Chapter 1: Brief Overview of Different Biosensors: Properties, Applications, and Their Role in Chemistry
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Published:20 Dec 2024
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Special Collection: 2024 eBook CollectionSeries: Detection Science Series
B. P. Suma and P. S. Adarakatti, in Biosensing Technology for Human Health
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Biosensors are cutting-edge instruments with the unique ability to identify biological molecules with extreme sensitivity and specificity, changing industries such as food safety, environmental monitoring, and healthcare. To transform the biological reaction into a detectable signal, their design usually combines a transducer with a biological detection element, such as enzymes, antibodies, or nucleic acids. They are extremely useful in various applications due to characteristics such as mobility, low detection limits, and rapid response. The principles of electrochemical processes, enzymatic reactions, and affinity-based interactions underpin the functioning of biosensors. Optical, electrochemical, and piezoelectric biosensors are among the types available; each has a unique set of benefits based on the intended use. Applications include pathogen detection, environmental pollution monitoring, medical diagnostics, and food safety and quality assurance. Further developments in disease diagnosis, tailored medication, and environmental monitoring are anticipated due to their adaptability and specificity, which are continuing to spur innovation.
1.1 Introduction
The field of bioelectronics focuses on using concepts of electronics to understand biology and medicine.1 A unique type of bioelectronic instrument frequently utilized in bioanalysis is the biosensor. The “primary component of a measurement chain that converts the input variable into a signal suitable for measurement” is often understood to be a sensor.2
Biosensors are small analytical instruments developed in response to the recognized need for monitoring key processes and parameters across various sectors. Drug discovery, disease diagnosis, biomedicine, food safety and processing, environmental monitoring, and defense and security are only a few of the areas where the development of these devices has provided new alternatives.3 Biosensors are commonly used as analytical tools to detect specific analytes in samples. These are autonomous, integrated devices that use a biological recognition element attached to a transduction element to deliver qualitative and semi-quantitative analytical data. These analytical instruments are only intended to rapidly deliver precise and trustworthy real-time information about an analyte of interest. Furthermore, numerous carbon substrates have been used in recent years for measuring various small biomolecules.4–14
The development of enzyme electrodes by scientist Leland C. Clark in 1962 marked the beginning of the history of biosensors.15 Since then, various research organizations have collaborated to develop increasingly complex, dependable, and advanced biosensing technologies. These groups include those in fields such as very-large-scale integration (VLSI), physics, chemistry, and materials science. These tools have been utilized in biotechnology, medicine, agriculture, and the military, as well as in the identification and combating of bioterrorism.16,17 Biosensors can also be referred to as optrodes, immunosensors, glucometers, chemical canaries, resonant mirrors, biochips, and biocomputers. According to S.P.J. Higson and D.M. Frazer, two commonly used definitions of biosensors are as follows: “a biosensor is a device for analysis, integrating a deliberate and intimate combination of a specific biological element, which generates a recognition event, and an external component, which transduces the recognition event,” and “a biosensor is a chemical sensing device in which a biologically derived recognition entity is coupled to a transducer, allowing the quantitative measurement of complex biochemical parameters.16 A biosensor is a device that combines a bioelement and a sensor element. Figure 1.1 illustrates the fundamental mechanism behind how a biosensor functions. The sensor element converts the change in the biomolecule into an electrical signal when a particular bioelement, such as an enzyme, detects a specific analyte. The bioelement has a very narrow range of responses to the analyte it is sensitive to, and it does not recognize additional analytes. Various biosensor types, including resonant, optical, thermal, and ion-sensitive field-effect transistor (ISFET) biosensors, can be used, depending on the transducing mechanism used.19 Conductometric, amperometric, and potentiometric electrochemical biosensors can be further categorized based on the characteristics being monitored.
A schematic representation of biosensors. Reproduced from ref. 18, https://doi.org/10.1155/2013/731501, under the terms of the CC BY 3.0 license, https://creativecommons.org/licenses/by/3.0/.
A schematic representation of biosensors. Reproduced from ref. 18, https://doi.org/10.1155/2013/731501, under the terms of the CC BY 3.0 license, https://creativecommons.org/licenses/by/3.0/.
1.1.1 Basic Principle of a Biosensor
A biosensor functions by signal transduction, as it consists of a biological receptor attached to a transducer and signal processing device. These elements work together to translate the biological response into an equivalent electrical reaction and, in the end, a quantifiable output.20 Put more simply, biosensors translate the biological activity of a molecule into a quantifiable signal, which allows for the quantitative assessment of the molecule. First, a physiological shift is caused by the binding of the molecule of interest in the test sample to or interacting particularly with the biological receptor. Consequently, the physicochemical characteristics of the transducer placed next to the biological receptor are further changed.21 This also leads to an alteration in optical or electronic characteristics of the transducer, which are then transformed into a detectable electrical signal. The response of the transducer can be either a voltage or a current, based on the type of biological receptor used. The transducer will transform any current that is produced into an equivalent voltage if the output is in that form. Additionally, a high-frequency noise signal masks the output voltage, which necessitates additional processing, amplification, and modification using various filters inside the signal processing unit. Ultimately, the biological quantity being evaluated ought to be equivalent to the output produced by the signal processing module.2
1.2 Biosensor Design
The two primary components of an effective biosensor are a transducer and a biological receptor, often known as a sensor element. Typically, a signal processing unit, which often includes a printer or display, is utilized with a biosensor, as shown in Figure 1.2.
A schematic representation of the biosensor design. Reproduced from ref. 22, https://doi.org/10.5772/intechopen.97576, under the terms of the CC BY 3.0 license, https://creativecommons.org/licenses/by/3.0/.
A schematic representation of the biosensor design. Reproduced from ref. 22, https://doi.org/10.5772/intechopen.97576, under the terms of the CC BY 3.0 license, https://creativecommons.org/licenses/by/3.0/.
1.2.1 Classification Based on the Biorecognition Element Used in the Development of Biosensors
A biosensing device consists of a bioreceptor component, which is an immobilized sensitive biological element such as an enzyme, DNA probe, or antibody, having specificity in recognizing the analyte of interest, like an enzyme substrate, complementary DNA, antigen, or antibody. Although antibodies and oligonucleotides are widely used, enzymes are by far the most commonly used biosensing elements in the fabrication of biosensors. Efforts have been achieved to develop biosensors based on oriented and site-specific immobilization of the biorecognition element.
The success of the developed biosensor and its analytical performances are strongly affected by the process of immobilization used during its development and fabrication.
Various immobilization techniques have been proposed to achieve a good analytical performance, such as the entrapment method, where enzymes can be immobilized in three-dimensional matrices, such as polymer films, neutral network of polymers [i.e. polydimethylsiloxane (PDMS)], silica gel, polysaccharides, or carbon-based composites.
Using the entrapment method of immobilization, an innovative biosensor for the detection of l-proline (L-Pro) in human plasma samples at physiological pH was developed by entrapping proline dehydrogenase on amine-functionalized dendritic fibrous nano-silica KCC-1. This fibrous nano-silica KCC-1–NH2 with its high surface area showed greater properties for the encapsulation of proline dehydrogenase. By combining KCC-1–NH2 with proline dehydrogenase and casting it onto a pre-treated glassy carbon electrode surface, a biosensor was fabricated to determine l-Pro.23
The adsorption method involves the use of solid supports on which the enzymes are adsorbed. It represents the easiest method of physical immobilization. The adsorption mechanisms are based on weaker Vander Waals forces or electrostatic or hydrophobic interactions. Essentially, the required enzyme is dispersed in a solution and the solid support is then placed in contact with the enzyme solution for a fixed period of time. The unadsorbed enzyme is then removed by washing with a suitable buffer solution.24
Another well-known method for immobilization of a biorecognition element is by cross-linking with glutaraldehyde or other bifunctional agents such as glyoxal or hexamethylenediamine to fabricate biosensors. In this case, the enzymes can be cross-linked either with each other or with a functionalized inert protein. This method is remarkable due to its strong chemical binding between biomolecules and its simplicity.
Bodur et al. developed a bienzymatic biosensor system using acetylcholine esterase (AChE) and choline oxidase (ChO) with carbon paste electrodes modified with a PAMAM–salicylaldehyde dendrimer. Later, acetylcholine esterase and choline oxidase enzymes were cross-linked with glutaraldehyde and immobilized onto modified carbon paste electrodes.25
Covalent immobilization, i.e. coupling of biorecognition element to polymeric supports through chemical bonding, is a popular chemical immobilization method frequently used to develop biosensors. For this purpose, biocatalysts are bound to the surface via different functional groups that they contain, which are not essential for their biocatalytic activity. The binding of enzymes to the solid support is generally carried out by the initial activation of the surface using multifunctional reagents.
The sensing platform of biomolecule-functionalized graphene has received extensive attention due to its high sensitivity and selectivity, especially in biosensors developed by combining antibodies, nucleic acids, or enzymes that efficiently recognize specific targets with graphene. Electrochemical biosensors based on graphene materials developed in recent years have been summarized. The methods of functional modification of graphene, graphene oxide, and reduced graphene oxide with antibodies, nucleic acids, or enzymes are briefly described. In addition, the advantages and disadvantages of the developed electrochemical biosensors in detecting pathogens and disease markers are also reviewed.26
Another strategy is to develop an affinity bond between an activated support (e.g. with lectin, avidin, or metal chelates) and a specific group in the protein sequence (e.g. carbohydrate residue, biotin, or histidine). This method facilitates control over biomolecule orientation, which prevents enzyme deactivation and blockage of the active site, thereby affecting selectivity and sensitivity.24
The best method for enzyme immobilization depends on whether the biosensor application requires maximum sensitivity rather than stability, as many enzyme-based biosensors are designed for single use and disposability. Sensitivity decreases if immobilization causes enzyme denaturation, conformational changes, deactivation, or modification, particularly on the active site, upon exposure to physiological conditions or the target analyte.
1.2.1.1 Enzyme-based Biosensors
Enzyme immobilization is a crucial factor in developing efficient biosensors with optimal performance, including good operational and storage stability, high sensitivity, high selectivity, short response time, and high reproducibility. Immobilized biorecognition elements must maintain their structure and conformation, preserve their specific function, retain their biological activity after immobilization, and remain tightly bound to the surface without being desorbed or deactivated during the biosensing process.
Since then, numerous enzyme sensors have been developed for determining various substances, such as glucose,27,28 cholesterol,29 or lactic acid in biological fluids (blood, serum, or urine),30 for toxicity analysis in environmental monitoring,31 for food and quality control,32 and in the biomedical and drug-sensing arena.
In this regard, Xue et al. developed a flexible epidermal microwave glucose biosensor integrated with nanostrip structures. These nanostrip-based microwave sensors were developed as epidermal glucose biosensors capable of conformal contact with the skin and non-invasive real-time monitoring of sweat glucose. Here, the GOx enzymes were integrated into the nanostrips for specific glucose sensing. To enable specific glucose sensing, the GOx enzyme was immobilized into nanowires of the proposed microwave sensor. When exposed to a glucose sample, the reaction between the GOx enzyme and glucose produces gluconolactone (lactone of gluconic acid) and hydrogen peroxide (H2O2) according to the following reactions.
GOx can be reused as an enzyme once the reaction reaches equilibrium, resulting in changes in product concentrations. Incorporating GOx ensures that any concentration change in the environment is specifically related to the presence and amount of glucose in a given sample. Concentration changes will result in a variation in the dielectric constant, which can be detected by the microwave sensor.33 Various types of electrodes for glucose detection are summarized in Table 1.1.
Comparison of various modified electrodes with existing sensors.
Electrode . | Technique . | Linear range . | Limit of detection . | Real sample . | Ref. . |
---|---|---|---|---|---|
ZnO–Pt–g-C3N4 | Chronoamperometry | 0.1 to 0.5 μM | 0.1 μM | Human blood, urine, and serum | 34 |
FcBA/glucose/3APBA/4MBA/AuNPs/ITO | DPV/EIS/CV | 0.5–30 μM | 43 μM | Human urine | 35 |
AuNP–MIPs | SWV | 1.25 nM to 2.56 μM | 1.25 nM | Human serum | 36 |
AuNP/PAB/FTO | CV | 2–50 μM and 50–250 μM | 0.40 μM | Human serum | 37 |
Cu-nanoflower@AuNPs–GO NFs | CV | 0.001–0.1 mM | 0.018 μM | NA | 38 |
Nafion/GOx/ZnO/rGO/ITO | CV/amperometry | 0–6.5 mM | 1 μM | Human serum | 39 |
Electrode . | Technique . | Linear range . | Limit of detection . | Real sample . | Ref. . |
---|---|---|---|---|---|
ZnO–Pt–g-C3N4 | Chronoamperometry | 0.1 to 0.5 μM | 0.1 μM | Human blood, urine, and serum | 34 |
FcBA/glucose/3APBA/4MBA/AuNPs/ITO | DPV/EIS/CV | 0.5–30 μM | 43 μM | Human urine | 35 |
AuNP–MIPs | SWV | 1.25 nM to 2.56 μM | 1.25 nM | Human serum | 36 |
AuNP/PAB/FTO | CV | 2–50 μM and 50–250 μM | 0.40 μM | Human serum | 37 |
Cu-nanoflower@AuNPs–GO NFs | CV | 0.001–0.1 mM | 0.018 μM | NA | 38 |
Nafion/GOx/ZnO/rGO/ITO | CV/amperometry | 0–6.5 mM | 1 μM | Human serum | 39 |
AuNP–MIPs, Gold nanoparticle–molecularly imprinted polymers; CV, Cyclic voltammetry; DPV, differential pulse voltammetry; FcBA/glucose/3APBA/4MBA/AuNPs/ITO, Ferrocene boronic acid/4-mercapto benzoic acid/gold nanoparticles/indium tin oxide; GO, Graphene oxide; PAB/FTO, Polyaniline blue/fluorine-doped tin oxide; SWV, square wave voltammetry; ZnO–Pt–g-C3N4, Zinc oxide–platinum–graphitic carbon nitride.
Enzymatic sensors for the determination of pesticides are most often emphasized in the food and agricultural industries. Sensing mechanisms are based on the inhibition of the activity of selected enzymes, such as cholinesterases, organophosphate hydrolase, alkaline and acid phosphatases, ascorbate oxidase, acetolactate synthase, and aldehyde dehydrogenase. These enzymatic biosensors are developed using various electrochemical signaling transducers, enzyme immobilization techniques, and measuring methodologies. The use of single-use and throw-type screen-printed biosensors in bulk measurements and continuous-flow injection monitoring with enzyme biosensors are the most intensively developed procedures. Not only improvement in detectability level but also multi-analyte determinations can be achieved using recombinant enzyme mutants. In several pesticide analyses, the target pesticide can be used as a substrate for enzymatic reactions.40
In the review article by Kaur and Singh, recent progress in the field of biosensors using optical detection of organophosphorus pesticides has been discussed. Specifically, biosensors developed using enzymes, such as acetylcholinesterase and organophosphate hydrolase, have been categorized based on the electrode material used for their fabrication.41
The use of nanoparticles as one of the components, along with the biorecognition moiety, in biosensor development has been explored recently. Owing to this, a one-step synthesis method for fabricating nanoparticle-based enzymatic biosensors has been proposed, which involves a simultaneous encapsulation of enzymes, like glucose and alcohol oxidases, a fluoropolymer like Nafion, and noble metal nanoparticles through co-deposition from a phosphate multiple electrolyte system in the simultaneous multi-analyte detection. The developed sensor showed good specificity along with high sensitivity.42
In this regard, multiscale simulations were utilized to study the adsorption behavior of the enzyme acetylcholinesterase from Torpedo californica (TcAChE) on different carbon substrates. The amino-functionalized carbon nanotube (CNT) (NH2–CNT), carboxyl-functionalized CNT (COOH-CNT), and pristine CNT surfaces are shown in Figure 1.3. Simulation studies showed that the active center and enzyme–substrate tunnel of TcAChE are close to each other and conformations oriented toward the surface when adsorbed on the positively charged NH2–CNT, which is beneficial to the direct electron transfer (DET) and accessibility of the substrate molecule, thereby increasing the charge transfer rate. NH2–CNT can also reduce the tunnel cost of the enzyme–substrate of TcAChE, thereby further accelerating the transfer rate of the substrate from the surface–solution interface to the active center. The results indicated that NH2–CNT is more suitable for the immobilization of TcAChE. The study provided an improved molecular understanding of the adsorption mechanism of TcAChE on functionalized CNT and also gave theoretical guidance for the ordered immobilization of TcAChE and the design, development, and improvement of TcAChE-OP biosensors based on different functionalized carbon nanomaterials.43
Various carbon substrates. Reproduced from ref. 43 with permission from the Royal Society of Chemistry.
Various carbon substrates. Reproduced from ref. 43 with permission from the Royal Society of Chemistry.
The review article by Verma summarizes nanobiotechnological advances in the field of the agricultural and food industry, with an emphasis on the selective detection of pesticides and food-borne contaminants encountered in day-to-day life. Nanoenzymatic biosensors display ultrasensitivity at the nanomolar-to-picomolar level and rapid detection time in real-time analysis.44
In the context of biosensing applications in the medicinal field, acid phosphatase (ACP) is a potential biomarker for detecting various diseases. Xiang et al. used DNA extension-based biological signal amplification and redox recycling of electrochemical species on carbon-modified electrodes. This approach holds great potential for the convenient detection of other biomolecules such as alkaline phosphatase, even at low levels. They established a highly sensitive, label-free, and electrode immobilization-free electrochemical assay for ACP in diluted human serum samples. The mechanism involves the target analyte ACP catalyzing the dephosphorylation of 3′-PO4 termini of ssDNAs into 3′-OH on magnetic beads (MBs), which consequently initiates the extension of strands into long guanine-rich ssDNAs by the terminal deoxynucleotidyl transferase with predefined ratios of guanine to adenine bases. Magnetic accumulation of beads on the rGO-modified electrode thus yields significantly enhanced current signals due to redox recycling oxidation of numerous guanine bases in the extended DNAs, mediated by [Ru(bpy)3]2+. Such significant current amplification can, therefore, lead to ultrasensitive electrochemical detection of ACP, with a low detection limit in diluted serum samples. This assay has the notable benefits of simplicity and sensitivity.45
Recently, there has an increasing demand for the development of tissue-interfaced chemical biosensors for on-site, real-time biological sample collection and analysis. Based on tissue structures in living animals, biosensors can be used in detecting and measuring biological sample matrices, such as drugs, hormones, and toxins. The main functions of these chemical biosensors are generally based on biochemical reactions or physical interactions between specific bioreceptors and target chemical analytes in complex biological samples. The biorecognition event occurring upon interaction can be translated into analytically measurable signals, which can then be analyzed semi-quantitatively or quantitatively. Biosensor cells transduce the concentration of the molecule being detected into a physical signal, which can be precisely measured during analysis.46 In this context, the review article by Peng et al. elaborates on the recent advances in tissue-interfaced chemical biosensors.47
Immunosensors were developed based on the fact that antibodies have a high and specific affinity toward their respective antigens, i.e. antibodies specifically bind to pathogens or toxins or interact with components of the host immune system.48 These immunosensors are profoundly used in disease detection and management.
Saber and Omidi reported the design and development of a novel electrochemical immunosensor for detecting cancer antigen 125 (CA125) oncomarker. In this study, polyamidoamine with gold nanoparticles (PAMAM/AuNPs) was used to increase the conductivity and enhance the number of antibodies immobilized onto the electrode surface. Three-dimensional reduced graphene oxide–multi-walled carbon nanotubes (3DrGO–MWCNTs) were used to modify the glassy carbon electrode surface to improve the electrode conductivity and active surface area. Ab and toluidine blue attached to O-succinyl-chitosan–magnetic nanoparticles (Suc-CS@MNPs) were used as a tracer. The poor solubility of chitosan (CS) was improved by adding succinic anhydride using a novel modification method. Under optimum condition, the developed immunosensor exhibited a wide linear range. The reliability of the engineered immunosensor for detecting CA125 was further compared to the enzyme-linked immunosorbent assay (ELISA) technique.49
Combining the increased sensitivity of electroanalytical methods with the inherent bioselectivity of the biological component of DNAs, DNA-based electrochemical biosensors hold the advantages of simplicity, rapid response, low cost, high sensitivity, miniaturization, and low sample volume requirements. They have been developed for detecting various biological molecules, including nucleic acid (e.g., DNA and RNA) and non-nucleic acid targets.50 The detection of nucleic acid targets mainly depends on the formation of double-helical structures based on the Watson–Crick base-pairing strategy, while the corresponding aptamers are usually utilized for the recognition of non-DNA targets with high affinity and specificity.51
In this regard, Panneerselvam et al. developed a fluorescent biosensor for the successive detection of Pb2+ ions and the cancer drug epirubicin (Epn) using the interactions between label-free guanine-rich ssDNA (LFGr-ssDNA), acridine orange (AO), and a metal–phenolic nanomaterial [i.e., nano-monoclinic copper–tannic acid (NMc–CuTA)]. An examination of the sensing mechanism showed that LFGr-ssDNA and AO strongly adsorb on NMc–CuTA through π–π stacking and electrostatic interactions, and this results in the fluorescence quenching of AO. To sense the target Pb2+ ions, initially, LFGr-ssDNA specifically binds with Pb2+ ions to form a G4 complex (G–Pb2+–G base pair), which was released from the surface of NMc–CuTA with strong AO fluorescence enhancement (Turn-ON). The subsequent addition of a biothiol, such as cysteine (Cys), to the G4 complex decreases the fluorescence, as the Pb2+ ions released from the G4 complex have a higher interaction affinity with the sulfur atoms of Cys; this further induces the unwinding of the G4 complex to form LFGr-ssDNA. Finally, Epn was added to this, which intercalates with LFGr-ssDNA to form a G4 complex via G–Epn–G, resulting in fluorescence recovery (Turn-ON).52
Another report focuses on the synthesis of high-quality graphene nanosheets obtained through the electrochemical exfoliation of biomass derived from corn cob. The conductive ink prepared from this exfoliated graphene was utilized for the preparation of paper-based graphene electrode for double-stranded DNA (dsDNA)-based sensor applications. In this study, two irreversible oxide peaks were obtained from paper-based printed graphene electrodes, corresponding to the oxidation of guanine (G) and adenine (A) in dsDNA. Furthermore, a small-scale printable circuit is fabricated using graphene shows good conductivity. The authors reveal that the method opens the possibility of direct electrochemical analysis without any prior sample preparation to preconcentrate the analyte.53
In recent developments in biosensors for cancer prevention and diagnosis, Kadhim et al. developed a nanobiosensor based on a DNA–GO nanohybrid to detect deletion mutations that cause lung cancer, as shown in Figure 1.4. In this method, mutations were detected using a fluorescein amidite (FAM)-labeled DNA probe with fluorescence spectrometry. For lung cancer detection, the DNA probe is adsorbed onto the GO surface, and the fluorescence intensity increases with the addition of healthy DNA. However, the fluorescence intensity shows no changes with mutated mDNA. In other words, the designed nanobiosensor responds differently to healthy and mutated DNAs, allowing for detecting mDNA in cancer patients.54
Responses of nanobiosensors to healthy DNA and mutated mDNA. Reproduced from ref. 54 with permission from the Royal Society of Chemistry.
Responses of nanobiosensors to healthy DNA and mutated mDNA. Reproduced from ref. 54 with permission from the Royal Society of Chemistry.
1.2.1.2 Aptamer-based Biosensors
Recently, aptamers have gathered increasing interest due to their inimitable properties, including inexpensive production, simple chemical modification, and long-term stability. Generally, aptamers are short, single-stranded nucleic acid molecular recognition mediators that fold into 3D conformations and bind specifically to targets such as proteins, peptides, small molecules, and metal ions. They behave similarly to antibodies. At the same time, aptamers possess similar binding affinity and specificity to their protein counterparts. Aptamers have a wide range of binding targets and applications, making them excellent tools among molecular recognition agents in the development of new-age biosensors.55 In line with this, Song et al. demonstrated the development of improved screen-printed carbon electrode (SPCE), which is applied for the detection of a wide range of bioactive molecules. Typically, they designed gap hybridization, aptamer “sandwich”, and aptamer competition reduction strategy for the detection of miRNA-141, thrombin, and ATP, respectively. The results showed that the DNA tetrahedron-modified SPCE worked well with serum samples. The carbon-based DNA framework nano–bio interface would expand the use of SPCE and make electrochemical biosensors more available and valuable in clinical diagnosis.56
1.2.2 Transducing Element
The key element in the development of a biosensor is a transducer. A substance that can change one type of energy into another is called a transducer.20 The role of a transducer in a biosensor is to transform biochemical signals that the biological receptor receives as a result of the interaction between the target analyte and the biological receptor into a tangible and quantifiable signal, which can be optical, piezoelectrical, electrochemical, or another type.57 The transformation that happens during biological receptor–analyte contact is detected and measured by the transducer. The pH sensor in a glucose biosensor serves as an example of a transducer. Subsequently, the pH sensor (transducer) recognizes the pH shift caused by the synthesis of gluconic acid and transforms it into a voltage shift.58 When designing a transducer, it is recommended to consider the following characteristics: reaction time, analyte concentration range, specificity to the target analyte, and realistic applicability. A transducer should ideally measure the lowest analyte concentrations in the shortest time while exhibiting good analyte specificity.22,59
Based on the type of transduction components the sensor contains, biosensors are most frequently categorized. Electrochemical biosensors, mass-based biosensors, and optical-based biosensors are the three major groups into which these biosensors are divided. As a result of their distinct operating principles, the three biosensors can be used in a wide range of circumstances. An overview of various types of biosensors and their function is provided later.
1.2.2.1 Electrochemical Biosensors
The inherent bioselectivity of the biological component combined with the sensitivity of electroanalytical techniques makes an electrochemical biosensor. The biological element of the sensor identifies its analyte, causing a catalytic or binding event. This, in turn, results in an electrical signal that is proportionate to the analyte concentration and can be tracked by a transducer. Numerous sensor devices have successfully made it to the market and are frequently used in industrial, agricultural, medical, and environmental settings.60–68
Multi-walled carbon nanotubes (CoPc–MWNTs) containing cobalt (ii) phthalocyanine have been developed and examined using transmission electron microscopy (TEM) and energy-dispersive X-ray spectroscopy (EDX). Furthermore, the following method has been used to synthesize the modified composite: by annulating dicyanobenzene on the surface of MWNTs using Co(ii) as a template, the CoPc–MWNTs were developed in situ, and 2.19 × 10−6 mol DBU (1,8-diazabicycloundec-7-ene) was mixed with 0.05 g MWNTs and 5.85 × 10 −4 mol CoCl2 after being dissolved in 30 mL of n-pentanol for two minutes in an ultrasonic bath. The blend was agitated for two hours at 170 °C while being nitrogen-suspended. After filtering the reaction mixture, the solids were cleaned with ethanol and vacuum-dried for four hours at 80 °C, and then using the drop-coating approach, the cobalt (ii) phthalocyanine–multiwalled carbon nanotube-modified glassy carbon (CoPc–MWNTs/GC) electrode was developed. For the oxidation of ascorbic acid (AA), the electrocatalytic performance of the chemically modified electrode was examined. The outcomes show that the modified electrode exhibits good electrocatalytic activity when it comes to oxidizing AA in a solution of 0.1 M phosphate buffer. The biosensor could withstand a broad linear concentration range for AA, from 1.0 × 10−5 M to 2.6 × 10−3 M, with a detection limit of 1.0 × 10−6 M. Fast reaction time, good repeatability, and stability are a few of the key qualities of the CoPc–MWNT-modified glassy carbon electrode.69
Additionally, an investigation aimed to develop nanoporous gold (NPG)-based electrochemical biosensors by using unique physical and chemical characteristics of NPG, which was produced merely by dealloying Ag from the Au/Ag alloy. While comparing the NPG-modified glassy carbon electrode (NPG/GCE) with the standard gold sheet electrode, there was a notable decrease in the overpotential of the electro-oxidation of hydrogen peroxide and nicotinamide adenine dinucleotide (NADH). This was due to the high electrocatalytic activity of NPG/GCE toward oxidations of these two substrates. The electrocatalytic performance is expected to be caused by the high density of edge-plane-like defective sites and the substantial particular surface area of NPG. Because of the electrocatalytic activity of NPG/GCE, alcohol dehydrogenase (ADH) or glucose oxidase (GOD) might be included in the three-dimensional NPG matrix to enable efficient low-potential amperometric biosensing of ethanol or glucose. As NPG has a clean, consistent, and evenly dispersed microstructure, ADH- and GOD-modified NPG-based biosensors performed well analytically in the biosensing of ethanol and glucose. The built-in biosensors were likewise extremely stable due to the stabilizing impact of NPG on the integrated enzymes. The initial current response of ADH- and GOD-based biosensors decreased by only 5.0% and 4.2% after a month of storage at 4 °C. Each of these suggested that NPG was a good electrode material to use for building biosensors.70
Fabiana et al. developed a screen-printed electrochemical electrode (SPE) for paraoxon, driven by the ability of the compound to block the enzyme butyrylcholinesterase (BChE). After dispersing carbon black nanoparticles (CBNPs) in dimethylformamide–water and drop casting it onto the electrode, BChE was immobilized onto the surface by cross-linking. The resultant biosensor was subjected to typical paraoxon solutions, and at a working voltage of +300 mV, the enzymatic product thiocholine was measured to track the enzymatic hydrolysis of butyrylthiocholine over time. Up to 30 μg L−1, the enzyme inhibition is linearly correlated with the paraoxon concentration, and the detection limit is 5 μg L−1. The biosensor can be stored at ambient temperature and in dry conditions for several days. It was used to detect paraoxon in wastewater samples that had been contaminated. The findings support the prospective application of CBNPs in electrochemical biosensors and show that they constitute a competitive option in addition to being a very cost-effective substitute for other carbon nanomaterials such as carbon nanotubes or graphene.71
To analyze uric acid in human blood serum in an easy and quick manner, a uricase (UOx) biosensor was developed. To detect hydrogen peroxide, which is produced by a reaction catalyzed by the UOx enzyme immobilized onto the surface of the working electrode, the proposed system uses an SPE modified with Prussian blue (PB) in conjunction with portable instruments. When it comes to electrochemical substances that are typically found in blood, including ascorbic acid and uric acid, PB is a mediator that is extremely stable at physiological pH and interference-free. An analysis of recovery was conducted following the immobilization of enzymes and the assessment of functions of biosensors. After analyzing 85 human serum samples with the biosensor, the outcomes had statistical significance and were compared to the findings of the spectrophotometric method.72
The uricase enzyme has been imprinted onto indium tin oxide scaffolds, which were coated with Prussian blue using the layer-by-layer approach, as shown in Figure 1.5. The uricase layers were substituted with poly(ethylene imine) or poly(diallyldimethylammonium chloride), and the resultant films served as uric acid amperometric biosensors. The best performing biosensors displayed a linear response between 0.1 and 0.6 μM of uric acid and a detection limit of 0.15 μA μmol l−1 cm−2, which is enough for application in clinical trials. As the measurement was done at 0.0 V vs. a saturated calomel electrode, biological activity remained elevated for weeks with very little interference.73
Schematic architecture of PEI or PDAC/uricase multilayers onto ITO modified with a PB layer (a). Mechanisms involved in the detection of uric acid with generation of H2O2 (b). Reproduced from ref. 73 with permission from Springer Nature, Copyright 2007.
Schematic architecture of PEI or PDAC/uricase multilayers onto ITO modified with a PB layer (a). Mechanisms involved in the detection of uric acid with generation of H2O2 (b). Reproduced from ref. 73 with permission from Springer Nature, Copyright 2007.
A simple amperometric uric acid sensor was developed by immobilizing immulose uricase on gold/amino acid nanocomposites. The resulting biosensor has a broad linear range from 0.02 to 2.5 mM, a sensitivity of 108 A mM−1 cm−2, and a quick reaction time of less than 4 s. Its apparent Michaelis–Menten constant is determined to be 1.78 mM, and its experimental detection limit is 7.0 M. It has also been found that the biosensor has favorable stability over a comparatively extended period of storage, good anti-interference capabilities, and great thermal stability. According to the findings, gold/amino acid nanocomposites offer a potentially useful material for the development of biosensors and other biological applications. Furthermore, the following strategy has been used to prepare the gold nanoparticle (GNP) superstructure; in a typical experiment, 0.4 mL of a freshly made, ice-cold 30 mM solution of NaBH4 was added immediately to 25.0 mL of 0.25 mM HAuCl4 in the presence of 30 mM glutamic acid. Upon adding NaBH4, the HAuCl4 was instantly converted into gold nanoparticles, as evidenced by an instantaneous shift in the color of the solution from colorless to red. By bringing the pH of the system to 5.0 and adding 1.0 M HCl to the GNP solution, the gold superstructure was established. When the remainder of the solution turned white after being left overnight, sedimentation of nanocomposites was completed. After removing the supernatant solution, the sample underwent three water washes before being allowed to dry in the open.74 Finally, the prepared modified electrode has shown improved bioanalytical responses toward the target analyte, which is due to the presence of gold nanoparticles.
The dihydrofolic acid reductase enzyme was used to develop a novel amperometric biosensor. An AuE surface was electrodeposited with c-MWCNT and TiO2 NP nanocomposites. DHFR, c-MWCNTs, TiO2NPs, and AuE were used as functional electrodes to build the suggested amperometric biosensor. These composites have TiO2 acting as a binder and MWCNTs providing catalytic characteristics. One-dimensional carbon nanotubes are characterized by their large surface area and unique combination of mechanical, chemical, physical, and biological capabilities. Because of their high thermal and electrical conductivity, CNTs have garnered significant attention from researchers studying nanomaterial chemistry in recent years. This newly developed biosensor was used to measure the FA levels in expectant mothers. The biosensor responded best at a pH of 7.5 and a temperature of 35 °C. Electrons generated at 0.125 V against the Ag/AgCl electrode were responsible for the maximal current observed in the developed FA biosensor. The suggested biosensor displayed good sensitivity, a broad linear range, and a low limit of detection (LOD).75
An innovative bioelectrode for developing a laccase-based amperometric biosensor was achieved through the immobilization of laccase enzymes and the modification of carbon paper electrodes with MoS2, which is a two-dimensional material (Figure 1.6). With a sensitivity of 108.3 nA µM−1 × cm2 and an LOD of 0.2 µM among the lowest values recorded for this oxidoreductase-based biosensor, the resultant bioelectrode demonstrated good analytical efficacy in the assessment of acetaminophen (ACE) in the citric acid buffer. Moreover, MoS2 was found to enhance sensitivity. It is crucial to look into enzymes with higher detection yields, as the investigation into differences in the nature of laccase enzymes also revealed a significant impact on the development of novel bioelectrodes. The use of bioelectrodes to measure ACE in groundwater samples was accomplished successfully, and LacII performed well despite naturally occurring inhibitors (major ions). This yielded a linear range of 1–98.04 µM or 151.13–14 819.9 µg L−1 and a limit of detection of 0.5 µM. The characteristics of this bioelectrode can be efficiently utilized to develop a monitoring instrument for detecting significant increases in guideline values, defined as 200 µg L−1, to investigate water contamination caused by this medication.76
Electron transfer mechanism of the developed CP-MoS2 Lac bioelectrodes for the detection of acetaminophen. Reproduced from ref. 76, https://doi.org/10.3390/s23104633, under the terms of the CC BY 4.0 license, https://creativecommons.org/licenses/by/4.0/.
Electron transfer mechanism of the developed CP-MoS2 Lac bioelectrodes for the detection of acetaminophen. Reproduced from ref. 76, https://doi.org/10.3390/s23104633, under the terms of the CC BY 4.0 license, https://creativecommons.org/licenses/by/4.0/.
In the framework that was developed, the two-dimensional structure of the boron nitride layer allowed uricase to interact with B-N groups. Thus, uricase was bound to B-N groups to form a distinct catalytic structure that was then shielded by Nafion. The developed enzyme-based biosensor demonstrated a low limit of detection and quantitation values (0.14 µM and 0.46 µM, respectively) for uric acid, along with a broad linear detection range (5–3000 µM). Additionally, the developed system demonstrated outstanding shelf-life features, low Km values, and great selectivity, sensitivity, repeatability, and reusability. Additionally, with a high recovery rate, an enzyme-based biosensor has successfully detected uric acid in commercial human serum. It has been advertised as a quick, highly sensitive, and accurate serum application for measuring uric acid levels.77
One useful marker for the early identification of diabetes is 1,5-anhydroglucitol (1,5-AG). This study built a high-performance light-addressable potentiometric sensor (LAPS) for the detection of 1,5-AG using pyranose oxidase (PROD) as the specific recognition substance and reduced graphene oxide–polyacrylamide–ferrocene/gold nanoparticles (rGO–PAM–Fc/AuNPs) as the sensing membrane. The rGO–PAM–Fc/AuNPs modification LAPS chip is one of the best at adsorbing more PROD and efficiently increasing the specific surface area and is shown in Figure 1.7. It additionally has the ability to enhance the detection sensitivity and the capacitive impact of LAPS. To convert divalent ferrocene ions in the rGO–PAM-Fc composite to trivalent ferrocene ions, 1,5-AG must be hydrolyzed by PROD. This hydrolysis results in hydrogen peroxide (H2O2), which shifts the I–V curve and changes the membrane potential on the surface of the LAPS chip, reflecting the change in 1,5-AG concentration. The proposed LAPS demonstrated satisfactory linearity with 1,5-AG concentration in the range of 100.0–1000.0 μg mL−1. The LOD was found to be 21.74 μg mL−1, and the sensitivity was calculated to be 0.1908 mV μg−1 mL−1.78
A sensing scheme of 1,5-anhydroglucitol (1,5-AG). Reproduced from ref. 78 with permission from Elsevier, Copyright 2023.
A sensing scheme of 1,5-anhydroglucitol (1,5-AG). Reproduced from ref. 78 with permission from Elsevier, Copyright 2023.
An essential energy source and metabolic intermediary for both plant and animal cells includes glucose (Glu). In this investigation, a LAPS and a specific sensitivity unit were combined to develop a high-performance biosensor for the selective recognition of Glu. The particular sensitive unit used a silicon-based capacitive framework, and a nanosensing membrane was developed by modifying the glucose oxidase/reduced graphene oxide–chitosan–ferroxidase/gold nanoparticles (GOX/RGO–CS–Fc/AuNPs) on the silicon-based working electrode. The combination of the particular sensitive unit and Glu produced a bilayer potential, and the resulting field effect altered the capacitance of the silicon structure, thereby modifying the photocurrent–voltage characteristics of the LAPS. The amount of Glu might be detected by the LAPS due to a potential shift in the photocurrent–voltage parameters. When all conditions were optimal, the potential shift and the Glu concentration (from 0.01 mg mL−1 to 4.00 mg mL−1) exhibited a linear relationship.79
Although methanol, or MeOH, is utilized as a solvent and cleaning agent in industry, it is toxic if consumed. For MeOH vapor, a release threshold of 200 ppm is advised. By grafting alcohol oxidase (AOX) onto electrospun polystyrene–poly(amidoamine) dendritic polymer-blend nanofibers (PS–PAMAM–ESNFs) on interdigitated electrodes (IDEs), researchers developed a unique, highly sensitive micro-conductometric MeOH biosensor. Using gaseous MeOH, ethanol, and acetone samples taken from the headspace above an aqueous solution with known concentrations, the analytical performance of the MeOH microsensor was assessed. As concentrations increase, the response time (tRes) of sensors varies from 13 to 35 s. The conductometric sensors possess a detection limit of 100 ppm in the gas phase and a sensitivity of 150.53 μS cm−1 (v/v) for MeOH.80
Using cell-imprinted polymer (CIP)-functionalized microwires (CIP-MWs), which have an affinity for E. coli, researchers were able to quantify the conductometric transduction of CIP–bacteria-binding events in a low-cost microfluidic sensor. The sensor was composed of two CIP-MWs placed opposite each other across a PDMS microchannel. The inter-wire electrical resistance of the microchannel was taken into account before, during, and after the exposure of CIP-MWs to bacteria. After the bacteria were incubated for 30 min, there was a decrease in the inter-wire resistance of the sensor. The limits of detection and quantification were 2.1 × 105 CFUmL−1 and 7.3 × 105 CFU mL−1, respectively, as determined by resistance change normalization and the subsequent analysis of the dose–response curve of the sensor across 0 and 109 CFU mL−1 bacteria. The bacterium counts could be statistically distinguished from one another within the dynamic range of the sensor, which was 104 to 107 CFU mL−1. The sensitivity of 7.35 μS CFU−1 mL−1 was obtained via a linear fit within this range. The selectivity of the sensor for the transfected E. coli cells was demonstrated in experiments utilizing competing Sarcina or Listeria cells.81
1.2.2.2 Optical-based Biosensors
The state-of-the-art combination of biotechnology with optics in optical biosensors provides a flexible and potent framework for biological molecule recognition and evaluation. These sensors are transforming various industries, including ecological surveillance and medical diagnostics, by providing real-time, label-free, and highly sensitive detection of biomolecular interactions by utilizing the principles of light–matter interactions. Transforming biological processes into quantifiable optical signals is the fundamental function of an optical biosensor. These sensors usually comprise a transducer surface that has an immobilized recognition element (e.g., enzymes, aptamers, or antibodies) on it. In real time, the optical properties of the system can be recognized and measured as a result of the interaction between the target analyte and the identification element.
One of the most extensively researched and applied types of biosensors is probably surface plasmon resonance (SPR). SPR biosensors are characterized by high specificity, sensitivity, and affordability. Figure 1.8 depicts a conventional surface plasmon resonance biosensor, with a gold layer on a prism and incident light reflected onto a detector.83 Free electrons at the metal–dielectric contact oscillate in their surface charge density, which results in surface plasmon resonance. Evanescent waves can be produced by an incident light beam that efficiently excites it. The incident optical signal is linked to the metal–dielectric surface using a glass prism with a high refractive index via the prism-based method. The sensitivity of the equipment can be determined by measuring the intensity of light reflected from the surface. For biosensing, the change in the refractive index at the detecting region/sensor surface is assessed as a change in the resonance angle.84,85
A typical surface plasmon resonance biosensor. Reproduced from ref. 82 with permission from the Royal Society of Chemistry.
A typical surface plasmon resonance biosensor. Reproduced from ref. 82 with permission from the Royal Society of Chemistry.
For instance, the performance characteristics of two types of optical fiber biosensors based on localized surface plasmon resonance (LSPR) and using gold nanospheres (GNSs) and gold nanorods (GNRs) were effectively assessed and cross-compared, and a schematic representation is shown in Figure 1.9. To circumvent the constraints of other intensity-based sensors, the authors of this study specifically chose to fabricate two representative sensor probes: GNSs with a diameter of 60 nm and GNRs with an aspect ratio of 4 : 1. These choices were made based on the outcomes of optimization obtained from each of these types of biosensors, which they have previously reported. To facilitate the development of efficient LSPR biosensors, GNSs and GRNs, respectively, were immobilized onto an exposed optical fiber surface before being functionalized with human IgG to produce a device that could detect anti-human IgG at varying concentrations. According to the test results, LSPR sensors based on GNSs and GNRs exhibit sensitivity levels of 914 and 601 nm/relative orientation unit (RIU), respectively, based on variations in refractive index. Nevertheless, as biosensors, they have proven to have the same detection limit of 1.6 nM when it comes to detecting anti-human IgG.86
Schematic diagrams of the structure of (a) GNS-coated and (b) GNR-coated LSPR optical fiber sensor probes. Reproduced from ref. 86 with permission from Elsevier, Copyright 2013.
Schematic diagrams of the structure of (a) GNS-coated and (b) GNR-coated LSPR optical fiber sensor probes. Reproduced from ref. 86 with permission from Elsevier, Copyright 2013.
To detect the DNA hybridization process, a D-shaped plastic optical fiber (D-POF) SPR biosensor based on the graphene/Au film (G/Au) was developed and practically shown. By directly evaporating the Au film onto the graphene formed on copper foil using the physical vapor deposition (PVD) process, the Au film served as a substitute for conventional polymethyl methacrylate (PMMA) and enhanced the detection performance of SPR sensors. Due to this technique, graphene and Au film formed a smooth contact. Subsequently, the G/Au was jointly deposited onto the D-shaped fiber. Using the finite-element technique (FEM), the authors investigated the G/Au SPR sensor and found the ideal material thickness for the configuration. The suggested sensor demonstrated a significant improvement in sensitivity over previous plastic optical fiber tests, with a computed value of 1227 nm/RIU in various glucose solutions. By monitoring the resonance wavelength shift, the suggested sensor concurrently detects single-nucleotide polymorphisms (SNPs) and hybridization with success. Additionally, it demonstrates a good linear response to the target DNA molecules at corresponding concentrations ranging from 0.1 nM to 1 µM, demonstrating the broad diagnostic potential of this technique.87
The research team has put forth a strict design for an SPR sensor coated with graphene that combines tungsten disulfide (WS2) to detect DNA hybridization. Gold (Au), WS2-graphene, prism (SF10 glass), and sensing medium make up the current setup. Researchers evaluate the performance characteristics of the suggested sensor in terms of sensitivity, precision, and quality. Here, scientists report a significant improvement in the overall result. The sensitivity is increased by adding graphene layers, but the other performance metrics are decreased. We apply WS2 between the metal and graphene layers to improve all performance characteristics. Additionally, this study analyzes the effect of the thickness of gold (Au). According to a numerical study, the difference in the SPR angle between complementing and mismatched DNA strands is significantly countable, but the difference between mismatched DNA strands is almost nonexistent. As a result, the suggested biosensor creates a new avenue for biomolecular interaction monitoring.88
Researchers have suggested a nanolayer biosensor enhanced with silicon and MoS2, based on surface plasmon resonance. First, researchers optimized the thickness of MoS2 to examine the sensitivity of the arrangement. Next, using the transfer matrix method, scientists computed the change in the resonance angle for a fixed 0.005 refractive index change of the sample layer based on the adjusted number of MoS2 layers. They also determined the optimal values by considering the maximum change in the resonance angle and the full width at half maximum (FWHM). Optimal results are obtained with a light wavelength of 633 nm, one layer of MoS2, and silicon and gold thicknesses of 35 nm and 7 nm, respectively. According to the final analysis of researchers using the optimized parameters, the suggested MoS2-enhanced architecture offers approximately 10% more sensitivity than the non-enhanced version. Based on the methodical computational work on the MoS2-enhanced SPR configuration, researchers believe that the MoS2 nanosheet will perform well in upcoming experimental investigations as well.89
To effectively detect DNA hybridization, a highly sensitive Au–MoS2–graphene hybrid-based SPR sensor is explored quantitatively. The suggested development demonstrates increased sensitivity by sandwiching an MoS2 monolayer between the traditional graphene-on-Au SPR sensor. The enhanced sensitivity of the suggested sensor is 89.29 and 87.8 degrees RIU−1 for single and double hybrid layers of graphene–MoS2, respectively. Compared to SPR sensors without an MoS2 layer, the sensitivity of the proposed sensor with a single MoS2 sensing layer is improved by roughly ten percent. Finally, applications for the suggested ultrasensitive biosensor include safety in food, environmental surveillance, healthcare diagnosis, enzyme recognition, and DNA hybridization analysis.90
Furthermore, the effect of the ZnO layer on the sensitivity parameters and electric field intensity enhancement factor (EFIEF) of graphene-based SPR biosensors is analyzed through a theoretical simulation using the angular interrogation approach. This SPR biosensor uses graphene, gold, and ZnO to improve the quality of the SPR signal, with molecular recognition sites binding tightly to biomolecules. This is the main characteristic of the sensing device. In comparison to graphene biosensors without ZnO, those with ZnO are far better designed. The sensitivity significantly increases with the addition of a ZnO layer at the base of the prism. Compared to other reported classical SPR biosensors for detecting bacteria similar to Pseudomonas, the suggested sensor exhibits higher sensitivity of 187.43 degrees RIU−1, detection accuracy of 2.05 degrees−1, a quality parameter of 29.33 RIU−1, and enhanced EFIEF. Conversely, as the number of graphene layers increases, reflectivity also increases for a fixed value of wavelength and refractive index. We have ensured that the suggested biosensor exhibits optimal sensitivity while maintaining the lowest feasible reflectance. Scientists anticipated that the best substitute for uses in biosensing would come from a graphene biosensor with a ZnO layer at the base of the prism.91
The use of fixed solid substrates immobilized with capture probes in affinity-based fluorescent biosensing systems for biomarker monitoring is limited, making it difficult to use these systems for either intermittent or continuous biomarker analysis. Additionally, there have been challenges in combining low-cost fluorescence detectors and microfluidic chips with fluorescence biosensors. Here, scientists showcased a mobile, highly effective fluorescence-enhanced affinity-based fluorescent biosensing technology that combines digital imaging with fluorescence amplification to overcome current limitations. Digital fluorescent imaging-based aptasensing of biomolecules with stimulated signal-to-noise ratios was achieved using fluorescence-enhanced moveable MBs embellished with zinc oxide nanorods (MB–ZnO NRs). By grafting bilayered silanes onto ZnO NRs, photostable MB–ZnO NRs with high stability and homogenous dispersion have been achieved. When contrasted with MBs without ZnO NRs, the fluorescence signal was enhanced by up to 2.35 times when ZnO NRs were produced on MBs. Furthermore, continuous biomarker assessments in an electrolytic environment were achieved by the incorporation of a microfluidic device for flow-based biosensing. The outcomes demonstrated the great potential of highly stable fluorescence-enhanced MB–ZnO NRs combined with a microfluidic platform for biological experiments, continuous or sporadic biomonitoring, and therapeutics.92
Long noncoding RNAs (lncRNAs) are essential for a wide range of physiological and pathological functions. Their abnormal expression can disrupt the normal regulation of gene expression, leading to various human disorders. This study developed a fluorescent light-up biosensor with minimal background for label-free lncRNA recognition by combining transcription-driven fluorogenic RNA aptamer-Corn synthesis with duplex-specific nuclease (DSN)-assisted target recycling amplification. The two linear probes that the authors developed are a linear template for RNA aptamer-Corn transcription and a capture probe for starting a cyclic cleavage sequence. Capture probes developed on MB surfaces bind to the target lncRNA to initiate a DSN-assisted cyclic cleavage reaction that releases a large amount of T7 promoter motifs. The process of producing fluorogenic RNA aptamer-Corns, which can light up small-molecule fluorogens such as 3,5-difluoro-4-hydroxybenzylidene-imidazolinone-2-oxime (DFHO), involves the free T7 promoter hybridizing with a linear template following magnetic separation. This transcription amplification is then efficiently induced with the help of T7 RNA polymerase. Notably, the use of MBs makes it easier to isolate and enrich target lncRNAs from the intricate biological matrix as well as to separate cleaved capture probes. This biosensor can detect lncRNA with a detection limit of 31.98 aM. It does this by making use of the high signal-to-background ratio of the Corn–DFHO complex and the high efficiency of DSN/T7 RNA polymerase-mediated cascade proliferation.93
1.2.2.3 Mass-based Biosensors
A class of analytical instruments, known as piezoelectric biosensors, operate on the basis of monitoring binding interactions. A sensor component that operates on the basis of changes in oscillations caused by mass bound on the surface of a piezoelectric crystal is known as a piezoelectric platform or piezoelectric crystal. Mass-based biosensors, known as piezoelectric biosensors, generate an electrical signal in response to the application of mechanical pressure. One instance of a piezoelectric biosensor is the quartz crystal microbalance (QCM) model. In the electronics sector, QCM is a widely used instrument. These instruments typically have a fundamental mode frequency of 1–20 MHz and are currently used as attenuators in electrical circuits. Although QCM with high frequencies offer significant prospects for a sensitive assay, various disadvantages have been reported, including their fragility and the requirement for technologically demanding equipment for their fabrication. The QCM sensor was developed using quartz crystal that has metal electrodes attached to it as its basic material.22
In this regard, using a quartz crystal resonator (QCR), a novel mass-sensitive biosensing method for identifying circulating tumor cells (CTCs) has been developed. To remove the Gaussian spatial distribution of response time in the first harmonic mode—a feature of QCRs—without sacrificing frequency response sensitivity, a ring electrode-based QCR was designed using a mathematical model. The ring electrode that was made using our model was validated using the ink-dot technique. Additionally, the capacity of the ring electrode QCR to collect circulating tumor cells was examined through experimentation, and the outcomes were contrasted with those of a QCR that is sold commercially and has keyhole electrodes. Through the use of silane, protein, and anti-EpCAM to modify the SiO2 surface of the ring electrode, an indirect way of surface immobilization technique was used. Compared to the keyhole QCR, which has shown nonuniform spatial responsiveness for the same cancer cell lines, the ring electrode significantly reduced the spatial nonuniformity of frequency response across three cancer cell lines: MCF-7, PANC-1, and PC-3. These findings hold potential for the development of QCR-based biosensors for cancer screening point-of-care diagnostics and early identification of cancer cells.94
Trichothecenes, zearalenone, and fumonisins are among the mycotoxins produced by Fusarium species that can co-contaminate food and feed across the whole supply chain, including cereal grains and animal feed. The need to better monitor mycotoxins across our food supply chain to improve global food security is growing. Rapid assays that can identify several poisons from a single sample are useful for time- and cost-efficient analysis. The aim of this research is to evaluate the viability of using both portable and non-portable mass-based biosensors for multiplex mycotoxin detection, taking into account the current trends in mycotoxin testing. Solidly mounted resonator (SMR) technology was utilized to develop miniaturized 4 × 16 mass-sensitive transducer pixels for the biosensor, known as a mass-sensitive microarray (MSMA). For the purpose of simultaneously and semi-quantitatively detecting three controlled mycotoxins—zearalenone (ZEN), fumonisin B1 (FumB1), and T2-toxin (T2)—individual pixels on the sensor surface were functionalized utilizing nano-spotting methodology. Competitive inhibition tests were developed following the integration of singleplex and multiplex calibration curves with portable and non-portable microfluidic devices for antibody and standard sample injections. Sensitivity, defined as the concentration that results in a 50% inhibition, was one of the features and performance of the MSMA that was assessed. For T2, FumB1, and ZEN, the corresponding sensitivity of singleplex assays conducted with the portable microfluidic device (PMD) was 1.3 ng mL−1, 2.0 ng mL−1, and 6.8 ng mL−1. The sensitivity of the multiplex assay using the PMD was 6.1 ng mL−1, 3.6 ng mL−1, and 2.4 ng mL−1 for T2, FumB1, and ZEN, respectively. The multiplex detection of three controlled mycotoxins has been demonstrated using the PMD, an extremely sensitive and user-friendly screening method. Real-time analysis was performed, and the output was entirely digital. The suggested technique of mycotoxin analysis was label-free and cost-effective, as the analyte detection was based on mass.95
A QCM electronic oscillator circuit was used to design a highly sensitive DNA biosensor. The oscillator circuitry had been modified to meet the Barkhausen requirement, despite quartz being submerged in a liquid medium and so having extremely poor quality factors. To ascertain the resolution of sensors, a study of the frequency noise of the developed QCM system was conducted. Under dynamic conditions (with liquid circulation), a mass resolution of 7.1 ng cm−2 was established. The DNA biosensor activity of the QCM oscillator has been shown. The high selectivity and sensitivity of the system in identifying complementary target DNAs in a solution are demonstrated by the results. The sensor can detect DNA targets at concentrations of more than 50 ng mL−1.96
Alpha synuclein protein (SYN alpha), a putative biomarker of Parkinson’s disease, was recently subjected to a sensitive, selective, and useful immunosensor developed using a quartz tuning fork (QTF) and an inventive biosensing fabrication technique. Two processes were involved in the functionalization of gold-coated QTFs: conjugation of gold nanoparticles (AuNPs) and the formation of a self-assembled monolayer containing 4-aminothiophenol (4-ATP). According to resonance frequency changes associated with a detection limit of 0.098 ng mL−1, the proposed biosensor system has a selective determination range of 1–500 ng mL−1 for SYN alpha.97
Using an enzymatic amplification procedure in conjunction with a composite coupling of a DNA probe, GNPs, and dendrimers, a mass-sensitive microRNA sensing surface is developed. Through covalent binding between the NH2 group in PAMAM or DNA and the COOH group on GNP, the DNA probe/GNP/dendrimer composite is developed. Target microRNA forms a heteroduplex when it attaches to DNA on magnetic NPs that has a stem–loop architecture. An abundance of linker sequences is generated on MNPs through amplified enzyme recycling. The assembly of DNA/GNP/dendrimer probe on the surface of gold chip, resulting in a sensitive frequency response, was achieved through the combination of a DNA probe, linker DNA on MNPs, and capture DNA on gold chip. With a detection linear range of 1.0 × 10–11 to 1.0 × 10–9 M, it delivers a very good microRNA-203 quantification capability. Future applications in early cancer diagnosis are suggested by the ability to distinguish between different expression levels of miR-203 in MCF-7 cells and the consistent results with conventional qRT-PCR measurement methodology.98
A novel approach was put forth by researchers to streamline the immobilization of bioactive compounds over a mass-sensitive transducer by using bovine serum albumin (AL-BSA) nanofibers as an infrastructure sensing layer on QCM surfaces (Figure 1.10). Consequently, scientists eliminated all synthetic or polymeric processes and used an entirely natural, incredibly thin substance with self-functional amino groups in the surfaces. The ideal settings for electrospinning were suggested to be an applied bias of 18 kV, a flow rate of 0.3 mL h−1, and a tip-to-collector distance of 10 cm. We were able to measure the frequency responses caused by the immobilization of the model Lys protein utilizing batch and continuous procedures with good sensitivity and reliability, thanks to the GA-activated ES/GA-QCMs, which eliminated the need for a further chemical functionalization phase. In the process of immobilizing lysozyme, the greatest frequency shifts and the least relative standard deviation were recorded as 1287.0 ± 176.2 Hz and 13.6% for the 60 s collected ES/GA-QCM nanofibers, respectively. Because of the sluggish reaction kinetics of the amino and aldehyde groups, frequency measurements revealed that the binding efficiency in the batch approach is nearly ten times higher than in the continuous procedure for all exposure times. Because biologically active molecules can attach to a natural matrix and the wet-chemistry step for immobilization is omitted, the current coating technique can be used to enhance the effectiveness of any type of immunosensor.100
A schematic diagram illustrating the electrospinning deposition of amyloid-like BSA nanofibers on the QCM gold electrode surface. Reproduced from ref. 99 with permission from Elsevier, Copyright 2017.
A schematic diagram illustrating the electrospinning deposition of amyloid-like BSA nanofibers on the QCM gold electrode surface. Reproduced from ref. 99 with permission from Elsevier, Copyright 2017.
1.3 Future Prospective
Intensive efforts must be made to develop successful immobilization strategies to ensure greater sensitivity and stability of biosensors. The challenges related to the denaturation of enzymes and the specificity and reproducibility of signals are of prime importance in the fabrication of biosensors. Another area of development of biosensors is the determination of pesticides, which will greatly benefit agriculture, food processing, and forensic analysis. The design of microbial and photosystem-based biosensors, along with electrochemical biosensors for target specificity, and also a significant growth of aptamers in biomedical applications in the near future are much appreciated.
1.4 Conclusions
Biosensors are the most significant and economically viable devices that humans have invented. The utilization of various biorecognition components in biosensors has expanded. The broad application of biosensors in the fields of food and agriculture, healthcare management, and drug and toxicity studies, especially in residual pesticide analysis, has become a noteworthy turning point. Nonetheless, to achieve high sensitivity, strong stability, and repeatability in a novel biosensor while maintaining its activity, efforts must be made to develop a dynamic immobilization technique. Their practical applications in clinical diagnostics and disease monitoring can now be elaborated, thanks to these breakthroughs.
Acknowledgements
The authors, Suma B. P. and Prashanth S. A., dedicate this chapter to Dr. M. Pandurangappa on his retirement. Additionally, Prashanth S. A. extends this dedication to Dr. P. S. Kandagal on his retirement.