4. Biosensors
Innovations in the development of biosensors integrated with biological elements and signal transducers enable the development of a new generation of sensor systems for POCT with rapid and precise detection of biological signals. Flexible and wearable sensor systems integrated with IoT sensing platforms have been developed to detect various biomolecules in the human body, such as glucose, lactate, uric acid, and bacteria (e.g., pathogenic
Escherichia coli) [
155,
156,
157]. In particular, novel glucose sensors are gaining considerable attention worldwide because of their applicability in the non-invasive diagnosis of diabetes mellitus through continuous glucose monitoring (CGM) [
158,
159,
160]. Early diagnosis of diabetes by monitoring glucose levels is of significant importance, considering that the total global diabetes population is expected to increase by over 50% in 2045 compared to the estimated number of diabetes patients worldwide in 2017 [
161]. Patients with diabetes experience uncontrolled blood glucose levels as a result of chronic hyperglycemia, causing various diabetic complications such as blindness, nerve damage, cardiovascular disease, and kidney failure [
162]. Therefore, adequate medical treatment and prevention of diabetes should be achieved by continuous real-time monitoring of blood glucose levels.
Electrochemical glucose sensors have been widely utilized by facilitating enzymatic reactions for CGM [
159]. In particular, the glucose oxidase (GO
x) enzyme has been commonly employed owing to several advantages such as high specificity toward glucose, stability over various pH levels, and temperature changes [
163]. The basic principle of glucose sensors using GO
x is based on the oxidation of glucose via an enzymatic reaction that produces gluconic acid and hydrogen peroxide (H
2O
2) in the presence of oxygen, as shown in the following reaction [
156]:
Toward the development of enzymatic glucose sensors using GO
x, three generations have been established depending on the mechanism of charge transfer to the sensing electrode [
156,
158]. The first generation of glucose sensors indicates the amount of glucose oxidation as a result of an enzymatic reaction, which is monitored by measuring either oxygen consumption or H
2O
2 production. The first-generation glucose sensors exhibit major advantages such as simplicity and potential for miniaturization; thus, they can be applied for in vitro and in vivo clinical trials [
164]. However, a high overpotential for the detection of H
2O
2 causes side reactions of electroactive species, resulting in low selectivity toward the target analyte. The second generation of glucose sensors involves the use of redox mediators with GO
x, wherein the mediators interact directly with enzymes and an electrical current signal is generated upon the addition of glucose as a result of the redox reaction of the mediator. For the third generation of GO
x-based glucose sensors, electron transfer occurs by direct interaction between the enzyme and the electrode without incorporating mediators. Generally, engineered enzymes are utilized to combine the electrode and GO
x through structural modification, resulting in direct electron exchange. For example, GO
x enzymes are coupled with porous polymeric membrane electrodes or nanostructured carbon nanotube electrodes to facilitate electron transfer [
165,
166,
167,
168].
Kang et al. demonstrated a wearable glucose-sensing system using GO
x-Nafion-composite-functionalized SWCNTs, which can be categorized as a third-generation GO
x-based glucose sensor [
169]. The multilayered structure of GOx-Nafion-composite-functionalized SWCNTs on a flexible substrate was achieved by an all-solution process (
Figure 13a). Specifically, a thin layer of PI with a thickness of 30 μm was coated on a Si wafer as a substrate, followed by the deposition of 1 μm of poly(methyl methacrylate) (PMMA) by spin-coating. Subsequently, the substrate was immersed in a 3-(aminopropyl)triethoxysilane (APTES) solution to form amine groups on the surface. A dispersion of SWCNTs (length ranging from 100 nm to 4 μm; diameter of 1.2–1.7 nm) in 1,2-dichlorobenzene (1 mg/100 mL) was deposited through spray-coating onto the APTES-modified PMMA/PI/Si substrate having a thickness of 3–7 nm followed by annealing at 150 °C for 30 min. Thus, dense SWCNT networks were formed as a result of Coulombic interactions between the SWCNTs and the amine groups from the APTES layer. Finally, a composite solution of GO
x and Nafion-117 was covered on the SWCNT networks through spin-coating. The composite layers were detached from the Si substrate resulting in a flexible and wearable glucose sensor, which can be directly attached to the human skin to monitor glucose concentration using a smartphone in real-time (
Figure 13a,b). A wearable glucose sensor system was established by integrating a small glucose sensor (1 cm × 1 cm in dimension) and an armband-type sensing module to transmit the sensing signal to a smartphone (
Figure 13c,d).
Material characterization and glucose-sensing performance of the fabricated sensor were investigated (
Figure 13e–g). The XPS survey analysis confirmed the surface functionalization of the SWCNT networks with the GO
x-Nafion composites, wherein peaks related to fluoride, oxygen, and sulfur were observed as a result of surface functionalization (
Figure 13e). On the other hand, the XPS survey spectrum of pristine SWCNTs exhibited no relevant peaks of GO
x-nafion composites. Real-time wireless glucose-sensing properties of the wearable SWCNT-based glucose sensor systems were evaluated by monitoring the response transitions defined by A/A
0, where A
0 and A are the initial current before exposure to glucose and measured current after the injection of glucose, respectively. The results revealed that there is a sudden increase in the current of the SWCNTs functionalized with GO
x-Nafion composites upon exposure to 50 μM glucose, whereas there were no changes in the current signal from the pristine SWCNTs (
Figure 13f). The current response transitions upon successive injection of glucose were investigated in the range of 50 μM–1 mM (
Figure 13g). Increasing current responses for the SWCNT-based glucose sensor functionalized with the GO
x-Nafion composite were observed with continuously increasing glucose concentrations. The glucose-sensing mechanism is based on the conductance of SWCNT networks affected by the enzymatic oxidation of glucose by GO
x. The fundamental principle of glucose oxidation can be explained by the formation of oxidized flavin adenine dinucleotide (FAD) as a sub-unit of the GO
x enzyme from the reduced form of FAD (i.e., FADH
2), while catalytically oxidizing glucose [
170]. The increasing current upon the injection of glucose is mainly attributed to the direct electron transfer to the SWCNT networks during the oxidation of FAD [
171].
Various flexible biosensor platforms for the detection of biological analytes have been developed on flexible substrates and applied for the point-of-care (POC) diagnosis [
172,
173]. For example, a flexible biosensor composed of a multilayered GO
x/gold/MoS
2/gold nanofilm on a PI substrate was demonstrated to be applicable for glucose detection [
173]. The multilayer structure was fabricated by sputtering gold on the PI film and subsequently depositing MoS
2 NPs through the spin-coating method. The gold sputtering process was performed again to form a gold/MoS
2/gold nanofilm on a PI substrate with the dimension of 2.5 × 20 mm. To induce glucose-sensing properties, GO
x was immobilized on a gold surface assisted by a chemical linker. The amperometric glucose-sensing result of the GO
x/gold/MoS
2/gold nanofilm revealed that a rapid increase in the current signal was obtained upon the addition of glucose with a limit of detection of 10 nM. The improved glucose-sensing response was mainly attributed to efficient electron transfer by the MoS
2 NPs during catalytic glucose oxidation.
The development of a biosensing platform composed of a paper substrate is advantageous considering its major advantages such as simplicity of fabrication, low cost, and large-scale production of sensor devices [
174]. Flexible biosensors with scalable and cost-effective strategies have been demonstrated using a disposable paper substrate. For example, a waste newspaper was employed as a sensor substrate for the detection of pathogenic
Escherichia coli O157:H7 (
E. coli O157:H7) using an electrochemical measurement technique [
172]. The disposable paper was coated with parylene C (P-paper) to enhance its mechanical properties and increase its hydrophobicity while maintaining its porous nature [
175]. After patterning the sensing electrodes on the P-paper, a self-assembled capture probe monolayer, i.e., single-strand probe DNA (ssDNA), was immobilized on a sensing electrode, followed by the formation of a 6-mercapto-1-hexanol (MCH) monolayer to block nonspecific binding to the bare gold electrode. Subsequently, a hybridization reaction was conducted by injecting synthetic cDNA or denatured amplicons of
E. coli O157:H7 as a model foodborne pathogen (
Figure 13h). CV and EIS were performed to investigate the step-by-step assembly process and the target cDNA detection capability. The CV characteristics of the aqueous 5 mM Fe(CN)
63−/4− electrolyte solution revealed that the peak current was significantly decreased with an increase in peak-to-peak separation (Δ
EP) from 110 mV for the bare Au electrode to 310 mV after the immobilization of the ssDNA probe and blocking with the MCH monolayer (
Figure 13i). Further decreased peak current and increased Δ
EP (330 mV) were achieved after hybridization with the target cDNA. EIS further confirmed the cDNA detection capability at different concentrations using P-paper-based sensors in the presence of Fe(CN)
63−/4− as an indicator (
Figure 13j). Nyquist plots revealed gradually increased charge transfer resistance (
Rct) upon increasing the target cDNA concentration. This result indicates binding between the ssDNA probe and cDNA, which results in a negatively charged surface leading to the attenuation of electron transfer.
Figure 13. (
a) Schematic illustration of a wearable SWCNT-based glucose sensor fabricated by an all-solution process. Camera images of (
b) a wearable SWCNT-based glucose sensor on a finger and (
c) integrated with a wearable sensing module. (
d) Wearable glucose sensor system integrated with an IoT-based sensing module to transmit glucose-sensing data to a mobile device. (
e) XPS spectra analysis to confirm the functionalization of the GO
x-Nafion composite on an SWCNT film. (
f) Real-time glucose-sensing property of pristine SWCNT and GO
x-Nafion-composite-functionalized SWCNT upon exposure to 50 μM of glucose. (
g) Real-time response changes of GO
x-Nafion-composite-functionalized SWCNT upon successive addition of glucose ranging from 50 μM to 1 mM. Reprinted with permission from Ref. [
169] Copyright (2019), Elsevier. (
h) Disposable-paper-based electrochemical sensors after coating parylene C (P-paper) and electrodes for the detection of foodborne pathogens (i.e., cDNA of
E.
coli O157:H7). (
i) CV curve to investigate the step-by-step assembly process of ssDNA probe immobilization, MCH blocking monolayer formation, and cDNA hybridization. (
j) Nyquist plots of the paper-based sensor upon exposure to different concentrations of target cDNA. Reprinted with permission from Ref. [
172] Copyright (2016), American Chemical Society.
As a different type of transduction mechanism, colorimetric sensing mechanism has been utilized to detect biomolecules and viruses because of their simple visual readout and their capability to rapidly screen multiple analytes with high portability [
176,
177,
178,
179,
180,
181]. Several studies have been conducted to fabricate colorimetric biosensors using a paper substrate paired with a smartphone-based reader for application in POCT [
182,
183,
184,
185,
186]. A paper/soluble polymer hybrid-based biosensing platform was developed for the diagnosis of myocardial infarction by detecting human cardiac troponin I (cTnI) as a standard biomarker [
182]. Among the various biosensing platforms for POCT applications, lateral flow assays (LFAs) are the most widely used because of their major advantages such as affordability, simplified device architecture, user-friendliness, ability for rapid detection, robustness, and long shelf life (~2 years) under ambient conditions [
186,
187,
188,
189]. However, because of the relatively low sensitivity of conventional LFAs, they cannot be effectively applied for biomarker detection in the concentration as low as sub-ng/mL. To overcome this limitation, a paper-based LFA with signal amplification (i.e., signal-amplification-based LFA) was proposed to facilitate biochemical reactions to further enhance sensitivity and promote quantitative analysis [
190,
191]. To this end, low-cost and mass-produced batch-type test strips were prepared to analyze cTnI, incorporated with a smartphone-based reader for high-performance POCT. The paper/polyvinyl alcohol (PVA) hybrid was patterned by dispensing the PVA solution on nitrocellulose (NC) membrane, which plays a key role in programmable fluid control and automated fluid switching (
Figure 14a). The test proceeds with an assay followed by a signal readout using a smartphone. The assay was performed by the injection of a sample solution containing a cTnI biomarker to induce immunoreaction and a reagent solution to activate the amplification reaction within 20 min (
Figure 14b). The mixture fluid injected through a reagent pad gradually dissolved the patterned PVA barrier, resulting in fluid switching from the sample fluid to the amplification fluid (
Figure 14c). As a result, the test platform realized automated signal amplification reactions at the test and control lines. Various amplification techniques have been demonstrated such as Au-ion amplification, wherein gold ions (Au
3+) were reduced to Au NPs in the presence of a reducing agent (H
3NO), thereby generating amplified colorimetric signal changes. The enhanced colorimetric signal was proportional to the amount of reduced Au NPs, which were formed after the immunocomplex reaction. The intensity of color changes was measured using a smartphone reader after 20 min of the assay (
Figure 14d). The result revealed excellent analytical sensitivity with a detection limit of 0.92 pg/mL cTnI and a coefficient of variation of <10% in serum or plasma samples comparable to those of commercially available standard analyzers, thereby demonstrating its potential application in POCT systems.
Figure 14. (
a) Dispensing of PVA on a nitrocellulose (NC) membrane for the LFA test platform with SEM images. (
b) LFA test process by the injection of sample and reagent solutions. (
c) Schematic illustration of colorimetric Au-ion amplification facilitating automated reaction/fluid switching mechanism by dissolving the PVA barrier. (
d) Colorimetric signal readout using a smartphone. Reprinted with permission from Ref. [
182] Copyright (2020), American Chemical Society. (
e) Schematic illustration of the mechanism of the colorimetric urea biosensors, generation of NH
3 by an enzymatic reaction between urea and urease, and pH-responsive reduction of Ag
+ ions to Ag NPs by tannic acid. (
f) Schematic illustration of the fabrication of a colorimetric urea biosensor. (
g) Analysis of colorimetric urea biosensor based on the calculated RGB ratio at different urea concentrations. (
h) Smartphone-assisted RGB ratio measurement in the urea concentration range of 0–500 mM. Reprinted with permission from Ref. [
183] Copyright (2021), Elsevier.