Patent Application
20250035581
Embodiments of the present invention relate to a biosensor, specifically a field-effect transistor biosensor having multiple functionalized chambers.
Within the recent COVID-19 outbreak, a lack high quality point-of-care systems that can detect specific biomolecules, such as hostile viruses or bacteria, became evident. The time between hosting the virus and developing the systems, getting tested through established methods, is long, and often crucial. Moreover, often the symptoms of different diseases of viruses (e.g. COVID-19 and Influenza) are similar, making hospital-style testing even more complicated. Therefore, there is a need for a portable biosensor system that is (a) rapid, (b) specific, (c) multi-targeted.
The pathology of upper respiratory viruses has regularly presented challenges to global health care systems and their resources. The emergence of new virus variants that can evade communal immunological memory can be rapidly transmitted through airborne mucosal droplets, often resulting in the emergence of sudden seasonal epidemics or pandemics. Over the last century, the most prominent of these viruses have been variants of influenza (Flu), which have been estimated to be responsible for approximately 400,000 deaths annually1. The emergence of the novel coronavirus SARS-COV-2 (COVID-19) in 2019 introduced a new upper respiratory virus that, as of now (September 2022) has led to at least 6.3 million deaths globally.
The COVID-19 pandemic has highlighted the need for new rapid point of care diagnostic systems for upper respiratory viruses, especially for high population density areas where the transmission can be the most potent and diagnostic availability and turnaround time the most limited. Significant challenges in respiratory diagnostics include the establishment of assays with a limit of detection (LoD) suitable for identifying early infections, minimizing false positive rates, and reducing the time to perform the assay. The current standard, the reverse transcription polymerase chain reaction (RT-PCR) isn't ideal for identifying early respiratory infections, as demonstrated by the United States' Center of Disease Control's (CDC) recommendation that these assays should be performed 5 days after an exposure to ensure maximal viral titera. Additionally, RT-PCR assays typically take a few hours to perform and often require transporting samples to professional laboratories, which can take a few additional days thus being a challenge during periods of high demand.
COVID-19 and Flu exhibit similar physiological symptoms underscoring the requirement for a rapid diagnostic tool capable of differentially diagnosing COVID-19 and Flu. An initial assessment of the potential cause of illness would allow a timely personalized treatment plan for the patient, thus not only aiding in curbing the spread, but also in utilizing medical resources in an efficient manner. As the recent COVID-19 pandemic spurred the rapid development of multiple COVID-19 detection platforms with varying degrees of usability and success, antibody-modified graphene field effect transistors (GFETs) have stood out due to their low LoDs and fast response time. Imbibing these GFETs with concurrent multiple target detection capability would increase their effectiveness not only during pandemics but also in instances where there is an urgent requirement to detect the cause of illness in a patient showing overlapping symptoms with another disease.
Accordingly, the present invention is directed to a multi-targeted, modular virus sensing platform based on graphene field-effect sensing that obviates one or more of the problems due to limitations and disadvantages of the related art.
In accordance with the purpose(s) of this invention, as embodied and broadly described herein, this invention, in one aspect, relates to a multi-target specific (e.g., COVID-19 and Influenza) biosensor technology utilizing field-effect transistors (FETs). For example, the FET may be a graphene or other 2D material based FET. The target-specific antibodies or aptamers (whenever available) will be used to add specificity to graphene channels. The technology will enable an early-stage diagnosis tool that differentiates between multiple analytes of choice, and to differentiate between Influenza and COVID-19, yet embodied into a single device for point-of-care monitoring.
In another aspect, the disclosure relates to a biosensor that includes a multiplexed array of electrolyte-gated 2D material based field effect transistors (FET); separating walls on an upper surface of the multiplex array, the walls fluidically separating each of the FETs from one another; and a perimeter wall on the upper surface of the multiplex array surrounding the separating walls and traversing the surface of all FETS in the multiplexed array, the perimeter all forming a fluidically tight well traversing multiple ones of the FETs, the fluidically sealed well subdivided into a plurality of chambers, each chamber corresponding to one of the FETs in the microarray. Each of the FETs includes monolayer of graphene as a conducting channel, wherein each conducting channel functionalized according to a predetermined agent.
Additional advantages of the invention will be set forth in part in the description which follows, and in part will be obvious from the description, or may be learned by practice of the invention. The advantages of the invention will be realized and attained by means of the elements and combinations particularly pointed out in the appended claims. It is to be understood that both the foregoing general description and the following detailed description are exemplary and explanatory only and are not restrictive of the invention, as claimed.
An advantage of the present invention is to provide a rapid response (e.g., within minutes) early-stage diagnosis tool that differentiates between multiple analytes of choice embodied into a single device, which allows for point-of-care monitoring.
Further embodiments, features, and advantages of the multi-targeted, modular virus sensing platform, as well as the structure and operation of the various embodiments of the multi-targeted, modular virus sensing platform, are described in detail below with reference to the accompanying drawings.
It is to be understood that both the foregoing general description and the following detailed description are exemplary and explanatory only and are not restrictive of the invention as claimed.
The accompanying figures, which are incorporated herein and form part of the specification, illustrate embodiments of the multi-targeted, modular virus sensing platform. Together with the description, the figures further serve to explain the principles of the multi-targeted, modular virus sensing platform described herein and thereby enable a person skilled in the pertinent art to make and use the multi-targeted, modular virus sensing platform.
Reference will now be made in detail to embodiments of the multi-targeted, modular virus sensing platform with reference to the accompanying figures. The same reference numbers in different drawings may identify the same or similar elements.
It will be apparent to those skilled in the art that various modifications and variations can be made in the present invention without departing from the spirit or scope of the invention. Thus, it is intended that the present invention cover the modifications and variations of this invention provided they come within the scope of the appended claims and their equivalents.
Throughout this application, various publications may have been referenced. The disclosures of these publications in their entireties are hereby incorporated by reference into this application in order to more fully describe the state of the art to which this invention pertains.
Described herein is a multi-target specific (e.g. COVID-19 and Influenza) biosensor technology utilizing graphene-based field-effect transistors (GFETs). The target-specific antibodies or aptamers (whenever available) may be used to add specificity to graphene channels. The technology yields a system that one can use at home, without doctor's or clinician's involvement. For example, a test device according to principles described herein provides a rapid response (e.g., within minutes) and enables an early-stage diagnosis tool that differentiates between multiple analytes of choice, and specifically to differentiate between Influenza and COVID-19, yet embodied into a single device, which allows for point-of-care monitoring.
According to principles described herein an exemplary device uses an FET biosensor having a monolayer graphene as the conducting channel. The channel is functionalized using anti-antibodies or aptamers to selectively capture a bioentity, antigen, or molecule of interest. For example, each individual graphene channels within a device may be functionalized with a different biomolecule. The biomolecules may be specific to bioentities that possess a potential threat to human organism, such as toxins, viruses (COVID-19 and Influenza, Ebola), and other proteins, DNAs, etc. or antigen.
One feature of the technology is multi-specificity and advanced target analyte selectivity. A chip comprising GFETs (or other 2D based FETs) can embody multiple target biomolecule specific counterpads. For example, antibodies or aptamers and provided into the graphene surface. The antibodies or aptamers can be localized within different chambers on the chip, the chambers corresponding to a different GFET channel, making multiple GFETs on a single chip sensitive to different biomolecules of interest. One non-limiting example is a GFET chip having dual biosensing specificity: for COVID-19 and for common Influenza viruses (such as, e.g. N1H1). One chamber may be functionalized as a control. By cross-correlating the response from the functionalized and the control sensors, the device will be able to quickly and decisively give its response. The GFETs may be electrolytic-gated GFETs.
For example, a concurrent rapid differential diagnosis platform using antibody-modified GFET can be fabricated according to principles described herein. The device is a holistic platform having 4 onboard GFETs isolated from each other using polydimethylsiloxane (PDMS) barriers yet enclosed in a higher perimeter PDMS wall so that they can be functionalized individually and tested using a shared biological sample without the assistance of complex microfluidics. Each GFET is modified with either an antibody of interest, i.e., COVID-19 or Flu or are used as a control. The device design enables isolated targeted functionalization of graphene channels while allowing a common medium for introducing the analyte, which then translates into common gating and a change in conductance of the GFET modified with the corresponding target/receptor12. In this case, the chip has two GFETs dedicated to antibody immobilization for COVID-19 and Flu each, while one GFET was only chemically passivated with Tween-20 (Tw20) and another left bare as a control.
Referring again to
In the example illustrated in
The example device illustrated in
In an aspect of the device, the device may include a single common electrode 110 to make four chambers (channels), which would use, for example, five leads. For a device with more chambers, the device would use a number of leads equal to the number of chambers (channels) plus one. For example, for a 30 chamber device (28 biosensors plus 2 control), 31 leads would be used. A number of source electrodes 112 may correspond to a number of channels.
Accordingly, an exemplary biosensor includes a multiplexed array of electrolyte-gated graphene based field effect transistors (GFET); separating walls on an upper surface of the multiplex array, the walls fluidically separating each of the GFETs from one another; and a perimeter wall on the upper surface of the multiplex array surrounding the separating walls and traversing the surface of all GFETS in the multiplexed array, the perimeter all forming a fluidically tight well traversing multiple ones of the GFETs, the fluidically sealed well subdivided into a plurality of chambers, each chamber corresponding to one of the GFETs in the microarray. Each of the GFETs includes monolayer of graphene as a conducting channel, wherein each conducting channel functionalized according to a predetermined agent. Examples of functionalizing agents may include DNA, aptamers, antibodies, viral particles (e.g., flu, covid, rotavirus, ebola, zika, etc.), small molecules, peptides, enzymes, bacteria (e.g., salmonella, e-coli, etc.), proteins, exosomes, toxins (e.g., ochratoxin, cholera, etc.), organic molecules (glucose, etc.) or the like.
While described here with a graphene based field effect transistor, other appropriate 2D materials, such as MoS2, hBN, WS2, WSe2, PtS2, PtSe2, could be used with an appropriate electrolyte. The electrolyte may be any one of or a combination of phosphate buffered saline, Hanks' Balanced Salt Solution, Tris, Saliva, nasal swab dissolved in DI water.
The perimeter wall and/or the separating walls may be made of any appropriate material, including a polymer such as polydimethylsilixane, EcoFlex, SU-8, PMMA, a combination thereof or the like. The perimeter wall and/or the separating walls may be SiO2/Si, sapphire, glass, SiC, and flexible polyimides, parylene, polyester, tape, polycarbonate, Teflon, PET, EVA/PET, etc-any substrate which survives chemical treatment with acetone or toluene a combination thereof or the like
The height of the separating walls may be less than height of the perimeter wall. The perimeter wall or the separating walls may be made of silicone, such as polydimethylsilixane. The predetermined agents may include flu antibody, COVID antibody, anti-ferritin antibody or other antigen-specific antibody. As discussed above, at least two of the GFETs can be functionalized differently.
According to the principles described herein, the technology includes:
The technological design allows one to (a) functionalize each graphene sub-channel individually, while (b) sample the target analyte with all channels simultaneously. Such design allows us to simplify the overall device fabrication and to bypass the necessity to build microfluidic chambers that would significantly complicate the overall technology and its at-home employment.
Strategy to functionalize the individual graphene channels with individual recognition elements at the same time, without cross-binding. At the same time, all individual channels within the array can be sampled in parallel, and when the target analyte is added into the chamber, it reacts with all individual graphene channels simultaneously.
The technology intends to solve the problem of point-of-care testing in ambulatory, home environments. The final device can be simplified down to a chip that could be ready via a mobile read-out system, and used at home or any place of convenience, utilizing such easy to access bioanalytes as saliva instead of blood.
The technology has the following advantages over current technologies:
Technologically, field-effect biosensing may be restricted with the Debye screening length, which is correlated to the molarity of the solution. The human saline has very high molarity, and if used directly, the Debye layer would be within 1 nm, restricting the ability to electrostatically detect large molecules. This limitation can be overcome by diluting the human saline in deionized water to yield lower molarity and lower concentration bio-analyte.
The technology can be used to detect perhaps any other viral bodies (e.g. ebola or others yet to come) or for detection of hormones, DNAs, proteins, etc.
Biosensing according to principles described herein begins with functionalization of the chambers/channels of the GFETs. Referring to
Each device consists of an array of 4 GFETs presenting 4 channels of operation (C-n), isolated from each other through PDMS enclosures (
To distinguish between viruses, we selected antibodies that recognize a unique antigen for each virus. For COVID-19, we used the antibody CR3022 to target the receptor-binding domain (RBD) region on the transmembrane Spike protein. For Flu, we selected the engineered antibody FI6v3 to bind to the conserved central stalk domain of transmembrane protein hemagglutinin (HA). The antibodies selected are each capable of binding to multiple variants of their respective virus. For COVID-19, the virus variants haven't mutated to the significance to elicit complete binding escape from antibodies targeting the original virus; thus, most COVID-19 antibodies such as CR3022 are capable of binding to recent variants such as omicron and delta. Additionally, FI6v3 was engineered to bind to all type 1 and 2 influenza A subtypes13. The diversity of variants that can be recognized gives this assay tremendous breadth among different subtypes of each virus. The interaction between the antibodies and their respective analyte proteins was validated through ELISA for each batch of antibodies.
The electric double layer (EDL) formed at the graphene electrolyte interface serves as a dielectric layer. The common electrolyte enabling the operation of the GFETs is a low ionic strength PBS set at 0.01×. The decision to employ PBS 0.01× was to counter the charge screening15-16 effect observed in high ionic concentration solutions, which reduces the observed signal strength17 resulting from the interaction of the target and analyte. It is imperative that EDL fall at the range suitable for IgG antibody interactions, around 4 to 14.5 nm18 as opposed to the low 0.7 nm above the surface EDL formed by PBS 1× 15. Through our experimentation, it was observed that PBS 0.01× served as the best concentration for signal detection while also maintaining bio-molecular integrity as observed through enzyme-linked immunoassay (ELISA).
To allow targeted detection, the GFET channels were modified through biochemical functionalization, starting with making CVD-grown graphene suitable for antibody immobilization. The lack of reactive sites or dangling bonds on CVD graphene offered no site for target immobilization, which was resolved through incubation of 1-pyrenebutanoic acid succinimidyl ester (PBASE) on the surface of graphene. PBASE is a pyrene-based succinimide ester that utilizes the π-π bonds extending out at the surface of graphene. The successful immobilization of PBASE on graphene was confirmed through Raman spectroscopy and electrical characterization.
The N-hydroxy succinimide (NHS) ester group in PBASE reacts with primary amine groups of the proteins, thus allowing antibody immobilization26. The PDMS enclosure allowed specific immobilization of the CR3022 and FI6v3 onto separate GFETs on the device. To ensure that the area of graphene that remained unoccupied by PBASE and the antibodies did not lead to any non-specific reaction, PEG-NH2 was introduced as the blocking reagent27. PEG-NH2 also plays an essential role in combatting the screening effect introduced by the electrolyte, since it increases the Debye length, thus making the EDL comparable to the dimensions of the antibodies28-29. To neutralize PBASE sites unoccupied by antibodies, ETA was used as the blocking agent to prevent any non-specific reaction initiated through the amine groups of analytes being tested. To ensure that the results observed are due to antibody-antigen interaction, rather than electronic drift or fluctuations, we deployed the third GFET as the comparative electronic control. The third graphene channel in this GFET was modified with Tw20 only, to serve as a blocking layer, with the expectation that it would not respond to introduction of any analyte into the solution. Each step of functionalization was characterized electrically and optically with all devices assembled, showing a consistent trend indicating successful immobilization and blocking.
To evaluate the sensing capability of the device, we performed a series of time trace measurements where all onboard transistors were exposed to varying concentrations of both COVID-19 S-protein (Spike) and Flu Hemagglutinin (HA) proteins at different intervals as outlined in the measurement protocol.
The antigen-antibody interaction utilizes the uniform turbulent diffusion of viral proteins delivered in low ionic strength PBS, entailing a facile operating procedure, where the user simply pipettes a drop of the viral protein solution onto the device and observes a response within seconds. Prior to testing the device against the target proteins, a negative control protein test was conducted with bovine serum albumin (BSA) as the analyte to verify its specificity. We established a precise dual detection of the two viruses without cross-reactivity of the signals; hence each time the characterized devices were exposed to control proteins to study cross-reactivity and specificity.
For all the time-resolved trace measurements, the gate voltage was set to the value which exhibited the highest transconductance (Vgmax) for the chip in PBS 0.01× post functionalization (
The derivative of the time series curve eliminates the impact of drift and other electronic artifacts observed in the real-time traces, as shown in
As observed in
The Kd value obtained through the Hill fitted (Eq. 1) data points is 0.147 nM, and the Hill coefficient (n) stands at 0.45. The Hill coefficient below 1 indicates that the interaction between the antigens and the antibodies follows negative cooperative binding33. This implies that the first instance of interaction between the antigen and the antibodies is the strongest, while the reaction at successively increasing concentrations is likely blocked by the presence of viral surface proteins already interacting with antibodies near the surface, leading to a diminished signal response.
When analyzing the device performance, we observed overall sensitivity of the devices is very high, above other emergent technologies. Sensitivity was calculated by performing a linear fit on the linear range of the (I/I0) % vs log (M) curve, achieving 2.4% change in signal per log (molar) concentration for COVID-19 (2.4%/log (M)) and 1.9% change in signal per log (molar) concentration (1.9%/log (M)) of Flu (
Our device's high sensitivity and low experimental LoD can be attributed to the deployment of low strength ionic buffer and PEG-NH2 in functionalization to combat the screening effect caused by short Debye length in high ionic strength buffers. Aiding the specific functionalization scheme is also the selection of the most sensitive Vgs corresponding to a high transconductance value. By virtue of the linear relationship (Eq. 2) between transconductance and W/L ratio, the high W/L ratio of 8.75 in the device architecture enables higher transconductance, imparting higher sensitivity in turn translating to ultra-low LoD.
Our device standing at 88 zM is already approaching breath sample detection levels (118.2 zM) while already surpassing the minimum LoD requirements for nasal (163 fM) and saliva sample (16.3 aM). Such low LoD, as exhibited by our device, allows versatility in selecting the type of sample and can potentially reduce the time for administering the test after exposure.
Owing to their molecular weights, theoretically, the lowest possible concentration with Spike and HA protein is ˜1.67 zM. Our device's lowest measured concentrations indicate the capability of almost approaching single molecule detection for each viral protein in their respective GFETs with essentially an immediate turnaround time.
High binding 96 well plates [Costar cat 07-200-721] were coated at 2 ug/mL with S protein or HA overnight at 4° C. Plates were washed three times with PBS 1× with 0.05% TW-20 (PBST) and were blocked with PBS 1×, 2% skim milk for 2 hours at room temperature. Antibodies in (1× or 0.01×) PBS, 0.05% TW-20, and 1% skim milk (PBSMT) were serially diluted across the 96 well plate before a 1-hour incubation. Goat Anti-Human-IgG with horseradish peroxidase (HRP) (Sigma-Aldrich™ cat A0293) diluted 1:5000 in PBSMT 1× was used as a secondary antibody and incubated for 30 minutes. 1-Step™ Ultra TMB-ELISA Substrate (Thermo Scientific™ cat 34029) was used to develop the plates and the reaction was quenched with 2M H2S04. Absorbance values were measured at 450 nM on a Synergy H1 microplate reader (BioTek™).
Gblocks ordered from Integrated DNA Technologies (IDT) containing antibody variable heavy or light chains were inserted into mammalian expression vector pcDNA3.4 by Golden Gate cloning and validated with sanger sequencing. Antibodies were expressed using the Expi-293™ Expression System (Thermo Scientific™ cat A14635) and purified with Pierce™ Protein G Plus Agarose (Thermo Scientific™ cat 22851). A stabilized version of the S protein, Hexapro was expressed using the Expi-293 expression system and purified using Ni-NTA agarose (Qiagen cat 30210). All proteins produced in house were validated on SDS-PAGE gels and quantified using the Pierce™ Coomassie Plus (Bradford) Assay Kit (Thermo Scientific™ cat 23236). Proteins purchased commercially included the HA strain H3N2 A/Singapore/INFIMH-16-0019/2016 (Native Antigen) and powdered BSA (Thermo Scientific cat BP9706100).
Photolithography and lift off techniques were employed to deposit gold on Si/SiO2 wafer as three terminals to create a 4-GFET array structure of the device. Cr/Au (10 nm/90 nm) layers were deposited through e-beam deposition and lift off techniques. Wet transfer method was utilized to transfer graphene onto the substrate.
Commercially obtained graphene sheet grown on copper (Grolltex) was spin-coated with Poly (methyl methacrylate) (PMMA) (PMMA 950 A4, MicroChem). After spin coating, the PMMA/graphene/Copper stack was baked at 150° C. for 10 minutes. The PMMA/graphene/Copper stack was upturned with the Cu side exposed and was subjected to Oxygen plasma for 30 sec at 30% flow rate. The copper sheet with PMMA/graphene film was then cut into 10 mm×10 mm pieces and placed into Ammonium Persulphate, (NH4)2S2O8, for 24 hours to dissolve the copper. Pieces were placed with PMMA side facing upwards to allow the copper to dissolve. PMMA/graphene film pieces were rinsed and allowed to soak in deionized (DI) water for a total of three consecutive times and then transferred to the silicon wafer with a gold deposit. PMMA/graphene transferred wafers were left to air dry overnight and then baked at 150° C. for 10 minutes. Wafers were then placed in an acetone bath for 24 hours to dissolve the PMMA layer. Bare graphene wafers were rinsed in ethanol and DI water and then dried with the air gun. Dried wafers were baked at 150° C. for 10 minutes. PDMS enclosures were made by cutting rectangular pieces of PDMS and using liquid PDMS to hold them together. The outer PDMS boundary was made with a taller height than the inside cross enclosure to allow overflow between channels on the top (during measurements) of the inside but to prevent leakage to the outside. Inter-leaking between channels was tested using isopropyl alcohol. Small lengths of copper wires were stripped at both ends and connected to the common source, the drain, and the ground through contact with the gold layer on the device and the use of silver epoxy (MG Chemicals 8331S Silver Epoxy Adhesive) to make sure the wires stayed attached to the device.
10 mM PBASE (Anaspec, AS-81238) solution in Dimethylformamide (DMF) (Thermo Scientific, 20673) was prepared. PBASE and DMF solution was added to both the COVID-19 and Flu-designated GFETs. Glass slide cleaned with ethanol was placed over the device during the 1-hour incubation period to mitigate the risk of DMF evaporating. Starting with one GFET at a time, the PBASE/DMF solution was taken out, and the GFET was rinsed with plain DMF once and DI water three times. Rinsing was performed quickly to avoid drying out the GFET. 50 μg/mL of COVID-19 (CR3022) antibodies were added to the GFET and incubated for an hour. Simultaneously, the Flu-designated GFET went through the same rinsing steps with DMF and DI water with 50 ug/mL of the Flu antibodies, FI6v3, being added with the same incubation time. After one hour of incubation, CR3022 and FI6v3 were taken out of GFET one at a time, and GFET was rinsed with PBS 1× three times. After the rinse, 3 mM PEG-NH2 (Broadpharm, P-22355) and PBS solution were added to the GFET and incubated for another hour. 1M ETA (Sigma Aldrich, 110167) solution was prepared by combining ETA with PBS 1× (pH8). After both GFETs had been incubated with PEG-NH2 for an hour, PEG-NH2 inside the GFET (one GFET at a time) was dispensed and rinsed with PBS 1× three times. The prepared solution of ETA was placed into the GFET and incubated for another hour. All ETA steps were repeated for the other GFET with antibodies. Tw20 (Sigma Aldrich) was placed into a third GFET that didn't contain any antibodies as a negative electronic control. After an hour of incubation with ETA, the ETA solution was dispensed from the GFETs with antibodies and rinsed with PBS 1×. Tw20 was also taken out of its designated GFET, and the GFET was rinsed with PBS 1×.
To ascertain the presence of PBASE and other functionalization reagents on graphene, Raman spectroscopy was performed using Witec Micro-Raman Spectrometer Alpha 300. Electrical functionalization was carried out using Keithley B2909A.
Device measurements were carried out using Keysight B2909 A source-meter for both I-V curve and time-resolved measurements. For functionalization step I-V curves, the PDMS chamber was filled with PBS 1×, and the gate voltage was swept over a range of −0.3 to 0.7 V with Vds=0.1V. For time series measurements against the proteins, the PDMS chamber was initially filled with PBS 0.01× at 400 ul and activated with the chosen gate voltage (voltage for highest transconductance) and Vds=0.1V. The chip was allowed to stabilize for at-least 500s. Before introducing the proteins of interest, a third-party test with Bovine Serum Albumin (BSA) was conducted by adding 25 ul of the BSA solution into the PDMS well. After the test, the chip was disconnected from the source meter and thoroughly rinsed and refilled with PBS 0.01× and reconnected to the source meter with the Vgs and Vds set at the same value as previously stated. Once the reconnected chip stabilized, protein samples were introduced at different concentrations. The samples of both Spike and HA proteins were prepared through serial dilution in PBS 1×. Since the buffer being used for testing is PBS 0.01×, the stock proteins prepared in PBS 1× were resuspended in PBS 0.01× (adding 10 ul of protein in 1×PBS into 990 ul of 0.01×PBS) and thoroughly mixed 5 seconds prior to introducing them to the chip (25 uL of the protein in 0.01×PBS added to 400 uL PBS 0.01× solution on the chip). The measurement was performed in pairs, the first 25 ul of Spike protein in PBS 0.01× was introduced into the chip. Once the current stabilized after reaction in the COVID-19 GFET, then 25 ul of HA protein in PBS 0.01× was added to the chip. This procedure was performed for each concentration of protein to be tested.
The proposed invention proves multi-specificity and advanced target analyte selectivity. We propose to embody multiple target biomolecule specific counterpads, often antibodies or aptamers into the graphene surface, making multiple GFETs per chip sensitive to those biomolecules of interest. As an example, a chip that will have a dual biosensing specificity, such as for COVID-19 and for common Influenza H3N2 viruses.
Due to the nature of field-effect biosensing and electrical recordings, the device's response time to target analytes is within 9-12 seconds. The current sensitivity stands at 2.4% change in signal per log (molar) concentration for COVID-19 (2.4%/log (M)) viral protein particles. 1.9% change in signal per log (molar) concentration (1.9%/log (M)) of Flu. The currently achieved level of detection is ˜47 ag/mL. Multifunctionality and multi-specificity gives additional potential to avoid false negative results, increasing the overall true positive and true negative percentages. The technology can be modular, where future sensing elements can be developed and attached “lego”-style to the existing devices and read-out systems.
One potential limitation is with regards to detection of large biomolecules, such as viruses. Technologically, field-effect biosensing is restricted with the Debye screening length, which is correlated to the molarity of the solution. The human saline has very high molarity, and if used directly, the Debye layer would be within 1 nm, restricting our ability to electrostatically detect large molecules. It can be overcome by diluting the human saline in deionized water to yield lower molarity and lower concentration bio-analyte.
Accordingly provided is a multi-target specific (e.g., COVID-19 and Influenza) biosensor technology utilizing graphene-based field-effect transistors (GFETs). The target-specific antibodies or aptamers (whenever available) will be used to add specificity to graphene channels. The technology will enable an early-stage diagnosis tool that differentiates between multiple analytes of choice, and specifically differentiated between Influenza and COVID-19, yet embodied into a single device, for point-of-care monitoring
By cross-correlating the response from the functionalized and the control sensors, the device will be able to quickly and decisively give its response.
Additional Information regarding the certain aspects of experimental results:
The current detected limit of detection is 47 attogram/ml, but this is not the real limit of the system, the device can go at least one order below this limit. The limit is on the order of a few attograms.
Detection range, based on the current experiments varies from a 0.01 mg/ml down to a few ag/ml. The ng/ml is not exactly the ceiling of our recording system. Any concentration above the ceiling will still be detected as “present”.
The specific chip design, large surface area and high-quality monolayer graphene are perhaps the contributing factors for the superior detection limit of the technology.
The technology can be used to detect from real-life samples (e.g., saliva or sweat, with a simple dilution of the samples in DI water).
Unlike the standard RT-PCR testing kits, which have a turn-around time of at least three hours, our device has a faster response time of a few seconds; hence enabling rapid diagnosis. Furthermore, the device is highly specific, as confirmed by cross-reactivity tests: the COVID-19 functionalized channels are not reactive to the Influenza virus, and vice-versa (
We anticipate (not proven yet) the fabricated biosensors to be able to store for >6 months at 4 C, and >24 hours at room temperature and average (50-100%) humidity.
We have developed design with N=4 compartments. This is a bare minimum to detect and differentiate between two targets. We can increase the number of parallel measurement compartments to any number, e.g., N=12 to differentiate between 10 analytes and N=32 to differentiate between 30 targets.
No special hardware or software has been developed.
Other 2D materials and their heterostructures can be used instead of graphene in the same configuration. MoS2, hBN, WS2, WSe2, PtS2, PtSe2, and others.
Arbitrary substrate can be used. SiO2/Si, sapphire, glass, SiC, and flexible polyimides, parylene, polyester, tape, polycarbonate, Teflon, PET, EVA/PET, etc-any substrate which survives chemical treatment with acetone or toluene.
Detectable items: DNA, aptamers, antibodies, viral particles (flu, covid, rotavirus, ebola, zika, etc.), small molecules, peptides, enzymes, bacteria (e.g., salmonella, e-coli, etc.) proteins, exosomes, toxins (e.g., ochratoxin, cholera, etc.), glucose, etc.
The system can be flipped, and target towards measuring antibodies [ref]
The device is biocompatible, or at least can be made on a biocompatible substrate. The large scale grown graphene alongside the used elements in making the biosensor are biocompatible.
The substrate is insulating, and the electric field over the graphene channel is applied throughout the electrolyte, forming the electrical double layer with high dielectric constant. We commonly use a diluted 0.01×PBS (Phosphate-buffered saline) yet the device can work with any electrolyte (Borate buffered saline, Tris-buffered saline, etc.), even the home tap water.
A single large area (almost around 2×2 cm) large monolayer graphene piece is transferred on top of the pre-fabricated chip with conductive, Au, Pt, or any other metal-based feedlines structured via shadow mask evaporation or photolithography, creating four graphene channels arranged radially with one common electrode and N=4 (or more) radial compartment. After graphene transfer, PDMS chamber with two different heights are placed on top: the inner compartments made of height A, while outer walls with height B. B>A in order to functionalize the compartments separately while measure all for elements for biosensing together at the last stage.
Antibody immobilized GFETs according to principles described herein have registered the lowest measured concentration of the COVID-19 Spike protein and the Flu surface protein, Hemagglutinin (HA), at around 88 zM and 227 zM, respectively. Combined with almost negligible cross-reactivity, we can claim a fast and specific response with the reaction time of ˜10 s depending on the antigen. Together, the performance of the proposed devices opens the possibility of diagnosing patient's conditions well ahead of the 5-day gap suggested by the CDC thus helping in curbing the spread of disease.
Designing for simultaneous and differential detection of COVID-19 and Flu, we describe a sensor platform consisting of an array of GFETs driven through a common gate and shared biological media with LoD at 88 zM for COVID-19 and 227 zM for Flu. These findings provide a proof-of-concept principle solution to the problem of rapidly differentiating two or more diseases with overlapping symptoms. The device enables immediate readout with a rapid turnaround time of around 10s. The differential sensing results from high specificity and sensitivity accorded by the specific immobilization of the antibodies on two GFETs accompanied by an electronic control in the form of passivated GFET. The device presents a highly specific, facile, and portable electronic point of care technology. It would especially benefit areas with high density and volume of patients and visitors such as clinics, nursing homes, universities, offices, etc.; mitigating the bottlenecks created due to high turnaround times and complicated testing procedures presented by conventional technologies. The multi-channel GFET device is also highly versatile since it can be repurposed with antibodies/receptors specific to other diseases, thus serving to track and mitigate future epidemic and pandemic threats.
The construction and arrangement of the systems and methods as shown in the various exemplary embodiments are illustrative only. Although only a few embodiments have been described in detail in this disclosure, many modifications are possible (e.g., variations in sizes, dimensions, structures, shapes and proportions of the various elements, values of parameters, mounting arrangements, use of materials, colors, orientations, etc.). For example, the position of elements may be reversed or otherwise varied, and the nature or number of discrete elements or positions may be altered or varied. Accordingly, all such modifications are intended to be included within the scope of the present disclosure. The order or sequence of any process or method steps may be varied or re-sequenced according to alternative embodiments. Other substitutions, modifications, changes, and omissions may be made in the design, operating conditions, and arrangement of the exemplary embodiments without departing from the scope of the present disclosure.
GFET-based biodetectors have been demonstrated by inventors listed here for the early detection of iron deficiency. These same principles are applicable to functionalizing a GFET for use in a biosensor for any desired antigen by replacing the anti-ferritin antibody with an antibody of choice.
Iron deficiency (ID) is the most prevalent and severe nutritional disorder globally and is the leading cause of iron deficiency anemia (IDA). IDA often progresses subtly symptomatic in children, whereas prolonged deficiency may permanently impair development. Early detection and frequent screening are, therefore, essential to avoid the consequences of IDA. In order to reduce the production cost and complexities involved in building advanced ID sensors, the devices were fabricated using a home-built patterning procedure that was developed and used for this work instead of lithography, which allows for fast prototyping of dimensions.
Described herein is the development of graphene-based field-effect transistors (GFETs) functionalized with anti-ferritin antibodies through a linker molecule (1-pyrenebutanoic acid, succinimidyl ester), to facilitate specific conjugation with ferritin antigen. The resulting biosensors feature an unprecedented ferritin detection limit of 10 fM, indicating a tremendous potential for non-invasive (e.g., saliva) ferritin detection.
Nutrition during the early years of life has a preeminent influence on the quality of health of an individual in their lifetime [1-3]. Specifically, micronutrients provide the essential building blocks for brain development, healthy growth, and a robust immune system [1,4-7]. The top three micronutrients of global health relevance are iodine, iron, and vitamin A, whereas iron deficiency is the most common nutritional disorder worldwide [8,9].
Iron deficiency (ID) refers to a condition of significantly low concentration of healthy red blood cells in the body due to the correspondingly low amount of iron [10,11]. The core function of iron in the body is oxygen transport in the blood. Iron deficiency, if not diagnosed and treated at the early stage, will lead to iron deficiency anemia (IDA). Although every age group is vulnerable, it is more prevalent in women and children [8,12]. However, it is often impossible to recognize ID in children until it degenerates to IDA. At that point, symptoms such as pale skin, frequent infections, fatigue/lethargy, pica, and poor appetite become apparent. ID impairs the cognitive development of children from infancy through to adolescence and is associated with increased morbidity rates [13,14]. It is, therefore, imperative to be able to promptly detect iron deficiencies in children, so that intervention programs are timely and better targeted.
Although iron status is best assessed by a combination of indicators, ferritin is established as the major iron-storage molecule; its production increases in cells as iron supplies increase. The serum ferritin level is, therefore, the most specific biochemical test that correlates with relative total body iron stores; hence, it is the most widely used iron status indicator [8]. However, since ferritin is an acute-phase reactant protein, its concentration is elevated in the presence of infection or inflammation. A child under five years of age is said to be iron-deficient if their serum ferritin level is <12 μg/L, while the threshold is <15 μg/L for children over five years old but rises to <30 μg/L in the presence of an infection [15]. Hence, ferritin tests should be taken very seriously when the results are abnormally low compared to when the measure is normal [10].
Investigating iron deficiency involves a continuous process of recording and assessing iron status in an individual to identify a drop in the indicator levels. Therefore, non-invasiveness becomes necessary, especially when children are involved. The use of saliva presents a non-invasive approach. Saliva is known to contain every information present in the blood but in significantly smaller quantities. Some research works demonstrated the use of saliva for micronutrient testing [16,17]. Moreover, research went into determining salivary ferritin concentrations in humans, as well as correlating serum (or plasma) and salivary ferritin concentrations, as presented in the Table 1, below.
From Table 1, it is clear that the lowest literature-reported iron-deficient salivary ferritin concentration is 0.186 μg/L, which is significantly lower than the 12 μg/L iron-deficient serum ferritin concentration level. The significantly low levels of ferritin in human saliva make it impossible to use the current micronutrient biosensors presented in the literature (presented in Table 2, below).
PAD to derive plasma;
0 ng/mL
-like
FET
)
indicates data missing or illegible when filed
The detection mechanisms of all sensors reported in Table 2, except the silicon nanowire type [29], are all optoelectronic. The biosensor detection range of the lateral flow immunoassay (LFIA)-based sensor is 3-556 μg/L in buffer and 5.78-888 μg/L in serum ferritin standard. However, the sensitivity and specificity reported are based on a detection limit of 18 μg/L. Ferritin concentrations lower than this cut-off resulted in a degradation of the sensitivity and specificity of the biosensor. The silicon nanowire detection mechanism is a departure from the rest in its utilization of a nano-field-effect transistor (FET). This presented the significant advantage of a lower detection limit as compared with others. From an extensive literature search, it was observed that there are significantly few studies on the detection of ferritin concentration using field-effect biosensors. Their method attained a ferritin detection limit down to 50 μg/mL using a horn-like polycrystalline-silicon nanowire (SiNW) FET. Even though their fabrication method is acclaimed to be simpler, the synthesis of silicon nanowires is generally non-trivial and expensive [30,31]. On the other hand, graphene synthesis is simple; graphene is widely commercially available and inexpensive. Moreover, unlike SiNWs, the two-dimensional (2D) planar surface structure of graphene facilitates ease of functionalization. Graphene fabrication and transfer to the substrate are also significantly simple compared to the procedure for SiNWs [32].
Since the first exfoliation of a single atomic layer of graphene in 2004 by Geim and Novoselov [33], of all other nanomaterials, it is known to be the most promising nanostructured material suitable for biosensing, under intense research for over a decade [34,35].
In this research work, we developed an FET biosensor using monolayer graphene as the conducting channel. We functionalized the channel using anti-ferritin antibodies to selectively capture the ferritin protein antigen, with a limit of detection about 10 fM. It is noteworthy that this performance was attained despite using our low-cost and straightforward shadow mask patterning procedure to derive the source and drain electrodes of the graphene-based FETs (GFETs), rather than the standard (ultraviolet (UV) or e-beam) lithography process [36]. This work is the first report of ferritin detection using graphene. It also offers the lowest ferritin detection limit obtainable by any reported sensor. This work demonstrates the enormous potential of using a GFET for non-invasive early detection of iron deficiency.
Graphene on 25-μm-thick copper foil (Gr/Cu) synthesized through chemical vapor deposition (CVD) was purchased from Chongqing Graphene Technology Co., Ltd. (also known as Chongqing Moxi Technology). The following materials were ordered from Millipore Sigma (formerly Sigma-Aldrich): ferritin, anti-ferritin antibody, dimethylformamide (DMF), Tween-20, ethanolamine (ETA), and ˜150 mM phosphate-buffered saline (1×PBS, pH 7.4 at 25° C.). Here, 1.5 mM PBS (0.01×PBS) was prepared by diluting 1×PBS appropriately with de-ionized water. Furthermore, 1-pyrenebutanoic acid, succinimidyl ester (PASE) was purchased from Thermofisher Scientific.
A 285-nm-thick SiO2 on Si wafer was used as a substrate. The source and the drain of the transistor used in this work were patterned according to an interdigitated electrode (IDE). The IDE-structured transistors were fabricated using a shadow mask to pattern the electrodes on top of the SiO2/Si substrate. The masks were fabricated via a simple yet robust technology, which allows for fast prototyping of desirable patterns at a fraction of time and cost, and which utilizes a commercially available, off-the-shelf tool, Silhouette Cameo, capable of providing resolution down to 200 um [37]. A detailed account of this process was reported elsewhere [36]. Although this is a more straightforward approach to patterning in contrast with lithography, there is a limit on the sizes obtainable due to the resolution of the mechanical cutting machine. Each SiO2/Si wafer yielded 28 transistors based on the pre-set dimensions. Since an IDE structure was used, the overall length of the channel was set to 1 mm, and the width was set to 68.8 mm, yielding a ˜69 W/L ratio. Using the CHA e-beam-assisted evaporator, a thin layer each of Ni (10 nm) and Au (90 nm) was deposited.
Nickel was deposited first to serve as an adhesion layer, while gold was the metal contact serving as the source and drain for the transistor. The purchased Gr/Cu was cut into desired sizes and stuck onto dummy silicon wafers. A protective polymer (poly(methyl methacrylate) (PMMA)) was drop-casted onto the Gr/Cu and spin-coated for even distribution. The resulting PMMA/Gr/Cu was then annealed at 150° C. for 5 min. This was thereafter transferred onto the etchant (0.1 M ammonium persulfate) to remove the underlying copper foil, leaving PMMA/Gr on top of the solution. The PMMA/Gr was then triple-washed with deionized (DI) water, followed by a careful transfer of a PMMA/Gr sheet onto each IDE-structured transistor to bridge the source and drain electrodes. After the transfer, PMMA/graphene was left to slowly dry out for 12 h at room temperature, followed by 5 min of 150° C. annealing in order to re-flow the PMMA and improve graphene-substrate adhesion. The devices were then left for 24 h in acetone in order to remove the protective PMMA layer, then washed with IPA, and dried with an oxygen gun. Polydimethylsiloxane (PDMS) chambers were then molded and attached to each chip to form an exposed well above the graphene sensing area, thereby creating the means for liquid-based measurements. The GFET fabrication process is summarized in
Transforming a GFET into a specific biosensor requires immobilization of the necessary biomolecules as seen in
After functionalization, the target analyte, ferritin, was prepared in 0.01×PBS to obtain the desired concentrations.
Prior to functionalization, we took the Raman spectra of the graphene on our IDE FET substrate via the Renishaw in Via Raman microscope, using the blue excitation laser wavelength of 442 nm and 4 mW power on the sample to verify the graphene quality and number of layers. The GFETs were electrically characterized at room temperature prior to functionalization, after antibody immobilization and after applying the blocking buffer (Tween-20 and ETA). All measurements were based on a liquid-gated FET set-up. We used 0.01×PBS as the electrolyte buffer solution, and a Keithley B2902A Source Measure Unit (SMU) coupled to a Wentworth Labs probe station. Ag/AgCl pellet electrodes (E-206, Science Products) were used as gate reference electrodes, and they were carefully washed between experiments in order to avoid any cross-contamination. The immobilization processes were characterized by monitoring the drain current changes for a drain-source voltage (VDS) of 0.2 V while sweeping the gate voltage (VGS) from −0.5 to 0.5 V. The sensor performance was determined by monitoring the drain current changes per time for a given drain-source voltage and gate voltage, as the GFET was exposed to the different concentrations of ferritin. The time-trace recordings were performed while keeping both VDS and VGS constant at a certain operational point. The point was set to be VDS=0.1 V and VGS=0.05 V to make sure there were no excessive currents through the graphene.
The used graphene was a high-quality monolayer, as verified by the I2D/IG ratio >2 [38]. The Raman spectrum also revealed a minimal D peak at 1350 cm−1, showing very low defect density. The quality of this graphene facilitated consistent GFET transport properties and confirmed the high fabrication yield of >99% as specified by the manufacturer.
Characterizations were based on transfer curves obtained by plots of drain current versus the gate voltage during stages of fabrication and functionalization of the GFET. The transfer curve obtained by characterizing the bare GFETs immediately after fabrication showed that the GFETs had an average and positive-valued Dirac voltage (Vdirac) of 211±60.4 mV (
The liquid-gated FET (LG-FET) measurement set-up is the primary measurement configuration for biosensors, where the “liquid” is the sample containing the analyte to be detected or quantified. In this LG-FET set-up, the gate voltage that triggers the modulations in the device is applied to a reference electrode through the liquid to the graphene channel. As this potential is applied, the ELECTRICAL DOUBLE LAYER (EDL) with a capacitance value of CEDL is formed just above the graphene channel. In effect, the CEDL in series with the air-gap capacitance due to graphene's hydrophobicity and the inherent quantum capacitance of graphene produce the total gate capacitance of the GFET. Therefore, a significant advantage of this set-up is the low operating voltage required for the device, typically within 1 V. The thickness of the EDL is a function of the Debye length (λD) as seen in Equation (1). When antigens bind to their antibodies immobilized on the FET surface, a change in surface charge is induced at the binding site. For the changes to be effectively captured, the binding site must be within the Debye length, defined by Equation (2) [41]. Therefore, changes that occur outside this length are subject to electrostatic charge screening.
where ε0 is the permittivity of free space, εr is the relative permittivity of the dielectric formed between the graphene surface and the liquid, and M (molarity) is the ionic strength of the sample (liquid). from Equation (2), it is evident that a higher molarity results in a shorter Debye length. This concept is of great concern because most biological interactions take place within high-ionic-strength solutions (e.g., 1× PBS ionic strength=˜150 mM). In effect, an attempt to sense these interactions electronically using FET-based sensors is severely impeded by the consequentially short Debye length (0.7 nm for 1×PBS). Therefore, although the binding efficiency of ferritin and its antibody is high due to its large molecular size [42], to ensure this binding is detected by the GFET biosensor, 0.01×PBS (M=1.5 mM, λD=7.3 nm) was used as the electrolyte to carry out the measurements.
The functionalization process incurs some height on the graphene surface that eats into the Debye length. However, the literature highlights that the incurred height from the sensor surface after a flat-on-orientation immobilization of the antibodies is typically about 4 nm [29,43]. Therefore, even for macromolecular antigens like ferritin, using 0.01×PBS will give room for detection of the antigen-antibody binding since the binding site will be within the Debye length of ˜7.3 nm.
For a p-type GFET device, the number of holes is greater than the number of electrons; hence, on the application of the gate voltage, decreased conductivity results. On the other hand, when the GFET is n-type, the application of the gate voltage leads to increased conductivity. However, the immobilization and the binding of charged target biomolecules to receptors on the channel yield specific channel modulation effects. For a p-type device, when a negatively charged biomolecule binds to the receptors on the graphene channel, holes accrue in the channel, leading to increased drain-source current [44]. This binding corresponds to a negative gating potential of the graphene channel and, hence, the reduced carrier density of graphene [45]. On the contrary, when a positively charged biomolecule binds to the receptors on the graphene channel, reduced drain-source current results [46]. Ferritin is a negatively charged molecule with a weight of 474 kDa [47-49]; therefore, with a GFET operated in hole-conduction mode, it is expected that the drain-source current increases (resistance decreases) as the antigen is immobilized on the device. Monitoring of current change is carried out at a certain working potential (0.05 V in this case), and the shift of current is a typical response of biomolecule attachment [40,50-53]. This expected trend can be observed in
Concerning the detection limit and range of the GFET biosensor, we started pipetting the ferritin antigen onto the chip from the smallest concentration of 10 ng/L, consequently increasing the ferritin concentration up to 8 μg/L. The initial concentration resulted in a significant rise in drain current, which suggests that the smallest analyte concentrations detectable by the developed GFETs are actually lower than 10 ng/L. Notably, the changes in drain current upon ferritin immobilization occurred within less than 10 s of pipetting the protein onto the GFETs, portraying real-time detection.
We consider A (antibody) and F (ferritin) to be two interacting bio-objects which can form a bound product, AF, and we let CA, CF, and CAF be their concentration in M (molarity). The time dependent rate equation for the formation of the product CAF is
Where the forward reaction rate constant kon, and the reverse reaction rate constant koff.
In equilibrium, the sum of all time-dependent derivatives is zero, which in fundamental interpretation obeys the law-of-mass-action equation in solution [54].
The strength of the interaction between A (antibody) and F (ferritin) can be linked to the affinity constant Ka via the concentration of bound ferritin molecules to the concentration of antibodies. However, it is also necessary to consider the dissociation constant KD, because it can be compared to the reactant ferritin concentrations. In solution, the total concentration of bound antibody-ferritin complex (CAF) depends on the concentration of both antibody and ferritin for biosensors with active surface areas where the law-of-mass-action applies [55].
Immobilized antibodies on the biosensor surface are fixed and, thus, the number of captured ferritin molecules will not change. To have an ideal experiment, the number of ferritin antigens should be in large excess with respect to the number of immobilized antibodies, such that the effective total concentration does not change when ferritin antigens adsorb from the solution to the surface.
By combining Equations (5) and (6), we can re-arrange and get the equivalent of the law-of-mass-action for active surface biosensors.
Equation (7) corresponds to the Langmuir isotherm [57], which is derived for the adsorption of the molecules onto surfaces (in this case, on the biosensor surface with attached antibodies) [53,58].
Compared to the law-of-mass-action, this method is simpler and only depends on the ferritin concentration CF and the equilibrium dissociation constant for the antibodies KD. In the specific case of antibodies binding to ferritin antigens, the affinity constant Ka should be calculated. With the known affinity constant, the binding isotherm for the antibody occupancy with the bonded ferritin antigen can easily be plotted (
In this work, we demonstrated the possibility of using graphene to develop an FET biosensor for the detection of serum ferritin protein, whose level gives reliable information about iron deficiencies in the human body. This is the first reported GFET biosensor for ferritin detection. These GFETs were fabricated using our innovative and low-cost method of preparing a shadow mask for patterning and evaporating metal contacts on the substrate. From our analysis, the ferritin detection limit of the GFET biosensor is 5.3 ng/L (10 fM), which is the lowest detection limit reported for ferritin in the literature, while the detection range is 5.3 ng/L (10 fM) to ˜0.5 μg/L (1 pM). These results show that there is excellent potential in using these GFETs for non-invasive ferritin sensing characterized by very low detection limits.
References discussed in paragraphs [0074]-[00108] are as follows:
While various embodiments of the present invention have been described above, it should be understood that they have been presented by way of example only, and not limitation. It will be apparent to persons skilled in the relevant art that various changes in form and detail can be made therein without departing from the spirit and scope of the present invention. Thus, the breadth and scope of the present invention should not be limited by any of the above-described exemplary embodiments, but should be defined only in accordance with the following claims and their equivalents.
This application claims the benefit of priority to U.S. Provisional Application No. 63/284,706 filed Dec. 1, 2021, which is hereby incorporated herein by reference in its entirety.
This invention was made with government support under Grant no. ECCS2033846 awarded by the National Science Foundation. The government has certain rights in the invention.
| Filing Document | Filing Date | Country | Kind |
|---|---|---|---|
| PCT/US2022/051427 | 11/30/2022 | WO |
| Number | Date | Country | |
|---|---|---|---|
| 63284706 | Dec 2021 | US |
