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Abstract
Since the outbreak of coronavirus disease 2019 (COVID-19), efficient real-time monitoring has become one of the challenges faced in SARS-CoV-2 virus detection. A compact all-fiber Mach-Zehnder interferometer optofluidic sensor based on a hollow eccentric core fiber (HECF) for the detection and real-time monitoring of SARS-CoV-2 spike glycoprotein (SARS-CoV-2 S2) is proposed, analyzed and demonstrated. The sensor is comprised of fusion splicing single mode fiber (SMF), hollow core fiber (HCF) and HECF. After the incident light passes through the HCF from the SMF, it uniformly enters the air hole and the suspended micrometer-scale fiber core of the HECF to form a compact all-fiber Mach-Zehnder interferometer (MZI). HECF is side polished to remove part of the cladding that the suspended fiber core can contact the external environment. Subsequently, the mouse anti SARS-CoV-2 S2 antibody is fixed on the surface of the suspended-core for the sake of achieving high sensitivity and specific sensing of SARS-CoV-2 S2. The limit of detection (LOD) of the sensor is 26.8 pM. The proposed sensor has high sensitivity, satisfactory selectivity, and can be fabricated at low cost making it highly suitable for point-of-care testing and high-throughput detection of early stage of COVID-19 infection.
© 2022 Optica Publishing Group under the terms of the Optica Open Access Publishing Agreement
Corrections
3 October 2022: Typographical corrections were made to the author list.
1. Introduction
The outbreak of COVID-19 at the end of 2019 and its continued spread has created an unprecedented health challenge to the global community and poses a significant threat to the lives and health of humans worldwide [1]. Symptoms of the severe acute respiratory syndrome coronavirus 2 (SARS-CoV-2) infection include fever, pneumonia, dyspnea, and even death. Long virus latency leads to the possibility of asymptomatic patients [2]. Therefore, the key to research work lies in early management and the development of several diagnostic methods, such as chest computed tomography (CT) [3], serological immunological assays (ELISA, colloid gold-based immunochromatographic assay, enzyme-linked immunosorbent assay) [4,5], and reverse transcription polymerase chain reaction (RT-PCR) [6,7]. RT-PCR has high sensitivity and specificity, and most large medical institutions are already equipped with relevant equipment to perform such tests. However, RT-PCR requires a fluorescent quantitative RT-PCR instrument, and samples need to be sent to the certified laboratory operated by professionals resulting in total testing time of the order of hours or even days, which is not suitable for resource-poor areas and point-of-care testing (POCT) [7]. More recently, new optical diagnostic sensors have been developed to support the SARS-CoV-2 diagnostic techniques, such as the opto-microfluidic sensing platform with gold nanospikes based on localized surface plasmon resonance (LSPR) [8], semiconductor-based Surface enhanced Raman scattering (SERS) substrates [9], and integrated nanoplasmonic biosensor [10]. Most of the optical platforms for detecting SARS-CoV-2 require precious metals and large-scale instrument preparation, and the spatial light detection system is easily affected by background light. At present, there is an urgent need for a high-sensitive and miniaturized detection platform that are capable of real-time SARS-CoV-2 diagnosis for asymptomatic and mild symptom cases, without sending samples to certified laboratory.
Optical fiber sensing platforms are well suited for biosensing applications, do not suffer from electromagnetic interference, can be fabricated at low cost, and can be remotely operated [11,12]. Biological recognition elements, such as antibodies, biological enzymes, and aptamers can be immobilized on the surface of optical fiber to enable specific biomolecule sensing [13–16]. Traditional optical fiber sensors are based on techniques such as whispering gallery mode (WGM) [17], Fabry-Perot interferometer (FPI) [18], Mach-Zehnder interferometer (MZI) [19], Michelson Interferometer [17,20], fiber grating [21] and surface plasmon resonance (SPR) [22], and are typically fabricated using tapering, etching, side polishing, femtosecond laser micromachining and coating. Fabrication methods such as tapering and etching tend to suffer from vulnerability and damage [23,24], while femtosecond laser micromachining and coating require expensive equipment and complicated processing protocols [25,26]. Recently, there have been many new research results of optical fiber sensors for SARS-CoV-2, such as an U-bent fiber optic probe with gold nanoparticle [27], a fiber sensor with a layer of graphene applied to the Ag-Au NPs alloy film [28], and a plasmonic D-shaped plastic optical fiber with gold nano-film [29]. Sensitivity of the sensors were improved by using noble metal nanoparticles, noble metal layers and two-dimensional materials such as graphene to avoid oxidation of the metal layer, which increase the process difficulty and production cost of sensors.
In addition, it is critical to select the appropriate biomarker for the sensitivity and specificity of the SARS-CoV-2 sensor. SARS-CoV-2 has a particle size in range of 50 to 200 nm and contains at least four structural proteins: spike (S) glycoprotein, small envelope (E) protein, matrix (M) protein, and nucleocapsid (N) protein [30–32]. It is the spike (S) glycoprotein that differentiates SARS-CoV-2 from SARS-CoV and SARS-related CoVs, and it is the key antigen biomarker for direct detection of SARS-CoV-2 [33–34]. Various techniques for detection of SARS-CoV-2 spike glycoprotein (SARS-CoV-2 S2) have been reported, such as electrochemical detection [35], colorimetry [36], field-effect transistor [37] and plasmonic fiber sensor [29].
In this work we report on an optofluidic sensor based on hollow eccentric core fiber (HECF) for SARS-CoV-2 S2 detection, which utilizes a unique microfluidic channel that facilitates integration of the optical fiber and the microfluidic platform. The sensor is fabricated through fusion splicing single mode fibers (SMFs), hollow core fibers (HCFs) and a side-polished HECF. The S2 antibody is immobilized on the sensor surface enabling specific detection of SARS-CoV-2 S2 in solutions at low concentrations. The responses of the sensor to different interfering substances, as phosphate-buffered saline (PBS), bull serum albumin (BSA), human serum albumin (HSA) and rabbit immunoglobulin G (IgG) antigen, were tested, and the results demonstrate that the sensor has sufficient specificity. The response time of the sensor is within 20-30 mins. The proposed sensor has high sensitivity, a compact structure, is simple to fabricate, can be produced as low cost, and has sufficient mechanical strength. The sensor furthermore does not require nucleic acid amplification and reverse transcription, nor does it require transfer of the samples to a laboratory making it suitable for POCT and high-throughput detection of early stage of COVID-19 infection.
2. Materials and methods
2.1. Chemicals
Deionized water (sterile), non-specific binding blocking solution and HAS were purchased from Shanghai Yuanye Bio-Technology Co., Ltd. The non-specific binding blocking solution mainly consists of tris (hydroxymethyl) aminoethane and bovine serum albumin. The antibody diluent was purchased from Beijing Solarbio Science & Technology Co., Ltd; and tween20 from neoFroxx GmbH. PBS, glutaraldehyde, ethanolamine, methyltriethoxysilane (MTES), BSA, 1-(3-(Dimethylamino)propyl)-3-ethylcarbodiimide hydrochloride (EDC), methanol and N-hydroxysuccinimide (NHS) were all procured from Shanghai Aladdin Biochemical Technology Co., Ltd; absolute ethanol, 98% sulfuric acid, hydrogen peroxide (30%) and aminopropyltriethoxysilane (APTES) from Sinopharm Chemical Reagent Co., Ltd; and rabbit IgG from Abcam (Shanghai) Trading Co., Ltd. SARS-CoV-2 S2 and mouse anti SARS-CoV-2 S2 antibody were obtained from the Native Antigen Company.
2.2. Optical sensor
The sensor is an all-fiber MZI consisting of a central HECF (Yangtze Optical Electronics Co., Ltd, China), which is side-polished to expose the hollow core and micrometer-scale fiber core such that they can contact the external environment, and two HCFs (Yangtze Optical Electronics Co., Ltd, China) that connect both ends of the HECF to SMFs (SMF-28, Corning, USA). A computer rendering of the sensor is shown in Fig. 1(a) and optimization is discussed in section 3.1. Figure 1(b) and (c) show micrographs of the sensor before and after side polishing, respectively. The fabrication process steps are shown in Fig. 1(d). First, observe the cross-section of the HECF under a microscope and confirm the direction of the HECF to avoid damage to the fiber core by subsequent side-polishing. In the second step, using an optical fiber fusion splicer (S178A, FITEL, Co., Japan), a piece of the central HECF is spliced with HCFs at both ends (HCF1 and HCF2) and the HCFs are in turn spliced with input and output SMFs (SMF1 and SMF2). The inner and outer diameters of the HCF are approximately 130 µm and 180 µm, respectively, and the ends of both the HECF and SMF extend about 75 µm into the HCF. In the third step, the central HECF is side-polished until the hollow core of the HECF is exposed. The schematic diagram of the side polishing processing system is shown in Fig. 1(e). The fiber fixed on the fiber holder is polished by a polishing wheel, and the polishing process is monitored by a microscope on the side of the fiber. Figure 1(f) and (g) show cross-sectional optical micrographs of the HECF before and after side polishing. The hollow core of the HECF has a diameter of 42 µm, the cladding a diameter of 125 µm, and the fiber core located on HECF inner wall has a diameter of 8.2 µm. It should be noticed that the polishing process just open a slit on the HECF, which does not affect the surface roughness of the effective sensing area of the HECF core.
Fig. 1. (a) Structural diagram of our all-fiber MZI optofluidic sensor based on a hollow eccentric core fiber. Micrographs of the sensor before (b) and after (c) side polishing. (d) The manufacturing processes. (e) The schematic diagram of the side polishing processing system. Cross-sectional optical micrographs of the hollow eccentric core fiber before (f) and after (g) side polishing. (h) Schematic diagram of the beam propagation paths in the sensor. SMF: single mode fiber; HCF: hollow core fiber; HECF: hollow eccentric core fiber.
The side-polished HECF serves as a microfluidic channel in which SARS-CoV-2 S2 antibodies are immobilized on the fiber core to specifically combine with SARS-CoV-2 S2. Figure 1(h) shows the schematic diagram of beam propagation path of the sensor. When the incident light propagates from the input SMF to the HCF1, the light beam rapidly expands due to the mode field mismatch between the interface of SMF and HCF. The expanded light beam further illuminates and splits into two transmission paths in the core and hollow area of HECF respectively. After travelling through the HECF, the two light beams with different paths couple together in the output SMF through the HCF2 due to the same mechanism of beam expanding and coupling, which finally forms a MZI structure. The MZI can be represented by Eq. (1), where $Ih$ and $Ic$ are the intensities of the light in the hollow core and the fiber core, respectively.
where ${\phi _0}$ is the initial phase difference and $\phi$ is the phase difference after the beams have passed through the respective propagation paths:The free spectral range (FSR) of the transmission spectrum of the sensor is:
2.3. Experimental setup
The experimental setup is shown in Fig. 2(a). A broadband source (BBS; DenseLight BX9) with a output wavelength range of 1250–1630 nm, is connected to the end of the sensor while the other end is connected to an optical spectrum analyzer (OSA; AQ6317, YOKOGAWA) that has a minimum wavelength resolution of 0.01 nm. During this measurement, the resolution, wavelength range and refresh time of OSA are 0.1 nm, 1525-1575 nm and about 10 s, respectively. The OSA is connected to a computer that stores and displays the measured spectral information via a GPIB IEEE 488/USB interface. To avoid measurement errors due to stress changes, the sensor is fixed inside a microfluidic channel made of polydimethylsiloxane (PDMS), as shown in Fig. 2(b). PDMS is the most used polymer for microfluidics fabrication due to its advantageous properties, such as excellent transparency and biocompatibility [39]. The channel includes the inlet and outlet to ensure sufficient contact between the sample and the sensor, as well as to prevent the sample from contamination. Each test requires 3 µL of sample solution into the 15 mm×1 mm×0.2 mm microfluidic channel.
Fig. 2. (a) Schematic diagram of the experimental setup. (b) Multi-channel microfluidic chip mount with the fiber optofluidic sensor. OSA: optical spectrum analyzer; BBS: broadband source.
3. Results and discussion
3.1. Simulation and optimization
The transmission spectrum and light intensity distribution of the sensor with and without the HCF were simulated by using commercial software RSoft and shown in Fig. 3 (a) and (b). Comparing the light intensity distributions in Fig. 3(a) and (b), one sees that the averaged value of light intensity in the fiber core of the HECF is much closer to that in the hollow core in the case with the HCF as compared to the case without HCF. As expected, such better balanced averaged values of light intensity in the two light paths of the interference structure result in a higher contrast ratio of transmission spectrum and a more uniform interference peak distribution. Transmission spectra of the sensors with different lengths of the HCF (50 µm, 100 µm, 350 µm, 500 µm and 750 µm) are shown in Fig. 3(c), and performance comparisons are listed in Table 1. When the length of the HCF is range from 350 µm to 500 µm, the contrast ratio of the transmission spectrum is large enough, the transmission loss is acceptable, and the interference peaks distribution is more uniform. Figure 3(d)-(f) show the transmission spectrum and the wavelength shift of the sensor at different HECF lengths with three magnitude orders (200 µm, 1.6 mm, and 1 cm), and performance comparisons are listed in Table 2. With increasing length of the HECF, the sensitivity of the sensor increases, and the FSR decreases, while resulting in a decrease in the dynamic range of the sensor. Furthermore, a too long HECF results in large energy attenuation in the air hole that impacts the quality of double-beam interference. It should be noticed that the RI sensitivity of surface nanolayer sensing is almost equal or even higher than volume RI sensing case according to earlier literature [40]. Considering the sensitivity, dynamic range and spectral quality, the optimal sensor design with an HCF length ranges from 350 µm to 500 µm and an HECF length of near 1.6 mm are chosen in this experiment.
Fig. 3. Simulated transmission spectrum of, and optical intensity distribution in, the HECF-based fiber optofluidic sensor with (a) and without (b) an HCF connector (the HCF length is 500 µm). (c) Simulated transmission spectrum with different lengths of the HCF. Simulated transmission spectrum and wavelength shift with varying lengths of the HECF: (d) 200 µm, (e) 1.6 mm, and (f) 1 cm.
Table 1. Performance comparisons of the HCF with different lengths
Table 2. Performance comparisons of the HECF with different lengths
Instead of costly femtosecond laser or focused iron beam machining, a slit on one side of the HECF cladding is fabricated simply by side polishing. Measured transmission spectra of the sensor and their fast Fourier transform (FFT) in air and in water before and after side-polishing the HECF are shown in Fig. 4(a) and (b). It can be clearly seen that the proposed sensor is insensitive to the surrounding RI before side-polishing HECF. After side polishing, according to Eq. (3) in Section 2.2, the different RI values of water (n = 1.333) and air (n = 1) cause the FSR of the transmission spectrum to change, which is more clearly shown in the FFT of the spectrum, proving that the sensor is sensitive to surrounding RI. Figure 4(c) and (d) show the measured transmission spectra and a linear fit of the resonant wavelengths corresponding to different surrounding RI values, which are consistent with the simulation results. Solutions of different RI values were prepared in methanol and deionized water and measured with the Abbe Refractometer (WAY-2WAJ, Zhejiang LICHEN Instrument Technology Co., Ltd). The concentration range of methanol (mass %) is 0-3.42%. The data was post-processed using a locally weighted scatterplot smoother (LOWESS), and the wavelength of the dip is located by direct searching location of the minimum value. The RI sensitivity and LOD of the sensor are calculated to be -14250 nm/RIU and 7.02 × 10−7 RIU, respectively. The sensitivity difference between experimental and simulated results may be caused by the deviation of practical material size and properties coefficient, such as fiber length, refractive index of fiber core etc.
Fig. 4. (a) Measured transmission spectra and (b) their fast Fourier transform (FFT) in air and water, before and after side-polishing of the HECF. (c) Measured transmission spectra and (d) the corresponding dip wavelengths as the surrounding RI varies.
3.2. Functionalization of the sensor surface
Prior to SARS-CoV-2 S2 measurement, the bare optical sensor surface needs to be functionalized [41,42]. The relevant chemical reagents used in this functionalization process are prepared as followed: Piranha solution used is a mixture of 98% sulfuric acid and hydrogen peroxide with a volume ratio of 7:3. APTES:MTES (1:30) solution is composed of APTES and MTES mixed in alcohol with a mol ratio of 1:30. EDC/NHS solution is composed of 16 mg EDC and 23 mg NHS in 50 ml PBS. The concentration of S2 antibody solution is 0.1 mg/mL, and the solvent is an antibody diluent. To activate the S2 antibody, an EDC/NHS solution at a ratio of 100:3 was added and ultrasonically stirred for 1 h. The non-specific binding blocking solution is supplemented with 1 M ethanolamine.
The specific processing steps are shown in Fig. 5(a). (1) cleaning with piranha solution for 30 mins to form hydroxyl group; (2) modification with APTES:MTES (1:30) solution for 1 h to form amide group; (3) reaction of 2.5% glutaraldehyde for 2 h to form aldehyde group; (4) immobilization with S2 antibody solution for 1 h; (5) treatment of non-specific blocking solution for 30 mins. Following each step, the sensor is washed with PBS for 30 mins to remove unreacted molecules remaining on the surface. Scanning electron microscopy (SEM; SU8010, HITACHI) is furthered used to confirm each step of the functionalization, the results of which is shown in Fig. 5(b), where differences in surface roughness and morphology can be observed, demonstrating the effectiveness of each step. The surface modified with MTES is medium hydrophobic, which reduces non-specific adsorption [40].
Fig. 5. (a) Functionalization process steps, (b) SEM images at each step of the sensor surface functionalization, including: (1) cleaning with piranha solution; (2) modification with APTES:MTES (1:30) solution; (3) reaction of 2.5% glutaraldehyde; (4) immobilization with S2 antibody solution and (5) treatment of non-specific blocking solution.
Figure 6(a) and (b) show evolution of the transmission spectrum and wavelength shift during the sensor surface functionalization at room temperature. It can be observed that the RI of the fiber core surface increases and the transmission spectrum in air blueshifts, as compared with the bare optical sensor, after the third step of aldehyde formation. After immobilization of the S2 antibody, the transmission spectrum blueshifts 8.2 nm when the sensor is immersed in PBS, compared to prior immobilization. This means that the RI of the fiber core surface increases, i.e., S2 antibody is successfully immobilized, consistent with what is seen in Fig. 5(b) and (c). In Fig. 6(b) we can see that, following the immobilization of S2 antibody solution for about 60 mins, the drift rate of the transmission spectrum tends to flat, indicating that the immobilization is completed. In practice, the S2 antibody may not completely fill the fiber core surface because of the incomplete chemical reaction. Impurities other than SARS-CoV-2 S2 will be adsorbed on the blank positions on the fiber core surface. Therefore, BSA in non-specific binding blocking solution is used to fill the blank positions to achieve non-specific blocking. Non-specific blocking is completed in about 10 mins, and the transmission spectrum blueshifts 2.5 nm, demonstrating that most of the binding molecules on the fiber core surface are effective S2 antibodies. Figure 6(c) and (d) show evolution of the transmission spectrum during immobilization of S2 antibody and treatment of non-specific blocking solution. The slight red shift of the transmission spectrum following PBS washing is attributed to the removal of molecules that are not firmly immobilized on the sensor surface during PBS washing. It should be noticed that as the splicing junction of the HECF and the HCF does not collapse, the solutions can flow into both the HECF and the HCF. Therefore, functionalization will take place both on the end-face of SMF, the inner surface of the HECF and the HCF. However, the HCF just acts as a beam splitter in the hollow core area, and the effective sensing area is the HECF core surface. Functionalization of the inner surface of the HCF and end-face of SMF maybe have a slight influence on the transmissivity of interface between SMF and HCF, resulting in a slight change in the power and contrast ratio of transmission spectrum of the sensor.
Fig. 6. (a) Evolution of the transmission spectra and (b) the shifted dip wavelengths during the sensor surface functionalization. Evolution of the transmission spectra during (c) immobilization of S2 antibody and (d) treatment of non-specific blocking solution.
3.3. Detection of SARS-CoV-2 spike glycoprotein S2
Following functionalization, the performance of the sensor for detection of SARS-CoV-2 S2 was tested. At room temperature 25 °C, SARS-CoV-2 S2 solutions with concentrations of 10 ng/ml, 100 ng/ml, 1 µg/ml, and 10 µg/ml were injected into the microfluidic channel in turn. Figure 7(a) shows the wavelength shift due to the varying SARS-CoV-2 S2 concentrations in the solutions. The SARS-CoV-2 S2 molecules fully reacted with the immobilized S2 antibodies, reaching a dynamic equilibrium within 20-30 mins. The variation and rate of the wavelength shift are proportional to the logarithm of the SARS-CoV-2 S2 concentration in the solution. Following each SARS-CoV-2 S2 concentration measurement, the sensor was washed with PBS for 20 mins to remove unreacted and unstable molecules from the sensor surface, which causes redshift by about 0.2-0.5 nm. After the complete reaction of the SARS-CoV-2 solution, which lasted for 40 mins, the wavelength shift was recorded. Figure 7(b) and (c) show the measured wavelength shift against the logarithmic and normal concentration of SARS-CoV-2 S2, respectively. The error bars correspond to the standard deviation obtained by three continuous independent measurements. The linear regression equation with 95% confidence bounds is Δλ(nm) = 0.8-1.09×log(C(ng/ml)), and R2 is 98.88%. The sensitivity of the sensor was calculated by taking the derivative of the calibration curve fitted to the data in Fig. 7(c), which is Δλ(nm)= -2.506×(C(ng/ml)^0.1026) + 2.76. The sensitivity of the sensor decreases with the concentration of SARS-CoV-2 S2 solution. This phenomenon maybe caused by the nonlinear binding process of SARS-CoV-2 S2 and S2 antibody, namely, the subsequent bonding reaction will be impeded by the molecules that had been bonded. At 10 ng/ml of SARS-CoV-2 S2 solution, the lowest concentration to calibrate the sensor, the sensitivity of the sensor is 26.9 pm/(ng/mL). Figure 7(d) shows the wavelength stability obtained by the sensor in PBS during 180 mins. The sampling interval is 5 mins. The standard deviation of the wavelength shift is σ=0.03 nm, and 3σ=0.09 nm. According to the International Union of Pure and Applied Chemistry [43,44], the estimated LOD = 3σ/ sensitivity = 3.35 ng/ml =26.8 pM.
Fig. 7. Experimental data of SARS-CoV-2 S2 detection at 25 °C by the HCEF-based optofluidic sensor: (a) the dip wavelengths; (b) measured wavelength shifts as the concentration varies in logarithmic coordinate and (c) in normal coordinate; (d) stability of the dip wavelength, (e) measured transmission spectra and (f) the bar chart of selectivity experiment from 0 to 10 ug/ml.
In addition, the specificity of the sensor for SARS-CoV-2 S2, PBS, BSA, HAS and rabbit IgG solutions with the same concentrations of 10 µg/ml were also investigated, the results of which are shown in Fig. 7(e) and (f). It can be clearly seen that the only significant wavelength shift occurs for the SARS-CoV-2 S2 solution, indicating that the sensor is highly selective for SARS-CoV-2 S2 and is thus suitable for SARS-CoV-2 determination. Variation of the contrast ratio of the transmission spectrum shown in Fig. 7(e) due to unstable fluctuations in the output power and spot uniformity of the light source over time.
All experiments are carried out in a constant temperature laboratory. The length of the HECF is a constant, according to Eq. (2) in section 2.2, the response of the sensor to temperature can be represented as:
Performance comparisons with previously reported SARS-CoV-2 S2 concentration sensors are listed in Table 3. Compared with a TiO2 nanotube-based electrochemical sensor [46],a plasmonic D-shaped plastic optical fiber [29], semiconductor-based SERS sensor [9] and nanoplasmonic biosensor [10], our proposed sensor shows significant improvement in performance with better sensitivity, LOD, commercial potential and lower cost.
Table 3. Performance comparison with previously reported SARS-CoV-2 S2 concentration sensors
4. Conclusion
In this work we have presented a highly sensitive fiber optofluidic sensor based on HECF that is suitable for SARS-CoV-2 S2 detection. The sensor design is based on an MZI structure and is fabricated by fusion splicing of SMF, HCF and HECF. A slit on one side of the HECF cladding is fabricated simply by side polishing, which serves as a microfluidic channel. Results of sensor functionalization were confirmed by SEM, as was real-time measurement by the optical sensor platform. The LOD of the sensor is 26.8 pM, which meets the requirements for SARS-CoV-2 early diagnosis. The response of the sensor to different disruptors furthermore confirms that the sensor has the desirable specificity. The area of contact between the micrometer-scale fiber core of the sensor and the external environment is sufficiently large to yield a large phase difference between the modes of the hollow core and fiber core, which is what allows for the sensor’s high sensitivity while being very compact in size. Excellent performance results suggest that the proposed fiber SARS-CoV-2 sensor is more suitable for high-throughput COVID-19 detection and POCT applications.
Funding
Key Research and Development Program of Zhejiang Province (2021C03178); Ningbo Science and Technology Project (2021Z030); Ningbo Science and Technology Plan Project-Key Core Technology Emergency Tackling Plan Project (2020G012); National Natural Science Foundation of China (11621101); China Postdoctoral Science Foundation (2018M642423); National Key Research and Development Program of China (2018YFC1407503); Key R & D Plan of Zhejiang Province (2019C03089); Shanghai Zhangjiang Science City Special Development Fund.
Disclosures
The authors declare that they have no known competing for financial interests or personal relationships that could have appeared to influence the work reported in this paper.
Data availability
No data were generated or analyzed in the presented research.
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