Accepted Manuscript Trends in Electrochemical Impedance Spectroscopy involving nanocomposite transducers: characterization, architecture surface and bio-sensing Jose Muñoz, Raquel Montes, Mireia Baeza PII:
S0165-9936(16)30287-4
DOI:
10.1016/j.trac.2017.08.012
Reference:
TRAC 14987
To appear in:
Trends in Analytical Chemistry
Received Date: 14 September 2016 Revised Date:
31 July 2017
Accepted Date: 17 August 2017
Please cite this article as: J. Muñoz, R. Montes, M. Baeza, Trends in Electrochemical Impedance Spectroscopy involving nanocomposite transducers: characterization, architecture surface and biosensing, Trends in Analytical Chemistry (2017), doi: 10.1016/j.trac.2017.08.012. This is a PDF file of an unedited manuscript that has been accepted for publication. As a service to our customers we are providing this early version of the manuscript. The manuscript will undergo copyediting, typesetting, and review of the resulting proof before it is published in its final form. Please note that during the production process errors may be discovered which could affect the content, and all legal disclaimers that apply to the journal pertain.
ACCEPTED MANUSCRIPT Trends in Electrochemical Impedance Spectroscopy involving nanocomposite transducers: characterization, architecture surface and bio-sensing Jose Muñoz,a* Raquel Montes,b and Mireia Baezac* a
Molecular Nanoscience and Organic Materials Group, Institut de Ciència de Materials de Barcelona (ICMAB-CSIC), Campus
UAB, 08193 Cerdanyola del Vallès (Bellaterra), Spain. b
UAB, 08193 Cerdanyola del Vallès (Bellaterra), Spain. c
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Departament d’Enginyeria Química, Biològica i Ambiental, Escola d’Enginyeria, Universitat Autònoma de Barcelona, Campus
Departament de Química, Facultat de Ciències, Entrada C-Nord, Universitat Autònoma de Barcelona, Campus UAB, 08193
Cerdanyola del Vallès (Bellaterra), Spain.
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*Corresponding author:
[email protected],
[email protected]
Abstract
Electrochemical Impedance Spectroscopy (EIS) has gained widespread application for the characterization of functionalized electrode surfaces and for the transduction of bio-
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sensing events. However, bio-sensors using EIS detection have to be carefully designed to minimize non-specific binding of the analyte. In this sense, surface engineering by using nanocomposite materials (NCs) is advantageous due to the increased electrode
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surface area, improved electrical conductivity of the sensing interface, chemical accessibility to the analyte and electroanalysis. Accordingly, this review summarizes the
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basis of the EIS technique as well as its implementation not only in common Faradaic EIS (impedimetric) bio-sensors using NCs as highly sensitive transducer platforms but also in not so conventional non-Faradaic EIS (capacitive) approaches. Finally, it is also highlighted the feasibility of EIS as an alternative characterization tool towards the optimization of NC electrodes in terms of loading ratios for electroanalytical improvements, summarizing the latest promising results in nanocomposite carbon paste electrodes.
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ACCEPTED MANUSCRIPT Keywords: Electrochemical Capacitance Spectroscopy; Carbon Paste Electrodes; SelfAssembled Monolayers; Bio-sensors; Nanomaterials
1. Introduction
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Electrochemical Impedance Spectroscopy (EIS) is a powerful method of analyzing the complex electrical resistance of a system and is sensitive to surface phenomena and changes of bulk properties. The term impedance was coined in 1886 by the electrical
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engineer, mathematician and physicist Oliver Heaviside, who adapted complex numbers to the study of electrical circuits [1]. The method of impedance measurements is widely
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used in many fields of electrochemistry, such as electrode kinetics, double layers studies, batteries, corrosion, solid-state electrochemistry and bio-sensing [2], [3], [4]. EIS technique has become a popular electrochemical tool for the detection of a wide range of chemical and biological targets in terms of electrochemical change of
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electrode interfaces, as it is revealed by the rapidly number increased of bio-sensors based on impedimetric measurements in the last years [5], [6], [7]. The use of EIS transducer signal in biological sensorial applications is mainly based on the interactions
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between the biological receptor and the target species selectively adsorbed from the solution. Such interactions cause a change on interfacial electron transfer kinetics
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between a redox probe in solution and the conducting electrode sites. This electrochemical change is then detectable by monitoring the charge-transfer resistance (Rct) that commonly increase in the same proportion given the increase in the quantity of targets bound to the receptive surface [8], [9]. However, while the selectivity of a bio-sensor mostly relies on the specificity of the molecular recognition element, several analytical parameters, such as sensitivity, response time and limit of detection, strongly
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ACCEPTED MANUSCRIPT depend on the physicochemical properties of the transducer, which can be improved by the utilization of proper materials and/or the design of new architecture surfaces. In this sense, functional nanomaterials, and especially nanocomposite (NC) materials, play a critical role in signal transduction since such nanosized compounds are
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shown to clearly increase the analytical performance [10], [11], [12], [13] providing not only a synergetic effect among catalytic activity, conductivity, and biocompatibility to accelerate the signal transduction but also amplify bio-recognition events with
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specifically designed signal tags. These improvements are mainly related to an increased surface area that enhances accessibility for the analyte to the receptor unit;
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this fact leads to the achievement of highly sensitive bio-sensing platforms [14], [15], [16]. In general, a NC material can be defined as the result of combining two or more different materials where, at least, one of the constituent parts has a nanometer scale dimension [17], keeping their individual identities but still impart their properties to the
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product resulting from their combination. Thus, building blocks with dimensions in the nanosize range enable us the possibility to design and develop new NCs with unprecedented versatility and improvement in their physical, chemical and
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electrochemical properties, being also able to act as transducers for the signal capture. Among the vast amount of examples where nanomaterials demonstrate their superiority
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to bulk materials, the development of NC materials based on a conductive phase as carbon nanoallotropes (e.g., carbon nanotubes, graphene and related) dispersed in an insulating polymeric matrix (such as epoxy, methacrylate, Teflon, etc.), has led to important advances in the analytical electrochemistry field, especially in the development of sensors and biosensors devices [18], [19], [20]. The strong demand to generate devices that respond selectively to specific biomolecules opens up alternative strategies associated with designing either Faradaic or
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ACCEPTED MANUSCRIPT non-Faradaic bio-interfaces, due to each approach needs specific surface architecture according to either resistive or capacitive transducer signal, respectively [21], [22], [23]. Regarding that, self-assembled monolayers (SAMs) have found widespread utility in the fabrication of electrochemical bio-sensors [24], [25] owing to the mainstream
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development of gold-thiol SAMs in the early 1980s [26]. In contrast to traditional Faradaic EIS (impedimetric) methods where the determination of Rct require the addition of a redox probe to the analytical solution prior to analysis and a subsequent
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fitting of data to an equivalent circuit, non-Faradaic (capacitive) methods provide an alternative strategy where no redox element is added to solution, being amenable to
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miniaturization [27], [28]. The signal of non-Faradaic EIS transducers is mainly due to capacitance changes on the electrode–electrolyte interface that can be easily monitored by double-layer capacitance (Cdl) means. Further, non-Faradaic technique can not only be applied in traditional analysis of dielectric films [29], [30] but also can be combined
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with surfaced tethered redox reporters in establishing highly sensitive label free assay based on Electrochemical Capacitance Spectroscopy (ECS), which measures “capacitive charging” (redox capacitance) [27], [31].
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Recently, EIS technique has also been exploited as promising characterization tool towards the optimization of NC-based electrodes by means of composition ratios;
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especially in the nano(bio)composite carbon paste electrodes field [32], [33]. Briefly, both the nature of the raw carbon filler and its loading throughout the insulating polymeric matrix directly affects the rate of electron transfer, the material stability and the background capacitance current [34], [35]. Therefore, an accurate loading characterization of each component is crucial in order to achieve high signal-to-noise ratios which ensure significant reproducibility, optimal sensitivity, fast response time as well as low detection limits.
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ACCEPTED MANUSCRIPT Accordingly, the framework of the current review tends to offer a comprehensive summary of the last trends in EIS approaches where NC materials are involved as highly sensitive transducer platforms, fact that might serve as a general outlook for planning further research. For this aim, this review overviews (i) the basis of the EIS
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technique; (ii) the latest development of EIS bio-sensors, which usually focus upon certain bio-recognition events in conjunction with NC-based electrodes; (iii) the demands and strategies associated with surface engineering for either Faradaic
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(impedimetric) or non-Faradaic (capacitive) EIS transducer development, highlighting those platforms based on SAMs; and (iv) the use of EIS technique as an alternative
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characterization tool for NC electrodes optimization (especially for NC-CPEs), fact that allows for an improvement on their electrochemical and electroanalytical performances.
2. General principles
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EIS technique provides electric information in the frequency domain. With this technique, a process occurring in an electrochemical cell can be modeled to equivalent circuits derived from using combination of resistors (R), capacitors (C) and/or inductors
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(L). Interpretation of EIS measurements is usually done by fitting the impedance data to an equivalent electrical circuit that is representative of the physical processes taking
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place in the system under investigation. By using equivalent circuits, the experimental spectrum can be fitted with the theoretical curve corresponding to the suitable circuit model, thus obtaining the values of electrical parameters [36], [37], [38], [39]. Impedance is defined as the ability of a circuit element to resist the flow of current and therefore, is simply the opposition force to electrical current in a circuit, which is measured in the same units as resistance (Ω). EIS generally measures the resulting of applying a sinusoidal alternating current (AC) potential to an electrochemical cell and
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ACCEPTED MANUSCRIPT measuring the AC current that crosses through the cell. When an oscillating potential is applied to an electrode surface, the resulting current has the same frequency as the applied potential but may be shifted in phase. This phase-shift is dependent on the relative resistive and capacitative features of the electrochemical system. Consequently,
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the phase-shift can be used to monitor physical processes at the electrode surface. As it is shown in Figure 1A, a sinusoidal AC voltage, E(t), of a small magnitude, Eº, is applied to an electrode over a range of frequencies (ω), and the resulting sinusoidal
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current is measured. The resulting AC current, I(t), is of the same frequency as the applied potential, but shifted in phase. The phase-shift is reported as the phase angle (θ)
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as it relates to the period of the sinusoidal waves.
This phase angle (or phase-shift) represents the relative capacitive, resistive and inductive character of the electrochemical system. Equations 1 and 2 describe the applied potential E(t) and resulting current I(t), respectively, where t is time and Iº is
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the magnitude of the current oscillations [3], [40]. E(t) = Eº sin (ωt)
Equation 1
I(t) = Iº sin (ωt – θ)
Equation 2
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The impedance of the system (Z) can be calculated by means of an expression
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analogous to Ohm’s Law, Equation 3: Z=
() ()
Equation 3
Using the Euler’s formula (Equation 4), and taking into account that the impedance
has an amplitude Zº = Eº / Iº and a phase-shift φ, it is possible to express Z in terms of complex numbers, following Equation 5, where Zreal is the real part, j = √−1 and Zimag is the imaginary part of the impedance. ejφ = cos (φ) + j sin (φ)
Equation 4
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ACCEPTED MANUSCRIPT Z = Zº [cos (φ) + sin (φ)] = Zreal + j Zimag
Equation 5
There are different ways to plot impedance. The most often used in the electrochemical literature when studying electron-transfer kinetics is the Nyquist diagrams because they allow for an easy prediction of the circuits elements. Nyquist
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diagrams, also known as a Cole-Cole plot or a complex impedance plane plot [41], [42], [43], consist of plotting Zimag as a function of Zreal, usually showing a semicircle profile and provide visual insight into system dynamics at the electrochemical interface. Every
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experimental point in the plot corresponds to a different frequency. However, Nyquist plot format has some disadvantages. For example, frequency does not appear explicitly
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and the electrode capacitance cannot be calculated until the frequency information is known. Thus, Bode plot (|Z| and φ vs. log ω) is an efficient alternative for evaluating how the impedance depends on the frequency since it appears as one of the axes. For further information on Bode plots, see Scully and Silverman [44].
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Therefore, usually EIS are analyzed both Nyquist and Bode plots. As was mentioned above, EIS data is commonly analyzed by fitting it to an equivalent electrical circuit model, to simulate the real impedance spectra, combining the resistor (R), which
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represents the resistor that charge carriers encounter in a specific process or material, the capacitor (C), which represents the accumulation of charged species and/or the
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inductance (L), which is used to represent the deposition of surface layers such as the passive layer. However, the major problem in using equivalent circuits resides in deciding which specific equivalent circuit out of an infinite number of possibilities should be chosen. The most important elements that can be used in equivalent circuits are: i) the solution resistance (RΩ) which is dependent on the ionic concentration, the type of ions and also the electrode area; ii) the charge-transfer resistance (Rct), which is inversely proportional to the electron transfer rate; iii) the double-layer capacitance
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ACCEPTED MANUSCRIPT (Cdl), which is directly related to the charging and background current; iv) the Warburg impedance (ZW), which arises from mass-transfer limitations and can be used to measure effective diffusion coefficients; and v) the constant phase element (CPE or Q), which is a general element which can represent a variety of elements such as inductance
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(n= –1), resistance (n= 0), Warbug (n= 0.5), capacitance (n= 1) or non-ideal dielectric behavior (–1 ≤ n ≤ 1) resulting from a distribution of relaxation times [45]. Thus, the impedance spectra allow a broad overview of the different processes taking place at the
dominating more at specific range of frequencies.
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FIGURE 1
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electrochemical interface (i.e., capacitive, resistive, diffusion effects) and which one is
It is important to highlight at this point that EIS experiments can be carried out by two different types of investigation: Faradaic and non-Faradaic processes. Faradaic
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processes are made utilizing redox markers and hence, particular importance is given to the Rct. Otherwise, since redox markers are not exploited by non-Faradaic systems, the
[46].
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analysis of the double layer capacitance Cdl is much more useful rather than Rct [30],
An electrochemical reaction on an electrode surface (Faradaic approaches) can
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occur by two limiting mechanisms: kinetically controlled (electronic transfer) and diffusion controlled (mass transfer). When the rate at which electrons can transfer across the interface is infinitely fast, the mass transport is the limitation process and it is said that the phenomenon is controlled by the diffusion. However, if the rate at which material gets from the bulk of solution to the electrode is extremely fast, the charge transfer kinetics is then the limitation process and therefore, the phenomenon is controlled by the kinetics.
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ACCEPTED MANUSCRIPT The Randles circuit, RΩ· (Rct·Cdl), is the simplest and most common electrical representation of an electrochemical system with kinetic-control process (see Figure 1B, inset). If the impedance spectrum is dominated by the semicircle feature, the electrochemical system is limited by the electron-transfer rate and it means that the
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electrochemical process is kinetically controlled. This indicates that the Nyquist plot for a Randles cell is a semicircle with two intercepts on the real axis in the high and low frequency regions, as it is depicted in Figure 1B. Herein, the beginning of the
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semicircle corresponds to the RΩ and the diameter of the semicircle corresponds to the Rct. In addition, the angular frequency is equal to the reciprocal of Rct·Cdl at the
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minimum value of Zimag. Nonetheless, the capacitor (Cdl) behavior frequently deviates from ideality mainly due to the roughness effect. In this case, different studies have shown that the distributed nature of the electrochemical double layer is better modeled with the electrical pseudo-element with constant phase known as CPE (or Q) [47], [48].
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Note that this RΩ· (Rct·Q) equivalent circuit introduces a depression of the semicircle (see Figure 1B, dotted line). But, such plot normally exhibits a kind of semicircle profile plus a linear region, demonstrating that Faradaic currents are affected by
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diffusion processes and therefore, the electron transfer is fast and the diffusion of the electroactive species is the limiting factor. In this case, it is said that the process is
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diffusion controlled. These diffusion processes can be modeled in a modified Randles equivalent circuit, noted as RΩ· (Cdl·[Rct·ZW]), by using a distributed element (ZW) in series with the charge transfer resistance (see Figure 1C, inset). Since the Warburg element takes into account the mass transfer processes, it is more predominant at lower frequencies. Figure 1C represents the ideal impedance spectra from an electrochemical system with mixed kinetic- and diffusion-control process [49].
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ACCEPTED MANUSCRIPT Regarding to EIS approaches based on non-Faradaic processes, Cdl calculation is based on the premise that both charge transfer resistance and diffusion contributions can be omitted, becoming then the electrical circuit as a mere [RΩ· Cdl], as it is shown in the inset from Figure 1D [50]. Considering that a generic EIS experiment can be expressed
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according to Equation 5 and taking into account that a complex capacitance can be expressed following the Equation 6, the capacitance of the electrochemical system in such cases results from combining both equations (Equation 7):
Zimag
2 –
ω|Z|
Zreal
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C = –
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C=
2 = Creal + j Cimag
ω|Z|
Equation 6
Equation 7
Then, the complex capacitance can be plotted as the common Nyquist plot, replacing Zreal and Zimag for Creal and Cimag, respectively. In contrast to traditional
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Faradaic EIS methods, this non-Faradaic-based EIS approach presents the benefit that no redox element is added to solution and therefore, no equivalent circuit is necessarily considered or fitted to, being simpler to calculate the total capacitance value at the real
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part of the capacitance from the relative minimum value of the imaginary part of
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capacitance, as it is indicated in Figure 1D (inset), [51], [52]. Since the physical interpretation of the distributed elements in an equivalent circuit
modeling solid electrodes is often somewhat more complicated than predicted by the Randles model, further information about other more specific equivalent circuit can be found in the following references: [43], [45], [53], [54], [55], [56], [57].
3. Trends in EIS nanocomposite bio-sensors
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ACCEPTED MANUSCRIPT EIS has been used since decades in non-biological applications such as corrosion detection, electrochemical activity of lithium ion cells or monitoring fuel cell performance as well as rudimentary characterization technique for confirming, in particular, the layer by layer fabrication of the sensors Nonetheless, EIS has gained
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recently much more popularity for bio-sensing a wide range of analytes owing to its label-free measurements and superb sensitivities, sometimes reaching down to the femto and attomolar ranges [58], [59], [60]. Common label based techniques are characterized
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by several disadvantages, such as time and cost consuming process, not real time measurements and common interaction problems between the probe and target
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molecules. Further, the fact that the label fluorescence signal must be measured by costly and bulky systems, their miniaturization is hindered. Thus, EIS represents one of the most promising label-free measurement techniques since greatly simplify the time and effort required for assay development, whereas removing experimental artefacts
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from quenching, shelf life, and background fluorescence [61], [62], [63]. Despite this, few drawbacks of label-free techniques may include low sensitivity and lack of specificity [14], [64], [30].
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EIS experiments for bio-sensing analysis have been traditionally performed under Faradaic conditions. Thus, the working electrochemical cell must contain both
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reduced and oxidized forms of a benchmark redox probe (i.e., anionic [Fe(CN)6]3-/4-, cationic [Ru(NH3)6]2+/3+ or neutral [Fc(MeOH)0/+] markers) that undergo redox processes at an electrode interface, being responsible of generating the well-known Faradaic currents [65]. The vast majority of the reported bio-sensing systems summarized in the present section have assumed that Faradaic current is larger higher than background current in order to properly quantify the chemical reactions by means of Faradaic currents. In this sense, Faradaic EIS is an effective method for probing bio-
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ACCEPTED MANUSCRIPT molecules binding events such as antibody-antigen interactions via supramolecular complex formation on the electrode surface. The typical Faradaic EIS signal transduction resulting from the adsorption of such bio-molecules on the electrode surfaces consists of recording the variations in the electrical current generated by the
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redox probes. These variations are mainly derived from i) the obstructing effect of the analyte after bio-affinity complex formation and/or ii) the electrostatic barrier of the electrode surface that repeals its interfacial electron transfer reaction. These phenomena
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commonly lead to an insulating layer formation that retards the interfacial electron transfer kinetics between the redox probes and the electrode due to the blocking effect
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generated by the “bulky” properties of the analyte and/or because of an increased electrostatic barrier that impedes the redox probes to approach the electrode surface. From an electrochemical point of view, both facts generate an increase on the Rct value. Some works have described alternative label-free bioelectronic strategies for
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transducting different biological recognition events based on switching from repulsion to attraction the redox marker diffusion via reversing the surface charge. Reductions in the impedance capability by Rct means (rather than the common increases) are attributed
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to a decrease of the electrostatic barrier over the sensor surface that facilitates access of the marker and therefore, its redox reaction [66], [67], [68]. Additionally, much more
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complicated situations can be found when signal changes are due to a combination of different effects [69], [70], [71]. The use of nanomaterials clearly already enhances the signal capture of EIS
transduction for bio-sensing due to their improved specific surface area which allows the immobilization of a larger amount of (bio)receptor units and therefore, the accessibility for the analytes. The emergence of NC materials has opened up enormous opportunities for tailoring the immobilization efficiency of a wide range of bio-
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ACCEPTED MANUSCRIPT recognition elements in/on/into the transducer platform [72], [73]. Given that the use of NC materials based on carbon nanoallotropes in modern electroanalytical chemistry [74], [75], [76], [77], [78], [79], [80], this section focuses mainly upon this approach in tandem with EIS for biological and chemical applications, resulting in different types of
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label-free bio-sensors, such as genosensors, immunosensors, aptasensors, cytosensors and enzymatic biosensors as well as some non-biological chemical-based sensors.
Genosensors: Recent advances in the electrochemical detection of nucleic acids
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have allowed the development of new types of genosensors. Faradaic EIS has been proven to be a most powerful and sensitive tool for probing the features of surface-
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modified electrodes and widely applied to the detection of target DNA sequences [81]. Zhu et al. recently reported a highly sensitive impedimetric biosensor based on bismuth sulfide nano (rBi2S3) and polyaniline nanocomposite film modified ionic liquid-carbon paste electrode [82]. In this work, EIS technique was used as an effective and sensitive
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tool to measure the impedance on the electrode surface during the immobilization of rBi2S3. Moreover, the hybridization performance using different DNA sequence of the constructed nanocomposite biosensor was also evaluated by EIS. It was found a
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remarkable enhancement of the impedance after the probe DNA modified electrode was hybridized with the target DNA, resulting in a detection limit of 4.37·10-16 M. Luo et al.
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proposed a novel simple protocol for preparing 1-aminopyrene/graphene (ApG) hybrids for fabricating label-free electrochemical impedance genosensor [83]. The DNA hybridization reaction of oligonucleotide probe with target DNA was monitored by EIS. By means of EIS technique it was possible to evaluate the changes on the charge transfer resistance due to the electrode modification from the bare electrode (GCE) until the incorporation of the complementary ssDNA (CDNA). From the Nyquist diagrams, compared to the bare electrode GCE, Ret value of GAApG/GCE increases from 68.1 Ω
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ACCEPTED MANUSCRIPT to 465.3 Ω revealing that GAApG has been successfully fixed on the surface of GCE. After covalently immobilized of probe ssDNA (PDNA) on the electrode surface, Ret of PDNA/GAApG/GCE continually increased, reaching a value of 727.3 Ω. When PDNA/GAApG/GCE has been hybridized with 10-9 M CDNA, the Rct of
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CDNA/PDNA/GAApG/GCE is further increased to 1059.6 Ω. After the good results in the CDNA immobilization, it was determined the DNA hybridization reaction of oligonucleotide probe with target DNA was monitored by EIS. Under optimum
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conditions, the proposed biosensor exhibited high sensitivity and a low detection limit for detecting the complementary oligonucleotide. The target oligonucleotide could be
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quantified in a wide range of 1.0·10-12 to 1.0·10-8 M with good linearity (R = 0.99) and low detection limit of 4.5·10-13 M (S/N = 3). Additionally, Bonanni et al. developed a novel impedimetric genosensor based on gold nanoparticles graphite–epoxy nanocomposite for the detection of triple mutation deletion in a cystic-fibrosis related
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human DNA sequence [84]. For first time, EIS technique was used for both characterizing the platform and DNA polymorphism detection correlated to the development of cystic fibrosis. Such platform allowed improving the detection limit of
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DNA polymorphism down to 2.25 fmol.
Immunosensors: Recently, the potential application of EIS technique to the
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immunosensors for a broad set of applications has been examined. EIS technique was found to be the best electrochemical tool to determine different bioanalytes, for example the immunoglobulin G (IgG). In this sense, Derkus et al. described a novel highly sensitive electrochemical impedimetric Protein A (PrA) nanocomposite immunosensor for the determination of IgG [85], which was developed by immobilization of PrA within a newly synthesized poly(maleicanhydride-alt-decene-1) polymer. In order to increase the sensitivity of the immunosensor, TiO2 nanoparticles were synthesized,
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ACCEPTED MANUSCRIPT resulting in better conducting properties than the commercial Degussa P25 TiO2 nanoparticles. The variation of the relative impedance (∆Rct = Rct (i) – Rct (0) / Rct (0)) was used to evaluate the interaction of immobilized PrA with the antibody, where Rct (0) is the charge-transfer resistance when PrA is immobilized on the electrode surface, and
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Rct(i) is the value of the charge transfer resistance after binding of antibody to PrA. The designed poly(MA-alt-D-1)-TiO2-PrA impedimetric immunosensor was capable to determine IgG with a low detection limit (0.57 ng·mL-1). On the other hand, Yu et al.
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proposed a sensitive and convenient EIS method for determining Aflatoxin B1 (AFB1), a toxin implicated in the etiology of human hepatocellular carcinoma, by using
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MWCNTs/RTIL nanocomposite films-based immunosensor [86]. Authors used EIS as an alternative tool compared to the conventional techniques such as high-performance liquid for investigating the electron-transfer properties of the modified electrodes. An increase on the Rct value was observed after the immobilization of the MWCNTs/RTIL-
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Ab on the GCE electrode (from 151.9 Ω to 605.6 Ω). A direct relationship between the increase of the Rct value after AFB1 immobilization and the amount of AFB1 was observed. This fact is caused by electrically insulating bioconjugates produced from
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specific interaction of AFB1 and Ab, which will block the electron-transfer process of the redox probe. The calibration curve for AFB1 was linear in the range of 0.1–10
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ng·mL-1 with 0.03 ng·mL-1 as the detection limit. The presence of MWCNTs warrant fast electron transfer, and the ionic liquid provides a benign microenvironment for antibody.
Aptasensors: EIS is also an ultrasensitive analytical technique that can rapidly detects small changes in reactive surfaces by a transducer and enlarged by an amplifier, making EIS the most suitable detection methods for aptamer-based biosensors. Aptamers possess several competitive advantages over antibodies, including accurate
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ACCEPTED MANUSCRIPT and reproducible chemical production. They are also stable in extreme conditions, and easy to be modified with functional groups. Thanks to these advantages, numerous aptasensors have been devised for the detection of a wide range of analytes targets. A highly sensitive and label-free EIS aptasensor for the detection of adenosine was
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fabricated by co-assembling thiolated aptamer, dithiothreitol (DTT) and 6mercaptohexanol (MCH) on gold electrode surface, forming Au/aptamer-DTT/MCH, reported by Wang et al [87]. EIS technique was performed in order to characterize the
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fabrication process of the proposed aptasensor. The impedance spectra of the different modification steps showed an increase on the electron transfer resistance which
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indicates the successful on the immobilization steps. Moreover, EIS was also used for biosensing adenoside with the proposed aptasensor. The interfacial electron transfer resistance of the aptasensor using [Fe(CN)6]3-/4- as the probe increased with adenosine concentration, and ∆Rct against the logarithm of adenosine concentration was linear
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over the range from 0.05 pM to 17 pM with a detection limit of 0.02 pM. Comparing this detection limit with the ones obtained by those aptasensors containing MCH or DTT alone as the backfiller (0.03 nM and 0.2 pM for Au/aptamer/MCH and
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Au/aptamer-DTT, respectively), the synergism environment derived from both materials dramatically improved the detection limits. Jia et al. used a nanocomposite aptasensor
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platform for the ultrasensitive detection of Staphylococcus aureus (S. areus) [88]. The novel nanocomposite, which increased electron transfer and electrochemical signal, was synthesized by introducing ssDNA to reduced graphene oxide (rGO) containing gold nanoparticles (AuNPs) on a GCE surface. By attaching S. aureus aptamer to the surface of rGO-ssDNA-AuNPs/GCE, it was achieved the direct detection of S. aureus in food samples without using bio-label. EIS was used to monitor the single steps of the electrode assembly process measuring the changes on the Rct. By means of this
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ACCEPTED MANUSCRIPT characterization, it was demonstrated an enhancement on the electronic transmission capability comparing the GCE electrode with rGO-ssDNA-AuNP nanocomposite, due to there is a decrease on the Rct value. The EIS response of the aptasensor to S. aureus in concentration of 105 cfu·mL-1 was checked, obtaining a RSD of 4.3%, which
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indicated that the nanocomposite biosensor presents excellent repeatability. Another example was reported by Ocaña et al. [89]. This work reports a sensitive aptamer– antibody interaction based assay for cytochrome c (Cyt c) using EIS. The employed
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transducer consisted on a multi-walled carbon nanotube (MWCNT) screen-printed electrode which surface allowed the immobilization of aptamer binding cytochrome c
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(APtCyt c) by covalent bond via prior electrochemical grafting. The interfacial Rct change experienced by redox marker was recorded to confirm the formation of a complex between aptamer and the target Cyt c, obtaining 12 pM as detection limit for Cyt c. The suitability of the nanocomposite biosensor for measurement in real samples
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was also tested by determining Cyt c in human serum samples, obtaining recovery values in the range of 94.6-107.8%. Simultaneously, Fan et al. reported the immobilization of the aptamer onto a gold electrode modified with gold nanoparticles
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[90]. The resulting formation of an acetamiprid–aptamer complex was evaluated following the Rct increasing, yielding a low detection limit of 1·10-9 M. The
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applicability of this aptasensor was also successfully evaluated by determining acetamiprid in tomatoes, achieving recovery values from 85.5% to 105.1% and RSD values lower than 5%.
Cytosensors: With the increasing cancer incidence and stubbornly high cancer mortality around the world, the early diagnosis and point-of-care monitoring of cancer therapy have drawn great attention. EIS technique has attached great interest due to their remarkably advantages. EIS is an effective tool to sensitively and rapidly monitor the
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ACCEPTED MANUSCRIPT minute changes of the electrode interface. Second, unlike large amplitude perturbation techniques (such as cyclic voltammetry and differential pulse voltammetry), EIS is a nondestructive technique due to the small amplitude perturbation. Moreover, it is not needed to introduce a label during impedance measurement. Therefore, by collecting
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EIS data of electrode interface at direct current the label-free electrochemical impedance cytosensors have been increasingly explored for clinical diagnosis of tumor, due to the inherent advantages such as label-free procedure, simple operation, high
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sensitivity, short analysis time, miniaturization, and potential of automation. In this sense, Han et al. reported label-free EIS cytosensor for cancer cells was developed by
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immobilizing the specific octapeptide-fused phages on the electrode surface [91]. The fabricated cytosensor was sensitive, selective, reliable and stable, which was benefitted from the inherent merits of phage-displayed specific octapeptides and the superiority of EIS. To monitor the fabrication process of biosensor and the recognition of analytes on
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the electrode surface, EIS technique was used. Nyquist diagrams changed incrementally with the successive modification of the Au electrode surface. Bare Au electrode surface was cleaned and activated, and exhibited a good electron transfer rate. After 3-
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mercaptopropionic acid (MPA) was covalently attached onto the Au electrode surface to
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form a carboxyl monolayer by the strong interaction of Au−S bonds, the Rct value increased from 79.9 ± 6.4 Ω to 578 ± 21 Ω, indicating the formation of an insulating layer of MPA. Afterwards, the carboxyl groups were immediately activated by EDC/NHS and the Rct value mildly increased to 703 ± 29 Ω. With the assembly of phage, the Rct value remarkably increased to 2245 ± 42 Ω, indicating the success of covalent immobilization of phage onto the surface of Au electrode by condensation reaction between carboxyl group and amine group. On the other hand, in order to investigate the performance of cytosensors for quantitative detection of tumor cells, EIS
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ACCEPTED MANUSCRIPT technique was used. The proposed cytosensor demonstrated a wide linear range (2.0·102 − 2.0·108 cells·mL−1), a low limit of detection (79 cells·mL−1, S/N = 3), high specificity, good inter-and intra-assay reproducibility and satisfactory storage stability. Recently, Song et al. has reported an electrochemical cytosesensor for the tumor marker MUC1
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determination [92]. MUC1-binding aptamer immobilized on Au NPs modified GCE provided the basis for the electrochemical assays was developed. By means of EIS, the charge transfer resistance is monitored in order to control the proteins immobilization
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onto the surface of the sensor. Good results were obtained with the proposed cytosensor in the determination id MUC1 in human serum samples.
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Enzymatic chemical sensors: Electrochemical impedance spectroscopy is now becoming extensively used for the characterization of electrochemical sensors and biosensors. EIS has been used to develop a methodology able to identify and quantify fermentable sugars present in the enzymatic hydrolysis phase of second-generation
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bioethanol production from pineapple waste. Conesa et al. developed a low-cost nondestructive system consisting of a stainless double needle electrode associated to an electronic equipment that allows the implementation of EIS [93]. These results allow
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them to introduce this EIS-based technique as an easy, fast, non-destructive, and in situ alternative to the traditional laboratory methods for enzymatic hydrolysis monitoring.
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Same authors developed a device called Advanced Voltammetry, Impedance Spectroscopy & Potentiometry Analyzer (AVISPA) equipped with a specific software application and a stainless steel double needle electrode for monitoring pineapple waste hydrolysis process. The validation of an EIS-based method to monitor the saccharification process of industrial pineapple waste was introduced as an innovative analytical procedure. On the other hand, Halliwell et al. developed a gold electrodes modified with self-assembled monolayers of the SNARE protein SNAP-25 which is
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ACCEPTED MANUSCRIPT selectively cleaved by active botulinum neurotoxin A [94]. It was possible to observe changes to the layer on addition of the toxin by EIS means. Moreover, the electrochemical impedance spectroscopy assay also allowed for detection of the active toxin at concentrations as low as 25 fg·mL-1. These results were obtained in under an
method for the detection of the active form of this toxin.
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hour outperforming the mouse bioassay which is considered to be the most reliable
Non-enzymatic chemical sensors: Glucose monitoring is one of the most
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common parameter in food industry, biotechnology and clinical diagnosis and it has been growing in recent years, accompanied by a steady increase in the number of
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proposed devices. Basically, glucose electrochemical sensors can be divided into two categories: enzymatic and non-enzymatic electrochemical sensors. Many investigators have been undertaken on non-enzymatic glucose sensors in order to overcome the disadvantages of enzyme and mediators in biosensors. In the last years the use of
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electrochemical impedance spectroscopy has been proposed as a transduction principle in the non-enzymatic sensors. Rinaldi et al. reported an impedimetric non-enzymatic glucose sensor based on hydroxide thin film onto gold electrode (EAuNi(OH)2) [95].
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The electrochemical oxidation of glucose on EAuNi(OH)2 electrode was evaluated by EIS in a concentration range from 0 to 14.8 mM of analyte. These analyses show a good
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linear range, 0–2 mM of glucose, with a slope of 484.7 Ω·mM-1 of glucose. Moreover, EAuNi(OH)2 sensor was applied to determine glucose en blood samples. According to the results, EIS offers good sensibility and selectivity for the glucose detection by nonenzymatic glucose sensor as an alternative to conventional methods. Another example of this type of sensor has been reported by Ensafi et al. [96]. Authors presented an efficient, fast and stable non-enzymatic glucose sensor prepared by decorating silver nanoparticles on organic functionalized multiwall carbon nanotubes (AgNPs/F-
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ACCEPTED MANUSCRIPT MWCNTs). Using the Nyquist plots obtained by EIS, electrical conductivity between the redox probe and the modified electrode was confirmed and it was confirmed that the electron transfer was increasingly improved by the modification of the GCE, using FMWCNTs and AgNPs/F-MWCNTs, as the semicircle that corresponds to the electron
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transfer-limited process (kinetic control) was decreased. This decrease on the Rct was attributed to the excellent conductivity of AgNPs/F-MWCNTs, which allows for an easier electron transfer. The results obtained from the amperometric analysis of glucose
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indicated an efficient performance of the electrode with a low detection limit of 0.03 mM and a high sensitivity of 1057.3 mA·mM-1, as well as a linear dynamic range of 1.3
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to 1000 mM. Furthermore, a practical application of this sensor was also examined by analyzing glucose in the presence of common interfering species that exist in a real sample of human blood serum and no significant interference effect was found. As it has been reported on this section, it is usually possible to find several
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examples in literature where NC-based EIS bio-sensors provide wider linear range and better detection limits for the analytes than any other type of electrochemical tool. A comparative study of the analytical performance for different electrochemical platforms
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towards the determination of DNA, IgG, S. areus and Cyt c using different detection methodologies are summarized in Table 1. Compared to other methods reported on the
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literature, Faradaic EIS bio-sensors allow obtaining lower detection limits and wider linear range. Thus, EIS has become a simpler and much sensitive detection method for a wide range of analytes than other alternative electrochemical (i.e., amperometry or voltammetry) and/or optical (i.e., fluorescence) instrumental technique.
TABLE 1
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ACCEPTED MANUSCRIPT 4.
Surface engineering for Faradaic and non-Faradaic bio-sensing
As was mentioned in Section 3, Faradaic EIS transducting is the most common used for bio-sensing applications but the fact that redox probes are required, their use would be challenging in real applications and could therefore be in part limited to laboratory
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studies. Thus, while in Faradaic EIS (impedimetric) transduction occurs through changes in the impediment presented to a solution phase probe (commonly assessed through impedance), in non-Faradaic EIS (capacitive), transduction occurs through
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changes in the impediment presented to surface dielectric, charge distribution or local conductance (commonly assessed through capacitance) [119], [120].
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Since the affinity aspect is often the limiting factor for the label-free biosensor performance, immobilization of biomolecules on the electrode surface has become an important feature in designing the recognition part of electrochemical bio-sensors [121]. Different strategies have been developed to attach the biorecognition elements, which
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can be incorporated on the sensor device via modified semiconductor surfaces [122], metal oxides surfaces [123], composite surfaces and SAMs on electrode surfaces [124]. Various immobilization strategies have been developed, for example, for enzyme
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entrapment on electrochemical biosensors [125], [126], including (i) entrapment, (ii) adsorption, (iii) cross-linking and (iv) covalent immobilization. In this sense, in order to
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design the most practical impedimetric aptasensor for thrombin detection, Ocaña and co-workers used these four strategies for the immobilization of aptamers on a graphite– epoxy composite sensor [127]. In spite of obtaining a an increase in the Rct in all cases after the aptamer–thrombin interaction in the presence of the [Fe(CN)6]3-/4- redox couple, the strategy (ii) resulted in the lowest detection limit of the probe (4.5 pM). The first article on an affinity biosensor based on a capacitive transducer for liquids was published in the late 1980s [128]. However, the main drawback of
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ACCEPTED MANUSCRIPT capacitive bio-sensors based on electrode-solution interfaces is the necessity to form proper insulating layers on the top of the electrode surface. Although many approaches for surface design have been discussed, modified electrodes with SAMs have been exploited for several electroanalytical applications since they were introduced in the
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early 1990s [129], [130], [131]. Such SAMs are mainly used to promote the immobilization of the bio-recognition element that yields to increase the sensitivity, selectivity or other aspects of the electrochemical sensor [132]. Such films are formed
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spontaneously by chemisorptions and self-organization from solution and are particularly useful since they allow electrochemical insulation of the surface of a
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working electrode as well as prevent leakage currents, the ones responsible for shortcircuiting the solid electrode and solution phases and therefore, producing a loss of sensitivity [133], [134].
The developed SAM-based bio-sensor can be built by assembling recognition
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elements in a thin layer grafted on an electrode and measuring changes in the dielectric properties when an analyte binds to the biorecognition sides, causing a decrease in the capacitance value. Changes may also be induced when a large receptor molecule is
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displaced by a smaller analyte. For example, if a protein on a surface changes its conformation after binding of an analyte, which has been demonstrated for heavy metal
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ions, it might also be detected by capacitance measurements [135]. Hence, an optimum design of the electrochemical transducer allows to maximize
the analyte–interface interactions as well as suppress the interface-interface interactions [136]. The recognition through SAMs on conducting surfaces provides a superb platform for exploiting electrochemistry bio-sensing [137], [138], [139]. Most impedimetric biosensors utilize SAMs to attach probes at the electrode–solution interface. The most simplest and common approach employed in surface engineering
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ACCEPTED MANUSCRIPT design for impedimetric bio-sensor focus on organic monolayers is based on thiols covalently grafted to gold surfaces [140], [141] because of the spontaneous attachment they form, resulting in a permanent strong bond (see Figure 2A) [142], [143], [144]. In the literature has been presented a few major approaches used for functionalizing the
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SAM interface with biological recognition agents. These include: i) SAM attachment followed by further functionalization (Figure 2B); ii) functionalization previously to SAM attachment (Figure 2C); and iii) SAM attachment followed by recognition agent
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incorporation (Figure 2D) [137]. According to the literature, binding events happening on the surface of a capacitive bio-sensor produce contradictory signal changes (both
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increases and decreases in the Cdl capability), fact that can be attributed to various factors, including changes in the dielectric properties, the displacement of water molecules, electrostatic repulsions between recognition elements and ions in solution, the change of the molecular conformation onto the surface electrodes, among others. If
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a tightly-packed SAM is desirable for non-Faradaic sensors, for Faradaic sensors the electrode surface needs to be accessible to the redox species but not to adsorption of other molecules [145]. For example, long-chain alkanethiols produces more dense and
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ordered monolayer, due to hydrophobic interaction of the chains, which mimics a membrane-like microenvironment, useful for immobilizing biological (antibodies,
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enzymes, nucleic acids, etc.) and organic molecules [146]. In this sense, Radecka et. al. embedded macrocyclic polyamine molecules with long alkyl chain within 1dodecanethiol SAM for the detection of adenine nucleotides [147]. Despite long-chain thiols can form stable and well-ordered SAMs, their transfer rate of electrons is much slower since they can effectively block the electrode surface from unwanted reactions, and are thus often used in capacitive bio-sensors. For this reason, usually short chain
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ACCEPTED MANUSCRIPT alkanethiols are used in Faradaic EIS approaches, which enable the electron transfer across the layer but reduce the stability of the interface [124].
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FIGURE 2
As it was recently reviewed by Justino et al. [148], the capacitance of a modifiedSAM platform carrying a biorecognition element in non-Faradaic approach can be
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described as a combination of three capacitances: the first capacitance, which consists of an insulating layer comprised by a SAM on the transducer (Cm), the second one
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(Crec), which is related to the formation of a receptor biofilm (biorecognition layer) and then, the binding event (target-receptor biorecognition) generates a third layer (Ca), being the total capacitance (C) the combination in series of these three components. There are different examples of capacitive biosensors capable of quantifying
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important biomarkers at trace levels [149], [150]. For example, Lin et al. developed a sensitive impedimetric non-Faradaic immunosensor to detect two cardiac biomarkers in human serum, the C-reactive protein (CRP) and the myeloperoxidase, by analyzing the
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change in capacitance [151]. Using this approach, the researchers were able to detect both biomarkers in pg·mL-1 levels in human serum. Another example of a capacitive
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SAM-based immunosensor was the developed by Limbut et al., which was able to detect 10 pg·mL-1 of carcinoembryonic antigen (cancer biomarker) [152]. The absence of a redox couple in solution during the analysis makes the associated non-Faradaic approach more translatable to point-of-care devices. On the other hand, the impedance-derived ECS is an alternative powerful technique that allows the development of ultra sensitive biosensors where the addition of a redox probe is unnecessary [153]. From an electrochemical point of view, redox-active SAMs
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ACCEPTED MANUSCRIPT on gold electrodes for ECS-based platforms are an excellent system for investigating electron transfer kinetics since the spatial and electronic coupling of these sites to the underlying electrodes can be tightly controlled. In this sense, Porter et al. suggested that any of the following three mechanism for electron transfer are valid for a SAM-based
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modified electrode: a) permeation of the film; b) charge transfer at defect sites or c) electron tunneling through the monolayer [154]. Nonetheless, electrochemical analyses on confined redox-active SAMs cause an additional non-Faradaic and uncompensated
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resistance effects that, though unresolved, can strongly distort redox analysis [22]. ECS approaches can also help to resolve all contributions derived from SAMs films,
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including those which are related to the layer dipolar/electrostatic relaxation characteristics [155], [156]. For example, Bueno et al. has recently developed different surface chemistries using CRP antigen and antibody interaction as biomolecular binding model aiming at constructing impedimetric and capacitive interfaces engineered by
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SAMs (see Figure 3) [28]. For the impedimetric SAM structure (Figure 3A), the Faradaic impedimetric measurements are carried out by using a [Fe(CN)6]3-/4benchmark redox probe. In such kind of systems the redox free receptive intermediates
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electron transfer, which depends on how the communication between the redox probe in solution and the metallic centers are impeded by the target binding. Otherwise, the
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redox capacitive biosensor approach (Figure 3B) is based on tethered redox centers within the SAM. In such kind of systems its electrochemical occupancy is affected by the change in its environment with the binding of the target, associated with Faradaic density of states. With this work, the authors demonstrated how both Faradaic (impedimetric) and non-Faradaic (capacitive) approaches are capable to achieve similar results, and the biochemistry affinity was shown to be independent of engineered surface or transducer signal. As it was mentioned in Section 2, ECS technique presents
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ACCEPTED MANUSCRIPT the benefits that no equivalent circuit is required to extract the capacitance value, making it highly suitable for use in multiplexing affinity arrays [51], [157]. FIGURE 3
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5. Optimization of NC-based bio-sensors by EIS characterization
Beyond bio-sensing applications, EIS technique has been recently used as a novel and powerful method of characterizing many of the electrical properties of materials and
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their interfaces with electronically conducting properties. It may be used to investigate the dynamics of bound or mobile charge in the bulk or interfacial regions of any kind of
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solid or liquid material [49], [158], [159], [160]. Since its introduction in bio-sensing applications, EIS has been mainly employed as a characterization technique for confirming the layer by layer fabrication of the electrochemical platforms [161], [162], [163].
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The incorporation of carbon nanoallotropes as filler conductive phase, such as nanohorns, carbon nanotubes (CNTs), graphene and graphite, have played a leading role in the analytical electrochemistry field because of their low cost, excellent electron
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kinetics, good chemical stability, facile tunability and high electroactive surface area [74], [164], [165]. The potential for NC materials reinforced with such carbon
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nanoallotropes, having extraordinary specific stiffness and strength, represent tremendous opportunity for application in the 21st century [20], [166], especially for Nanocomposite Carbon Paste Electrodes (NC–CPEs) development [72], [167], [168], [169]. Comparing to conventional solid electrodes, the integration of carbon nanoallotropes in a polymeric matrix confers to the final electrode a series of benefits, including robustness from a mechanical point of view, renewable surface, chemical inertness, stable response, low ohmic resistance, biocompatibility, printability for
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ACCEPTED MANUSCRIPT miniaturizing and suitability for a variety of bio-sensing applications, being a nontoxic and environmentally friendly type of electrode. But, above all, they are well-known for the possibility of modifying and fine tuning their composition to achieve the requirements of specific analytical applications [170].
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To date, the characterization and optimization of carbon/polymer nanocomposite and nanobiocomposite bio-sensors (either containing or not a biological material) had been done from the point of view of maximum conductivity and maximum electron-
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transfer [171], [172], [173]. However, it is known that high carbon loadings can increase the background current and smear the Faradaic signal response, being this fact
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especially important when the electroactive species are present in low concentrations. Seeking to deduce the optimum carbon/polymer NC composition to avoid such drawbacks, Olivé-Monllau and co-authors [174] developed in 2010 a characterization protocol via EIS which is focused on optimizing several physical parameters, such as
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RΩ, Rct and Cdl for different compositions of CNT/epoxy NC sensors. These physical parameters, which were previously described, are directly related with the system sensitivity, response time and signal-to-noise ratio, respectively; each one being
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important in the overall electroanalytical performance of the final NC sensor, also affecting the reproducibility of the system [175]. In terms of electroanalytical response,
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this approach has been demonstrated by means of voltammetric detection of paracetamol, both in synthetic and pharmaceutical preparations [176], and also for the amperometric determination of free chlorine in water [177], [178]. Further, EIS technique also allowed to explain the direct electrochemical effect of semi-conducting nanocrystals, as CdS quantum dots, when they were integrated in such kind of CNTbased NC sensors [179]. After doping CNTs with CdS quantum dots, a superb synergistic electrocatalytic effect was observed, decreasing drastically the Cdl value and
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ACCEPTED MANUSCRIPT thus, to an improved signal-to-noise ratio. This effect was mainly related to the distribution of these semiconductors within the CNT-based NC that forms much more conductive microzones exposed on the electrode surface, behaving like a microelectrode array. This setup yielded to an excellent improvement in sensitivity and detection limits
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for the amperometric determination of both ascorbic acid and hydrogen peroxide [180]. Regarding to CNT-based NC–CPEs, their main drawback resides in the lack of homogeneity of the different commercial CNT lots due to the different amounts of
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impurities contained in the tubes as well as the dispersion in their diameter/length and state of aggregation [35], [181]. Nevertheless, Muñoz et al. developed an accurate EIS
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characterization protocol to enhance the electroanalytical response of CNT/polymer nanocomposite sensors regarding to the nature of the raw carbon nanofiller material [33], [182]. Preparing different series of NCs (from 1% to 20% w/w in conducting nanofiller just changing the nature of the CNTs), it was demonstrated that meanwhile
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the diameter is a key parameter that determines the electrochemical behavior of CNTbased NC sensors, purity and length are two parameters that remained unaltered since no significant changes in the optimum composition ratios were observed. Interestingly,
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the best electrochemical and therefore, electroanalytical performances were obtained using non-purified CNTs, thanks to the electrocatalytic activity shown by the metal
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impurities, fact that was also verified by decorating CNTs with different functional metal nanoparticles (i.e., Ag, Au, Pd, Pt and CuO nanoparticles) [183], [184], [185]. Rapidly, the EIS technique was spread to the characterization of NC sensors and
biosensors based on other carbon nanoallotropes, such as graphite [19] and graphene [186], demonstrating that EIS is a useful tool to optimize carbon-based electrochemical nanocomposite sensors from 1D (carbon nanotubes) to 2D (graphene) and 3D (graphite) filler nanomaterials. As it is depicted in Figure 4, the same trend in terms of RΩ, RCT
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ACCEPTED MANUSCRIPT and Cdl is observed for different carbon-based nanocomposite electrodes, independently of the filler nature. This fact was demonstrated for (a) CNT–, (b) graphene– and (c) graphite–based nanocomposite devices. Briefly, RΩ and Rct values decrease and Cdl value increases with the carbon filler loading. It can be explained since: i) whereas at
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low carbon loads the RΩ is dominated by ohmic composite resistance, at higher carbon loads it is dominated by the solution resistance; ii) Rct decreases with increasing carbon loading because of the rise of electroactive sites and thus, the electroactive area; iii) Cdl
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value increases for those electrodes comprising high surface area of conducting material due to the electrode capacitance can be determined nearly exclusively by the exposed
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carbon on the electrode surface. According to EIS characterization, the optimum composition is the one that just falls in the near-percolation composition zone [187], since that composition ensures a compromise between low resistances (quick electron transfer) and low capacitances (low background currents and high signal-to-noise
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ratios). Thus, some of the benefits of decreasing the maximum carbon filler loading in the NC for achieving the near-percolation composition are better detection limits, wider lineal range and increase of the stability as well as the repeatability of the analytical
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signal. Besides this approach produces a decrease of sensitivity of the system, the analytical response is well compensated by an increase in the signal-to-noise ratio. For
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example, a graphene/epoxy NC based on near-percolation composition demonstrated to be the most suitable for determining amperometrically thyroxine, achieving the most sensitive electrochemical device with respect to any other electrochemical bio-sensor previously described on the literature [188]. This electrochemical platform is based on functionalized properly the graphene sheets with Au nanoparticles, which were then used as nanotemplates for biorecognition agent immobilization. Furthermore, this work also reveals the importance of engineering surface since the electrochemical platform
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ACCEPTED MANUSCRIPT presented superb response towards thyroxine when biorecognition agent was immobilized upon Au nanoparticles but no response when biorecognition agent was immobilized directly into the NC paste. Comparing the different carbon-based NC sensors dispersed throughout the epoxy
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resin referenced above, using ascorbic acid as a model analyte, a highly decrease in the detection limits was observed when graphene was used as the carbon filler [186]. While the optimized sensors based on graphite (15% w/w) and CNTs (10% w/w for the ones
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with: diameter of 30–45 nm, length of 0.5–2.0 µm and purity > 99%) presented a LOD of 17 µg·L-1 and 7.0 µg·L-1, respectively, this parameter was significantly decreased
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(3.0 µg·L-1) for the optimized graphene-based NC sensors (13% w/w). Further, if graphene is compared with graphite, the simple bulk graphite exfoliation via Hummers method [189] for graphene synthesis presents important benefits from an analytical point of view, resulting in enhanced transducers because of they provide lowest
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detection limits. Otherwise, if graphene is compared with CNTs, the 2D nanomaterial present significant advantages since it does not contain metallic impurities and is produced from flacked graphite, which is economical and accessible. Based on these
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results, the graphene seems to be the most suitable carbon filler for the development of NC-based transducers for electrochemical bio-sensing approaches.
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Regarding graphite-based nanobiocomposite sensors, which incorporates an
enzyme model on the matrix, the first report on the characterization and optimization of the nanobiocomposite composition using EIS technique was performed in 2015 [32]. In this work, glucose oxidase (GOD) was immobilized as an ezyme model throughtout an epoxy resin containing graphite as the filler carbon material. Two series of biocomposite were constructed varying the conducting filler material from 13% to 20% and GOD between 1% and 2% in w/w. The optimization of the conductive particle
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ACCEPTED MANUSCRIPT distribution and the amount of the biological material inside the biocomposite allowed a significant improvement of different electrochemical properties, resulting in an enhancement of several electroanalytical parameters, such as signal stability,
FIGURE 4
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reproducibility and limit of detection.
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In addition, it was also demonstrated that the amount of filler loading in the nanobiocomposite is more critical than the enzyme loading to obtain lower limit of
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detections. Thus, this EIS-based characterization and optimization strategy was also recently extended to NC immunosensors based on graphite–epoxy containing antibodies (IgG) as biological agents [190]. For first time, and using the anti-rabbit Immunoglobulin G (RIgG) antibody as the proof, EIS technique allowed for optimizing
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the antigen–antibody ratio using a competitive assay, resulting in a reduction of the immunological species used in the competitive assay for achieving lower detection limits. Consequently, an appropriate characterization of the composition ratios is a sine
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qua non condition to improve the detection limit of immunospecies.
Conclusions and future perspectives
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6.
One of the main advantages of working with electrochemical bio-sensors based on EIS measurements is the small amplitude perturbation from steady state, which makes EIS a non-destructive technique, as well as its label-free measurements and excellent sensitivities. Engineering the bio-electrochemical sensing interface is of paramount importance to achieve highly sensitive and selective EIS transducers. Synergetic effect
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ACCEPTED MANUSCRIPT of different functional nanomaterials, especially carbon nanomaterials (i.e., carbon nanotubes, graphene, graphite and related), became promising approaches for highly sensitive bio-detection applications. Faradaic EIS methodology is at the moment the most promising example where carbon-based nanocomposite materials were shown to
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be particularly versatile when combined with nanosized acceptors. These carbon-based nanocomposite materials as core transducer platforms have shown several advantages in EIS bio-sensing because of their good electron conductivity, electrocatalytic properties,
sensitivity and stability of electrochemical devices.
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biocompatibility and high surface area, which are beneficial for improving the
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Despite the research done at the academic level and the significant achievements reached in the design and development of highly sensitive and selective Faradaic EIS platforms, these developments are not reflected enough in the availability of widely used commercial products due to the fact that redox probes are required. In this sense,
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surface chemistry approaches by self-assembled monolayers towards the development of capacitive bio-sensors via immobilizing recognition elements over the electrode surface have lead recently new trends in non-Faradaic-derived EIS bio-sensing. This
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strategy, which is based on measuring changes in the dielectric properties when an analyte binds, seems to be a simple, viable and promising alternative for the
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development of economical, ultra sensitive and at point-of-use devices where the addition of a redox probe is unnecessary. Besides the impressive progress that has been recently made in the field of
surface chemistry architecture to develop switchable EIS (redox probe in solution) and ECS (redox probe confined) transducers, many of the bio-sensors described in the literature are still far from real-life application, especially regarding the stability and reproducibility. In this sense, several challenges and obstacles remain to be overcome,
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ACCEPTED MANUSCRIPT such as their application beyond buffer solution, owing to the sample matrix effects should be considered carefully in real sample analysis, and additional problems like non-specific binding which will increase the detection limits are then likely to occur. Thus, new multifunctional nanocomposites, which integrate different physico-chemical
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properties within a single nanoscale complex, would be discovered for novel bio-nano interactions, leading to alternative conjugating EIS method with improved performance. Accordingly, future development for the production of either Faradaic (impedimetric) or
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non-Faradaic (capacitive) EIS bio-sensors with better quality includes investigation of
as measurements on real samples.
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new immobilization procedures, new nanocomposite-based transducer materials as well
Beyond bio-sensing purposes, EIS technique has also demonstrated its feasibility on the characterization of modified-surface, electrode processes and determining diffusion kinetics and mass-transport parameters or reaction interfaces. As was
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discussed above, the potential for nano(bio)composites reinforced with nanostructured carbon materials represent tremendous opportunity for application in the 21st century. However, it has been discussed with specific examples using different 1D, 2D and 3D
factors
that
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conducting carbon fillers how both the nature and the ratio of each component are key determines
the
electroanalytical
performance
of
the
resulting
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electrochemical nanocomposite bio-sensors. Such reasons advocates that an optimum characterization of the raw nanocomposite material before being used for electrochemical purposes is a must to achieve transducer platforms with enhanced electroanalytical parameters. Herein, this review has demonstrated the usefulness of EIS technique as an alternative and novel characterization tool to achieve the aforementioned goal in carbon-based nanocomposite sensors, which might be easily extended to alternative NC materials.
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Acknowledgements Jose Muñoz and Raquel Montes thank Universitat Autònoma de Barcelona (UAB) for
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the PhD fellowships and mobility grants during PhD studies.
References
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[1] P. Yang, Q. Zheng, H. Xu, J. Liu, L. Jin, A Highly Sensitive Electrochemical Impedance Spectroscopy Immunosensor for Determination of 1‐Pyrenebutyric Acid Based on the Bifunctionality of Nafion/Gold Nanoparticles Composite Electrode, Chinese Journal of Chemistry, 30 (2012) 1155-1162. [2] B.-Y. Chang, S.-M. Park, Electrochemical impedance spectroscopy, Annual Review of Analytical Chemistry, 3 (2010) 207-229. [3] E. Barsoukov, J.R. Macdonald, Impedance spectroscopy: theory, experiment, and applications, John Wiley & Sons2005. [4] E.P. Randviir, C.E. Banks, Electrochemical impedance spectroscopy: an overview of bioanalytical applications, Analytical Methods, 5 (2013) 1098-1115. [5] F. Lisdat, D. Schäfer, The use of electrochemical impedance spectroscopy for biosensing, Analytical and bioanalytical chemistry, 391 (2008) 1555. [6] A. Bonanni, M. Del Valle, Use of nanomaterials for impedimetric DNA sensors: a review, Analytica chimica acta, 678 (2010) 7-17. [7] C. Zhu, G. Yang, H. Li, D. Du, Y. Lin, Electrochemical sensors and biosensors based on nanomaterials and nanostructures, Analytical chemistry, 87 (2014) 230-249. [8] R. Elshafey, C. Tlili, A. Abulrob, A.C. Tavares, M. Zourob, Label-free impedimetric immunosensor for ultrasensitive detection of cancer marker Murine double minute 2 in brain tissue, Biosensors and Bioelectronics, 39 (2013) 220-225. [9] R. Ohno, H. Ohnuki, H. Wang, T. Yokoyama, H. Endo, D. Tsuya, M. Izumi, Electrochemical impedance spectroscopy biosensor with interdigitated electrode for detection of human immunoglobulin A, Biosensors and Bioelectronics, 40 (2013) 422-426. [10] L. Wang, J. Zhao, X. He, J. Gao, J. Li, C. Wan, C. Jiang, Electrochemical impedance spectroscopy (EIS) study of LiNi1/3Co1/3Mn1/3O2 for Li-ion batteries, Int. J. Electrochem. Sci, 7 (2012) 345-353.
35
ACCEPTED MANUSCRIPT
AC C
EP
TE D
M AN U
SC
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[11] M. Karttunen, P. Ruuskanen, V. Pitkänen, W.M. Albers, Electrically conductive metal polymer nanocomposites for electronics applications, Journal of Electronic Materials, 37 (2008) 951-954. [12] X. Jia, S. Dong, E. Wang, Engineering the bioelectrochemical interface using functional nanomaterials and microchip technique toward sensitive and portable electrochemical biosensors, Biosensors and Bioelectronics, 76 (2016) 80-90. [13] F. Arduini, L. Micheli, D. Moscone, G. Palleschi, S. Piermarini, F. Ricci, G. Volpe, Electrochemical biosensors based on nanomodified screen-printed electrodes: Recent applications in clinical analysis, TrAC Trends in Analytical Chemistry, 79 (2016) 114-126. [14] J. Li, N. Wu, Biosensors based on nanomaterials and nanodevices, CRC Press2013. [15] M. Holzinger, A. Le Goff, S. Cosnier, Synergetic Effects of Combined Nanomaterials for Biosensing Applications, Sensors, 17 (2017) 1010. [16] J. Jordan, K.I. Jacob, R. Tannenbaum, M.A. Sharaf, I. Jasiuk, Experimental trends in polymer nanocomposites—a review, Materials Science and Engineering: A, 393 (2005) 1-11. [17] A. Mamalis, L. Vogtländer, A. Markopoulos, Nanotechnology and nanostructured materials: trends in carbon nanotubes, Precision Engineering, 28 (2004) 16-30. [18] S. Alwarappan, C. Liu, A. Kumar, C.-Z. Li, Enzyme-doped graphene nanosheets for enhanced glucose biosensing, The Journal of Physical Chemistry C, 114 (2010) 12920-12924. [19] R. Montes, J. Bartrolí, F. Céspedes, M. Baeza, Towards to the improvement of the analytical response in voltammetric sensors based on rigid composites, Journal of Electroanalytical Chemistry, 733 (2014) 69-76. [20] H.J. Salavagione, A.M. Diez-Pascual, E. Lazaro, S. Vera, M.A. Gomez-Fatou, Chemical sensors based on polymer composites with carbon nanotubes and graphene: the role of the polymer, Journal of Materials Chemistry A, 2 (2014) 14289-14328. [21] P.R. Bueno, G. Mizzon, J.J. Davis, Capacitance spectroscopy: a versatile approach to resolving the redox density of states and kinetics in redox-active self-assembled monolayers, The Journal of Physical Chemistry B, 116 (2012) 8822-8829. [22] C.G. Vayenas, S. Bebelis, S. Neophytides, Non-Faradaic electrochemical modification of catalytic activity, The Journal of Physical Chemistry, 92 (1988) 5083-5085. [23] S.H. Kazemi, M. Shanehsaz, M. Ghaemmaghami, Non-Faradaic electrochemical impedance spectroscopy as a reliable and facile method: Determination of the potassium ion concentration using a guanine rich aptasensor, Materials Science and Engineering: C, 52 (2015) 151-154. [24] N. Phares, R.J. White, K.W. Plaxco, Improving the stability and sensing of electrochemical biosensors by employing trithiol-anchoring groups in a six-carbon self-assembled monolayer, Analytical chemistry, 81 (2009) 1095-1100. [25] P. Solanki, S. Arya, Y. Nishimura, M. Iwamoto, B. Malhotra, Application of Self-Assembled Monolayer of 10-Carboxy-1-Decanethiol for Cholesterol Biosensor, Journal of Biomedical & Pharmaceutical Engineering, 2 (2008) 7-13. [26] R.G. Nuzzo, D.L. Allara, Adsorption of bifunctional organic disulfides on gold surfaces, Journal of the American Chemical Society, 105 (1983) 4481-4483. [27] A. Santos, J.J. Davis, P.R. Bueno, Fundamentals and Applications of Impedimetric and Redox Capacitive Biosensors, Journal of Analytical & Bioanalytical Techniques, 2015 (2014). [28] F.C. Fernandes, A. Santos, D.C. Martins, M.S. Goes, P.R. Bueno, Comparing label free electrochemical impedimetric and capacitive biosensing architectures, Biosensors and Bioelectronics, 57 (2014) 96-102. [29] M.S. Goes, H. Rahman, J. Ryall, J.J. Davis, P.R. Bueno, A Dielectric Model of Self-Assembled Monolayer Interfaces by CapaciMve Spectroscopy†, Langmuir, 28 (2012) 9689-9699. [30] V. Tsouti, C. Boutopoulos, I. Zergioti, S. Chatzandroulis, Capacitive microsystems for biological sensing, Biosensors and Bioelectronics, 27 (2011) 1-11. [31] A. Santos, J.P. Piccoli, N.A. Santos-Filho, E.M. Cilli, P.R. Bueno, Redox-tagged peptide for capacitive diagnostic assays, Biosensors and Bioelectronics, 68 (2015) 281-287.
36
ACCEPTED MANUSCRIPT
AC C
EP
TE D
M AN U
SC
RI PT
[32] R. Montes, J. Bartrolí, M. Baeza, F. Céspedes, Improvement of the detection limit for biosensors: Advances on the optimization of biocomposite composition, Microchemical Journal, 119 (2015) 66-74. [33] J. Muñoz, J. Bartrolí, F. Céspedes, M. Baeza, Influence of raw carbon nanotubes diameter for the optimization of the load composition ratio in epoxy amperometric composite sensors, Journal of Materials Science, 50 (2015) 652-661. [34] R.L. McCreery, Advanced Carbon Electrode Materials for Molecular Electrochemistry, Chemical Reviews, 108 (2008) 2646-2687. [35] M. Morcom, K. Atkinson, G.P. Simon, The effect of carbon nanotube properties on the degree of dispersion and reinforcement of high density polyethylene, Polymer, 51 (2010) 35403550. [36] B. Pejcic, R. De Marco, Impedance spectroscopy: Over 35 years of electrochemical sensor optimization, Electrochimica Acta, 51 (2006) 6217-6229. [37] M.E. Orazem, B. Tribollet, Electrochemical impedance spectroscopy, John Wiley & Sons2011. [38] Y. Xu, X. Xie, Y. Duan, L. Wang, Z. Cheng, J. Cheng, A review of impedance measurements of whole cells, Biosensors and Bioelectronics, 77 (2016) 824-836. [39] A. Lasia, Electrochemical impedance spectroscopy and its applications, Modern aspects of electrochemistry, Springer2002, pp. 143-248. [40] J.R. Macdonald, E. Barsoukov, Impedance spectroscopy: theory, experiment, and applications, History, 1 (2005). [41] K.S. Cole, R.H. Cole, Dispersion and absorption in dielectrics I. Alternating current characteristics, The Journal of chemical physics, 9 (1941) 341-351. [42] B. Maundy, A. Elwakil, Extracting single dispersion Cole–Cole impedance model parameters using an integrator setup, Analog integrated circuits and signal processing, 71 (2012) 107-110. [43] Y.-T. Tsai, D. Whitmore, Nonlinear least-squares analyses of complex impedance and admittance data for solid electrolytes, Solid state ionics, 7 (1982) 129-139. [44] J.R. Scully, D.C. Silverman, Electrochemical impedance: analysis and interpretation, ASTM International1993. [45] A. Amirudin, D. Thierry, Application of electrochemical impedance spectroscopy to study the degradation of polymer-coated metals, Progress in Organic Coatings, 26 (1995) 1-28. [46] P.M. Levine, P. Gong, R. Levicky, K.L. Shepard, Real-time, multiplexed electrochemical DNA detection using an active complementary metal-oxide-semiconductor biosensor array with integrated sensor electronics, Biosensors and Bioelectronics, 24 (2009) 1995-2001. [47] J. Bisquert, G. Garcia-Belmonte, P. Bueno, E. Longo, L. Bulhoes, Impedance of constant phase element (CPE)-blocked diffusion in film electrodes, Journal of Electroanalytical Chemistry, 452 (1998) 229-234. [48] B. Hirschorn, M.E. Orazem, B. Tribollet, V. Vivier, I. Frateur, M. Musiani, Determination of effective capacitance and film thickness from constant-phase-element parameters, Electrochimica Acta, 55 (2010) 6218-6227. [49] J. Bisquert, Theory of the impedance of electron diffusion and recombination in a thin layer, The Journal of Physical Chemistry B, 106 (2002) 325-333. [50] N. Couniot, A. Afzalian, N. Van Overstraeten-Schlögel, L. Francis, D. Flandre, Capacitive biosensing of bacterial cells: analytical model and numerical simulations, Sensors and Actuators B: Chemical, 211 (2015) 428-438. [51] A. Santos, F.C. Carvalho, M.-C. Roque-Barreira, P.R. Bueno, Impedance-derived electrochemical capacitance spectroscopy for the evaluation of lectin–glycoprotein binding affinity, Biosensors and Bioelectronics, 62 (2014) 102-105. [52] J. Cecchetto, F.C. Fernandes, R. Lopes, P.R. Bueno, The capacitive sensing of NS1 Flavivirus biomarker, Biosensors and Bioelectronics, 87 (2017) 949-956.
37
ACCEPTED MANUSCRIPT
AC C
EP
TE D
M AN U
SC
RI PT
[53] V.F. Lvovich, Impedance spectroscopy: applications to electrochemical and dielectric phenomena, John Wiley & Sons2012. [54] C. Gabrielli, Electrochemical impedance spectroscopy: principles, instrumentation, and applications, Physical Electrochemistry: Principles, Methods, and Applications (1995) 243-292. [55] F. Mansfeld, Electrochemical impedance spectroscopy (EIS) as a new tool for investigating methods of corrosion protection, Electrochimica Acta, 35 (1990) 1533-1544. [56] H. Cesiulis, N. Tsyntsaru, A. Ramanavicius, G. Ragoisha, The Study of Thin Films by Electrochemical Impedance Spectroscopy, Nanostructures and Thin Films for Multifunctional Applications, Springer 2016, pp. 3-42. [57] G. Song, Equivalent circuit model for AC electrochemical impedance spectroscopy of concrete, Cement and concrete research, 30 (2000) 1723-1730. [58] A.R. Cardoso, F.T.C. Moreira, R. Fernandes, M.G.F. Sales, Novel and simple electrochemical biosensor monitoring attomolar levels of miRNA-155 in breast cancer, Biosensors and Bioelectronics, 80 (2016) 621-630. [59]R. Akter, B. Jeong, Y.M. Lee, J.S. Choi, M.A. Rahman, Femtomolar detection of cardiac troponin I using a novel label-free and reagent-free dendrimer enhanced impedimetric immunosensor, Biosensors and Bioelectronics, 91 (2017) 637-643. [60] Z. Gao, H. Deng, W. Shen, Y. Ren, A label-free biosensor for electrochemical detection of femtomolar microRNAs, Analytical chemistry, 85 (2013) 1624-1630. [61] A. Brecht, G. Gauglitz, Optical probes and transducers, Biosensors and Bioelectronics, 10 (1995) 923-936. [62] B.T. Cunningham, L.G. Laing, Advantages and application of label-free detection assays in drug screening, Expert opinion on drug discovery, 3 (2008) 891-901. [63] H.K. Hunt, A.M. Armani, Label-free biological and chemical sensors, Nanoscale, 2 (2010) 1544-1559. [64] M. Vestergaard, K. Kerman, E. Tamiya, An overview of label-free electrochemical protein sensors, Sensors, 7 (2007) 3442-3458. [65] W.T. Van Den Beld, M. Odijk, R.H. Vervuurt, J.-W. Weber, A.A. Bol, A. Van Den Berg, J.C. Eijkel, In-situ Raman spectroscopy to elucidate the influence of adsorption in graphene electrochemistry, Scientific reports, 7 (2017) 45080. [66] M.C. Rodriguez, A.-N. Kawde, J. Wang, Aptamer biosensor for label-free impedance spectroscopy detection of proteins based on recognition-induced switching of the surface charge, Chemical Communications (2005) 4267-4269. [67] P. Ćwik, U.E. Wawrzyniak, M. Jańczyk, W. Wróblewski, Electrochemical studies of selfassembled monolayers composed of various phenylboronic acid derivatives, Talanta, 119 (2014) 5-10. [68] H. Yan, J. Zhang, C. You, Z. Song, B. Yu, Y. Shen, Influences of different synthesis conditions on properties of Fe 3 O 4 nanoparticles, Materials Chemistry and Physics, 113 (2009) 46-52. [69] N. Sekar, R.P. Ramasamy, Electrochemical impedance spectroscopy for microbial fuel cell characterization, J Microb Biochem Technol S, 6 (2013). [70] D.D. Macdonald, Reflections on the history of electrochemical impedance spectroscopy, Electrochimica Acta, 51 (2006) 1376-1388. [71] V.M. Huang, S.-L. Wu, M.E. Orazem, N. Pébère, B. Tribollet, V. Vivier, Local electrochemical impedance spectroscopy: A review and some recent developments, Electrochimica Acta, 56 (2011) 8048-8057. [72] J. Muñoz, M. Baeza, Customized Bio-Functionalization of Nanocomposite Carbon Paste Electrodes for Electrochemical Sensing: A Mini Review, Electroanalysis, 29 (2017) 1660-1669. [73] A. Warsinke, B. Nagel, Towards separation-free electrochemical affinity sensors by using antibodies, aptamers, and molecularly imprinted polymers—a review, Analytical letters, 39 (2006) 2507-2556. [74] P. Ramnani, N.M. Saucedo, A. Mulchandani, Carbon nanomaterial-based electrochemical biosensors for label-free sensing of environmental pollutants, Chemosphere, 143 (2016) 85-98.
38
ACCEPTED MANUSCRIPT
AC C
EP
TE D
M AN U
SC
RI PT
[75] R. Zhang, W. Chen, Recent advances in graphene-based nanomaterials for fabricating electrochemical hydrogen peroxide sensors, Biosensors and Bioelectronics, 89 (2017) 249-268. [76] Y. Song, Y. Luo, C. Zhu, H. Li, D. Du, Y. Lin, Recent advances in electrochemical biosensors based on graphene two-dimensional nanomaterials, Biosensors and Bioelectronics, 76 (2016) 195-212. [77] Y. Fang, E. Wang, Electrochemical biosensors on platforms of graphene, Chemical Communications, 49 (2013) 9526-9539. [78] E.B. Bahadır, M.K. Sezgintürk, Applications of graphene in electrochemical sensing and biosensing, TrAC Trends in Analytical Chemistry, 76 (2016) 1-14. [79] L. Wang, M. Pumera, Electrochemical catalysis at low dimensional carbons: Graphene, carbon nanotubes and beyond–A review, Applied Materials Today, 5 (2016) 134-141. [80] A. Martín, M.Á. López, M.C. González, A. Escarpa, Multidimensional carbon allotropes as electrochemical detectors in capillary and microchip electrophoresis, Electrophoresis, 36 (2015) 179-194. [81] J.-Y. Park, S.-M. Park, DNA hybridization sensors based on electrochemical impedance spectroscopy as a detection tool, Sensors, 9 (2009) 9513-9532. [82] Q. Zhu, F. Gao, Y. Yang, B. Zhang, W. Wang, Z. Hu, Q. Wang, Electrochemical preparation of polyaniline capped Bi 2 S 3 nanocomposite and its application in impedimetric DNA biosensor, Sensors and Actuators B: Chemical, 207 (2015) 819-826. [83] L.-q. Luo, Z. Zhang, Y.-p. Ding, D.-m. Deng, X.-l. Zhu, Z.-x. Wang, Label-free electrochemical impedance genosensor based on 1-aminopyrene/graphene hybrids, Nanoscale, 5 (2013) 58335840. [84] A. Bonanni, M.I. Pividori, M. del Valle, DNA polymorphism sensitive impedimetric detection on gold-nanoislands modified electrodes, Talanta, 136 (2015) 95-101. [85] B. Derkus, K.C. Emregul, H. Mazi, E. Emregul, T. Yumak, A. Sinag, Protein A immunosensor for the detection of immunoglobulin G by impedance spectroscopy, Bioprocess and biosystems engineering, 37 (2014) 965-976. [86] L. Yu, Y. Zhang, C. Hu, H. Wu, Y. Yang, C. Huang, N. Jia, Highly sensitive electrochemical impedance spectroscopy immunosensor for the detection of AFB 1 in olive oil, Food chemistry, 176 (2015) 22-26. [87] Y. Wang, J. Feng, Z. Tan, H. Wang, Electrochemical impedance spectroscopy aptasensor for ultrasensitive detection of adenosine with dual backfillers, Biosensors and Bioelectronics, 60 (2014) 218-223. [88] F. Jia, N. Duan, S. Wu, X. Ma, Y. Xia, Z. Wang, X. Wei, Impedimetric aptasensor for Staphylococcus aureus based on nanocomposite prepared from reduced graphene oxide and gold nanoparticles, Microchimica Acta, 181 (2014) 967-974. [89] C. Ocaña, S. Lukic, M. del Valle, Aptamer-antibody sandwich assay for cytochrome c employing an MWCNT platform and electrochemical impedance, Microchimica Acta, 182 (2015) 2045-2053. [90] L. Fan, G. Zhao, H. Shi, M. Liu, Z. Li, A highly selective electrochemical impedance spectroscopy-based aptasensor for sensitive detection of acetamiprid, Biosensors and Bioelectronics, 43 (2013) 12-18. [91] L. Han, P. Liu, V.A. Petrenko, A. Liu, A Label-Free Electrochemical Impedance Cytosensor Based on Specific Peptide-Fused Phage Selected from Landscape Phage Library, Scientific reports, 6 (2016) 22199. [92] J. Song, Y. Zhou, B. Chen, W. Lou, J. Gu, Development of Electrochemical Aptamer Biosensor for Tumor Marker MUC1 Determination, Int. J. Electrochem. Sci, 12 (2017) 56185627. [93] C. Conesa, E. García-Breijo, E. Loeff, L. Seguí, P. Fito, N. Laguarda-Miró, An Electrochemical Impedance Spectroscopy-Based Technique to Identify and Quantify Fermentable Sugars in Pineapple Waste Valorization for Bioethanol Production, Sensors, 15 (2015) 22941-22955.
39
ACCEPTED MANUSCRIPT
AC C
EP
TE D
M AN U
SC
RI PT
[94] J. Halliwell, A.C. Savage, N. Buckley, C. Gwenin, Electrochemical impedance spectroscopy biosensor for detection of active botulinum neurotoxin, Sensing and Bio-Sensing Research, 2 (2014) 12-15. [95] A.L. Rinaldi, R. Carballo, Impedimetric non-enzymatic glucose sensor based on nickel hydroxide thin film onto gold electrode, Sensors and Actuators B: Chemical, 228 (2016) 43-52. [96] A.A. Ensafi, N. Zandi-Atashbar, B. Rezaei, M. Ghiaci, M.E. Chermahini, P. Moshiri, Nonenzymatic glucose electrochemical sensor based on silver nanoparticle decorated organic functionalized multiwall carbon nanotubes, RSC Advances, 6 (2016) 60926-60932. [97] H. Chang, Y. Yuan, N. Shi, Y. Guan, Electrochemical DNA biosensor based on conducting polyaniline nanotube array, Analytical chemistry, 79 (2007) 5111-5115. [98] Y. Bo, H. Yang, Y. Hu, T. Yao, S. Huang, A novel electrochemical DNA biosensor based on graphene and polyaniline nanowires, Electrochimica Acta, 56 (2011) 2676-2681. [99] D. Zheng, Q. Wang, F. Gao, Q. Wang, W. Qiu, F. Gao, Development of a novel electrochemical DNA biosensor based on elongated hexagonal-pyramid CdS and polyisonicotinic acid composite film, Biosensors and Bioelectronics, 60 (2014) 167-174. [100] K.-J. Huang, Y.-J. Liu, H.-B. Wang, T. Gan, Y.-M. Liu, L.-L. Wang, Signal amplification for electrochemical DNA biosensor based on two-dimensional graphene analogue tungsten sulfide–graphene composites and gold nanoparticles, Sensors and Actuators B: Chemical, 191 (2014) 828-836. [101] C.-X. Xu, Q.-G. Zhai, Y.-J. Liu, K.-J. Huang, L. Lu, K.-X. Li, A novel electrochemical DNA biosensor construction based on layered CuS–graphene composite and Au nanoparticles, Analytical and bioanalytical chemistry, 406 (2014) 6943-6951. [102] M. Díaz-González, D. Hernández-Santos, M.B. González-García, A. Costa-García, Development of an immunosensor for the determination of rabbit IgG using streptavidin modified screen-printed carbon electrodes, Talanta, 65 (2005) 565-573. [103] G. Liu, J.J. Gooding, Towards the fabrication of label-free amperometric immunosensors using SWNTs, Electrochemistry Communications, 11 (2009) 1982-1985. [104] Z.-M. Liu, H.-F. Yang, Y.-F. Li, Y.-L. Liu, G.-L. Shen, R.-Q. Yu, Core–shell magnetic nanoparticles applied for immobilization of antibody on carbon paste electrode and amperometric immunosensing, Sensors and Actuators B: Chemical, 113 (2006) 956-962. [105] H. Zarei, H. Ghourchian, K. Eskandari, M. Zeinali, Magnetic nanocomposite of anti-human IgG/COOH–multiwalled carbon nanotubes/Fe 3 O 4 as a platform for electrochemical immunoassay, Analytical biochemistry, 421 (2012) 446-453. [106] J. Wang, D. Song, H. Zhang, J. Zhang, Y. Jin, H. Zhang, H. Zhou, Y. Sun, Studies of Fe 3 O 4/Ag/Au composites for immunoassay based on surface plasmon resonance biosensor, Colloids and Surfaces B: Biointerfaces, 102 (2013) 165-170. [107] S. Saha, S.K. Arya, S. Singh, K. Sreenivas, B. Malhotra, V. Gupta, Zinc oxide–potassium ferricyanide composite thin film matrix for biosensing applications, Analytica chimica acta, 653 (2009) 212-216. [108] F.-C. Loo, S.-P. Ng, C.-M.L. Wu, S.K. Kong, An aptasensor using DNA aptamer and white light common-path SPR spectral interferometry to detect cytochrome-c for anti-cancer drug screening, Sensors and Actuators B: Chemical, 198 (2014) 416-423. [109] I.P. Lau, E.K. Ngan, J.F. Loo, Y. Suen, H. Ho, S. Kong, Aptamer-based bio-barcode assay for the detection of cytochrome-c released from apoptotic cells, Biochemical and biophysical research communications, 395 (2010) 560-564. [110] J.-M. Liu, X.-P. Yan, Ultrasensitive, selective and simultaneous detection of cytochrome c and insulin based on immunoassay and aptamer-based bioassay in combination with Au/Ag nanoparticle tagging and ICP-MS detection, Journal of Analytical Atomic Spectrometry, 26 (2011) 1191-1197. [111] M. Pandiaraj, T. Madasamy, P.N. Gollavilli, M. Balamurugan, S. Kotamraju, V.K. Rao, K. Bhargava, C. Karunakaran, Nanomaterial-based electrochemical biosensors for cytochrome c using cytochrome c reductase, Bioelectrochemistry, 91 (2013) 1-7.
40
ACCEPTED MANUSCRIPT
AC C
EP
TE D
M AN U
SC
RI PT
[112] M. Pandiaraj, A.R. Benjamin, T. Madasamy, K. Vairamani, A. Arya, N.K. Sethy, K. Bhargava, C. Karunakaran, A cost-effective volume miniaturized and microcontroller based cytochrome c assay, Sensors and Actuators A: Physical, 220 (2014) 290-297. [113] J. Yuan, Y. Yu, C. Li, X. Ma, Y. Xia, J. Chen, Z. Wang, Visual detection and microplate assay for Staphylococcus aureus based on aptamer recognition coupled to tyramine signal amplification, Microchimica Acta, 181 (2014) 321-327. [114] N. Duan, S. Wu, C. Zhu, X. Ma, Z. Wang, Y. Yu, Y. Jiang, Dual-color upconversion fluorescence and aptamer-functionalized magnetic nanoparticles-based bioassay for the simultaneous detection of Salmonella Typhimurium and Staphylococcus aureus, Analytica chimica acta, 723 (2012) 1-6. [115] T. Miao, Z. Wang, S. Li, X. Wang, Sensitive fluorescent detection of Staphylococcus aureus using nanogold linked CdTe nanocrystals as signal amplification labels, Microchimica Acta, 172 (2011) 431-437. [116] G.A. Zelada-Guillén, J.L. Sebastián-Avila, P. Blondeau, J. Riu, F.X. Rius, Label-free detection of Staphylococcus aureus in skin using real-time potentiometric biosensors based on carbon nanotubes and aptamers, Biosensors and bioelectronics, 31 (2012) 226-232. [117] M. Ruan, C.-G. Niu, G.-M. Zeng, P.-Z. Qin, X.-Y. Wang, D.-W. Huang, J. Huang, C.-Z. Fan, Rapid detection of Staphylococcus aureus via a sensitive DNA hybridization assay based on a long-lifetime luminescent europium marker, Microchimica Acta, 175 (2011) 105-112. [118] Y.J. Sung, H.-J. Suk, H.Y. Sung, T. Li, H. Poo, M.-G. Kim, Novel antibody/gold nanoparticle/magnetic nanoparticle nanocomposites for immunomagnetic separation and rapid colorimetric detection of Staphylococcus aureus in milk, Biosensors and Bioelectronics, 43 (2013) 432-439. [119] C. Ocaña, A. Hayat, R.K. Mishra, A. Vasilescu, M. del Valle, J.-L. Marty, Label free aptasensor for Lysozyme detection: A comparison of the analytical performance of two aptamers, Bioelectrochemistry, 105 (2015) 72-77. [120] J.S. Daniels, N. Pourmand, Label-Free Impedance Biosensors: Opportunities and Challenges, Electroanalysis, 19 (2007) 1239-1257. [121] A. Sassolas, L.J. Blum, B.D. Leca-Bouvier, Immobilization strategies to develop enzymatic biosensors, Biotechnology Advances, 30 (2012) 489-511. [122] A. Barraud, H. Perrot, V. Billard, C. Martelet, J. Therasse, Study of immunoglobulin G thin layers obtained by the Langmuir-Blodgett method: application to immunosensors, Biosensors and Bioelectronics, 8 (1993) 39-48. [123] A. Gebbert, M. Alvarez-Icaza, W. Stoecklein, R.D. Schmid, Real-time monitoring of immunochemical interactions with a tantalum capacitance flow-through cell, Analytical Chemistry, 64 (1992) 997-1003. [124] V. Ganesh, S.K. Pal, S. Kumar, V. Lakshminarayanan, Self-assembled monolayers (SAMs) of alkoxycyanobiphenyl thiols on gold—A study of electron transfer reaction using cyclic voltammetry and electrochemical impedance spectroscopy, Journal of colloid and interface science, 296 (2006) 195-203. [125] R.A. Sheldon, Enzyme immobilization: the quest for optimum performance, Advanced Synthesis & Catalysis, 349 (2007) 1289-1307. [126] B. Brena, P. González-Pombo, F. Batista-Viera, Immobilization of enzymes: a literature survey, Immobilization of Enzymes and Cells: Third Edition (2013) 15-31. [127] C. Ocaña, M. del Valle, A comparison of four protocols for the immobilization of an aptamer on graphite composite electrodes, Microchimica Acta, 181 (2014) 355-363. [128] A. Newman, K. Hunter, W. Stanbro, The capacitive affinity sensor: a new biosensor, Proceedings of the Second International Meeting on Chemical Sensors, Bordeaux, France, 1986, pp. 596-598. [129] J. Sagiv, Organized monolayers by adsorption. 1. Formation and structure of oleophobic mixed monolayers on solid surfaces, Journal of the American Chemical Society, 102 (1980) 9298.
41
ACCEPTED MANUSCRIPT
AC C
EP
TE D
M AN U
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[130] E. Marchante, M. Maglione, N. Crivillers, C. Rovira, M. Mas-Torrent, A four-state capacitance molecular switch based on a redox active tetrathiafulvalene self-assembled monolayer, RSC Advances, 7 (2017) 5636-5641. [131] E. Marchante, N. Crivillers, M. Buhl, J. Veciana, M. Mas-Torrent, An Electrically Driven and Readable Molecular Monolayer Switch Based on a Solid Electrolyte, Angewandte Chemie, 128 (2016) 376-380. [132] N. Crivillers, M. Mas-Torrent, J. Vidal-Gancedo, J. Veciana, C. Rovira, Self-assembled monolayers of electroactive polychlorotriphenylmethyl radicals on Au (111), Journal of the American Chemical Society, 130 (2008) 5499-5506. [133] M. Riepl, V.M. Mirsky, I. Novotny, V. Tvarozek, V. Rehacek, O.S. Wolfbeis, Optimization of capacitive affinity sensors: drift suppression and signal amplification, Analytica Chimica Acta, 392 (1999) 77-84. [134] K. Wadu-Mesthrige, N.A. Amro, G.Y. Liu, Immobilization of proteins on self-assembled monolayers, Scanning, 22 (2000) 380-388. [135] I. Bontidean, C. Berggren, G. Johansson, E. Csöregi, B. Mattiasson, J.R. Lloyd, K.J. Jakeman, N.L. Brown, Detection of heavy metal ions at femtomolar levels using protein-based biosensors, Analytical chemistry, 70 (1998) 4162-4169. [136] N.F. Atta, E.H. El-Ads, A. Galal, Self-Assembled Monolayers on Nanostructured Composites for Electrochemical Sensing Applications, Handbook of Nanoelectrochemistry: Electrochemical Synthesis Methods, Properties, and Characterization Techniques, (2016) 417478. [137] D. Mandler, S. Kraus-Ophir, Self-assembled monolayers (SAMs) for electrochemical sensing, Journal of Solid State Electrochemistry, 15 (2011) 1535-1558. [138] A.L. Eckermann, D.J. Feld, J.A. Shaw, T.J. Meade, Electrochemistry of redox-active selfassembled monolayers, Coordination chemistry reviews, 254 (2010) 1769-1802. [139] R.K. Mendes, R.S. Freire, C.P. Fonseca, S. Neves, L.T. Kubota, Characterization of selfassembled thiols monolayers on gold surface by electrochemical impedance spectroscopy, Journal of the Brazilian Chemical Society, 15 (2004) 849-855. [140] E. Marchante, N. Crivillers Clusella, M. Buhl, J. Veciana Miró, M. Mas-Torrent, An electrically driven and readable molecular monolayer switch using a solid electrolyte, 128 (2016) 376-380. [141] J. Casado-Montenegro, E. Marchante, N. Crivillers, C. Rovira, M. Mas-Torrent, Donor/Acceptor Mixed Self-Assembled Monolayers for Realising a Multi-Redox-State Surface, ChemPhysChem, 17 (2016) 1810-814. [142] S. Arnold, Z.Q. Feng, T. Kakiuchi, W. Knoll, K. Niki, Investigation of the electrode reaction of cytochrome c through mixed self-assembled monolayers of alkanethiols on gold (111) surfaces, Journal of Electroanalytical Chemistry, 438 (1997) 91-97. [143] A. Tlili, A. Abdelghani, S. Hleli, M.A. Maaref, Electrical characterization of a thiol SAM on gold as a first step for the fabrication of immunosensors based on a quartz crystal microbalance, Sensors, 4 (2004) 105-114. [144] M. Guardingo, E. Bellido, R. Miralles-Llumà, J. Faraudo, J. Sedó, S. Tatay, A. Verdaguer, F. Busqué, D. Ruiz-Molina, Bioinspired Catechol-Terminated Self-Assembled Monolayers with Enhanced Adhesion Properties, Small, 10 (2014) 1594-1602. [145] R.Y. Lai, D.S. Seferos, A.J. Heeger, G.C. Bazan, K.W. Plaxco, Comparison of the Signaling and Stability of Electrochemical DNA Sensors Fabricated from 6- or 11-Carbon Self-Assembled Monolayers, Langmuir, 22 (2006) 10796-10800. [146] N.K. Chaki, K. Vijayamohanan, Self-assembled monolayers as a tunable platform for biosensor applications, Biosensors and Bioelectronics, 17 (2002) 1-12. [147] J. Radecki, I. Szymańska, L. Bulgariu, M. Pietraszkiewicz, Covalent and embedment immobilization of macrocyclic polyamines on gold electrodes and their voltammetric responses towards ethene dicarboxylic acids, Electrochimica acta, 51 (2006) 2289-2297.
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ACCEPTED MANUSCRIPT
AC C
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[148] C.I.L. Justino, A.C. Freitas, R. Pereira, A.C. Duarte, T.A.P. Rocha Santos, Recent developments in recognition elements for chemical sensors and biosensors, TrAC Trends in Analytical Chemistry, 68 (2015) 2-17. [149] M. Dijksma, B. Kamp, J. Hoogvliet, W. Van Bennekom, Development of an electrochemical immunosensor for direct detection of interferon-γ at the attomolar level, Analytical chemistry, 73 (2001) 901-907. [150] W. Limbut, P. Kanatharana, B. Mattiasson, P. Asawatreratanakul, P. Thavarungkul, A comparative study of capacitive immunosensors based on self-assembled monolayers formed from thiourea, thioctic acid, and 3-mercaptopropionic acid, Biosensors and Bioelectronics, 22 (2006) 233-240. [151] K.-C. Lin, V. Kunduru, M. Bothara, K. Rege, S. Prasad, B. Ramakrishna, Biogenic nanoporous silica-based sensor for enhanced electrochemical detection of cardiovascular biomarkers proteins, Biosensors and Bioelectronics, 25 (2010) 2336-2342. [152] W. Limbut, P. Kanatharana, B. Mattiasson, P. Asawatreratanakul, P. Thavarungkul, A reusable capacitive immunosensor for carcinoembryonic antigen (CEA) detection using thiourea modified gold electrode, Analytica chimica acta, 561 (2006) 55-61. [153] C. Berggren, B. Bjarnason, G. Johansson, Capacitive Biosensors, Electroanalysis, 13 (2001) 173-180. [154] M.D. Porter, T.B. Bright, D.L. Allara, C.E. Chidsey, Spontaneously organized molecular assemblies. 4. Structural characterization of n-alkyl thiol monolayers on gold by optical ellipsometry, infrared spectroscopy, and electrochemistry, Journal of the American Chemical Society, 109 (1987) 3559-3568. [155] J. Lin, J. Szymanski, P.C. Searson, K. Hristova, Electrically addressable, biologically relevant surface-supported bilayers, Langmuir, 26 (2010) 12054-12059. [156] L. Romaner, G. Heimel, C. Ambrosch-Draxl, E. Zojer, The Dielectric Constant of SelfAssembled Monolayers, Advanced Functional Materials, 18 (2008) 3999-4006. [157] F.C. Bedatty Fernandes, A.V. Patil, P.R. Bueno, J.J. Davis, Optimized Diagnostic Assays Based on Redox Tagged Bioreceptive Interfaces, Analytical chemistry, 87 (2015) 12137-12144. [158] R.B. Schoch, J. Han, P. Renaud, Transport phenomena in nanofluidics, Reviews of modern physics, 80 (2008) 839. [159] R.P. Janek, W.R. Fawcett, A. Ulman, Impedance spectroscopy of self-assembled monolayers on Au (111): evidence for complex double-layer structure in aqueous NaClO4 at the potential of zero charge, The Journal of Physical Chemistry B, 101 (1997) 8550-8558. [160] J. Kang, J. Wen, S.H. Jayaram, A. Yu, X. Wang, Development of an equivalent circuit model for electrochemical double layer capacitors (EDLCs) with distinct electrolytes, Electrochimica Acta, 115 (2014) 587-598. [161] M. Ramanathan, M. Patil, R. Epur, Y. Yun, V. Shanov, M. Schulz, W.R. Heineman, M.K. Datta, P.N. Kumta, Gold-coated carbon nanotube electrode arrays: Immunosensors for impedimetric detection of bone biomarkers, Biosensors and Bioelectronics, 77 (2016) 580-588. [162] X. Li, J. Guo, Q. Zhai, J. Xia, G. Yi, Ultrasensitive electrochemical biosensor for specific detection of DNA based on molecular beacon mediated circular strand displacement polymerization and hyperbranched rolling circle amplification, Analytica chimica acta, 934 (2016) 52-58. [163] N. Wang, M. Lin, H. Dai, H. Ma, Functionalized gold nanoparticles/reduced graphene oxide nanocomposites for ultrasensitive electrochemical sensing of mercury ions based on thymine–mercury–thymine structure, Biosensors and Bioelectronics, 79 (2016) 320-326. [164] I.I. Suni, Impedance methods for electrochemical sensors using nanomaterials, TrAC Trends in Analytical Chemistry, 27 (2008) 604-611. [165] S. Kurbanoglu, S.A. Ozkan, A. Merkoçi, Nanomaterials-based enzyme electrochemical biosensors operating through inhibition for biosensing applications, Biosensors and Bioelectronics, 89 (2017) 886-898.
43
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[166] A. Alam, C. Wan, T. McNally, Surface amination of carbon nanoparticles for modification of epoxy resins: plasma-treatment vs. wet-chemistry approach, European Polymer Journal, 87 (2016) 422-448. [167] M. Pacios, M. Del Valle, J. Bartroli, M. Esplandiu, Electrochemical behavior of rigid carbon nanotube composite electrodes, Journal of Electroanalytical Chemistry, 619 (2008) 117-124. [168] F. Gutierrez, F.N. Comba, A. Gasnier, A. Gutierrez, L. Galicia, C. Parrado, M.D. Rubianes, G.A. Rivas, Graphene paste electrode: analytical applications for the quantification of dopamine, phenolic compounds and ethanol, Electroanalysis, 26 (2014) 1694-1701. [169] A. Afkhami, H. Khoshsafar, H. Bagheri, T. Madrakian, Construction of a carbon ionic liquid paste electrode based on multi-walled carbon nanotubes-synthesized Schiff base composite for trace electrochemical detection of cadmium, Materials Science and Engineering: C, 35 (2014) 8-14. [170] S. Sharma, S. Hussain, S. Singh, S. Islam, MWCNT-conducting polymer composite based ammonia gas sensors: A new approach for complete recovery process, Sensors and Actuators B: Chemical, 194 (2014) 213-219. [171] S.K. Vashist, D. Zheng, K. Al-Rubeaan, J.H. Luong, F.-S. Sheu, Advances in carbon nanotube based electrochemical sensors for bioanalytical applications, Biotechnology advances, 29 (2011) 169-188. [172] P.P. Joshi, S.A. Merchant, Y. Wang, D.W. Schmidtke, Amperometric biosensors based on redox polymer-carbon nanotube-enzyme composites, Analytical chemistry, 77 (2005) 31833188. [173] G.A. Zelada-Guillén, A. Tweed-Kent, M. Niemann, H.U. Göringer, J. Riu, F.X. Rius, Ultrasensitive and real-time detection of proteins in blood using a potentiometric carbonnanotube aptasensor, Biosensors and Bioelectronics, 41 (2013) 366-371. [174] R. Olivé-Monllau, M.J. Esplandiu, J. Bartrolí, M. Baeza, F. Céspedes, Strategies for the optimization of carbon nanotube/polymer ratio in composite materials: applications as voltammetric sensors, Sensors and Actuators B: Chemical, 146 (2010) 353-360. [175] J. Muñoz, F. Céspedes, M. Baeza, Modified multiwalled carbon nanotube/epoxy amperometric nanocomposite sensors with CuO nanoparticles for electrocatalytic detection of free chlorine, Microchemical Journal, 122 (2015) 189-196. [176] R. Olivé-Monllau, F.X. Muñoz-Pascual, E. Baldrich, Characterization and optimization of carbon nanotube electrodes produced by magnetic entrapment: Application to paracetamol detection, Sensors and Actuators B: Chemical, 185 (2013) 685-693. [177] R. Olivé-Monllau, A. Pereira, J. Bartrolí, M. Baeza, F. Céspedes, Highly sensitive CNT composite amperometric sensors integrated in an automated flow system for the determination of free chlorine in waters, Talanta, 81 (2010) 1593-1598. [178] R. Olivé-Monllau, C.S. Martínez-Cisneros, J. Bartrolí, M. Baeza, F. Céspedes, Integration of a sensitive carbon nanotube composite electrode in a ceramic microanalyzer for the amperometric determination of free chlorine, Sensors and Actuators B: Chemical, 151 (2011) 416-422. [179] J. Bastos-Arrieta, J. Muñoz, A. Stenbock-Fermor, M. Muñoz, D.N. Muraviev, F. Céspedes, L.A. Tsarkova, M. Baeza, Intermatrix Synthesis as a rapid, inexpensive and reproducible methodology for the in situ functionalization of nanostructured surfaces with quantum dots, Applied Surface Science, 368 (2016) 417-426. [180] J. Muñoz, J. Bastos-Arrieta, M. Muñoz, D. Muraviev, F. Céspedes, M. Baeza, CdS quantum dots as a scattering nanomaterial of carbon nanotubes in polymeric nanocomposite sensors for microelectrode array behavior, Journal of Materials Science, 51 (2016) 1610-1619. [181] F.Y. Castillo, R. Socher, B. Krause, R. Headrick, B.P. Grady, R. Prada-Silvy, P. Pötschke, Electrical, mechanical, and glass transition behavior of polycarbonate-based nanocomposites with different multi-walled carbon nanotubes, Polymer, 52 (2011) 3835-3845.
44
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AC C
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[182] J. Muñoz, F. Céspedes, M. Baeza, Effect of Carbon Nanotubes Purification on Electroanalytical Response of Near-Percolation Amperometric Nanocomposite Sensors, Journal of The Electrochemical Society, 162 (2015) B217-B224. [183] J. Bastos-Arrieta, J. Muñoz, N. Vigués, D.N. Muraviev, F. Céspedes, J. Mas, M. Baeza, M. Muñoz, Intermatrix synthesis of Ag, AgAu and Au nanoparticles by the galvanic replacement strategy for bactericidal and electrocatalytically active nanocomposites, New Journal of Chemistry, 40 (2016) 10344-10352. [184] J. Muñoz, J. Bastos-Arrieta, M. Muñoz, D. Muraviev, F. Céspedes, M. Baeza, Simple green routes for the customized preparation of sensitive carbon nanotubes/epoxy nanocomposite electrodes with functional metal nanoparticles, RSC Advances, 4 (2014) 44517-44524. [185] J. Muñoz, J. Bastos-Arrieta, M. Baeza, Carbon/polymer nanocomposite (bio)sensors containing nanoparticles: functionalization and applications, Polymer Nanocomposites: Synthesis, applications and research 2017, pp. 23-58. [186] J. Muñoz, L.J. Brennan, F. Céspedes, Y.K. Gun’ko, M. Baeza, Characterization protocol to improve the electroanalytical response of graphene–polymer nanocomposite sensors, Composites Science and Technology, 125 (2016) 71-79. [187] C.-W. Nan, Y. Shen, J. Ma, Physical properties of composites near percolation, Annual Review of Materials Research, 40 (2010) 131-151. [188] J. Muñoz, M. Riba-Moliner, L.J. Brennan, Y.K. Gun’ko, F. Céspedes, A. González-Campo, M. Baeza, Amperometric thyroxine sensor using a nanocomposite based on graphene modified with gold nanoparticles carrying a thiolated β-cyclodextrin, Microchimica Acta, 183 (2016) 1579-1589. [189] J. Chen, B. Yao, C. Li, G. Shi, An improved Hummers method for eco-friendly synthesis of graphene oxide, Carbon, 64 (2013) 225-229. [190] R. Montes, F. Céspedes, M. Baeza, Highly sensitive electrochemical immunosensor for IgG detection based on optimized rigid biocomposites, Biosensors and Bioelectronics, 78 (2016) 505-512.
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ACCEPTED MANUSCRIPT Figure captions Figure 1. A) Excitation sinusoidal signal applied on the system and the sinusoidal current registered shifted in phase. Nyquist plot for an electrochemical Faradaic system with B) kinetic-control process and C) mixed kinetic and diffusion-control process (inset: their corresponding equivalent Randles circuits). D)
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Nyquist plot for an electrochemical non-Faradaic system (inset: example of a complex capacitance plot, where it is indicated the relative minimum of the imaginary part of capacitance at which the real part of the capacitance is sampled, with its corresponding equivalent circuit).
Figure 2. A) General scheme of the SAM formation process of a modified-thiolated biomolecule on a
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gold electrode and their consequently biorecognition process. Schematics of different strategies to functionalize Au electrodes through thiolated biomolecule: B) SAM attachment followed by biomolecule
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recognition process; C) SAM attachment followed by its funtionalization and target recognition; and D) functionalization followed by SAM attachment and target recognition.
Figure 3. Scheme of different surface chemistry architectures for electrochemical biosensing. For the EIS architecture (A) the impedimetric analysis is used to promptly obtain ∆Rct as a function of target concentration (typical Nyquist plot). No apparent changes on the Nyquist capacitive diagram are expected
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due to the electrochemical energy storage is impossible so that the Faradaic capacitance (Cr) is absent. Otherwise, for ECS architecture (B), the analogous capacitive plot may be obtained since the molecular redox levels are tethered to the metal electrode surface and therefore, the electrochemical energy storage capability of this interface between the electrode and the molecular layer. Accordingly, the Cr values can
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Figure 4. Electrochemical values of different carbon/epoxy electrode compositions: ohmic resistance
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(RΩ, black, ■), charge transfer resistance (RCT, light gray, ■) and double-layer capacitance (Cdl, dark grey, ■), with their corresponding standard deviation for A) CNT/epoxy, B) rGO/epoxy and C) graphite/epoxy nanocomposite electrodes. EIS experiments were carried out using a 0.01 M [Fe(CN6)]3-/4- redox solution containing 0.1 M KCl (freq. range: 100 kHz–100 mHz; bias potential: +150 mV and AC amplitude: 5 mV).
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IgG
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Cyt c
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Linear range (nM) 3.8·10-6 – 7.8·10-4 2.12·10-3 – 2.12·103 1.0·10-5 – 1.0 1.0·10-5 – 5.0 1.0·10-3 – 1.0 1.0·10-6 – 1.0·10-2
Linear range (ng·mL-1)
Detection methods Amperometry Amperometry Amperometry SPR Fluorescence EIS EIS Detection methods – Biorecognition element SPR – Aptamer Fluorescence – Aptamer ICP-MS/TEM – Antibody EIS – Aptamer CV – Enzyme CV – Enzyme
2 – 2.5·103 5.1·102 – 30.17·103 30 – 1000 1.5·102 – 40·103 10 – 5·103 3·102 – 7·103 6.2 – 5.0·105
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DNA
Detection Method – Probe immobilization matrix DPV – Avidin-PANI/Pt DPV – Graphene/PANI/GCE DPV – PIA/eh-CdS/GCE DPV – AuNPs/WS2-graphene/GCE DPV – AuNPS/CuS-graphene/GCE EIS – PANI/rBi2S3/IL-CPE
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Table 1. Summary of the comparative analytical biosensing performance using different methodologies for detection.
Linear range (pM) 80 – 8.0·104 2 1.0·10 – 2.0·104 50 – 5.0·104 1.0·106 – 1.0·109 1.0·104 – 5.0·108
Detection limit (nM) 1.0·10-6 3.25·10-4 3.9·10-6 2.3·10-5 1.0·10-4 4.37·10-7 Detection limit (ng·mL-1) 30 180 25 150 10 300 0.57 Detection limit (pM) 50 266 1.5·10-3 63.2 5·105 1.0·104
Refs. [97] [98] [99] [100] [101] [82] Refs. [102] [103] [104] [105] [106] [107] [85] Refs. [108] [109] [110] [59] [111] [112]
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S. aureus
Visual detection – Tyramine single amplification Luminescent bioassay – MNPs/UCNPs Fluorescent bioassay – AuNPs/CdTe nanocrystals Potentiometric aptasensor / SWCNT Acid hybridization assay – Long life time biomarkers and two-probe tandem DNA hybrid Colorimetric – AuNPs/MNPs EIS – rGO/AuNPs
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** CV: Cyclic Voltammetry ***SPR: Surface Plasmon Resonance
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* DPV: Differential Pulse Voltammetry
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Detection methods – Materials used
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EIS – Aptamer-antibody
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[89] Refs. [113] [114]
8 50
[115] [116]
2
8·10
1.03·103 3
1.5·10 10
[117] [118] [88]
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EIS technique to optimize the filler /polymer ratio of nanocomposite (bio)sensor Impedance technique for the quantification of bioanalytes EIS technique as a characterization strategy to enhance the electroanalytical response Advanced electrochemical nanocomposite platforms using impedance measurements Electrochemistry biorecognition using EIS or ECS
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