Nanoporous Layer Integration for the Fabrication of ISFET and Related Transistor-Based Biosensors

Abstract

More and more chemosensors and biosensors are turning to electronic transistors, as they are ideal transducers, precise in current response, miniaturized in size and capable of providing sub-picomolar detection limits. Among these devices, ISFET transistors—Ion-Sensitive Field-Effect Transistors—have the capacity of integrating ion-sensitive layers together with field effect transistors of ultimate generations. Recent studies have indicated that nanoporous materials deposited or grown within the transistor gate space offer a dual advantage—a favorable environment for an optimal capture of liquid state receptors through capillary effects, but also of direct anchoring of these nanoporous structures on a Si wafer. This article aims to review the constructive evolutions of ISFET transistors, along with some newer nanowire devices, as well as their co-integration techniques with nanoporous materials, which are beneficial in the optimization of many chemosensors but of enzymatic biosensors in particular.

1. Introduction

In the last five years, electronics has proposed the smallest Field-Effect Transistors (FETs) ever manufactured, generating 2 nm technology [1]. Their miniaturization in the coming years will be limited by the dimension of one atomic layer, up to 0.2 to 0.5 nm [2]. For this reason, the roadmap of semiconductor devices has also been directed towards integrated sensors, especially chemo- and biosensor applications [2,3]. From the transducer point of view, the next biosensor families belong to the bioreceptors integrated near nano-transistors [3,4,5], sometimes Artificial Intelligence (AI)-assisted [6], to produce different kinds of Biosensor Field-Effect Transistors (BioFETs) [5,6,7,8].

Therefore, chemoscience pays attention to those sensors that are technologically compatible with micro- and nano-electronic devices [9]. What must be emphasized from the beginning concerns these smallest transistors, with dimensions of 1 to 20 nm, which are mainly used for the manufacture of memories and computer processors, where the fight for supremacy in processing speed and packaging density is the strongest [10]. But the 2 nm technology is not compulsory in use nowadays for analog electronics, or electronics for interfacing with real phenomena from the environment, such as the field of transducers for sensors. Especially for low-cost biosensors, more relaxed and very well-experienced technological nodes are frequently used, such as the 150 nm Metal Oxide Semiconductor Field-Effect Transistor (MOSFET) [11], or the 300 nm Ion-Sensitive Field-Effect Transistor (ISFET) [12], or even more than 300 nm for organic transistors [13,14]. There is no linear dependence or a perfect proportionality between smaller or larger sizes, which would generate better performances of biosensors. As proof, there is a huge number of articles annually published, but they frequently return to the detection of approximately the same group of analytes (where glucose is at the top) and do not systematically better obtain the entire set of parameters for sizes.

But, in MOSFET downscaling, there are laws that require a correlation between smaller widths and lengths of the conduction channel and higher doping concentrations, thinner gate oxides, and the use of high-k dielectrics [1,2,3]. These principles can also be extrapolated to FET-based biosensors. Thus, transistors with smaller lengths and widths must be co-integrated with thinner sensitive elements of a comparable aspect ratio. In addition, aiming for uni-molecular detection, the conditions of feeding the biosensors with such tiny quantities of sample liquids of the order of micro-nano-pico-liters can be accomplished only by Micro-Electro-Mechanical System (MEMS) engineering.

A specific sub-field of chemosensors involves the use of feeding tubes—capillary structures designed to deliver minute volumes of liquid to Lab-On-a-Chip (LOC) platforms [15]. In these systems, MEMS integration techniques are frequently adapted from micro- and nano-electronics technologies [16,17].

This capillarity domain is related to the objectives of this article in two directions: (i) in the immobilization technology of liquid bioreceptors, which can be captured through capillary phenomena if porous materials are integrated in the transistor’s vicinity; and (ii) these MEMS feeding channels are particularly described in some LOC biosensors, in agreement with the transistor technology.

Another relevant sub-field of chemosensors is focused on the fabrication of nanoporous layers directly onto glossy surfaces such as Si, SiO₂, or Si₃N₄, which are commonly used in conventional transistor technologies [18,19,20,21]. The pore sizes can be precisely controlled through fabrication processes, enabling effective binding of bioreceptors in the liquid phase during immobilization steps. When ion-sensitive electrodes (ISEs) are employed as receptors and FET transistors serve as transducers, the resulting configuration functions as an ion-sensitive field-effect transistor, ISFET [22].

The main goal of this paper is to fill the gap between two essential constructive parts of an ISFET chemosensor: (i) the receptor element, which can vary from a simple inorganic ion-sensitive layer to enzyme, aptamers or complicated proteins, and (ii) the transducer element, which is the transistor—mainly an ISFET, but also a related FET. These two parts would be disjointed if no intermediate layer is used. This intermediate layer, which is made of micro-nanoporous materials, is of crucial importance for the proper functioning of the final bio- or chemosensor. Hence, the overarching objective is focused on the fabrication technology and properties analysis of these key intermediate layers, which should be compatible with the transistor technology, be strongly anchored on the gate area and be versatile to offer controlled porosity to capture various bioreceptor solutions.

From the location point of view, this intermediate layer can be integrated on the top of the Si wafer, in the device proximity [23] or on a separate extended gate [24,25], as an Extended Gate Field-Effect Transistor (EGFET) [26,27,28,29], aspects that will be approached in this paper, too.

As mentioned above, one direction of study of this review concerns technologies for immobilizing bioreceptors by capturing them on nanoporous intermediate layers [30,31,32]. Few particular methods to immobilize enzymatic membranes for Enzyme-FET biosensors [33,34], even paper-based devices [34], still exist.

Sometimes, liquid enzyme solutions mixed with high concentrations of cross-linkers can lead to proper immobilization. Other times, the enzymatic membrane is strongly anchored to the substrate but accompanied by alleviated activity [35] or, in other cases, it is accompanied by cracks, fissures, and peeling off [36]. The first generations of biosensors used enzyme immobilization in gels [37,38], but later they switched their capture on nanoporous or nanostructured materials [39,40]. Therefore, a better resolution is expected to result from a combination of chemical immobilization methods and fluid attraction by capillarity through nanopores.

A final topic of this paper concerns the location of these intermediate nanoporous layers in chemosensors. The ideal location would be in proximity to the conduction channel of the transistor to influence as sensitively as possible the current through this device. Some authors have proposed an optimum location just above the gate insulator [41,42]. Technological inconsistencies immediately arise: if we intend to use nano-transistors with extremely small gate areas, then the amount of enzyme that can be immobilized becomes negligibly small, which reduces the sensitivity of the sensor. In addition, during measurements with liquid droplets above the sensitive layer, contamination of the gate insulator with ions by the flat-band effect or accidental gate-source current leakage via liquid drop may occur [43]. On moving the entire sensitive area to the so-called extended gate, advantages such as electrical isolation successfully occur, but it is accompanied by disadvantages, such as poor influence of the ion charges generated in the extended gate area on the current through a distant channel [44]. However, a general opinion suggests attaching the sensitive layers just above the gate space [45]. Alternative transistors with Solution Gate FETs (SGFETs) in contact with the liquid analyte solution would be a possibility for integration of a solid-state transducer with a liquid analyte solution [46,47]. However, evaporation phenomena and possible secondary reactions at an electrode make the SGFETs an imprecise and more degradable device than a silicon-based FET. ISFET transistors have accumulated over 50 years of experience [48], being the most resilient over years of use.

However, the discovery of a technology able to directly capture nanomaterials near low-size transistors, technology that is also capable of immobilizing enzymes or other liquid bioreceptors and is also compatible with standard CMOS or ISFET technologies, is still a challenge [49,50]. This is the further argument for which this review is carried out, finally emphasizing the possible development directions for the future.

2. ISFET Work Principle and Technological Challenges

2.1. From ISFET Theory to Materials Motivation

Bergveld, the ISFET’s father, describes the evolution of this transistor after thirty years of existence [51] in conjunction with its integration of sensitive layers. ISFET is essentially a Metal-Oxide Semiconductor Field-Effect Transistor (MOSFET) [52], biasing the gate by a separate connection to a standard electrode placed in an aqueous electrolyte in connection with the gate oxides (Figure 1a). The electronic symbol of an ISFET transistor is presented in Figure 1b and comprises four terminals: S (Source), D (Drain), G (Gate) and B (suBstrate). The red dotted line is used to suggest a frequent situation, when the S and B terminals are connected together inside the transistor capsule, in which case, only three terminals rest outside the capsule: S, D and G [52]. When this red dotted line is not materialized as a short circuit inside the capsule, it must be deleted from the symbol, and the device possesses four distinct terminals: S, D, G and B.

The operation of an ISFET transistor can be explained by two principles: (i) an electrochemical and (ii) an electronic principle.

The electrochemical principle applies to the liquid–solid system at the top of the ISFET (colored part in Figure 1a), which includes the liquid carrying ionic charges, placed between two interfaces: I1 metal–solution, and I2 solution–gate oxide. In practice, only one of I1 and I2 is the predominant interface, the one whose potential depends most sensitively on the concentration of an ion species in the solution. Usually, the I1 interface is the one that causes the ISFET sensitivity by choosing a specific metal, which defines the Ion-Sensitive Electrode (ISE), together with the liquid. At the level of the I1 interface, the Nernst law offers the first conversion from the ionic concentration into electric potential V_I1, as the transduction curve of the ISFET sensor. The article [51] shows that the I2 interface can also influence how the gate potential depends on the targeted ion concentration by the V_I2 term; we choose an appropriate gate insulating material, which comes in contact with the ionic liquid, according to the most general law:

$${\mathrm{V}}_{\mathrm{G}\mathrm{B}}={\mathrm{V}}_{\mathrm{I}1}+{\mathrm{V}}_{\mathrm{I}2}+{\mathrm{V}}_{\mathrm{S}}+2{\mathsf{\phi }}_{\mathrm{F}}+{\mathrm{V}}_{\mathrm{F}\mathrm{B}}$$

(1)

where V_S is the source voltage, V_FB is the flat-band voltage of the transistor given by the FET technological node, ϕ_F is the Fermi potential in p-type semiconductor, and V_GB is the applied voltage between G and B terminals, where V_GB = V_G−V_B in the most general case.

The FET transistor is the part drawn in black and white under the gate oxides in Figure 1a. The FET operating principle is the following: the external voltage applied on the gate must be positive and high enough so that the semiconductor surface potential takes the positive value of 2ϕ_F. From this moment, a non-null drain current arises, I_D, a current that flows through the inversion channel established at the semiconductor surface between the source and the drain. It is known from the MOSFET’s theory [52] that after V_G has increased, so that its surface has entered a strong inversion regime (at which moment V_G = V_T), the gate voltage can still increase above this threshold voltage, V_T, but the potential of the semiconductor surface remains constant at 2ϕ_F.

Considering the biasing model of the FET part as for any n-channel MOSFET in threshold conditions [52], with the source and substrate grounded together so that V_S = V_B = 0, and while the gate voltage is positive biased to attract an electron channel at the semiconductor surface, the following expression of the threshold voltage results from (1):

$${\mathrm{V}}_{\mathrm{T}}={\mathrm{V}}_{\mathrm{I}1}+{\mathrm{V}}_{\mathrm{I}2}+2{\mathsf{\phi }}_{\mathrm{F}}+{\mathrm{V}}_{\mathrm{F}\mathrm{B}}$$

(2)

If only the I2 interface is taken into account for the ISFET sensor and any effect at the I1 interface is neglected (i.e., V_I1 = 0 in Equation (2)), then the threshold voltage of ISFET depends on the concentration of ions in the solution, according to a Nernst model reported in [8], but is corrected by an α parameter [53]:

$${\mathrm{V}}_{\mathrm{T}}=2\cdot {\mathsf{\phi }}_{\mathrm{F}}+{\mathrm{V}}_{\mathrm{F}\mathrm{B}}+{\mathrm{V}}_0+2.3\cdot \mathsf{\alpha }\cdot {\displaystyle \frac{\mathrm{k}\mathrm{T}}{\mathrm{q}}}\cdot \lg \mathrm{c}$$

(3)

Here, V₀ is the standard electrode potential, k is the Boltzmann’s constant, q is the elementary electric charge, T is temperature, c is the analyte concentration and α is a sensitivity factor depending on the ability of the oxide surface to supply or to acquire protons by the capacitance of the double layer, while ΔV is the last term from Equation (3). The α parameter is ideally 1 for a maximum sensitivity, or it is subunitary for realistic compounds such as SiO₂, Si₃N₄, Al₂O₃ and Ta₂O₅ deposited over the usual gate oxide of a transistor [51].

The last term from Equation (3) was extracted from experimental ISFETs for each previous oxide [51] and computed for HfO₂ oxide according to the model depicted in [8], considering NaCl as the electrolyte (aq. Solution at pH = 5.8) (Figure 2).

The α parameter depends on the capacitance of the double layer at the electrolyte/oxide interface C_D and the surface capacity C_S [53]:

$$\mathsf{\alpha }={\displaystyle \frac{{\mathrm{q}}^2}{2.3\cdot \mathrm{k}\mathrm{T}}}\cdot {\displaystyle \frac{{\mathrm{C}}_{\mathrm{S}}}{{\mathrm{C}}_{\mathrm{D}}+{\mathrm{C}}_{\mathrm{S}}}}$$

(4)

The C_s value can be expressed in terms of the acidic or basic equilibrium constants of the related surface reactions [8,51]. For the above oxides, the α parameter takes sub-unitary values that are equivalent to sub-Nernstian sensitivities. Hence, ISFET transistors can use not only conductive material ISEs (Ion Sensitive Electrodes) as sensitive elements for ion detection but also different gate oxides that translate analyte concentration variations through the buffer capacity of the insulator surface.

2.2. ISFET Particular Technology

One dissimilarity between the ISFET and the MOSFET, besides some different gate oxide stacks, is that the S and D metal contacts are laterally spaced from the gate electrode. This requires particular long lateral S and D diffusions that extend from the gate oxides far toward the drain metallic contact. This effect is presented by the technology simulation in Figure 3a. These diffusions produce large series resistances that will lead to a decrease in the transconductance, g_m of an ISFET versus g_m of a MOSFET [52], diminishing the device sensitivity.

Another specific technological process of the needle ISFET transistors that is usable in vivo is silicon etching, first described in 1978 by Esashi et al. [54]. Later, the etching process was transferred to a silicon on sapphire (SOS) wafer, which is a particular silicon on insulator (SOI) technique [55]. The technological advantage of the buried insulator usage, as an excellent etching stop layer for silicon, has propagated until now [56]. The further SOI advantage is fulfilled: a completely isolated chip occurs because the source and drain metallization are placed on the top side, while the back gate is placed on the bottom side of the wafer [56]. To emphasize these particularities, an example of SOI structure is presented in Figure 3b. The current vectors arise at the bottom of the Si-p film if electrons are driven by the back gate. Another technological incompatibility arises: a back gate integration with ion-sensitive material on the top side of any extended gate structure becomes impossible [24,25,26,27,28,29].

However, a full isolation of a Si-chip is expensive for a typical disposable sensor. If all terminals are on the same side of the chip, usually as top electrodes, there still exists another drawback of ISFET—its short-term stability. When it is in touch with aqueous ionic solutions from the test samples, some corrosion processes occur near the gate electrode, reducing its lifetime [57]. To eliminate this effect, the ion-sensitive part was extended outside the electronic chip as a longer electrode. This new device architecture is known as Extended-Gate Field-Effect Transistor (EGFET) [25,26,27,28,29] and was recently optimized [58,59]. The extended gate is essentially an elongation of the work electrode on the same chip but outside of the electronic area. In this way, only the extended gate is in contact with the tested liquid, while the FET part rests completely dry.

3. Nanostructured Oxides Used for Sensitive Layer Integration

3.1. Porous Silicon on Si Wafer

It is expected that the materials most compatible with Si technology would be the first candidates to be grown, deposited or converted in nanoporous layers directly onto the Si wafer. Therefore, among the first nanostructured materials to be directly anchored to Si was the Si itself by porous silicon (PS). A very accurate study was performed by Hoover Tiffany on the technology of converting Si into porous Si, with the pores’ measured distribution of 67 µm in the center of the wafer and 102 µm on the edges, and with a pore width of 1.6 µm [60]. The author proved that pure hydrofluoric acid (HF) is not the ideal electrolyte due to a difficult infiltration through formed pores in the presence of the capillary effects. The best solvent was an organic one—Di-Methyl-Formamide (DMF). It was able to enhance the anisotropic etching per pore, acting as an optimum surfactant that permits a uniform liquid infiltration in pores. Other authors reported the porous silicon manufacturing by electrochemical anodization of the top mono-crystalline silicon surface, by electrolyte of HF and ethanol as organic solvent, for photovoltaic (PV) applications [61]. The pore diameter ranged from 8 nm to 2 µm, according to the Si wafer properties, currents during anodization and the electrolyte formula [61].

More recently, further improvements were described within some optimized technologies [62]. Porous silicon was prepared by electrochemical anodization of a monocrystalline silicon p-type wafer, <100> crystalline orientation. Macro-porous silicon was fabricated for the electrolyte composed of Di-Methyl-Sulfoxide (DMSO) combined with HF 45% at the volume ratio of 10:46, and the current density was 8 mA/cm². Meso-porous silicon was fabricated for an aqueous–alcoholic electrolyte composed of isopropanol, HF 45% and deionized water at a volume ratio of 1:1:3, when the current density was 30 mA/cm², 60 mA/cm², and 100 mA/cm² for times of 20 s, 22 s, and 20 s, respectively (Figure 4). Different pore sizes between 2 nm and 50 nm are visible by scanning electron microscopy (SEM).

For the PS application in the construction of gas sensors, a crystalline Si substrate <100> orientation p- or n-type was the start wafer to produce PS with 45% porosity. The PS structures were fabricated in a Teflon electrochemical cell by hydrofluoric acid (HF) (48%) and almost pure ethanol (EtOH) (99.98%) as electrolytes under a current density of 13.6 mA/cm² for p-type and 10 mA/cm² for n-type. The anodization time ranged between 64 and 825 s to achieve the proposed layer thicknesses between 1 and 10 μm [63]. For the sensor fabrication, Al electrodes were deposited according to a geometric predefined pattern of two electrodes. The electrodes contained two concentric rings of 1 mm width and a trace connector of 1 mm width. Figure 5a shows the sizes of the geometric shape, while Figure 5b shows a photograph of the sensor with deposited aluminum on the Si-porous regions.

A few years later, the metal-assisted chemical etching (MACE) method was applied to PS film on Si wafers, making it able to be transferred to another substrate [64]. Few windows were opened on a Si wafer covered by thin silver film by lift-off, which was subsequently immersed in an etch MACE cell using a mixture of HF, H₂O₂ and deionized DI water. After the samples were completely etched and dried in the nitrogen environment, horizontal cracks occurred; consequently, the PS film suffered self-peeling from the Si substrate. The Si pore size was dependent on the ratio R of HF to H₂O₂ etchant: R = 5:5, 6:4, 7:3, 8:2 and 9:1 for 30 min, shown below through SEM imaging (Figure 6). The great advantage of this method is to define a robust technology of vertical transfer for large-area PS films onto inorganic and rather organic substrates.

In 2024, a step forward in PS development was the transition to porous silicon nanotubes (Psi-NTs) [65]. The synthesis of gold nanoparticles (Au-NPs) functionalized with 4-Merca-Ptophenyl-Boronic Acid (MPBA) on porous Si-nanotubes was possible using a nanowire template initiated from ZnO seeds deposited on Si wafers [65]. Finally, the ZnO was removed in NH₄Cl at 500 °C for 120 min using a helium environment. The accomplished typical nanotube morphology of the processed Psi-NTs is demonstrated by the transmission electron microscopy (TEM) images shown in Figure 7a. Then, the Psi-NT surface is functionalized with 3-aminopropyltriethoxysilane (APTES) acting as a reducing agent. APTES also facilitates the Au-NP synthesis, followed by the incubation of the Au-precursor solution. Different TEM images of the Psi-NT structures are presented in Figure 7b–e for Au-NP uniform distribution. Figure 7f indicates the size distribution of the Au-NPs. Applications are still valid in glucose biosensors [65] or lithium storage batteries [66]. Ultimate developments in PS-based biosensors show PS films able to immobilize amino-terminated anti-LF aptamers [67] for the detection of some protein biomarkers, like lactoferrin (LF) [68]. The optimal pore width was 50 nm, accomplished at a current density of 375 mA/cm² during 10–35 s of etching [67].

3.2. Porous Al₂O₃ on Si Wafer

One of the pioneering articles in the field of nanostructured Al₂O₃ anchoring directly on a Si wafer was reported by the co-authors of this article in 2010 [21]. The proposed multilayer structure, Si/SiO₂/Si₃N₄/Al₂O₃, successfully entrapped a glucose-oxidase (GOX) enzyme within the work electrode. The device works as a glucose biosensor. It was designed as a four-electrode electrochemical cell, (Figure 8), and although it is integrated in Si, it is not a transistor yet. The mono-crystalline form of Al₂O₃ is known as sapphire.

Some experimental tests occurred on discrete structures of porous Al₂O₃ anchored on a Si wafer, still in the absence of transistor co-integration. The Al film was deposited on a Si substrate over the previous multilayer structure Si/SiO₂/Si₃N₄ under vacuum conditions at a low deposition rate, with a thickness between 80 and 100 nm. This process resulted in the formation of γ-type catalytic Al₂O₃. Figure 9a presents an SEM image of the porous Al₂O₃ synthesized layer, while Figure 9b shows the experimental current–time dependence during the Al film conversion into porous Al₂O₃ by anodization at a constant voltage 80 V in diverse aqueous electrolytes: 1—(HCOO)₂, 2—(H₃PO₄), 3—(HCOO)₂: H₃PO₄ [69].

The anodic γ-type aluminum oxide offers two key advantages for chemosensors and biosensors: its high adsorptive capacity facilitates effective immobilization of enzyme solutions, and its favorable catalytic properties support the electro-oxidation of certain organic compounds.

Only relatively recently, some FET-based biosensors have integrated Al₂O₃ [24] on MoS₂ semiconductor and bioreceptor elements [70]. Despite achieving a very low detection limit for the prostate-specific antigen (PSA) analyte (LOD of 1 fg/mL), the use of atomic layer deposition (ALD) technology, multiple functionalization and the use of antigen receptors make it an expensive sensor. On the other hand, the exposure to moisture led to the hydration of aluminum oxide, demonstrating a rapid aging process. This phenomenon was observed in a sapphire gate ISFET [71]. A recent review paper on the ISFET [72] evolution stated that an optimized process uses Al₂O₃ sensing film deposition, configured by the lift-off technology to laterally define this sensing layer.

To improve the adsorbent properties of the porous oxide, in 2025, some authors reported a novel transistor with a 10 nm Al₂O₃ layer on a Ga₂O₃ layer of different thicknesses between 10 nm and 230 nm, exploring the device performances and revealing efficient reliability for those transistor-based biodetectors, with sub-10 nm Ga₂O₃ thickness [73]. In general, there has been no particular focus on nanoporous Al₂O₃ in ISFETs. Looking for a metal biocompatible with living tissues, we find Ti in almost all prostheses [74,75]. Therefore, we will discuss the nanostructured titanium oxide, which is more widely used in the construction of transistor-based biosensors.

3.3. Nanostructured TiO₂ Grown on Si Wafer

Another oxide material that can be converted by anodizing into a nanoporous or nanostructured material is titanium oxide. Our research group had used nanostructured TiO₂ for enzyme immobilization since 2011 [36]. It was later proven that the defects and cracks that appeared above the TiO₂ in the enzyme membrane were not caused by the underlying oxide but rather by the adopted cross-linking method [76]. An SEM image of the glucose-oxidase enzyme immobilized on a Si wafer covered by successive layers of Si/SiO₂/TiO₂/nafion is presented in Figure 10: (a) top view 1 mm scale for the entire wafer covered by enzyme; (b) detail at 0.1 mm scale, revealing cracks and fissures in the enzyme membrane.

It is not at all efficient to purchase TiO₂ nanoparticles from which to create a film in liquid or gel states, which is then deposited on a Si wafer, because the adhesion of such a layer is poor. Instead, it is much more efficient to deposit Titan directly on a Si wafer, even if sometimes Si is coated with SiO₂ or Si₃N₄ [36]. In all these cases, the Ti surface can then be converted into nanoporous TiO₂ or nanostructured TiO₂ by anodization in different electrolyte agents at different electrolysis currents. By adopting such a method, nanoporous TiO₂ film was created in anatase phase, and it is visible by SEM in Figure 11a. In this case, a Ti film of 90 nm thickness was strongly anchored to the substrate by sputtering deposition on a Si wafer, p-type, initially doped by 10¹⁵ cm⁻³. The anodization cell had a calomel reference electrode, against which the working electrode and counter electrode were polarized at voltages from 0 to 10 V at a degree of 0.2 to 0.4 V/s. The electrolyte was an aqueous solution of phosphoric acid mixed with oxalic acid, keeping the pH at 6.8 ÷ 7.4 by a phosphate buffer [76]. Just after the electrochemical process, the TiO₂ layer was amorphous.

After a thermal process at 500 °C, the amorphous film gets TiO₂ anatase phase with deep vertical nanotubes and homogenous nano pores of 10–30 nm diameter from a top view (Figure 11a). An ATR-FTIR spectrometer (attenuated total reflection—ATR, within a Fourier Transform Infrared—FTIR spectrometer) that works in the spectral area of 4000–370 cm⁻¹ confirms the nanostructured TiO₂ presence (Figure 11b).

In order to immobilize the GOX enzyme above, the cross-linker was changed from nafion to glutaraldehyde, keeping the same nanostructured TiO₂ formula. The present spectra were recorded by FTIR to prove the GOX–glutaraldehyde anchoring on the Si/SiO₂/TiO₂ surface (Figure 12a). The bands of 1161 cm⁻¹ and 1118 cm⁻¹ are allocated to the stretched Ti-O-C bond, demonstrating a strong anchoring of the GOX cross-linker to substrate. A top uniform film without defects is visible in an SEM view (Figure 12b).

But nanostructured TiO₂ was also studied outside the goal of its enzyme binding. Compared to the case of nanoporous TiO₂ annealed at 500 °C from Figure 11a, a study of modifying the technological parameters was continued. If the anodizing potential between the electrodes is varied in a larger range of 2–25 V, keeping 500 °C as the post-annealing temperature, the surface morphology at the wafer edge presents 40 nm pores (Figure 13a).

If, additionally, the post-anodization temperature was increased to 800 °C in the nitrogen atmosphere, the surface morphology proves better uniformity but similar pore sizes (Figure 13b).

The nanoporous aspect on the surface should be explained by the nanotube roots anchored deep inside the oxide after anodization [77]. Cross-section images of anchored TiO₂ nanotubes achieved by anodization in aqueous and glycerol electrolytes at four voltages are presented in Figure 14.

In the case of glycerol as the electrolyte, the TiO₂ surface is not so rough. It is more superficially anchored, but it is more ordered. These observations lead to the recommendation of aqueous electrolytes for chemo- and biosensor applications.

In 2019, a pH sensor that was ISFET-based with extended gate area and covered by TiO₂ film was reported [78]. The TiO₂ nanostructures were prepared on a fluorine-doped tin oxide as substrate, resulting in so-called TiO₂ nanoflowers (Figure 15).

The TiO₂ precursor solution was premixed with DI water and 36% hydrochloric acid, while titanium(IV) n-butoxide (>99%, 2 mL) was added. The best variant of this sensor measures pH between 2 and 12, with an almost perfect linearity of 99.91% and a maximum sensitivity of 46 mV/pH [78], while other nanoporous TiO₂ sensors, created in the same period, offered only 19.3 mV/pH sensitivity [79].

Very recent technological studies describe a 1.5 µm thick titanium layer sputtered onto the SiO₂ surface, which was previously grown by thermal processes [80]. The electrolyte for anodization comprised ethylene glycol with 0.35 wt% NH4F and 1 wt% deionized water. Finally, titanium dioxide nanotubes get a height of 1.2 µm and an inner diameter of the tube of 46 nm (Figure 16). Usual post-anodization thermal treatments were applied at 450 °C to obtain the crystalline state.

4. Other Oxides Used in ISFET and Related-Transistor Construction

Other nanostructured oxides have been used, too, as a sensitive element for ISFET transistors. Here are a few examples. A transistor that avoids the construction of a reference electrode inside the gate space uses ZnO prepared by the sol-gel method to provide a pH-sensitive element and offers a sensitivity of 27 mV/pH [81]. However, to be able to immobilize an enzyme within an EGFET biosensor, the oxide must be nanostructured in a so-called ZnO nano-array for glucose detection [82].

Although created in 2009, a pH ISFET based on vanadium oxide and placed over a hexadecylamine membrane within the gate space offered superior sensitivity of 38 mV/pH [83]. However, in 2025, the vanadium oxide combined with graphene oxide (GO) nano-compound is coming back to the attention of glucose biosensor manufacturers due to its excellent selectivity and sensitivity [84]. In all cases, the synthesis of Mn(V₂O₆)/GO started from NH₄VO₃ as a precursor, following a simple and cheap technology.

Another oxide borrowed from modern transistor technology that is part of the class of high-k dielectrics was applied in the construction of an EGFET biosensor for blood marker proteins. This is HfO₂.

By recent technology, the deposition of a thin HfO₂ film in the gate space can be performed at a thickness of 20 nm by high-power pulsed magnetron sputtering [85]. The performances allow the clinical detection of Parkinson’s biomarkers: a linear response of 99.45% degree in the detection range of 0.0001–1 ng/mL, a limit of detection of 198 fg/mL and a sensitivity of 12.11 mV/decade.

In order to fabricate a high-sensitivity FET-based dopamine biosensor, another oxide plays the role of a semiconductor. Consequently, an ultrathin In₂O₃ film of approximately 4 nm thickness was fabricated using a sol-gel method under an inexpensive, large-area spin-coating process [86]. The technology consisted of chemical lift-off lithography utilizing self-assembled monolayers of alkane-thiols on gold. The ingenious integration technique allowed a bottom-gate, bottom-contact FET coupling with an oxide surface functionalization, capable of immobilizing DNA aptamers as dopamine receptors [86].

The semiconductor properties of this nanostructured In₂O₃ material, which can effectively bind certain analytes, have been previously demonstrated, too [87,88]. The transfer characteristics reveal high drain currents stimulated by moderate operation gate voltages, a blocking current I_OFF = 2 × 10⁻¹¹ A at V_GS < 0 V to a conduction current I_ON = 70 μA at V_GS = 30 V, while V_DS = 30 V. The transistor parameters, like I_ON/I_OFF ratio of 10⁷, carrier mobility sub-10 cm²/Vs and threshold voltage around 4 V are in agreement with the transistor size—35 μm channel width and 15 μm length—plus constituent materials—Au electrodes on SiO₂ on Si wafer and heavily doped Si substrates as the back-gate electrode. Also, the output characteristics of this In₂O₃ transistor show a saturation of the drain current to 15 μA when V_GS was 20 V and V_DS was greater than a saturation voltage of 16 V, which is in agreement with the usual values for a thin film transistor that is bottom-gate commanded [89,90,91,92].

From the bio-sensing point of view, the dopamine biosensor with ultrathin In₂O₃ film used a bottom-gate configuration with interdigitated drain and source electrodes to increase the contacted area [76,81,93]. The DNA aptamers were the dopamine receptors. They were entrapped on In₂O₃ top film by intermediate linkers (3-aminopropyl)trimethoxysilane and 3-maleimidobenzoic acid N-hydroxysuccinimide) [86]. The aptamers’ attachment onto the channel surfaces produced a tenfold decrease in the transistor current [86]. The effect can be clarified by the electrostatic attraction/repulsion induced at the channel surface by the negatively charged DNA molecules. Finally, they repel electrons and push the n-type In₂O₃ film in a depletion work regime [94].

On the other hand, environmental parameters can hardly affect the preservation of sensitive electrodes, making them vulnerable to moisture. Therefore, humidity sensors have recently been proposed, either based on nanoporous materials [95], polymers [96] or combinations [97]. Nanotechnology has significantly improved the performance of humidity sensors by enabling a highly porous structure with a large specific surface area, which enhances the adsorption of water molecules. For this purpose, tantalum oxide, TaO₂, was recently used to protect the SiO₂ oxide in the gate of an ISFET dedicated to ion or free radical measurements in agriculture from soil [98]. Figure 17a,b presents the structure of the Front-Gate Field-Effect Transistor (FGFET) using functionalized Multi-Walled Carbon Nanotubes (MWCNTs). Bioactivation of MWCNTs by chemical bonding between Micro-Cystin-Leucin-aRginine (MC-LR) targeted aptamers (MCTAs) and carboxylated MWCNTs is shown in Figure 17a.

Figure 17c shows the fabricated FGFET sensor possessing interdigitated source and drain electrodes as in other practices [82,95], but here distanced at 20 µm [98]. Figure 17d shows details of MCTA-MWCNT nanotubes with 26 nm width by SEM imaging [98].

An impact parameter is the sensor stability, as the time after which the sensor response remains stable over time. For example, measurements of H₂PO₄⁻ ions have revealed stability times of 50 s for the ISFET sensor [98].

Aided by nanotechnology, ISFET transistors are in continuous development, both in terms of modeling [8,51,52,53,54], nanoporous materials comprehension [99], technology [100] and applications [101]. Combining new nanoporous materials that increase sensor sensitivity with ISFETs that are fabricated by ultimate standard technologies (e.g., 180 nm for analog applications), some researchers produced a non-invasive sweat sensor for four different ionic analytes, consuming only 2 pico-Watts [101]. To form a sensitive layer to the hydrogen ions, the deposited metal was Al, followed by anodization to form nanoporous Al₂O₃, a material previously described. In order to make it selective to the other three ions, the nanoporous oxide was separately functionalized, embedding an ionophore receptor in a polyvinyl-chloride/bis (2-ethylhexyl) sebacate selective membrane [101]. The analyte detection (H⁺, Na⁺, K⁺ and Ca²⁺) was achieved with the following sensitivities: 58 mV/dec for H⁺, 57 mV/dec for Na⁺, 48 mV/dec for K⁺ and 26 mV/dec for Ca²⁺ [101].

In order to capture the main properties for the proposed nanostructured materials, the fabrication processes, biosensing properties, advantages and limitations of each material are comparatively evaluated in

Table 1. Different nanostructured materials applied in sensing applications.

Nanomaterial	Reported Year	Film Properties/ Specific Parameters	Technological Hints	Applications	Final Parameters/Limitations
PS (porous Si)	2005	Pores 1 ÷ 2 µm [60]	Anodizing HF: DMF	Human breath capacitive biosensor	2.5 × 10−10 F/1 exhale, Time recovery = 20 min
	2007	Pores 8 nm ÷ 2 µm [61]	Anodizing HF–Ethanol or HF:DMF	PV cells	Reflectance range 5–12%
	2022	Pores 2 ÷ 50 nm [62]	Anodizing HF in DMSO or isopropanol, PS at 8 mA/cm2	Platforms for surface-enhanced Raman scattering (SERS) spectroscopy	SERS verifications for PS-based sensor at 1 uM dye rhodamine as analyte
PSi-NTs (porous silicon nanotubes)	2024	Pores 20 ÷ 50 nm [65]	AuNPs functionalized to MPBA on PSi-NTs using ZnO seeds on Si wafer	Glucose biosensor, other SERS platform	Linear range 0.5 ÷ 80 mM, LOD = 0.5 mM
Nanostructured Al2O3	2010	Nanoporous Al2O3 pores of few µm [21,69]	Anodization at 80 V, in electrolytes: (HCOO)2: H3PO4	Glucose biosensor by GOX enzyme entrapping	Poor enzyme adherence
	2023	Biomolecular binding on nanoporous MoS2 channel using Al2O3 grafting layer [70,71]	ALD of Al2O3 on MoS2 semiconductor	PSA biosensor	Excellent LOD of 1 fg/mL Expensive technology
Nanostructured TiO2	2011 2017	TiO2 pores of 7 ÷ 40 nm [36,76]	Ti deposition followed by anodization at 10 V in phosphoric acid mixed with oxalic acid	Glucose biosensor by GOX enzyme entrapping	Cross-linker is crucial for GOX immobilization; optimal TiO2 porosity for GOX 5 ÷ 50 nm
	2021	TiO2 pores~20 nm [80] Post-anodizing annealing film at 500 ÷ 800 °C [39]	Technol. recommendations: avoid glycerol electrolyte [77]	Glucose biosensor by GOX enzyme entrapping	Linear range was 0.1 µM-1 µM if ISFET works in linear regime at VDS = 0.2 V; 0.001 mM-100 mM if ISFET works in saturation at VDS = 2 V [39]
	2019	TiO2 nano-flowers of 2 ÷ 10 µm size [78]	TiO2 was prepared on a fluorine-doped tin oxide as substrate, using precursor mixed in DI water, HCl, titanium(IV) n-butoxide [78]	pH sensor, ISFET-based with extended gate	Excellent linear range pH 2–12, with 99.91% linearity; sensitivity of 46 mV/pH
Nanostructured vanadium oxide MnV2O6/GO	2025	2D graphene oxide/manganese vanadium oxide nanocomposite synthesized via eco-friendly, microwave-assisted method [84]	MnV2O6/graphene oxide (GO) composite, nanosheet-like structure	Tailored for the photoelectrochemical (PEC) detection of glucose	Limit of detection, LOD of 0.13 µM for glucose

.

5. Enzyme-FET and EGFET as Closest Relatives of ISFET

The first Enzyme-FET was called a “smart ISFET” in the 1970s because it selectively measures the ion concentrations by a key enzyme from only one biological component [102]. The method is actually a reduction of a more complicated problem to a problem that we already know how to solve. It was difficult to detect glucose among so many components in the blood, but if a key enzyme like GOX is added, then in the next minute, only the glucose is oxidized and generates ions that can be easily detected with an already invented ISFET [102]. In any Enzyme-FET, a key enzyme extracts from an analyte of the living world some ions, usually H⁺ or O⁻, or NH₄⁺, for which concentrations can be monitored by ISE or ISFET.

However, two elements give specificity to the Enzyme-FET concept: (i) the technology of immobilizing the enzyme near a transistor; and (ii) the ideal location to place the enzyme membrane. The first Enzyme-FETs placed the enzyme on top of the gate oxides, but separated by a membrane. Among these detectors, the first urea sensor was an ISFET that immobilized urease enzyme in acrylamide polymer on the surface of a cationic electrode sensitive to NH₄⁺ ions, using only ISE principles and free of any transistor [103]. Twenty years later, a urease-based ISFET biosensor was depicted, which was affected by interferents with sodium and potassium ions. In general, the polymeric membrane was semipermeable and acted as a first physical filter for blood components to reach the enzyme itself. But those membranes were much thicker than the entire thickness of the transistor, and the analyte–enzyme signal transmission to the transducer was very poor, so the membrane was eliminated.

The current Enzyme-FETs are ISFET structures with an enzyme membrane on top (Figure 18a), or EGFET with the separate sensitive area, which is connected with metal traces to the transistor gates (Figure 18b) [80]. Both solutions have advantages and disadvantages: the sensitivity of the sensor decreases the further we move the enzyme–droplet system from the gate area, but as we move the system closer to the gate, transistor damage can occur more easily.

In the case of the transistors in Figure 18a,b. the generated ions from the enzyme–droplet system act as an induced charge in the gate that influences the current through the transistor: negative ions will deplete the n-channel of electrons, while positive ions will enhance the n-channel. In all these cases, the current will be sensitively dependent on the amount of generated ions, which in turn depends on the consumed analyte via an enzyme reaction mole by mole.

In 2024, a recent work reports the experimental results of GOX immobilization on a separate area of nanoporous TiO2 of 10 nm porosity on a Si wafer, with proper uniformity and anchoring [80]. Figure 19a presents the SEM image of the top surface of the synthesized nanoporous TiO₂ layer achieved by anodization, not directly on a Si wafer but over a 90 nm SiO₂ that was previously thermally grown. The best process was accomplished for an anodizing potential of 8–10 V at a pH of 6.6, achieving a porosity of 10 nm and successfully used for enzyme entrapping. Figure 19b confirms the uniformity of the GOX layer over nanoporous TiO₂ by SEM.

Other authors describe enzyme layer immobilization on Nanoporous Gold (NPG) [104]. NPG has the advantage of stability and highly conductive properties for biosensors containing oxidoreductases and hydrolases. Terminal functional groups induced by Self-Assembled Monolayers (SAMs) in NPG materials allow bio-conjugation reactions for enzyme binding. Due to the exceptional oxidation rate of NPG towards adenine dinucleotide (NADH), a co-catalysis with nicotinamide adenine dinucleotide (NAD+) allows an amperometric biosensor that detects NADH with good linearity from 0.05 mM to 2 mM with an LOD of 15 μM and a sensitivity of 1.58–1.72 μA/mM [105].

In 2024, an enzyme-free glucose biosensor was fabricated using nanoporous TiO₂ material doped with Vanadium [106] but in the absence of co-integration with a small-sized transistor. However, the direction of study of nanoporous TiO₂ structures combined with hydrothermally derived metals Cobalt (Co), Nickel (Ni) and Vanadium (V) of the type M_xTi_1-xO₂, where x can be 0.01–0.03 [107], has captured the attention of the chemosensor community in order to obtain more stable, more active inorganic catalytic receptors than enzymes. The V_0.03Ti_0.97O₂ material offered a high sensitivity of 1129.3 μA/mMcm² and a limit of detection of 1.8 μM for glucose [106].

Very recently, special attention has been paid to Carbon-based nanostructured materials applied in chemosensors for humidity and free of enzymes, too. A breath humidity sensor based on cellulose nano-paper, was successfully used to detect both breathing and skin moisture desorption at a distance of 2 cm [108]. An improved sensitive nanoporous film composed of a cellulose nanofiber/Carbon nanotube combination was proposed to increase water molecule adhesion onto the sensor surface [109]. Another study in 2025 reminded us how biomimetic materials like silk fibroin (SF) hydrogel were integrated within a Graphene Field-Effect Transistor (GFET) for a glucose biosensor [110]. The enzymatically cross-linked SF was an excellent carrier for GOX and offered improvements, such as a long lifetime for GOX and reduced adhesion for interferents. The enzymatic FET biosensor proved an LOD sub-200 nM and a linear detection range from 1 µM to 10 mM for glucose [111].

From the above discussions, it can be observed that in the last five years, advanced experience has been accumulated in integrating enzymes into various nanoporous materials in isolated areas. Even if this co-integration has often taken place in the absence of a transistor (e.g., in the case of electrochemical biosensors), this know-how nevertheless opens up future development opportunities for EGFETs.

6. Discussions About Future Directions in ISFET Development

Beneficial results, such as Nernstian detection limits, sub-picomolar sensitivities and sub-pico-Watts consumed power, can only be successfully achieved by combining nanoporous or nanostructured materials and miniaturized ISFET transistors, which are currently reaching from 300 nm down to 2 nm [112]. The voltage sensitivity Sv, called the Nernstian limit, is defined as the derivative: Sv = dV_G/dlog₁₀ I, where V_G is the gate voltage and c is the ion concentration, Sv being measured in mili-Volts per decade of ion concentration, which reaches an upper limit value [101]. If, additionally, the ISFET transistor will be operated in the subthreshold regime, where an exponential drain current, I_D, dependence versus the gate voltage, V_G, occurs, I_D~exp (qV_G/kT) [101], an almost limit-Nernstian sensitivity, Sv, of 59 mV/decade will be reached. Only a few ISFETs exceed the Nernstian limit and become super-Nernstian devices, as published in a few papers [113,114]. Other theoretical demonstrations prove that super-Nernstian sensitivity is based on a circuit amplification rather than nanomaterial optimization [115]. The pH sensitivity can be directly computed from the log(I_D)−V_G curves in terms of shift of V_G for a decade, decreasing the drain currents in weak inversion [101]. The pH ISFET for sweat offered a subthreshold slope of 85.3 mV/dec when the drain current was represented at a logarithmic scale [101].

Recently, another oxide has been investigated to become a sensitive layer within an ISFET, namely, tin oxide SnO₂. However, the resulting ISFET device combined multiple oxides to decorate both top gate (TG) and bottom gate (BG) space [115]. A high-k dielectric material was selected in the TG space: SiO₂ 20 nm covered by Ta₂O₅ 80 nm, to improve the silicon/oxide interface, especially near the channel location. The SnO₂ 50 nm was used as a sensitive insulator for an extended gate configuration. To achieve a 5:1 amplification ratio, the top oxides were engineered to possess an equivalent oxide thickness less than one-fifth that of the SiO₂ film from the bottom gate (Figure 20a).

The starting substrate for the sensitive part was a glass wafer covered by 300 nm indium tin oxide (ITO), then coated with 50 nm SnO₂ by radio-frequency magnetron sputtering. Further functionalization added some OH groups on the surface, followed by APTES exposure by the vapor-phase reaction method [116]. The final ISFET structure is presented in Figure 20b, and the Bovine Serum Albumin (BSA) biosensor, aided by ISFET with extended gate, is shown in Figure 20c. When the ISFET transistor was operated without analyte in the Single Gate (SG) configuration, it offered slightly better electrical device parameters than when it was Double Gate (DG) operated, giving just one example: SS|_SG = 137 mV/dec < SS|_DG = 213 mV/dec [115]. From the sensing point of view, the situation reversed for the SnO₂-based biosensor that was DG actioned. DG ISFET with sensitive SnO₂ reached a maximum super-Nernstian sensitivity of 141.19 mV/decade of BSA concentration, versus the optimal SG variant that reached 29.08 mV/decade of BSA concentration (Figure 20d). However, Figure 20d also reveals inferior sensitivities for the DG ISFET, with sensitive SiO₂ reaching 71.59 mV/dec, while a minimum sub-Nernstian sensitivity of 14.83 mV/dec is offered by the biosensor with a SiO₂ layer acted by a single gate.

The 10-fold sensitivity amplification, only in the presence of the SnO₂ layer, opens new visions of biosensors without any additional external electronic circuits.

Another technological issue faced with the reliable realization of ISFETs is the occurrence of traps and defects at the semiconductor/oxide interfaces. These defects considerably reduce the current and the sensitivity of the sensor. Usually, thermal treatments up to 800 °C are applied for standard MOS transistors [52,117], but these methods are no longer compatible with the co-integration of sensitive materials, especially biomaterials. Recent state-of-the-art approaches employing UV radiation to remove oxide-trapped charges have yielded satisfactory results [118]. However, the time processing for a device-to-device calibration took 17 h of UV radiation per ISFET transistor.

Nanostructured materials have also been cleverly integrated to achieve super-Nernstian sensitivity by a label-free sensor, accompanied by a double gate CMOS transistor. The sensor contained silicon nanowire arrays co-integrated on a silicon on insulator (SOI) wafer to detect C-reactive protein (CRP). The sensor shows excellent stability (20 pA/min) and an outstanding current sensitivity up to 1.2 nA/decade for CRP proteins in the linear range of detection [119]. Silicon nanowire-based sensors were applied for the detection of other proteins, too, like troponin [120], deoxyribonucleic acid (DNA) [121], micro-RNA [122] and viruses [123], and it has also been used for CRP detection, but for a poorer concentration range [124]. The linear interval of detection was 60 ng/mL–100 μg/mL, covering the whole concentration range of human CRP from the clinical samples [119].

Another target for the near future is reaching a sub-pico-molar detection limit. Sensors based on silicon junctionless transistors in nanowire configuration were able to detect concentrations as low as 580 zM zeptomolar of the streptavidin protein, at the closest level to single-molecule detection [125,126,127].

7. Conclusions

The paper reviewed the ISFET evolution, which integrates a sensitive intermediate layer made by nanoporous, nanostructured or specific oxides near the electronic device. One of the main technological goals was to highlight those methods capable of anchoring the nanoporous layer to the substrate, usually a Si wafer. The general techniques consist of depositing metal layers with high adhesion to the substrate (semiconductors or oxides) by classical and reliable methods from microelectronics, such as chemical vacuum deposition, lift-off, sputtering and laser deposition. The usual thicknesses of the deposited metal vary in the range of 4 nm–100 nm. Then, the surface of the metal electrode is converted into nanoporous oxide, with deep anchors of nanotubes from the constituent material. For this purpose, an electrochemical cell must be used, as inert as possible, as the material of the cell, in which the surface of the metal that must become nanoporous plays the role of the working electrode. Using various electrochemical currents, various polarization voltages and a wide range of electrolytes, various porosities, various geometries, which have an exponentially multiplied surface, and different sensitive areas to an analyte were obtained. The key element for future biosensor scalability is the co-integration of these nanoporous intermediate layers with extremely small transistors. This solution opens the door for uni-molecular detection. Nowadays, the use of Si nanowire transistors has pushed this detection limit to the atto-molar (10⁻¹⁸ M) and even zepto-molar (10⁻²¹ M) range for streptavidin detection. On the other hand, the ISFETs biased by two gates, which include sensitive oxide of SnO₂,too, offer a maximum super-Nernstian sensitivity of 141.19 mV/dec. This 10-fold sensitivity amplification opens the further development of biosensors free of additional external electrical circuits, with a huge advantage for the device price.

Some technological gaps still exist. Only a short list of biological receptors, like enzymes, aptamers and antibodies, can be integrated through nanoporous intermediate layers with classical transistors. Consequently, biosensors for a very wide variety of analytes at affordable prices do not exist on the market. Some technological gaps have been avoided, for example, by skipping bioreceptors that are easily degradable over time, such as enzymes or antibodies, and manufacturing biosensors free of enzyme using inorganic catalysts. Nanoparticles are only usable in large electrochemical cells, as they are incompatible with size scaling. ISFET transistors and other related sub-300 nm FETs are essential for this co-integration because they are standardized and can generate mass production.

In conclusion, the emerging trends of ISFET sensors in the future will necessarily be interdisciplinary, co-integrating micro- and nano-electronic devices, ultimately scaled nano-transistors, intermediate nanomaterials, nanoporous layers, nanostructured films and more refined receptors.