Immunosensors in Clinical Laboratory Diagnostics

CHAPTER THREE

Immunosensors in Clinical Laboratory Diagnostics Celine I.L. Justino*,†,1, Armando C. Duarte*, Teresa A.P. Rocha-Santos* *

CESAM and Department of Chemistry, University of Aveiro, Aveiro, Portugal ISEIT/Viseu—Instituto Piaget, Estrada do Alto do Gaio, Galifonge, Lordosa, Viseu, Portugal Corresponding author: e-mail address: [email protected]

† 1

Contents 1. Introduction 2. Classification of Immunosensors 2.1 Electrochemical Immunosensors 2.2 Optical Immunosensors 2.3 Piezoelectric Immunosensors 3. Applications of Immunosensors for Clinical Diagnostics 3.1 Cancer Biomarkers 3.2 Cardiac Biomarkers 3.3 Hormones 3.4 Pathogenic Bacteria 3.5 Virus 4. Final Remarks and Perspectives Acknowledgments References

66 68 69 71 72 72 73 86 96 98 99 102 103 103

Abstract The application of simple, cost-effective, rapid, and accurate diagnostic technologies for detection and identification of cardiac and cancer biomarkers has been a central point in the clinical area. Biosensors have been recognized as efficient alternatives for the diagnostics of various diseases due to their specificity and potential for application on real samples. The role of nanotechnology in the construction of immunological biosensors, that is, immunosensors, has contributed to the improvement of sensitivity, since they are based in the affinity between antibody and antigen. Other analytes than biomarkers such as hormones, pathogenic bacteria, and virus have also been detected by immunosensors for clinical point-of-care applications. In this chapter, we first introduced the various types of immunosensors and discussed their applications in clinical diagnostics over the recent 6 years, mainly as point-of-care technologies for the determination of cardiac and cancer biomarkers, hormones, pathogenic bacteria, and virus. The future perspectives of these devices in the field of clinical diagnostics are also evaluated.

Advances in Clinical Chemistry, Volume 73 ISSN 0065-2423 http://dx.doi.org/10.1016/bs.acc.2015.10.004

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2016 Elsevier Inc. All rights reserved.

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1. INTRODUCTION The use of biosensors in the clinical diagnostics has been reported in literature as alternative technologies from traditional methodologies. In such scientific area, immunosensors are immunological biosensors, since they are based on the transduction of signals (with electrochemical, optical, or piezoelectric source) generated in the interactions between antibodies or antigens in body fluids (mainly in serum). Immunosensors have been developed for continuous monitoring of analytes through point-of-care devices, which provide low cost, full automation, portability, fast response, high sensitivity, accuracy, and precision [1]. Immunoassays are biochemical tests for quantifying a substance based on its ability to bind to an antibody (immunoreaction), but they require expensive instrumentation, sample pretreatment protocols, laborious work, and rigid operational procedures. Besides, they are time consuming and the detection is indirect, since they are based on optical transduction and one of the reagents is labeled [2]. Thus, the immunosensors have many advantages over the corresponding traditional analytical techniques (immunoassays), since they can provide fast screening results, they are small, semiautomated, and increasingly portable [3]. The increase in portability has been the consequence of improvements in sensor miniaturization. In the majority of cases, immunosensors have a unique use, that is, they are disposable but the regeneration of the immunosensor surface is also highly requested due to the cost-effectiveness of regenerated sensors and the washing with an appropriate solution of high ionic strength or low pH can be considered for this aim. According to Bahadir and Sezgintu¨rk [4], even though the antibody–antigen linkage can be broken under drastic conditions (e.g., in alkaline or acidic solutions), the immobilized immunoreagents could also suffer from the functional damage or even be released from the immunosorbents. Some works have yet tested the regeneration of immunosensors with good results. For example, Zhuo et al. [5] rinsed immunosensor with 0.05 M NaOH and then used 0.05 M HCl, 4 M urea, or 0.2 M glycine-HCl buffer solution (pH 2.8) to break bonds between carcinoembryonic antigen (CEA) and anti-CEA and the regenerated immunosensors were used to detect the same concentration. After the first seven times regeneration by urea solution, the immunosensor kept 95.97% of the original amperometric responses [relative standard deviation (RSD) of 0.44%], after the first four times regeneration by glycine-HCl solution, the immunosensor kept 96.21%, and after the first

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three times regeneration by HCl solution, the immunosensor kept 95.28%, while after the first two times regeneration by NaOH solution, the immunosensor kept 96.5% of the original responses. Thus, Zhuo et al. [5] chose urea solution as a suitable reagent for recycling the immunosensors. The miniaturization of immunosensors and sensors in general is due to the impact of nanotechnology in the design, materials, and methods for construction, and the development of nanosensors has been attracted considerable attention in the field of clinical applications. Nanomaterials such as nanowires synthesized from metals (e.g., Ni, Cu, Au, Pt), metal oxides (ZnO, SnO2, Fe2O3), and silicon/indium/gallium semiconductors (Si, InP, GaN), quantum dots based on CdSe, CdTe, or CdSeTe, carbon nanotubes (CNT), and metal nanoparticles (gold nanoparticles, nanoparticles of copper, palladium, cobalt, silver, or platinum) have been incorporated in biosensors [6]. Large improvements in the detection and quantification of compounds such as proteins, neurotransmitters, and cancer biomarkers by biosensors have been reported due to the integration of nanoscience and nanotechnology, as reviewed by Justino et al. [2,7]. For example, the association of nanostructured materials with various functional groups such as carboxyl and amine groups, or with metallic nanoclusters such as gold and silver, has been used for the improvement of signal transduction, by amplifying the binding event, in the case of immunosensors. Point-of-care testing involves rapid diagnostic tests carried out at or near the site of patient care to obtain immediate results [8]. Thus, point-of-care sensors and biosensors have significant applications as diagnostic tools in the point-of-care testing. For example, the incorporation of nanomaterials such as CNT in point-of-care biosensors is a promising research area, due to their sensitivity, their adequate response time, and their fitness for purpose, since they can be hand-held and used by patients, as recently reviewed by Justino et al. [2]. The application of immunosensors in clinical diagnosis has been emphasized in recent, interesting reviews: (a) Chikkaveeraiah et al. [9] highlighted the advances in the development of electrochemical immunosensors for detection of cancer biomarkers [e.g., prostate-specific antigen (PSA), prostate-specific membrane antigen, platelet factor 4, interleukin 6 (IL-6), interleukin 8 (IL-8), and CEA], with emphasis on opportunities for further improvement in cancer diagnostics and treatment monitoring; (b) Hasanzadeh et al. [10] and Pedrero et al. [11] focused their attention on the description of electrochemical nanoimmunosensing of cardiac biomarkers for acute myocardial infarction (e.g., cardiac troponin I (cTnI)

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and cardiac troponin T (cTnT), myoglobin, and thrombin); (c) Ronkainen and Okon [6] reviewed the nanomaterial-based electrochemical immunosensors for clinically significant biomarkers; and (d) Bahadir and Sezgintu¨rk [4] focuses on recent development for tumor and cardiac markers testing and monitoring of early clinical diagnosis with electrochemical immunosensors. This chapter provides a state-of-the-art of the immunosensors employed in the area of clinical diagnostics; practical examples are discussed corresponding to literature covering the recent 6 years. The classification of immunosensors is also reported.

2. CLASSIFICATION OF IMMUNOSENSORS In an immunosensor, the interaction between antibody and antigen origins an analytical signal, which is converted through a transducer into a physicochemical response. The transduction mechanisms employed in immunosensors can be based on signal generation (electrochemical or optical signal) or on changes in physicochemical properties such as mass charges, which means that the respective immunosensors are called electrochemical, optical, or piezoelectric, respectively. The major advantage of immunosensors is related to the selectivity and affinity of the antibody–analyte binding reaction. Figure 1 shows the different types of transduction mechanisms and respective response.

Surface acoustic wave

Quartz crystal microbalance

Elasticity or density

Mass

Reflectometric interference spectroscopy

Differential interferometry

Piezoelectric

Optical path or refractive index

Refractive index

Grating couplers

Surface plasmon resonance

Total internal reflection fluorescence Fluorescence

Conductimetry Conductance or resistance

Potentiometry Electrode potential or voltage

Amperometry Current

Transduction mechanism

Response

Optical

Electrochemical

Immunosensor

Figure 1 Types of transduction mechanisms in immunosensors and respective response.

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A

Enzyme

B

yme

Enz

yme

Enz

Antigen Antibody

Figure 2 Schematic diagrams of: (A) sandwich-type immunosensor and (B) competitivetype immunosensor.

Another classification of immunosensors is based on their format, that is, in the type of antibody–antigen interaction such as capture, sandwich, competition, and inhibition. The sandwich-type immunosensors tend to be more sensitive and robust, and they are the most commonly used together with the competitive-type immunosensors, which are schematically shown in Fig. 2. In sandwich-type immunosensors (Fig. 2A), two antibodies are used. The capture antibodies first bind to the transducer material, followed by the addition of antigens, and finally, the enzyme-labeled antibodies react with the substrate to form the antigen–antibody complexes. The enzyme most used is the horseradish peroxidase (HRP). In competitive-type immunosensors (Fig. 2B), a competitive interaction of labeled and unlabeled antigens occurred with the capture, that is, the binding of the capture antibodies to the transducer material is followed by the addition of enzymelabeled antigens to form antigen–antibody complexes.

2.1 Electrochemical Immunosensors The´venot et al. [12] defined “electrochemical biosensor” as a self-contained integrated device capable of providing specific quantitative or semiquantitative analytical information using a biological recognition element that is retained in direct contact with an electrochemical transduction element (e.g., a pair of electrodes or field-effect transistors). When the recognition element is an antibody immobilized on the transduction surface, the biosensors are antibody–antigen-based affinity biosensors, that is, immunosensors, used for the detection of an antigen, which has a corresponding specific antibody. Thus, the technique for immobilization of the antibody is important

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for the performance of the immunosensor. Physical and chemical adsorption can be used; for example, the self-assembly technique is a chemical adsorption method for the immobilization of antibody, in which a self-assembly monolayer (SAM) is fixed on the surface through chemical bonding, and the antibody is covalently attached to the monolayer by using crosslinkers [4]. The event of the formation of antigen–antibody complex is converted into an electrochemical signal: an electric current (for amperometric immunosensors), a voltage difference (for potentiometric immunosensors), or a resistivity change (for conductimetric immunosensors). Significant differences in magnitude of such electrochemical signals occurred when the antigen–antibody complex is formed. For example, in FET, the formation of antigen–antibody complex at the gate terminal modulates the charge carrier flow between source and gate, generating an increase in current [13], as shown in Fig. 3. The main advantage of electrochemical immunosensors is the ease for miniaturization. With the progress of nanoscience and nanotechnology, the assemblage of new electrochemical immunosensors has been reported, for example, with the fabrication of electrochemical immunosensors including nanomaterials such as CNT, carbon nanofibers, and carbon nanorods which have large surface area, good chemical stability, and excellent biocompatibility [2].

VD

VS

VG Before recognition

Drain current

VD

VS

VG After recognition

Figure 3 Schematic of an electrochemical immunosensor based on FET principle (© 2012, Moina and Ybarra [13]. Originally published in “Fundamentals and applications of immunosensors” under CC by 3.0 license. Available from doi:10.5772/36947).

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2.2 Optical Immunosensors Optical biosensors are powerful detection instruments and versatile tools for analytical purposes, not only due to their low signal-to-noise ratio and low reagent volume requirements, but also because they are immune to electromagnetic interference and capable of performing remote sensing, and can provide multiplexed detection within a single device [14]. In optical immunosensors, the optical signal is generated (color or fluorescence) or changed in optical properties of the environment according to the formation of antibody–antigen complex. The biological-sensitive element is immobilized onto the surface of the transducer and the optical signals are collected by a photodetector and converted into electrical signals that are further electronically processed [13]. For example, the fluorescence is one of the most widely methods employed in immunosensors, where the conjugate antibody is labeled with a fluorescent probe. Figure 4 shows the schematic of an optical sensor based on internal reflection fluorescence. This technique can enhance the sensitivity of the biosensor; when light is transported by total internal reflection in an optical guide, an evanescent field is generated at the interface between the guide and the external medium [13]. As explained by Moina and Ybarra [13], on reflection at dielectric interface, light penetrates into the second phase that has a lower refractive index than that of the core; the intensity further decreases exponentially over the penetration (which typically is of the order of the incident wavelength), and any labeled antibodies located close to the glass-aqueous medium interface are excited to produce fluorescence, while those located further into the solution will not.

Figure 4 Schematic of an optical immunosensor based on the principle of total internal reflection fluorescence (©2012, Moina and Ybarra [13]. Originally published in “Fundamentals and applications of immunosensors” under CC by 3.0 license. Available from doi:10.5772/36947).

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Figure 5 Schematic of a piezoelectric immunosensor based on the principle of a microcantilever. (A) Before antibody addition; (B) after antibody addition. The formation of the antigen–antibody complexes provokes a surface stress and, consequently, a deflection of microcantilever, which is detected optically (© 2012, Moina and Ybarra [13]. Originally published in “Fundamentals and applications of immunosensors” under CC by 3.0 license. Available from doi:10.5772/36947).

2.3 Piezoelectric Immunosensors Piezoelectric sensors employ materials that resonate under the application of an external alternating electrical field. Typically, quartz crystals are utilized, producing an oscillating electric field in which the resonant frequency of the crystal depends on its chemical nature, size, shape, and mass [15]. The mass changes that take place after the formation of antigen–antibody formation can be measured by means of piezoelectric transducers, such as quartz crystal microbalances (QCM) and microcantilevers. In QCM, the immobilized antibody–antigen interaction occurs at the surface of the crystal, which is within an oscillating circuit, so the changes in mass lead to a decrease in resonant frequency. Such devices have been used for a wide range of applications in the medical field due to their sensitive response, low cost, and ease of operation [16]. In microcantilevers, the formation of the antigen–antibody complexes provokes a surface stress and, consequently, a deflection of microcantilever, which is detected optically, as shown in Fig. 5. Microcantilevers are especially attractive for immunosensors because of the possibility of microfabrication at low cost.

3. APPLICATIONS OF IMMUNOSENSORS FOR CLINICAL DIAGNOSTICS The application of immunosensors for clinical diagnostics, and based on electrochemical, optical, and piezoelectric principles, is mainly reported for detection of: (a) cancer biomarkers [e.g., PSA, interleukins (IL-6 and IL-8), matrix metalloproteinases (MMP-2 and MMP-3), α-1-fetoprotein, and CEA];

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(b) cardiac biomarkers [e.g., C-reactive protein (CRP), cTnT and cTnI, myoglobin, and amino-terminal pro-B-type natriuretic peptide (NT-proBNP)]; (c) hormones (e.g., cortisol and estradiol); (d) pathogenic bacteria (e.g., Escherichia coli and Streptococcus pyogenes); and, (e) virus (e.g., dengue virus, hepatitis C virus, and avian influenza virus).

3.1 Cancer Biomarkers In ideal conditions, the cancer biomarkers would be proteins or protein fragments that can easily detected and quantified in the biological fluids of patients, such as blood or urine, with non- or minimally invasive collection [6]. According to Malhotra et al. [17], the development of devices for sensitive and reliable point-of-care measurement of biomarker proteins for early cancer detection and treatment monitoring is a significant challenge. Since in the absence of a tumor, the levels of the tumor marker are low, rising even small tumor forms, the limit of detection (LOD) of an analytical methodology should be as lowest as possible for the early screening of a small tumor [4]. The potential of such devices is very important, since the point-of-care analyses would reduce costs, minimize sample decomposition, facilitate on-the-spot diagnosis, and alleviate patient stress. On the other hand, they require minimal technical expertise and system maintenance. Table 1 shows the normal concentrations of cancer biomarkers in blood. 3.1.1 Prostate-Specific Antigen PSA is a single-chain glycoprotein and it is the most typical tumor marker used as screening tool for clinical diagnosis of prostate cancer, which could provide direct information about the diagnosis of the prostate cancer via its Table 1 Normal Concentrations of Cancer Biomarkers in Blood Cancer Biomarker Normal Concentration

PSA

0.5–2 ng mL

IL-6

6 pg mL
1

IL-8

<20 pg mL

MMP-3

14 ng mL

α-1-Fetoprotein

10 ng mL
1

CEA

3 ng mL

References

[18] [17]


1


1


1


1

in women; 34 ng mL

[19]
1

in men

[20] [4] [21]

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concentration in blood [22,23]. Efficient methods have been developed for the detection of PSA based on immunoreaction such as enzyme-linked immunosorbent assay (ELISA), time-resolved fluorescence assay, chemiluminescence immunoassay, bioluminescence immunoassay, and electrochemical immunoassay [23]. However, electrochemical immunosensors exhibit significant advantages, including high sensitivity, simple instrumentation, and easiness of miniaturization [23]. Yang et al. [24] and Li et al. [25] developed electrochemical immunosensors for PSA detection. Yang et al. [24] used quantum dots functionalized graphene sheets as labels of the electrode and Li et al. [25] have used graphene–cobalt hexacyanoferrate nanocomposite to modify the electrode surface. In the work of Yang et al. [24], the immunosensor displayed a wider range for the linear response (0.005–10 ng mL
1) and a lower LOD (3 pg mL
1) and in the work of Li et al. [25], the amperometric signal decreased linearly with PSA concentration in the range of 0.02–2 ng mL
1 with a low LOD of 0.01 ng mL
1. To evaluate the reproducibility of the immunosensors, a series of five electrodes were prepared for the detection of 1 ng mL
1 of PSA and the RSD of their measurements were 7.9% and 6.7%, respectively. In the work of Yang et al. [24], five patient serum samples were used for detection of PSA both with the immunosensor and the reference methodology (ELISA). According to the results, the PSA contents determined by the immunosensor agreed well with the ELISA measurements with a correlation of 89.8%. In the work of Li et al. [25], six clinical serum samples were tested and the results were compared with those obtained by standard ELISA. There is no significant difference between the results of the two methods with RSD less than 9%, indicating the proposed method could be used for routine clinical testing. Zani et al. [22] proposed an electrochemical immunosensor for PSA based on antibody-modified paramagnetic microparticles, which were coupled to the multiplexed electrochemical platform. Such platform consists of 8-sensors screen-printed arrays (based on carbon working electrodes) as the electrochemical transducers and a simple target capturing step by means of anti-PSA and magnetic beads functionalized with protein G. The calibration curve was constructed of PSA concentrations between 0 and 400 ng mL
1, obtaining an LOD of 1.4 ng mL
1. Real samples of serum were tested for detection of PSA and the reproducibility obtained was similar to that found for standard solutions (RSD of 8%), indicating that the system can be suitable for real samples analysis [22]. The authors also concluded that the electrochemical immunosensor is a sensitive and multiplexed tool for fast and easy PSA analysis. Tian et al. [26] also

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fabricated an electrochemical immunosensor for PSA detection but based on glassy carbon electrode modified with a nanocomposite containing gold nanoparticles supported with starch-functionalized multiwalled CNT (MWCNT). The nanocomposite exhibits high current response intensity, good biocompatibility, and high stability, which are demonstrated to be well competent for the development of electrochemical immunosensors [26]. Under optimal conditions (pH 6.0; 25 °C), an LOD of 7 pg mL
1 was obtained within a concentration range of 0.01–0.5 ng mL
1. A regeneration step of the immunosensor was also tested by Tian et al. [26]; after detection of PSA, the electrode was retreated with 0.2 M glycine-HCl (pH 2.0) solution for 1 min to break the antibody–antigen linkage. Then, the electrode was repeated six times of consecutive measurements of one sample and an RSD of 6.2% was obtained at PSA concentration of 0.5 ng mL
1. The performance of the immunosensor was tested for the detection of PSA in six normal serum samples and the recovery of PSA detected by the immunosensor ranged from 91.0% to 106.3%. According to the authors, the developed immunosensor exhibits good sensitivity and reproducibility (RSD of 4.1%) and may become a promising technique for the diagnosis and monitoring of prostate cancer [26]. Recently, Li et al. [23] also employed nanomaterials to the transduction of immunosensors applied to detect PSA, and fabricated a label electrochemical immunosensor based on graphene- (amino-functionalized graphene sheet–ferrocenecarboxaldehyde composite materials) and silver-hybridized mesoporous silica nanoparticles. Under optimal conditions (pH 7.4; concentration of amino-functionalized graphene sheet of 0.8 mg mL
1; concentration of ferrocenecarboxaldehyde of 1.0 mg mL
1), the fabricated immunosensor showed a wide linear range with PSA concentration (0.01–10.0 ng mL
1) and a low LOD (2 pg mL
1). In addition, the immunosensor demonstrated good stability (RSD lower than 10% after 3 weeks of stand-by) and reproducibility (RSD of 6.4%). The authors concluded that the structure of the nanocomposites can be applied for the immobilization of biomolecules other than PSA and in clinical diagnostics [23].

3.1.2 Interleukins ILs are a major class of biologically active protein mediators, known as cytokines. They are secreted proteins that release from activated cells and bind to their specific cell surface receptors in order to induce growth and differentiating functions in target cells [27]. ILs are assigned to each family (total of

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37 families) based on sequence homology and receptor chain similarities or functional properties [27]. For example, IL-6 is produced by endothelial cells, fibroblasts, monocytes, and macrophages during systemic inflammation and it has a critical role in the inflammatory response, such as rheumatoid arthritis, cardiovascular disease (CVD), and inflammatory bowel disease [27,28]. It is also a suitable biomarker overexpressed by several types of cancer, including head and neck squamous cell carcinoma [28]. It can also be overexpressed by several types of cancer such as head and neck squamous cell carcinoma [17]. IL-8 is produced by a variety of cells, such as monocytes and macrophages, neutrophils, and lymphocytes and increased concentrations of IL-8 were found in inflammatory sites in patients with psoriasis, respiratory syncytial virus infection, or chronic obstructive pulmonary diseases [27]. Malhotra et al. [17] applied an ultrasensitive electrochemical immunosensor for the detection of IL-6 in serum. Single-walled CNT (SWCNT) with attached antibodies for IL-6 coupled with multienzyme labels was used as a strategy for such detection. Malhotra et al. [17] obtained good analytical performance in terms of accuracy and reproducibility. An excellent LOD (0.5 pg mL
1) and high sensitivity (19.3 nA mL cm
2 (pg IL-6)
1) were obtained when compared with that of their previous work [29], which LOD obtained was of 30 pg mL
1 and sensitivity of 3.6 nA mL cm
2 (pg IL-6)
1; the LOD was much lower than that of standard ELISA of 0.1 ng mL
1 [30]. A concentration range of 1–1000 pg mL
1 in complex matrix samples was obtained, which is representative of IL-6 in the serum of cancer disease and diseases-free patients. It demonstrates the potential of the immunosensor for detecting IL-6 in research and clinical applications such as point-of-care testing. Wang et al. [28] fabricated an electrochemical immunosensor based on gold nanoparticles–graphene–silica sol–gel as immobilization biointerface using gold nanoparticles–polydopamine– CNT as the label of IL-6 antibodies. The sensing platform provides a stable network for the immobilization of antibody and exhibits a dynamic working range of 1–40 pg mL
1 with a low LOD of 0.3 pg mL
1 (at 3 s). Wang et al. [28] compared the results of serum samples by the immunosensor with the results of an ELISA method, obtaining an acceptable agreement. Effective materials, such as the polydopamine, are novel, simple, inexpensive, “green,” and stable films for labeling; in this case associated with gold nanoparticles and CNT [28]. According to the authors, the “green” operation and ultrasensitivity of the developed methodology provide a promising potential in clinical diagnosis. An optical immunosensor based on

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electrochemiluminescence was constructed by Liu et al. [31] for the determination of IL-6. The immunosensor was based on graphene oxide nanosheets/polyaniline nanowires/CdSe quantum dots nanocomposites, which were synthesized due to their excellent biocompatibility and solubility [31]. The immunosensor has a sensitive response to IL-6 in a linear range of 0.0005–10 ng mL
1, exhibiting high specificity, long-term stability (no statistically difference of signal after 14 days with RSD of 7.4%), and excellent reproducibility (RSD of 5.9%) with an LOD of 0.17 pg mL
1, and could provide a promising application for the detection of proteins in clinical samples [31]. Recently, Ojeda et al. [32] and Lou et al. [33] also fabricated immunosensors for IL-6 but based on different transduction principles. Ojeda et al. [32] prepared an electrochemical magnetoimmunosensor (covalent immobilization of anti-IL-6 antibodies onto carboxyl-functionalized magnetic microparticles) using poly-HRP streptavidin conjugates as labels for signal amplification for the determination of IL-6 in saliva and urine. A schematic diagram of the fundamentals of the developed magnetoimmunosensor is shown in Fig. 6. An LOD of 0.39 pg mL
1 was obtained within a concentration range of 0–1000 pg mL
1. The anti-IL-6 magnetic beads conjugates exhibited excellent storage stability providing amperometric responses with no significant loss during at least 36 days [32]. The authors also referred that the time needed for the assay, including the capture antibody immobilization on the activated magnetic beads, the blocking step, the formation of the

Anti-IL6

Ethanolamine

IL-6

Biotin-anti-IL-6

H2O

H2O2

HRPox

HRPred

HQred

HQox e−

Poly-HRP-strept

Magnet

Figure 6 Schematic display of the different steps involved in the preparation and the functioning of magnetoimmunosensor. Reprinted from Ojeda et al. [32], with kind permission from Springer Science and Business Media.

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immunocomplex, and the amperometric measurement, is 3 h, which is shorter than the preparation of other electrochemical immunosensors [32]. In addition, the obtained results were statistically in agreement with those provided by a commercial ELISA kit (no significant differences were found between both methodologies at a significance level of 0.05). Regarding the work of Lou et al. [33], an electrochemical immunosensor was constructed and based on the electrically heated carbon electrode and silver nanoparticles functionalized labels. The heated carbon electrode is a great way to accelerate reaction kinetic without changing the bulk solution temperature while it improves the mass transport by changing temperature of electrode, thus leading to an enhanced electrochemical signal together with a higher signal-to-background ratio [33]. The immunosensor for IL-6 exhibited a wide linear response to IL-6 ranging from 0.1 to 100,000 pg mL
1 with an LOD of 0.059 pg mL
1. The immunosensor was also used to detect IL-6 in serum samples with satisfactory results, since RSD of 6.3% was obtained when the results of the immunosensor were compared with those of ELISA. Regarding the IL-8, Wan et al. [34] constructed an electrochemical immunosensor for the simultaneous detection of PSA and IL-8 with a screen-printed carbon electrode array (with 16 channels) based on MWCNT as the detection platform. The immobilization of capture antibodies on this platform was considerably easy and robust, which involved an electrochemical activation of the carbon working electrode to generate carboxylate groups [34]. LOD of 5 pg mL
1 for PSA and 8 pg mL
1 for IL-8 was reported within a concentration range tested of 5–4000 pg mL
1 and 8–1000 pg mL
1, respectively, suggesting that the multiplexing immunosensor is a promising approach for point-of-care testing in clinical diagnostics. Another work reporting the construction of an immunosensor for IL-8 was proposed by Munge et al. [35]. Superparamagnetic particles were combined with gold nanoparticles to promote the attomolar detection of IL-8 (LOD of 1 fg mL
1, which corresponds to 100 aM). According to Munge et al. [35], such LOD is 30,000-fold lower than that of the conventional ELISA and about 1000-fold lower than that of commercial bead-based protein assays. Primary antibodies were attached in the dense film of gold nanoparticles that capture human IL-8 from the sample; then, when coupled to superparamagnetic beads, 500,000 HRP labels and secondary antibodies were loaded to develop an amperometric signal proportional to the amount of IL-8 in the sample.

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3.1.3 Matrix Metalloproteinases MMPs are zinc-dependent proteases and they are associated with the invasion and the metastasis of cancer cells, and also involved in the progression of atherosclerosis and aneurysm formation [36]. In particular, MMP-2, also known as gelatinase A, is one of the crucial MMPs in tumor growth, invasion, and metastasis, and plays a key role in physiological and pathological states including morphogenesis, reproduction, and tissue remodeling due to its ability to degrade type VI collagen [37]. Yang et al. [37] have developed an electrochemical immunosensor for detection of MMP-2, which consists in an assembly of well-defined gold nanoparticles on nitrogen-doped graphene sheets. Such composite facilitated robust immobilization of antibodies, promoted electron transfer, and exhibited excellent electrochemical activity [37]. A concentration range from 0.0005 to 50 ng mL
1 was tested, obtaining an LOD of 0.11 pg mL
1. According to the authors, the immunosensor exhibited good stability (more than 90% of the initial response of the immunosensor could be remained after 1 week stored at 4 °C) with adequate reproducibility (RSD of 5.7%) and accuracy (RSD of 6.7%), and also demonstrated efficiency in the detection of MMP-2 in real samples. The MMP-3 was detected in calf-serum samples, as a clinically related medium, with an electrochemical immunosensor based on vertically aligned SWCNT [38]. Munge et al. [38] obtained an LOD of 0.4 ng mL
1, applying the multienzyme detection principle, which is lower than the normal range of MMP-3 levels in subjects with cancer. They indicated that this immunosensor could constitute a point-of-care device, taking into account its short time for analysis, simplicity, and cost-effectiveness, and high sensitivity and selectivity in detecting the cancer biomarker. Recently, Yin et al. [39] have constructed a label-free electrochemical immunosensor for determination of MMP-3 and based on the covalent immobilization of MMP-3 antibody on a 3D graphene oxide/polypyrrole-ionic liquid composite film. The 3D film was constructed by one-step electrodeposition of graphene oxide/polypyrrole in 1-butyl-3-methylimidazolium tetrafluoroborate onto an indium tin oxide electrode modified with a silica opal template. Such composite film showed high surface area, good biocompatibility, and enhanced electronic conductivity [39]. The relative increased impedance values are proportional to the logarithmic value of MMP-3 concentrations in a linear range of 1–1000 pg mL
1 with a low LOD of 1 pg mL
1. According to the authors, such a 3D film could have widespread applications in clinical screening of cancer biomarkers for point-of-care diagnosis.

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3.1.4 α-1-Fetoprotein The α-1-fetoprotein is an oncofetal glycoprotein that can be widely used as a tumor marker and its concentration increases in response to hepatocellular carcinoma and liver carcinoma [4,40]. Che et al. [40] have constructed an amperometric immunosensor for the determination of α-1-fetoprotein. The immunosensor was based on MWCNT–silver nanoparticle composite on the surface of a glassy carbon electrode. Gold nanoparticles were electrodeposited on the electrode due to their large specific surface area, strong adsorption ability, and good conductivity, in order to immobilize anti-α1-fetoprotein. The amperometric detection was based on the change of the peak current response before and after antigen–antibody reaction. The system was optimized to realize a reliable determination of α-1-fetoprotein in the range of 0.25–250 ng mL
1 with an LOD of 0.08 ng mL
1. The proposed immunosensor showed a rapid and highly sensitive response to α-1-fetoprotein with acceptable stability (after 2 weeks and 1 month, the immunosensor retained 94.1% and 85.3% of its initial current response, respectively) and reproducibility (RSD of 3.3%). According to the authors, as no significant difference was observed between the proposed immunosensor and a typical ELISA method when applied for detection of α-1-fetoprotein in real serum samples, the immunosensor is able to be applied for clinical determination of this cancer biomarker. Du et al. [41] described an electrochemical immunosensor for sensitive detection of the cancer biomarker α-1-fetoprotein. The immunosensor is a graphene sheet sensor platform which employs functionalized carbon nanospheres labeled with HRP-secondary antibodies. The functionalized graphene sheets used for the sensor platform increased the surface area to capture a large amount of primary antibodies, thus amplifying the detection response [41]. A sevenfold increase in detection signal was verified in the immunosensor due to the signal amplification strategy of graphene sheets and the multienzyme labeling compared to the immunosensor without graphene modification and carbon nanospheres labeling [41]. In addition, the immunosensor could respond to 0.02 ng mL
1 α-1-fetoprotein with a linear calibration range from 0.05 to 6 ng mL
1. According to the authors, this amplification strategy is a promising platform for clinical screening of cancer biomarkers and point-of-care diagnostics. Gan et al. [42] also developed an immunosensor of α-1-fetoprotein but based on a magnetic transduction principle. The amperometric immunosensors were based on a glassy carbon electrode modified with a nanomagnetic Fe3O4/ZrO2/nanogold composite membrane for the immobilization of antibodies. The modified electrode was sensitive to

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α-1-fetoprotein with a linear relationship between 0.05 and 10 ng mL
1 with an LOD of 0.01 ng mL
1 under the optimal conditions (pH 7.0; applied potential of
0.45 V). The immunosensor was tested with five human serum samples by determining the concentration of α-1-fetoprotein with the proposed immunosensor and ELISA; there was no significant difference between the results and the ELISA, which means that the immunosensor could be satisfactorily applied to the clinical determination of α-1-fetoprotein levels in human serum. Similarly with the work of Du et al. [41], Li et al. [43] also employed a dual-signal amplification strategy for photoelectrochemical immunosensing of α-1-fetoprotein. Titanium dioxide coupled with α-1-fetoprotein–CdTe quantum dots–glucose oxidase bioconjugate was synthesized for signal amplification. The fabrication procedure of the immunosensor is shown in Fig. 7. The authors reported a greatly enhanced sensitivity for α-1-fetoprotein due to dual-signal amplification strategy. The immunosensor has a linear detection range from 0.5 pg mL
1 to 10 μg mL
1 with an LOD of

Figure 7 Schematic fabrication process of photoelectrochemical immunosensor. A greater amount of CdTe can lead to much more efficient light absorption and consequently increase the photocurrent response by more electron injection from the excited CdTe. ITO, indium tin oxide; CS, chitosan; anti-AFP, antibodies specific to α-1-fetoprotein; BSA, bovine serum albumin; AFP, α-1-fetoprotein; AFP–CdTe–GOx, α-1-fetoprotein–CDTe quantum dots–glucose oxidase bioconjugate. Reprinted with permission from Li et al. [43]. Copyright (2012) American Chemical Society.

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0.13 pg mL
1. In addition, a good long-term stability (when the anti-α-1fetoprotein antibody-modified immunosensor was stored in a dark and moisturizing environment at 4 °C over 2 weeks, no apparent change in photocurrent response in the detection of 10 ng mL
1 α-1-fetoprotein was found) and reproducibility (RSD of 1.0–1.5%) were reported [43]. Recently, Wang et al. [44] reported an electrochemical immunosensor for α-1-fetoprotein detection, where graphene sheets and thionine were used as electrode materials (for promoting electron transfer and immobilization of primary antibody of α-1-fetoprotein), and mesoporous silica nanoparticles loaded with Fe3O4 nanoparticles (for immobilization of secondary antibody of α-1-fetoprotein) and HRP were employed as labels for signal amplification. Figure 8A shows the fabrication procedure of the immunosensor. The immunosensor could detect α-1-fetoprotein over a wide concentration range from 0.01 to 25 ng mL
1 with an LOD of 4 pg mL
1; calibration plot is shown in Fig. 8B.

3.1.5 Carcinoembryonic Antigen CEA is a cell surface glycoprotein and it is one of the most widely used tumor makers responsible for clinical diagnosis of colorectal, pancreatic, gastric, and cervical carcinomas [45]. Ho et al. [46] proposed a sensitive electrochemical immunosensor for CEA detection based on a carbon nanoparticle-modified screen-printed electrodes covered with anti-CEA antibodies. Cadmium sulfide nanocrystals were used as biotracers for signal amplification and as platform for the secondary antibodies and carbon nanoparticles were used to enhance electron transfer. According to the authors, the immunosensor based on stripping voltammetry has shown potential for point-of-care or disposable home-care self-diagnostic tool, due to its high sensitivity and low LOD (32 pg mL
1) into a linear range of 0.032–10 ng mL
1. Urine samples were used in order to determine urinary CEA levels, which are very important for the early detection of urothelial carcinoma and also a better marker than serum CEA. The method was found as precise and sensitive, obtaining a coefficient of variation (CV) of 17% for the peak current from the five replicates and acceptable recovery rates (108.8%). A sensitive electrochemical immunosensor has been proposed by Huang et al. [47] to detect CEA using a composite containing gold nanoparticles, MWCNT, and chitosan. The mixture was dripped on the glassy carbon electrode and anti-CEA antibodies were immobilized on the resulted

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A

50 nm

100 nm

2

H

NH2 N

NH 2

2

2

NH2 N H

O4 Fe 3

NH

Br

Ab2

HR

P

NH 2

NH 2

NH

HN

2

NH2

NH N H NH

NH

NH2

NH 2

HN

MSNs

GS-TH

Ab1

BSA

AFP

GCE

B Current change (µA)

25

20

15

10 0

5

10

15

Concentration (ng

20

25

mL–1)

Figure 8 (A) Schematic representation of the electrochemical immunosensor for detection of α-1-fetoprotein. TEM images of mesoporous silica nanoparticles (MSNs) and MSNs–Fe3O4 nanoparticles are also observed and (B) calibration curve of the α-1-fetoprotein immunosensor. Ab2, secondary antibodies; HRP, horseradish peroxidase; GS-TH, graphene sheets and thionine; Ab1, primary antibody; BSA, bovine serum albumin; AFP, α-1-fetoprotein. Reprinted from Wang et al. [28], Copyright (2014), with permission from Elsevier.

modified electrode to construct the immunosensor. Under optimal conditions (pH 6.0; incubation time of 20 min and incubation temperature of 37 °C to antibody immobilization), an LOD of 0.01 ng mL
1 was obtained, which would be valuable for diagnosis and monitoring of carcinoma. The specificity of the immunological reaction, the sensitivity obtained with

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the signal amplification due to the gold nanoparticles and MWCNT, and the stability of the proposed immunosensor (RSD of 3.8% after 60 measurements and no signal decline after storage of 30 days) are also other advantages of such immunosensor [47]. In addition, Huang et al. [47] tested the regeneration of the immunosensor by treating the electrode with 4 M urea solution for 8 min to break the antibody–antigen linkage and an RSD of 4.2% was obtained after repeating the analysis six times, which demonstrates the success of regeneration. Human serum samples were used for detection of CEA with the proposed method and ELISA. No significant difference was found between the results and ELISA method, which means that the developed immunosensor could be satisfactorily applied to the clinical determination of CEA levels in human serum [47]. Zhong et al. [48] developed an electrochemical immunosensor for CEA detection based on nanogold-enwrapped graphene nanocomposites, which were used for label of secondary antibodies and for sensitivity enhancement. A working range of 0.05–350 ng mL
1 was shown with an LOD of 0.01 ng mL
1. The precision and reproducibility of immunosensors was investigated through intra- and interassays, respectively. The intraassay precision of the analytical method was evaluated by analyzing four concentration levels and five times per run, and the CV obtained was between 4.1% and 6.5%. Similarly, the CV of the interassay was between 5.3% and 7.8% [48]. The immunosensor provides a promising strategy for clinical applications, since acceptable agreement (99%) was obtained between the levels of CEA with the sensor and with reference values (ELISA), when 54 serum samples were assayed. Gao et al. [49] proposed an amperometric immunosensor for CEA detection, which is based on a uniform nanomultilayer film of CNT, where gold nanoclusters were electrodeposited to immobilize antibodies specific for CEA. Under optimal conditions (pH 6.5; incubation time of 20 min and incubation temperature of 37 °C to antibody immobilization), the proposed immunosensor could detect CEA in two linear ranges from 0.1 to 2.0 ng mL
1 and from 2.0 to 160.0 ng mL
1, with an LOD of 0.06 ng mL
1. Compared with the results obtained by an ELISA, the developed immunosensor showed acceptable accuracy. The main advantage of such immunosensor is the possibility of regeneration by simply immersing in regeneration solution (4 M urea solution or 0.2 M glycine-HCl solution at pH 2.8) for about 5 min followed by a rinse with double-distilled water. An RSD of 3.4% and 4.2% was obtained, respectively, for each regeneration solution (deviation obtained after five regenerations).

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Wang et al. [45] proposed a simple label-free immunosensor for detection of CEA. Glassy carbon electrode surface was first drop-coated with a mixture of thionine (electron mediator) and chitosan, then covered with nafion membrane containing SiO2 nanoparticles, where silver nanoparticles were deposited to facilitate electron transfer process [45]. The label-free immunosensor had a linear range of 1 fg mL
1 to 100 pg mL
1 with a low LOD (1 fg mL
1). When tested without silver nanoparticles, the immunosensor had a narrow linear range (100 fg mL
1 to 100 pg mL
1) and a higher LOD (100 fg mL
1). In order to evaluate CEA levels in real samples, human serum samples were analyzed by the immunosensor and a microparticle enzyme immunoassay. The results obtained with the immunosensor are consistent with the data determined by the microparticle enzyme immunoassay with relative errors between
8% and 5%. The authors referred that together with the low cost and ease of preparation, this immobilization method could be widely applied to clinical research of other important biological molecules [45]. Recently, Lin et al. [50] and Sun et al. [51] proposed two immunosensing strategies to determine the levels of CEA in real serum samples. Lin et al. [50] proposed an electrochemical immunosensor based on a glassy carbon electrode with graphene oxide/ chitosan film and enhanced with nanogold functionalized mesoporous carbon foam to label the antibodies. Under optimal conditions (0.5 mg mL
1 of graphene oxide; silver solutions diluted 20 times; incubation time of 40 min for the saturated binding between the analyte and the capture antibody), the proposed immunoassay method showed wide linear range from 0.05 pg mL
1 to 1 ng mL
1 and an LOD down to 0.024 pg mL
1, which was much lower than those reported previously using nanomaterial-based multienzyme amplification strategies [50]. In addition, acceptable reproducibility (RSD of 7.6%) and stability were obtained since when the immunosensor was stored in dry at 4 °C, over 92% of the initial response was remained after a storage period of 15 days. The assay results of clinical serum samples using the proposed immunosensor were compared with the reference values obtained by commercial electrochemiluminescence tests. The results showed an acceptable agreement with relative errors less than 11.9%, which indicated an acceptable accuracy of the proposed method for the detection of CEA in clinical samples. Regarding the work of Sun et al. [51], a microfluidic paper-based electrochemical immunosensor was fabricated for the first time using novel three-dimensional flower-like gold nanoparticles (which has a large surface area for the immobilization of primary antibodies) modified paper working electrode as the sensor platform

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Figure 9 Fabrication process of the microfluidic-based paper immunosensor. a, cellulose fibers in the paper sample zone; b, 3D flower-like gold nanoparticles; c, screenprinted carbon working electrode; Ab1, primary antibody; Ab2, secondary antibodies; BSA, bovine serum albumin; PWE, paper working electrode. Reprinted from Sun et al. [51], Copyright (2014), with permission from Elsevier.

and gold-silver bimetallic nanoparticles as tracer for binding of signal antibodies (secondary antibodies). Figure 9 shows the fabrication process of the immunosensor. The microfluidic paper-based analytical devices are new nanodevices currently proposed and optimized as point-of-care devices, since they provide quantitative analytical analysis at low cost, and they can be adapted to clinical biosensing [2]. Under the optimized experimental conditions (pH 7.4; incubation time of 200 s and incubation temperature of 37 °C for the immunoreaction), the proposed paper-based electrochemical immunosensor exhibits excellent analytical performance for enzyme-free detection of CEA, ranging from 0.001 to 50 ng mL
1 with a low LOD of 0.3 pg mL
1 [51]. In addition, Sun et al. [51] verified that there was no significant difference (relative errors between
2.2% and 3.2%) between CEA levels of real human serum samples obtained with the immunosensor and a reference methodology (based on electrochemiluminescence), which means that the developed biosensor could be applied in the clinical determination of CEA levels in human serum, providing point-of-care testing applications. All the works reported for the determination of CEA in biological fluids are able to be used in the routine clinical analysis, since the LOD observed is below the normal level of CEA, which is of 3 ng mL
1 (Table 1).

3.2 Cardiac Biomarkers The CVDs are classified mainly as coronary heart disease and cerebral vascular accident being mostly associated with smoking, hypertension, elevated

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Table 2 Normal Concentrations of Cardiac Biomarkers in Blood Cardiac Biomarker Normal Concentration

CRP

1.0–5.0 μg mL


1


1

cTnT

0.01 ng mL

cTnI

0.5–2.0 ng mL
1

Myoglobin NT-proBNP

100 ng mL
1

1 ng mL


1

References

[55] [56] [57] [58] [59]

low-density lipoprotein cholesterol levels, and sedentary habits [52]. According to Packard and Libby [53], the diagnosis of such biomarkers in real-time and with rapid analysis is required for the adequate evaluation and prediction of risk to CVDs. The immunosensors with potential to a point-of-care testing for cardiac biomarkers such as cTnT are highly desirable, since these devices can be applied to cardiology emergencies, allowing a rapid screen of the patients and consequently the best clinical management [54]. Table 2 shows the normal concentrations of cardiac biomarkers in blood. 3.2.1 C-Reactive Protein The CRP is a positive acute-phase protein that levels increases in blood in response to a wide variety of disease states including CVDs and systemic inflammation conditions [60]. Vermeeren et al. [61] proposed an electrochemical immunosensor based on nanocrystalline diamond surface for real-time and label-free CRP detection. Diamond has been studied as a transducer material for biosensor fabrication due to their advantageous physical, optical, chemical, and electrical characteristics, as well as its biocompatibility allowing in vitro applications [61]. The calibration curve ranges from 1 to 1000 nM obtaining an LOD of 10 nM. Sensitivity experiments in real time showed a clear discrimination between 1 μM, 100 nM, and 10 nM of CRP after 10 min, with the concentration at 10 nM of CRP clearly distinguishable from buffer solution; thus, the immunosensor reaches a sensitivity within the physiologically relevant concentration range of this biomarker in healthy controls (8–10 nM) and CVD patients (>10 nM). Gan et al. [62] described a disposable screenprinted immunosensor for rapid determination of CRP in human serum.

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HRP-labeled anti-CRP functionalized Fe3O4/gold magnetic nanoparticles were used as biorecognition probes and attracted to a Fe(III) phthalocyanine/chitosan membrane-modified screen-printed carbon electrode by an external magnetic field. After incubation with CRP antigen solution, the access of the activity center of the HRP electrode was partially inhibited leading to a linear decrease of the catalytic efficiency of the HRP to the reduction of immobilized Fe(III) phthalocyanine by H2O2 in the CRP concentration range from 1.2 to 200 ng mL
1 with an LOD of 0.5 ng mL
1. The immunosensor was used for the determination of this cardiac marker in real serum samples of heart disease patients, with results consistent with those of an ELISA method. It was reusable once constructed and could be regenerated by adding new modified nanoparticles to the surface of ´ vila the basal electrode with the help of a magnet. Esteban-Ferna´ndez de A et al. [63] also proposed a disposable electrochemical immunosensor based on gold screen-printed electrode for CRP quantification using the magnetic principle. The covalent immobilization of the capture antibody was made onto carboxylic-modified magnetic beads and an incubation step with a streptavidin–HRP conjugate was employed to allow the monitoring of the affinity reaction. The CRP magnetoimmunosensor exhibited a wide range of linearity between 0.07 and 1000 ng mL
1 with a low LOD (0.021 ng mL
1), when compared to that obtained by ELISA (0.002–100 ng mL
1). A good reproducibility of the analytical responses was obtained (RSD of 6.5%) with different magnetoimmunosensors constructed following the same protocol [63]. A sensitive label-free and nonamplified electrochemical immunosensor for the determination of CRP in blood has been recently described by Bryan et al. [64]. It was based on controlled and coverage-optimized antibody immobilization on standard polycrystalline Au electrodes. These interfaces were capable of quantifying CRP across a clinically relevant range of concentrations in whole blood serum (60 μg L
1–6.0 ng L
1) with an LOD of 19 μg L
1 and, moreover, they were reusable. The assays could be carried out with only 5 mL of sample and results reported in just 10–15 min, while surface regeneration was achieved in approximately only 5 min by immersion of the used surfaces in 6 mM NaOH with 0.6% ethanol. A different immunosensor, based on piezoelectric principle, was described by Zhou et al. [65]. The piezoelectric immunosensor was based on Fe3O4/SiO2 magnetic capture nanoprobes and HRP–antibody coimmobilized nanogold as signal tags. The change in frequency response was proportional to CRP concentration from 0.001 to 100 ng mL
1 with

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Figure 10 The calibration curve of the immunosensor modified with the different signal tags. (a) HRP–CRP–antibody, (b) gold nanoparticles–HRP–CRP–antibody, and (c) gold nanoparticles–HRP/HRP–CRP–antibody. Reprinted from Zhou et al. [65], Copyright (2013), with permission from Elsevier.

a low LOD of 0.3 pg mL
1. In order to verify the amplification effect of the immunosensor, Fig. 10 shows the calibration curves using the three signal tags tested. As shown in Fig. 10, the use of HRP coupled with the HRP-linked secondary CRP antibody coimmobilized on gold nanoparticles as the signal tag exhibited much larger slope coefficient in calibration curve than the other two labels, which means that the detection signal can be greatly amplified through the introduction of more amount of HRP in the signal tags [65]. Recently, Justino et al. [60] proposed disposable immunosensors based on field-effect transistors with SWCNT (NTFET) for the determination of CRP in clinical samples of undiluted blood serum and saliva. The immunosensor can detect CRP between 10
4 and 102 mg L
1 and the LOD obtained was about 10
4 mg L
1, which is far better by two to three orders of magnitude than the traditional immunochemical assays (LOD from 0.03 to 0.2 mg L
1) [60]. The authors also reported that the LOD is much lower than the clinical relevant range of concentrations in serum (Table 2), and it is better than other measurement systems used for the detection of the CRP based on optical, chemiluminescence, and magnetic sensors [60]. The NTFET showed comparable analytical performance with the ELISA when applied to clinical samples, since there is no statistically significant difference between the NTFET and the ELISA when they are used to determine the CRP levels (p ¼ 0.912). Furthermore, as a strong correlation was found

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between the salivary and the serum CRP both determined by the NTFET and the ELISA, it is suggested that the developed NTFET constitute an alternative for the use of point-of-care testing methodology to detect the risk of developing a CVD, based on a noninvasive sampling [60]. 3.2.2 Cardiac Troponins cTnT and cTnI have been recommended as the biomarkers of choice for the serological diagnosis and prognosis of acute myocardial infarction because of their high sensitivity and specificity [54]. Silva et al. [54] prepared disposable immunosensor for human cTnT, where the surface of screen-printed electrodes was modified with streptavidin polystyrene microspheres in order to enhance the sensitivity. The modified screen-printed electrode was fabricated using rigid conducting graphite–epoxy silver composite. The integration of streptavidin microspheres through glutaraldehyde as homofunctional agent allowed a stable immobilization of biotinylated cTnT antibodies on the electrode surface. The use of silver–epoxy composite for electrode manufacture has the advantage of easily preparing microelectrodes with different shapes and sizes compared to the carbon-paste microelectrode [54]. The use of streptavidin microspheres has significantly increased the analytical sensitivity of the electrode in 8.5 times, showing a curve with a linear response range of 0.1–10 ng mL
1 of cTnT and an LOD of 0.2 ng mL
1. This immunosensor coupled with a portable electrochemical analyzer shows great promise for point-of-care quantitative testing of necrosis cardiac proteins. Also, based on the work of Silva et al. [54], the concentration of cTnT detectable by using the proposed immunosensor meets the requirements of clinical analysis. The method was validated by comparing the results obtained in serum samples with the values given by a Roche Elecsys 2010 immunoassay analyzer based on electrochemiluminescence (RSD of 5.4%). However, more detailed studies with real samples for the direct determination of cTnT concentrations in human serum are necessary to indicate that this system is an alternative approach to cTnT assay in clinical routine and aiming at point-of-care testing development [54]. The same research group has recently described a nanostructured immunosensor using carboxylated CNT supported on a gold electrode by a conductive polymer film of polyethyleneimine (a highly cationic polymer) for the detection of cTnT [66]. A sandwich configuration with anti-cTnT-HRP was employed in the cyclic voltammetric detection of this cardiac biomarker. The nanostructured surface increased the immunoreactive electrode area allowing a response to be obtained within a linear range between 0.1 and 10 ng mL
1,

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with an LOD of 0.033 ng mL
1, which is significant for acute myocardial infarction diagnosis. Fonseca et al. [67] proposed an immunosensor for cTnT based on piezoelectric principle (QCM-flow immunosensor). The immunosensor was fabricated with gold nanoparticles coimmobilized on a dithiol-modified surface, where anti-cTnT antibodies were covalently immobilized by thiol–aldehyde linkages. According to Fonseca et al. [67], the gold nanoparticles have been widely used to enhance the sensitivity of biomolecular detection assays through the improvement of antibody immobilization, since they increase the surface contact of the electrode. The QCM-flow immunosensor exhibited good reliability, measuring concentrations of cTnT from 0.003 to 0.5 ng mL
1 in human serum with an LOD of 0.0015 ng mL
1. The advantages of the immunosensor are its speed and high sensitivity. As the desirable blood level of cTnT is equal or less than 0.01 ng mL
1, the immunosensor proposed by Fonseca et al. [67] is a sensitive method for measuring the concentration of cTnT below the cutoff point for the diagnosis of myocardial infarction and it is faster than conventional methods for the diagnosis of acute myocardial infarction. Recently, Silva et al. [56] have described a label-free immunosensor based on CNT screen-printed electrodes for the detection of cTnT. The disposable device was fabricated by tightly squeezing an adhesive carbon ink containing CNTs onto a polyethylene terephthalate substrate forming a thin film. The analytical response was obtained by differential pulse voltammetry after the interaction of antibody–antigen at the surface of the CNT screen-printed electrodes. The analytical response was the difference between the peak current of the CNT screen-printed electrodes with and without cTnT at +0.25 V versus Ag/AgCl. A rapid detection of cTnT with high sensitivity and an LOD of 0.0035 ng mL
1 were obtained within a linear range of 0.0025 and 0.5 ng mL
1, which demonstrates the potential of the disposable immunosensor to be applied as a tool for point-of-care acute myocardial infarction [56]. The concentration range was comparable to the clinical range, which is associated with a cutoff of 0.01 ng mL
1. Mattos et al. [68] proposed an immunosensor based on an o-aminobenzoic acid film for the detection of cTnT in human serum. According to the results, the calibration curve presented a good linear response range from 0.05 to 5.0 ng mL
1 cTnT with an LOD of 0.016 ng mL
1. In addition, the immunosensor showed a good stability upon the analytical responses retaining 91.6% of its initial response after 18 days, when stored at 4 °C. Regarding the cTnI, Zhou et al. [69] reported clinical applications of an electrochemical immunosensor for simultaneous detection of cTnI and

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CRP based on microfluidic chips. The microfluidic chips were based on a poly(dimethylsiloxane)–gold nanoparticles composite. The CdTe and ZnSe quantum dots were bioconjugated with antibodies specific for cTnI and CRP for the immunoassay. In this work, square-wave anodic stripping voltammetry was used to detect metal ions (Cd2+ and Zn2+) dissolving from the quantum dots in a microchannel through the flow-injection mode. Zhou et al. [69] verified the simultaneous detection of cTnI and CRP in linear range of 0.01–50 and 0.5–200 μg L
1, with LOD of 0.004 and 0.22 μg L
1, respectively. Other advantages are the excellent selectivity and the absence of nonspecific adsorption effects. In another work [57], gold nanoparticles were electrodeposited on indium tin oxide and applied to detect molecular interaction between human cardiac cTnI and specific antibody by measuring the open-circuit potential as a new detection principle. In this work, a new strategy has been adopted to obtain an electrical signal due to the enzyme-based immune catalytic reaction by measuring the changes of open-circuit potential in the range of 1–100 ng mL
1 of cTnI. Such electrochemical sensor has great potential to be applied in clinical diagnostic due to the detection of cTnI in the clinical borderline of cTnI in normal and disease patients. Recently, Singal et al. [70] proposed an immunosensor for cTnI fabricated with a composite film of platinum nanoparticles and graphene deposited on a glassy carbon electrode. The immunosensor exhibited a linear response to cTnI over the concentration range of 0.01–10 ng mL
1 with LOD of 4.2 pg mL
1 cTnI. In addition, the authors referred that the stability of the immunosensor was examined by monitoring its performance after 30 days storage in a refrigerator at 4 °C, retaining about 96% of its initial sensitivity for 1.0 ng mL
1 cTnI detection, which indicates a good shelf-life stability. According to Singal et al. [70], the results showed with the immunosensor indicate that the graphene–platinum composite may be useful for further clinical research for cTnI detection. 3.2.3 Myoglobin The main advantage of myoglobin as a cardiac marker is that it is released earlier from damaged cells than other cardiac markers, allowing early detection of acute myocardial infarction [71]. Suprun et al. [58] have reported an electrochemical immunosensor based on metal nanoparticles (gold, silver, and copper) for cardiac myoglobin detection in human blood plasma. The detection of cardiac myoglobin was based on direct electron transfer between the Fe(III)-heme and the transducer surface (gold nanoparticles-modified

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graphite electrode) which was stabilized by the specific antibodies. Although the method does not require signal enhancement or amplification, the electrode needs a 15-min immunoreaction time with the target myoglobin sample, followed by further 5 min incubation in phosphate-buffered saline (PBS) and washing of nonspecifically binding molecules before electrochemical measurement. Low LOD of 5 ng mL
1 was obtained in a broad range of concentrations (10–400 ng mL
1). In addition, the immunosensors are portable and cheap with potential to be applied in the diagnosis of acute myocardial infarction. Recently, Zapp et al. [71] proposed a label-free immunosensor (glassy carbon electrode) based on ionic liquid crystal and gold nanoparticles for the electrochemical detection of myoglobin. The immunosensor displays a linear response over the range of 9.96–72.8 ng mL
1 with an LOD of 6.29 ng mL
1 and a good reproducibility (RSD of 4.3%). The authors referred that the main aim of the work was the investigation of new materials for electrochemical sensors, contributing to their simplicity in design state, good precision, high accuracy, and good selectivity for myoglobin detection. Recently, Kim et al. [72] have reported the continuous immunosensing of myoglobin in human serum with a surface plasmon resonance (SPR) immunosensor. According to Kim et al. [72], the continuous monitoring of biomarkers in biological samples is possible if an antibody specific to the marker is combined with a label-free sensor such as SPR sensor, which eliminates separation of the binding complexes. The continuous sensor reproducibly traced the varying doses of myoglobin over about 8 h with periodic one-point calibration every 3 h [72]. Acceptable variations of the analytical response were reported with CV of 4.91% in average, and an LOD of 31.0 ng mL
1. The continuous biosensor can provide an early warning sign on, for example, cell phone as soon as the concentration exceeds a predetermined cutoff [72]. According to Kim et al. [72], the immunosensor could be used as a companion diagnostic technique with real-time electrocardiographic measurement, significantly enhancing the sensitivity of acute myocardial infarction diagnosis and thereby enabling treatment at an early stage. 3.2.4 Amino-Terminal Pro-B-type Natriuretic Peptide The cardiac biomarker NT-proBNP is an important marker for heart failure that reflects ventricular volume expansion, ventricular overload, and the degree of cardiac injury [73]. The use of immunosensors can facilitate the point-of-care testing and undertake molecular analysis without the need for state-of-the-art laboratories [73].

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Some electrochemical immunosensors have been developed for the determination of NT-proBNP involving sandwich formats. Capture antibodies were immobilized on nanostructured gold and CNT composite platforms and bioconjugates of gold nanochains, while HRP-labeled secondary antibodies were employed for signal amplification [74]. The detection of NT-proBNP was based on cyclic voltammetry and the linear range was from 0.02 to 100 ng mL
1 with an LOD reaching 6 pg mL
1. The specificity, regeneration, and stability test (during 30 days) demonstrated the feasibility of the developed immunoassay, which gives the attractive characteristics to be a candidate for the detection of NT-proBNP and other proteins of interest in both fundamental and applied research [74]. A regeneration-free electrochemical immunosensor with only a single antibody/antigen pair (anti-NT-proBNP and NT-proBNP) has been proposed for the detection of NT-proBNP in serum samples by Yi et al. [75]. Biotinylated fragments of monoclonal antibodies specific to NT-proBNP and avidin-modified magnetic nanoparticles were used for such detection. The sensor detected NT-proBNP from 0.04 to 2.5 ng mL
1 with an LOD of 0.03 ng mL
1. The immunosensor is therefore a simple, costeffective method to detect NT-proBNP, and the proposed immunoassay system would also enable other proteins to be detected and open new opportunities for protein diagnostics. An amperometric magnetoimmunosensor for the sensitive detection of NT-proBNP using an indirect competitive configuration has recently been described by Esteban-Ferna´ndez de A´vila [73]. The antigen was covalently immobilized onto activated carboxylic-modified magnetic beads and further incubated in a mixture solution containing variable concentrations of the antigen and a fixed concentration of an HRP-labeled detection antibody. The NT-proBNP in the sample (human serum) competed for binding to a fixed amount of the specific HRP-labeled secondary antibody. The immunoconjugate-bearing magnetic beads were captured by a magnet under the surface of a disposable gold screen-printed electrodes. The amperometric response was used to monitor the affinity reaction. The magnetoimmunosensor showed an excellent analytical performance for the determination of this cardiac biomarker at the concentration levels clinically relevant in human serum (0.12–42.9 ng mL
1) and an LOD of 0.02 ng mL
1, which can be used in clinical diagnosis of chronic heart failure in the elderly and for classifying patients at risk of death after heart transplantation.

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More recently, a disposable multiplexed electrochemical magnetoimmunosensor has been reported by the same research group [59] for the simultaneous determination of NT-proBNP and CRP in human serum. The quantification of NT-proBNP and CRP was performed by using an indirect competitive and a sandwich assay configuration, respectively. The use of dual screen-printed carbon electrodes allowed simultaneous independent amperometric readouts for each cardiac biomarker using capture antibodies (anti-CRP) and antigen-(NT-proBNP) modified magnetic beads for the determination of CRP and NT-proBNP, respectively. Figure 11 shows the picture of screen-printed carbon electrodes (center) and the schematic display of the fundamentals involved in the development of the disposable dual magnetoimmunosensor for the simultaneous determination of NT-proBNP (left) and CRP (right). Very low LOD was achieved (0.47 ng mL
1) and the dual magnetoimmunosensor also exhibited excellent analytical performance in terms of selectivity together with a wide range of quantifiable antigen concentrations (2.0–100 ng mL
1). The whole multiplexed assay for the two cardiac biomarkers could be completed in approximately 60 min once the anti-CRP magnetic beads and NT-proBNP magnetic beads were prepared. Despite the big difference between the clinically relevant concentration ranges for these two cardiac markers (1–5 μg mL
1 for CRP and 1 ng mL
1 for NT-proBNP), the developed dual magnetoimmunosensor allowed the

Strep–HRP Anti-NT-proBNP–HRP NT-proBNP

Biotin–anti-CRP CRP

HOOC-MBs Anti-CRP

HOOC-MBs

Figure 11 Picture of screen-printed carbon electrodes (center) and schematic display of the fundamentals involved in the development of the disposable dual magnetoimmunosensor for the simultaneous determination of NT-proBNP (left) and CRP (right). HOOC-MBs, carboxylic acid-modified magnetic beads. Adapted from Esteban-Fernández de Ávila et al. [59], Electroanalysis, doi:10.1002/elan.201300479. Copyright © (2014).

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simultaneous determination to be performed in a single assay, without need for splitting or using a different sample dilution for each quantification. According to the authors, the low cost, easy automation, and miniaturization of the employed instrumentation, together with the use of disposable mass-produced electrodes make the developed approach a promising, attractive, and user-friendly alternative diagnosis tool for the development of point-of-care devices for on-site clinical diagnosis [59].

3.3 Hormones Various immunosensors have also been used in the determination of hormones in human samples, such as cortisol [76,77], human growth hormone (hGH) [78], and estradiol [79]. The cortisol is a steroid hormone found in blood, saliva, urine, and interstitial fluids. It is also considered as a biomarker for numerous diseases and is important for the regulation of blood pressure, glucose levels, and carbohydrate metabolism, within the physiological limit; abnormal increase in cortisol level inhibits inflammation, depresses immune system, and increases fatty and amino acid levels in blood [77]. Arya et al. [77] have fabricated an immunosensor for cortisol detection based on gold microelectrode arrays, which were functionalized with a self-assembled monolayer of dithiobis(succinimidyl propionate). The formation of self-assembled monolayer has the advantage of immobilization of biomolecules at the surface of electrodes, which allow enhanced sensitivity, fast response, low cost, and portability of electrochemical devices [77]. The cortisol-specific monoclonal antibody was covalently immobilized on the surface of gold microelectrode array in order to determine cortisol concentration. The immunosensor was applied to human interstitial fluids and the measurements were compared to ELISA. Arya et al. [77] found that the immunosensor enabled cortisol detection between 1 pM and 100 nM within the 40 min of analysis time. The deviation in measured values of cortisol with the immunosensor from ELISA values was between 0.3% and 14.1%. Another immunosensor was fabricated for cortisol detection but using a different type of transduction. Moreno-Guzman et al. [76] have employed functionalized magnetic particles (with protein A) to fabricate a disposable immunosensor based on screen-printed carbon electrodes. Concentrations of cortisol between 0.005 and 150 ng mL
1 were determined obtaining an LOD of 3.5 pg mL
1. Real samples of human serum were tested in order to demonstrate the usefulness of the immunosensor, which can be further used for

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rapid determination of cortisol in sport medicine and other clinical applications [76]. Recently, Pasha et al. [80] developed an electrochemical immunosensor for detection of cortisol in saliva samples. Apart its role in screening psychological stress and health monitoring, the detection of cortisol in saliva is considered as an important screening tool for the diagnosis of Cushing syndrome, Addison disease, and posttraumatic stress disorder [80] and the development of rapid devices can be useful for the point-of-care detection of cortisol. In the work of Pasha et al. [80], microfabricated interdigitated microelectrodes were used as electrochemical platform and anticortisol antibodies were covalently immobilized on a dithiobis(succinimidylpropionate) SAM. Through cyclic voltammetry, the sensor exhibited a detection range from 10 pg mL
1 to 100 ng mL
1 and an LOD of 10 pg mL
1, which is below the typical cortisol concentration in saliva of healthy adults (1–8 ng mL
1) [80]. Pasha et al. [80] verified that obtained cortisol concentrations from the electrochemical immunosensor are correlated with those obtained using ELISA (within 2–5%). Another electrochemical immunosensor was proposed by Vabbina et al. [81] for the detection of cortisol and based on sonochemically synthesized ZnO, under two morphologies (1D nanorods and 2D nanoflakes). The selective detection of cortisol was processed through cyclic voltammetry and achieved by immobilizing anticortisol antibody on the ZnO nanostructures. Vabbina et al. [81] obtained an LOD of 1 pM with the immunosensor, which is 100 times better than conventional ELISA. The immunosensor was tested on real saliva samples and the levels of cortisol obtained with immunosensor were compared with those of an enzyme immunoassay, with a good agreement between them (R2 of 0.9997). According to the authors, the developed sensors can be integrated with microfluidic system and miniaturized potentiostat for point-of-care detection of cortisol. The hGH is a polypeptide hormone synthesized and secreted by the anterior pituitary gland, and it is essential for body growth since it stimulates the production of insulin growth factor (IGF-1), which in turn stimulates the production of cartilage cells, resulting in bone growth [78]. The determination of hGH is necessary to diagnose various disorders occurring in childhood and the existence of pituitary tumors [78]. Serafı´n et al. [78] proposed the preparation and application of an electrochemical magnetoimmunosensor for the determination of hGH, where tosyl-activated magnetic microparticles were used to immobilize the antibodies specific for hGH. Concentrations of hGH between 10
3 and 103 ng mL
1 were used to construct the calibration plot, with linear range observed between 0.01

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and 100 ng mL
1. An LOD of 0.005 ng mL
1 was obtained, which is lower than that obtained with a reference methodology based on SPR (LOD of 4 ng mL
1) and other immunoassay methods (LOD between 0.01 and 0.03 ng mL
1) [78]. The estradiol is a natural estrogen with particular impact on the reproductive and sexual functioning, and its quantification in serum and urine is important in investigations of fertility treatments, postmenopausal status, hyperandrogenism, and breast cancer [79]. Ojeda et al. [79] proposed the preparation of an amperometric immunosensor for estradiol based on the surface modification of a screen-printed carbon electrode with grafted p-aminobenzoic acid followed by covalent binding of streptavidin and immobilization of biotinylated anti-estradiol. The calibration curve for estradiol exhibited a linear range between 1 and 250 pg mL
1 and an LOD of 0.77 pg mL
1 was achieved [79]. The authors have applied the immunosensor to the analysis of clinical samples such as human serum and urine with good results (recoveries ranged from 96% to 102%).

3.4 Pathogenic Bacteria Clinical immunosensors have been developed for detection of pathogenic bacteria, mainly in urine samples. For example, Yang et al. [82] have separated and detected E. coli in a microfluidic channel in order to detect urinary tract infections. It is known that E. coli is responsible for up to 80% of such infections [82]. The traditional methods of detection include 1 or 2 days of cultivation and labor-intensive procedures; dip-stick methods for fast E. coli detection are also available but they do not provide sufficient sensitivity [82]. The lab-on-a-chip device consists of two chambers for concentration and sensing connected in series and an integrated impedance detector. The concentration chamber contains microsized magnetic beads conjugated with anti-E. coli antibody to separate E. coli from the urine sample. The immobilized E. coli is transferred to a sensing chamber for the measurement. Yang et al. [82] obtained a clear distinction of signal between 104 and 105 CFU mL
1, which covers the threshold concentration known typically considered for urinary tract infection (105 CFU mL
1). The sensitivity of the lab-on-a-chip device is at least 104 CFU mL
1. Thus, the authors considered that the obtained data show promising potential for application in portable lab-on-a-chip devices for detection of urinary tract infection [82]. Ahmed et al. [83] proposed an electrochemical immunosensor for detection of pathogenic bacteria S. pyogenes in human saliva. Commercial screen-

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printed gold electrodes were used as sensing platform and polytyramine was electrodeposited for immobilization of biotin-tagged whole antibodies against S. pyogenes. Ahmed et al. [83] reported a detection range of 100–105 cells per 10 μL and 100–104 cells per 10 μL of bacteria in PBS, respectively. This range covers the pathogenic load of S. pyogenes infection, which is around 106 cells mL
1. Sensors were also able to specifically detect S. pyogenes in 50% (v/v) human saliva in PBS with good selectivity. Recently, Barreiros dos Santos et al. [84] fabricated a label-free indium tin oxide-based immunosensor for the detection of very low concentrations of pathogenic bacteria (E. coli). Anti-E. coli antibodies were immobilized onto indium tin oxide electrodes using a simple, robust, and direct functionalization methodology (through silane monolayers using trifunctional silanes). A very low LOD was obtained (1 CFU mL
1) over a large linear working range (10–106 CFU mL
1). Barreiros dos Santos et al. [84] proposed that the functionalization of indium tin oxide is a potential alternative for the development of highly sensitive and selective immunosensors for other pathogenic bacteria.

3.5 Virus The diagnostics of diseases can also be attained through the application of immunosensors, for example, for the determination of dengue virus, hepatitis C virus, and avian influenza virus in human blood. 3.5.1 Dengue Virus Dengue disease is caused by females of Aedes spp. mosquitoes infected with RNA-containing dengue virus and it is the most important arthropodborne disease in the present, mainly in tropical and subtropical regions of the world [85]. Fang et al. [86] prepared a label-free immunosensor for diagnosis of dengue infection with simple electrical measurements with the following steps: (a) the preinactivated dengue virus (sensing probe) was firstly immobilized onto the immunosensor surface, precoated with the sol–gel-derived barium–strontium–titanate thin film over the interdigitated electrodes. The modified sensor surface served as selective sensing probe to capture/conjugate the dengue antibody molecules present in patient serum and (b) direct correlation was obtained between the signal outputs with respect to serum concentrations of dengue antibody; the measured signal changes in impedance/current without/with the presence of dengue antibody were attributed to the surface conductivity change upon biomolecules immobilization. By monitoring the impedance or current

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change, the antibody molecules in the patient serum could be positively detected [86]. Fang et al. [86] reported that the promising sensing response results (ability to detect dengue antibody in human serum even after 50,000 times of dilution) indicate that the immunosensor has the potential to be developed into a hand-held device for clinical point-of-care screening or molecular diagnosis of dengue infection in the field. In another work, a label-free and real-time SPR-based immunosensor for serological diagnosis of dengue virus infection was employed by measuring an increase in the resonance angle of the surface-deposited sample in the presence of virus [87]. SAM-based covalent immobilization of a bioreceptor conjugate of a dengue antigen and bovine serum albumin (BSA) was performed onto a gold chip surface. The regeneration was achieved by pepsin solution in glycine-HCl buffer (pH 2.2) and sensor surface displayed a high level of stability during repeated immunoreaction cycles. The proposed biosensor was simple, effective, and based on natural antigen–antibody affinity, contributing to the development of biosensors for diagnosis of dengue and dengue hemorrhagic fever which continues to be a major health problem in various regions of world [88].

3.5.2 Hepatitis C Virus The hepatitis C virus is the major causative agent of chronic viral hepatitis which can develop into cirrhosis and hepatocellular carcinoma [88]. Electrochemical immunosensors have created a high interest in clinical diagnostics and will be expected to provide fast and highly sensitive detection of hepatitis C virus core antigen [88]. Ma et al. [88] reported the construction of a label-free electrochemical immunosensor for detecting the core antigen of the hepatitis C virus. A glassy carbon electrode was modified with a nanocomposite made from gold nanoparticles, zirconia nanoparticles, and chitosan. The sandwich-type immunosensor displayed high sensitivity to the hepatitis C virus core antigen in the concentration range of 2–512 ng mL
1, with an LOD of 0.17 ng mL
1. This immunosensor exhibited a wide linear range, good stability (30 days), good reproducibility (RSD of 4.2%), and high sensitivity for the detection of hepatitis C virus core antigen. The assay was convenient and cost-effective and had promising potential for the early clinical diagnosis of hepatitis C virus infection. This immunosensor provides an alternative approach toward the diagnosis of hepatitis C virus. The same research group [89] has proposed another

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ultrasensitive and selective electrochemical immunosensor for the detection of hepatitis C virus core antigen. This one consists of graphitized mesoporous carbon–methylene blue nanocomposites as electrode-modified material and HRP–DNA-coated carboxyl MWCNT as a secondary antibody layer. After modification of the electrode with the nanocomposite, gold nanoparticles were electrodeposited onto the electrode to immobilize the captured antibodies. The reduction current of methylene blue was generated in the presence of hydrogen peroxide and it was monitored by squarewave voltammetry. The amperometric signal increased linearly with the core antigen concentration (range of 0.25–300 pg mL
1). The immunosensor had an LOD as low as 0.01 pg mL
1, and showed high selectivity. The new protocol presented acceptable stability (10 days) and reproducibility (RSD between 2.8% and 5.4%), as well as favorable recovery (95–106%) for hepatitis C virus core antigen in human serum. The proposed immunosensor has considerable potential for use in clinical applications and provides a promising universal multi-HRP–DNA–carboxyl MWCNT label for different analytes [89].

3.5.3 Avian Influenza Virus Pathogenic influenza virus has caused international human infections, spreading easily by air transmission, and infection through the respiratory system is quickly acquired [90]. Thus, rapid and reliable analytical methods have been requested, and other than diagnostic test kits or ELISA, which are time consuming, expensive, or require a laboratory and a trained technician [90]. A magnetic bead-based bienzymatic electrochemical immunosensor was fabricated for the determination of H9N2 avian influenza virus [91]. In such work, the bienzymatic strategy was proposed by using the first enzyme as tracer tagged on immunomagnetic beads which could be accumulated on the magneto-gold electrode and the second enzyme was immobilized on the gold electrode by layer-by-layer assembly technique. H9N2 avian influenza virus was captured using immunomagnetic beads based on a sandwich-type immunoassay and then detected on the HRP-modified magneto-gold electrode [91]. The linear range of this method was from 0.05 to 2 μg mL
1 with an LOD of 1 ng mL
1. According to Zhou et al. [91], the reported immunosensor displayed a rapid detection of H9N2 virus (1 h) with high sensitivity and good reproducibility (RSD of 4.8%) and could be used in complex samples such as serum samples,

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and thus expanded to other multienzymatic amplification systems in order to construct more sensitive immunosensors for on-site medical diagnostic application. Recently, Singh et al. [90] proposed a label-free immunosensor based on dielectrophoretically deposited SWCNT for detection of influenza virus H1N1. The resistance of the SWCNT-based electrode channels increased with the binding of the influenza virus to the antibodies, and the immunosensors showed a linear behavior when the virus concentration varied from 1 to 104 PFU mL
1 along with a detection time of 30 min [90]. According to the authors, the SWCNT-based immunosensor has potential applications in a point-of-care test kit for rapid and simple clinical diagnosis or a component of a portable lab-on-a-chip system [90].

4. FINAL REMARKS AND PERSPECTIVES Electrochemical immunosensors are the class of immunosensors most used in clinical diagnostics as point-of-care devices, since they are portable, simple, easy to use, cost-effective, and disposable in most cases. Studies of electrochemical immunosensors for cardiac and cancer biomarkers have demonstrated the potential usefulness of these approaches in public health applications. In particular, the sensitivity of biomarkers is important for early diagnostics of diseases. The nanotechnology applied to biosensors has improved the methods for construction of these devices, with miniaturization increasing portability, accuracy, and reliability. An ideal electrochemical biosensor should be both integrated and highly automated in order to improve the efficiency of diagnostic testing. In addition, the use of a variety of unique nanomaterials can maximize the detection capabilities by improving the analytical performance and increase the stability of the biosensors useful for detection of biomarkers at extremely low concentrations. It should be highlighted that the immunosensors applied in clinical diagnosis are in majority employed as prototypes, which were developed and tested in laboratory conditions but the final goal is their use as point-of-care technologies. Thus, future research should be directed in optimizing the stability of immunosensors and eliminating undesirable interferences in order to constructed devices applicable for commercial use. Moreover, the integration of multiple devices on a single disposable platform should lead to significant advantages in terms of cost and speed of the detection of a specific disease.

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ACKNOWLEDGMENTS This work was funded by Portuguese Science Foundation (FCT) through scholarships (ref. SFRH/BPD/95961/2013) under QREN-POPH funds and cofinanced by the European Social Fund and Portuguese National Funds from MEC. The authors also acknowledge Portuguese National Funds through FCT/MEC (PIDDAC) under project IF/00407/2013/CP1162/CT0023 and by FCT under project UID/AMB/50017/2013.

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