MEA is an electrochemical biosensor developed to detect the action potential (AP) in the extracellular microenvironment of cells. On an MEA, a thin metallic film is fabricated between a substrate of glass or silicon and a passivation layer with several electrode sites exposed for sensing the extracellular field potential changes generated by the objective cells.

When spreading on the microelectrodes, cultured cells adhere to the substrate. But there is still a minute volume of electrolyte between the cells and the microelectrodes; thus, a solid-liquid interface on the electrode surfaces is formed. The electrochemical properties of the interface are the basis of the sensing mechanism of MEA.

According to the electric double layer (EDL) theory, when a metal is placed into ionic liquid, an equilibrium condition is established once the charge transfer between the metal and the solution is equal. The electric field on the interface generated by electron transfer causes the formation of an inner Helmholtz plane (IHP) and an outer Helmholtz plane (OHP). The net reaction induces the creation of an electric double layer, which is also an electrified interface describing the interphase region at the electrolyte boundary [30].

The equivalent circuit of metal-electrolyte interface can be explained with the Randles model, as shown in Figure 1(a). In the circuit, an interfacial capacitance (C_I) is in parallel with charge transfer resistance (R_t) and diffusion related Warburg element (R_W and C_W). The spreading resistance (R_S) represents the effect of current spreading from the localized electrode to a distant counter electrode.

An equivalent circuit (Figure 1(b)) of the signaling pathway illustrates how a biological signal is converted into an electrical one [31]. There are mainly three components in the equivalent circuit. Firstly, the transmembrane potential functions as an electric source (V_in). Secondly, the parallel circuit of C_M and I_M expresses an overall effect of the cellular double-lipid layer structure, ion channels and cellular signal pathway. Finally, the electrolyte between the cells and the electrodes induces part of the current to leak out to the bulk medium. This effect is expressed as the sealing resistance (R_seal). The current flowing through the basal membrane is the sum of the ionic current (I_ionic) and the capacitive current (I_M). Thus, the total transmembrane current is given by:

(1)

Accordingly, the voltage at node A is proportional to the second derivative of the transmembrane potential. This voltage is also proportional to the magnitude of R_seal [32]. The larger R_seal is, the better the recorded signal can reflect the transmembrane potential. In the extreme condition of an infinite seal resistance, the voltage at node A would correspond to the intracellular potential, thereby simulating a whole-cell patch configuration. This trend is shown in Figure 2 [32,33,34].

In MEA design, the ‘sandwich’ structure has been employed since it was first pioneered by Thomas in 1972 [16]. Firstly, a thin metallic layer is deposited on an insulative substrate as the sensing layer. Then a passivation layer is coated on it with electrode sites and pads exposed for electric sensing and signal output.

The flow chart in Figure 3 is a general fabrication process of glass-based MEA. In the beginning, Cr (30 nm) and Au (300 nm) are sequentially sputtered or evaporated onto the glass substrate. The use of the Cr layer can enhance the adhesion of the Au layer onto the substrate. Then the electrodes and traces are patterned from this composited metallic layer by conventional lithography and etching techniques. After that, an insulative layer of SiO₂ is deposited onto the chip surface by plasma-enhanced chemical vapor deposition (PECVD). Finally, the electrodes and pads are exposed from the silicon nitride layer by reactive ion etching. The fabrication of 3D MEA shares the design idea of planar MEA [35,36].

Figure 4 shows a typical fabricated MEA chip. The electrode is designed to match the size of a cell, with a diameter of 10–100 μm. To reduce the electric interference between electrodes, the center-to-center distance is usually more than 100 μm. To lower the impedance and improve the performance of the electrodes, usually a T_iN or platinum black layer is deposited on electrodes [37].

Before growth and propagation, cell adherence to substrate or surface is always the first step in a cell-based experiment. ECIS is one of the electrochemical techniques that can be applied to monitor cell adhesion, spreading, and motility in real time [39]. Figure 5 shows a schematic overview of the measurement setup pioneered by Giaever and Keese [40]. In this setup, microelectrodes are constructed beneath the cell attachment platform. An alternating current is applied on the electrodes and the voltage is monitored using a lock-in amplifier. If no cells are cultured on the surface, electric current can flow freely from the surface to the electrodes. Cells growing on the electrode will impede the current flow and thus increase the resistance of the system (measured in ohms) because cell membrane acts as an insulator. The magnitude and time course of the increase is found to become more pronounced when the cell density during seeding approaches confluence. Electrodes without any cells will produce a minimal ‘base-line’ resistance. The addition of analytes that either rupture the cell monolayer or cause disturbances will lead to changes in impedance.

Typically, two structures are used in impedance sensing: the monopolar electrode [41,42] and the IDEs [21,43]. With monopolar structure, the area excluding the working electrode is relatively large, only a few cells on the working electrode contribute to the measured impedance, which results in fluctuations among experiments. Moreover, large counter electrode hinders further miniaturization, which is essential in a high-throughput screening (HTS) system. In contrast, the interdigitated structure shows improved performance as the current flows in the vicinity of the electrodes surface. A much higher sensitivity to impedance change is demonstrated compared with the conventional design. In addition, the IDEs commonly cover up to about 50% of the area of each well, lowering the well-to-well differences in the HTS system. However, no preference can be given to these two prototypes without specific applications.

Generally, the fabrication process of ECIS is the same as MEA, with a metallic layer deposition on the substrate and a following lithography and etching process, as shown in Figure 3.

LAPS is a semiconductor device proposed by Hafeman et al. in 1988 [27]. With a structure of Si/SiO₂/Si₃N₄, it can be excited by a modulated light source to generate a photocurrent which corresponds to the potential on the Si₃N₄ surface. In principle, any effect that results in the change of surface potential can be detected by LAPS, including the ionic change [44], redox reaction [45], etc. The fabrication process of LAPS is easier due to the simple structure, and the encapsulation of LAPS is much less critical. Besides, the extremely flat surface makes it convenient to incorporate into the micro-volume chamber, which facilitates trace measurement. Therefore, LAPS holds promise as a preferred platform for future cell-based biosensors.

LAPS is a potentiometric semiconductor device that could have an EIS structure or an electrolyte-metal-insulator-semiconductor (EMIS) structure, as shown in Figure 6. Generally, it has the layer sequence of an Si/SiO₂/sensitive layer. An external DC bias voltage is applied to form an accumulation layer, a depletion layer or an inversion layer at the interface of insulator (SiO₂) and semiconductor (Si). When an AC modulated light illuminates the LAPS chip, the semiconductor absorbs energy, leads to energy band transition and produces electron-hole pairs. Electron and hole would combine soon and photocurrent cannot be detected by peripheral circuit without light illumination. When LAPS is biased in depletion, an internal electric field is produced across the depletion layer, and the width of the depletion layer is a function of the local surface potential. When a modulated light illuminates the bulk silicon, light induced charge carriers are separated by the internal electric field and thus photocurrent can be detected by the peripheral circuit. The amplitude of the photocurrent depends on the local surface potential. Therefore, by detecting the photocurrent of LAPS, local surface potential can be obtained [27].

When cells change the surface potential of LAPS through the ions (H⁺, Na⁺, K⁺, Ca²⁺, etc.) or redox couples like CySS/CyS (cystine/cysteine) and GSSG/GSH (oxidized/reduced glutathione), the actual bias voltage varies, and fluctuations are generated in the corresponding photocurrent (Figure 6(a)). Therefore, by focusing the light spot on the surface cultured with cells, ionic changes or redox potential can be detected through the local surface potential measurement. Thus LAPS is applied in cellular metabolism detection, cytotoxicity evaluation and drug analysis.

For pH detection, a layer of Si₃N₄ is fabricated on the surface of LAPS. This thin layer of silicon oxynitride is used as the insulating layer that can effectively separate the silicon substrate from the electrolyte. According to the site-binding theory [30,46] which is also demonstrated in the semiconductor sensor based on field-effect [47,48], a potential difference related to H⁺ concentration of the electrolyte arises on the insulator (Si₃N₄/SiO₂)-solution interface. The insulating layer interacts with protons in the solution to form groups of silanol (Si-OH) and silamine (Si-NH₂). Therefore, the sensor surface potential is determined by the protons in the solution. As is shown in Figure 7, if a voltage is applied to the sensor, an electric field is formed at the silicon-insulator interface. By illuminating the rear side of the sensor, a photocurrent is generated [49]. This photocurrent corresponds to a hole-electron pair created by radiation absorption at an atomic level. Since a field potential is formed in the chip, holes and electrons move in opposite directions, resulting in the creation of local current which can be output from the backside aluminum layer.

The amplitude of the photocurrent depends on the applied bias voltage and the sensor surface potential determined by the pH value of the electrolyte. The insulator surface is partially covered by silanol and silamine, both of which can be considered as a function of pH. The surface is neutral around pH 3.5. Therefore, under physiological conditions (pH 7.4), the surface becomes negatively charged, which makes the surface potential pH-dependent. Such a dependence is Nernstian and linear from pH 2 to 11 [27]. Generally the characteristic current-voltage curve (I–V curve) is S-shaped (Figure 6(b)). In the continuous measurement, the applied bias voltage is fixed at a constant value, which is selected as the inflection point of the I-V curve. By doing this, the recorded photocurrent change can reflect the pH variation most sensitively. Then the variation of photocurrent ΔI represents the pH change.

Electron exchange accompanies oxidation-reduction reactions, which can be detected by a noble electrode (Pt or Au electrode) through potentiometric measurement. The potential on the electrode (also redox potential) is determined by the ratio of the concentrations of the solution species [50] and the value is stated by the Nernst equation. Redox potential measurements can also be obtained by depositing a metal layer (often Au) on the insulator surface of the EIS structured sensor. When in contact with an aqueous solution which contains a redox pair, the surface potential of the metal layer reflects the ratio of the redox pair’s concentration. The Au layer deposited functions as an electron transfer electrode similar to a traditional Au electrode, and the equilibrium potential formed on the surface of the layer is measured by photovoltage technology. As the EMIS structured sensor and EIS structured sensor have similar structure, fabrication and measurement technology, they can share the same measuring system.

It is worth mentioning that measurements of reversible, Nernst potentials are accompanied by a number of requirements, among which the most important is the need of fast electron exchange between the metal layer and the analyte. Slow electron transfer is not propitious to the formation of stable equilibrium potential, and this is responsible for the failure of potentiometric measurements of biological samples. To obtain meaningful analytic information, mediators are usually used so that stable redox potential measurements could be made. The mediators can be redox couples directly added to the samples, like Fe(III)/Fe(II) [51] or electron transport promoter, such as 4-pydidinethiol (4-PySH) and bis(4-pyridyl)disulfide [52] that are to be modified on the metal surface (Figure 8).

The LAPS device usually consists of the heterostructure of Si/SiO₂/Si₃N₄. Fabrication of LAPS is easy and fully compatible with the standard microelectronics facilities (Figure 9). A silicon chip is chosen as the substrate with a resistance of about 5–10 Ω∙cm. Then a SiO₂ layer is formed with a thickness of 30–50 nm. A 50–100 nm layer of Si₃N₄ was deposited by PECVD as a sensitive layer on the upper side of the bulk. For the EMIS structured sensor, a deposition step of 30 nm Cr and 100 nm Au on the Si₃N₄ surface is necessary. Finally, an aluminum membrane about 300 nm in thickness is evaporated on the backside of the silicon chip to form an ohmic contact.

MEA is a noninvasive and long-term method enabling stimulation and recording of bioelectricity with high spatial and temporal resolution in cell and tissue cultures. As the MEA technology can be applied to any electrogenic tissue or cell, that is, central and peripheral neurons, cardiomyocytes, and retinas, the MEA biosensor is an ideal in vitro system to study the dynamics and physiology of these cells or tissues and to monitor both acute and chronic effects of drugs and toxins.

As the basic physiological feature of life, heterotrophic cells absorb various nutrients, produce energy and secrete acidic wastes for growth and development. Metabolic energy is generated by carbon sources such as sugars, amino acids and fatty acid. The schematic of cellular metabolism and physiological processes is displayed in Figure 18. Under normal conditions, glucose is taken up by the cells and degraded into energy and acidic products. Under natural aerobic conditions, glucose is converted into CO₂ and energy via glycolysis, the citric acid cycle and oxidative phosphorylation. While under anaerobic conditions, glucose is converted into lactate with energy [110] via glycolysis and combining lactate dehydrogenase. The ATP generation per glucose molecule under aerobic conditions is 19 times higher than that under anaerobic conditions (38 ATP/glucose versus 2 ATP/glucose), while the acidic product is much less than anaerobic conditions (0.167 H⁺/ATP versus 1 H⁺/ATP).

Lactic acid and CO₂ are the main products from aerobic and anaerobic glucose degradation. They can be hydrolyzed into lactate/H⁺ and HCO₃⁻/H⁺ respectively, both in the cell and outside the cell after passing through the plasma membrane in the form of an unhydrolyzed state. Lactate is excreted with facilitated transport by monocarboxylate carriers and anion exchange proteins [112]. Meanwhile HCO₃⁻ can be transported outside the cell by an antiport. During the hydrolysis process, intracellular and extracellular protons are generated. The intracellular ones have to be transported outside by a Na⁺/H⁺ exchanger. Besides, when the cell receives external stimulation, e.g., exposure to a toxic agent or receptor activation, the metabolic activity or the ionic equilibrium in the cell can be interfered with, and the consumption of cellular ATP may change. Since ATP hydrolysis is closely related to the production of acidic metabolites, it can cause changes of extracellular acidification rate (ECAR).

As mentioned above, LAPS has many advantages for constructing cell-based biosensors. Since the first publication of the Cytosensor^TM Microphysiometer, it has been widely used. Several reviews [28,44,49,113,114] have been published to introduce the application of the microphysiometer. Besides, the newly proposed cell-semiconductor LAPS device for extracellular detection is considered as a useful tool for cell electric biology study. With the microphysiometer, functional characteristics of ligand/receptor can be studied, usually by monitoring the time-dependent and dose-dependent response of ligand/receptor binding. The time-dependent response shows the dynamic characteristics of ligand/receptor binding, while the dose-dependent response indicates the activation dose of the receptor.

The identification of specific and functional orphan receptors will facilitate studies on their physiological roles and the search for receptor agonists and antagonists. Due to the lack of a proper radioactive label, screening this type of ligand and receptor with radio receptor assay (RRA) is too difficult. It is also difficult to solve this problem with the biochemical methods, as the details in the mechanism of this type of receptor are not very clear. When using the microphysiometer, a generic characterization of ligand/receptor can be obtained. But it is usually insufficient for the identification of ligand/receptor. Therefore, other methods are usually used along with the microphysiometer to give precise identification of ligand/receptor.

ECAR of CHO cells expressing orphan GPCRs was measured [116]. When treated with synthetic compounds, such as bioactive peptides, it was found that CHO cells expressing an orphan GPCR, FM-3 were responsible to neuromedin U in the microphysiometric assay. It was also discovered that bradykinin (BK) activated extracellular signal-regulated protein kinase 1 and 2 (ERK) in the human embryonic kidney (HEK) 293 cells [115]. ECAR was a reflection of the Na⁺/H⁺ exchange (NHE). When combined with intracellular Ca²⁺ concentration from fluorescent measurements, and ERK activation assessment through western blotting with a phosphor-specific ERK antibody, it was found that signals obtained by exposure of HEK 293 cells to BK, were blocked by HOE-140 (B₂ receptor antagonist) but not by des-Arg¹⁰-HOE-140 (B₁ receptor antagonist) (Figure 19). Thus it was concluded that HEK 293 cells express endogenous functional BK B₂ receptors.

Cell metabolism is accompanied by various oxidation-reduction reactions, such as hydrogen transfer reactions and electron transfer reactions in the respiratory chain. Maintaining the proper concentrations of oxidants and reductants may be essential for cell health. The excretion of acidic waste from intra-cellular fluid has been used as a parameter to evaluate the metabolism state of cells. Similar to the acid-base mechanism, the redox potential in tissue maintains within a narrow range and injury may be brought about at the cellular, subcellular, and molecular level when the balance breaches. The concept of redox potential balance is not new but has not received the attention of that associated with acid-base chemistry. One of the reasons may be the lack of effective and direct measuring technology for redox potential when involved in biological systems.

The early measuring method is based on the dyes (such as mythylene blue) that indicate the redox potential through their color change [117]. The earliest application of redox potential for the extra-cellular phenomenon was the measurement of oxidation potential of in vitro plant oxidase using a gold electrode electrochemically. Then, electrodes (mainly platinum or gold electrode) had been used as the major effective way of redox potential detection in several decades. The cells and tissues measured included yeast, various bacteria [118,119,120], thin slices of liver and muscle suspended in buffer solutions, and tumors in vivo [121]. Until now the potentiometric measurement using noble electrodes has been widely used in redox potential monitoring of cell culture [122,123]. Usually, the redox potential electrode is worked as a working electrode and connected to a reference electrode through a potentiometer. Besides, Ward et al. [52] developed an universal redox electrode by modifying a Au wire electrode with electron transport promoter 4-pydidinethiol and bis(4-pyridyl)disulfide, which improved the performance of the electrode and avoided biofouling from biological samples. As the electrode technology and manufacturing process approaches maturity, commercial combined redox potential electrodes have come within our range of vision. The combined electrodes are mostly applied in biotechnology or fermentation because of their convenient usage in comparison with the traditional two-electrode system in electrochemical cells.

In biotechnology and microbiology, redox potential is one of the most complex indicators of the physiological state of microbial cultures and its measurement is useful for qualitative and quantitative determination of the microbial contamination [124]. Besides, the significance of redox potential is not limited to functioning as an indicator to reveal related state about cultures, another no less important application is to externally control the redox potential to lead the system developing as premeditated. Zheng et al. [125] adopted extracellular oxidation-reduction potential (ORP) of −170 mV and obtained a maximal 176 g/L lactic acid production compared to other different ORPs. Koo et al. [126] found that the production of coenzyme Q₁₀ can be improved by increasing the NADH/NAD⁺ ratio in agrobacterium tumefaciens, which is also an example of redox potential working as a control factor. The indicator and controller roles are just like the two sides of a coin, which also can be applied to cell-based research.

For decades, the redox potential has become an important parameter monitored in cell culture medium. The study focuses on cell metabolism detection. Considering the complexity of redox potential in cell culture medium, other parameters like ATP, dissolved oxygen, glucose and glutamine consumption are also monitored to obtain a general analysis of cellular physiological states. Hwang and Sinskey [127] used the redox potential measurement of the medium using a platinum electrode and empirical correlation of the cell concentration with the redox potential for on-line estimation of viable cell concentration. It has been demonstrated that this method is applicable to various cell lines. Later, Eyer and Heinzle [122] used a method based on an ATP balance with ATP steady-state assumption to estimate viable cells in a hybridoma culture and confirmed the validity of viable cell estimation using redox potential measurement. Cell metabolism is always accompanied by substance consumption or depletion, which can be indicated by a decreasing redox potential. Higareda et al. [123] studied the relationship between culture redox potential (CRP) and glucose depletion in hybridoma cultures, and found that simultaneous measurement of oxygen uptake rate (OUR) and CRP can permit the discrimination between operational eventualities and real metabolic events.

However, the measured redox potential of cell culture, also extracellular redox potential, is a comprehensive assessment of all redox pairs existing in the culture medium, and the measurement can be easily affected by pH and dissolved oxygen. Recent evidence suggests that the concentrations of specific redox couples may play a role in the regulation of other cellular functions, including gene expression. It is expected that the intracellular concentrations of redox couples like NAD⁺/NADH, NADP⁺/NADPH, CySS/CyS and GSSG/GSH may change selectively with changes in the cellular environment. Biochemical techniques exist for direct quantitation of most import cellular redox couples listed above, such as reverse-phase high performance liquid chromatography (rpHPLC) [128] or cellulose high performance thin-layer chromatography (HPTLC) [129], which have the ability to separate these species and realize measurement of specified species based on the reactivity of the thiol moiety, and a less specific detection method using membrane-permeable compounds which are fluorescent upon thiol conjugation [130]. Nevertheless, these biochemical techniques are invasive for cells and often based on endpoint detection. Electrochemical detection is expected to avoid the drawbacks and realize noninvasive and continuous monitoring of the intracellular redox state. An important difference between the electrochemical and biochemical methods is that the former measures the activity of cellular reactions rather than the steady-state concentration of a particular biochemical species.

The first application of electrochemical-based biosensor for intracellular redox potential detection is reported by Rabinowitz et al. [51] who employed an EMIS structured sensor to probe intracellular redox activity in live cells. The sensor structure and measurement system adopted a modified Cytosensor microphysiometer. The electrochemical potential at the gold-electrolyte interface contributes to the reverse-bias potential acting across the photosensitive silicon semiconductor. Illumination of the silicon with a light-emitting diode generates an AC photocurrent in the external circuit, determined by the voltage applied through the controlling electrode. Changes in the redox potential at the gold surface are reflected through this current. In addition, there are no measurable direct current flows through the circuit during these measurements, which enables stable, rapid, low-noise measurements on microliter volumes.

The most essential point in their study lies in measurement of the redox activity of the interior of cells noninvasively from extracellular microenvironment change. A carrier mediator that exchanges electrons rapidly with both the exterior and interior couple and links the extracellular electrolyte solution to intracellular redox pairs was used. As shown in Figure 20, the couple menadione/menadiol was used as an effective carrier mediator, because both of them are lipid soluble. Menadione and menadiol can be oxidized and/or reduced by intracellular enzymes. The major enzyme had been demonstrated to be NADPH, produced by the pentose phosphate pathway. In the meantime, the external medium bathing the cells contains a ferricyanide/ferrocyanide couple which reacts rapidly with menadiol. Thus the monitored redox potential of the ferricyanide/ferrocyanide couple is a measure of intracellular redox activity. Using this method they studied three cell lines (L929, CHO, and CH27), and the effect of reactions of mitochondrial electron transport chain and signal transduction cascade on the cell reduction rate.

Cellular life is characterized by dynamic molecular processes whose highly interconnected regulation is essential for the whole organism. Cells continuously integrate different sources of chemical and physical signals originating from both internal and external environments. The ‘output’ of this cellular signaling network may be manifested as a decision about growth and mitosis, activation of distinct metabolic pathways, the production and release of proteins or the initiation of programmed cell death. Conventional methods, for instance fluorescence techniques, are most often used to dissect the molecular basis of cell functions, either with fluorescent protein analogs or with synthetic dyes. However, no truly appropriate fluorescence techniques have been developed so far for the analysis of cell electric activity, morphology and metabolic rates. Functional cellular assays are employed to monitor several kinds of signals at the same time. Thus the signal-processing of living cells can be profiled in a more detailed and thorough way.

Brischwein et al. [160] first reported on multi-parametric cell based assays with data obtained solely with integrated sensors on silicon chips. Extracellular acidification rates (with ISFETs), cellular oxygen consumption rates (with amperometric electrode structures) and cell morphological alterations (with impedimetric electrode structures, IDES) were monitored simultaneously for up to several days. Ceriotti et al. [163] employed a multi-parametric chip-based system to measure cell adhesion, metabolism, and response to metal compounds which was previously classified as cytotoxic in immortalized mouse fibroblasts (BALB/3T3 cell line). In a parallel, online, and in a label-free way, the system measured extracellular acidification rates (with ISFETs), the cellular oxygen consumption (with amperometric electrode structures (Clark-type sensors)), and cell adhesion (with IDESs) (Figure 24(a)). A group from the biosensor national special lab in China has established an automatic multi-parametric platform based on an integrated biochip which measures the cell electric activity (with MEA), the cell morphology (with ECIS) and metabolic rates (with LAPS) (Figure 24(b)). Using the integrated multi-parametric sensors, some cell-based assays are conducted, including cytotoxicity and drug assessment, cell development and stem cell differentiation research.

From another point of view, extracellular acidification and redox potential are important indicators of cell metabolism activity. As the EIS and EMIS sensor have similar structure and working principle, it makes sense to integrate the two sensors on the same chip. The related work has been carried out by Wang et al. [170]. The integrated potentiometric sensor (Figure 25) was constructed by depositing a gold metal layer on a partial surface of silicon dioxide. The sensitivity for pH and redox potential measurement is approximately 41.6 mV/pH and 53.2 mV/log[Fe(III)/Fe(II)], respectively. The latter is in accordance with the work of Adami et al. [171]. The integrated sensor was used for nephrotoxicity assay under drug stimulation.

The physiological state shows an acidified extracellular microenvironment and a reduced redox potential. It is revealed from this study that gentamicin inhibits the regular cellular metabolism, and the toxicity enhances with the concentration, but the toxicity is reversible when washed with PBS solution, which is in accordance with reversible or partially reversible kidney failure induced by gentamicin in clinical use, while 5-FU exhibits an anti-metabolism effect both intra and post 5-FU administration. The change of extracellular redox potential is significantly different from that of acidification, which indicates that the drugs are toxic for kidney cells mainly through acidification inhibition but induce no significant change to the concentration of extracellular redox pairs. However, this study was just a preliminary application of integrated photovoltage-based biosensors for cellular toxicity evaluation and a complete dose-response analysis and EC₅₀ determination was not presented. Anyhow, the concept of an integrated sensor for acidification and redox potential detection has potential in drug toxicity analysis and can work as a supplementary means for preclinical drug screening.

Microfluidic technology is a powerful tool to facilitate the miniaturization and integration of analysis systems and can even address many issues that traditional methods cannot deal with. It has a cell-comparable small size and can precisely confine cells in the specific microchannel. By employing micro pumps and valves, mechanical or chemical stimulus can be introduced to cells and the corresponding cellular level biological responses can be monitored by integrated electrochemical electrodes. Even single cell and small cell group analysis can be achieved. Besides, microfluidics can realize high-throughput parallel biological monitoring and analysis.

The microfluidic chip with specific micro-fabricated electrodes has broad applications in cell morphology research and has many dramatic advantages over traditional assays, such as simple operation, rapid detection, and less invasiveness. Cells are cultured in the microfluidic chip and the culture media solution can be introduced automatically through the microfluidic channel. Using suitable electrochemical detection methods, cell morphology information, such as adhesion, migration, proliferation, and apoptosis, can be monitored in real time. One typical example is the automatic transwell assay, in which two microfluidic chambers separated by a porous membrane are designed to monitor cell migration [172]. As illustrated in Figure 26(a), cells are seeded in the upper chamber through microfluidic channels and an EIS sensor is placed in the lower chamber. Once cells migrate from the upper chamber into the lower one, great impedance changes can be detected. Figure 26(b) shows the impedance changes at different times. In addition, the influence of different ECM components can be also detected by this assay.

The microfluidic channel also provides great convenience for cellular metabolism analysis, especially in terms of single cells and small cell groups. Cells can be cultured in a tiny space and their physiological and electrical responses will be captured easily. Currently, many researches are covered in analyzing extracellular microenvironment [173] and intracellular components [174], monitoring ligand-induced cell secretion [175], recording the extracellular potential [176] and cell dynamic responses [177,178], etc. Besides, the cell patterning technique is an important issue in fundamental cell biology, tissue engineering, and neuron network research. Microfluidic technology plays a great role in this field. The chemical and cellular environment can be controlled precisely through the microchannels modified with diverse chemical molecules [179,180,181].

Assisted with microfluidic technology, cell-based electrochemical biosensors will have broad commercial and medical applications in the future. Low cost, disposable chips can be produced in large amounts and point-of-care diagnosis can be realized in the near future.

In recent years, nanotechnology has been greatly improved and widely applied in the biosensing fields. The most commonly used nanomaterials include nanoparticles, nanotubes, nanowires, and nanobars. Due to attractive electronic, optical, magnetic, thermal and catalytic properties, they have important applications in many interdisciplinary fields.

Nanomaterials play important roles in many aspects of biosensors. Owing to the large surface-volume ratio, nanogold particles are employed to modify the sensitive surface and enlarge the effective working area. Magnetic nanoparticles are usually used to accumulate or separate some important biomolecules conveniently [182]. Since carbon nanotubes (CNTs) have high electrocatalytic effect and fast electro-transfer rate, they are always used in the detection strategy to assist signal detection [183]. As shown in Figure 27(a), an indium tin oxide (ITO) electrode modified with CNTs was used as the working electrode and SU8 was patterned to limit the dimension of the ITO electrode. PC12 cells were cultured in a limited chamber (hundreds of μm²). Once extracellular stimuli induced the exocytosis of cells, the molecules can be detected by the CNTs-ITO electrode, as illustrated in Figure 27(b). The results demonstrated that the sensitivity of the electrochemical sensor after CNTs surface modification was improved by 2.5–3 orders of magnitude and a typical amperometric response from a single cell was shown in Figure 27(c). The distinct optical performance of nanomaterials is also very attractive in the development of optical biosensors. Many kinds of nanomaterial-based biosensors are investigated to detect DNA, protein, and cellular substance by diverse detection technologies [184].

As an attractive trend, the distinct performance of nanomaterials can contribute much to cellular diagnosis. Nano-dimension electrodes can be used to probe intracellular compounds and monitor cell metabolism and electrophysiological activities. Nanomaterials can help to simplify and miniaturize biosensors for biomolecular trace detection.