Biofunctionalization of Graphene-Based FET Sensors through Heterobifunctional Nanoscaffolds: Technology Validation toward Rapid COVID-19 Diagnostics and Monitoring

Abstract

The biofunctionalization of graphene field-effect transistors (GFETs) through vinylsulfonated-polyethyleneimine nanoscaffold is presented for enhanced biosensing of severe acute respiratory-related coronavirus 2 (SARS-CoV-2) spike protein and human ferritin, two targets of great importance for the rapid diagnostic and monitoring of individuals with COVID-19. The heterobifunctional nanoscaffold enables covalent immobilization of binding proteins and antifouling polymers while the whole architecture is attached to graphene by multivalent π-π interactions. First, to optimize the sensing platform, concanavalin A is employed for glycoprotein detection. Then, monoclonal antibodies specific against SARS-CoV-2 spike protein and human ferritin are anchored, yielding biosensors with limit of detections of 0.74 and 0.23 nm, and apparent affinity constants ( $K_D^{GFET}$ ) of 6.7 and 8.8 nm, respectively. Both biosensing platforms show good specificity, fast time response, and wide dynamic range (0.1-100 nm). Moreover, SARS-CoV-2 spike protein is also detected in spiked nasopharyngeal swab samples. To rigorously validate this biosensing technology, the GFET response is matched with surface plasmon resonance measurements, exhibiting linear correlations (from 2 to 100 ng cm^-2) and good agreement in terms of K _D values. Finally, the performance of the biosensors fabricated through the nanoscaffold strategy is compared with those obtained through the widely employed monopyrene approach, showing enhanced sensitivity.

Keywords: COVID‐19; ferritin; field‐effect transistors; graphene; severe acute respiratory‐related coronavirus 2; spike protein; surface plasmon resonance.

Figure 1

A) Photograph of the portable and wireless field‐effect measurement station and scheme of a GFET sensor. B) Histogram of electrical resistance for 240 GFETs prepared with a wafer‐scale method and randomly selected from independent production batches. Different sensing features were obtained from electrolyte‐gated GFET measurements for nine randomly selected sensors: C) transconductance (g _m) normalized by V _DS = 0.05 V for the hole (orange) and electron (turquoise) branches; D) integrated I _DS current noise (I _DS ^rms); and E) gate‐source potential noise (V _GS ^rms). Data are presented with a boxplot (values within 25th and 75th percentile) and the average value is represented with a solid line. F) Scheme of the VS‐PEI nanoscaffold on a GFET sensor. G) Transfer characteristic curve for a bare GFET (solid line) and the same sensor after its surface modification with VS‐PEI (blue line), and subsequently with both mAb‐spike protein and PEG (red line). H) ΔV _CNP as a function of time for three GFETs modified with the mAb specific to SARS‐CoV‐2 spike protein. Electrolyte‐gated measurements were performed with PBS buffer pH 7.4 and V _DS = 0.05 V.

Figure 2

A) Scheme of the biosensors functionalization process employing the VS‐PEI strategy. B) Relative change in I _DS for a ConA‐GFET upon the addition of different GOx concentrations (V _DS = 50 mV, V _GS = −250 mV, HEPES buffer pH 7.4 with 0.5 mm CaCl₂). C) Results obtained at different ionic strengths and Hill (n = 1) fitting.

Figure 3

A) Scheme of the SPR measurement setup. B) Change in the surface plasmon resonance minimum reflectivity angle shift (Δθ_SPR) upon increasing GOx concentration. C) Correlation between I _DS response and GOx surface mass density obtained from GFET and SPR measurements, respectively. Both measurements were carried out in HEPES buffer pH 7.4 with 0.5 mm CaCl₂.

Figure 4

A) Scheme of the VS‐PEI approach employed for the fabrication of the SARS‐CoV‐2 S1 protein biosensors. B) I _DS relative change for a VS‐PEI‐MAb‐GFET upon the addition of increasing spike protein concentrations (V _DS = 50 mV, V _GS = −250 mV, PBS ×0.1 at pH 7.4). C) Results obtained from the selectivity assay employing the spike MERS protein compared with SARS‐CoV2. D) Change in the SPR angle upon flowing spike solution through the SPR cell for a VS‐PEI‐modified SPR sensor and scheme of the SPR measurement setup. E) I _DS relative change comparing both approaches (n = 2). F) Relative changes in I _DS as a function of the SARS‐CoV‐2 spike protein surface mass density obtained from SPR measurements for both strategies. G) SARS‐CoV‐2 S1 detection results in spiked nasopharyngeal swab samples (error corresponds to three independent measurements).

Figure 5

A) Scheme of the fabrication of ferritin biosensors, by VS‐PEI (left) and PBSE (right) approaches. B) Relative change in I _DS for a MAb‐VS‐PEI‐GFET upon the addition of different ferritin concentrations (V _DS = 50 mV, V _GS = −250 mV, HEPES buffer ×0.1 at pH 7.4). C) Results obtained for the two different approaches (n = 2). D) Relative changes in I _DS for deposited ferritin surface mass density obtained from SPR measurements for both strategies. E) Specificity experiment for ferritin sensing by BSA and mAb‐modified GFETs.

Figure 6

A) Comparison of obtained K _D values from fittings of the SPR data and the data obtained with FETs. B) Comparison of the slopes obtained with both approaches for the different systems evaluated; and a scheme of the main features presented in the biosensors that use VS‐PEI nanoscaffolds. The error bars correspond to the values obtained from the fitting.