key: cord-0989520-fj81l3hp
authors: Chen, Mingrui; Cui, Dingzhou; Zhao, Zhiyuan; Kang, Di; Li, Zhen; Albawardi, Shahad; Alsageer, Shahla; Alamri, Faisal; Alhazmi, Abrar; Amer, Moh. R.; Zhou, Chongwu
title: Highly sensitive, scalable, and rapid SARS-CoV-2 biosensor based on In(2)O(3) nanoribbon transistors and phosphatase
date: 2022-03-28
journal: Nano Res
DOI: 10.1007/s12274-022-4190-0
sha: 301675e17b2ab212cef31acce12d1c6d8d9c22ef
doc_id: 989520
cord_uid: fj81l3hp

Developing convenient and accurate SARS-CoV-2 antigen test and serology test is crucial in curbing the global COVID-19 pandemic. In this work, we report an improved indium oxide (In(2)O(3)) nanoribbon field-effect transistor (FET) biosensor platform detecting both SARS-CoV-2 antigen and antibody. Our FET biosensors, which were fabricated using a scalable and cost-efficient lithography-free process utilizing shadow masks, consist of an In(2)O(3) channel and a newly developed stable enzyme reporter. During the biosensing process, the phosphatase enzymatic reaction generated pH change of the solution, which was then detected and converted to electrical signal by our In(2)O(3) FETs. The biosensors applied phosphatase as enzyme reporter, which has a much better stability than the widely used urease in FET based biosensors. As proof-of-principle studies, we demonstrate the detection of SARS-CoV-2 spike protein in both phosphate-buffered saline (PBS) buffer and universal transport medium (UTM) (limit of detection [LoD]: 100 fg/mL). Following the SARS-CoV-2 antigen tests, we developed and characterized additional sensors aimed at SARS-CoV-2 IgG antibodies, which is important to trace past infection and vaccination. Our spike protein IgG antibody tests exhibit excellent detection limits in both PBS and human whole blood ((LoD): 1 pg/mL). Our biosensors display similar detection performance in different mediums, demonstrating that our biosensor approach is not limited by Debye screening from salts and can selectively detect biomarkers in physiological fluids. The newly selected enzyme for our platform performs much better performance and longer shelf life which will lead our biosensor platform to be capable for real clinical diagnosis usage. [Image: see text] ELECTRONIC SUPPLEMENTARY MATERIAL: Supplementary material (materials and methods for device fabrication, functionalization of In(2)O(3) devices, photographs of the liquid gate measurement setup, mobilities of the nine devices labeled in Fig. 1(b), family curves of I(DS)−V(DS) with the liquid gate setup and current change after bubbling the substrate solution (current vs. time curve for S1 antigen detection)) is available in the online version of this article at 10.1007/s12274-022-4190-0.

On the 11th of March 2020 the World Health Organization (WHO) declared Coronavirus disease 2019 (COVID- 19) outbreak, caused by the severe acute respiratory syndrome coronavirus (SARS-CoV-2), as a pandemic [1] . SARS-CoV-2 can cause severe respiratory distress, and as of November 2021, cumulative number of infected people is over 250 million, resulting in over 5 million death all over the world [2, 3] . Owing to the high infectivity of the virus and lack of specific treatments against it, early diagnosis of COVID-19 is essential in pandemic prevention and control, which mandates a highly sensitive and yet more scalable, cost-efficient, and rapid detection method of SARS-CoV-2, particularly for developing countries carrying a heavy economic burden caused by COVID-19 pandemic [4] .

SARS-CoV-2 mainly contains 4 types of proteins: spike(S), membrane(M), nucleocapsid(N), and envelop(E) proteins ( Fig.   1(a) ). Among these proteins, the S protein is on the viral envelope and its S1 subunit is the outermost component of the virus. During SARS-CoV-2 infection, S1 proteins recognize and bind with the cellular receptor on the surface of the target cells, facilitating the virus particles to enter the target cells and triggering the infection [5] [6] [7] [8] . Therefore, S1 protein is an important target for SARS-CoV-2 diagnostic detection. Moreover, detecting S1 protein specific antibody in human blood is also important for COVID-19 diagnosis in terms of determining previous viral exposure and immunity status. Antibody tests (also known as serology test) can be nearly 100% accurate for blood samples collected 20 days after infection or the onset of symptoms [9] [10] [11] . In this work, we developed a versatile electronic biosensing platform detecting both SARS-CoV-2 S1 protein and S1 protein IgG antibody, whose detection target can be easily switched between the two biomarkers.

Thus far, based on previously established laboratory protocols, diagnosis of COVID-19 depends mainly on nucleic acid amplification tests (NAATs), specifically real-time reverse transcription-polymerase chain reaction (RT-PCR) [12] [13] [14] . Although this method is sensitive and shows good specificity for COVID-19 detection, this technique has a long processing time (usually 4-5 h) and depends on advanced instruments, expensive reagents, and skilled technicians.

Among recent advances in rapid screening methods, field-effect transistor (FET)-based biosensors have many advantages such as small size, real-time detection, high-sensitivity, and capability for integrated multiplexing [15] [16] [17] [18] [19] [20] . Seo et al. reported graphenebased FET biosensors for spike protein of SARS-CoV-2 [21] , and Shao et al. demonstrated single-walled carbon nanotube (SWCNT)-based FETs to probe spike and nucleocapsid proteins of SARS-CoV-2 [22] . As it is widely known, graphene is semimetallic while semiconducting carbon nanotubes usually exhibit ptype conductivity. To reduce the probability of getting false positive and false negative diagnosis, it is desirable to have complementary sensing capability by using both p-type and ntype semiconductors [23] . In 2 O 3 is a well-known n-type semiconductor and has been successfully applied as the channel material for a variety of biosensors [24] [25] [26] [27] [28] [29] . In our work, we fabricated In 2 O 3 transistors using a lithography-free process. Our simplified fabrication method made In 2 O 3 biosensors cost-efficient and suitable for large-scale COVID-19 clinical diagnosis.

FET-based biosensors can be combined with enzymatic reactions because many enzymatic reactions can release acidic or basic products, in which process the pH change of the environment will generate electronic signal of the FETs [30, 31] . Urease is the most common enzyme incorporated with FET-based biosensors [32] [33] [34] [35] [36] [37] . However, urease displays poor stability [38] [39] [40] [41] [42] , which would lose most of the enzymatic activity after being stored at 4 ℃ for 10 days [42] . In our work, we applied alkaline phosphatase as the enzyme reporter for our electronic biosensors. During the sensing process, alkaline phosphatase produces proton by enzymatically removing a phosphate group from the substrate compound, and the decreased pH value of the medium will then lead to an increase in the current of our In 2 O 3 FETs. Phosphatase is widely used as the enzyme in commercial colorimetric ELISA kits (LoD: 16 pg/mL [43] , 0.55 mg/mL [44] ). Compared to urease, phosphatase is much more stable and suitable for long-term storage while maintaining high activity. Consequently, our biosensing platform with phosphatase overcomes the limit of electronic biosensing technique, whose application used to be constrained by the poor stability of enzyme reporter. Our biosensors display a prolonged shelf life and is advantageous for practical COVID-19 diagnosis.

In this work, Our In 2 O 3 biosensors were able to detect S1 antigen of SARS-CoV-2 in both phosphate-buffered saline (PBS) buffer and universal transport medium (UTM) with a reliable performance and similar LoD of 100 fg/mL. Besides antigen detection, S1 protein IgG antibody detection in both PBS and human whole blood (WB) were also carried out. The detection performance in both mediums displayed LoD of 1 pg/mL. The similar detection results in different mediums indicate the capability of our electronic biosensors to eliminate the interference of salts, cells, and other particles in detected specimens. Taken together, our biosensors are highly sensitive and scalable, showing good potential for clinical SARS-CoV-2 diagnosis. Our findings shed light on a new generation of biosensors for rapid and highly sensitive COVID-19 screening, getting us closer to winning the fight against COVID-19 pandemic. Figure 1 (b) shows the photograph of the In 2 O 3 nanoribbon FET devices on a Si/SiO 2 (500 nm) substrate. The lithography-free fabrication of the FETs includes two steps. First, In 2 O 3 nanoribbons were defined by the first shadow mask and deposited using radio frequency (RF) sputtering technique. The channel length and width of the In 2 O 3 nanoribbons were 500 and 25 µm, respectively ( Fig. 1(c) ), and the thickness was 18 nm (Fig. 1(d) ). Next, the source and drain electrodes were defined by the second shadow mask. After alignment of the second mask, 1 nm Ti and 50 nm Au were deposited by electron beam evaporation. Twentyeight biosensing chips can be fabricated on a 3-inch substrate, and each chip contains an array of 4 separate In 2 O 3 FET devices. The electrical performance of the devices was first characterized in ambient environment using an Agilent Semiconductor Analyzer 1500B. Nine different In 2 O 3 FET devices were randomly picked over the substrate (labelled in Fig. 1(b) ), and their drain current−back gate voltage (I DS −V GS ) curves with the drain voltage fixed at 1 V are shown in Fig. 1 (e). To further demonstrate the uniformity of our devices, I DS −V GS curves were measured for more devices at different locations on our substrate with drain voltage fixed at 150 mV and 5 V (Figs. S2(a) and S2(b) in the ESM), both of which show that our devices are uniform. Moreover, to demonstrate the stability of our devices, we have measured I DS −V GS curves of a single device for 20 times over a period of 1 week, as shown in (Fig. S2 (c) in the ESM), which showed very little variation. Based on these measurements, we can deduce that our fabricated devices exhibit high uniformity and stability.

The I DS −V GS curves of the FET devices illustrate n-type transistor behavior with an average mobility of 108±4.3 cm 2 /V·s ( Fig. S3 in the ESM) and on-off ratio of ~1×10 5 . Owing to the thick SiO 2 dielectric, the devices can be turned on only at high back gate voltage. Figure 1 (e) demonstrates the electrical uniformity of In 2 O 3 nanoribbon devices over the entire wafer.

Figure 1(f) shows the schematic diagram depicting the functionalization of our In 2 O 3 devices for detection of S1 protein.

The photograph of the experimental set up is shown in Fig. S1 in the ESM, a Teflon cell with an opening at the bottom was used to contain the medium in contact with the In 2 O 3 biosensors. Prior to the electronic biosensing process, the In 2 O 3 devices were first immersed in boiling acetone and isopropyl alcohol for 5 min each to clean the surface, followed by an O 2 plasma treatment to generate hydroxyl groups on the surface of In 2 O 3 . 3-Phosphonopropioninc acid, as the linker molecules, was then applied to bind with the hydroxyl groups on the surface. Subsequently, devices were functionalized with N-(3dimethylaminopropyl)-N′-ethylcarbodiimide hydrochloride/Nhydroxysuccinimide (EDC/NHS), creating activated carboxyl groups to bind with amino groups from proteins. Next, capture antibodies that can specifically target S1 antigens of SARS-CoV-2 were applied and immobilized via reaction with the EDC/NHS groups on the surface. This immobilization process is followed by a washing step to rinse off unbounded antibodies as well as a blocking step by 10% bovine serum albumin (BSA) solution to deactivate the surface of In 2 O 3 in order to avoid non-specific binding during consecutive processes. Target biomarkers were then introduced to our biosensors. As shown in Figure 1 (f), S1 antigen solutions with known concentrations were dissolved in either PBS or UTM followed by a 1-hour incubation period. Next, biotinylated secondary antibody were introduced and incubated for 1 hour followed by a 40-min incubation of streptavidin conjugated phosphatase. Finally, phosphatase substrate solution was applied and an amperometric signal was recorded as the result of the detection. The total time required since applying biomarkers until obtaining final sensing result is nearly 3 hours, with actual detection time ranges between 10 to 15 min.

Our electronic detection of SARS-CoV-2 is based on enzymatic reaction of alkaline phosphatase. As shown in Fig. 2(a) , when catalytic alkaline phosphatase is present, the phosphate group of the substrate compound would get removed during the reaction and protons would get produced, leading to a change in the solution's pH value. To prepare the substrate solution, 10 mg of 4-Nitrophenyl phosphate disodium salt hexahydrate, which is the commercial phosphatase substrate for colorimetric ELISA, was dissolved in 10 mL of 0.5 M MgCl 2 solution. The pH value of the substrate solution was then adjusted to 9.7 by NaOH solution. We monitored the pH change after manually mixing 5 µL of 1.5 mg/mL alkaline phosphatase solution with 3 mL of the substrate solution, and the pH value decreased by 1.3 in 5 minutes ( Fig.  2(b) ). Our In 2 O 3 nanoribbon FET devices served as a pH sensor in response to the pH change caused by the phosphatase enzymatic reaction. Hence, devices are required to operate in a wet environment ( Fig. 1(f) ). An Ag/AgCl reference electrode was used to apply liquid gate bias. Figure 2 (c) shows the drain current−liquid gate voltage (I DS −V GS ) curve of the devices immersed in 0.01×PBS buffer with the drain voltage fixed at 200 mV. Figure S4 in the ESM shows the family curves of drain current−drain voltage (I DS −V DS ) where In 2 O 3 -FETs display good current saturation at high V DS . Based on these measurements, we can deduce that our In 2 O 3 -FETs can be reliably and efficiently controlled using the liquid gate with phosphatase was directly attached to the FET device by EDC/NHS chemistry, while the detailed incubation process was similar to the S1 antigen detection described above.

To ensure the detection of the final amperometric signal, prior to adding substrate solution, the channel region of the In 2 O 3 FET was immersed in MgCl 2 and NaOH solution with the on/off ratios (I on /I off ) of 10 4 .

When the enzymatic reaction starts, the deprotonated hydroxyl groups on the surface of In 2 O 3 nanoribbons become protonated as the pH decreases. Therefore, the negative charges on the surface of In 2 O 3 nanoribbons decrease and the conduction of the material increases. Figure 2 (d) shows the current change of the FET devices in response to pH change, with a fixed liquid gate voltage at 150 mV and a fixed drain voltage at 200 mV, and the I DS of the device increased by a factor of 14 as the pH changed from 10 to 4.

In order to demonstrate that In 2 O 3 nanoribbon FET devices and phosphatase enzymatic reaction can be integrated for electronic biosensing, we performed a positive control experiment. Figure 2 (e) depicts the schematic diagram of the sequence of molecule binding. The streptavidin-conjugated same pH value as the substrate solution. A liquid gate voltage of 150 mV was applied while the drain voltage was fixed at 200 mV to generate the observed baseline. Substrate solution was then added to trigger the enzymatic reaction and I DS of the device was recorded as a function of time. As shown in Fig. 2(f) , the I DS started to increase as soon as the substrate was introduced with a current enhancement of 110% after 8 minutes.

Furthermore, to examine the interference in our biosensing results caused by potential electrochemical reaction happening in the substrate solution and the change in solution volume during the electrical measurement, we performed a negative control experiment. With fixed liquid gate voltage at 150 mV and the substrate solution (pH = 9.7), which is used in our biosensing setup, we measured I DS over time, as shown in Fig. S5 in the ESM. We observed a very stable I DS vs. time curve, indicating that no electrochemical reaction that can cause observable current change occurred within the rather small bias of 150 mV we used. In addition, we varied the volume of the substrate solution from 200 to 400 µL during this process and no current change was observed. Consequently, we can deduce that the volume change of the substrate solution does not affect the result of our biosensors.

We performed electronic biosensing targeting at S1 antigen of SARS-CoV-2 spike protein in 1 × PBS with known concentrations ranging from 100 fg/mL to 1 ng/mL. Figure 3(a) shows the normalized responses for each concentration. After taking the average of sensing results from 5 different biosensors and subtracting the negative control, we obtained responses of 5%, 11%, 28%, 54% and 78% of current increase for antigen concentrations of 100 fg/mL, 1 pg/mL, 10 pg/mL, 100 pg/mL and 1 ng/mL S1, respectively (Fig. 3(c) ). The negative control was performed by merely replacing the antigen biomarker with BSA solution during the functionalization process. The slight increase in the signal of the negative control can be caused by the nonspecific binding of the phosphatase to the In 2 O 3 nanoribbons.

The detection of S1 protein dissolved in UTM with the same range of concentrations was carried out as well. Universal transport medium is a common solvent for collection, transport, maintenance, and long-term freeze storage of clinical specimens containing viruses. It's widely used for cell culture, rapid antigen detection, PCR, and nucleic acid amplification assays. Hence, antigen detection in UTM helps to assess our biosensor to be used for clinical diagnosis of SARS-CoV-2. In order to simulate a real clinical setting, we have adopted a nasopharyngeal swab sample taken from a healthy person in our UTM-based detection setup. The swab was suspended in the UTM along with S1 protein. Real-time responses monitored at 100 fg/mL, 1 pg/mL, 10 pg/mL, 100 pg/mL and 1 ng/mL of S1 proteins in UTM. (c) Plot of normalized responses at each S1 protein concentration, in nanograms per milliliter, n = 5, error bars represent standard deviation. (d) Real-time responses monitored using ten biosensors from four batches of fabrication at 10 pg/mL of S1 protein, showing the reproducibility of our technique. and 78% of current changes (average of 5 trials) have been generated in response to antigen concentrations of 100 fg/mL, 1 pg/mL, 10 pg/mL, 100 pg/mL and 1 ng/mL, respectively (Fig.  3(c) ). The results of detections in PBS (black squares) and UTM (red dots) are plotted in Fig. 3(c) . The results illustrated that the ingredients of UTM, such as Hank's balanced salts, BSA and sucrose, would not interfere with the sensing results. This finding reveals that our electronic biosensing approach can circumvent Debye screening from salts in the fluid, indicating that our biosensors can be used for clinical diagnostic testing without any special sample preparation or pretreatment. The LoD of our biosensor reached 100 fg/mL, which is 2 orders of magnitude lower than conventional colorimetric ELISA [43, 44] . This could be caused by the local reaction nature of our sensing technique. We expected the local pH changes of the solution surrounding the In 2 O 3 nanoribbon to generate the amperometric signal instead of changing the pH of the entire solution, which would require greater number of phosphatase molecules anchored to the surface of In 2 O 3 and take a prolonged period of time for detection. This hypothesis can be supported by the measurement shown in Fig. S6 in the ESM, as bubbling the solution extensively to mix it thoroughly, the current dropped back to a point close to the baseline, indicating the pH of the entire solution didn't change much. Once the bubbling was stopped, the current increased again as the enzymatic reaction changed the local pH of the solution.

To investigate the reproducibility of our detection, we performed biosensing of 10 pg/mL S1 antigen in different mediums using devices from 4 different batches of fabrication. As shown in Fig. 3(d) , 5 rounds of detection in PBS and 5 rounds of detection in UTM generated similar sensing responses. Hence, it is safe to say that our fabricated biosensors exhibit a remarkable reproducibility for large scale clinical diagnostic testing.

To demonstrate the versatility of our electronic biosensing platform, we applied our In 2 O 3 devices to detect SARS-CoV-2 spike protein IgG antibody. Figure 4 (a) shows schematics of the structure of surfaced functionalized In 2 O 3 nanoribbon for S1 protein specific antibody detection. In this structure, the sequence of molecule binding along with the conjugated antibody and S1 protein serve as the capturing component. After forming this capturing component, target antibody was incubated for 1 h. Biotin conjugated anti-Human IgG antibody was then introduced as the secondary antibody, and then streptavidin-conjugated phosphatase was introduced. Streptavidin would bind to biotin, and thus phosphatase got anchored to our biosensor surface, and the device became ready for electronic biosensing. Finally, amperometric signal was recorded while the device is immersed in liquid gate setup. Figure 4(b) shows the amperometric signals generated as the result of detecting antibodies dissolved in PBS buffer solution. 1, 10, 100 pg/mL and 1 ng/mL concentrations of antibodies generated 18%, 37%, 65%, and 82% of current enhancement (average of 5 trials), respectively (Fig. 4(d) ). It's worth noting that the anti-Human IgG antibody could potentially bind with the primary antibody. Therefore, to investigate the effect of this undesired binding, we performed a negative control experiment, where BSA solution was used to replace the target antibody. Figure 4 (b) shows the result of this negative control (green curve) where a 4% current increase was generated. This indicates that the anti-Human IgG antibody could hardly bind with the primary antibody, which is already anchored to the surface of In 2 O 3 , and hence, would not affect the sensing results.

Furthermore, our biosensors exhibit high selectivity when detecting S1 protein specific antibody in physiological solutions, as demonstrated in Fig. 4(c) . Here, we spiked target antibodies with known concentrations into human whole blood (WB) and carried out the detection. We varied S1 protein specific antibody concentration to include 1, 10, 100 pg/mL and 1 ng/mL, which led to current enhancement (average of 5 trials) of 16%, 29%, 58% and 77%, respectively ( Fig. 4(d) ). The results of the detection in PBS (black squares) and human whole blood (red dots) are plotted in Fig. 4(d) . As depicted in the figure, antibodies spiked in PBS and WB generated analogous responses while the magnitude of sensing responses detecting antibody in blood is slightly lower than that in PBS. These miniscule differences could stem from the high viscosity of the whole blood, where precipitated blood cells can attach to the surface of the biosensor during the incubation period, and thus could induce interference with the sensing results. Nevertheless, we can safely say that the sensing behaviors in both mediums are mainly attributed to the binding of the target antibody.

The results of detecting SARS-CoV-2 S1 antigen in PBS and UTM, and S1 protein specific antibody in PBS and human whole blood reveal that our In 2 O 3 devices and electronic biosensing approach can be highly sensitive, dual-functional biosensors for COVID-19 diagnosis. The ability of our biosensors to switch from antigen detection to antibody detection by simply adding an additional step of antigen incubation and changing the secondary antibody, along with their high immunity to interference caused by ions, cells, and other particles from complex fluids, makes our cost-efficient biosensors desirable for worldwide rapid screening of SARS-COV-2 virus with high sensitivity.

The demonstrated electronic biosensing platform based on phosphatase and shadow mask fabrication process is not limited to In 2 O 3 material and SARS-CoV-2 biosensing. This technology has the potential to be applied to other semiconducting materials such as silicon nanowires [16, 45] , metal oxide nanowires/nanoribbons [46] , carbon nanotubes [47] , graphene [48] , and other two-dimensional materials [49] for biosensing applications. In addition, the electronic biosensing technique based on phosphatase has the potential to be generalized as a standard protocol of electronic biosensors for highly sensitive detection of various proteins with prolonged shelf life.

Since the enzyme and other reagents used in the functionalization of our biosensors are also used for commercial colorimetric ELISA kits, it is possible to integrate our biosensors with conventional colorimetric ELISA for a dual-readout biosensor platform. Such integration would improve the sensing characteristics with high reliability. In this platform, colorimetric signal can be used for quick and preliminary detection, while the electrical signal can subsequently be used for a more sensitive diagnosis.

We have demonstrated sensitive, scalable, and cost-efficient COVID-19 biosensors using electronic biosensing platform based on In 2 O 3 FET devices functionalized with phosphatase. The devices were fabricated by a simple and cheap shadow mask method. Our biosensors were able to detect SARS-CoV-2 spike protein in UTM (LoD: 100 fg/mL) and S1 protein specific IgG antibody in human whole blood (LoD: 1 pg/mL), indicating its potential for clinical diagnostic testing. Our results can be instrumental for the management and control of the current pandemic and can possibly prevent further community transfer through early, rapid, and cost-efficient screening of COVID-19, giving us the upper hand to win the fight against this pandemic.

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We would like to acknowledge the financial support of this research by King Abdul-Aziz City for Science and Technology (KACST) through The Center of Excellence for Nanotechnologies (CEGN).

Supplementary material (materials and methods for device fabrication, functionalization of In 2 O 3 devices, photographs of the liquid gate measurement setup, mobilities of the nine devices labeled in Fig. 1(b) , family curves of I DS −V DS with the liquid gate setup and current change after bubbling the substrate solution (current vs. time curve for S1 antigen detection)) is available in the online version of this article at https://doi.org/10.1007/10.1007/s12274-022-4190-0.