1

7. Prussian Blue and analogues: Biosensing applications in health care

Salazar P1,2

, Martín M1,3

, O’Neill RD4, Lorenzo-Luis P

5, Roche R

1,2, and González-

Mora JL1

1Neurochemistry and Neuroimaging group, Faculty of Medicine, University of La Laguna,

Tenerife, Spain. 2Informática y Equipamiento Médico de Canarias S.A., Tenerife, Spain

3Atlántica Biomédica S.L., Tenerife, Spain.

4UCD School of Chemistry and Chemical Biology, University College Dublin, Belfield, Dublin

4, Ireland. 5Department of Inorganic Chemistry, Faculty of Chemistry, University of La Laguna, Tenerife,

Spain.

Abstract:

Prussian Blue (PB), Fe4[Fe(CN)6]3, belongs to a transition metal hexacyanometallate

family. Its electrochemical properties were revealed in 1978 when Neff reported the

successful deposition of a thin layer on platinum foil. After that, numerous publications

have appeared exploring its electrocatalytic properties and its applications in

biomedical science. During the last decade, a great number of studies involving PB

have appeared, using different biosensor substrates and different oxidase enzymes.

Together with the facile modification of the electrode substrate and the low cost of

production, this has led to an on-going replacement of the more common enzymatic

detection method involving horseradish peroxidase. Its high electrocatalytic activity

and low operating overpotential have contributed to the diversification of its use in

enzyme-based biosensors and immunosensors. Based on these results, it is clear that PB

and its analogues will have important roles in the future development of biomedical

devices for next generation health-care strategies.

Keywords: Prussian Blue, biosensors, immunosensors, health care.

Table of Contents:

7.1. Introduction

7.2. General aspects of Prussian Blue and other hexacyanoferrates

7.2.1. Overview

7.2.2. Chemical and structure of Prussian Blue and its analogues

7.2.3. pH stability and deposition method

7.3. Prussian Blue: hydrogen peroxide electrocatalysis

7.4. Prussian Blue: Biosensor applications

7.4.1. Prussian Blue and analogues enzyme system

7.4.1.1. Glucose oxidase

7.4.1.2. Lactate oxidase

7.4.1.3. Cholesterol oxidase

7.4.1.4. Alcohol oxidase

7.4.1.5. NADH oxidase

7.4.1.6. Diamine oxidase

7.4.1.7. Choline oxidase

7.4.1.8. Acetilcholinesterase and butyrylcholinesterase

7.5. Prussian Blue: Immunosensor applications

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7.5.1. α-fetoprotein antigen

7.5.2. Carcinoembryonic antigen

7.5.3. Carbohydrate antigen 19-9

7.5.4. Neuron-specific enolase antigen

7.5.5. Carcinoma antigen 125

7.5.6. Human chorionic gonadotropin antigen

7.5.7. Prostate specific antigen

7.5.8. Hepatitis B antigen

7.6. Conclusions

Date: 2013-06-07

7.1. Introduction

According to the International Union of Pure and Applied Chemistry (IUPAC) a

biosensor is defined as a self-contained integrated device, which is capable of providing

specific quantitative or semi-quantitative analytical information using a biological

recognition element, which is retained in direct spatial contact with a transduction

element [1] (see Figure 1). In electrochemical approaches to biosensor design, the

chemical reactions produce or consume ions or electrons which in turn cause some

change in the electrical properties of the solution and/or transducer surface. There is a

growing demand for new analytical devices for health applications, especially highly

selective and non-invasive methods. In this context, biosensors are ideally small and

portable devices, allowing the selective quantification of chemical and biochemical

analytes. Today, they are replacing other, more sophisticated, techniques such as

chromatographic and spectroscopic methods, and biosensors are already an important

tool in health-care applications [2-3]. Taking account of the general definition of health

care (diagnosis, treatment, and prevention of disease, illness, injury, and other physical

and mental impairments), biosensors may be applied from prevention and early

diagnosis to the progression and monitoring of treatment. Another important advantage

in biosensor design is that different transduction elements may be used such as

electrical, optical, mass, thermal, etc. [1]. Nevertheless, electrochemical biosensors

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(which will be discussed in this chapter) are particularly attractive due to their many

advantages over other detection methods. These benefits include high sensitivity and

selectivity, low cost, real-time output, simplicity of starting materials, possibility to

develop user-friendly and wireless integrated devices, and ready-to-use biosensors. The

last two advantages enable even non-qualified patients to measure and control their own

metabolite levels at home.

Undoubtedly, biosensors for detecting glucose have received more attention than other

devices due to the significance of diabetes as a global health problem, which has

generated considerable interest in the development of an efficient glucose biosensor. In

2000 it was estimated that about 171 million people worldwide were diabetic and that

this value will reach 370 million in 2030 [4]. It is not surprising, therefore, that there are

a great number of commercial devices for determining glucose [5]. Nevertheless, it is

also possible to find commercial biosensors for other metabolites, such as cholesterol,

lactate, urea, creatine, uric acid, ascorbic acid, choline and glutamate, related to

emerging health issues [2-3].

Figure 7.1 Schematic display of different biosensor configurations, illustrating different recognition

materials being coupled to different signal transducers.

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Although biosensors emerged in the early 1960s, when Clark and Lyons [6] coupled an

enzyme (glucose oxidase) to an amperometric electrode for detecting O2, and Prussian

Blue (PB) electrochemical properties have been known since the late 1970s [7], the first

work on biosensors involving the use of a PB-modified electrode was not reported until

1994 [8]. Due to its high activity and selectivity towards H2O2 reduction, PB has been

called an “artificial enzyme peroxidase” [9-10]. During the last decades, a great number

of studies involving PB have appeared using different biosensor substrates (carbon

paste, screen-printed electrodes, glassy carbon, etc.) and oxidase enzymes (glucose

oxidase, lactate oxidase, glutamate oxidase, etc.) [11]. Together with the facile

modification of the electrode substrate and the low cost of production, this has led to an

on-going replacement of the more common enzymatic detection method involving

horseradish peroxidase, which is more expensive, with inferior temporal stability and is

more complicated than PB-modified substrates. In addition, other useful PB properties

such as non-toxicity, high electrocatalytic activity and low operating overpotential have

contributed to the diversification of its use in enzyme-based biosensors and label-free

immunosensors.

Figure 7.2 Detection scheme for a glucose biosensor based on a Prussian Blue (PB) modified electrode.

Glucose is converted to gluconolactone, catalyzed by glucose oxidase (GOx) immobilized on the

electrode surface. Secondary to this reaction is the production of hydrogen peroxide that can be reduced

amperometrically at low applied overpotentials, electrocatalyzed by the PB.

5

More recently, new applications for PB and its analogues [12] have appeared in the

literature for a variety of important analytes in biomedical fields, mainly due to their

electrocatalytic properties. In this way, new applications for the determination of

biomolecules [12] such as dopamine, epinephrine, norepinephrine, morphine, cysteine,

methionine, thiocholine and other thiols, ascorbic acid, nitric oxide, nitrite, isoprenaline,

and vitamin B-6 have already been described (however, in the present chapter we

discuss only biosensing and immunosensing applications). Based on these results, it is

clear that PB and its analogues will have important roles in the future development of

biomedical devices for next generation health-care strategies.

7.2. General aspects of Prussian Blue and other hexacyanoferrates

7.2.1. Overview

PB can be considered as the oldest known hexacyanometallate compound, and has

attracted the attention of many scientists seeking an understanding of its formation,

composition, structure and physical properties. The linking of two different metal ions

by the cyanide ligand is the basis of these properties. In fact, the question of cyanide–

isocyanide isomerism is as old as the discussion of its structure [13, 14], and since then

the electronic interactions between two metals across a cyanide bridge have proven to

be a fertile area of research [11, 15-18].

7.2.2. Chemical and structure of Prussian Blue and its analogues

The cyanide ligand (CN–) is a reactive ligand in important organometallic catalytic

reactions, and is an ancillary ligand in coordination and bioinorganic chemistry. Like

carbon monoxide, the cyanide ion can function as a -acid ligand, but because of its

negative charge, the cyanide ion can also form strong -bonds. This behavior allows

CN– to stabilize both high and low oxidation states of metals. The cyanide ion can bind

6

to metals in both terminal and bridging M-CN-M´ modes; the bridges are commonly

linear, and are present in many polymeric metal types of cyanide [19], and in particular

in PB [20]. Depending on the specific conditions of the preparation, several methods

have typically been used to prepare these cyanide complexes. Addition of [Fe(CN)6]3-

to

Fe2+

aq gives the deep blue complex Turnbull´s Blue (TB), while if [Fe(CN)6]4-

is added

to aqueous Fe3+

, the deep blue complex PB is produced [14, 19]. Both PB and TB are

hydrated salts of formula Fe4III

[FeII(CN)6]3·xH2O (x 14), and related to them is

KFe[Fe(CN)6] – soluble PB [11]. As depicted in Figure 7.3, the zeolite-like structure

possesses extended lattices containing cubic arrangements of Fen+

centres linked by CN–

bridges; each Fen+

(high- and low-spin) is an octahedral anionic building block

[Fen(CN)6]

n-6 with a cubic unit cell of 10.2 Å along the Fe

III-NC-Fe

II-CN-Fe

III-sequence

[13,14]. The selective diffusion of low molecular weight molecules (such as O2 and

H2O2) and some ions with small hydrated radius (such as Cs+, K

+ and NH4

+) is due to its

channel diameter of about 3.2 Å [11]. Consequently, the degree of hydration, as well as

the size of ion, are basic factors for the diffusion of ions through the channels of the PB

lattices [15].

Figure 7.3. One-eighth of the unit cell of KFe[Fe(CN)6] soluble Prussian blue. The K+ ions

and the remaining interstitial or zeolitic water in the cubic sites have been omitted for clarity

from the scheme

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7.2.3. pH stability and deposition method

The chemical literature reports that to achieve a regular structure of electro-deposited

PB, two main factors have to be considered: the pH of the initial growing solution and

the deposition potentials [11, 12]. In this way, pH stability of PB film seems to be

dependent also on the different modes of deposition of the PB layer. For this reason, the

solution pH is a critical point not only during deposition, but also during its applications

in real samples. The reason for this behavior has been ascribed to the strong interaction

between the Fe3+

from PB and hydroxide ions (OH−) which form Fe(OH)3 at pH higher

than 6.4, thus leading to the solubilization of the PB film [11, 15]. For the second factor,

the applied potential should not be lower than 0.2 V vs SCE, where ferricyanide ions are

intensively reduced (vide infra). Different strategies have been described in order to

obtain stable PB films, such as galvanostatic or cyclic voltammetric methods, chemical

deposition or PB microparticle synthesis. In this context, cyclic voltammetric methods

and a heat-treatment step are commonly used to activate and stabilize the PB film,

respectively [11].

7.3. Prussian Blue: hydrogen peroxide electrocatalysis

Although H2O2 is electrochemically ambivalent in that it can be either oxidized to

molecular oxygen or reduced to hydroxide ions, depending on the applied potential used

[21], the former (anodic) mode of electroactivity has been by far the more common

approach for the detection of enzyme-generated H2O2 in first-generation biosensors

[22]. However, an intrinsic problem associated with the relatively high applied anodic

potentials needed to oxidize H2O2 efficiently on most electrode materials (0.4 to 0.7 V

vs SCE [23,24]) is that many substances, including ascorbic and uric acids, in the

biosensor target medium (blood, fat, neural tissues, etc.) also oxidize at these potentials,

thus interfering with the biosensor signal. In recent decades, one strategy being explored

8

to limit this interference is the use of PB and its analogues to electrocatalytically reduce

H2O2 at mild applied potentials (~0 V vs SCE).

In 1984, Itaya et al. [7] showed that the reduced form of PB (Prussian White, PW; see

Figure 7.4) displayed catalytic activity for the reduction of O2 and H2O2. The zeolite

structure of PB, with its small channel diameter (see Section 7.2.2), allows the diffusion

of low molecular weight molecules (such as O2 and H2O2) through the crystal structure

[11]. Nowadays its electrochemical behavior is well understood with cyclic

voltammograms (CVs) of PB-modified electrodes showing two quasi-reversible redox

couples [11,25] (Figure 7.4). The first peak pair corresponds to the interconversion of

PW and PB forms, and the second pair from PB to Prussian Green (PG). Also shown in

Figure 7.4 are the electron transfer reactions in the presence of potassium chloride as

supporting electrolyte, with corresponding formal electrode potentials at ~0.1 V and

~0.85 V vs SCE, respectively [12].

Electrochemical properties such as electrode potential, sensitivity, stability and electron

transfer rate constants of the PB/PW and PB/PG conversions depend on deposition

method, pH, nature and concentration of the supporting electrolyte, etc. As illustrated in

Figure 7.4, these reduction and oxidation reactions involve diffusion of cationic and

anionic species, respectively, through the PB matrix, and the size of the ionic

constituents of the solution medium exerts a major influence on the electrochemical

properties of the PB layer [16,17].

The main drawback of PB as an electrocatalyst for peroxide reduction is its gradual

degradation in solutions with pH values close to neutrality [26]. PB is unstable in

alkaline solutions, so OH– formed during peroxide reduction (Figure 7.4) can cause a

loss of electrocatalytic activity. Because of the prevalence of neutral media in the

context of sensor applications in biological environments, several strategies have been

9

used to improve the stability of PB films in this pH region. For example, enhanced

stability has been achieved by coating electro-deposited PB with a cast Nafion layer

[15], and by deposition of PB in the presence of a variety of surfactants [16,17].

A recent paper by Araminaitė et al. [26] has revealed details of the mechanism and pH

dependence of H2O2 catalytic activity of electro-deposited PB. The results were

interpreted within the framework of a 2-step reaction mechanism, involving dissociative

adsorption of H2O2 with the formation of OH radicals, followed by 1-electron reduction

of these radicals to OH–. At a higher concentration of H2O2, and especially at a higher

pH (pH 7.3), the second process appears rate limiting. The analytical implications are

that a linear dependence of cathodic current on H2O2 concentration should be observed

within the narrow peroxide concentration range associated with biosensor applications

in neutral solutions, as has been reported in practice [17,15].

Figure7.4 CV and redox reactions associated with the interconversion of surface-bound PB to the fully

oxidized form, Prussian Green (PG; Eo ≈ 0.85 V vs SCE), and a more reduced form, Prussian White (PW;

Eo ≈ 0.1 V vs SCE). In biosensor applications, PB-modified electrodes are often poised at ~0 V vs SCE, a

potential where the PW form predominates which can electrocatalytically reduce H2O2 to hydroxide ions.

+ H2O2

– 4K+

– 4OH–

(PW) K4Fe(II)4[Fe(II)(CN)6]3

↓↑+4e– + 4K

+

(PG) Fe(III)4[Fe(III)(CN)6]3Cl3

(PB) Fe(III)4[Fe(II)(CN)6]3

↑↓–3e– + 3Cl

–

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In a parallel paper [27], electrocatalytic reduction of H2O2 at electrodes modified by

electro-deposited layers of PB were studied with an in-situ Raman spectro-

electrochemical technique. During the cathodic reduction of H2O2, PW appeared to turn

partially into its oxidized form (PB; see Figure 7.4) even at electrode potentials

corresponding to the reduced form of a modifier. The ratio of PB/PW within the

modifier layer was shown to depend on H2O2 concentration, indicating that

electrocatalysis proceeds within the modifier layer rather than at an outer modifier–

electrolyte interface. In contrast, electrooxidation of AA did not affect the in-situ Raman

spectra, indicating an outer interface as the most probable site for AA oxidation.

More recently, PB and its analogues have been combined with a variety of novel

materials for electrocatalytic detection of H2O2. For example, a PB composite with

graphene oxide and chitosan (Chi) gave a detection limit of 100 nM H2O2 [28] and

cobalt hexacyanoferate nanoparticles (CoHCF/NPs) modified with carbon nanotubes

(CNT) showed a synergic effect toward H2O2 detection [29]. In this way Han et al.,

2013, reported a composite of CoHCF and platinum nanoparticles on carbon nanotubes

provided a sensor with a linear response up to 1.25 mM H2O2, also with a detection

limit of 100 nM, and a fast response time (< 2 s) [30]. Controlled synthesis of mixed

nickel-iron hexacyanoferrate nanoparticles (~35 nm average size) has been shown to be

an excellent material for selective electroanalytical applications for H2O2 and

glutathione sensing [31]. A number of these novel, mostly nanoparticle-based materials,

have been exploited to develop sensitive and selective devices for electroanalysis,

including glucose biosensors [32,33] and immunosensors [34].

7.4. Prussian Blue: Biosensor applications

During the last two decades some authors have suggested the use of electrocatalytic

films to detect H2O2 in biosensing applications [8, 10]. Based on this approach,

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Karyakin and Chaplin [8] proposed to modify the transduction element with a thin film

of PB, allowing the detection of the H2O2 generated enzymatically by GOx at a

potential close to 0 V vs SCE. That year, Jaffari and Turner [35] presented a patent in

the UK (later extended to an international patent) for an amperometric biosensor for the

determination of blood glucose using a PB-modified graphite electrode [36]. After that,

PB-modified biosensors for other molecules such as lactate, sucrose, galactose,

cholesterol, choline, oxalate, lysine, acetylcholine, ethanol, glutamate and NADH have

been reported in the literature [11, 12].

Although the first PB-modified transducers were carbon paste, glassy carbon and

platinum electrodes, recently screen-printed electrodes (SPEs) have been used because

they are inexpensive, simple and quick to prepare, versatile, and are the most

economical method for large-scale production and for the assembling of spot-test kits

for clinical and environmental analysis. First reports were based on the chemical

synthesis of PB and subsequent bulk modification of the carbon ink by PB

microparticles [37] or the in-situ modification of glassy carbon or graphite powder with

PB [38]. Another method, proposed by Ricci et al. [39], involved the direct chemical

synthesis of PB onto SPEs, by placing a drop of precursor solution onto the working

electrode area. An important advance in the context of the electro-deposition of PB (and

other hexacyanoferrates) was the addition a cationic surfactant such as CTAB (acetyl-

trimethyl-ammonium-bromide), BZTC (benzethonium chloride) or CPC

(cetylpyridinium chloride). With this approach, a significantly enhanced film growth,

efficient charge transfer kinetics, and high stability and sensitivity toward H2O2

detection have been reported for SPEs [16, 17, 40, 41].

Nowadays, Dropsens SL (Oviedo, Spain) commercializes screen-printed carbon

electrodes (SPCEs) modified with PB (Figure 7.5), being recommended for the

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development of enzymatic biosensors based on oxidases, for working with

microvolumes and for decentralized assays.

Figure 7.5 The structure of commercial Prussian Blue-modified screen printed electrode commercialized

by Dropsens SA, Spain, including a scanning electron micrograph on the right.

7.4.1. Prussian Blue and analogues enzyme system

Due to the electrocatalytic properties of PB toward the reduction of H2O2 at mild

applied potentials (see Figure 7.3) most of the enzymes employed with this approach

have been oxido-reductases [11]. However, the recent discovery of more generic

electrocatalytic properties of PB for other compounds [12] has led to new classes of

enzymes being incorporated into biosensors, such as hydrolases (e.g.,

acetylcholinesterase and butyrylcholinesterase) (see Section 7.4.1.8).

7.4.1.1. Glucose oxidase

Glucose oxidase enzyme (GOx) (EC 1.1.3.4) is an oxido-reductase that catalyses the

oxidation of glucose to H2O2 and D-glucono-δ-lactone. In cells, it aids in breaking the

sugar down into its metabolites. It is highly selective for β-D-glucose and does not act

on α-D-glucose. GOx, which is often extracted from Aspergillus niger, is widely used

for the determination of free glucose in body fluids, and in the chemical,

pharmaceutical, food, beverage, biotechnology and other industries.

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Glucose biosensors based on PB have been successfully applied to blood, serum, saliva

and urine samples. First data were presented by Deng et al. [42], where the serum

samples obtained from healthy and diabetic persons were diluted 1/50 in phosphate

buffered saline and showed a good agreement with the reference method. In 2003,

Wang et al. [43] presented a glucose biosensor based on Chi/PB film and compared

their results obtained in whole-blood samples with those obtained by a

spectrophotometric method; results from 100 samples were in excellent agreement, with

a correlation coefficient of >0.99.

In recent years, Salazar et al. [44, 45] have been working on PB-modified carbon fiber

microelectrodes (CFEs) to detect enzyme-generated H2O2 at low applied potentials as

an alternative to first- and second-generation biosensors used for physiological

applications. Thanks to this approach, the glucose biosensor reached very low

dimensions (~10 μm diameter) and displayed excellent in-vitro and in-vivo responses

based on criteria relevant to applications in neuroscience. Using last approach, Roche et

al. [46] studied different aspects of the relationship between oxygen and glucose

supplies during neurovascular coupling by detecting the temporal and spatial

characteristic of hemoglobin states and extracellular glucose concentration, combining

the use of glucose PB-modified microbiosensors with 2-dimension optical imaging

techniques.

7.4.1.2. Lactate oxidase

Lactate oxidase (LOx, EC 1.13.12.4) is classified as a flavoenzyme, which is an enzyme

containing a flavin nucleotide (FMN or FAD). LOx is a member of the FMN-containing

enzymes which catalyze the oxidation of α-hydroxyacids without the formation of any

intermediates. Historically, lactate was considered a dead-end metabolite of glycolysis

or a sign of hypoxia and anaerobic energy metabolism. However, a body of evidence

14

has been accumulated to indicate that large amounts of lactate can be produced in many

tissues under fully aerobic conditions, including neural activations [47].

In 2001, Garjonyte et al. [48] presented a PB-modified GCE where LOx was

immobilized in Nafion. Biosensors operated at –50 mV (vs Ag/AgCl, 0.1 M KCl) in a

flow injection analysis (FIA) system showed a linear range up to 0.8 mM with a

detection limit of <1 µM. However, biosensor stability and reproducibility were limited.

Using a similar approach, Lowinsohn and Bertotti [49] developed a lactate biosensor

and measured the lactate concentration in blood, demonstrating that PB-modified

biosensors were suitable for monitoring changes in the lactate levels during physical

exercise.

Recently, Salazar et al. [50] presented a lactate microbiosensor with low dimensions

(∼10 µm diameter, 250 µm length) to allow its use in neuroscience applications. CFEs

were modified with PB for the detection of enzyme-generated H2O2 at a low applied

potential. In this way, theses authors electrodeposited PB in presence of the surfactant

BZTC in order to improve its stability and sensitivity toward H2O2 (see above). This

lactate microbiosensor design displayed a sensitivity of 42 nA mM-1

cm-2

with a

detection limit of ~6 µM and a linear range up to 0.6 mM. Furthermore, the linear range

was extended up to 1.2 mM with an additional Nafion film. Finally, the microbiosensor

response was checked under physiological and electrical stimulation conditions in rat

brain and exhibited good results for in-vivo applications.

7.4.1.3. Cholesterol oxidase

Cholesterol oxidase (ChOx, EC 1.1.3.6) is a monomeric flavoenzyme that catalyzes the

oxidation and isomerization of cholesterol to cholest-4-en-3-one using O2 as electron

acceptor. Two forms of the enzyme are known, one containing the cofactor non-

covalently bound to the protein and one in which the cofactor is covalently linked to a

15

histidine residue. It is the most commonly studied enzyme for the construction of

biosensors for cholesterol assessment in biological samples. Preliminary determination

of cholesterol is clinically important because abnormal concentration is related to

disorders such as hypercholesterolemia, high blood pressure, type-2 Diabetes,

peripheral vascular diseases, stroke and coronary diseases.

In 2003 Li et al. [51] reported a cholesterol biosensor prepared by immobilizing ChOx

in a silica sol-gel matrix on the top of a PB-modified electrode. The ChOx in the sol-gel

layer maintains its activity for a long time (35 days half-life). Biosensors gave a

detection limit of ~0.2 μM and were free of the most common interference effects.

Finally, the authors determined dissociated cholesterol in serum with excellent results.

Later, Tan et al. [52] developed an amperometric cholesterol biosensor based on CNTs

and a sol-gel Chi/SiO2 organic/inorganic hybrid material. The biosensor exhibited high

sensitivity, good reproducibility and selectivity, and long-term stability. The authors

compared the free cholesterol concentration in human serum obtained with their

cholesterol biosensor against a spectrophotometric method, and found a good

correlation between the two approaches.

In 2013 Liu et al. [53] reported a novel cholesterol biosensor based on a hydrophobic

ionic liquid (IL)/aqueous solution interface on a PB-modified GCE. According the

authors the hydrophobic IL thin film played a signal amplification role because it not

only partitioned the cholesterol from the aqueous solution, but also served as an

immobilization matrix for the ChOx. The fabricated IL-ChOx/PB/GCE exhibited a

linear response to cholesterol in the range of 0.01–0.40 mM with a detection limit of ~4

μM.

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7.4.1.4. Alcohol oxidase

Alcohol oxidase (AOx, EC 1.13.13) is an oligomeric enzyme consisting of eight

identical sub-units arranged in a quasi-cubic arrangement, each containing a strongly

bound cofactor, FAD molecule. It is produced by methylotrophic yeasts (e.g.

Hansenula, Pichia, Candida) in subcellular microbodies known as peroxisomes. AOx is

the first enzyme involved in the methanol oxidation pathway of methylotrophic yeasts

and although its physiological role is the oxidation of methanol, it is also able to oxidise

other short-chain alcohols, such as ethanol, propanol and butanol. AOx is thus

responsible for the oxidation of low molecular weight alcohols to the corresponding

aldehyde, using O2 as the electron acceptor. Due to the strong oxidizing character of O2,

the oxidation of alcohols by AOx is irreversible.

The detection and quantification of alcohols with high sensitivity, selectivity and

accuracy is required in many different areas. Accurate and rapid measurement of

ethanol is very important in clinical and forensic analysis in order to analyze human

body fluids, e.g. blood, serum, saliva, urine, breath and sweat, among others.

The first alcohol biosensors based on PB were reported by Karyakin et al. in 1996 [54],

where they immobilized AOx within a Nafion layer onto PB-modified GCEs. Different

alcohols were checked such as methanol, ethanol, n-propanol, i-propanol and i-butanol.

The authors found that alcohol sensitivity decreased with increasing carbon chain length

and was higher for primary than secondary alcohols. The detection limits for methanol

and ethanol were 1 and 50 µM, respectively.

Recently, Costa et al. in 2012 [55] presented a comparative study of different alcohol

biosensors based on SPCEs modified with three different mediators (Prussian Blue,

ferrocyanide and Co-phthalocyanine). In addition, AOx from three different yeasts

(Hansenula sp., Pichia pastoris and Candida boidinii) were employed also. The authors

17

found the highest sensitivity value with PB-modified SPCEs and AOx from Hansenula,

although the background currents were very high which seriously affected the

reproducibility.

7.4.1.5. NADH oxidase

NADH oxidase (EC 1.6.99.3) is a dimeric flavoprotein and carries out oxidoreduction

reactions. NADHOx is very active at room temperature, catalyzing the proton transfer

from NADH to an electron acceptor, such as FAD, ferricyanide, oxygen and others.

Moreover, the enzyme is able to catalyze the electron transfer from NADH to various

other electron acceptors such as PB, Methylene blue, cytochrome c, p-nitroblue

tetrazolium, 2,6-dichloroindophenol, even in the absence of flavin shuttles. NADH

plays a central role in mitochondrial respiratory metabolism, stimulating the energy

production in all living cells (notably, brain, heart and muscles). NADH detection is of a

great importance because it is produced in reactions catalyzed by more than 250

dehydrogenases.

In 2007 Raoi et al. [56] developed a biosensor for the determination of reduced NADH

using a recombinant enzyme NADHOx from Thermus thermophilus covalently

immobilized on PB bulk-modified SPEs. FIA was selected to optimize the biosensor

configuration and other analytical parameters such as cofactor (FMN) concentration,

flow rate, buffer types, pH dependence, response time and operational stability. The

biosensor showed the highest response at pH 5.0, for which the detection and

quantification limits were 0.1 and 0.4 μM, respectively, with a linear working range

between 1 and 400 µM. Finally, the proposed biosensor was stable for 2 months.

Another interesting approach was presented by Gurban et al. in 2008 [57], where

NADH oxidase was immobilized on PB-modified SPEs. The amperometric detection of

18

NADH was performed at +0.25 V vs Ag/AgCl. Two different approaches were

employed: either adding FMN to the reaction medium or immobilized on the biosensor.

The optimal configuration was obtained when FMN was entrapped with NADHOx in

the biocatalytic layer using a sol–gel matrix, and displayed sensitivity, linear range and

detection limits of 4.6 mA M−1

cm−2

, 1.6 mM and 1.2 µM, respectively. Finally,

biosensors showed good long-term and operational stability.

7.4.1.6. Diamine oxidase

Diamine oxidase (DaOx, DAO, EC 1.4.3.6) catalyzes the degradation of histamine and

other biogenic amines. The enzyme belongs to the class of copper-containing amine

oxidases which catalyze the oxidative deamination of primary amines by O2 to form

aldehydes, NH3 and H2O2. These copper amine oxidases are characterized by possessing

the active-site cofactor topa quinone, formed post-translationally by modification of a

conserved tyrosine residue. Biogenic amines (histamine, putrescine, cadaverine,

spermine) are volatile amines that are produced as a result of the breakdown of amino

acids. Histamine has been identified as the causative agent of the disease

Scombrotoxicosis or scombroid poisoning, which can, in severe cases, cause symptoms

such as headache, nausea, vomiting, diarrhoea, itching, oral burning sensation, red rash

and hypotension. Biogenic amines may also be considered as carcinogens because of

their ability to react with nitrites to form potentially carcinogenic nitrosamines

In 2010 Piermarini et al. [58] presented a simple and rapid method for the analysis of

biogenic amines in human saliva by using DaOx immobilized on a PB-modified SPE.

The biosensor response was investigated for different amines such as putrescine,

cadaverine, spermine, histamine, etc. The results obtained during the evaluation of

saliva showed that the developed electrochemical biosensor can be considered a valid

19

point-of-care testing method for the determination of salivary polyamines, as well as

being suitable for biomedical studies.

7.4.1.7. Choline oxidase

Choline oxidase (ChlOx, EC 1.1.3.17) is an enzyme that catalyzes the oxidation of

choline to generate glycine betaine via betaine aldehyde with H2O2 generation. The

enzyme acts on the CH-OH groups of donor with O2 as electron acceptor and FAD as

cofactor. Choline and its metabolites are needed for three main physiological purposes

such as structural integrity for cell membranes, cholinergic neurotransmission and a

major source for methyl groups via its metabolite, betaine. On the other hand, choline,

as a marker of cholinergic activity in brain tissue, is very important in biological and

clinical analysis, especially in the clinical detection of neurodegenerative disorders.

In 2006, Shi et al. [59] reported an amperometric choline biosensor based on the

immobilization of ChlOx in a layer-by-layer (LBL) multilayer film on a PB-modified Pt

electrode. The authors suggested that the high sensitivity and fast response time

observed may be due to the efficacy of the enzyme immobilization and to the ultrathin

nature of the LBL film, in which mass-transport problems were minimized. The

analytical values of choline in serum samples obtained by this choline biosensor agreed

satisfactorily with those by a spectrophotometric method. Finally, the choline biosensor

retained ~86% of its initial current response to choline after ca. 2 months.

In 2012, Zhang et al. [60] developed an electrochemical approach for the detection of

choline based on PB-modified iron phosphate nanostructures (PB–FePO4). These

nanostructures showed a good catalysis toward the electro-reduction of H2O2, and

allowed the construction of an amperometric choline biosensor immobilizing ChlOx on

the PB–FePO4 nanostructures. The biosensor exhibited a rapid response (ca. 2 s), low

detection limit (0.4 μM), wide linear range (2 μM to 3.2 mM), high sensitivity (∼75 μA

20

mM−1

cm−2

), as well as good stability and repeatability. Also, the common interfering

species, such as ascorbic acid, uric acid and 4-acetamidophenol did not cause

observable interference.

During recent decades, organophosphorus (OP) compounds have been received much

attention due to their harmful effects on human health. Therefore, the development of

fast and sensitive detection methods has become more urgent. In 2009, Sajjadi et al.

[61], developed a PB-modified SPEs coupled with ChlOx for detection of paraoxon as

inhibitor. The concentration of H2O2 produced by ChlOx was electrochemically

determined by the PB-modified electrode poised at −50 mV versus the internal screen-

printed Ag pseudo-reference electrode. The decrease in current caused by the addition

of inhibitor was used for evaluation of paraoxon concentration. For an incubation time

of 5 min, the biosensor response was linear from 0.1 to 1 μM of paraoxon with a

detection limit of 0.1 μM.

7.4.1.8. Acetilcholinesterase and butyrylcholinesterase

Acetylcholinesterase (AChE, EC 3.1.1.7) is a serine protease that hydrolyzes the

neurotransmitter acetylcholine, and belongs to the carboxylesterase family. The active

site of AChE comprises two subsites – the anionic site and the esteratic subsite. The

structure and mechanism of action of AChE have been elucidated from the crystal

structure of the enzyme. AChE is found at mainly neuromuscular junctions and

cholinergic brain synapses, where its activity serves to terminate synaptic transmission.

Butyrylcholinesterase (BuChE, EC 3.1.1.8) is a non-specific cholinesterase enzyme that

hydrolyses many different choline esters. In humans, it is found primarily in the liver

and is encoded by the BCHE gene. It is very similar to the neuronal

acetylcholinesterase. Recently, biosensor techniques based on the inhibition of AChE

and BChE activity by OPs and toxins have gained considerable attention due to the

21

advantages of simplicity, rapidity, reliability and low cost devices. In addition, the

challenges of biohazards and bioterrorism, especially the need for early detection of

nerve agents have contributed to the use of such approaches.

In 2010, Sun and Wang [62] developed a novel acetylcholinesterase AChE biosensor

based on dual-layer membranes (a Chi membrane and a PB membrane) modifying a

GCE. Before biosensor operation, the Chi enzyme membrane was quickly fixed on the

surface of PB/GCE with an O-ring to prepare an amperometric GCE/PB-AChE sensor

for OP pesticides. The proposed biosensor exhibited extreme sensitivity to OP

pesticides compared to the other kinds of AChE biosensor and provided good results for

dichlorvos, omethoate, trichlorfon and phoxim.

In 2012, Zhang et al. [63] demonstrated a facile procedure to efficiently prepare PB

nanocubes/reduced graphene oxide (PBNCs/rGO) nanocomposite by directly mixing

Fe3+

and Fe(CN)63-

in the presence of GO in a polyethyleneimine aqueous solution.

Later, thanks to the high electrocatalytic activity of PBNCs/rGO towards the oxidation

of thiocholine, this nanocomposite was employed to develop a novel

acetylcholinesterase (AChE) biosensor for detection of OP pesticides.

Recently, Arduini et al. [64] immobilized BChE onto PB-modified SPEs, and nerve

agent detection was performed by measuring the residual activity of the enzyme. The

optimized biosensor was tested with two common nerve agent standard solutions (sarin

and VX), and showed detection limits of 12 and 14 ppb (10% of inhibition),

respectively. The enzymatic inhibition was also checked by exposing the biosensors to

sarin in gas phase at a concentration of 0.1 mg m−3

. Finally, Arduini et al. [65], have

reported a portable prototype for nerve agent detection based on an electrochemical

AChE biosensor; tests with paraoxon gave satisfactory results.

7.5. Prussian Blue: Immunosensor applications

22

Radioimmunoassay (RIA), enzyme-linked immunoassays (ELISA) and

fluoroimmunoassay (FISA) have been successfully used for great numbers of

applications, such as detection of polypeptide hormones (insulin, glucagon), detection

of steroid and amino acid/fatty acid-derived hormones (aldosterone cortisol, melatonin),

detection of therapeutic agents (amikacin, chloramphenicol, gentamicin), drugs of abuse

(amphetamines, barbiturates, canabinoids, cocaine, opiates) and disease markers

(thyroid disease, cancer, hypercalcemia and bone disease, hirsutism, virilism,

infertility). However, these methods involve complicated, time-consuming assay

processes, need specially equipped personnel and sophisticated instrumentation.

Moreover, implementation of Point of Care (POC) testing, which is very important

especially for developing countries where access to medical and analytical resources is

limited, is difficult with conventional immunoassays because rather large and expensive

equipment is necessary. Therefore, new techniques for simple, rapid and reliable

detection of disease markers are strongly desirable. Recently, electrochemical

immunosensors, which combine the high efficiency of enzyme catalysis, specificity of

antibody-antigen binding and high sensitivity of electrochemical response, have gained

much attention and been applied for clinical immunoassays, not least because this

technique combines simple, portable, low-cost electrochemical measurement systems.

A primary strategy for amplifying the electrochemical signal in immunosensors is

associated with a labeled antigen or antibody. Enzyme labeling method is most

commonly employed to amplify the number of signal-reporting molecules per

biospecific binding between a target biomolecule and an enzyme-labeled biomolecule.

An enzyme such as horseradish peroxidase (HRP) or alkaline phosphatase (AP) is

usually conjugated to an antibody or an antigen to generate electrochemically detectable

23

species and offered biocatalytic signal amplification in a competitive or noncompetitive

assay; others such as GOx may be used too (see Figure 7.6)

Figure 7.6 Detection scheme for a Prussian Blue (PB)-modified electrode in immunosensing. Glucose

oxidase-labeled antibody converts glucose to gluconolactone on the electrode surface. Secondary to this

reaction is the production of hydrogen peroxide that can be reduced amperometrically at low applied

overpotentials, electrocatalyzed by the PB.

In addition, the label-free configuration has emerged over recent years as an easy and

promising method. Below we present some immunosensor devices under these two

approaches (enzyme-labeled and label-free) using some common antibodies selective

against a variety of antigens, including: α-fetoprotein antigen (anti-AFP),

carcinoembryonic antigen (anti-CEA), carbohydrate antigen 19-9 (anti-CA19-9),

carcinoma antigen 25 (anti-CA125), and prostate specific antigen (anti-PSA).

7.5.1. α-fetoprotein antigen

α-fetoprotein antigen (AFP) is a glycoprotein with a molecular weight of approximately

70 kDa. It is normally excreted during fetal and neonatal development by the liver, yolk

sac, and in small concentrations by the gastrointestinal tract. It is one of the most

extensively used clinical tumor markers. The concentration of AFP in healthy human

serum is typically below to 25 ng mL-1

. Nevertheless, elevated AFP concentration in

serum may be an early indication of some cancerous diseases such as hepatocellular

24

cancer, liver metastasis from gastric cancer, testicular cancer and nasopharyngeal

cancer, etc.

In 2004, Guan et al. [66] reported one-step immunoassay using PB-modified SPEs.

Here anti-human-AFP monoclonal IgG was inmobilized onto the electrode and used as

the recognition element, and polyclonal anti-human-AFP IgG labeled with GOx was

employed to produce the electrochemical signal. The PB-modified SPE catalyzed H2O2

from the reaction of GOx which allowed quantification of AFP in the sample. Using

FIA, the detection range was in the range from 5 to 500 ng mL-1

. Finally, the authors

compared their results in real serum samples against ELISA. Both methods were found

to give similar results; however, the new approach was less time consuming (30 min)

compared to ELISA (2 h).

In 2007, Yuan et al. [67] presented a label-free amperometric immunosensor for the

determination of AFP by immobilizing TiO2 colloids on a PB-modified electrode. AFP

responses showed two concentration ranges from 3 to 30 ng mL-1

and from 30 to 300 ng

mL-1

with a detection limit of 1 ng mL-1

and exhibited high selectivity, good

reproducibility, long-term stability (>2 months) and good repeatability. Finally, the

author compared these results, obtained for real serum samples, against

chemiluminescence immunoassays (CLIA) and found good agreement between the two

methods.

Lately, Hong et al. [68] developed PBNPs and coated them with bovine serum albumin

(BSA) to improve their stability. Then gold colloids were loaded on the BSA-coated

PBNPs to construct a core-shell-shell nanostructure. Finally, AFP antibody was

attached to GNPs and PBNPs/BSA/GNPs/anti-AFP was used to AFP detection. The

dynamic range of the resulting immunosensor for the detection of AFP was from 0.02 to

25

200 ng mL-1

with a detection limit of 0.006 ng mL-1

and displayed good selectivity,

stability and reproducibility.

In 2010, Jiang et al. [69] reported an amperometric immunosensor based on the

sequential electrodeposition of PB and GNPs on a CNT/GCE surface. Finally, anti-AFP

was immobilized onto the GNP surface and BSA employed to block possible remaining

active sites of GNP monolayer and avoid any nonspecific adsorption. Under optimal

experimental conditions, the immunosensor showed an ultralow limit of detection of 3

pg mL-1

and a linear range from 0.01 to 300 ng mL-1

. Moreover, the immunosensor, as

well as a commercially available kit, were both used in the determination of AFP in real

human serum and showed excellent correlation.

Finally, in 2011, Dai et al. [70] developed a sandwich electrochemical immunosensor

for the sensitive determination of AFP based on PB-modified hydroxyapatite

(PB@HAP) modified with HRP and secondary anti-AFP antibody (Ab2) to fabricate

the electrochemical immunosensor label (PB@HAP/HRP/Ab2). The results indicated

that the immunosensor fabricated using PB@HAP/HRP/Ab2 label had high sensitivity,

much higher than other labels such as PB@HAP/Ab2, PB/HRP/Ab2 or HAP/HRP/Ab2.

Optimized amperometric signals increased linearly with AFP concentration in the range

of 0.02 to 8 ng mL-1

with a low detection limit of 9 pg mL-1

.

7.5.2. Carcinoembryonic antigen

Carcinoembryonic antigen (CEA) is a glycoprotein involved in cell adhesion. It is

normally produced during fetal development, but the production of CEA stops before

birth. CEA is one of the most widely used tumor markers worldwide. Its main

application is mostly in gastrointestinal cancers, especially in colorectal malignancy,

colorectal carcinoma, pancreatic carcinoma, lung carcinoma and breast carcinoma.

26

In 2009, Ling et al. [71] developed an immunosensor for CEA based on the electrostatic

adsorption between the positively charged MnO2 nanoparticles and Chi composite

membrane (nano-MnO2/Chi) and the negatively charged PB. Thus, nano-MnO2/Chi

membrane was adsorbed onto PB-modified electrode surface and GNPs were

electrodeposited to immobilize anti-CEA. Under the optimized conditions,

determination of CEA displayed a linear response in two ranges, from 0.25 to 8.0 ng

mL-1

and from 8.0 to 100 ng mL-1

, with a detection limit of 0.08 ng mL-1

. Real serum

samples were measured with the immunosensor approach and compared against ELISA,

with no significant difference observed between the two methods.

Then Zhuo et al. [72] presented a bienzyme functionalized three-layer composite

magnetic nanoparticles for electrochemical determination of AFP and CEA. Authors

developed a three-layer magnetic nanoparticle composed of a Fe3O4 magnetic core, a

PB interlayer and a gold shell (Fe3O4/PB/Au). In addition, they used a new signal

amplification strategy based on a bienzyme system using HRP and GOx functionalized

nanoparticles for electrochemical immunosensing using anti-CEA and anti-AFP,

respectively.

Finally, in 2011, Tang et al. [73] showed the utility of a sensitive electrochemical

immunoassay for CEA detection with signal dual-amplification using GOx and PB. The

first signal amplification was based on the labeled GOx on the anti-CEA-gold-silver

hollow microspheres (anti-CEA/GOx/Au@Ag@HS) toward the catalytic oxidation of

glucose. Thus the enzymatic-generated H2O2 was catalytically reduced by the

immobilized PB on the graphene nanosheets. With a sandwich-type immunoassay

format, optimized electrochemical immunosensor exhibited a wide dynamic range of

0.005 to 50 ng mL−1

with a low detection limit of 1.0 pg mL−1

. Finally, immunosensors

27

were evaluated for clinical serum specimens, and showed good correlation with those

obtained by the electrochemiluminescent method (ECL).

7.5.3. Carbohydrate antigen 19-9

Carbohydrate antigen 19-9 (CA19-9) is one of the most important carbohydrate tumor

markers associated with biliary, liver, lung and breast cancers, and various other gastro-

intestinal malignancies. Furthermore, CA19-9 is also used as a diagnostic marker for

hepatic cyst infection in the autosomal dominant polycystic kidney, and the serum

concentrations of CA19-9 increase remarkably in patients with severely damaged renal

function

In 2010, Liu et al. [74] proposed a novel signal amplification strategy based on PB and

Pt hollow nanospheres (Pt/HN) for developing a highly sensitive label-free CA19-9

immunosensor. The authors combined the excellent electrocatalytic activity of PB and

Pt/HN toward H2O2 reduction to amplify the electrochemical signal. The resulting

immunosensors showed a high sensitivity and broad linear response to CA19-9 in two

ranges from 0.5 to 30 and 30 to 240 U mL-1

with a low detection limit of 0.13 U mL-1

.

Real samples analyzed with this approach and ELISA methodology revealed good

agreement; the authors suggested that the developed immunoassay may be applied for

clinical determination of the CA19-9 level in human serum specimens.

7.5.4. Neuron-specific enolase antigen

Neuron-specific enolase (NSE) is a glycolytic isoenzyme which is located in central and

peripheral neurons and neuroendocrine cells. It has been detected in patients with

certain tumors, namely: neuroblastoma, small cell lung cancer, medullary thyroid

cancer, carcinoid tumors, pancreatic endocrine tumors, and melanoma.

In 2010, Zhong et al. [75] developed a label-free electrochemical immunosensor based

on PB doped silicon dioxide (SiO2/PB) nanocomposite. SiO2/PB nanocomposite

28

(produced by using a microemulsion method) was used to obtain a nanostructural

monolayer on a GCE surface. Next 3-aminopropyltriethoxy silane (APTES) was self-

assembled in order to obtain an amino-functionalized interface and Chi-stabled gold

nanoparticles (Chi/GNP) were subsequently attached. Finally, anti-NSE was loaded via

the adsorption of gold nanoparticles. The immunosensor exhibited good linear behavior

in the concentration range from 0.25-5.0 and 5.0-75 ng mL-1

with a limit of detection of

0.1 ng mL-1

. The authors demonstrated the application of this proposed immunosensor

for the determination of NSE in human serum samples and obtained good results when

compared against ELISA

7.5.5. Carcinoma antigen 125

Cancer antigen 125 or carbohydrate antigen 125 (CA125) is a member of the mucin

family glycoproteins expressed in coelomic epithelium during embryonic development.

CA-125 has found application as biomarker for ovarian cancer detection, although it

may also be elevated in other cancers, including endometrial, fallopian tube, lung, breast

and gastrointestinal cancers.

In 2008, Chen et al. [76] immobilized anti-CA125 onto GNPs-modified PBNPs and

used this to develop a highly sensitive amperometric immunosensor for the detection of

CA125. Firstly PBNPs, synthesized using Chi and poly(diallyldimethylammonium

chloride) (PDDA) as a protective matrix, were cast onto a GCE surface. Then, GNPs

were assembled by the interactions between GNPs and amino groups of Chi and

electrostatic interactions between oppositely charged GNPs and PDDA. Finally, anti-

CA125 was assembled onto the surface of GNPs. The proposed immunosensor showed

a high sensitivity, broad linear range and low detection limit for CA125 determination.

DPV peak current was proportional to the CA125 concentration in two ranges from 2.0

to 40 and 40 to 100 U mL-1

and the detection limit close to 0.7 U mL-1

. Simultaneous

29

analysis of human serum samples with the present approach and with the ELISA

protocol suggested acceptable agreement between these two methods.

7.5.6. Human chorionic gonadotropin antigen

Human chorionic gonadotropin (HCG) is a hormone produced during pregnancy that is

made by the developing placenta after conception, and later by the placental component

syncytiotrophoblast. Some cancerous tumors produce this hormone; therefore, elevated

levels measured when the patient is not pregnant can signal some type of cancer, such as

seminoma, choriocarcinoma, germ cell tumors, hydatidiform mole formation, teratoma

with elements of choriocarcinoma, and islet cell tumor.

In 2011, Yang at al. [77] developed an electrochemical immunosensor for HCG based

on HRP-functionalized PB-carbon nanotubes-gold nanocomposites (HRP/GNPs/PB/

CNTs) as labels for signal amplification. Using this approach, Chi hydrogel and TiO2

nanocomposites were first coated onto a GCE surface for the immobilization of primary

antibodies (Ab1). Then, the detectable signal was recorded and amplified based on a

sandwich-type immunoassay by the employment of HRP-labeled secondary antibodies

(HRP-Ab2) bioconjugate (HRP-Ab2/GNPs/PB/CNTs). Under optimized conditions, the

immunosensor showed a linear range from 0.05 to 150 mIU mL-1

and a low detection

limit of 0.02 mIU mL-1

HCG. The feasibility of the immunoassay for clinical

applications was investigated by analyzing several real samples and comparing with the

ELISA method. The author found no significant difference between the two methods

and anticipated that this immunosensor could be reasonably applied in the clinical

determination of HCG.

7.5.7. Prostate specific antigen

Prostate specific antigen (PSA) is an androgen-regulated serine protease. PSA is

secreted by the epithelial cells of the prostate gland. PSA is produced for the ejaculate,

30

where it liquefies semen in the seminal coagulum and allows sperm to swim freely. It is

also believed to be instrumental in dissolving cervical mucus, allowing the entry of

sperm into the uterus. When PSA enters in the circulatory system, it is rapidly bound by

protease inhibitors, primarily α1-antichymotrypsin (ACT), although a fraction is

inactivated in the lumen by proteolysis and circulates as free PSA (f-PSA). Total PSA

(T-PSA) refers to the sum of f-PSA and PSA/ACT complex in serum. T-PSA levels

significantly increase in serum during prostate cancer and other prostatic diseases.

Han et al. [78] described in 2012 a sandwich-type electrochemical immunosensor for

simultaneous sensitive detection of PSA and fPSA. First, GNPs modified PB and nickel

hexacyanoferrates nanoparticles were prepared and used to decorate onion-like

mesoporous graphene sheets (O-GS/PBNPs/GNPs and O-GS/NiNPs/GNPs). Then, O-

GS/PBNPs/GNPs and O-GS/NiNPs/GNPs were modified with anti-fPSA and anti-PSA

respectively, and streptavidin and biotinylated alkaline phosphatase (bio-AP) were

employed to block active sites. Then dual catalysis amplification was achieved by

catalysis of the ascorbic acid 2-phosphate to AA in the presence of bio-AP, and then the

enzyme-generated AA was further catalytically oxidized by O-GS/PBNPs/GNPs and O-

GS/NiNPs/GNPs nanohybrids at ~0.2 and ~0.4 V vs SCE, respectively. The experiment

results showed that the linear range of the proposed immunosensor for simultaneous

determination of fPSA was from 0.02 to 10 ng mL−1

with a detection limit of 7 pg mL−1

and PSA was from 0.01 to 50 ng mL−1

with a detection limit of 3.4 pg mL−1

. To

evaluate the performance of the proposed immunosensor, clinical analysis with serum

samples were evaluated and compared against ELISA. The authors found good

correlations between those results that confirm the viability of the developed method for

real sample assay.

31

7.5.8. Hepatitis B antigen

Hepatitis B (HB) is an infectious inflammatory illness of the liver caused by the

hepatitis B virus (HBV). The virus is transmitted by exposure to infectious blood or

body fluids such as semen and vaginal fluids, while viral DNA has been detected in the

saliva, tears, and urine of chronic carriers. Perinatal infection is a major route of

infection; other risk factors for developing HBV infection include working in a

healthcare setting, transfusions, dialysis, acupuncture, tattooing, etc.

Recently, He et al. [79] developed a label-free amperometric immunosensor based on

the electro-deposition of GNPs over PB film. In this way, HB antibody was

immobilized onto a modified sensor (GCE/PB/GNPs) surface. Under optimal conditions

the peak current response was inversely proportional to the HB antigen concentration.

Current changes were proportional to HB antigen concentration ranging from 2 to 300

ng mL-1

with a detection limit of 0.4 ng mL-1

. Finally, the authors measured and

compared 50 clinical samples; their results were in good agreement with those using

ELISA as a reference method.

7.6. Conclusions

In the present chapter we discussed the main advantages of biosensors in biomedical

applications. Numerous benefits are driving the replacement by biosensors of other

conventional, more sophisticated, analytical methods. Nowadays, biosensors are

focused to Point of Care applications, glucose biosensors being the best example due to

the serious global problem of diabetes. However, great efforts are being made with a

wide range of other illnesses and, in this context, PB and its analogues are promising

materials due to their electrocatalytic properties. Among different approaches, oxidase-

based biosensors are still preferred, and PB is an excellent material for use in the

fabrication of biosensors because it acts as an “artificial peroxidase”. Together with the

32

facile modification of the electrode substrate and the low cost of production, this has led

to an on-going replacement of the common enzymatic detection method – horseradish

peroxidase (HRP) – which is more expensive and complicated than PB-modified

substrates. In addition, over recent years, new applications have appeared in the form of

immunosensors, with an important and new creative concept. In this way, novel hybrid

configurations have appeared, using PB and analogues as well as other modern

materials such as graphene, carbon nanotubes, magnetic beads, gold nanoparticles, etc.

Based on these results, it is clear that during the next decades, PB and its analogues will

have important roles in the future development of biomedical devices for health care.

Acknowledgment

The funds for the development of this contribution have been provided by subprograma

INNCORPORA TU (INC–TU–2011–1621) from Ministerio de Economía y

Competitividad, Ministerio de Industria, Turismo y Comercio (TSI–020100–2011-189

and TSI–020100–2010–346); Ministerio de Ciencia e Investigación (TIN2011–28146)

and FEDER.

References

1. D.R. Thévenot, K. Toth, R.A. Durst and G.S. Wilson, Biosens. Bioelectron,

Vol.16, p. 121-131, 2001.

2. P. D'Orazio, Clin. Chim. Acta, Vol.334, p.41-69, 2003.

3. P. D'Orazio, Clin. Chim. Acta, Vol.412, p.1749-1761, 2011.

4. S. Wild, G. Roglic, A. Green, R. Sicree and H. King, Diabetes Care, Vol.27(5),

p.1047-1053, 2004.

5. J. Wang, Electroanal, Vol. 13, p. 983-988, 2001.

6. L.C. Clark and C. Lyons, Ann. N. Y. Acad. Sci, Vol. 102, p. 29-45, 1962.

7. K. Itaya, N. Shoj and I. Uchida, J. Am. Chem. Soc, Vol.106, p. 3423-3429, 1984.

8. A.A Karyakin and M. Chaplin, J. Electroanal. Chem, Vol. 56, p. 85-92, 1994.

9. A.A Karyakin, O.V Gitelmacher and E.E Karyakina, Anal. Chem, Vol. 67, p.

2419-2423, 1995.

10. A.A Karyakin and E.E Karyakina, Sens. Actuators B Chem., Vol. 57 p. 268-273,

1999.

11. F. Ricci and G. Palleschi, Biosen. Bioelectron, Vol. 21, p. 389–407, 2005.

12. A.A Karyakin, Electroanal, Vol.13, p. 813-819, 2001.

33

13. H.J. Buser, D. Schwarzenbach, W. Petter, and A. Ludi, Inorg. Chem., Vol. 16, p.

2704, 1977.

14. F. Herren, P. Fischer, A. Ludi, and W. Halg, Ludi, Inorg. Chem., Vol. 19, p. 956,

1980.

15. P. Salazar, M. Martín, R.D. O´Neill, R. Roche, and J.L. González-Mora,

Electrochim Acta, Vol. 55, p. 6476, 2010

16. P. Salazar, M. Martín, R.D. O´Neill, R. Roche, and J.L. González-Mora, Colloids

Surf B Biointerfaces, Vol. 92, p. 180-189, 2012.

17. P. Salazar, M. Martín, R.D. O´Neill, R. Roche, and J.L. González-Mora, J. Electroanal. Chem., Vol. 674, p. 48-56, 2012.

18. A. Peeters, P. Valvekens, R. Ameloot, G. Sankar, C. E. A. Kirschhock, and D. E.

De Vos, ACS Catal., Vol. 3, p. 597, 2013.

19. J.F. Hartwig, Organotransition Metal Chemistry: From Bonding to Catalysis,

California, University Science Books, Mill Valley, Chapter 3, p. 102-03, 2010.

20. J.E. Huheey, E.A. Keiter, and R.L. Keiter, Inorganic Chemistry: Principles of

Structure and Reactivity, New York, Harper Collins College Publishers, p. 519-

20, 1994.

21. J.P. Lowry and R.D. O'Neill, Electroanal, Vol.6, p. 369-379, 1994.

22. W. Chen, Q.Q. Ren, Q. Yang, W. Wen and Y.D. Zhao, Anal. Lett., Vol. 45, p.

156-167, 2012.

23. R.D. O'Neill, S.C. Chang, J.P. Lowry and C.J. McNeil, Biosens. Bioelectron.,

Vol. 19, p. 1521-1528, 2004.

24. M.A. Rahman, N.H. Kwon, M.S. Won, E.S. Choe and Y.S. Shim, Anal. Chem.,

Vol. 77, p. 4854-4860, 2005.

25. V.D. Neff, J. Electrochem. Soc., Vol.125, p. 886-887, 1978.

26. R. Araminaite, R. Garjonyte and A. Malinauskas, J. Solid State Electrochem.,

Vol. 14, p. 149-155, 2010.

27. R. Mazeikiene, G. Niaura and A. Malinauskas, J. Electroanal. Chem., Vol. 660,

p. 140-146, 2011.

28. H.Q. Gong, M.H. Sun, R.H. Fan and L. Qian, Microchim. Acta, Vol. 180, p.

295-301, 2013.

29. L. Qian and X. Yang, Talanta, Vol. 69(4), p. 957-962, 2006.

30. L.J. Han, Q. Wang, S. Tricard, J.X. Liu, J. Fang, J.H. Zhao and W.G. Shen, RSC

Adv., Vol. 3, p. 281-287, 2013.

31. P.C. Pandey and A.K. Pandey, Analyst, Vol. 138, p. 952-959, 2013.

32. C.Y. Wang, S.H. Chen, Y. Xiang, W.J. Li, X. Zhong, X. Che and J.J. Li, J. Mol.

Catal. B - Enz., Vol. 69, p. 1-7, 2011.

33. L.Y. Jin, X. Gao, H. Chen, L.S. Wang, Q. Wu, Z.C. Chen and X.F. Lin, J.

Electrochem. Soc., Vol. 160, p. H6-H12, 2013.

34. G. Volpe, U. Sozzo, S. Piermarini, E. Delibato, G. Palleschi and D. Moscone,

Anal. Bioanal. Chem., Vol. 405, p. 655-663, 2013.

35. S.A. Jaffari, and A.P.F. Turner. Hexacyanoferrate (III) modified carbon

electrodes, its application in enzyme electrodes, UK Patent Application GB

9402591.3, 1994.

36. A.P.F Turner and S.A. Jaffari. Hexacyanoferrate modified electrodes,

International Patent, WO95/21934, 1995.

37. M.P. O’Halloran, M. Pravda and G.G. Guilbault, Talanta, Vol. 55, p. 605, 2001.

38. F. Ricci, A. Amine, C.S. Tuta, A.A. Ciucu, F. Lucarelli, G. Palleschi and D.

Moscone, Anal. Chim. Acta, Vol. 485, p. 111, 2003.

34

39. F. Ricci, A. Amine, G. Palleschi and D. Moscone, Biosens. Bioelectron., Vol.

18, p. 165, 2003.

40. R. Vittal, K.J. Kim, H. Gomathi and V. Yegnaraman, J. Phys. Chem. B., Vol. 112

p. 1149, 2008.

41. S.M. Senthil Kumar and K. Chandrasekara Pillai, J. Electroanal. Chem., Vol.

589 p.167, 2006.

42. Q. Deng, B. Li and S. Dong, Analyst, Vol. 123(10), p. 1995-1999, 1998.

43. Y. Wang, J. Zhu, R. Zhu, Z. Zhu, Z. Lai and Z. Chen , Meas. Sci. Technol., Vol.

14, p. 831-836, 2003.

44. P. Salazar, R.D. O’Neill, M. Martín, R. Roche and J.L. González-Mora, Sens

Actuators B Chem., Vol. 152, p. 137-143, 2011.

45. P. Salazar, M. Martín, R. Roche, J.L González-Mora and R.D. O'Neill, Biosens

Bioelectron., Vol. 26(2), p. 748-753, 2010.

46. R. Roche, P. Salazar, M. Martín, F. Marcano and J.L. González-Mora, J.

Neurosci Methods., Vol. 202(2), p.192-198, 2011.

47. L. Pellerin, Mol. Neurobiol., Vol. 32, p 59-72, 2005.

48. R. Garjonyte, Y. Yigzaw, R. Meskys, A. Malinauskas and L. Gorton, Sens

Actuators B Chem., Vol.79, p.33-38, 2001.

49. D. Lowinsohn and M. Bertotti M, Anal Biochem., Vol. 365(2), p. 260-265, 2007.

50. P. Salazar, M. Martín, R.D. O’Neill, R. Roche, J.L González Mora., Int. J.

Electrochem. Sci., Vol. 7, p. 5910-5926, 2012.

51. J. Li, T. Peng and Y. Peng, Electroanal, Vol. 15(12), p. 1031-1037, 2003.

52. X. Tan, M. Li, P. Cai, L. Luo and X. Zou, Anal. Biochem., Vol. 337, p. 111-

120, 2005.

53. X. Liu, Z. Nan, Y. Qiu, L. Zheng and X. Lu, Electrochim. Acta., Vol. 90, p. 203-

209, 2013.

54. A.A. Karyakin, E.E. Karyakina and L. Gordon, Talanta, Vol. 43 (9), p. 1597-

1606, 1996.

55. E. Costa, J. Biscay, M.B. González, J. Reviejo, J.M. Pingarrón and A. Costa,

Anal. Chim. Acta., Vol. 728, p. 69-76, 2012.

56. A. Radoi, D. Compagnone, E. Devic and G. Palleschi, Sens Actuators B Chem.,

Vol. 121, p. 501-506, 2007.

57. A.M Gurban, T. Noguer, C. Bala and L. Rotariu, Sens Actuators B Chem., Vol.

128, p.536-544, 2008.

58. S. Piermarini, G. Volpe, R. Federico, D. Moscone and G. Palleschi, Anal. Lett.,

Vol. 43, p. 1310-1316, 2010.

59. H. Shi, Y. Yang, J. Huang, Z. Zhao, X. Xu, J. Anzai, T. Osa and Q. Chen,

Talanta, Vol.70, p.852-858, 2006.

60. H. Zhang, Y. Yin, P. Wu and C. Cai, Biosens Bioelectron., Vol. 31, p.244-250,

2012.

61. S.Sajjadi, H. Ghourchian and H. Tavakoli, Biosens Bioelectron., Vol. 24 (8), p.

2509-2514, 2009.

62. X. Sun and X. Wang, Biosens Bioelectron., Vol. 25, p. 2611–2614, 2010.

63. L. Zhang, A. Zhang, D. Du and Y. Lin, Nanoscale, Vol. 4 (15), p. 4674-4679,

2012.

64. F. Arduini, A. Amine, D. Moscone, F. Ricci and G. Palleschi, Anal. Bioanal.

Chem., Vol. 388, p. 1049–1057, 2007.

35

65. F. Arduini, D. Neagu, S. Dall’Oglio, D. Moscone and G. Palleschi, Electroanal,

Vol. 24 (3), p. 581-590, 2012.

66. J. Guan, Y. Miao and J. Chen, Biosens Bioelectron., Vol. 19, p. 789–794, 2004.

67. Y. Yuan, R. Yuan, Y. Chai, Y. Zhuo, Y. Shi, X. He and X. Miao, Electroanal,

Vol. 19 (13), p. 1402-1410, 2007.

68. C. Hong, R. Yuan, Y. Chai and Y. Zhuo, Electroanal, Vol. 20 (20), p. 2185-

2191, 2008.

69. W. Jiang, R. Yuan, Y. Chai and B. Yin, Anal. Biochem., Vol. 407, p. 65-71,

2010.

70. Y. Dai, Y. Cai, Y. Zhao, D. Wu, B. Liu, R. Li, M. Yang, Q. Wei, B. Du and H.

Li, Biosens Bioelectron., Vol. 28, p.112-116, 2011.

71. S. Ling, R. Yuan, Y. Chai and T. Zhang, Bioprocess. Biosyst. Eng., Vol. 32, p.

407-414, 2009.

72. Y. Zhuo, P. Yuan, R. Yuan, Y. Chai and C. Hong, Biomaterials, Vol. 30, p.

2284-2290, 2009.

73. J. Tang, D. Tang, Q. Li, B. Su, B. Qiu and G. Chen, Anal. Chim. Acta, Vol.

Volume 697 (1-2), p. 16-22, 2011.

74. H. Liu, R. Yu, K. Peng, H. Zhao, L. Li and X.Wu, Electroanal., Vol.22 (21), p.

2577-2586, 2010.

75. Z. Zhong, J. Shan, Z. Zhang, Y. Qing and D. Wang, Electroanal., Vol. 22 (21),

p.2569-2575, 2010.

76. S. Chen, R. Yuan, Y. Chai, Y. Xu, L. Min and N. Li, Sens Actuators B Chem.,

Vol. 135, p. 236-244, 2008.

77. H. Yang, R. Yuan, Y. Chai, H. Su, Y. Zhuo, W. Jiang and Z. Song, Electrochim.

Acta., Vol. 56 (5), p. 1973-1980, 2011.

78. J. Han, Y. Zhuo, Y. Chai, R. Yuan, W. Zhang and Q. Zhu, Anal. Chim. Acta., V.

746, p. 70-76, 2012.

79. X. He, R. Yuan, Y. Chai, Y. Zhang and Y. Shi, Biotechnol. Lett., Vol. 29, p.149-

155, 2007.