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WIREs Nanomed Nanobiotechnol
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Performance limitations for nanowire/nanoribbon biosensors

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Field‐effect transistor‐based biosensors (bioFETs) have shown great promise in the field of fast, ultra‐sensitive, label‐free detection of biomolecules. Reliability and accuracy, when trying to measure small concentrations, is of paramount importance for the translation of these research devices into the clinical setting. Our knowledge and experience with these sensors has reached a stage where we are able to identify three main aspects of bioFET sensing that currently limit their applications. By considering the intrinsic device noise as a limitation to the smallest measurable signal, we show how various parameters, processing steps and surface modifications, affect the limit of detection. We also introduce the signal‐to‐noise ratio of bioFETs as a universal performance metric, which allows us to gain better insight into the design of more sensitive devices. Another aspect that places a limit on the performance of bioFETs is screening by the electrolyte environment, which reduces the signal that could be potentially measured. Alternative functionalization and detection schemes that could enable the use of these charge‐based sensors in physiological conditions are highlighted. Finally, the binding kinetics of the receptor–analyte system are considered, both in the context of extracting information about molecular interactions using the bioFET sensor platform and as a fundamental limitation to the number of molecules that bind to the sensor surface at steady‐state conditions and to the signal that is generated. Some strategies to overcome these limitations are also proposed. Taken together, these performance‐limiting issues, if solved, would bring bioFET sensors closer to clinical applications. WIREs Nanomed Nanobiotechnol 2013, 5:629–645. doi: 10.1002/wnan.1227 This article is categorized under: Diagnostic Tools > Biosensing Diagnostic Tools > Diagnostic Nanodevices

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Wafer map showing the distribution of threshold voltages for various devices at different positions on a 4″ silicon‐on‐insulator (SOI) wafer. The tight control over the threshold voltage (indicated by the dashed lines) means that devices can be easily multiplexed and biased at the same operating point using a global gating scheme. (Reprinted with permission from Ref . Copyright 2011 Elsevier Ltd)
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Real‐time sensor responses of HMGB1–DNA binding. Each curve represents the measurement of a different DNA concentration from the same device, and sensor responses are plotted using (Ids − I0)/gm. Dashed lines represent the fits obtained from the fitted values of the association and dissociation rates.
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Theoretical plot of the surface coverage as a function of analyte concentration for different equilibrium dissociation constant (KD) values. For a certain charge of the analyte, if the detection limit occurs at 40% of surface coverage (dashed line), the minimum detectable concentration of analyte (ρmin) depends on the KD of the system.
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Schematic of nanoelectronic enzyme‐linked immunosorbent assay (ELISA) showing the device surface as unmodified, with the gold leads functionalized with an ELISA assay. The assay terminates with the enzyme urease, which breaks down urea and consumes protons (H+) in the reaction process. The subsequent change in pH is then detected by the charging of the hydroxyl groups on the nanowire surface, leading to a change in the conductance.
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(a) Primary antibodies to multiple biomarkers, here prostate‐specific antigen (PSA) and cancer antigen 15.3 (CA15.3), are bound with a photocleavable crosslinker to the MPC. The chip is placed in a plastic housing and a valve (pink) directs fluid flow exiting the chip to either a waste receptacle or the nanosensor chip. (b) Whole blood is injected into the chip with the valve set to the waste compartment (black arrow shows the direction of fluid flow) and, if present in the sample, biomarkers bind their cognate antibodies. (c) Washing steps follow blood flow, and the chip volume (5 µl) is filled with sensing buffer before UV irradiation (orange arrows). During UV exposure, the photolabile crosslinker cleaves, releasing the antibody–antigen complexes into solution. (d) The valve is set to the nanosensor reservoir (black arrow shows the direction of fluid flow) and the 5 µl volume is transferred, enabling label‐free sensing to be performed to determine the presence of specific biomarkers.
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(a) Normalized current noise power density at f = 1 Hz is plotted against drain current. The noise profile does not change significantly with changes in PBS (phosphate‐buffered saline) concentration or by changing the electrolyte to KCl (potassium chloride). The proportionality to (gm/Id)2 confirms that the data are well fitted by the number fluctuation model. (b) Signal‐to‐noise ratio (SNR) is plotted against solution gate voltage to highlight the regime at which SNR is maximized. (c) Transconductance values extracted from I–V measurements are also plotted against solution gate voltage to point out that maximum SNR occurs close to the point of peak transconductance.
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The response of silicon nanowire field‐effect transistor‐based biosensors (bioFETs) for the detection of the antigen Troponin T as a function of ionic strength for devices functionalized with full antibodies (red), the double Fab regions (black), and the single Fab domain (blue). Inset shows real‐time sensing data.
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The schematic outlines the generation of only part of the antibody (Fab fragment) by using the enzyme pepsin. Mercaptoethylamine (MEA) can be used to further cleave the disulfide bonds between the two Fab fragments. The size of the receptor is thus reduced while the binding sites remain functional, resulting in bound molecules being closer to the sensor surface and resulting in an enhanced signal.
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(a) Measured Hooge parameters for three sets of devices. Each set was etched using either tetramethylammonium hydroxide (TMAH) or Cl2 or CF4. The box plot shows the 25th percentile, the median, and the 75th percentile (the mean is indicated by a square marker). The average values of αH were 0.0021 for the TMAH devices, 0.015 for the Cl2 devices, and 0.017 for the CF4‐etched devices. (b) Measured subthreshold swing for three sets of devices, etched using either TMAH or Cl2 or CF4. The average value for the TMAH devices was 1.0 V/decade, Cl2‐etched devices was 2.6 V/decade, and CF4 devices was 3.0 V/decade.
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Summary of Hooge's parameter (αH) for different nanowire materials as well as submicron MOS structures utilizing high‐k dielectrics. Included is our best silicon‐on‐insulator (SOI) nanowire device (red circle). The dash‐dotted line shows the International Technology Roadmap for Semiconductors (ITRS) specification of αH for the 45‐nm technology node.
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Signal‐to‐noise ratio (SNR) of single‐walled carbon nanotubes (SWNTs) plotted as a function of solution gate bias for bare devices and PMMA passivated devices. A significant degradation of the SNR is observed for the PMMA passivated SWNT devices.
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(a) Comparison of the normalized current noise power density for indium oxide field‐effect transistors (FETs), showing that with a surface passivation of octadecanethiol (ODT), the noise is significantly reduced. This is attributed to the passivation of surface states/defects. (b) Transfer characteristics of the unpassivated and ODT‐functionalized devices, showing a threshold voltage shift and negligible degradation of the transconductance.
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Comparison of the normalized current noise power density for silicon nanowire field‐effect transistors (FETs), showing that as the channel diameter is reduced, the normalized noise power also decreased. This reduction in noise is attributed to a volume inversion effect, where the conducting channel is formed in the middle of the nanowire cross‐section, resulting in less surface scattering.
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(a) Calibration curve for prostate‐specific antigen (PSA) detection at different concentrations, including a blind measurement, showing accuracy of the calibration. (b) Similar calibration curve for cancer antigen 15.3 (CA15.3) detection. (c) Normalized response in conductance for different concentrations of PSA and for different sizes of devices. (d) Change in conductance for various glucose concentrations, focusing on the linear region of the calibration curve. Inset in the lower right shows the full response curve.
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