Some researchers presented ultrasensitive DPV-based detection technology using calixarene functionalized graphene oxide for targeting RNA of SARS-CoV-2 [14]. It was affirmed that the technology identifies RNA of SARS-CoV-2 avoiding amplification and reverse transcription stages by employing a portable electrochemical smartphone. The biosensor consists of a capture probe, target sequence, label probe, and an auxiliary probe [15]. The capture probe is complementary to the 5′-terminal of the target sequence, while the label probe is complementary to the 3′-terminal; two different label probe areas have complementary sequences to 5′- and 3′-regions of the auxiliary probe [15,16]. The LOD in the clinical sample is 200 copies/mL, from which it follows that only two copies (10 μL) of all viral RNA copies are needed per analysis. The sensitivity for samples from confirmed COVID-19 patients was 85.5% [14].

It is worth noting that there are some investigations concerning the potential use of G-quadruplex-based biosensors in COVID-19 diagnosis [17]. G-quadruplex (GQ) is a guanine-rich DNA/RNA sequence, which is folded into four-stranded secondary structures. Recently, 25 putative G-quadruplex-forming sequences (PQSs) in the genome of SARS-CoV-2 virus were recognized [18]. The PQSs are situated in the ORF1ab, ORF3a, S-, M-, and N-genes of SARS-CoV-2 [17]. Some of the found PQSs are observed in a wide range of coronaviruses, while the main two PQSs, which generate RNA G-quadruplex structures, are strictly observed only in a limited range of viruses. Moreover, a straight interaction between G-quadruplex of coronavirus and viral helicase (nsp13) was obtained by microscale thermophoresis. The results of molecular docking-based modeling suggest that nsp13 alters the G-quadruplex structure. The helicase allows the guanine bases to go out of the guanine quartet planes, therefore, simplifying their unfolding [18]. Thus, RNA G-quadruplex sequences of SARS-CoV-2 could be used for the design of affinity-sensors, which are based on the identification of the viral helicase protein, nsp13 [18].

Fluorescence quenching is a powerful technique for the design of affinity biosensors [19]. One type of biosensor for the determination of enzymes based on fluorescence quenching by G-quadruplex has been reported recently [20]. 

2.1.2. Plasmonics-Based Affinity Biosensors

The primary concept of plasmonic biosensors is based on the distribution of surface plasmons lengthwise the interface of the thin metallic layer (usually noble metals), and dielectric [21]. Most plasmonic biosensors are built on the basics of surface plasmon resonance (SPR) [21,22]. Interactions occur on the surface that is suitable for observation SPR-based signals in two different modes: (1) bulk SPR signal and (2) localized SPR (LSPR) signal. Both effects rely on the refractive index of the ambient media to evoke spectral shifts. 

It was reported that a dual-functional plasmonic biosensor incorporating the plasmonic photothermal (PPT) effect and LSPR sensing transmission enables the development of an alternative approach for SARS-CoV-2 virus detection, where the detection is provided through the hybridization of complementary NA with one NA immobilized on the surface of the gold nanoislands (AuNIs). The LSPR and PPT effects were utilized mutually to increase the signal. The LOD of this assay for the RdRp gene was 0.22 pM. The specificity, the discrimination between the RdRp gene of SARS-CoV and SARS-CoV-2, can be precisely established by onsite PPT improvement on gold AuNIs-based chips [23]. Plasmonic biosensing has technological benefits including the possibility of a combination of SPR with electrochemical, and electroassisted chemiluminescence methods [24]. 

2.2.1. Field-Effect Transistor Based Immunosensors

It was reported that a Field-effect transistor (FET)-based biosensor enables the real-time detection of SARS-CoV-2 in clinical specimens. The device was manufactured by covering the graphene plates of the FET with an antibody produced as a response to the SARS-CoV-2 S-protein. The antibody was fixed on the surface of the biosensor by 1-pyrenebutyric acid N-hydroxysuccinimide ester (PBASE) (Figure 2). The cultured virus, antigen protein, and nasopharyngeal swab samples from an infected person have been utilized for the assessment of the efficiency of the immunosensor. It was determined that the FET immunosensor enables the detection of the S-protein at a concentration of 1 fg/mL in PBS and 100 fg/mL in the transport medium, whereas LOD for SARS-CoV-2 was 1.6 × 10¹ pfu/mL in culture medium and 2.42 × 10² copies/mL in clinical specimens [25].

Figure 2. Schematic representation of field-effect transistor-based immunosensor for SARS-CoV-2 detection [1].

2.2.2. Quartz Crystal Microbalance Based Biosensors

The quartz crystal microbalance (QCM) can be successfully applied for the development of affinity biosensors [26]. In QCM-based approach, the binding with the viral S-protein occurs on the engineered quartz crystal surface covered by self-assembled monolayer (SAM), and the detection is carried out by QCM. The main working principle of the QCM is altering (decreasing) the frequency of the vibrating quartz crystal with the increasing the adsorbed mass [27]. 

One more type of affinity biosensor, which is very promising for the determination of virus-induced diseases, is the ultrasound transducer-based immunosensors, e.g.: capacitive micromachined ultrasound transducer (cMUT) was applied in immunosensors for the detection of specific antibodies against some virus proteins [28]. Moreover, the ultrasound-based test allows performing the SARS-CoV-2 virus detection in the gas phase (ultrasonator-produced viral aerosol) [29], while the vast majority of the assays are designed for the solution.

2.2.3. Molecularly Imprinted Polymer Based Electrochemical Affinity Sensors

In affinity sensors, the target protein is detected on the surface of the device, thus the design of the surface with appropriate protein recognition properties is required for the development of such sensors. For this purpose, molecularly imprinted polymers (MIPs) can be very efficiently applied [7,30-33]. The advantage of molecularly imprinted sensors is that they are cheaper and more stable, and can be based on protein-imprinted polymers such as polypyrrole [34] and some other electrochemically deposited polymers [35-37]. Various signal determination methods can be applied in the design of MIP-based sensors, but mostly potentiodynamic electrochemical techniques [34] or QCM-based [26,38] approaches are used for this purpose.

Development and application of MIPs in sensor design is reasonable because MIPs can be developed for small and low molecular weight molecules [12,39]. The efficiency of MIPs for the determination of some virus proteins was also demonstrated [34] and this technology recently was applied for the development of a molecularly imprinted poly-m-phenylenediamine based electrochemical sensors for the detection of SARS-CoV-2 proteins, namely, N-protein [40]. The sensor represented a disposable MIP-modified thin film electrode possessing selectivity to N-protein. Electrochemical signal was observed by DPV and a linear response to N-protein was up to 111 fM with a detection and quantification limit of 15 fM and 50 fM [40].

It should be noted that even some short DNA-based oligomers can be determined by MIP-based sensors [31], which makes MIP-based sensors attractive for DNA and probably for RNA fragment determination. Due to the rather low price of MIPs in comparison to that of antibodies, MIP-related research area is of particular interest and, therefore, MIPs potentially can replace antibodies during the design of various bioanalytical systems and immunosensors.