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
3
(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.
4
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,
11
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.
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