Encyclopedia of Nanotechnology

2016 Edition
| Editors: Bharat Bhushan

Nanostructure Field Effect Transistor Biosensors

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DOI: https://doi.org/10.1007/978-94-017-9780-1_338
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Synonyms

 Bio-FET;  Chem-FET;  CNT biosensor;  CNT-FET;  DNA FET;  Nano-FET;  Nanowire FET biosensor

Definition

Nanostructure field effect transistor (nano-FET) biosensors are nanoscale electronic devices used to detect and measure the presence and concentration of specific biological molecules within a given sample solution. These sensors utilize a quasi-one-dimensional semiconducting nanostructure as the active sensing element to detect the presence of local electric fields produced by charged biological molecules.

Overview

Biomolecular Detection Technologies

Methods for detecting and measuring the presence, abundance, and activity of biological molecules, such as proteins, DNA, and RNA, enable the investigation of basic biological function, molecular diagnosis of diseases, and development of therapeutic treatments. The combination of polyacrylamide gel–based electrophoresis separation (and related blotting techniques) and absorption-based chromogenic dye staining (e.g., Coomassie brilliant blue or silver staining) has traditionally formed the core technologies for protein and nucleic acid detection. Subsequent development of fluorescence- and luminescence-based labeling stains offered the opportunity for multicolor labeling, making multiplexed analysis possible. However, these methods were limited to the detection of biological molecules on gels and blots. Immuno-labeling of fluorescent and luminescent probes allowed specific protein detection and facilitated the emergence of immunoassays able to detect and quantify the abundance of a target protein molecule in solution. These optical detection methods suffer from large sample volume consumption, high reagent costs, and long assay times which, when combined with the drawback of using labeling molecules, preclude the application of these technologies to many laboratory and clinical applications (e.g., real-time protein quantification and point-of-care clinical diagnostics).

A number of novel label-free protein-detection strategies have since been developed in an effort to overcome some of the aforementioned limitations. Many of these strategies leverage nonoptical transduction modalities to circumvent the limitations imposed by optical probes. Examples include micro-/nano-surface plasmon resonance sensors [1], micro- and nano-cantilevers that translate biomolecular interactions into mechanical deformations [2], nanostructured field effect transistors (FETs) that measure intrinsic biomolecular charge [3], and electrochemical sensors that translate biomolecular adsorption to changes in redox current [4]. Of these technologies, nano-FET biosensors offer unique device characteristics including ultrasensitive, real-time, and label-free measurement capability, compact physical form, CMOS-compatible and simple on-chip integration with circuitry and microfluidic systems, direct electrical readout, and multiplexed detection capability.

Structure

Nano-FET biosensors are composed of a semiconducting quasi-one-dimensional nanostructure (e.g., nanowire or nanotube) bridging the source and drain electrodes (Fig. 1). The nanostructure serves as the transistor’s channel and functions as the sensing element of the device. The biosensor structure deviates slightly from that of the transistor in the placement and structure of the gating electrode. While the gate electrode in a nano-FET is typically located directly above the channel, separated by a thin dielectric layer, the nano-FET biosensor uses a solution gate electrode that is typically composed of platinum or Ag/AgCl and is immersed within the aqueous solution being measured. Furthermore, the nanostructure surface is functionalized with a layer of analyte-specific receptors to impart a degree of specificity to the sensor. This may be accomplished via chemical bonding using bifunctional linker molecules, or through physical interaction due to adsorption, van der Waals forces, and weak electrostatic forces.
Nanostructure Field Effect Transistor Biosensors, Fig. 1

(a) Schematic of a single nano-FET biosensor. (b) Schematic of a nano-FET biosensor chip with integrated microfluidic sample delivery system. The nano-FET biosensor array is housed within the microfluidic system (Adapted with permission from Macmillan Publishers Ltd. [5], copyright 006)

While the specific nanostructure employed may vary in shape, size, and composition, it is important that these structures be quasi-one-dimensional and exhibit semiconductive material properties. Typical nanostructures used in the construction of nano-FETs include single and multi-walled carbon nanotubes, graphene, and silicon nanowires. Devices constructed from other semiconductive materials such as metal-oxides and nanostructures of various shapes including nanotubes, nanobelts, nanorods, nanoribbons, and nanobars have also been reported.

Physics

Nano-FET biosensors utilize semiconductive nanostructures to detect the presence of charged molecules in the immediate vicinity around the nanostructure surface. The sensing mechanism is based on the observation that an externally applied electric potential is capable of modulating the number of free charge carriers within a semiconductive material, thus leading to a measurable change in electrical conductance. For a cylindrical semiconducting nanostructure with radius R and length L, the charge carrier concentration may be described as a function of the applied potential:
$$ {n}_0=\frac{C\left({V}_G-{V}_t\right)}{q\uppi {R}^2L} $$
(1)
where C is the gate capacitance, q is the elementary charge, V G is the gate voltage, and V t is the threshold voltage [6]. As the conductance of the semiconducting nanowire is described by:
$$ G=nq\mu \left(\frac{\uppi {R}^2}{L}\right) $$
(2)
the current passing through the nanowire (I D ) for a given source-drain voltage (V DS ) is therefore:
$$ {I}_D=\frac{\uppi {R}^2}{L}{n}_0q\mu {V}_{DS} $$
(3)
where μ is the carrier mobility. Charged functional groups present on the surface of bound biological molecules such as proteins and nucleic acids provide the necessary electric potential required to induce a conductance change in the semiconductive nanostructure, thus enabling the direct electrical detection of such molecules using nano-FET devices (Fig. 2). The strength of a molecule’s surface charge is a function of electrolyte pH and is characterized by the molecule’s isoelectric point.
Nanostructure Field Effect Transistor Biosensors, Fig. 2

Real-time nanowire FET sensing results. (a) Conductance versus time data recorded following alternate delivery of prostate-specific antigen (PSA) and pure buffer solution (1 μM phosphate (potassium salt) containing 2 μM KCl, ph 7.4). Subsequent PSA concentrations were 5 ng ml−1, 0.9 ng ml−1, 9 pg ml−1, 0.9 pg ml−1, and 90 fg ml−1, respectively. (b) Complementary sensing of PSA using p-type (NW1) and n-type (NW2) nanowire FET devices. Vertical lines correspond to addition of PSA solutions of (1, 2) 0.9 ng ml−1, (3) 9 pg ml−1, (4) 0.9 pg ml−1, and (5) 5 ng ml−1. Arrows on the bottom represent the injections of sensing buffer solution (Adapted with permission from Macmillan Publishers Ltd. [5], copyright 2006)

The simplified analysis qualitatively illustrates the physical sensing mechanism behind nano-FET biosensors. However, this treatment does not accurately describe the complexities of the nano-FET biosensor system. Equations 1, 2, and 3 describe the electrical behavior of a cylindrical semiconducting crystal subjected to a uniform charge density over the entire nanostructure surface. This assumption fails to take into consideration the nonuniform nanostructure surface charge distribution resulting from discrete biomolecular binding events. The discrete gating mechanism by individual molecules in solution introduces several additional phenomena that significantly influence the sensor’s response including depletion of analyte from solution, analyte transport mechanisms to the nanostructure surface, and analyte-receptor binding and unbinding kinetics. These effects become particularly important in the detection of rare analytes.

Equation 3 also neglects the influence of electrostatic screening – the spatial arrangement of mobile ions around a charged molecule to effectively neutralize that charge. Importantly, target analytes must be sufficiently close to the sensor surface in order to elicit a conductance change due to electrostatic screening by mobile charges around the analyte molecule (e.g., ions in solution and charge carriers in the semiconducting crystal). The Debye screening length, the distance away from a charged molecule at which the electrostatic potential is diminished to zero due to screening, defines the required proximity.

Debye length (aqueous electrolyte):
$$ {k}^{-1}=\frac{1}{\sqrt{8\uppi {\lambda}_B{N}_AI}} $$
(4)
where I is the ionic strength of the electrolyte (mole/m3), N A is the Avogadro number, and λ B is the Bjerrum length of the medium (0.7 nm for water).
Debye length (silicon):
$$ {L}_D=\sqrt{\frac{\upvarepsilon_0{\upvarepsilon}_r{k}_BT}{q^2{N}_D}} $$
(5)
where ε r is the dielectric constant of the semiconducting material, k B is the Boltzmann’s constant, T is the absolute temperature in Kelvin, q is the elementary charge, and N d is the density of donors in a substrate. The implications of charge screening and a finite Debye length make several important predictions on nano-FET biosensing performance. Namely, these relations suggest a conductance dependence on charge carrier concentrations (ionic strength of an electrolyte solution and doping density of a semiconductor), size of the semiconducting crystal, strength of the molecule’s net surface charge, and the nanostructure-analyte separation distance. Inclusion of these phenomena into a unified analytical model that accurately describes experimental nano-FET biosensor response is an active area of research.

Performance Parameters

Nano-FET Sensitivity

Nano-FET sensitivity is the induced change in device conductance (ΔG sd ), upon exposure to a certain biomolecular stimulus for a constant source-drain voltage (Vds). Sensitivity is defined as:
$$ \mathrm{Sensitivity} = \frac{\Delta {G}_{sd}}{G_0} $$
(6)
where ΔG sd is the direct conductance change observed in sensing experiments and G 0 is the initial device conductance [7]. Because ΔG sd depends on the specific nanowire parameters (e.g., diameter, mobility, etc.) this value does not reflect the intrinsic sensitivity. Instead, it is more meaningful to characterize the sensitivity using a dimensionless parameter (6). Normalization enables comparison across devices with different physical dimensions.

Sensitivity is closely related to three additional parameters: detection limit (the smallest analyte concentration that can be measured, which is dictated by the system signal-to-noise ratio); measurement resolution (the smallest detectable change in analyte concentration, which is dictated by noise levels); and measurement range (the measurable analyte concentrations, with the upper and lower bounds defined by nanostructure saturation and the detection limit, respectively). A higher sensitivity directly translates into lower detection limit, finer measurement resolution, and a broader measurement range. Typical protein and DNA detection limits reported in literature are in the pM (picomolar) to fM (femtomolar) range [8, 9]. Single virus detection has also been achieved [10].

Specificity

Specificity is a measure of a biosensors relative responsiveness to the target analyte as compared to all other molecules present within the sample solution. Very specific sensors should possess high sensitivity toward the target analyte and low sensitivity toward all other biomolecules. This attribute is imparted onto a nano-FET biosensor by functionalizing the nanowire surface (either covalently or noncovalently) with a confluent monolayer of analyte-specific receptors/ligands that bind specifically to the molecule of interest.

Response Time

Response time is the time it takes for the biosensor system to obtain a stable biomolecular concentration measurement upon introduction of the sample solution to the sensor surface. As with any dynamic system, this characteristic is captured in the step response of the system by the settling time. For a nano-FET biosensor, this parameter is influenced by the analyte-receptor binding and unbinding kinetics, transport of analyte molecules to the sensor surface, concentration and degree of mixing of the analyte solution, and the electrical response behavior of the measurement system. The typical response time of nano-FET biosensors is on the order of minutes.

Key Research Findings

Active areas of research in the field of nano-FET biosensing include: (1) optimization of biosensor performance based on the systematic examination of how fundamental device parameters affect sensitivity, specificity, and response time; (2) development of robust large-scale fabrication strategies for the manufacturing of biosensor arrays; and (3) development of an analytical model capable of accurately describing and predicting nano-FET biosensor performance.

Performance Optimization

Minimizing the Effects of Charge Screening

As the effect of charge screening acts to minimize the total nanostructure volume gated by surface charges, thus reducing device sensitivity, strategies that increase screening length are expected to improve device sensitivity. The maximum device sensitivity may be obtained in a situation where the effective screening length is much larger than the radius of the nanowire. Improved device sensitivity can thus be obtained through careful design and optimization of several device parameters that influence the effective screening length:

Diameter

Reduction of the nanostructure diameter will dramatically increase the surface-to-volume ratio, thus improving device sensitivity. This is the primary advantage of decreasing the channel size and explains why nanostructure FET sensors exhibit significantly higher performance as compared to traditional planar FET devices.

Dielectric Thickness

Silicon nanowires form a native oxide layer (∼1–2 nm) on their surface that serves as a dielectric layer reducing the electric field strength acting on the nanowire itself. This layer also increases the separation distance between surface-bound molecules and the semiconductive core of the nanowire. Thus, a thinner oxide layer is expected to improve sensitivity. Complete removal of the native oxide layer surrounding silicon nanowires has been shown to increase device sensitivity [11, 12].

Functionalization Scheme and Receptor Size
Decreasing the separation distance between the receptor-bound analyte molecule and the nanostructure surface minimizes the effect of electrolyte screening. While antibodies are commonly used to selectively bind analyte molecules to the biosensor surface, the use of smaller capture probes such as antibody fragments or aptamers in place of antibodies can increase device sensitivity without the loss of selectivity (Fig. 3) [13, 14]. Similarly, utilizing shorter bifunctional linker molecules to attach the capture probes to the nanostructure surface may also increase device sensitivity. The choice of functionalization strategy may also dictate the capture probe density and coverage on the nanostructure surface. A poor capture probe density can result in the nonspecific adsorption of molecules onto unmodified areas of the nanowire surface and may also limit the upper detection limit of the sensor due to capture probe saturation.
Nanostructure Field Effect Transistor Biosensors, Fig. 3

Schematic representation of nano-FET biosensors functionalized with (a) antibodies and (b) aptamers. Aptamers (and antibody fragments) are smaller than whole antibodies. The use of these capture probes thus results in increased nano-FET sensitivity. For certain experimental conditions, aptamers may serve as the ideal probe candidate as they are shorter than the system Debye length while antibodies are not (Reprinted with permission from [13]. Copyright 2007 American Chemical Society)

Electrolyte Ion Concentration and pH

The ionic strength of the electrolyte solution influences the efficiency of charge screening as the density of free ions is directly related to size of the electric double layer formed around a charged molecule. Higher ionic strength electrolytes are thus able to screen a charged molecule over a shorter distance. It is therefore advantageous to perform biosensing experiments in low-ion conditions to maximize sensitivity [15]. However, this is not always feasible as many sensing applications are performed with physiological samples (e.g., serum, whole blood, etc.).

The pH of a solution also influences device sensitivity by changing the effective electrical charge on a biological molecule. Biological molecules such as proteins and nucleic acids contain both acidic and basic functional groups that may be positively or negatively charged depending on the availability of protons (H+) present in solution. The pH for which the negative and positive charges on the molecule are balanced, resulting in zero net charge, is defined as the isoelectric point (pI). The net charge on a molecule is therefore a function of the electrolyte pH and can become increasingly more positively or negatively charged as the term |pI - pH| increases. While tuning this parameter may be useful in specific applications, the pH of a sample is usually fixed at physiological levels (pH = 7.4).

Carrier Concentration

The screening length within a semiconductive nanowire dictates the depth of penetration into the nanostructure for which surface-bound charged molecules are able to gate. For example, the screening length in a silicon nanowire operating with a typical carrier concentration of 1018–1019 cm−3 is ∼ 1–2 nm, suggesting that only a small portion of the total nanowire volume (a ∼ 1–2-nm thick layer at the nanowire surface) is affected by surface charges. Decreasing the charge carrier concentration will therefore increase device sensitivity by increasing the screening length and the total gated volume. This can be achieved by varying either of two device parameters: the nanostructure doping density or the bias voltage applied through a gate electrode on the nano-FET biosensor. In terms of semiconductors, lower doping densities result in longer screening lengths and therefore produce higher sensitivity devices [16]. It is instructive to note that the optimum biosensor sensitivity is obtained when the device is operating in the subthreshold regime, as under these conditions, the transconductance is at a maximum [7, 17, 18].

Optimization of Analyte Delivery Efficiency

Biomolecular detection using nano-FET sensors requires the delivery of analyte to the sensor surface. The efficiency of this process can therefore limit the performance of nano-FET biosensors. Several phenomena that influence this process include diffusive and convective molecular transport mechanisms of analyte to the biosensor surface, depletion of free analyte from solution, and nonspecific analyte adsorption and binding.

Molecular Transport
As with many surface-based biosensor systems, the transport of analyte in solution to the sensor surface plays a crucial role in governing binding kinetics and ultimately sensor performance. This is especially true when analyzing samples with dilute concentrations of target molecules and/or employing microfluidic systems for efficient and automated handling of small sample solution volumes. A number of competing physical processes influence target transport to the sensor surface. Analyte molecules suspended in solution may diffuse randomly within the solution, be convected along with flowing fluid, bind to adjacent surface-bound receptors, or subsequently unbind to reenter solution. The binding (or collection) of biomolecules onto the sensor surface simultaneously depletes the surrounding solution to form a so-called depletion region around the sensor (Fig. 4). As the depletion region grows in size, the analyte diffusive flux decreases and thus the collection rate becomes slower. Through finite element analysis, Sheehan and Whitman determined that the size and shape of the sensor profoundly affects the total analyte flux to the sensor surface (Fig. 5a) and further argues that without directed transport of biomolecules, individual nanoscale sensors would be limited to picomolar order sensitivity for practical time scales (hours to days) [20].
Nanostructure Field Effect Transistor Biosensors, Fig. 4

The steady-state depletion zone around a nanowire sensor from a pure mass-transport problem. The depletion zone is thick compared to the nanowire, but is substantially thinner than the channel itself (δs ∼ LPes−½ = 100 nm). Reaction-limited binding onto the nanowire follows the Langmuir binding curve (inset top right) (Adapted with permission from Macmillan Publishers Ltd. Nature Biotechnology [19], copyright 2006)

Nanostructure Field Effect Transistor Biosensors, Fig. 5

(a) Theoretical calculation of the time required for a 10-μm-long hemicylindrical sensor to accumulate 1, 10, and 100 molecules due to diffusive molecular transport. The sensor lies at the bottom of a channel whose width is equal to the sensor’s length and which is filled with a 1 fM analyte solution. (b) Finite element analysis (points) and theoretical calculation (lines) of the total flux of molecules to a hemicylindrical sensor in a microchannel (800 μm wide, 100 μm high). The total flux plotted is the steady-state value at the given volumetric flow rate for a 1 fM analyte concentration (Adapted with permission from [20]. Copyright 2005 American Chemical Society)

While the diffusive depletion zone grows indefinitely, albeit at an ever-decreasing rate, introduction of convective flow (assumed to be laminar) halts this growth resulting in a steady-state depletion zone with a length scale defined by the balance between the convective analyte flux delivered to the depletion zone boundary and the diffusive flux out of the depletion zone. By adjusting the convective flow rate through the channel, it is possible to tune the size of the depletion layer and thus control the rate at which molecules are delivered to the sensor surface (Fig. 5b). However, Squires et al. argue that the total mass transport varies only weakly with the flow rate as flow rates must be increased 1,000-fold to enhance flux by a factor of 10 [19]. The treatment thus far has assumed instantaneous analyte-receptor coupling upon transport of target molecules to the sensor surface. This represents one of two sensing regimes that can be described as being “transport-limited” as the availability of analyte adjacent to the sensor surface limits sensor response.

A second biosensing regime arises when considering the kinetics of analyte-receptor association and dissociation (assumed to be first-order Langmuir kinetics) [19]. If delivery of target molecules to the sensor surface occurs at a faster rate than the net analyte-receptor association rate, then the transport is said to be “reaction limited.” In general, the reaction kinetics is determined by the fidelity of the immobilized reagents. While each of the two regimes imposes an independent maximum target collection rate, the “reaction-limited” condition is most desirable for nano-FET operation as, for a given device, biomolecular sensing is purely a function of analyte concentration. The sensor size and flow rate should thus be adjusted to ensure nano-FET operation within this regime. Alternatively, introduction of turbulent flow or fluid mixing through innovative microfluidic design may also be employed to overcome the limitations imposed by convective and diffusive molecular transport [8].

Depletion of Free Analyte

As silicon and metal-oxide nanostructures share similar surface chemistries with several popular microfabrication substrates (e.g., silicon wafers, glass, quartz), surface modification techniques used to attach receptor molecules to the nanostructure surface can also functionalize the device substrate. In this situation, the total device surface area exposed to the sample solution during biomolecular detection experiments is proportional to the number of receptors available for reaction. Analyte binding events with substrate-bound receptors do not induce a change in nanostructure conductance. Rather, these reactions serve only to deplete free analyte molecules from solution, thus reducing the number of analyte molecules available for reaction with nanostructure-bound receptors. Minimizing the total surface area exposed to sample solutions is, therefore, important for maximizing device sensitivity. This effect may be circumvented using nanostructure-specific surface modification chemistries.

Nonspecific Adsorption
Random adsorption of biological molecules onto the nano-FET sensor produces a false-positive measurement signal. This effect may be minimized by grafting a dense layer of long nonreactive polymer chains (e.g., polyethylene glycol) onto the nanostructure surface alongside receptor molecules (Fig. 6). The polymer chains prevent molecular adsorption directly onto the nanostructure surface by acting as a spacer while the co-functionalized receptors are able to bring analyte molecules into intimate contact with the nanostructure through molecular binding. The length of the polymer chains used for this application should exceed the screening length of the system.
Nanostructure Field Effect Transistor Biosensors, Fig. 6

Schematic of a nanotube field effect transistor coated with a polymeric functional layer. Co-functionalization with a molecular receptor (biotin) and a nonreactive polymer (e.g., polyethylene glycol or polyethylene oxide) prevents nonspecific adsorption of molecules onto the nanotube surface (Adapted with permission from [21]. Copyright 2003 American Chemical Society)

Nonspecific adsorption of random molecules is especially problematic when working with samples containing numerous types of biological molecules such as whole blood. One strategy that enables detection of a specific analyte within a whole blood sample involves the use of a sample pre-purification procedure before nano-FET measurement [22]. In this method, the whole blood is introduced into a chamber in which analyte-specific receptors have been immobilized on to the chamber walls. Upon binding of free analytes to the surface-bound receptors, the whole blood solution is flushed out of the chamber and discarded. The receptors are then detached from the chamber surfaces and transported to the nano-FET biosensor for analysis.

Large-Scale Methods of Fabrication

Reliable large-scale fabrication of nano-FET biosensors remains to be a significant technological challenge, impeding the adoption of this technology in medical diagnostic and biomolecular detection applications. Unlike traditional microfabrication methods for producing micrometer-level structures, the difficulties encountered in nano-FET biosensor fabrication arise from the process of establishing electrical contact with nano-sized structures in a high-throughput and reproducible manner.

Early methods for manufacturing nano-FETs involve (1) the growth of individual nanostructures using a chemical vapor deposition process, (2) transfer of as-grown nanostructures from the growth substrate onto a device substrate (a process which results in the random placement and orientation of the deposited nanostructures), and (3) identification of a suitable nanostructure on the device substrate followed by the patterning of metallic electrodes onto the ends of that nanostructure using a mask-less nano-lithographic tool (e.g., electron beam lithography or focused ion beam lithography), which has a low throughput and is only suitable for proof-of-principle use. Recently devised batch-fabrication strategies for the wafer-scale production of nano-FET arrays can be broadly categorized into “bottom-up” or “top-down” fabrication methods.

Bottom-Up Fabrication Methods

One strategy for bottom-up fabrication is to make use of pre-synthesized freestanding nanostructures using epitaxial growth methods such as chemical vapor deposition. This strategy affords several advantages over top-down and other bottom-up approaches including (1) strict control over nanostructure dimensions and electrical properties, (2) the ability to assemble nano-FET devices onto flexible and transparent substrates, and (3) the ability to incorporate nanostructures with unique chemical compositions and architecture into nano-FET devices. Successful incorporation of these nanostructures into nano-FET devices relies primarily on the ability to transfer, align, and position the as-grown nanostructures onto a device substrate. Contacting electrodes may then be patterned using conventional photolithography to form individual sensors or sensor arrays.

Flow Alignment

Flow alignment is a method that makes use of flowing fluid to align suspended nanostructures along a single orientation in the direction of flow [23]. By passing a nanostructure suspension through a microfluidic structure formed between a poly(dimethylsiloxane) mold and the device substrate, free-flowing nanostructures assemble onto the device substrate due to surface forces and remain relatively aligned parallel to the direction of flow. The resulting degree of alignment is a function of the flow rate as increased flow velocities result in a narrower angular distribution among deposited nanostructures. Furthermore, varying the total flow duration dictates the nanostructure deposition density, with longer flow durations favoring higher nanostructure densities. This method is useful in controlling the degree of nanostructure alignment and the deposition density, thus allowing subsequent electrode formation. Nano-FET fabrication can therefore be achieved by patterning a pair of electrodes parallel to the flow direction using conventional microfabrication processes. For this physical arrangement, minimizing the electrode pair separation distance increases the likelihood of a bridging nanostructure between the electrode pair; however, the exact number of bridging nanostructures resulting from this process is probabilistic.

Nanostructure Contact Printing
Nanostructure contact printing is a method that provides similar results to the flow alignment strategy. This method, however, relies on mechanical shear forces to detach freestanding nanostructures from the growth substrate, and friction to align the deposited the nanowires onto the device substrate [24]. In this method, the growth substrate is inverted and placed on top of the device substrate such that the nanostructures are sandwiched in between (Fig. 7). Weights are added on top of the growth substrate to control the applied pressure between the two substrates in order to optimize the deposited nanostructure density. The growth substrate is then pulled over the device substrate at a continuous velocity in order to deposit and align the nanostructures onto the device substrate. The shear velocity and the amount of friction between the two substrates influence the resulting degree of nanostructure alignment. The latter may be adjusted with the use of a lubricant.
Nanostructure Field Effect Transistor Biosensors, Fig. 7

Contact printing of nanowires. (a) Schematic of the contact printing process for producing well-aligned nanowire arrays. (b) Dark-field optical and (c) SEM images of 30-nm nanowires printed on a Si/SiO2 substrate showing highly dense and aligned monolayer of nanowires. The self-limiting process limits the transfer of nanowires to a single layer, without significant nanowire bundling (Adapted with permission from [24]. Copyright 2008 American Chemical Society)

Thin-Film Nanowire Suspension
Thin-film nanowire suspension is a method that makes use of the observation that nanostructures suspended within a liquid thin film will align along a single direction when that film is elongated. Therefore, this method involves the formation of a large balloon from a nanostructure suspension (Fig. 8) [25]. The aligned nanostructures within the thin film can then be transferred to a substrate by placing the substrate in contact with the balloon. Nano-FET fabrication would subsequently proceed as in the case of flow alignment.
Nanostructure Field Effect Transistor Biosensors, Fig. 8

Blown bubble film approach. Nanostructures are dispersed in a polymer solution. A volume of solution is expanded as a bubble using a die to produce well-aligned nanostructures suspended within a thin film. The film is then contacted with a substrate to deposit well-aligned nanostructures onto the substrate ([25] Reproduced with permission of The Royal Society of Chemistry)

Directed Self-Assembly
Directed self-assembly relies on the minimization of surface energy to assemble suspended nanostructures into predefined locations on a substrate [26]. In this method, surface modification strategies are used to render certain portions of the device substrate either hydrophobic or hydrophilic (Fig. 9). Upon immersion of the substrate into a nanostructure suspension, individual nanostructures will self-assemble onto the hydrophobic portions of the substrate. This strategy can be used to deposit individual nanostructures at specific locations over a large area on a substrate. A similar strategy using analyte-ligand interactions has also been employed in which suspended nanostructures are functionalized with biotin molecules. Upon immersion of a substrate with bound streptavidin molecules at predefined locations, biotin-streptavidin conjugation forces the nanostructures to assemble onto the device substrate at predefined locations.
Nanostructure Field Effect Transistor Biosensors, Fig. 9

Linker-free directed self-assembly. (a) The patterning process of octadecyltrichlorosilane self-assembled monolayer on a solid substrate via photolithography. (b) Self-assembly suspended nanostructures onto the bare surface regions on the substrate. Subsequent photolithography and metallization can be used to pattern electrodes onto the ends of the deposited nanostructures (Adapted with permission from Macmillan Publishers Ltd. Nature Nanotechnology [26]: copyright 2006)

Dielectrophoresis

Dielectrophoresis (DEP) utilizes electrical fields to manipulate nanostructures. In this method, metallic contact electrodes are patterned on the device substrate prior to the assembly of nanostructures onto the device. By applying a biased alternating voltage across the electrodes, a local nonuniform electric field is produced, exerting a force on suspended semiconducting nanostructures. This force causes the nanostructures to assemble across the two electrodes. This process is easily amenable to the large-scale production of nano-FET arrays and can be used to fabricate single nanostructure FET devices; however, typical DEP trapping processes yield networked multi-nanostructure FET devices [27]. This process is highly susceptible to the experimental conditions including the size, shape, and properties of the nanostructure to be manipulated, and parameters of the electrical signals, as well as the electrical properties of the surrounding medium.

In Situ Growth of Nanostructures

In situ growth of nanostructures is a method that utilizes patterned catalysts to selectively synthesize nanostructures at certain regions on the device substrate. These catalysts (e.g., iron or gold particles) are required to initiate and/or propagate nanostructure growth. In some cases, the direction of nanostructure growth can be influenced with the application of an electric field, therefore enabling control of the resulting nanostructure orientation. This method usually produces numerous nanostructures and is, therefore, useful in fabricating nano-FET sensors that utilize a network of nanostructures.

Top-Down Fabrication Method

Anisotropic Lateral Wet Etching
Top-down methods are typically based on the anisotropic lateral wet etching of nanometer-thin SOI (silicon-on-insulator) wafers to produce well-defined nanostructures from the device layer (Fig. 10) [8]. Micro- or nano-sized etch masks are patterned via conventional or e-beam photolithography and anisotropically time-etched to produce nanometer-wide nanostructures. Source and drain electrodes are degenerately doped, rendering them conductive and unaffected by the etchant. This approach enables wafer-scale formation of semiconducting nanostructures at precise locations on a wafer-scale, making subsequent electrode fabrication relatively straightforward using conventional photolithography. As this method relies on time-controlled anisotropic etching for the removal of material, it is highly susceptible to small changes in processing conditions such as etch time, processing temperature and mixing, and device-layer doping density and crystal orientation.
Nanostructure Field Effect Transistor Biosensors, Fig. 10

(a) Schematic of nano-FET device after anisotropic etching. The silicon-on-insulator active channel (yellow, width w and thickness t) is undercut etched, whereas degenerate leads (red) are etch-resistant. The source (S), drain (D), and underlying backgate (G) are labeled. (b) Scanning electron micrograph of a complete device (Adapted with permission from Macmillan Publishers Ltd: Nature [26], copyright 2007)

Future Directions

At this point in time, the development of nano-FET biosensors is still in the proof-of-concept phase. The promise of highly sensitive, label-free electrical sensors with real-time measurement capability remains attractive for numerous medical and basic science research applications. While great progress has been made in nano-FET fabrication and detection, continued research into large-scale fabrication methods for batch-manufacturing nano-FET sensors will be essential for the ultimate commercial success of this technology and its application to clinical and research applications. In addition to low cost and high yield, these fabrication methods must produce devices with consistent sensing performance across the nano-FET sensors such that measurements made with these devices can be standardized. Another critical hurdle that must be overcome is the issue of poor analyte specificity and nonspecific binding. While some strategies have been put forth to remedy these issues, these solutions must be further improved upon before nano-FET biosensors are able to reliably analyze whole blood, serum, and other specimens. Lastly, efforts in the area of nano-FET biosensor integration with microfluidic and lab-on-chip devices will facilitate point-of-care diagnostic applications and real-time closed-loop drug delivery systems. Successful lab-on-chip integration will also enable the creation of multiplexed nano-FET biosensor arrays, which could hold great promise in the areas of basic chemical and biological research, high-throughput screening systems for drug development, and novel in vitro biology experimentation. The future will likely see many point-of-care biosensors using electrical-based detection, and nano-FET biosensing technologies can possibly play an important role.

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Copyright information

© Springer Science+Business Media Dordrecht 2016

Authors and Affiliations

  1. 1.Advanced Pharmaceutics and Drug Delivery Laboratory, Leslie Dan Faculty of PharmacyUniversity of TorontoTorontoCanada
  2. 2.Department of Electrical and Computer EngineeringUniversity of TorontoTorontoCanada
  3. 3.Department of Mechanical and Industrial Engineering and Institute of Biomaterials and Biomedical Engineering and Department of Electrical and Computer EngineeringUniversity of TorontoTorontoCanada
  4. 4.Department of Mechanical and Industrial Engineering and Institute of Biomaterials and Biomedical EngineeringUniversity of TorontoTorontoCanada