A resonant biosensor is generally used to detect bio-relevant molecules in aqueous media. The sensing focuses on the change of mass, viscosity, or conductivity of the substrate surface. Many biosensors fall into this category, which includes surface acoustic wave sensors, magnetoelastic sensors, quartz crystal (piezoelectric) biosensors, etc.

The prototype of a surface acoustic wave sensor (SAW) was established based on a strongly confined acoustic energy detected by interdigitated transducers [4,19,20]. A SAW sensor has one piezoelectric material positioned in between two transducers. An input transducer electrically excites acoustic wave of the piezoelectric material, in which the acoustic wave is received at the output transducer. The wave energy radiates into the aqueous bulk due to the perpendicular displacement of the wave propagation in aqueous environment. This causes a high attenuation of the received signal, which hinders the application transferred into a biosensor [4,21,22]. The shear-horizontal (SH) surface wave, generally referred to as a Love wave, is generated using a deposited elastic layer to guide the direction of the acoustic resonance. The elastic material significantly reduces the spreading loss of acoustic energy [23–25]. Now the operation of a SAW-based biosensor is driven by the coupled wave transducer and antibody on a piezoelectric substrate. The antibody used as a bio-element is immobilized on the device that catches analytes from the aqueous medium. The bonded analytes will change the velocity of the SAW, which alters the output signal generated by the integrated electronics. The variation of the output signal can be used to evaluate the concentration of the analytes. The schematic setup of a SAW-based biosensor is presented in Figure 2 [26].

The theoretical principal of quartz crystal biosensors relies on the piezoelectric effect of the crystal. When a piezoelectric material is subjected to an AC potential, a mechanical oscillation of the material is excited. The frequency change of such oscillation is correlated to the mass change on the material surface. The mathematical relationship (Equation (1)) between the resonant frequency and the mass variation was firstly described by Sauerbrey [27]: (1) Δ f m = f 0 2 F q ρ q m s where f₀ is the resonance frequency of the quartz resonator, F_q is the constant of crystal frequency, ρ_q is the material density, and m_s is the mass/area. In the quartz biosensor, the adsorption of target analytes onto the quartz surface causes the shift of resonant frequency that can derive information for quantitative analysis.

The operation of magnetoelastic sensors relies on the mechanical vibration of a magnetoelastic material when the material is subjected to a magnetic field. Such responses of the sensor not only can be detected acoustically, but also magnetically. The vibration generates an elastic wave within the magnetoelastic material that results in a detectable magnetic flux [28]. In a magnetoelastic sensor, the vibrational frequency (f) can be expressed as[29]: (2) f = 1 2 L E ρ where L is the length, E is elasticity, and ρ is the density. When a magnetoelastic sensor is subjected to a mass change (Δm), the variation of frequency can be expressed as[29]: (3) Δ f = − f 0 Δ m 2 m 0 where f₀ is the initial resonance frequency. When the sensor is subjected to a damping effect created by the surrounding medium, the shift in resonance frequency can be expressed as [30]: (4) Δ f = π f 0 2 π ρ s d η ρ 1 where η is viscosity of medium, s_l is the density of medium, ρ_s is the density of sensor material, and d is the thickness of the sensor. At present, the characteristics of magnetoelastic materials have been employed in different kinds of bio-sensing such as glucose, avidin, Escherichia coli, and platelet-rich plasma [30–33], among others.

Surprisingly, even with their poor reputation for weak sensitivity and non-specific heating effects, the applications of thermal biosensors still draws considerable attention. A thermal biosensor is a promising analytical tool due to the following advantages: -

No chemical contact between transducer and sample leading to long-term stability.

Novel thermal biosensors based on enzymatic conversion have been developed for monitoring enzyme reactions. This is mainly because most biochemical reactions have an exothermic character. The principle of thermal biosensor measurement is based on the first law of thermodynamics: (5) Q = − n p ∑ Δ H (6) Q = Cp Δ T where Q is the total energy of heat, H is the enthalpy, Cp is the heat capacity. The measureable local temperature shift (ΔT) generated by the heat production is dependent on the heat capacity of the system [34]: (7) Δ T = − Δ Hn p / Cp

Generally, the Cp of most organic media is lower than that of the aqueous solvent [35]. The sensitivity and detection limit of the sensor are determined by the organic solvents. Figure 3 shows a schematic diagram of the principle set-up for an enzyme thermistor (ET). The thermostated box controls the physiological temperature. Samples and the buffer are injected in the ET where the aluminum thermostates the buffer stream. The heat generated by the enzymatic conversion reduces thermistor resistance and the bridge amplifier registers the signal.

In this type of biosensor, the measured output transduced signal is the light intensity. More than 75% of the research papers for optical biosensors focus on using the surface plasmon resonance (SPR) [37] effect. The SPR essentially is a diffraction anomaly due to the surface excited plasma [38]. The electrons resonate when the wavelength of oscillating mobile electrons (plasma) matches the wave vector of incident light. The resonating plasma is associated with the electromagnetic waves propagating in a direction parallel to the interface of two media, and decaying evanescently, i.e., evanescent wave. Due to the limited propagation length of a surface plasma wave (SPW), the detection of SPR sensor is conducted where the SPW is excited by the incident light source. A SPR sensor is constituted by an incident light, a transducer with a gold side contacted with the detection apparatus and the other side contacted with microfluidic system (flow side), and an electronic system for processing the output signal. Fixed wavelength is shot to the gold side and is reflected, which induces an evanescent wave penetrating into the flow side. During the measurement, the analyte is introduced through the microfluidic channel and bound with the sensor, which changes the dielectric constant of the medium. This will lead to the changes of refractive index near the surface hence affecting the refracting SPR angle (Figure 4) [39–41].

The sensing mechanims of other photometric biosensors are based on the optical properties of analytes, which include absorbance/scattering and fluorescence. The utilization of the absorbance properties of analytes is to detect the amount of laser light being blocked or transmitted by the target cells [43,44]. Such a measuring system provides a rapid, simple, repeatable, and label-free assay for the immobilized analytes. Due to the limited sensitivity of the conventional methods in absorbance studies, an optical waveguide has been introduced in order to enhance the efficiency of absorbance biosensor. Optic fiber probes with different geometries have been employeed in absorbance biosensors that include coiled and tapered fibers, among others [45,46]. The fluorescent staining technique is also introduced in photometric biosensors. Such a technique is usually requred in microfludic-platform biosensors for the purpose of bacterial counting [47].

The development of ISFETs, usually used for pH and ion concentration measurements [48–50], started in the 1970s. To integrate the sensing circuit, most ISFETs were produced through MEMS fabrication based on silicon substrates. This also makes ISFETs a perfect transducing element for biosensing. The reactive silanol (SiOH) groups on the SiO₂ surface provide a stem for covalent attachment of bio-molecules by using H⁺ and OH⁺ as binding sites: SiOH ‐ ‐ ‐ ‐ ‐ ‐ ‐ ‐ ‐ ‐ ‐ SiO − + H + SiOH + H + ‐ ‐ ‐ ‐ ‐ ‐ ‐ ‐ ‐ SiOH 2 +

Through silicon surface modification, the analytes can be successfully immobilized onto the gate surface [51–55]. Figure 5 outlines the schematic configuration of ISFETs.

An ISFET consists of a sensing electrode coated with a polymer selectively permeable to analyte ions, and a field effect transistor structure. The current flow in the gate voltage is regulated by the potential difference between source and drain. When the analyte ions diffuse through the ion-sensitive polymer, the charged biomolecules cause the depletion or accumulation of charge carriers. This will lead to the potential difference at the detecting interface. This interfacial potential difference is regulated by the ions concentration in the solution; therefore, the change of the current in the transistor is a measure of analyte concentration.

The constitutional concept of an electrochemical biosensor is based on a matrix bound bioactive material coupled with an electrochemical transducer. Essentially, it is a surface modified electrical conductor for different electrochemical functions. This type of sensor targets those biological reactions that derive ionic production and consumption. This will cause the charge transfer across the double layer of the physio-chemical transducer that generates the measureable signal [56–58]. Based on the measured signal characterizations, the electrochemical biosensor has three main classifications of silicon based chips. Figure 6 is the illustration of the electrochemical biosensor.

By definition, SAM is a single layer of biological and/or chemical molecules formed through self-assembly. It is a part of molecular nanotechnology that attracts interest due to its simple production and versatility in molecular and reaction selection. The first study on SAM was reported in 1946 by Bigelow et al. who used a metal surface to absorb surfactant molecules and made a layer with monomolecules [63].

Under suitable conditions, SAMs can be directly formed on a sensor substrate. A substrate is placed in a solution containing assembling molecules, where SAM forms spontaneously. In some cases, UV irradiation, heat, voltage, and other conditions need to be applied in order to complete SAM formation. Self-assembling molecules contain three parts: the head group, which is open for bond formation with surface substrate; the alkyl chain, which is important in stabilization and order, interacts with neighboring chains through van der Waals and electrostatic forces; the functional terminal, which is open for functionalization and is used to determine the chemistry properties of SAM surface with the outer environment [64]. The idealized SAM system is shown in Figure 7(A).

As shown in Figure 7(B), there are four phases during the SAM formation, including vapor phase, intermediate phase, low density phase, and high density phase. In the vapor phase, the surface is randomly deposited with isolated molecules. In the intermediate phase, the surface is deposited by adsorbate molecules with disordered conformation. In the low density (liquid) phase, adsorbate molecules are lying down on the surface. In the high density (solid) phase, molecules are standing in order and packed with tilting angle less than 30°. SAM formation would first go through the vapor phase when the temperature is lower than the triple point, followed by the intermediate phase, and the high density solid phase last. Otherwise, the formation would incorporate a low density liquid phase if the temperature is higher than the triple point. Both processes involve the formation of solid phase islands surrounded by isolated vapor phase molecules and the nucleation and growth of these islands till the entire surface is covered. The Langmuir model Equation (8)) shown below is used to explain SAM formation: (8) d θ d t = k ( 1 − θ ) where θ is the proportional coverage of the surface, and k is the rate constant [65]. The equation indicates that the SAM formation rate is proportional to the uncovered surface. The initial formation of alkanethiol SAMs is 80–90% complete during the first several minutes, and molecules reorganize themselves over 12–16 hours [66]; however, changes may still happen weeks later [67].

Increasing numbers of SAM systems are presented due to the accumulating knowledge in chemistry. SAM systems are separated into four types based on their chemical characteristics (Table 2). They are alkanethiol and organosulfur on a metal surface, organosilane on a hydroxylated oxide surface, alkyne and alkene on a hydride-silicon surface, and aryl diazomium salts on a carbon, metal, metal oxide, silicon surface, etc. In the following sections, they will be discussed in detail.

Biomolecules can be attached to the functional terminals of modified electrodes by covalent and non-covalent bonds, as summarized in Table 3. Non-covalent bonds, which includes hydrogen bonds and electrostatic interactions, are widely applied in attachment of biomolecules. The attachment is relatively weak compared to a covalent bond. Nevertheless, it only needs simple reaction steps and usually is reagentless. Covalent bonds provide stronger immobilization, but are restricted to certain reactions.

Graphene was first discovered by Novoselov in 2004 [109]. The superior electron mobility, biocompatibility, and flexibility of graphene qualify it as an ideal material for biosensors. In a glucose biosensor application, graphene was modified by polyvinylpyrrolidone (PVP) and added to a certain ionic liquid (IL). The solution of PVP-graphene-IL was dropped on a glass carbon electrode (GCE) and was allowed to dry before detecting glucose. A linear response for glucose detection from 2 mM to 14 mM was recorded [110]. In a similar report, a graphene-chitosan solution was dropped onto a GCE and allowed to dry. Glucose oxidase was then coated on the graphene-chitosan-GCE. The glucose detection range was from 0.08 mM up to 12 mM [111]. Zeng functionalized graphene with sodium dodecylbenzenesulphonate (SDBS). Horseradish peroxidase (HRP) and SDBS-graphene self-assembled on the surface of a GCE. The resulting sensor showed high sensitivity and a H₂O₂ linear response [112]. Another graphene-based biosensor for H₂O₂ detection incorporated metallic nanoparticles. The solution, which contained graphene, HRP, and chitosan, was casted on a GCE. Au was electrodeposited on the surface of the modified GCE and clusters of Au nanoparticles were later formed. The H₂O₂ detection range was from 0.005 mM to 5.13 mM with a detection limit as low as 1.7 μM [113]. A screen printed electrode (SPE) was used as a disposable sensor system shown by Song (Figure 15). 1-Pyrenebutanoic acid succinimidyl ester (PASE) has a pyrenyl group, which interacted strongly with graphene, and a succinimidyl ester group, which reacted highly with the amines substitution. Tyrosinase functionalized Au NPs were mixed with PASE modified graphene oxide (GO). The solution of the mixture was dropped on the working SPE for catechol monitoring. The resulting sensor showed high stability and sensitivity [114].

Some graphene biosensors are electrodeless. ssDNA molecules were connected to certain dyes. GO bonded to the ssDNA labeled dye and quenched the dye's fluorescence. When the targeted ssDNA hybridized to the ssDNA-dye-GO complex, the double stranded DNA-dye was released from the GO and the fluorescence was enhanced [115]. In a biological environment, it is difficult to differentiate dopamine from its coexisting ascorbic acid (AA) and uric acid (UA) because of the overlapping voltammetric responses. A graphene nanoflake film was synthesized on a Si substrate by microwave plasma enhanced CVD. The film was shown to be capable of determining dopamine in the presence of AA and UA with high sensitivity [116].

CNTs are another group of organic materials used in biosensors due to their low detection limit and good electron transfer properties. In the case of a CNT glucose biosensor, the CNTs were conjugated with glucose oxidase and was then mixed with polypyrroles (PPy). After the electropolymerization of PPy on a GCE, the biosensor showed a linear response to the glucose up to 50 mM [117]. The CNT-sulfuric acid solution was dropped on a GCE and allowed to dry. The NADH was detected through cyclic voltammograms [118]. The mixture of CNTs, chitosan, MDB, and glutamate dehydrogenase was dropped onto the surface of a GCE and left to dry before use. Thus, the resulting glutamate sensor is sensitive and stable [119]. In the complex H₂O₂ sensor system presented in Figure 16, activated CNTs with carboxylic groups were coated on a GCE. Dopamine functionalized Pt NPs were connected to the surface of the activated CNTs [120]. Berti synthesized a CNT thin film using the CVD method and covalently connected the CNT surface with ssDNA probes. In this specific case, the ssDNA probe was inosine-modified and guanine free. After hybridization, the guanine from the target ssDNA was easily oxidized and the oxidation signal was detected [121]. A single wall CNT-FET was reported for detecting bacterial cells. The CVD method was used to generate single wall CNT connections between catalyst islands patterned on a Si substrate. The conduction of CNT was shown to drop 50% when an E. coli cell stayed on the CNT [122,123]. In a biosensor for biotin-connected molecule detection, the pyrolytic graphite electrode was oxidized in order to generate carboxylic groups. Poly-L-lysine, CNT, and anti-biotin were covalently connected to the carboxylic groups. An amperometric signal was detected when biotin-connected molecules were bound to the sensor [124]. A paste electrode was fabricated using specpure carbon, CNT, Cu₂O, and paraffin oil, which was used to differentiate amino acids with a signal to a noise ratio of 3 [125].

Conductive polymers are widely used in biosensors due to their sensitivity, selectivity, and the ability of integration for low-cost microfabrication [126,127]. In 1977, the high conductivity of halogen derivatives of polyacetylene was reported by Heeger, MacDiarmid, and Shirakawa [128]. The alternating singles and double bonds in conjugated polymers provide delocalized electrons, and the charge is carried by the same. The common conductive polymers include polyaniline, polythiophene, polyacetylene, and PPy [129]. The porous structure of polymers can be used as a bio-analyte immobilization matrix coupled with an electronic conduit. The most common technique for conductive polymer film fabrication is electrochemical polymerization, where the process is carried out in a monomer and bio-active species solution. In research done by Wallace, the negatively charged bio-active species were entrapped in the polymer during the electrochemical oxidation [127]. The enzyme, which serves as glucose oxidize, immobilized in the polymer matrix triggered the redox reaction of glucose. The electronic signal generated by the glucose oxidation/reduction can be relayed back to the detector through the conductive matrix. In a neuron recording application, the polymer SU-8 was used in the electrode fabrication in order to increase biocompatibility and to eliminate the tissue-electrode gap filled with a passivation layer [130]. In order to anchor the biotinylated proteins and the DNA, the NTA functionalized PPy has been coated on a Pt electrode. This reaction was proven to be as efficient as biotin-avidin reaction but avoiding the avidin layer [131]. In the example shown in Figure 17, either glucose hydrogenase or glucose oxidase was entrapped in PPy electropolymerized film on a graphite substrate. The linear detection limit of alkaline phosphate was 10⁻⁶ nM for glucose hydrogenase and 10⁻³ nM for glucose oxidase [132]. In another enzyme entrapment electrode, sulfite oxidase was immobilized in PPy film on a Pt electrode. The detection range for sulfite was from 0.9 to 400 mM [133]. PPy was electrogenerated on a GCE. Graphene and glucose oxidase were incorporated in the film matrix. It is shown that the enzyme-doped graphene sensor had high sensitivity towards glucose [134].