Perspectives on and Precautions for the Uses of Electric

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Perspectives on and Precautions for the Uses of Electric Spectroscopic Methods in Label-free Biosensing Applications Beatriz Lucas Garrote, Adriano Santos, and Paulo Roberto Bueno ACS Sens., Just Accepted Manuscript • DOI: 10.1021/acssensors.9b01177 • Publication Date (Web): 09 Aug 2019 Downloaded from pubs.acs.org on August 10, 2019

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Perspectives on and Precautions for the Uses of Electric Spectroscopic Methods in Label-free Biosensing Applications Beatriz L. Garrote, Adriano Santos and Paulo R. Bueno* Institute of Chemistry, São Paulo State University (UNESP, Universidade Estadual Paulista), CP 355, 14800900, Araraquara, São Paulo, Brazil

*[email protected], tel: +55 16 3301 9642, fax: +55 16 3322 2308

ABSTRACT Label-free approaches to molecular diagnostic applications are appealing because of their inherent point-of-care advantages. Nonetheless, technical challenges impose a limit on use of these methods as will be discussed in the present paper. Electrochemical spectroscopic methods, such as impedance and impedance-derived methods, are highly effective in the development of label-free diagnostic assays, but they require careful control of the dynamics of the sensing interface. We herein report the strength and challenges of current methodologies associated with the applications of impedance and impedance-derived methods by focusing on their principles of operation. We demonstrate that the uses of their potentialities are not based on a know-how of these methods, but on how to combine the spectroscopic features with the required chemical design for the associated sensing interfaces. Predominantly, we illustrate how to use the resistive and capacitive terms of the interface to improve its sensitivity to the target. For instance, with the proper signal amplification strategy, limitations related to target-to-receptor size ratio can be overcome. The target-to-receptor ratio is one of the difficulties that we use as an example to illustrate how the sensing of an electric signal can be improved by controlling the properties of the interface at a nanometer scale.

Keywords: molecular assaying, electrochemical impedance spectroscopy, capacitive biosensors, electrochemical capacitance, charge transfer resistance, biosensors, label-free diagnostic, point-ofcare diagnostics.

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In general, electrochemistry is receiving considerable attention, and this is associated with the usefulness of the concepts existing therein, which include those pertinent to the development of important devices such as energy storage1,2, photocatalystic3 and biosensing4,5 devices. These applications have been focused on because they involve the daily use of a variety of mobile and portable electronic devices6,7 in which electrochemical components have been embedded, such as rechargeable batteries.8 Furthermore, though nowadays molecular diagnostics are constrained to clinical laboratories, in the near future, routine medical diagnostics could be implemented in an easy-to-use portable device9; indeed, this is already the case with regards to the monitoring of blood sugar levels using pinprick methodologies combined with electrochemical signals.10 Because of the success that has been achieved by glucose biosensors11, electrochemistry has been considered an obvious target platform for portable diagnostic devices.12 Accordingly, we herein intend to demonstrate that electrochemical impedance spectroscopy (EIS) can be a powerful and useful tool in helping to develop electrochemical biosensors13–17 for molecular diagnostics and healthcare applications. For the most part, we focus on label-free assays18 where impedance methods have great advantages over traditional transient methods. In general, label-based methods have high resolution, but their obvious disadvantage is their limitation with regards to achieving competitive and low-cost point-of-care applications when compared with label-free methods.18 Nonetheless, in a label-free platform, EIS needs to be used prudently in order to avoid false positive analyses which are associated with the dynamics of the interface and its inherently low signal-to-noise ratio. Control of the stability of the signal-to-noise ratio, which controls the dynamics of the interface, should be the first planned requirement rather than the detection of signal changes in the interface per se. The first requirement – and one which takes precedence over detecting signal changes in the interface – is to establish control over the signal-to-noise ratio such that it remains stable. This is because the signal-to-noise ratio controls, in turn, the dynamics of the interface. The control of the signal-to-noise ratio is associated with the fact that a detection of a signal may not always be related to a specific target binding process occurring at the sensing interface and is occasionally interpreted incorrectly as a positive signal (in spite of the control of a blank signal); rather than this, a detected signal could be solely a consequence of changes in the system’s state, induced, for instance, by the exposition of the sensing interface (comprising its structure inclusive of receptors) to an incubation environment. Therefore, in these particular cases, it is the change in the state of the external potential imposed on the interface, rather than the detection of a target per se, that is the cause of signal variations. The external potential (variation in the ionic content in a receptive interface during incubations, for instance) induces structural or charge state variations, which are due to a drift instead of a positive signal. Thus, the charge state is not necessarily correlated to the detection of a biological target species in a label-free platform. Therefore, judicious use of EIS in a label-free biosensing platform requires certain particularities to be taken into account, especially when EIS is applied to studying the signal response of nanoscale


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(lower than 10 nm) interfaces.6,7 In general, EIS at the nanoscale is both uncommon and underused by those electrochemists and electro-analysts who are not familiar with the physical principles on which EIS methods are based and their intrinsic dynamics. Researchers use EIS routinely owing to its power as a method, but EIS is not an indicated technique for these types of procedures, unless a deep knowledge of the stability of the system is well-established and optimized; furthermore, the steadiness of the interface to the assaying environment is highly recommended (though not strictly required – see below). EIS can provide an associated equivalent circuit analysis of the interface, which is connected to its physical and chemical properties. Basic principles of impedance spectroscopic methods. Though EIS offers an overview of the timedependent or frequency response of electrochemical interfaces, realistic analysis depends on a careful control of the dynamics of these interfaces. If the dynamics is properly controlled, EIS methods are very sensitive and signals in the spectroscopic mode can be resolved, which is not allowed for by transient methods such as amperometry, voltammetry, and so on. Actually, DC transient methods rely on a certain “unknown” proportionality (or relationship) between current density and voltage, i.e., 𝐽 ∝ 𝑉 and the function that relates 𝐽 to 𝑉 is generally inferred from a Ohm’s “law”19; phenomenologically, this means that models such as the Buttler–Volmer method are applied in explaining (empirically) the relationship between 𝐽 and 𝑉. On the other hand, EIS methods measure the function that establishes an association between a sinusoidal perturbation of current density and the potential19, that is, 𝐽(𝑡) = [1/𝑍 ∗ (𝜔)]𝑉(𝑡), where 𝑌 ∗ (𝜔) = 1/𝑍 ∗ (𝜔) is the admittance and 𝑍 ∗ (𝜔) is the impedance, and both are functions of frequency as required by the inherently spectroscopic characteristics of the method.19 Therefore, the measurement of 𝑍 ∗ at different frequencies, 𝜔, provides a spectrum of impedances, 𝑍 ∗ (𝜔). Furthermore, it makes available the function that relates 𝐽(𝑡) to 𝑉(𝑡). EIS approaches also allow for the separation of different contributions contained in 𝑍 ∗ (𝜔), such as those energetic contributions associated with resistive and capacitive phenomena (to which we can attribute suitable chemical meanings using macroscopic or nanoscopic physics). Therefore, 𝑍 ∗ (𝜔) is a function that provides the elements of a circuit (when this is the case) that shape a measured 𝐽 versus 𝑉 transient DC curve. This illustrates the superiority of AC to DC methodologies, i.e., they provide users with functions that govern the chemical properties of the interface. In this case, the function that relates 𝐽 to 𝑉 is not inferred; it is undeniably measured. The difficulties rely on providing a physical interpretation for this function. Nonetheless, this is not a problem related to EIS methodologies; rather, it requires certain extra-care and effort of their users. In the latter context, we demonstrated that different immittance functions20,21, which are functions obtained from an impedance spectrum as measured for different concentrations of the target, have different sensitivities. We also verified that immittance functions can be optimized for different interfaces aiming solely to analytical uses of the functions in diagnostic assays in which their meaning is dissociated from a physical meaning. Additionally, immittance methodologies allow for the setup of optimized single frequencies for each biological receptive interface and sensorial environment, which makes the analysis faster.21 Immittance analysis, when correctly


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used, is a powerful means of measuring and of calibrating electrochemical interfaces for multiplexable uses. Multiplexing is important, especially in a label-free platform, for simultaneous measurements of different targets. Following the analysis of molecular diagnostic assays using the immittance function, the alternative use of the impedance-derived capacitive method17,22–25 constitutes a specific (or merely a different) way of using the principles of immittance function analysis, but where the physical meaning of the obtained function is taken into account.6,7,17,22–25 Therefore, the effectiveness of the impedance-derived capacitive method is based on the physical meaning of capacitive components, which is established in terms of the first-principles of quantum mechanics26; we demonstrated that a redox capacitance of an interface can be used as a transducer signal and that capacitive interfaces can be designed in different ways21–25,27,28 in which the measured changes in capacitance is related to the changes in the energy of the interface.26 This involves not only the designing of different capacitive interfacial signals, but also involves a way of extracting and analyzing a particular capacitive element obtained from the impedance spectra of the sensing interface.21–25,27,28 For instance, changes in the electrochemical capacitance of the interface can be analyzed as a function of target concentration and an analytical curve can be constructed aiming to diagnostic applications.21–25,27,28 Signal dependence on the design of the interface and the redox probe arrangement. Conventionally, EIS is known to resolve different interfacial phenomena and to provide information of ion diffusion, electron transfer rate, redox and double layer capacitances19 of mainstream electrochemical interfaces. Because EIS is able to sense tiny chemical changes in the interface, it has been applied to the detection of biomarkers of great interest12,17,29 in suitable analytical ways with both high sensitivity and label-free characteristics.12,29,30 Nonetheless, there are different types of electrochemical interfaces, and the sensitivity of these interfaces depends on the transducer signal associated to the type of interface as shown in Figure 1. The simplest electrochemical interface is formed when a conductive electrode is placed in contact with an electrolyte as revealed in Figure 1a. Therefore, when a metallic foil (electrode) is placed in a liquid electrolyte, an equilibrium of charge (electronic and ionic) is established in that interface12 (Figure 1a), and this charge equilibrium can be resolved in details by EIS. An EIS study of this interface can be conducted in the absence (non-faradaic response) or in the presence of a redox probe (faradaic response). The non-faradaic mode (Figure 1a) corresponds to resistive and capacitive contributions that can be ideally modelled by a resistance combined in series with a capacitance; in this case, the resistance of the solution, 𝑅𝑠, is placed in series with the capacitance of the double layer, 𝐶𝑑𝑙. Realistically, the response of the interface is non-ideal, and heterogeneities can be considered by modelling 𝐶𝑑𝑙 by using a constant phase element.29 Figure 1 illustrates two types of faradaic modes. In the most common faradaic mode, a redox probe, e.g., [Fe(CN)6]3-/4- (Figure 1b), is added to an electrolyte that is in contact with the electrode; the addition of a redox probe provides extra circuit elements to the EIS response. These additional elements of the circuit are the charge transfer resistance, 𝑅𝑐𝑡, and the Warburg circuit element12,


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𝑍𝑤, which accounts, respectively, for the electron transfer (redox reaction) occurring between the redox probe and the electrode, and for a diffusional process associated with a limited-diffusional transport process due to the mass transport of the redox probe species from the bulk of the electrolyte to the interface of the electrode. The electric circuit that considers all these elements, in this faradaic mode of operation, is known as circuit of Randles (Figure 1b). Both faradaic and nonfaradaic modes reported above have been used in biosensing applications.12,31

Figure 1. Non-faradaic a) and faradaic b-c) processes occurring in an electrode-electrolyte modified interface and their representation in terms of equivalent electric circuit elements. The non-faradic situation is, in terms of equivalent circuit, modelled by considering a double layer capacitance, 𝐶𝑑𝑙, and the solution resistance, 𝑅𝑠. There are two faradaic possible schemes, that is, b) the faradaic with redox probe in solution that is modelled by the well-known Randles circuit, where 𝑅𝑐𝑡 is the charge transfer resistance and 𝑍𝑤 is the Warburg element, and c) the faradaic situation in which the redox probe is covalently attached to the interface. The latter is modelled by a “revised” Randles circuit, where 𝐶𝑟 is the redox capacitance replacing the 𝑍𝑤 term as the limiting kinetic control term. The insets in each of the schemes are examples of impedance and capacitive Nyquist plot responses expected in EIS measurements.

Another type of faradaic mode that is less known but quite functional in label-free assays15,27,28,32 is when a redox probe is covalently attached to the electrode. In this case, in place of a Warburg diffusional element, there is another circuit element (Figure 1c) that accounts for the charging of the redox centers and how they communicate with electrode states.33 This is a type of pseudo-capacitive element that is fundamentally better whether stated as an electrochemical capacitance.33 In Figure 1c, this is referred to as 𝐶𝑟 (redox capacitance), which is a specific type of


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electrochemical capacitive phenomena, where the capacitance is strictly associated with a redox reaction.33

Figure 2. Representation of a) non-faradaic and b-c) faradaic interfaces for electrochemical biosensing applications. The interface consists of a mixed self-assembled monolayer (SAM) containing 11-mercaptoundecanoic acid, MUA, and 6mercaptohexanol, 6C-OH, or a SAM constituted of a designed redox-active peptide. The receptive centers (biological receptor, specific target of interest, and the block agent, which prevent unspecific interaction) are coupled over these SAMs. Modifications in the a) non-faradaic interface are monitored by changes in the double layer capacitance (𝐶𝑑𝑙), whereas modifications in a b) faradaic configuration with a redox probe in solution are monitored by changes in charge transfer resistance (𝑅𝑐𝑡); finally in the case of c) faradaic interface where the redox probe is confined on the electrode surface, modifications of the interface are monitored by changes in the redox capacitance (𝐶𝑟), which depends on changes in the redox density-of-states associated with redox moieties and 𝐶𝑟 usually decreases after the binding of the target in the interface.

Coming back to the situation described in Figure 1b, where the interface of the electrode is modified with biological receptors and the 𝑅𝑐𝑡 of this interface is measured by EIS methods12,29, we will observe that an electrochemical assay can be constructed because of the impedimetric response of the interface, which enables establishment of a sensible interface by the recruitment of targets by the receptor that cause changes in the 𝑅𝑐𝑡 of the interface. Hence, the impedimetric term arises because the interface offers (upon the occupation of the receptive centers) an electronic impediment for charges to be transferred from the electrode to the redox probe (contained in the electrolyte) and vice-versa. The electric impediment (measured as changes in 𝑅𝑐𝑡) for an exchange of charge between the electrode and electrolyte increases as a function of target binding to the interface. This impedimetric characteristic of the interface has been a common adopted strategy to attain label-free biosensing interfaces (see Figure 2b).12 On the other hand, the capacitive sensing of the interface (Figure 2c) is based on monitoring the changes in the interface by using an electrochemical capacitive signal associated with 𝐶𝑟 (Figure 2c). Thus, the changes are due to variations in the redox activity of the interface upon the binding of a target to receptive centers. Ultimately, it is the energy of the interface associated with 1/𝐶𝑟 that is effectively monitored33 as a function of a target binding to the interface. In this case, there is no redox probe in the electrolyte. The entire electrochemical signal is self-contained in the interface (redox states are confined in a molecular length that is less than 3 nm in thickness), demonstrating that the


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chemical design of the interface (with or without redox probe) determines the origin of the transducer signal.34 Among the types of non-faradaic and faradaic sensing interfaces (as shown in Figure 1) studied by EIS13–17 or derived-EIS15,27,28,32 methods in a label-free format are those constructed by modifying gold electrodes (to attach a receptive biological center) with a thiolated self-assembled monolayer (SAM)35, as illustrated in Figure 2. The biological recognition is achieved by using biological receptors attached to these SAMs, which are specific to detect a given target of interest such as antibodies, enzymes, nucleic acids, etc. What are the limitations of impedance or impedance-derived label-free methods? Although faradaic-based EIS methods present unquestionable features for biosensing applications13–17, there are concerns to be taken into account that can compromise the functioning of the devices. For instance, these are problems that can not only affect reproducibility, sensitivity and specificity, but also invalidate the analysis. In the present text, we discuss and illustrate important issues associated with the quality of the experiments that we believe can be helpful to those interested in using EIS methods on a daily basis and in developing molecular diagnostic assays for real samples. It is important to start by mentioning that impedance spectroscopy, as a technique or methodology, must be interpreted or used only if its principles apply to the system of interest. In the present case, our system of interest is a modified interface that is aimed to be used as a biological sensor. The transducer of the sensor works on the principle that the response of a biologically receptive interface can be converted in a biological binding event associated with a biomarker of interest into an easy interpretable analytical signal. In using the resistive (Figure 1b) or capacitive (Figure 1c) characteristics of the interface as the transducer signal, the impedance of the interface must be measured so that the resistance or capacitance can be obtained from the impedance spectra. To reliably measure the frequency response of a sensing interface, we need to assume that a quantitative measurement of the output signal of the interface (electric current, for instance) as induced by a stimulus or perturbation (potential, for instance) can be used to quantify the steady-state dynamics of the interface. Therefore, to generate an impedance spectrum, we need to measure the magnitude and the phase of the output (with respect to the perturbation) as a function of the frequency. To be able to do so, the biosensing interface must possess a linear response with respect to the perturbation and be time-invariant during the acquisition of the spectrum. These are the two main requisites for using impedance spectroscopic methods in electrochemical platforms. In other words, a system is considered linear if, by doubling the amplitude of the input, the output signal is equivalently doubled. The system is a time-invariant regime if spectra do not vary in time for a given steadystate. The requirement for linearity is easier to achieve in electrochemical interfaces than that for time-invariances. The linearity is generally achieved by how the users set the magnitude of the perturbation (small as possible is better in obtaining the desired linearity, though the ratio of signalto-noise decreases). The time-invariance depends on the chemical stability at a particular steadystate, which is inherently difficult to control in a heterogeneous interface within a nanometer scale.


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In the case of biosensing applications, the time-invariance of assays depends on the invariance of the interface under the conditions of different incubation steps. For instance, some designed interfaces, when incubated in buffered saline environments, presents a 𝑅𝑐𝑡 drift36 owing to the variations in steady-state; the incubation acts as an external potential environment with a chemical potential that induces changes in chemical states of the interface during the incubation. This is the situation where time-invariance is not attained. In this circumstance, obviously, changes in the impedance are not associated with the biological binding event per se; they are a result of a variance associated with the reactivity of the interface to the environment of incubation, which sometimes is mistakenly interpreted as an useful signal response of 𝑅𝑐𝑡 to the biological binding event (this can, for instance, artificially enhance sensitivities). Provided that an interface is optimized, EIS assays are effective, but the optimization and control of the interface to enable sensing can be obviously laborious. This is certainly more challenge when the aiming is to use complex samples. Owing to the fact that real biological samples are composed by a pull of different proteins, ions and interferent molecules, the obtention of a specific and reproducible response of the assay is challenging. Nonetheless, the control of a biosensing interface is possible by using nonfouling surfaces which allows to prevent non-specific surface protein adsorption.5 Several EIS biosensors providing specific detection in real samples are described in the literature. For instance, Bryan and co-workers37 developed a faradaic EIS biosensor for the detection of α-synuclein in real serum samples aiming to the diagnostic of Parkinson’s disease. They achieved high specificity by using a thiolate polyethylene glycol (PEG) self-assembled monolayer, which provided a useful antifouling characteristic of the surface, enabling the specific detection of α-synuclein in complex biological samples. Similarly, Luo et al.38 designed a specific non-faradaic EIS biosensor for the detection of insulin in real blood serum. They increased the specificity of the assay by using a chemisorbed zwitterionic polymer film with additional anti-fouling characteristics in the place of a PEGylated surface. Additionally, it is important to mention the fabrication steps that affect the quality of the interfaces. For instance, the quality control of the SAMs is primordial to achieving suitable performances of resistive (impedimetric) or capacitive-based biosensors. It is known that wellpacked thiolated SAMs can prevent redox probe (cyanide) etching. This is because well-formed SAMs are both well-ordered and structurally compact;39 however, to obtain a SAM with an appropriate crystallinity is a challenging task. Thiol purity, incubation time, surface roughness of the electrode, and the length of the thiolate precursors, which control the thickness of the molecular film, influence the self-assembling of monolayers over the electrode and consequently the SAM’s properties. The presence of defects or pinholes can also compromise the stability of the assays, while enhancing cyanide etching.40 Using the 𝑅𝑐𝑡 signal of the interface, it can be demonstrated that pinholes reduce the insulating properties of the monolayers, allowing sometimes for an unrestricted electronic and ionic migration between the electrolyte and the electrode41,42, which is sensed (especially in DC methods) as an unblocking effect that is a percolation of the blocking effect to the probe redox with significant decreases in 𝑅𝑐𝑡 values. Consequently, the target recognition event (e.g., antibody-antigen coupling) could not appropriately change 𝑅𝑐𝑡 values because of the


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percolative effect and the interaction of the target in the interface is not detectable by measurements in the 𝑅𝑐𝑡. In the present study, we illustrate a case of a drift where an assay using a drifting interface leads to unreliable sensitivities and cases of stable interfaces which are though unresponsive owing to the fact that a small target was aimed to be sensed with an inappropriate control of the target-toreceptor ratio. As a proof of concept, the small target used was interleukin-6 (IL-6 protein). The latter case was studied in two different faradaic situations, i.e., resistive and capacitive platforms. Both approaches were insensitive. Further, on a nanometer capacitive designed interface, which deliberatively amplifies the capacitive signal, we were finally able to detect the small IL-6 protein, demonstrating not only that limitations associated with detecting sensitively the binding of lowerweighted targets can be overcome but also that the magnitude of a transducer signal can be chemically planned. Within this sequence of intentional events and by exemplifying EIS interfacial responses, the purpose of the present paper is to a) illustrate as much as possible the types of faradaic impedance-derived prospects that are reported in the literature, b) introduce simple ESI platforms and how they can be used for the designing of label-free assays, and c) illustrate how electrochemical interfaces can be chemically and appropriately designed to achieve suitable EISbased responsive and reliable assays. EXPERIMENTAL PROCEDURE Chemical reagents and proteins. Sodium hydroxide, potassium hydroxide, 1-Ethyl-3-(3dimethylaminopropyl) carbodiimide (EDC), N-Hydrosuccinimide (NHS), sodium chloride, trisodium citrate dihydrate, monopotassium phosphate, sodium phosphate dibasic dodecahydrate, potassium chloride, sodium carbonate, sodium bicarbonate, sodium nitrate, potassium ferricyanide, potassium ferrocyanide, perchlorate tetrabutylammonium (TBA-ClO4), 11-mercaptoundecanoic acid (MUA), 6-mercapto-1-hexanol (6-COH), BSA (Bovine Serum Albumin), human C-reactive protein (CRP), and human CRP polyclonal antibody (Ab) were purchased from Sigma-Aldrich. Antibody and IL-6 protein were purchased from Rhea Biotech (Brazil). The recombinant dengue virus non-structural 1 (NS1) protein and monoclonal antibody (ab64456) were purchased from Abcam. The redox peptide used as support of some for the receptors was Fc-Glu-Ala-Ala-Cys, manually synthesized by solid phase peptide synthesis using Fmoc protocols on rink amide resin (0.48 mmol g-1). Coupling was performed at a 2-fold molar excess relative to the amino component in the resin, using diisopropylcarbodiimide (Dic)/ 1-hydroxybenzotriazole (HOBt). Fmoc groups were deprotected using 20% 4-methylpiperidine/dimethylformamide (DMF) for 80 min. The ferrocene redox probe was introduced at the N-terminus by reaction with one molar equivalent 3ferrocenylpropionic anhydride in 5 mL DCM/DMF (1:1) for 24 h. Peptide cleavage was performed by removing side chain protecting groups with 94% trifluoroacetic acid (TFA), 2.5% 1,2-ethanedithiol, 2.5% H2O and 1% triisopropylsilane for 2 h. The peptide was then precipitated with diethyl ether and separated from the reaction solution by centrifugation. Further details of the synthesis are given by Piccoli et al.28


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The solutions used were as follows: 12 mM Phosphate buffer (PB); 1 mM K3[Fe(CN)6] , 1 mM K4[Fe(CN)6] · 3H2O in 0.5 M KNO3 in 12 mM PB pH 7.4 (for simplicity throughout this text we will refer to this electrolyte as 1mM/1mM [Fe(CN)6]3-/4-); carbonate/bicarbonate buffer pH 9 (1.59 g L-1 sodium carbonate, 2.39 g L-1 sodium bicarbonate); PBS-T (8 g L-1 sodium chloride; 0.2 g L-1 monopotassium phosphate; 1.15 g L-1 sodium phosphate dibasic dodecahydrate; 0.2 g L-1 potassium chloride; 0.2 g L-1 sodium nitrate; 0.5% tween 20). All the solutions were prepared using ultrapure water from Milli-Q system (18.2 MΩ.cm at 25°C). Electrode pre-treatment. Gold electrodes were cleaned by mechanical polishing with aluminum oxide pads (1 µm, 0.3 µm and 0.05 µm) for 3 min each. After being polished, the electrodes were sonicated for 2 min in ethanol. Then, cyclic voltammetry (CV) was performed in 0.5 M NaOH (100 cycles from -1.7 to -0.7 V versus Ag|AgCl at 0.1 V s-1). Posteriorly, the electrodes were immersed for 30 min in ethanol with agitation. Finally, the electrodes were electrochemically polished by CV in 0.5 M H2SO4 (50 cycles from -0.2 to 1.5 V versus Ag|AgCl at 0.1 V s-1). The area of the reduction peak obtained in the sulphuric acid CV was used to calculate the electro-active area of the polished gold surface by using a conversion factor of 410 µC cm-2.43 The ratio of the electro-active area to the geometric electrode area (0.03142 cm-2) corresponded to the electrochemical surface roughness. The surface roughness was maintained below 1.5 to ensure reproducible and stable molecular film properties. Electrochemical measurements. Electrochemical measurements were performed in an AUTOLAB potentiostat (from METROHM), equipped with a frequency response analysis (FRA) module using a three-electrode system: a gold disk working electrode (2.0 mm diameter from METROHM, pretreated as described above), a platinum plate counter electrode, and a silver/silver chloride reference electrode (Ag|AgCl filled with 3.0 M KCl). All potentials reported are relative to this reference electrode. CV in resistive faradaic assays (with redox probe in solution) were measured from -0.2 to 0.7 V versus Ag|AgCl at a scan rate of 0.1 V s-1 in 0.5 M KNO3 in 12 mM PB pH 7.4 containing 1mM/1mM [Fe(CN)6]3-/4- to determine the formal potential of the redox probe in solution. CV in capacitive faradaic assays (redox probe covalently assembled within the monolayer) were measured from -0.0 to 0.7 V versus Ag|AgCl at a scan rate of 0.1 V s-1 in 20 mM of TBA-ClO4 dissolved in acetonitrile and water (1:4 v/v) as a supporting electrolyte to determine the formal potential of the electroactive SAM. EIS data for both configurations (resistive and capacitive) were acquired at the formal potential calculated by the CV measurements (c.a. 0.25 V versus Ag|AgCl in the resistive configuration and c.a. 0.36 V versus Ag|AgCl in the capacitive configuration) in a frequency range from 1 MHz to 0.1 Hz with a 20 mV amplitude (peak-to-peak) sinusoidal perturbation. The gold surface modifications in the resistive assay were monitored by the charge transfer resistance (𝑅𝑐𝑡), calculated by fitting a variation in the Randles equivalent circuit (Figure 1b) in which the double layer capacitance (𝐶𝑑𝑙) was replaced by a constant-phase element (𝑄).40 In the capacitive assay, modifications of the gold electrode surface were monitored by the inverse of the redox capacitance (1/𝐶𝑟) calculated by fitting a variation in the Randles equivalent circuit (Figure 1c).


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Biosensing interface design. The SAM was formed on the gold surfaces by immersing freshly cleaned gold electrodes in thiols solution (for the resistive assays, the SAM consisted of 0.05 mM of 11-mercaptoundecanoic acid (MUA) and 1 mM of 6-mercapto-1-hexanol (6-COH) in ethanol, while for the capacitive assays it was formed by adding 2 mM redox peptide diluted to acetonitrile and water [1:1 (v/v)]) and allowing the solution to incubate for 16 h at room temperature. After the chemical modification, the electrodes were briefly rinsed in ethanol or acetonitrile/water and ultrapure water and dried in nitrogen. The modified electrodes were immersed for 30 min in 200 µL of an aqueous solution containing 0.4 M EDC and 0.1 M NHS (v/v) at room temperature to activate the carboxylic groups of the SAM molecules. After that, the electrodes were incubated for 1 h at room temperature in an antibody solution (1 µM) in PB pH 7.4. Then, the electrodes were briefly rinsed in PB, ultrapure water, and dried with nitrogen. The unbound carboxylic groups were blocked by a 30 min incubation in 0.1% BSA solution in PB pH 7.4. Then, the electrodes were rinsed in PB, ultrapure water, and dried with nitrogen. Each modification was characterized by CV and EIS measurements. The stability of the biosensing interface was tested by immersion for 30 min in PB pH 7.4 and then measuring its electrochemical impedance (EIS measurement). This process was repeated three times in each assay to calculate the drift of each electrode. Those with a drift lower than 5% were considered stable and ready to use. To test the assay with the specific target, solutions were prepared in increasing concentrations in PB pH 7.4 and dropped on the modified gold electrode surfaces, where they were incubated for 30 min. Then, the electrodes were briefly rinsed in PB, ultrapure water, and dried with nitrogen. For the negative control, fetuin was used. Fetuin is an abundant serum protein found in mammals. It is a glycoprotein member of the cysteine protease inhibitors that is expressed during embryogenesis in multiple tissues (it is considered the major protein during the fetal period) and secreted into the blood by an adult liver.44 In normal serum, its concentration is around45 450–600 µg mL-1. 𝑅𝑐𝑡 or 1/ 𝐶𝑟 variations between measurements were presented as a relative response in percentage (RR%) calculated as 𝑅𝑅(%) = [(𝑋[𝑡𝑎𝑟𝑔𝑒𝑡] ― 𝑋𝑛)/𝑋𝑛]·100,

(1)

wherein 𝑋[𝑡𝑎𝑟𝑔𝑒𝑡] is the variable used to monitored the variation of 𝑅𝑐𝑡 or 1/𝐶𝑟 for a given target concentration and 𝑛 refer to the response of the blank (which is the sample in which the biomarker is absent). IL-6, the main target of interest in this study, is a small 26 kDa glycoprotein with an important role in the immune process and is involved in homeostatic and neuroendocrine functions. The overexpression of IL-6 is associated with cardiovascular disease, osteoporosis, diabetes, arthritis and tumors. Increasing interest has focused on the association of IL-6 with cardiovascular disease.46 Here, IL-6 was used as a proof of concept, in order to analyze the influence of the molecular weight of the biological receptor and the target on the biosensor response. Nonetheless, the use and the test of this important receptor in a label-free platform in real patient samples is still to be demonstrated.


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RESULTS AND DISCUSSION Case of 𝑹𝒄𝒕 drifting and evidence of the importance of controlling ionic “strength” effects. A critical aspect to be controlled in an electrochemical interface that is aimed at label-free assay applications is the possibility of buffer-induced EIS drift. It has been demonstrated36,47,48 that the type of electrolyte and incubation in PBS affects the stability of assays, generally inducing false sensitivities. The stability of SAMs in different buffers is, obviously, dependent on the chemistry and defects contained in the SAMs, as was discussed in the introduction. Based on a previous assessment, well-formed SAMs based on mixed thiolate structures, such as 11mercaptoundecanoic acid (MUA) and 6-mercapto-1-hexanol, drift after immersions in PBS (pH 7.4). Though there are different causes for the EIS drift during an assay, one of the sources is the presence of chloride ions, which are present in PBS. Under a drift condition, false analytical curves can be obtained using 𝑅𝑐𝑡 variations, as shown in Figure 3, which are merely associated with the immersions of SAM functionalized electrodes in PBS (incubation steps), as the PBS is allowed to “react” with the SAM. Therefore, each of the incubation steps increases 𝑅𝑐𝑡 and the increase is also dependent on the duration of the incubation (there is no time-invariance). The increase in 𝑅𝑐𝑡 can thus be associated with the incubation’s paces and non-biological drifts during the assays. If this is combined with some sort of specific adsorption, it causes a false enhancement in the sensitivity of the assay that is not only dependent on the binding but also on the duration of the incubation. Such a type of drift impacts the sensitivity and cannot be solely corrected by measuring a blank signal and using it as a reference signal. The increase in the 𝑅𝑐𝑡 signal was approximately 50%, 75%, and 25% subsequently after the first, second, and third incubation in PBS (without any specific protein), respectively, as shown in Figure 3b. As mentioned previously, chloride ions are considered one of the ions responsible for drifting progressions owing to the fact that chloride ions are supposed to interact (or absorbs) strongly with gold surfaces. Because, when chloride ions penetrate SAMs to attach to the gold, they modify the interface. Therefore, a receptive interface constructed over certain types of SAMs on gold is expected to change upon each assay measurement (or incubation steps) and thus cause “artificial” changes in 𝑅𝑐𝑡; thus, presumably, only part of the changes in 𝑅𝑐𝑡 can not be associated with a recognition element-target coupling (i.e. increase in the insulating film feature due to target binding). This is a type of instability (and an example of invariance in time) that has been observed in other gold-based interfaces by our research group. These problems have been overcome by constructing suitable surface chemistry and by an appropriate engineering of the interface. For instance, by using our own designed monolayers based on peptides28 among others optimized interfaces15,27,32, we constructed reliable EIS-based label-free assays. Figure 3b confirms the effect of chloride ions on EIS data. Note that when EIS data is recorded in 1mM/1mM [Fe(CN)6]3-/4-in KNO3 (a chloride ion free electrolyte) with an incubation in PB (pH 7.4), the drift process decreases and can thus be controlled to undamaged levels for the assaying process (though better SAM stabilities can be obtained in terms of SAM; please see below for the case of SAMs based on peptide designed structures).


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Figure 3. Analysis of the PBS-induced EIS drift associated with the presence of chloride ions. a) Impedimetric Nyquist plot showing the drift caused by incubations of the working electrode in PBS (the inset shows the Nyquist capacitive response). The measurements were made in the presence of 1mM/1mM [Fe(CN)6]3-/4- in PBS. b) Relative response in percentage (RR%) of 𝑅𝑐𝑡 [Eq. (1)] measured in b) in comparison with a relative variation in 𝑅𝑐𝑡 measured in 1mM/1mM [Fe(CN)6]3-/4- in KNO3 and incubations made in PB.

In a resistive faradaic platform, i.e., by using a reagent-dependent redox probe in solution to measure 𝑅𝑐𝑡 (see Figure 1b), it is obviously important to select a proper redox probe. For instance, several studies have been carried out in order to understand instabilities associated with the reactiveness of the redox probe with SAMs. Some studies confirmed that the [Fe(CN)6]3-/4- redox probe reacts with the surface of the electrode even when it is constituted with platinum instead of gold. It has been demonstrated that iron-cyanide based complexes can be converted into other compounds such as Prussian Blue. This conversion or reaction with the electrode is sensed by EIS as an increase in 𝑅𝑐𝑡.49,50,51 For instance, Lazar et al.48 studied the influence of cleaning protocols, electrolytes, and incubation times in buffer solutions containing [Fe(CN)6]3-/4- as the redox couple. They made continuous EIS measurements and utilized atomic force microscopy (AFM) and scanning electron microscopy (SEM).48 All measurements combined were very indicative of the reaction of [Fe(CN)6]3-/4- with the SAMs. In short, they confirmed that (𝑖) a cleaned and an appropriate electrode surface is decisive in achieving reproducible and stable EIS results, (𝑖𝑖) PBS is not a suitable electrolyte for measuring 𝑅𝑐𝑡 because of chloride ions as discussed above, (𝑖𝑖𝑖) the formation of cyanide ions (CN-) by decomposition of the redox probe complexes by UV irradiation can lead to surface deterioration, increasing the surface roughness and thus decreasing 𝑅𝑐𝑡, and (𝑖𝑣 ) cyanide etching effects, both in the presence or in the absence of electrochemical measurements (or of the driving forces associated with imposed voltages), were demonstrated as an issue to be taken into account. The interaction of cyanide ions with gold was also reported by Vogt et al.47 They measured 𝑅𝑐𝑡 of a dsDNA SAM every 30 min over 60 h in 1 mM ferri/ferrocyanide and observed that 𝑅𝑐𝑡 decreased by about 40%, which was associated with the generation of CN- during EIS measurements. The latter process was attributed to the etching of the gold surface and to the removal of adsorbed molecules.


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Sensing by using redox probe in solution (faradaic resistive assay). In this section we demonstrate the case of a faradaic impedance study (faradaic resistive assay) with a standard [Fe(CN)6]3-/4- redox probe contained in the electrolyte wherein the 𝑅𝑐𝑡 is measured. Hence, 𝑅𝑐𝑡 variations were used as the transducer signal to detect the IL-6 protein. The interface of the sensor was designed by modifying a gold electrode with a mixed thiolated SAM, i.e., by using 11-mercaptoundecanoic acid (MUA) and 6-mercapto-1-hexanol, following a similar mixed SAM structure as was discussed in the previous section. After the formation of the SAM, a specific IL-6 antibody was attached to the MUA component of the SAM by using the standard carbodiimide EDC/NHS activation process of the chemical groups of the SAM. Figure 4 shows the obtained CV curves (4a) and the impedimetric Nyquist plots (4b) for each step involved in the construction of the interface. Therefore, Figure 4 shows the responses of CV and EIS for the bare gold surface (black) after the assembling of the SAM (red) over this gold surface, the coupling of the antibody (blue) to the SAM formed over gold, and, finally, the blocking of the SAM interface with a non-specific protein (yellow). The goal of the latter step is to eliminate the remaining unreacted and non-specific sites to specific IL-6 antibody. The presence of unreacted sites can increase the non-specificity of the surface during assays and thus is undesirable. It is well-known that SAMs have an electronic insulating effect when immobilized onto a gold surface.52 Consequently, as expected, it can be observed in Figure 4b that 𝑅𝑐𝑡 increases and the current density peaks decrease (related with [Fe(CN)6]3-/4- redox probe) when compared with bare gold (Figures 4a and 4b), which presents a relative small 𝑅𝑐𝑡 and higher reduction/oxidation current peaks. The biological receptor (antibody) is coupled to the SAM by using the standard carbodiimide EDC/NHS protocol.53 The EDC/NHS reaction yields an ester that eliminates the negative charges on the SAM (i.e., deprotonated carboxylic groups at pH 7.4) with a consequent decrease in the insulating effect due electrostatic attraction between SAM and negative redox ions. The charge variation increased the current density as shown in Figure 4a and it decreased the 𝑅𝑐𝑡 value by 70%, as shown in Figure 4b, when compared to those values of the SAM before the EDC/NHS protocol is used. This is a clear indication of how impedance is sensitive to small changes in the electronic structure of monolayers that represent large changes in the charge density of the monolayer. Because of the atomic or molecular scale, the charge densities can change significantly due to a single change in the charge component of the monolayer. A variation of 70% is a quite significant in comparison to those caused by an antibody-target interaction in which 𝑅𝑐𝑡 varies approximately in the order of 5 to 10%.14,20 Both the antibody and the immobilization of a block agent increases 𝑅𝑐𝑡 by inducing an electric insulation of the electrode that, in turn, prevents charge transfer. The physical immobilization of the antibody was confirmed (in the experiments associated with Figure 4) through a QCM analysis. The chemical adsorption of the molecules over the gold electrode surface is shown in the Supplementary Material (Figure SM1 and Table SM1).


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Figure 4. Electrochemical responses of each step of modification of gold electrodes to form a biosensing impedimetric interface. The measurements were acquired in a 0.5 M KNO3 electrolyte containing 1mM/1mM [Fe(CN)6]3-/4- buffered in 12 mM PB with a pH of 7.4. a) Cyclic voltammetry patterns and b) impedance Nyquist plots during each of the fabricating steps.

The sensitiveness of a receptive biological interface to variations in 𝑅𝑐𝑡, as obtained in a faradaic resistive EIS configuration, depends on, quite intuitively, the size of the target. This can be intuitively inferred because one would imagine that the higher the target, the larger the electronic blocking effects caused in the interface. Thus, the higher size of the target, the higher the increase in 𝑅𝑐𝑡 values. Nonetheless, this is not supported in a straightforward way experimentally because this also increases 𝑅𝑐𝑡 depending on the length of the conductive layer and how the resistive layer and its length are combined or entangled with the receptive layer. It is graphically represented in Figure 5.


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Figure 5. Structural conformation of a faradaic impedimetric-based biosensing interface to illustrate the influence of the target-to-receptor size ratio on the sensitivity of the assay: a) Configuration based on a high size target and receptor that allows for prediction of the assay’s intermediate sensitivity; b) best expected sensitivity based on a small receptor and a large target; and c) worst based sensitivity based on a large receptor and small target.

The structural conformation of the IL-6 assay presented here follows that shown shown in Figure 5c. It consisted of a high molecular weight receptor, i.e., the IL-6 antibody – 150 kDa, and a small size target, i.e., IL-6 – 26 kDa. As shown in Figure 6a, when the receptive platform was incubated with two concentrations of IL-6, 38.5 nM and 385 nM (1 and 10 µg mL-1), the differential response observed with the negative control and baseline 𝑅𝑐𝑡 values was not observed (all the variations were < 5%). The insensitiveness was presumably caused by the low target-to-receptor size ratio that negatively affects the sensitiveness, as previously reported in faradaic assays based on a redox probe confined in the electrode surface (Figure 6b).


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Figure 6. Result of the a) resistive and b) capacitive assays for IL-6 detection. In a) a EIS Nyquist plot was recorded in 1mM/1mM [Fe(CN)6]3-/4- in 0.5 M KNO3 and 12 mM PBS at a pH of 7.4. C(-) was the negative control (fetuin 10 µg mL-1). In b) an ECS Nyquist plot was recorded in 20 mM TBA-ClO4 in acetonitrile/water (1:4, v/v). c) Comparison of the IL-6 analytical curves in resistive (based on 𝑅𝑐𝑡) and capacitive (based on 1/𝐶𝑟) assays with the analytical curves of NS1 (resistive assay) and CRP (capacitive assay) showing the influence of the target-to-receptor size ratio in the sensitiveness of both types of configurations. The target-to-receptor size ratios of the NS1 and CRP assays were 0.30 and 0.79, respectively. The analytical curves were obtained from the relative response in percentage (RR%) of either 𝑅𝑐𝑡 or 1/𝐶𝑟 for each target concentration [see Eq. (1)].

Sensing by using redox probe moieties confined in monolayer (faradaic capacitive assay). In a faradaic capacitance assay platform, the assumption of a target-to-receptor size ratio becomes evident. It was demonstrated, after several consistent experimental studies, that the sensitiveness of the receptive platform depends on the molecular weight of the specific target. IL-6 was purposely used here because it is difficult to detect owing to its small size. IL-6 is a 26 kDa protein whose weight is 4.5 times lower than that of the C-Reactive protein (CRP)25,28 and 1.8 times lower than that of the non-structural protein 1 (NS1)14,27, both of which are proteins previously detected by resistive or capacitive assays. The influence of the size of the target on the electrochemical capacitive sensitiveness was recently reported.34 This influence was demonstrated in the context of redox-active SAMs and allowed to access, as discussed in the introduction, the 𝐶𝑟 term34 of the equivalent circuit of the interface for different situations of target-to-receptor size ratios. A clear relationship between the target-to-receptor size ratio and the sensitiveness of the assay was observed. The relationship was explained based on the fact that the capacitive response of the interface, which is associated with the redox activity, is dependent on a length scale that is associated with a type of electric fieldeffect that corresponds to a phenomenon (Tomas Fermi screening) distinct from that provided by Debye length, a screening that is the mainstream analysis associated with the double layer phenomena. Thus, the length scale associated with a redox capacitive signal is defined as 𝜅 ―1 = 1/2 (𝜀𝑟𝜀0 𝐶 ) , where 𝜀 is the relative permittivity of the dielectric environment and 𝜀 is the 𝑟

𝑟

0

vacuum.34,54

dielectric permittivity of the Therefore, it is common for assays with higher target-toreceptor size ratios to show higher sensitivities than those with lower values, as 𝜅 ―1 is proportional 1/2 to (1 𝐶 ) . Note that 1/𝐶 incorporates both Debye and Tomas Fermi screening phenomena. 34,54 𝑟

𝑟

This explains why the engineered receptive interface used to detect different targets was not suitable for detecting IL-6. The receptive-IL-6 interface possessed a 0.17 target-to-receptor size


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ratio, which is quite low compared to other studied receptive layers and thus was not suitable for detection of IL-6 using redox monolayers. The same supportive redox monolayers for the receptors were able to detect other targets when the target-to-receptor size ratio was significantly higher and thus a favorable size ratio for the detection of the target could be achieved. For instance, it was possible, using the same redox supportive monolayer, to detect the C-reactive protein, CRP (Figure 6c). The target-to-receptor size ratio in this case (0.79) was almost 7 times higher than that used to detect IL-6. The same behavior was observed in faradaic impedance configurations (please see Figure 6c). The sensitiveness of the impedance assay presented here was not sufficient enough to allow detection of IL-6 owing to its lower target-to-receptor size ratio. However, when a similar mix SAM (different proportion of MUA and 6-mercapto-1-hexanol) was used with a higher target-to-receptor size ratio, the assay interface was able to detect the protein. This is the case for NS114, whose target-to-receptor size ratio is 0.30, almost 2 times higher than that of IL-6. In summary, it was not possible to detect the IL-6 target when a resistive (using 𝑅𝑐𝑡 signal) or capacitive (using 𝐶𝑟 signal) faradaic EIS configuration was used. The target-to-receptor size ratio appears to be important in different types of faradaic EIS platforms. This is because the process is based on how the charge transfer resistance between the solution and the metallic electrode surface is impeded (resistive configuration) or based on changes in the electrochemical occupancy of the redox moieties within the SAM (capacitive configuration), wherein both receptor and target sizes will proportionally affect the signal.34 Amplifying sensitiveness by a suitable engineering of the interface. To overcome the lower sensitiveness observed in low target-to-receptor size ratio assays, an ideal solution is to increase the target-to-receptor size ratio by using a smaller molecular weight receptor. However, in some cases, the number of available receptors of a certain target is limited. For those cases, changes in the assay configuration that enhance the assay’s sensitiveness can be made. In resistive faradaic configurations (with redox probe in solution), as was discussed, the molecular weight of the biomolecules involved in the assay has an effect on the charge transfer resistance. Nonetheless, the charge of those biomolecules also has an effect on 𝑅𝑐𝑡 that presumably is even higher than that of the massif, as was observed in the EIS result after EDC/NHS activation of the SAM (Figure 3). Thus, using a negatively/positively charged target or receptor makes 𝑅𝑐𝑡 more sensitive to variations in the biosensing interface. For instance, by using this approach, Radi et al.16 developed a resistive faradaic biosensor for the detection of ochratoxin A (OTA), a 404Da mycotoxin. Although OTA is 65 times smaller than IL-6, Radi and collaborators achieved a relative response (RR) of 96% when the specific receptive platform was incubated with 10 ng mL-1 of the target. This was because, at neutral pH, OTA is deionized – OTA2-; therefore, the interaction AbOTA/OTA2- allowed for the generation of an electrostatic repulsion effect over the redox probe in solution, increasing 𝑅𝑐𝑡, even with small mass variation.


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On the other hand, in capacitive faradaic configurations (with redox moiety covalently assembled within the monolayer), sensitiveness can be improved by increasing the number of accessible redox states in the transducer. This has been recently demonstrated by Oliveira et al.55 The structure of the assay contained in a glassy carbon electrode modified with a first layer of active Prussian Blue nanoparticles (PBNP), responsible for the redox capacitive response, and a second layer of graphene oxide (GO), responsible for the covalent assembling of Ab-IL-6 (receptor), was successfully fabricated for the detection of IL-6. Prussian Blue is a well-known electroactive material. Oliveira and collaborators used it as redox transducer because of its intrinsic pseudocapacitance characteristics. The redox activity of the PBNP-GO layer was compared with a redox peptide SAM, in terms of redox density (Figure 7a), redox capacitance, and sensitiveness (Figure 7b). They observed that the PBNP-GO layer showed a redox density of 1.3 x 1015 states per cm2, while the redox peptide SAM presented a redox density of 1.2 x 1014 states per cm2, which was one order of magnitude lower. Accordingly, the value of redox capacitance was quite different between them. The redox peptide SAM presented a redox capacitance of 260 µF cm-2, while the PBNP-GO layer presented a value of 3.2 mF cm-2, which is around 12 times higher. In the previous section it was discussed that the sensitiveness of the assay depends on the length scale (𝜅 ―1) associated with 1/2 the redox capacitance and that 𝜅 ―1 was proportional to (1 𝐶 ) . Thus, the higher the redox 𝑟

capacitance, the lower the length scale for the shielding of the electric-field, and therefore, the higher the sensitiveness of the assay. Because of the supercapacitance of the PBNP-GO layer contribution, the assay introduced by Oliveira et al. was more suitable for detecting IL-6 than the capacitive assay based on the redox peptide layer34 ( please see Figure 7b).

Figure 7. a) Comparisons of CV measurements for PBNP over GO (PBNP-GO) and self-assembly monolayers of redox peptides over gold (r-Pep-SAM). b) Analytical curves corresponding to the detection of IL-6 by redox capacitance variations in a PBNP-GO film compared with the performance of a redox peptide film. The analytical curves were obtained from the relative response in percentage (RR%) of either 𝑅𝑐𝑡 or 1/𝐶𝑟 for each target concentration [see Eq. (1)]. Reprinted (Adapted and Reprinted in part) with permission from Oliveira, R. M. B.; Fernandes, F. C. B.; Bueno, P. R. Pseudocapacitance Phenomena and Applications in Biosensing Devices. Electrochimica Acta 2019, 306, 175–184 – reference 55. Copyright 2019 - Elsevier Ltd.

FINAL REMARKS AND CONCLUSIONS


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In summary, although EIS is a sensitive technique for biosensing, thoughtful designing of the receptive interface and electric properties of the interface is necessary to achieve both sensitive and reliable assaying.34 In the absence of a laborious experimental control associated with an understanding of EIS concepts, changes due to drift or a non-specific binding event can be misinterpreted as a specific biological interaction and hence provide false-positive results. This is quite notorious in label-free platforms where the resolution of the signal-to-noise ratio is low. EIS provides a sensitiveness that resolves signal-to-noise ratio issues, but the signal must be interpreted carefully. Otherwise, the potency of using impedance as an electrical transfer function, where control of the biosensor and its variability by comparing it with a desired control value (a blank response, for instance), is statistically lost. Finally, EIS methods are powerful tools for molecular diagnostics if their concepts are practically used in conjunction with a suitable chemical design of the interfaces containing the transducer (resistive or capacitive) signal. Thus, we suggest the readers to follow the flowchart presented in Figure 8, where we summarized the decisions that should be made during the development of electrochemical biosensors based on either resistive or capacitive approaches. The “algorithm” constructed based on the experience of our research group. This means that this is a preliminary “proposal” which can be inputted as the research in this field, made by our and other groups, processes.


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Figure 8. This is a flowchart suggested for the optimization of an electrochemical biosensors based on a resistive or a capacitive response of the interface. Note that flow of information proposed is based majorly on our group experience, so the information is far to be complete and will requires continuous revisions. (*) shows the values that were obtained by our own data and experience; Fr refers to roughness factor and t/r refers the target-to-receptor size ratio.

SUPPORTING INFORMATION AVAILABLE The Supporting Information is available free of charge. SI-vf.docx: Quartz Crystal Microbalance (QCM) experimental procedure and results, sensograms (Figure SM1) and variable values (Table SM1), demonstrating the formation of the receptive platform of the biosensors.


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AUTHOR INFORMATION Corresponding Author *E-mail: [email protected], tel: +55 16 3301 9642, fax: +55 16 3322 2308

ORCID Beatriz L. Garrote: 0000-0002-4744-1036 Adriano dos Santos: 0000-0001-6812-5609 Paulo R. Bueno: 0000-0003-2827-0208 Notes The authors declare no competing financial interest ACKNOWLEDGMENTS We gratefully acknowledge the Royal Society and FAPESP (São Paulo Research Foundation) for their financial support. REFERENCES (1) (2) (3) (4) (5) (6) (7) (8) (9) (10) (11) (12) (13) (14) (15) (16)

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