*3.1. SARS-CoV-2*

The SWCNT-based optical sensing approach was employed by Pinals et al. [16] to detect the SARS-CoV-2 spike protein. The angiotensin-converting enzyme 2 (ACE2), which has a strong affinity with the SARS-CoV-2 spike protein, was used to functionalize SWCNTs. A 2-fold increase in fluorescence was observed within 90 min of SARS-CoV-2 spike protein exposure, which exhibited a limit of detection (LOD) of 12.6 nM for the device. Shao et al. [17] functionalized a high-purity semiconducting SWCNT surface with an anti-SARS-CoV-2 spike protein antibody and used it as a channel in an FET-type biosensor, to detect SARS-CoV-2 antigens in clinical nasopharyngeal samples. The fabricated device exhibited an LOD of 0.55 fg/mL (Figure 5). For efficient and precise identification of the SARS-CoV-2 S1 antigens in fortified saliva samples, Zamzami et al. [18] created a rapid, simple-to-use, inexpensive, and quantitative CNT-based antibody-functionalized p-type depletion FET biosensor. Through a non-covalent interaction with the linker 1-pyrenebutanoic acid succinimidyl ester (PBASE), the SARS-CoV-2 S1 antibody was bound to the CNT surface in the FET channel region. The CNT-FET biosensor successfully identified the SARS-CoV-2 S1 antigen in 10 mM AA buffer at a pH of 6.0, at concentrations ranging from 0.1 fg/mL to 5.0 pg/mL, with an LOD of 4.12 fg/mL. On a flexible Kapton substrate, Thanihaichelvan et al. [19] created CNT-FETs and immobilized the reverse sequence of the SARS-CoV-2

RNA-dependent RNA polymerase gene onto the CNT channel, to develop a biosensor for coronavirus detection with an LOD of 10 fM. The primary signal generation depended on RNA hybridization, while the main signal transducer was a liquid-gated CNT-FET. Cardoso et al. [20] used carboxylated CNTs and screen-printed (SP) them onto carbon electrodes (CNT-SPE). Later, the electrode was modified with EDC/NHS coupling chemistry to produce an amine layer, to adsorb the SARS-CoV-2 spike protein antibody. The sensor exhibited a linear response between 1.0 pg/mL and 10 ng/mL, with an LOD of ∼0.7 pg/mL. A platform for electrochemical sensing was fabricated by Curti et al. [21] employing SWCNT-SPEs, which were functionalized with a DNA aptamer that was already redox-tagged. In the presence of the SARS-CoV-2 spike protein S1, the concentration-dependent folding of a DNA aptamer occurred, which resulted in a change in the amperometric signal with an LOD of 7 nM. Monoclonal antibodies were employed for the functionalization of SWCNT by Li et al. [22], to develop an extremely sensitive immuno-resistive sensor for the detection of SARS-CoV-2. To minimize the contact resistance, silver electrodes were screen-printed onto SWCNTs, and the complete arrangement was mounted on polyethylene terephthalate (PET) film. The LOD of the developed sensor was 350 genome equivalents/mL. Through systematic analysis, Kim et al. [23] examined the relationships between different thin-film characteristics and the sensitivity of CNT thin-film-based immunosensors for the rapid detection of the SARS-CoV-2 virus. They found that smaller surface roughness and better CNT alignment resulted in improved sensitivity at a given film thickness, with a LOD value of 5.62 fg/mL, because of the enhanced bioreceptor-binding surface area. Table 1 summarizes the earlier discussion of the creation of CNT-based biosensors to identify the SARS-CoV-2 virus.

**Figure 5.** The detection of SARS-CoV-2 Ag, using SWCNT-based FET biosensors. (**a**) Schematic structure of SARS-CoV-2 to demonstrate the targeting proteins. (**b**) Schematic illustration of a liquidgated SWCNT FET for the detection of SARS-CoV-2 SAg and NAg. Interdigitated gold electrodes (yellow blocks) are configured as the source (So) and drain (Dr) (Reprinted with permission from Ref. [17]).

#### *3.2. Dengue Virus (DNV)*

Based on CNT-SPE, an immunosensor for non-structural protein 1 (NS1) of the DNV was created by Dias et al. [24]. They employed a uniform mixture of carboxylated CNTs distributed in carbon ink to make the CNT-SPE and used an ethylenediamine film to covalently attach anti-NS1 antibodies to CNT-SPE. The developed biosensor was able to detect the DNV NS1 protein, with an LOD of 12 ng/mL. A robust poly (allylamine) (PAH) sandwich-based immunosensor was fabricated by Silva et al. [25] for detecting DNV NS1. A thin coating of PAH on carboxylated CNTs helped to immobilize anti-NS1 antibodies on the electrode surface. To strongly bond CNTs to the electrode surface, as well as anti-NS1 antibodies, through their Fc terminus to prevent random immobilization, PAH, a cationic polymer, was used. The fabricated immunosensor had a linear range of operation between 0.1 g/mL and 2.5 g/mL, with an LOD of 0.035 g/mL. An SWCNT-based, inexpensive, label-free chemiresistive biosensor was designed by Wasik et al. [26], where, for the first time, heparin was utilized as a biorecognition component as opposed to a conventional antibody. The biosensor revealed clinically significant sensitivity for people infected by *Aedes aegypti*, concerning the detection of whole DNV, with an LOD of 8 DNV/chip. Later,

they modified the biosensor by employing a network of anti-dengue NS1 monoclonal antibodies for the functionalization of SWCNTs, instead of using herpin [27]. The modified biosensor exhibited a linear response in the range of 0.03–1200 ng/mL, with an LOD of 0.09 ng/mL. Almost all laboratory and commercial dengue NS1 diagnostic measures include a blood-drawing procedure, which limits the advantage of point-of-care (POC) diagnostics and reduces patient readiness. Instead of blood, NS1 can be extracted from human saliva for the early detection of dengue infection, which is a straightforward, non-invasive, pain-free, and economical process that can be performed by even untrained/less-trained workers. This pathway was also explored by Wasik et al. [28] through the fabrication of a label-free chemiresistive immunosensor employing a network of anti-dengue NS1 monoclonal antibody-functionalized SWCNTs. The biosensor exhibited a detection range of ~1 ng/mL to 1000 ng/mL for the DNV NS1 protein (Figure 6). To detect DNV antibodies, a high-performance impedimetric immunosensor was developed by Palomar et al. [29] via the deposition of CNT on electrodes and was later functionalized with polypyrrole-NHS to immobilize the DNV 2 NS1 glycoprotein via covalent amide coupling. This biosensor exhibited high linearity after optimization, in a broad range of concentrations (10−<sup>13</sup> to 10−<sup>5</sup> g/mL). For the detection of dengue toxin, the CNT/Au nanoparticle (AuNP) composite was deposited on a homemade Au electrode by Palomar et al. [30], to immobilize dengue antibodies on AuNPs through covalent bonding. The electrochemical signal enhancement and improvement in overall performance were achieved due to the porosity of the tridimensional network of Au-CNT, with an LOD of <sup>3</sup> × <sup>10</sup>−<sup>13</sup> g/mL. Mendonça et al. [31] fabricated a label-free immunosensor employing a thin film of CNT-ethylenediamine. The covalent immobilization of the anti-NS1 monoclonal antibodies on CNTs allowed for great measurement stability. Finally, differential pulse voltammetry (DPV) was used to analyze the responses to dengue NS1. The measurement showed that the linear range for the operation of the biosensor was 20 to 800 ng/mL, with an LOD value of 6.8 ng/mL. The preceding discussion on the development of CNT-based biosensors for DNV detection is summarized in Table 1.

**Figure 6.** Schematic of the functionalization of SWNT networks with the Dengue Virus (DNV) NS1 antibody (Reprinted with permission from Ref. [28]).

#### *3.3. Influenza Virus*

An MWCNT-cobalt phthalocyanine nanocomposite and poly (amidoamine) (PAMAM) dendrimer were deposited on a glassy carbon electrode (GCE) by Zhu et al. [32] to electrochemically detect the avian influenza virus (AIV) genotype in a label-free format. The DNA probes were, then, effectively immobilized on the modified electrode using the coupling agent, G4 PAMAM dendrimer, and were monitored with DPV to achieve an LOD of 1.0 pg/mL. The surface of MWCNTs was utilized by Tam et al. [33] to immobilize the DNA probe by covalent interactions between the DNA sequence's amine and phosphate groups. Changes in conductance of the sensor surface were used to detect the hybridization of the DNA probe and the influenza viral DNA for the label-free detection of the influenza virus with an LOD of 0.5 nM. CNT-SPEs were used by Bonanni et al. [34] as an electrochemical-sensing platform for the detection of influenza viral DNA. The CNT-SPEs were functionalized with carboxylic molecules to covalently immobilize the oligonucleotide probe, employing facile carbodiimide chemistry. A direct coupling and sandwich scheme, two distinct techniques for the impedimetric detection of DNA hybridization, were employed and compared. The sandwich scheme revealed better results, with an LOD value of 7.5 fM. A self-assembled SWCNT thin film was prepared by Lee et al. [35] for the fabrication of an affordable, label-free biosensor to detect the swine influenza virus (SIV). The anti-SIV antibodies were bounded with the SWCNT thin film for the detection of SIV, with an LOD of 180 TCID50/mL, by monitoring the change in resistance (up to 12%). Dielectrophoretic and electrostatic forces were used to deposit COOH-functionalized SWCNTs on a selfassembled monolayer of the polyelectrolyte, polydiallyl dimethyl-ammonium chloride (PDDA), by Singh et al. [36]. Viral antibodies were immobilized, utilizing biotin-avidin coupling, after avidin was coated on the PDDA-SWCNT channels (Figure 7). Changes in the channels' resistance were monitored to detect influenza viruses, with a detection limit of 1 PFU/mL. AuNP-decorated CNTs (Au-CNTs) were produced, utilizing phytochemical composites at room temperature, and were employed by Lee et al. to build a plasmon-assisted fluoro-immunoassay (PAFI) for the detection of the influenza virus [37]. The surfaces of Au-CNTs and CdTe quantum dots (CdTe-QDs), the photoluminescence intensity of which varied according to viral concentration, were conjugated with specific antibodies against the influenza virus to achieve an LOD of 1 ng/mL (for Beijing/262/95 (H1N1)), 0.1 pg/mL (for New Caledonia/20/99IvR116 (H1N1)), and 50 PFU/mL (for Yokohama/110/2009 (H3N2)). Later, a two-step method was adopted by Lee et al. [38] to decorate CNTs with Au/magnetic nanoparticles, which offered superior magnetic properties and high electrical conductivity. Later, the Au nanoparticle's surface was conjugated with thiol-group-functionalized probe DNA to detect influenza viral DNA with a detection limit of approximately 8.4 pM. AuNPs were bonded to the CNT surface via in situ accumulation under mild conditions by Ahmed et al. [39]. The improved peroxidase-like activity of the Au-CNT nanohybrid was employed to develop a supersensitive colorimetric optical sensor to detect the influenza virus. The sensor comprised a test system containing specific influenza antibodies, Au-CNT nanohybrids, 3, 3 , 5, 5 -tetramethyl-benzidine (TMB), and H2O2. The color of the test system turned blue upon the addition of the influenza virus and the LOD was 3.4 PFU/mL. Fu et al. [40] fabricated chemiresistor-type biosensors to detect AIV, employing semiconducting SWCNTs or nitrogen-doped MWCNTs and noncovalently functionalizing them with DNA probe sequences. Complementary DNA target sequences of AIV, with concentrations ranging from 2 pM to 2 nM, could be detected by the constructed biosensor after 15 min at room temperature. To detect influenza type A viral DNA, a CNT-FETs-based DNA sensor was developed by Tran et al. [41]. The initial probe DNA, the hybridization period, and the reaction temperature were some of the aspects that were examined since they affected the sensing data. The DNA sensor demonstrated a quick response time of under a minute, with a very low (1 pM) detection limit and a broad linear detection range of 1 pM to 10 nM. Wang et al. [42] compared the effectiveness of the aptamer and antibody, concerning the detection of influenza A virus (California/07/2009 (pdmH1N1)), employing an MWCNT-Au conjugated sensing surface with a di-electrode. They found that the electric response and affinity of aptamers were much stronger than those of antibodies. Wang et al. also reported that the aptamer could provide an LOD in the range of 10 fM, whereas the antibody showed an LOD of 1 pM in detecting the influenza A virus. Using flexible and stable SP-CNT-polydimethylsiloxane electrodes, a paper-based immunosensor for the label-free hemagglutinin antigen (HA) detection of several AIVs (H5N1, H7N9, and H9N2) was demonstrated by Lee et al. [43], whereby immune responses were measured via DPV. The LOD values for the different viruses were 55.7 pg/mL for H5N1 HA, 99.6 pg/mL for H7N9 HA, and 54.0 pg/mL for H9N2 HA. In Table 1, an overview of the CNT-based biosensors for influenza detection is provided.

**Figure 7.** Schematic illustration of the SWCNT-based immunosensor system for H1N1 virus detection. The inset shows an optical image of di-electrophoretically deposited SWCNTs on a PDDA selfassembled monolayer (Reprinted with permission from Ref. [36]).

#### *3.4. Human Immunodeficiency Virus (HIV)*

Mahmoud et al. [44] immobilized thiol-terminated ferrocene-pepstatin (ThFcP) conjugate on an SWCNT/AuNP-modified Au electrode. Electrochemical impedance spectroscopy (EIS) was used to track the nature of the interaction between HIV-1 protease and the ThFcP conjugate, which resulted in the change of the interfacial characteristics of Au electrodes. The electrochemical biosensor was able to detect HIV-1 protease, even at a 10 pM level. Later, thiolated SWCNT/AuNPs were used to modify the disposable SP-Au electrode surface, while ThFcP was subsequently self-assembled on those surfaces by Mahmoud et al. [45] to fabricate a sensitive electrochemical biosensor to detect HIV-1 protease. A nanocomposite of AuNP, amino-functionalized MWCNT, and acetone-extracted propolis (AEP) was prepared by Kheiri et al. [46] and deposited in the same way on an Au electrode for immobilization of the p24 antibody (anti-p24 Ab) to create an immunosensor for the detection of the HIV antigen. The immunosensor demonstrated high electrochemical sensitivity in detecting p24 in a range of concentration from 0.01 to 60.00 ng/mL, with an LOD of 0.0064 ng/mL. An advanced molecularly imprinted polymers (MIPs) electrochemical sensor was fabricated on MWCNT-modified GCE by Ma et al. [47] through the polymerization of the surface, using acrylamide (AAM), N,N -methylene bisacrylamide (MBA), and ammonium persulphate (APS) as the functional monomer, cross-linking agent, and initiator, respectively. The developed sensor was capable of detecting HIV-p24 in human

serum samples, with an LOD of 0.083 pg/cm3. A chitosan/glutaraldehyde crosslinking system was employed by Giannetto et al. [48] to immobilize the target protein on disposable CNT-SPE for the maximum exposure of p24, to interact with a mouse anti-p24 IgG1. The linear operating range of the immunosensor was 10 pM to 1 nM, with an LOD of 2 pM, for HIV-related p24 capsid protein in human serum. Harvey et al. [49] observed that denatured proteins can improve the optical responsiveness of CNTs to nucleic acids. Their study revealed that following hybridization, hydrophobic regions of the denatured protein interact with the surface of the CNTs, which results in a larger shift in the nanotube emission. They later employed this strategy for the detection of intact HIV in serum. A nickel-organic composite/AuNP/CNT/polyvinyl alcohol (PVA) substance was used to fabricate a flexible paper-based electrode by Lu et al. [50] for the detection of HIV DNA (Figure 8). The methylene blue was employed as a redox indicator for DNA hybridization on the electrode. The large surface area and the presence of π-electron, donated by the Ni-Au composite, facilitate the higher loading of target DNA. With a linear range of 10 nM–1 μM and an LOD of 0.13 nM, this flexible paper electrode demonstrated good sensing capability. The previous discussion on the creation of CNT-based biosensors to identify HIV is summarized in Table 1.

**Figure 8.** Schematic of the fabrication process for a flexible Ni-Au composite/CNT/PVA film electrode and the detection of the target Human Immunodeficiency Virus (HIV) DNA (Reprinted with permission from Ref. [50].

#### *3.5. Hepatitis Virus*

To detect the short DNA sequences associated with the Hepatitis B virus (HBV), a label-free electrochemical DNA biosensor was developed by Li et al. [51], employing 4,4 diaminoazobenzene (4,4 -DAAB) and MWCNT-modified GCE. The CNT carboxyl groups were covalently linked to the oligonucleotides and the DPV was used to monitor the hybridization reaction, with an LOD of 1.1 × <sup>10</sup>−<sup>8</sup> M. To detect HBV, Oh et al. [52] created an FET-based biosensor, using CNTs consisting of a microfluidic channel with immobilized hepatitis B antibody on it. The electrical conductance changed over time, owing to the presence of the hepatitis B antigen in the channel. The change in the channel conductance was proportional to the hepatitis B antigen concentration. Ly et al. [53] immobilized bovine IgG, employing cyclic voltammetry on a DNA-linked CNT electrode to fabricate an electrochemical biosensor for the detection of human HBV in non-treated blood. The relative standard deviation of 0.2 mL HBV was 0.04 (*n* = 4) within the working limits of 0.035–0.242 mg/mL

anti-bovine IgG. A CNT-conducting polymer (CP) network was created by Hu et al. [54] via drop-casting a CNT solution on a GCE, followed by the electrochemical polymerization of a poly (pyrrole propionic acid) (pPPA) film for crosslinking and stabilizing the CNTs. The CNTs served as the network's structural foundation and provided excellent specific surface areas for immobilizing antibodies. Moreover, owing to its self-limiting growth characteristic, the conducting film facilitated CNT in forming a stable network and offered ample carboxyl groups to immobilize the probe proteins for the detection of hepatitis B surface antigen in serum, with an LOD of 0.01 ng/mL. Amino-CNT and hyaluronic acid (HA) were bonded with amide groups and assembled onto the surface of GCE by Cabral et al. [55]; the response of the electrode in the presence of hepatitis B core protein antibodies was measured by square-wave voltammetry (SWV). The immunosensor response was linear up to 6.0 ng/mL, with an LOD of 0.03 ng/mL. For the detection of the core hepatitis B antigen, an electrochemical immunosensor based on polytyramine (PTy)-CNT composite was developed by Trindade et al. [56]. Because of the substantial creation of NH3+ ionic species, the composite possesses high catalytic activity. The HBV was electrochemically identified by SWV in a label-free and reagent-free manner. The immunosensor exhibited a LOD of 0.89 ng/mL while operating in a linear range of 1.0 to 5.0 ng/mL. An AuNPs/chitosan-ferrocene-ammoniated MWCNT (CS-Fc-AMWNT) nanocomposite was prepared by Chen et al. [57] via the Schiff base reaction to achieve a large specific surface area, adequate conductivity, and exceptional biocompatibility. The electrochemical deposition was used to modify the AuNPs for the screen-printed electrode (SPE), while physical adsorption was used to adhere the CS-Fc-AMWNTs composite to the electrode surface. Later, through glutaraldehyde cross-linking, hepatitis B surface antibodies were immobilized on the surface of the electrode. The biosensor can be operated in the range of 1–250 ng/mL for the detection of hepatitis B antigen, with an LOD of 0.26 ng/mL. The SP carbon electrode was modified by Upan et al. [58], through the addition of CNTs that were embellished with AuNP and AgNP (Figure 9). The AuNPs offered biocompatibility and a wide surface area for the immobilization of the hepatitis B surface antibody, which aided the signal improvement. Subsequently, in DPV detection, AgNPs served as a sensing probe to detect the target antigen in the linear range of 1–40 ng/mL with an LOD of 0.86 ng/mL. Using a peptide nucleic acid-functionalized SWCNT-FET biosensor, Dastagir et al. [59] also exhibited the explicit and label-free detection of a hepatitis C virus RNA sequence, with a detection limit of 0.5 pM. The deposition precipitation approach was used by Pusomjit et al. [60] to create Pt nanoparticles that were later coated onto SWCNTs. The generated nanocomposite was finally used as a substrate to immobilize antibodies on a paper-based, SP graphene electrode surface for the purposes of analyzing the hepatitis C virus. DPV was used to measure the target antigen in the range of 0.05 to 1000 pg/mL, with an LOD value of 0.015 pg/mL. Table 1 summarises the preceding discussion on the development of CNT-based biosensors and their use to detect several types of human viruses HBV.

**Figure 9.** The stepwise fabrication process of the immunosensor for the detection of Hepatitis B virus (HBV) (Reprinted with permission from Ref. [58].


**Table 1.** CNT-based biosensors for human virus detection.


**Table 1.** *Cont.*

#### **4. Current Challenges and Future Perspectives**

Carbon nanomaterials are frequently employed in biomedical applications, due to their multifunctionality and minimal complexity in surface modification, which, in turn, improves their biophysical characteristics. CNT is the sp2 hybridized allotrope of carbon, with a hollow cylindrical tubular structure, and has a high aspect ratio. The efficiency and precision of detection in sensing viral genomes, proteins, and other viral cellular biological components using CNT are governed by the exotic characteristics of the nanotubes. CNTs differ from conventional nanomaterials in that they possess many novel physiochemical properties that hold great promise for a variety of applications, including biosensing. It is envisaged that the utilization of CNT's special capabilities in a biological setting would lead to significant improvements in disease diagnosis, monitoring, and treatment. Consequently, CNT-based biosensors offer many benefits over other types of sensors, such as those based on metal oxides or silicon, including high sensitivity, quick response times, reduced redox reaction potential, and longer lifetimes with greater stability. Although CNTs have many desirable qualities and benefits, dispersion, which is triggered by the high surface energy of the CNTs, stands in the way of moving forward. Because of their extreme hydrophobicity, CNTs cannot be dissolved in water or other common solvents. In order to increase their solubility and other functional qualities, CNTs must be functionalized, depending on the application. The benefits and drawbacks of employing CNTs in biological applications are listed in Table 2.


**Table 2.** Pros and cons of using CNTs for biomedical applications. (Reprinted with permission from Ref. [61].

To establish a faithful nano-bio interface, CNTs may be engineered by means of covalent or non-covalent modification, EDC/NHS chemistry, click chemistry, and lengthlocation tuning. Numerous viruses can go into a dormant stage called latency, wherein they remain inactive inside the host cell before activation [62–66]. Often, it is necessary to identify the virus, even in the latent stage, to eliminate the probability of infection/reinfection. Thus, the pathways may be searched through the functionalization or surface modification of CNTs, to enable CNTs to detect the virus in both the dominant and latent stages. Although the great sensitivity and extended durability of CNTs-based viral biosensors make them a promising candidate for viral detection, their accessibility and the economic perspective are essential for rapid diagnosis. The supervision of economic feasibility is an important component of commercialization. The manufacturing complexity levels should be kept to a minimum from the beginning of the design process, by means of the selection of facile fabrication methods that can be easily scaled up. The selection of materials is also very important, particularly in terms of balancing the cost and the desired material features for the required application. The CNT exhibits tremendous economic potential, especially in light of recent developments that demonstrate a substantial decrease in the cost of making high-quality CNT, which fosters a silver lining in CNT-based nanodevice fabrication. Wearable biosensors with wireless communication facilities should be introduced to address this problem [67], enabling the patient to benefit from swift analysis and report the findings, to ensure a timely diagnosis. The integration of biosensors with electronic gadgets possessing smart read-out capabilities is turning into a top requirement in contemporary state-of-the-art living, in line with the current technological evolution [68]. Conversely, the toxicity of CNT and the greenness of biosensors are currently the major challenges regarding biocompatibility and sustainability. The toxicity of CNT can be reduced with the use of several techniques, such as tuning the surface defects [69], utilizing native small-molecule drugs [70], and attenuating the CNT length [71]. The greenness of the biosensor may be achieved through the green synthesis of CNT; the entire sensing operation must be performed in a green manner, i.e., using green solvents, green waste management, green power management, etc. [72]. In many cases, it was observed that the viruses mutated very rapidly (e.g., SARS-CoV-2 [73]); thus, it is vital to create dependable and effective methods based on integrated multiple biosensor technology for the quick detection of several mutations of a virus at once [74]. Moreover, CNT can also be used as an antiviral agent to inhibit viruses. By adding protoporphyrin IX to acid-functionalized MWCNTs, Banerjee et al. [75] created porphyrin-conjugated MWCNTs, which significantly

reduced the capacity of the Influenza A virus to infect mammalian cells when exposed to visible light. Iannazzo et al. [76] investigated the anti-HIV efficacy of several functionalized MWCNTs. The findings demonstrated that the antiviral activity of functionalized MWCNTs was regulated by their hydrophilic functionality and water dispersibility. Thus, CNT can also be employed in a bimodal "detection-inhibition" role to fight against viruses. Finally, further research into the nanotechnology used in virus detection is needed to achieve novel platforms that might completely revolutionize the current viral identification systems used in clinics.

**Author Contributions:** Conceptualization, C.M.H.; methodology, J.S.; writing-original draft preparation, J.S.; writing-review and editing, C.M.H. All authors have read and agreed to the published version of the manuscript.

**Funding:** This research received no external funding.

**Institutional Review Board Statement:** Not applicable.

**Informed Consent Statement:** Not applicable.

**Data Availability Statement:** Not applicable.

**Conflicts of Interest:** The authors declare no conflict of interest.
