents used in the experiments were of analytical grade. Deionized water (resistivity ≥18 MΩ cm at 25° C.) was obtained from a Milli-Q Advantage-0.10 purification system (Millipore). Human ACE2 Fc Chimera was obtained from GenScript. SP was kindly donated by Dr. Scott Hensley from the University of Pennsylvania. Graphene oxide conjugated with polyethylene glycol (G-PEG) amine-functionalized, N-(3-dimethylaminopropyl)-N-ethylcarbodiimide hydrochloride (EDAC), and N-Hydroxysuccinimide (NHS) with a degree of purity ≥98% and phosphate buffer saline solution, pH=7.4, were purchased from Sigma-Aldrich. Carbon and Ag/AgCl conductive inks and a dielectric ink were acquired from Creative Materials.
Fabrication of bacterial cellulose (BC) substrate. BC substrates were produced by G. hansenii (ATCC 53582), schematically illustrated in
The electrochemical devices were manufactured by the screen-printing method with 3-electrode configuration cells (dimensions: 2.5×2.0 cm) on the biodegradable BC substrate. Carbon conductive ink was used to fabricate the WE and counter electrode (CE), and Ag/AgCl conductive ink was used to fabricate the reference electrode (RE). To cure the conductive tracks, the printed BC substrates were placed in a thermal oven at 70° C. for 30 minutes. After the curing step, the devices were cut into small pieces (2.5×2.0 cm). To delimit the electrode area, a non-conductive ink was used, and the devices were submitted to an additional curing step under the same conditions as described above.
Modification of BC electrodes. To prepare the electrochemical biosensor, 5.0 μL of 2 mg mL−1 G-PEG solution was dropped on the carbon WE and allowed to dry for 60 min at 37° C. Next, 5.0 μL of a mixture of 0.33 mg mL−1 ACE2 receptor containing 25 mmol L−1 EDAC and 50 mmol L−1 of NHS solution were drop-casted on the surface of the WE and incubated at 37° C. for 60 min. The unmodified zones of the WE were blocked with 5.0 μL of 1% (m/v) BSA solution, and the WE was stored for 30 min at 37° C. to dry, to avoid non-specific interactions of other biomolecules present in the sample with the biosensor's surface. Finally, 5.0 μL of 1.0% (m/v) Nafion® was deposited onto the WE, and the WE was incubated at 37° C. for 60 min. The biosensor was then washed with PBS 0.1 mol L−1 (pH=7.4) before use.
To evaluate the electrochemical behavior of the sensor modification, cyclic voltammetry (CV) and electrochemical impedance spectroscopy (EIS) were used to record measurements of each functionalization step of the biosensor (
The bare carbon screen-printed electrode on the biodegradable BC substrate presented a defined redox process with peak currents (ip) of 148.5 μA and resistance to charge transfer (RCT) of 40.2 S2 (
Analytical performance of the biosensor. Highly specific interactions between the SP and the ACE2-coated electrode induce a potential variation. In these interactions, the output voltage is logarithmically correlated to the concentration of the target species in the solution, similarly to traditional ion-selective electrodes (ISE). Thus, when the analyte (SP) is present in the analyzed sample, the binding of the analyte to the functional membrane receptor (ACE2) produces an excess of surface charge on the electrode surface. Consequently, a potential change develops at the electrode, which can be used for diagnostic purposes. The presently disclosed reagentless electroanalytical method is based on these interactions.
An evaluation was performed regarding the electrochemical signal (potential difference) provided by the inventive potentiometric biodegradable BC-based biosensor through dose-response curves with low concentrations of virus.
The measurements were recorded by dropping 10 μL of SARS-CoV-2 SP or clinical samples onto the surface of the biosensor and incubating it for 7 minutes before each measure. The biosensor required at least 30 s to provide a stable potential difference response in the presence of SARS-CoV-2 SP to stabilize the accumulated charge (
ΔE=Esample−Eblank (Eq.1)
where Esample is the potential measured in the presence of SARS-CoV-2 and Eblank is the potential obtained for the blank, i.e., 0.1 mol L−1 PBS at pH=7.4. The electrical potential was sampled at 3 minutes for quantitative purposes to ensure a stable response. The signal for ΔE increased with the increase in the concentration of SARS-CoV-2 SP over the concentration range studied of 10.0 zg mL−1 to 1.0 μg mL−1 in 0.1 mol L−1 PBS at pH=7.4 (
Next, titered samples with B.1 SARS-CoV-2 concentrations ranging from 1×10−1 copies μL−1 to 1×105 copies μL−1 were analyzed (
The limits of detection (LOD) and quantification (LOQ) of the electrochemical device were calculated based on the four-parameter logistic (4PL) method, which is commonly employed for bioassays that use binding interactions. Applying equations 2 and 3, we obtained an LOD of 4.26×10−18 g mL−1 and an LOQ of 1.42×10−17 g mL−1 for SARS-CoV-2 SP (
L
C=μblank+t(1−α,n−1)σblank (Eq. 2))
where LC is a value of blank limit, μblank is the mean of signal intensities for n blank (negative control) replicates, σblank is the standard deviation of blank replicates, and t(1−α, n−1) is the 1—α percentile of the t-distribution given n−1 degrees of freedom, α=β=0.05 significance levels.
LOD=L
C
+t[1−β,m(n−1)]σtest (Eq. 3)
where Ld is the LOD in the signal domain, σtest is the pooled standard deviation of n test replicates and t[1−β, m(n−1)] is the 1−β percentile of the t-distribution given m(n−1) degrees of freedom. Again, the evaluation was set to σ=β=0.05, but these significance levels can be chosen properly for each study.
A comparison of sensing methods for SARS-CoV-2 was performed, and the inventive device provided the lowest LOD for SARS-CoV-2 SP solution (LOD=4.26×10−18 g mL−1) and gave results in a short time. The testing time was set as 10 minutes, which included 7 minutes for incubation of the sample and 3 minutes for potentiometric analysis (ΔE sampled at 3 min).
Cross-reactivity, reproducibility, and potential stability assays. To investigate the specificity of the instant biosensing electrochemical device for SARS-CoV-2, it was applied to other viruses and viral antigens under the same optimized experimental conditions. In addition to SARS-CoV-2, we tested four other viruses (H1N1, Influenza—A/California/2009; Influenza B—B/Colorado; MHV—murine hepatitis virus; and HSV2—herpes simplex virus-2) and three antigenic preparations (corresponding to heat-inactivated Zika virus and yellow fever, and gamma-irradiated Ebola virus). All experiments were carried out in triplicate in 0.1 mol L−1 PBS at pH=7.4 for 300 seconds of analysis. 10 μL of each virus or viral antigen was incubated on the biosensor surface for 7 minutes before the potentiometric measurements were taken (
Reproducibility assays were carried out to ensure that different test batches of SARS-CoV-2 performed similarly. For this study, potentiometric measurements were recorded of 1×101 copies μL−1 of SARS-CoV-2 prepared in a virus transportation medium (VTM) over 7 minutes of incubation time. The relative standard deviation (RSD) obtained with 10 biosensors representing different fabrication batches was 3.78%, indicating that the present fabrication method and functionalization protocol were highly reproducible (
The stability of the biosensor was evaluated potentiometrically using 0.1 mol L−1 PBS (pH=7.4) and VTM for 60 minutes (
Detection of SARS-CoV-2 in clinical samples. Using the optimized experimental conditions, the present biodegradable electrochemical biosensor was applied to the analysis of 15 OP/NP clinical samples, 5 of which had the original SARS-CoV-2 strain and 10 of which had the SARS-CoV-2 delta variant.
Viral loads in the samples ranged widely, with cycle threshold (Ct) values varying from 14.0 to 27.3 (
The biosensor detected SARS-CoV-2 samples and delta variants in all 15 clinical samples analyzed. The ACE2-based biosensor provided a higher analytical response, i.e., increased potential difference, for the SARS-CoV-2 delta variant samples compared to original strain samples with similar Ct values, which may be associated with the higher affinity of their mutated SP with the ACE2 receptor.
In order to evaluate the efficacy and robustness of the biosensor for COVID-19 diagnosis, another set of 50 NP/OP clinical samples was tested, 25 of which were positive NP/OP samples containing 12 SARS-CoV-2 variants of different lineages and 25 of which were negative NP/OP clinical samples (Table 2) obtained, after heat-inactivation, from patients from the Hospital of the University of Pennsylvania (HUP).
All the lineages were confirmed by RT-PCR. The cut-off value of the inventive biosensor was set as ΔE>0.025 V as positive for SARS-CoV-2, and ΔE<0.025 V as negative (
New variants of SARS-CoV-2 are likely to continue to emerge in the months and years ahead, and that inexpensive sensors for the detection of this virus will be needed to gather data on outbreaks and to diagnose cases. Here, the robustness and accuracy of a BC-based biosensor were evaluated by analyzing 65 NP/OP clinical samples (40 positive NP/OP samples from 13 SARS-CoV-2 lineages and 25 negative NP/OP samples; Tables 1 and 2). The accuracy of detection of this range of samples suggests that the inventive device would not require additional adaptation to detect emerging SARS-CoV-2 variants, as long as the newly mutated virus interacted with ACE2 to enable its entry into human cells. Based on its outstanding analytical parameters (high selectivity, reproducibility, specificity, and accuracy), low cost, simplicity, and biodegradability, the present device is well suited for frequent testing at the point of need. Thus, the inventive devices may help to prevent outbreaks in countries where the SARS-CoV-2 vaccination rates are low but frequent testing is feasible and sanitary practices are adequate.
i. Electrochemical measurements. For electrochemical characterization of the electrodes in each step of modification, the CV technique was used in a potential window ranging from 0.7 to −0.3 V and with a scan rate of 50 mV s−1. EIS experiments were carried out at frequencies ranging from 1×105 Hz to 0.1 Hz using an amplitude of 10 mV, and under open circuit potential (OCP). The electrochemical studies were recorded using 0.1 mol L−1 KCl solution containing 5.0 mmol L−1 of the redox probe [Fe(CN)6]3−/4− solution. Potentiometric measurements were carried out in a time interval of 300 seconds using 0.1 mol L−1 PBS (pH=7.4). A MULTI AUTOLAB M101 potentiostat with six channels, controlled by the NOVA 2.1 software, was used for all the electrochemical measurements. Experiments were carried out at room temperature, 25±3° C.
ii. SARS-CoV-2 biosensing. For SARS-CoV-2 biosensing, 10.0 μL of 0.1 mol L−1 PBS (pH=7.4) or VTM containing either SP or SARS-CoV-2 samples was applied to the biosensor surface and the device was incubated at room temperature for 7 minutes. Following incubation, the electrochemical cell was gently washed with 0.1 mol L−1 PBS (pH=7.4) to remove the unbound virus and sample. Then, 200 μL, of the 0.1 mol L−1 PBS (pH=7.4) was used for potentiometric measurements and the potential value (E) was obtained. The calibration curves were obtained in the concentration range from 10.0 zg mL−1 to 1.0 μg mL−1 SARS-CoV-2 SP in 0.1 mol L−1 PBS (pH=7.4).
iii. Reproducibility, stability, and cross-reactivity studies. To carry out the reproducibility study, the potential response was obtained by exposing 6 electrodes (from different batches) to 1×101 copies μL−1 of SARS-CoV-2 prepared in VTM for 7 minutes. The stability of the electrode response was potentiometrically evaluated in both 0.1 mol L−1 PBS and VTM for 1 hour. Cross-reactivity studies were performed with the following viral strains, all at 105 PFU mL−1: H1N1, Influenza—A/California/2009; Influenza B—B/Colorado; MHV—murine hepatitis virus; HSV2—herpes simplex virus-2. Cross-reactivity studies were also performed with heat-inactivated antigenic preparations of Zika virus (viral genome copy number: 1.1×107 copies μL−1), yellow fever virus (viral genome copy number: 1.8×104 copies μL−1), and Ebola virus (viral genome copy number: 1.1×107 copies μL−1), obtained from BEI Resources®. All the experiments were carried out by combining the viral sample with 0.1 mol L−1 PBS for 300 seconds of analysis, and 10 μL, of each virus (or antigenic preparation) was incubated on the biosensor surface for 7 minutes before the potentiometric measurements were made.
iv. Clinical sample analysis. NP/OP swab patient samples were heat-inactivated prior to analysis. Of the 65 NP/OP samples analyzed in this study, 40 were positive and 25 were negative for SARS-CoV-2 when tested by the RT-PCR method. The 25 negative clinical samples were acquired from the Hospital of the University of Pennsylvania (IRB protocol 844145). The 40 positive SARS-CoV-2 samples containing 13 variants: B.1.350, B.1.340, B.1, B.1.291, B.1.369, B.1.240, B.1.243, B.1.311, B.1.1.304, B.1.1.317, B.1.2, B.1.1.7 (alpha variant), and B.1.617.2 (delta variant) were obtained from under IRB protocol 823392. WA cut-off value of potential response (ΔE) was set to higher than 25 mV to express a positive diagnostic result, in accordance with the analytical response obtained for the lowest detected concentration of SARS-CoV-2 (104 copies μL−1) in the dose-response curve (
An HSV biosensor that was functionalized with nectin-1 was prepared. Electrochemical impedance spectroscopy (EIS) was used for the transduction of biosensor response, i.e., the selective binding between the nectin-1 bioreceptor immobilized on the electrode surface and the gD2 glycoproteins from HSV-2. The binding between nectin-1 and gD2 changes the interfacial electron transfer kinetics between ferricyanide/ferrocyanide (i.e., the redox probe used) and the electrode. The altered kinetics, in turn, can be detected by monitoring the increase in resistance to charge transfer (RCT), indicating a positive diagnostic result for HSV-2 infection (
All data from the optimization studies and analytical curves were plotted using the normalized RCT response, as defined by the following equation:
where Z is the RCT value obtained after incubating the electrode surface with gD2 or HSV-2 samples, and Z0 is the RCT value of the analytical blank solution [i.e., PBS or Dulbecco's Modified Eagle Medium (DMEM) with 5% fetal bovine serum (FBS)]. The normalization process of RCT corrects variation in the sensor response, which may be caused by analyst operation and temperature fluctuations when testing. Thus, normalization facilitates the eventual use of the sensor at decentralized testing sites.
The electrochemical sensors (3-electrode configuration) were manufactured by a screen-printing technique on phenolic paper circuit board material, as a low-cost and convenient platform. Electrically conductive carbon and Ag/AgCl inks (Creative Materials, USA) were employed to construct the working (WE)/auxiliary (ΔE) and reference (RE) electrodes, respectively. After a curing step of 30 min at 100° C., the material was cut into 2.5×2.0 cm pieces, and their geometrical area was delimited using dielectric tape.
Initially, to generate a robust and sensitive biosensor, two strategies were evaluated to modify the working electrode (WE) and enable the anchoring of the nectin-1 bioreceptor. In the first approach, the WE surface was coated with glutaraldehyde (GA), a dialdehyde used to anchor biomolecules through their N-terminal groups; for the second approach, the WE was modified with PEI, a conductive polymer containing amino functional groups enabling the attachment of biomolecules through their carboxylic acid and ester groups.
Using GA as a modifier did not provide significant discrimination of the analytical signal (RCT) at the concentrations of gD2 tested (10−12-10−9 g mL−1,
The main fabrication, modification, and functionalization steps of the biosensor using PEI was investigated. First, the working electrode was modified with 4.0 μL of 1.0 mg mL−1 PEI solution, by drop-casting, and incubated for 60 min at 37° C. This procedure generates —NH functional groups on the carbon electrode surface. Then, 4.6 μL of 0.13 mg mL−1 of the nectin-1 receptor, containing a mixture of 25.0 mmol L−1 EDC+50.0 mmol L−1 NHS, was deposited on the surface of the PEI-modified WE, and the biosensor was incubated for 30 min at 37° C. The carboxyl groups on nectin-1, when exposed to EDC-NHS, are activated to form a stable ester, which undergoes a nucleophilic addition with the amino groups on the PEI-modified WE, such that a stable amide bond is formed between the PEI-modified carbon electrode and nectin-1. Subsequently, the remaining unmodified sites of the electrode surface were blocked with 4.0 μL of a 1.0% (m/v) BSA solution. In the last step, 4.0 μL of 0.5% (m/v) chitosan was dropped on the surface of the nectin-1-modified WE.
After selecting PEI as the immobilization strategy for nectin-1, the use of two types of permeable membranes, namely Nafion® and chitosan, was investigated. Analytical curves ranging from 1×10−12 to 1×10-10−10 g mL−1 of gD2 in 0.1 mol L−1 of PBS (pH=7.4) were constructed. Experiments were performed in triplicate to compare 3 strategies: i) without a permeable membrane, ii) with 0.5% Nafion, and iii) with 0.5% chitosan (
Subsequently, the optimal incubation time of either gD2 or viral samples with the surface of the biosensor was investigated to obtain a compromise between analytical frequency and sensitivity for HSV-2 detection. The evaluation was based on the analytical sensitivity (slope) parameter obtained by analytical curves, determined in triplicate, at concentrations of gD2 ranging from 10−12 to 10−10 g mL−1 (
Electrochemical Characterization. For each functionalization step (
All electrochemical measurements were carried out using a mixture of 5.0 mmol L−1 [Fe(CN)6]3− and [Fe(CN)6]4−, as a redox probe, in 0.1 mol L−1 KCl solution. All functionalization steps of the biosensor were characterized by electrochemical Impedance Spectroscopy (EIS), which was also used to quantify the HSV-2 and gD2 concentrations. The frequencies used ranged from 1×105 Hz to 0.1 Hz, and the open circuit potential was applied with an amplitude of 10 mV (vs. Ag/AgCl). For cyclic voltammetry (CV) experiments, the potential ranged from −0.3 to 0.7 V (vs. Ag/AgCl) using a scan rate of 50 mV s−1.
Analytical Performance. EIS was used to quantify free gD2 and HSV-2 virus in 0.1 mol L−1 PBS (pH=7.4). Dose-response curves were built with the previously described experimental conditions (i.e., 1 mg mL−1 PEI, 0.5% chitosan, and 5 minutes of incubation time), and the analytical results were normalized according to Eq. 1.
Collectively, these experiments highlight the excellent sensitivity displayed by the inventive biosensor, which should provide an early diagnosis of HSV-2 infection in human clinical samples. Another advantage for diagnostic purposes is the short testing time, i.e., 9 minutes, consisting of a 5-minute incubation of the sample on the electrode surface and an additional 4 minutes for the EIS measurements of both the analytical blank and the sample of interest.
In comparison to other approaches reported in the literature, the present disclosure provides the first approach that uses a moiety for detecting the viral glycoprotein gD2 instead of genosensor technology using genetic material for the recognition of HSV. In addition, the presently disclosed HSV sensors presents the fastest testing time, with a very low LOD and a large interval concentration range to detect HSV-2. Furthermore, the devices can be produced inexpensively. Considering the cost of nectin-1 ($800/mg), for example, the final cost to assemble each HSV biosensor was exactly $1.00: $0.12 for electrode fabrication+$0.40 for all the chemicals used in the functionalization step (PEI+EDC+NHS+BSA+Chitosan)+$0.48 for nectin-1. Because the present biosensors are low-cost, their production is potentially highly scalable.
The effect on the biosensor's electrochemical response of adjusting the pH of the medium to a pH that is close to physiological conditions was also studied (
To verify the reproducibility of the proposed method, i.e., to assess whether different batches of biosensors performed similarly, 6 biosensors from different fabrication rounds were evaluated using the same optimized protocol. Briefly, the RCT measures were recorded by EIS using 5 mmol L−1 [Fe(CN)6]−3/−4 after incubating the biosensor with 1×10−9 g mL−1 of gD2 prepared in 0.1 mol L−1 of PBS (pH=7.4) (
The stability of the electrochemical biosensor, stored in sealed Petri plates at various temperatures (−20° C., 4° C., and 25° C.), was evaluated over 7 days. Analytical curves were built at concentrations ranging from 1×10−12 g mL−1 to 1×10−9 g mL−1 gD2 in 0.1 mol L−1 PBS, pH 7.4 (
The ability of the biosensor to detect HSV-2 in pre-clinical samples was assessed. Tested blindly, in triplicate (n=3), were 9 HSV-2 positive and 11 negative biofluid samples collected from the vagina of guinea pigs (
The biosensors achieved 88.9% sensitivity, 100% specificity, and 95% accuracy for the set of 20 samples evaluated, i.e., the biosensors correctly diagnosed 19/20 samples tested. There was a response variation between the proposed method and the titrated method for sample analyses (
Cross-reactivity experiments were performed to rule out any potential off-target effects of the nectin-1-modified electrode with viruses other than HSV. Selectivity was studied for 5 viruses: H1N1 (A/California/2009), Influenza-B/Colorado, H3N2, MHV-murine hepatitis virus, and SARS-CoV-2. All experiments were performed using the same optimized conditions as those used for HSV-2 detection. No significant cross-reactivity was detected with any of the viruses tested, as revealed by a relative RCT percentage of up to 12%, which is lower than the cut-off value of 22% established for a positive diagnosis of HSV-2 infection in biofluid samples (
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.
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 105 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.
Inventors 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) (
As illustrated in
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.
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.
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) (
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 (
The optimized protocol generated the best analytical signal for the detection of SARS-CoV-2 in human biofluid samples (
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 (
The key steps required for the electrode's functionalization were optimized and characterized (
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 (
Taking into account the optimal analytical conditions evaluated (Table 1′), we built calibration curves for free SP (
where Z is the RCT of the sample and Z0 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.
The dose-response curve for the free SP in PBS solution ranged from 1 fg mL−1 to 10 μg mL1 (
The RCT values of Nyquist plots were extracted by the application of an equivalent circuit (
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 101 to 106 PFU mL−1. 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 μL−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 RealStar® 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% (
Next, we evaluated the stability of DETECT at different temperature storage conditions (25° C., 8° C., and −20° C.) over 10 days (
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 (
We also observed that centrifuging the samples did not lead to increased impedimetric detection of the SP (
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 COVID-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 RNA copies μL−1). It is worth noting that although the heat inactivation protocol decreased the response of our biosensor due to the inactivation of SP (
DETECT demonstrated high sensitivity, specificity and accuracy (96.2%, 100% and 97.4%, respectively; Table 4′).
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.
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.
Several key analytical features were used to compare the performance of DETECT with respect to other electrochemical methods reported in the literature (Table 6′).
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.,
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.
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 (i.e., 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 (
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 antibodies17. Our device presented higher accuracy, specificity, and selectivity than most existing methods available for SARS-CoV-2 detection11. 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 tests10,11,18, and could potentially be lowered even more by using engineered versions of human ACE2 with enhanced selective binding towards SARS-CoV-2 SP′. The use of such ACE2 variants would also help reduce the rate of false positives in complex biofluids such as saliva7,19,20.
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′).
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).
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.
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 μL 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 μL at 0.32 mg mL−1), then incubated at 37° C. for 1.5 hours. Next, 4 μL 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.
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 (
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 (
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),
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.
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.
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 (KCl), potassium ferricyanide K3[Fe(CN)6], potassium ferrocyanide K4[Fe(CN)6], 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).
Fabrication of electrochemical devices. The electrochemical sensors were screen-printed in a three-electrode configuration cell (dimensions: 1.8×1.2 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 (1.1×1.7 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 screen-printing 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 (2×2 cm) and a non-conductive layer was applied to delimit the electrode area.
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 μL of GA solution (25% in water) was added on the surface of the working electrode using the drop-casting method. After 1 hour, 44 of ACE2 solution (0.32 μg 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, 44 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, 44 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.
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 KCl 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 μL of a mixture of 5 mmol L−1 ferricyanide/ferrocyanide prepared in 0.1 mol L−1 KCl solution added after the sample incubation on the electrode (104 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 105 and 10-1 s−1 using a typical open circuit potential of 0.15 V and an amplitude of 10 mV at room temperature.
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 μg mL−1) of angiotensin II (AngII), 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 electronic surface) when compared to NP/OP swab samples.
Cross-reactivity experiments. Cross-reactivity assays were carried out by exposing the sensor to three coronaviruses (MHV—murine hepatitis virus at 108 PFU mL−1, HCoV-OC43—human coronavirus OC43 at 104 PFU mL−1, and human coronavirus 229E at 107 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 105 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 μL of virus sample, and EIS measurements as specified above (Electrochemical Measurements section).
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.
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.
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) (
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 M D T, 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 104M−1s−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 101 to 106 PFU mL−1.
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 COVID-19 Holds Promise to Combat Current and Future Pandemics. ACS Appl Bio Mater 3(11):7326-7343).
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 V M, 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 (Boum 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.
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, COVID-19 related symptoms and other medical conditions (Table 9′, below).
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%;
Additional details concerning the performance of the present cohort study are as follows.
The testing platform comprised two components: the electrochemical sensor and a potentiostat. The electrochemical sensors were prepared following established protocols (Torres M D T, 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 (2×2 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 μL of 25% glutaraldehyde (GA) solution was added for 1 hour at 37° C. to allow the formation of a cross-linked polymer, which enabled subsequent anchoring of ACE2 (4 μL at 0.32 mg mL−1). ACE2 was then incubated at 37° C. for 1.5 hours. Next, 4 μL 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 K A, Moore R B (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 R F, et al. (2020) Simple and inexpensive electrochemical paper-based analytical device for sensitive detection of Pseudomonas aeruginosa. Sensors Actuators B Chem 308:127669).
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.
SquidStat Plus (Admiral Instruments) and MultiAutolab M101 (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 KCl 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 105 to 10−1 Hz using a sinusoidal signal with 10 mV of amplitude at room temperature, respectively.
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 charge-transfer 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 2004 of a mixture of 5 mmol L−1 ferricyanide/ferrocyanide prepared in a 0.1 mol L−1 KCl solution added after incubating the clinical sample (104 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 105 and 10-1 s−1 using a typical open circuit potential of 0.15 V and an amplitude of 10 mV.
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 (i.e., mixture of 5 mmol L−1 ferricyanide/ferrocyanide prepared in 0.1 mol L−1 KCl 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).
For the RT-PCR assays, RNA was extracted and purified using the QIAmp DSP Viral RNA Mini Kit (Qiagen) from a 140 μL 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 (ABI). 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.
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).
The RCT values of Nyquist plots obtained using Squidstat Plus (Admiral Instruments) and Multi Autolab M101 (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˜10Ω) (Uygun Z O, Ertu{hacek over (g)}rul 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).
To diagnose a given sample, the normalized RCT, defined by the following equation, was used:
where Z is the RCT of the sample and Z0 is the RCT of the blank solution (VTM).
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.
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 K K, et al. (2020). Science (80-) 369(6508):1261-1265).
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 M J, et al. (2020) N Engl J Med 383(22):e120).
The following publications may be relevant to the presently disclosed subject matter relating to SARS-CoV-2 biosensors described under Section I above:
The following publications may be relevant to the presently disclosed subject matter relating to HSV biosensors described under Section II above:
The following publications may be relevant to the presently disclosed subject matter relating to SARS-CoV-2 biosensors described under Section III above:
The present application is the U.S. national stage and is also a continuation-in-part of PCT/US2021/071789, filed Oct. 8, 2021, which claims priority to U.S. Provisional Application No. 63/089,905, filed Oct. 9, 2020, U.S. Provisional Application No. 63/134,690, filed Jan. 7, 2021, and U.S. Provisional Application No. 63/155,963, filed Mar. 3, 2021. The present application also claims priority to U.S. Provisional Application No. 63/489,494, filed Mar. 10, 2023. The entire contents of each of the above-cited patent applications are incorporated herein by reference.
Number | Date | Country | |
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63089905 | Oct 2020 | US | |
63134690 | Jan 2021 | US | |
63155963 | Mar 2021 | US | |
63489494 | Mar 2023 | US |
Number | Date | Country | |
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Parent | PCT/US2021/071789 | Oct 2021 | US |
Child | 18296613 | US |