Anal Bioanal Chem (2002) 372 : 664–682
DOI 10.1007/s00216-002-1235-9
REVIEW
María Dolores Marazuela · María Cruz Moreno-Bondi
Fiber-optic biosensors – an overview
Received: 31 July 2001 / Revised: 20 December 2001 / Accepted: 21 December 2001 / Published online: 21 February 2002
© Springer-Verlag 2002
Abstract This article reviews progress and developments
during the past five years in the field of optical fiber biosensors. Because of the expense and time constraints associated with modern laboratory analysis, there is a growing need for real-time, low-cost technology that can be
used industrially, environmentally, and clinically, and to
monitor food processing. Miniaturization, integrated systems, and multianalyte determination have become key
aspects of sensor development and efforts in this direction
will also be discussed, with some pointers to likely directions of future research in the area. The review will provide information about the analytical characteristics and
applications of fiber-optic biosensors classified depending
on the biorecognition element employed – enzymes, whole
cells, antibodies, nucleic acids, and biomimetic polymers.
Keywords Biosensor · Optical sensors · Fiber-optic
Introduction
Biological recognition elements have attracted extraordinary interest in recent years, because of the key role they
play in the development of highly sensitive and selective chemical analysis. According to a recently proposed
IUPAC definition [1], “A biosensor is a self-contained
integrated device which is capable of providing specific
quantitative or semi-quantitative analytical information using a biological recognition element (biochemical receptor) which is in direct spatial contact with a transducer element. A biosensor should be clearly distinguished from a
bioanalytical system, which requires additional processing steps, such as reagent addition. Furthermore, a biosensor should be distinguished from a bioprobe which is
either disposable after one measurement, i.e. single use,
M.D. Marazuela · M.C. Moreno-Bondi (✉)
Optical Sensors Group, Department of Analytical Chemistry,
Faculty of Chemistry, Universidad Complutense de Madrid,
28040 Madrid, Spain
e-mail: [email protected]
or unable to continuously monitor the analyte concentration”. Biosensors that include transducers based on integrated circuit microchips are known as biochips [2].
Specificity and sensitivity should be the main properties of any proposed biosensor. The first depends entirely
on the inherent binding capabilities of the bioreceptor
molecule whereas sensitivity will depend on both the nature of the biological element and the type of transducer
used to detect this reaction [3]. In general, depending on
the recognition properties of most biological components,
two biosensor categories are recognized [2, 3, 4, 5, 6]:
1. Catalytic biosensors. These are also known as metabolism sensors and are kinetic devices based on the
achievement of a steady-state concentration of a transducer-detectable species. The progress of the biocatalyzed reaction is related to the concentration of the analyte, which can be measured by monitoring the rate of
formation of a product, the disappearance of a reactant,
or the inhibition of the reaction. The biocatalyst can be
an isolated enzyme, a microorganism, a subcellular organelle, or a tissue slice.
2. Affinity biosensors. In these the receptor molecule
binds the analyte “irreversibly” and non-catalytically.
The binding event between the target molecule and the
bioreceptor, for instance an antibody, a nucleic acid, or
a hormone receptor, is the origin of a physicochemical
change that will be measured by the transducer.
Biosensor development is driven by the continuous need
for simple, rapid, and continuous in-situ monitoring techniques in a broad range of areas, e.g. medical, pharmaceutical, environmental, defense, bioprocessing, or food technology. Several reviews and books published in recent
years [7, 8, 9, 10, 11, 12, 13, 14] summarize the main
achievements of biosensor research in these areas. These
devices can, because of their sensitivity, selectivity, versatility, ruggedness, and capability of simultaneous multianalyte monitoring, be regarded as an interesting alternative
to conventional techniques for these applications.
Modern optical biosensors have evolved from developments in the communication industry, in information tech-
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nology, and in molecular biology. Although different optical measurement techniques could be used for this purpose,
their application in sensor development has often been
possible as a result of the use of optical fibers. It is, in
fact, usually said that application of optical fibers enables
the scientist “to bring the spectrometer to the sample”.
The development of optical-fiber sensors (OFS) during
recent years is related to two of the most important scientific advances of the 1960s, the laser (1960) and modern
low-cost optical fibers (1966) [15]. During the late 1960s
and early 1970s [16, 17] some of these low-loss optical
fibers were used in the development of the first chemical
sensors; since then their application has continued to
progress and spread to very different areas, e.g. clinical,
environmental, industrial, food, and military applications
[14, 18, 19, 20, 21, 22, 23, 24].
The great success of optical-fiber sensors is because
they can be used to tackle difficult measurement situations
where use of conventional sensors is not appropriate. The
sensors are usually compact and lightweight, minimally
invasive, and can be multiplexed effectively on a single
fiber network. They are immune to electromagnetic interference, because there are no electrical currents flowing at
the sensing point. OFS can survive in difficult environments, e.g. in the presence of high doses of radiation, and
so might have potential in the nuclear industry [25].
The purpose of this review is to outline the basics of
optical fiber biosensors and provide specific information
on recent developments in this field and on challenges
for the future. Most of the references included result from
a computer search of Chemical Abstracts from 1996 to
2001. The review focuses on journal articles and books,
and did not include patents, conference proceedings, reports, or dissertations.
Table 1 Classification of the
biological recognition elements
and signal transducers employed in biosensor development [2], [3], [13], [26]
Biosensor classification
Biosensors can be classified according to either the nature
of the bioreceptor element or the principle of operation of
the transducer. As shown in Table 1 the main types of
transducer used in the development of biosensors can be
divided into four groups [26]: 1. optical, 2. electrochemical, 3. mass-sensitive, and 4. thermometric. Each group
can be further subdivided into different categories, because of the broad spectrum of methods used to monitor
analyte–receptor interactions.
The bioreceptor component can be classified into five
groups [2, 3, 13]:
1. Enzymes, proteins that catalyze specific chemical reactions. These can be used in a purified form or be present in a microorganism or in a slice of intact tissue.
The mechanisms of operation of these bioreceptors can
involve: 1. conversion of the analyte into a sensor-detectable product, 2. detection of an analyte that acts as
enzyme inhibitor or activator, or 3. evaluation of the
modification of enzyme properties upon interaction
with the analyte.
2. Antibodies and antigens. An antigen is a molecule that
triggers the immune response of an organism to produce an antibody, a glycoprotein produced by lymphocyte B cells which will specifically recognize the antigen that stimulated its production [12].
3. Nucleic acids. The recognition process is based on the
complementarity of base pairs (adenine and thymine or
cytosine and guanine) of adjacent strands in the double
helix of DNA. These sensors are usually known as genosensors. Alternatively, interaction of small pollutants
with DNA can generate the recognition signal [2, 27].
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Fig. 1 (A) Total internal reflection of
light at the interface between phases
of refractive indices n1 and n2, where
n1>n2. Incident rays at an angle
greater than the critical angle, θc,
will be totally internally reflected at
the interface. (B) Electric field amplitude E, on both sides of the core/
cladding interface of a waveguide. In
the lower index medium (n2, cladding) the electric field amplitude of
the evanescent wave decays exponentially, with a penetration distance, dp,
that depends on λ, θ, n1, and n2 (see
text)
4. Cellular structures or whole cells. The whole microorganism or a specific cellular component, for example a
non-catalytic receptor protein, is used as the biorecognition element.
5. Biomimetic receptors. Recognition is achieved by use
of receptors, for instance, genetically engineered molecules [28], artificial membranes [29], or molecularly
imprinted polymers (MIP), that mimic a bioreceptor.
The most recent investigations in artificial receptors
include application of a combined approach of computer (molecular) modeling and MIP and the application of combinatorial synthesis for the development of
new sensing layers.
Fundamentals of fiber optics
Optical fibers transmit light on the basis of the principle
of total internal reflection (TIR). When this phenomenon occurs the light rays are guided through the core of
the fiber with very little loss to the surroundings. The optical fiber is formed by a core with a refractive index n1
and a cladding with a refractive index n2 (Fig. 1). For light
propagation by TIR the refractive index of the core (n1)
must be larger than that of the cladding (n2), i.e. n1>n2.
When a ray of light strikes the boundary interface between these transparent media of different refractive index and the angle of incidence is larger than the critical
angle, defined by the Snell’s law (θc=sin–1[n2/n1]), it will
be totally internally reflected and propagated through the
fiber.
When the incident light is totally internally reflected,
its intensity does not abruptly decay to zero at the interface. A small portion of light penetrates the reflecting medium by a fraction of wavelength, far enough for recognition of the different refractive index. This electromagnetic field, called the evanescent wave, has an intensity
that decays exponentially with distance, starting at the interface and extending into the medium of lower refractive
index. The penetration depth (dp), defined as the distance
required for the electric field amplitude to fall to 1/e (0.37)
of its value at the interface, increases with closer index
matching and it is also a function of the wavelength of the
light and the angle of incidence.
The evanescent wave can interact with molecules within the penetration depth, thereby producing a net flow of
energy across the reflecting surface in the surrounding
medium (i.e. that with refractive index n2) to maintain the
evanescent field. This transfer of energy will lead to attenuation in reflectance which can be used to develop absorption sensors based on evanescent waves (attenuated
total reflection (ATR) sensors). When the evanescent light
selectively excites a fluorophore, the fluorescence emitted
can be directed back into the fiber and guided to the detector. This principle is called total internal reflection of
fluorescence (TIRF) and has been widely applied in the
design of immunosensors, as will be described later.
Optical fiber biosensors can be used in combination
with different types of spectroscopic technique, e.g. absorption, fluorescence, phosphorescence, Raman, surface plasmon resonance (SPR), etc. When absorbance is measured
the biological receptor can be immobilized close to the optical fiber or directly on its surface. On interaction with the
analyte, variation of the absorbance properties of the sensitive layer (Beer’s law) will occur and will be related to the
concentration of the species analyzed. One or several fibers
can be used to guide the light from the source to the sensitive tip, and the emerging radiation back to the detector.
Although fluorescence measurements can be used whenever a naturally fluorescent analyte is detected, this is not
common in biosensor development and the technique is
usually applied in combination with artificially labeled
compounds, for instance in competitive immunosensors,
or with fluorescence quenching measurements. This type
of transduction is usually applied in enzymatic biosensors
in which the analyte is biocatalytically converted to a product, or reacts with a compound with optical properties, or
that induces an optical signal. This is so, for instance, with
oxidase-type enzymes, for which enzymatic consumption
of oxygen, as a result of the transformation of the analyte,
is evaluated by measuring the decrease in the luminescent
signal of a fluorescent dye formed by this molecule. Bioor chemiluminescent sensors, in which the analyte induces emission of light on interaction with a bioreceptor,
can also be used.
The measurement scheme for an extrinsic fluorescence
sensor will be similar to that used in absorbance measurements. In this instance the same, or a different, fiber will
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Fig. 2 (A) Representation of
the Kretschmann prism arrangement for the excitation of
surface plasmon resonance
(θ=angle of incidence of light
with the metal surface). (B) Illustration of the sharp minimum produced in the reflected
intensity as a function of the
incident angle, because of SPR
excitation at the resonant angle. (Adapted with permission
from reference [30])
collect the light emitted by the sensing element containing
the fluorescent indicator; this light will then be filtered and
delivered to the detector. Alternatively, molecular recognition can occur at the surface of the fiber core, accompanied by binding or release of a chromophore or luminophore that can be excited by the evanescent wave. In all
instances light intensity, decay time, polarization, or phase
of the emitted radiation can be selected as the analytical
property used to evaluate the concentration of the analyte.
Surface plasmon resonance (SPR) is an optical phenomenon caused by charge density oscillation at the interface of two media with dielectric constants of opposite
sign, for example a metal and a dielectric [30]. When light
of an appropriate wavelength interacts with the dielectric–
metal interface at a defined angle, called the resonance
angle (Fig. 2), there is a match of resonance between the
energy of the light photons and the electrons at the metal
surface. In this situation the photon energy is transferred
to the surface of the metal as packets of electrons called
plasmons and the reflection of light from the metal film
will usually be attenuated. This resonance is experimentally observed as a sharp minimum of light reflectance
when the angle of incidence is varied. Alternatively, SPR
can also be generated by use of a fixed angle, white light,
and spectral detection [31].
The resonance angle will depend on different factors
[32, 33] – the wavelength of the incident light, the metal,
and the nature of the media in contact with the surface.
Any change in the refractive index of the adsorbed layers
at the metal surface will affect the SPR coupling conditions, and produce a shift in the resonance angle. The most
common configurations used to couple light rays into a
surface plasmon mode that exists on the surface of a thin
metal solid film are prisms and diffraction gratings, but
SPR sensors can also be based on optical fibers or integrated optical waveguides [32, 34].
Advantages and disadvantages of fiber-optic sensors
As shown in Table 1, optical biosensors can be used in
combination with different types of spectroscopy – absor-
bance, refraction, luminescence, Raman, etc. – and different light properties can be used for sensor development –
amplitude, energy, polarization, phase, or decay profile.
The incorporation of an optical fiber into a biochemical sensor results in several advantages, e.g.:
1. An enormous background of optically based methods
is available for chemical analysis. Almost every chemical analyte can be determined by use of its spectroscopic properties.
2. Fibers can be used to transmit light over long distances
and the bioreceptor need not be in intimate contact with
the optical fiber, enabling a wider range of non-invasive
configurations.
3. Proper adjustment of the refractive indexes of the waveguide and the surrounding media enables the performance of surface-specific spectroscopy [35, 36]; this is
difficult to apply in condensed media.
4. Fibers have a multiplex capability. Because they can
guide light of different wavelengths at the same time
and in different directions, multiple analyte determinations or single-analyte monitoring in single locations
can be performed with a single central unit.
5. They can be used in harsh environments and are immune to electric or magnetic interference, and so can
be safer than electrochemical biosensors. No reference
electrode is needed and multiwavelength measurements
can be used to correct for drift in optical and electrical
components.
6. They can be easily miniaturized at low cost, thus finding application for in-vivo measurements.
7. A light guide can carry more information than electric
wire.
8. The temperature-dependence of the fiber is lower than
that of electrodes.
Some drawbacks of optical-fiber sensors can limit their
applicability:
1. Interference of ambient light, although this can be
avoided by use of suitable light isolation or modulated
light sources.
2. Possible photobleaching or indicator wash out when
indicator phases are employed, and limited stability of
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the immobilized biological component, although this
would also be a drawback or other types of transducer.
3. Background absorbance, fluorescence, or Raman scatter of the fiber.
4. Long response times if mass transfer to the reagent
phase is needed.
5. Limited availability of optimized commercial accessories for use with optical fibers.
Optical sensors based on the use of fiber optics can be
classified into two different categories:
1. intrinsic sensors, where interaction with the analyte occurs within an element of the optical fiber; and
2. extrinsic sensors, in which the optical fiber is used to
couple light, usually to and from the region where the
light beam is influenced by the measurand. This region
is external to the fiber but can be suitably attached to it,
e.g. by fusion-splicing, gluing, or mechanical connection, that can be easily uncoupled.
Enzyme-based biosensors
Enzymes are the most commonly used biological components of fiber-optic biosensors. The main reasons are:
1. the large number of reactions they catalyze; 2. the possibility of detecting a broad range of analytes (substrates,
products, inhibitors, and modulators of the catalytic activity); and 3. the different transduction principles that can
be used to detect the analyte of interest.
Because enzymes are natural proteins which transform
a specific substrate molecule into a product without being
consumed in the reaction, they can easily be used for continuous biosensing of a specific compound. Among other
advantages [37] enzymes are highly selective and sensitive compared with chemical reactions, they are fairly
fast-acting in comparison with other biological receptors,
and they can be used in combination with different transduction mechanisms. The stability of the enzyme, on the
other hand, determines the lifetime of the biosensor. Many
enzymatically catalyzed reactions require the presence of
other molecules, e.g. inorganic ions (e.g. Fe2+, Mg2+, Mn2+,
or Zn2+), called cofactors, or more complex organic molecules (e.g. nicotinamide adenine dinucleotide, NAD+),
called coenzymes, that can change chemically during the
course of the reaction. Such changes in the optical or redox properties of the co-reactants can also be used to
monitor the course of the enzymatic reaction.
Most enzymes do not have intrinsic optical properties
that change when they interact with the analyte, so most
fiber-optic enzymatic biosensors rely on the use of a chemical transducer (O2, pH, CO2, NH3, etc) the signal from
which can be related to the concentration of analyte in the
sample. The enzymes commonly applied in optical biosensors are: 1. oxidases and oxidoreductases that catalyze
the oxidation of compounds using oxygen or NAD, 2. esterases that produce acids, 3. decarboxylases that produce
CO2, and 4. deaminases that produce NH3. Table 2 summarizes some examples of reported enzymatic fiber-optic
biosensors based on different types of enzyme and transduction schemes.
Most optodes involving oxidase type enzymes described
in the literature are based on the use of oxygen transduction, via luminescence quenching of transition metal com-
Table 2 Fiber-optic enzyme sensors based on different types of transduction mechanisms
Analyte
Enzyme (s)
Solid support
Transducer Indicator/Reagent
Ref.
Glucose
Glucose
Bilirubin
Phosphatidylcholine
Cholesterol
Penicillin
Creatinine
Glucose
Urea
Urea
Pesticides
Carbamates
Paraoxon
Organophosphates
Glutamate
Ethanol
Pyruvate
AMP, ADP, ATP
Acrylamide gel
Controlled-pore glass
Acrylamide gel
Nylon mesh
Graphite powder
Polyvinyl alcohol matrix
Polyvinyl alcohol matrix
Polyvinyl alcohol matrix
Polyvinyl alcohol matrix
Polyacrylamide/PPY film
Isothiocyanate glass
Controlled-pore glass
Langmuir–Blodgett film
Sol–gel glass
Optical fiber surface
Langmuir–Blodgett film
Polyvinyl alcohol matrix
Collagen membrane
Oxygen
Oxygen
Oxygen
Oxygen
Oxygen
pH
pH
pH
pH
pH
pH
pH
pH
pH
NADH
NADH
NADH
ATP
Ru(phen)3Cl2
Pt-octaethylporphine
Ru(dpp)3Cl2
Ru(dpp)3Cl2
Ru(dpp)3Cl2
Aminofluorescein
Aminofluorescein
Aminofluorescein
Aminofluorescein
Polypyrrole (PPY)
Thymol blue
Chlorophenol red
Litmus dye
Indoxyl
Luciferin/Mg2+
[38]
[42]
[39]
[40]
[41]
[58]
[58]
[58]
[58]
[59]
[62]
[63]
[64, 65]
[66]
[48]
[49]
[50]
[53]
Chlorophenols
Lactate
Glucose oxidase
Glucose oxidase
Bilirubin oxidase
Phospholipase/Choline oxidase
Cholesterol oxidase
Penicillinase/penicillin-G amidase
Creatinine iminohydrolase
Glucose dehydrogenase
Urease
Urease
Acetylcholinesterase
Acetylcholinesterase
Acetylcholinesterase
Cholinesterase
Glutamate dehydrogenase
Alcohol dehydrogenase
Lactate oxidase/dehydrogenase
Adenylate/creatine kinase/
firefly luciferase
Horseradish peroxidase
Lactate oxidase
H2O2
H2O2
Luminol
Luminol
[52]
[56]
Choline
Choline oxidase
Collagen membrane
UltraBind/Immunodyne
memb.
DEAE Sepharose-PVA gel
H2O2
Luminol
[57]
669
plexes [38, 39, 40, 41, 42]. These compounds, in particular ruthenium diimine complexes such as tris(1,10phenantroline) ruthenium(II) dichloride (Ru(phen)3Cl2) or
tris(4,7-diphenyl-1,10-phenantroline)ruthenium dichloride
(Ru(dpp)3Cl2), have become very popular as oxygen indicators, because of their high photostability, long excitedstate lifetimes (1–5 µs), high quantum yields (0.1–0.6),
and high Stern–Volmer quenching rate constants in the
presence of oxygen [18, 43, 44, 45].
One of these complexes, Ru(phen)3Cl2 has been used
in the fabrication of the first micron-sized fiber-optic biosensor for glucose [38]. The device, based on the wellknown enzymatic oxidation of glucose by glucose oxidase, was prepared by co-immobilizing the ruthenium complex and the enzyme on an acrylamide polymer that was
attached by photopolymerization to the fiber-optic. Because of its small size, the response time of the sensor was
only 2 s and the absolute detection limit was of the order
of 10–15 mol. The development of small devices is currently one of the main trends in biosensor research [46].
Li et al. [39], have reported the development of a
miniaturized fiber-optic sensor for analysis of bilirubin in
serum samples. The analyte is detected by measuring the
fluorescence quenching of Ru(dpp)3Cl2 by dissolved oxygen during the reaction catalyzed by bilirubin oxidase
(BOx). The enzyme and the oxygen-sensitive dye were
immobilized on an acrylamide polymer that was covalently attached to the tip of a silanized optical fiber by photopolymerization. Miniaturization of the sensor has enabled significant reduction of the response time (10 s) compared with other bilirubin sensors [47] and sub-microliter
samples can be measured accurately. Adsorption of bilirubin and biliverdin on the sensor tip, however, causes passivation of the device, which affects both reproducibility
and sensor lifetime.
Optodes based on Ru(dpp)32+ luminescence quenching
by oxygen have also been used to monitor choline-containing phospholipids in serum samples [40]. The enzyme
is covalently immobilized on a nylon membrane placed
in dip contact with an oxygen-sensitive layer at the tip of
an optical fiber bundle. Sensor response was evaluated, by
use of flow-injection analysis (FI), in the range 0.08 to
3 mg mL–1 phosphatidylcholine; the sensor has been successfully applied to determination of phospholipid in control serum samples. By use of the same measurement
scheme Marazuela et al. [41] developed an optode for analysis of free cholesterol; they used the enzyme cholesterol
oxidase, covalently immobilized in graphite powder deposited on the oxygen sensitive membrane. The chemical
measurement conditions were optimized by use of a supermodified simplex method that has proven very useful in
selection of the optimum working conditions for the device.
A phosphorescent octaethylporphine-ketone platinum
complex dissolved in polystyrene has also been used as an
oxygen-sensitive dye for detection of glucose in a flow system [42]. Two different sensor configurations have been
used – direct immobilization of the enzyme glucose oxidase either on the oxygen-sensitive membrane or on controlled-pore glass in a micro-column reactor. A simple in-
strument equipped with a phosphorescence phase detector
and a fiber-optic probe was used to follow changes in the
oxygen concentration during the enzymatic reaction.
Several fiber-optic sensors have been based on detection of the fluorophore NADH [48, 49, 50]. Glutamate,
one of the most intensely studied neurotransmitter molecules, can be quantified by use of a submicron fiber-optic
sensor containing the enzyme glutamate dehydrogenase
(GDH) directly immobilized on an optical fiber by conventional covalent binding [48]. Glutamate analysis has
been performed by monitoring the fluorescence of NADH
produced during the enzymatic reaction between glutamate and NAD+. A detection limit of 0.2 µmol L–1 was
achieved with an absolute mass detection limit of 3 attomol.
Because of the small size of the optode and its extremely
fast response time (50 ms), it can potentially be used to
monitor the release of glutamate ions in brain cells.
Alcohol dehydrogenase (ADH) immobilized in Langmuir–Blodgett (LB) films has been used in combination
with optical fibers for ethanol measurement [49]. The
Langmuir–Blodgett technique enables fabrication of ultra-thin films of highly ordered enzyme layers on the molecular level; this leads to short response times compared
with other methods of immobilization. In this example
ethanol reacts with NAD+, catalyzed by the ADH-immobilized LB film, and the fluorescence of the product,
NADH, is followed at 455 nm. The ethanol concentration
can be related directly to NADH fluorescence intensity.
Negatively charged arachidic acid was chosen as the most
suitable lipid for adsorption of ADH, on the basis of its
electrostatic force and hydrophobicity compared with
other lipids. The response of the sensor was linearly dependent on ethanol concentrations up to 40 mmol L–1 for
a 20-layer ADH-immobilized LB film.
Zhang et al. [50] have developed a dual-enzyme fiberoptic biosensor for pyruvate by immobilizing lactate oxidase (LOx) and lactate dehydrogenase (LDH) at the sensitive tip. The detection mechanism is based on monitoring
of the consumption of the fluorophore NADH during the
reaction sequence:
LDH

→ Lactate + NAD
Pyruvate + NADH ←

LOx
Lactate + O2 −→ Pyruvate + H2 O2
(1)
(2)
The biosensor showed an 8.0-fold sensitivity; a 5.4-fold
improvement in detection limit and response times shorter
than those observed for single enzyme-based biosensors
[51].
A bioluminescence method has been applied by Jocoy
et al. [52] for determination of NADH at picomolar to atomolar concentration levels. The fiber-optic biosensor uses
two enzymes, NADH oxidoreductase and bacterial luciferase (BL); the latter occurs naturally in photobacteria and
produces light by oxidation of fatty aldehydes, according
to the reactions:
NADH
NADH−oxidored
+ Flavinmononucleotide (FMN)
−→
NAD+
+ FMNH 2
(3)
670
BL
FMNH 2 + O2 + decanal −→ decanoic acid
+ FMN + light(λ = 490nm)
(4)
A more complex system based on the tri-enzymatic sequential reaction of adenylate kinase (AK), creatine kinase
(CK), and firefly luciferase (FL) has been used to determine three adenylic nucleotides, ATP, ADP, and AMP
[53]. The two kinases were covalently co-immobilized on
a collagen membrane and firefly luciferase was bound to
a separate membrane. Both membranes were placed at the
tip of a fiber-optic bundle, with the ATP-generating membrane facing the measurement solution. The sensitivity of
the biosensor for each of the monitored nucleotides can be
improved by proper assembly of the enzymatic membranes. The popular chemiluminescence reaction between
luminol and hydrogen peroxide catalyzed by horseradish
peroxidase (HRP) has served as the basis for different biosensors [54, 55, 56, 57, 58]. Thus, in a fiber-optic biosensor for H2O2 detection [54] HRP was immobilized by microencapsulation in a sol–gel matrix and attached to the
end of an optical fiber. The biosensor was used to measure
H2O2 in a contact-lens-disinfectant solution. Irreversibility is the main disadvantage of this biosensor – it must be
discarded after a single use.
Blum et al. [55] have developed a fiber-optic chemiluminescence sensor for detection of chlorophenols. This sensor is based on the capacity of some phenolic compounds
to enhance the chemiluminescence reaction of luminol catalyzed by HRP. In this device the enzyme is immobilized
on a collagen membrane. Ten chlorophenols were tested
and the greatest sensitivity was obtained for 4-chlorophenol-3-methylphenol. For the other assayed compounds,
the sensitivity decreased in the order 4-chlorophenol>2,4dichlorophenol>2-chlorophenol=3-chlorophenol>2,4,5-trichlorophenol. Sensitivity could be improved by use of purified crystallized luminol.
The electrochemiluminescence of luminol in the presence of hydrogen peroxide has also been recently explored
in the design of biosensors based on oxidase enzymes immobilized on preactivated membranes [56]. Such biosensors have been shown to be sensitive, stable, and free
from matrix interference. This transduction mechanism
has been used in the fabrication of biosensors for glucose,
lactate, and choline [56, 57], with detection limits of the
order of picomol.
Other metabolites of clinical interest, for instance penicillin, creatinine, glucose, and urea have been determined
by use of multianalyte optical sensors based on pH-transduction [58]. The sensitive layers were fabricated by coimmobilization of the corresponding enzymes and a pHsensitive fluorophore, aminofluorescein, in a poly(vinyl
alcohol) matrix.
De Marcos et al. [59] developed a new optical sensor
for determination of urea, on the basis of its enzymatic reaction with urease, which was photoimmobilized with
polyacrylamide on a polypyrrole (PPy) film. The main advantage of this sensor is that no pH indicator dye is needed, because PPy itself acts as support and as indicator. The
need for membrane reconditioning, however, limits the
practicability of this sensor.
Gong et al. [60] have proposed a simplified enzyme
fiber-optic biosensor for assay of uric acid in serum and
urine. Uricase and horseradish peroxidase (HRP) were coimmobilized on bovine albumin, via glutaraldehyde, and
thiamin was used as enzymatic fluorimetric substrate. The
sensing membrane can selectively catalyze the oxidation
of uric acid by dissolved oxygen to produce H2O2, which
oxidizes the nonfluorescent thiamin to a highly fluorescent product, thiochrome, with a wavelength of maximum
emission at 440 nm. The sensor has higher sensitivity than
previously reported sensors; the linear dynamic range is
0.5–5.0 µg mL–1 and the detection limit 0.15 µg mL–1.
Biocatalytic fiber-optic biosensors can also be based
on the use of enzymes that are specifically inhibited by
the analyte. One advantage of this type of sensor is that,
usually, a large number of compounds of a particular class
of chemical can inhibit the enzyme at very low concentrations. Sometimes, however, this interaction between analyte and enzyme is irreversible and requires the presence
of a substrate and sometimes a cofactor or mediator, complicating the assay format.
Most sensors exploiting this mechanism are based on
the use of acetylcholinesterase, an enzyme involved in
neurochemical reactions which is inhibited by pesticides
[61]. Several authors [62, 63, 64, 65, 66] have described
the use of immobilized acetylcholinesterase (AChE) in
optical fiber biosensors used for detection of carbamate
and organophosphate pesticides by use of enzyme-inhibition measurements. These devices are based on detection
of pH changes caused by release of acetic acid during the
enzymatic reaction (Eq. 5):
AChE
Acetylcholine + H2 O −→ Acetic acid + Choline (5)
Different solid supports and pH-sensitive dyes have been
used in the development of these biosensors. Andrés et al.
[62] obtained a bioreactive layer consisting of AChE covalently bound to isothiocyanate glass and mixed with a
pH colorimetric indicator (thymol blue) covalently bound
to aminopropyl glass. The mixture was placed at the tip of
a bifurcated fiber-optic head, integrated in a flow-through
cell. The response time of the sensor to the substrate, acetylcholine, was rather long – approximately 16–22 min
were needed to obtain the maximum signal and recover
the baseline level. The optode was used to determine carbofuran and paraoxon and regeneration of the paraoxoninhibited sensor was achieved by use of 2-pyrimidine aldoxime. The detection limit of the sensor was 3.1 and
24.7 µg L–1, for carbofuran and paraoxon, respectively
(sample volume 5 mL).
By use of the same working principle Xavier et al. [63]
developed an optode for analysis of carbaryl and propoxur
(Baygon), two of the carbamate insecticides most commonly used on vegetable crops. The enzyme AChE was
covalently immobilized on controlled-pore glass beads
(CPG) which were packed in a thermostatted reactor connected to a flow-through cell that contained CPG-immobilized chlorophenol red, placed at the common end of a
bifurcated fiber-optic bundle. Linear dynamic ranges for
determination of carbaryl and propoxur were from 0.8 to
3.0 mg L–1 and from 0.03 to 0.50 mg L–1, respectively, and
671
the detection limits were 25 ng for carbaryl and 0.4 ng for
propoxur, for a pesticide injection volume of 50 µL and
6 min incubation. The AChE-based biosensor has been
successfully applied to the analysis of propoxur in spiked
vegetable samples (lettuce and onion) after ultrasonic extraction of buffered solutions.
Choi et al. [64, 65] used the Langmuir–Blodgett (LB)
film technique to develop a fiber-optic biosensor which
uses an AChE-immobilized viologen LB film for detection of paraoxon. AChE was absorbed on a viologen monolayer by electrostatic forces and the resulting AChE–LB
film was placed in a reactor integrated into a flow system
for reagents and sample introduction. The pH indicator
(litmus dye) was selected on the basis of its absorption
spectra and the transmissibility of the fiber. The response
time and the linear dynamic range of the sensor towards
paraoxon were 5 min and 0–2.0 µg mL–1, respectively.
Silicate glasses obtained by the sol-gel method can be
a promising host matrix for entrapping protein molecules
such as enzymes. The sol–gel material is chemically inert,
thermally stable, and transparent, enabling the use of
fiber-optic transducers. Microencapsulation in the pores
of the sol-gel matrix does not affect the functional groups
responsible for the activity of the biomolecules.
Navas Díaz et al. [66] have developed a sol–gel cholinesterase (ChE) biosensor for determination of organophosphorus pesticides. The sensitive tip, prepared using
tetramethyl orthosilicate (TMOS)-derived crystals doped
with ChE, was attached directly at the end of an optical
fiber. Organophosphorus pesticides were monitored by fluorimetric detection, by measuring inhibition of the enzymatic hydrolysis of a non-fluorescent synthetic substrate
(indoxyl-acetate) to a highly fluorescent product (indoxyl). Fenitrothion, azinphos-ethyl, methidation, naled, and
mecarbam were analyzed, at concentrations ranging from
1.21–11.99 (naled) to 4.9–328.9 µg mL-1 (methidation),
after incubation for 5 min. Detection limits were in the
range 0.12 (naled) to 57.6 µg mL–1 (methidation). The
main disadvantage of the biosensor was rapid loss of activity of the immobilized enzyme with continuous use (50%
of initial activity after seven uses).
Pesticide analysis performed with cholinesterase-based
biosensors usually requires long and sometimes tedious
protocols involving long incubation times in the presence
of the inhibitors and the need for partial regeneration and/
or recovery of enzyme activity, by use of pyridine-2-aldoxime, after exposure to organophosphate pesticides. Also,
because AChE is inhibited not only by organophosphatetype pesticides but also by carbamates, triazines, heavy
metals, and many other compounds, these biosensors are
not selective and are therefore unsuitable for quantification of either a single pesticide or a class of pesticides,
which might be required for monitoring of detoxification
processes.
In contrast, the enzyme organophosphorus hydrolase
(OPH) selectively hydrolyzes a range of organophosphate
esters to form chromophoric products; these can be measured and correlated with the pesticide concentration in
the samples. This is the basis of the fiber-optic enzyme
biosensor reported by Mulchandani et al. [67] which enables rapid, direct, and selective measurement of three
organophosphorus agents (parathion, paraoxon, and coumaphos) without interference from carbamates and triazines. The enzyme OPH was immobilized on a Biodyne
nylon membrane and attached to the common end of a bifurcated optical fiber bundle. Absorption of the products
of the enzymatic reaction – p-nitrophenol (paraoxon and
parathion) and chlorferon (coumaphos) was monitored at
400 and 348 nm, respectively. The sensitivity of the biosensor is lower for coumaphos than for parathion and paraoxon, which is the preferred substrate, and the short response time (2 min) is far superior to those obtained with
AChE-based fiber-optic biosensors [24, 25, 26, 27, 28].
The optode has potential for use as a detector for any
chromatographic separation process used to determine
concentrations of organophosphate nerve agents, either
individually or as a class.
Chen et al. [68] have described a new technique for obtaining a multilayer enzyme assembly on the surface of an
optical fiber and application of the fiber in the development of a chemiluminescence-based biosensor for organophosphorus pesticides. The immobilization approach involves building a multilayer enzyme assembly of alkaline
phosphatase (AP) on the surface of a glass or silica fiber
by use of a bifunctional amino coupling reagent, bis(sulfosuccinimidyl) suberate. The optode has been applied to
the detection of paraoxon, achieving sub-µg mL–1 detection limits when optical fibers with three layers of immobilized enzyme are used.
Whole-cell-based biosensors
The use of whole living cells, rather than isolated biological components, for biosensor development has several
advantages [61, 69, 70]: 1. whole cells are more tolerant
of changes in pH or temperature than purified enzymes;
2. some microorganisms (i.e. bacteria, fungi, yeast, etc.)
can easily be isolated from natural sources (river water,
sediments, soil, activated sludge, etc.); 3. a single cell can
contain all the enzymes and co-factors needed for detection of the analyte; otherwise these would have to be provided separately; 4. measurement is frequently possible
without extensive preparation of the sample, and 5. biosensors can easily be regenerated by letting the cells regrow.
Limitations of this type of biosensor are, however,
longer response times and poorer selectivity compared
with enzyme-biosensors, although this can sometimes be
turned into an advantage for certain applications, for instance pollution screening, where a broad sensitivity spectrum is required [71].
Most of the whole-cell-based biosensors described in
the literature employ microorganisms, animal or plant tissues, or cell receptors as bioreceptors. Some typical examples, described in the literature, are presented below in
more detail.
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Microorganisms
In recent years the use of recombinant DNA techniques in
analytical chemistry has led to the development of the socalled “bioluminescent microbial biosensors”. The general approach consists of fusing an inducible promoter to
a reporter gene that codes for bioluminescent proteins
(e.g. luciferase and β-galactosidase) that emit visible light
in response to the presence of specific substances [72].
Several biosensors using this promoter–reporter–gene concept have been developed for the detection of heavy metals [73, 74, 75] and toxic organic compounds [76, 77, 78].
Corbisier et al. [73] have described a biosensor for
quantification of biologically available heavy metal ions
in environmental samples. The biosensor was prepared
from genetically engineered soil bacteria containing a geneencoding firefly luciferase, which produces a bioluminescent signal in the presence of specific metal ions. Preliminary results have demonstrated the high specificity of the
biosensor for lead and chromate ions. Alginate-immobilized cells are much more stable than those immobilized on
agarose, which lost 84% of their activity within six days.
Luminescent microbial biosensors for Hg2+ have been
developed [74, 75] by use of Escherichia coli cells encoded with firefly luciferase luc as reporter gene under the
control of a mercury(II)-inducible promoter. The lowest
detectable concentration of Hg2+ was 0.1 fmol L–1 and no
interference from other heavy metal ions has been observed, even at concentrations a million times those of
mercury. The biosensor has been successfully applied to
the determination of Hg2+ in urine samples of subjects
with dental amalgam restorations.
Ikariyama et al. [76] have measured highly toxic benzene derivatives by means of a bacterial fiber-optic sensor
containing genetically modified Escherichia coli cells
bearing a firefly luciferase (bioluminescent) gene-fused to
the TOL plasmid, immobilized on a dialysis membrane.
The TOL plasmid of Pseudomonas putida encodes a series of enzymes for degradation of benzene and its derivatives and the firefly luciferase gene is used as a convenient reporter system. In this way transformed E. coli cells
generate luminescence in the presence of aromatic compounds. It should be noted that E. coli cells must be treated with EDTA if they are to produce sufficient luminescence, and that 60 min are required to induce a linear luminescence response. Detection limits of 5 µmol L–1 for
m-xylene can be achieved by use of this biosensor, but the
microbe membrane must be replaced after a single use.
Benzene, toluene, xylene, and their derivatives can be monitored.
Pseudomonas fluorescens HK44 is the first recombinant microorganism approved for field testing in the United States for polycyclic hydrocarbon (PAH) bioremediation purposes. Strain HK44 contains a combination of a
lux (bioluminescent) gene, genetically engineered with a
naphthalene catabolic plasmid, that enables the bacteria to
bioluminesce as they degrade specific PAH, e.g. naphthalene. Fiber-optic biosensors utilizing alginate-encapsulated HK44 cells have been used to monitor the presence of
naphthalene in soil bioremediation studies [77]. The estimated lifetime of these biosensors is approximately one
week; after that they must be replaced with newly encapsulated cells.
The construction and evaluation of a fiber-optic microbial biosensor based on recombinant E. coli cells expressing organophosphorus hydrolase (OPH) has recently been
described for the determination of organophosphate nerve
agents [78]. E. coli cells were immobilized in an agarose
gel inside a nylon membrane and then attached to a common end of a fiber-optic bundle. OPH-expressing E. coli
cells catalyze the hydrolysis of organophosphorus pesticides to form chromophoric products which absorb light.
The biosensor has the potential for selective measurement
of different organophosphates directly in a mixture, without interference from triazine or carbamate pesticides, although the detection limit of 3–5 µmol L–1 is still not low
enough for environmental monitoring. The sensor is, however, very stable (for approximately 1 month) when stored
in buffer at 22 °C and when the sensor was used more than
75 times the response did not decline.
Animal and plant tissues
Whole mammalian tissue slices or mammalian cells cultured in-vitro have been used as biosensing elements.
Plant tissues can also be used in plant-based biosensors –
they are effective catalysts, because of the enzymatic
pathways present [61], and plant tissues have the advantage of low cost compared with mammalian tissue slices.
A fiber-optic biosensor based on immobilized living
algae cells (Scenesdesmus subspicatus) has been developed for the determination of herbicides in wastewater
[79]. Most herbicides inhibit the electron-transport involved in the photosynthetic processes responsible for
ATP production in plants. As a consequence an increase
in chlorophyll fluorescence can be measured and correlated with the concentration of the pollutant. The green algae cells were first cultivated and grown in a miniaturized
LED-photobioreactor. A biosensor was constructed by immobilizing the cells on filter-paper disks covered with a
thin alginate layer and placed at the tip of an optical fiber
bundle. By use of this biosensor herbicides such as atrazine and endrine can be measured in the µg L–1 concentration range with response times of the order of 10 min.
Although an hour in the nutrient medium is required for
biosensor regeneration after use, the immobilized organisms can be stored at 4 °C for approximately 6 months without significant loss of fluorescent properties. The main
application of this device could be for preselection of suspected samples which can then be sent to environmental
laboratories for standard analyses, thus reducing the costs
of monitoring programs.
Naessens et al. [80] have proposed another algal biosensor, based on kinetic measurements of chlorophyll fluorescence in Chlorella vulgaris cells, for detection of herbicides. The microalgae were simply immobilized on removable glass microfiber filter membranes and placed at
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the common end of an optical fiber bundle inside a homemade microcell. The response of the biosensor was evaluated by monitoring the chlorophyll fluorescence intensity
of C. vulgaris after 1 min illumination with 470-nm excitation light. A decrease in the response of the biosensor,
because of loss of biological activity of the algae, was observed after two weeks. If reproducible membranes were
to be obtained the physiological state of the algae was
very important and it was necessary to observe strict conditions during their culture.
The device was particularly useful for the herbicides
(e.g. diuron, simazine, and atrazine) known as photosystem II (PSII)-inhibitors, which could be reversibly detected below the concentration limits required by European Community legislation, although it could also be applied to the assessment of other compounds, e.g. alachlor
and glyphosate. The biosensor could act as an early-warning device for monitoring any of these pollutants in natural waters.
C. vulgaris cells immobilized within alginate beads have
been used in a fluorescence-based fiber-optic biosensor for
toxicity measurements [81]. When the colony of immobilized algal cells is exposed to fluorescein diacetate (FDA)
an intracellular esterase enzymatic reaction occurs which
hydrolyzes the compound, liberating fluorescein, a highly
fluorescent compound. A patented waveguiding system
enables separation of the excitation and emission wavelengths without the need for optical filtering. The rate of
hydrolysis and, thus, the fluorescence intensity are affected by the overall toxicity of the sample. The working
range was linear between 0.25 and 10 mg mL–1 FDA and
the fiber-optic biosensor was capable of detecting organic
and metallic pollutants at levels of environmental interest.
Other types of microorganism can be chosen for exact
matching of the final application required.
Cell receptors
Many proteins found within cells act as bioreceptors for
intracellular reactions that will occur later or in another
part of the cell. These proteins provide a means of molecular recognition by a variety of mechanisms.
On the basis of this biorecognition principle Barker et
al. [82, 83] constructed selective micro- and nanosensors
for determination of nitric oxide, using dye-labeled cytochrome C′ or guanylate cyclase as bioreceptors. The receptors were immobilized at the distal end of an optical
fiber in a polyacrylamide matrix, or by covalent binding
to gold colloids on the fiber. The response of the sensors
was rapid (<1 s) and the response range was a reversible
and linear up to 1 mmol L–1 nitric oxide. The detection
limit, of the order of 1 µmol L–1 nitric oxide, enabled the
analysis of extracellular nitric oxide released by macrophages.
Detection of lipopolysaccharide (LPS) endotoxin by use
of a protein bioreceptor-based evanescent wave fiber-optic biosensor has been reported by James et al. [84]. LPS
is a causative agent in the clinical syndrome known as
sepsis which is responsible for more than 100,000 deaths
per year. A covalently immobilized protein, polymyxin B,
is used as bioreceptor element and the optode is based on
a competitive assay using fluorescently labeled LPS.
Lipopolysaccharide concentrations of 10 ng mL–1 can be
measured in 30 s.
Ignatov et al. [85] have recently described a novel fiberoptic biosensor for detection of lactate. It is based on the
combination of bacterial cytoplasmic membranes (CPM)
expressing lactate oxidase activity and an oxygen-sensitive ruthenium dye as the transducer. To construct the biosensing membranes CPM were adsorbed on to a cellulose
disk mechanically fixed over an oxygen-sensitive siloxane layer, at the distal end of an optical fiber bundle. This
optode specifically detects lactate over the clinical range
of interest, without interference from other common substances found in biological samples, e.g. glucose, fructose, and glutamic acid.
Immunosensors
Immunosensors are an important class of biosensor based
on selective bioaffinity interactions between an antibody
and a specific compound, or antigen, or a closely related
group of antigens. The unique capacity of antibodies to
bind specifically the analyte of interest is the key factor in
their usefulness in immunosensor design. The higher the
affinity of the antibody for the analyte, the better the sensitivity achieved; the selectivity is, on the other hand, a
specific property of the antibody applied, rather than depending on the assay format.
Immunosensors are based on the principles of solidphase immunoassays and the assay formats commonly
employed in immunosensors are shown in Fig. 3 [86],
[87], [88], [89]. In a direct assay, the antigen (for instance,
a naturally fluorescent compound) is incubated with excess amounts of an immobilized antibody and the interaction is detected. The measured signal is directly proportional to the amount of antigen present.
Competitive assay formats are based on the competition between an analyte derivative, either labeled or immobilized, and the analyte in the sample for a limited
number of antibody binding sites (Fig. 3(B)) [13], [89].
An alternative heterogeneous test format is the so-called
binding inhibition assay (Fig. 3(C)) []. In this assay the antibody and the analyte are first pre-incubated and after
equilibration the solution is placed in contact with the immobilized antigen. Only the non-inhibited antibodies (unbound) will bind to the transducer and originate a detectable signal.
In the sandwich assay format the antigens are incubated with an excess of a primary antibody and the resulting
antigen-antibody complex is incubated with a second labeled antibody, which binds to a second antigenic site
(Fig. 3(D)). The amount of labeled antibody bound is related to the analyte concentration. The fact that two antibodies recognize different epitopes on the analyte molecule decreases the chance of interference by other similar
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Fig. 3 Assay formats commonly employed in immunosensors: (A) direct
assay, (B) competitive assay, and (C)
binding inhibition assay. (D) Sandwich assay (for explanation see text)
Table 3 Applications of fiber-optic evanescent wave immunosensors
Analyte
Application(s)
Assay format
Label(s)
Ref.
Explosives (TNT and RDX)
Cyclodiene insecticides
Herbicides
Isoproturon
2,4-D
Cocaine
Coca alkaloids
Benzo[a]pyrene
Protein C
D-dimer
Salmonella spp.
Staphylococcus aureus
Enterotoxin B
E. coli O157:H7
Fumonisin B1
Environmental analysis
Environmental analysis
Environmental analysis
Environmental analysis
Environmental analysis
Clinical analysis
Clinical analysis
Clinical analysis
Clinical analysis
Clinical analysis
Clinical and food analysis
Clinical and food analysis
Clinical and food analysis
Food analysis
Food analysis
Competitive
Competitive
Competitive
Competitive
Competitive
Competitive
Competitive
Direct
Sandwich
Competitive
Sandwich
Sandwich
Sandwich
Sandwich
Competitive
Cy5-labeled antigens
Fluorescein-labeled antigens
HRP-labeled antibodies
Cy5.5-labeled antibodies
Cy5/FITC-labeled antigens
Cy5-labeled antibodies
Fluorescein-labeled antigens
Cy5-labeled antibodies
Fluorescein-labeled antibodies
Cy5-labeled antibodies
FITC-labeled antibodies
Cy5-labeled antibodies
Cy5-labeled antibodies
FITC-labeled antigens
[91, 92, 93, 94]
[98]
[99]
[100]
[101]
[103]
[104]
[105, 106]
[107, 108]
[110]
[115]
[116]
[117]
[119]
[120]
species, but limits the application of this technique to sufficiently large antigens that fulfil this requirement.
Use of antibodies for sensor development has some
limitations: 1. strong dependence of the antibody-binding
capacity on the assay conditions, for instance pH and temperature, and 2. the irreversible nature of the antibody–
antigen interaction. Chaotropic reagents, organic solvents
alone or in combination with acidic buffers, or even ultrasonic radiation [90], have been used as effective agents to
disrupt antibody–analyte association [88].
Optical detection has a clear advantage over electrochemical methods in the development of immunosensors,
because it can be used to monitor binding reactions directly. Optical techniques such as evanescent wave (EW)
and surface plasmon resonance (SPR) have been widely
applied in recent years, providing a way for direct evaluation of the antigen–antibody interactions occurring at the
surface–solution interface. Optical immunosensors have
found application in numerous areas – pollution control,
clinical diagnoses, bacteriological or food analysis, among
others [91]. Some of the latest achievements in this area
are described below and listed in Table 3.
Evanescent wave immunosensors
The main advantage of evanescent wave sensors is their
suitability for measurement of molecules located on or
near the outer surface of the optical fiber, even in the presence of turbid or absorbing solutions. In this type of device the selective receptor molecules are immobilized on
the sensor surface and, in the presence of the analyte of
interest, a biochemical reaction occurs on the surface of
the waveguide and induces a change in its optical properties which is detected by the evanescent wave. Most of the
immunosensors based on this principle (using fiber-optics
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or planar optical waveguides) make use of fluorescencelabeled molecules that can re-emit the absorbed evanescent photons at a longer wavelength by total internal reflection fluorescence (TIRF), as was pointed out in the
section on the fundamentals of fiber optics.
A portable multichannel fiber-optic immunosensor for
on-site detection of the explosives 2,4,6-trinitrotoluene
(TNT) and hexahydro-1,3,5-trinitro-1,3,5-triazine (RDX)
has been optimized on the basis of this sensing scheme
92, 93, 94], [95. The antibodies are covalently immobilized on the surface of an optical fiber by conventional
covalent-coupling techniques. Competition between both
TNT or RDX and their labeled analogues Cy5-EDA(TNB or RDH) for the immobilized antibodies is used as
the basis for the detection. An instrument has been developed which can monitor the response of four different
biosensors simultaneously. Regeneration of the optical
probes is readily achieved by exposure of the fibers to a
solution of 50% ethanol in PBS for 1 min. This enables
analysis of three samples with the same RDX probes and
six with the same TNT probes. The optode has been successfully applied to the simultaneous detection of TNT
and RDX in ground water samples [92, 93, 94] and soil
extracts [95]. The detection limits of the multi-analyte
sensor (5 ng mL–1 TNT and 2.5 ng mL–1 RDX for ground
water and 0.1 µg mL–1 for soil samples) are of the same
order as those approved by the USEPA for the HPLC
method and enable rapid, simple, and cost-effective onsite detection of explosives.
Compared with traditional chromatographic and spectroscopic methods immunosensors have significant advantages in the field of pesticide analysis [96, 97]: 1. because of their high specificity, low detection limits can be
achieved without the need for expensive, time-consuming, and laborious clean-up steps, and 2. immunosensors
are suitable for real time, in-situ or on-line monitoring of
pollutants. For determination of pesticide residues in drinking water, only antibody-based methods conform with
European Community legislation (EEC Directive 80/778)
that set a maximum admissible concentration of 0.1 ng
mL–1 for individual pesticides and 0.5 ng mL–1 for total
pesticides [98]. Immunosensors do, however, suffer from
limitations such as their high prices and, above all, their
limited application to the detection of one particular pesticide or a limited group of structurally similar cross-reacting compounds.
Brummel et al. [99] developed a fiber-optic evanescent
fluoroimmunosensor for detection of cyclodiene insecticides by covalently immobilizing polyclonal antibodies
on quartz fibers. A fluorescent conjugate of the target compound, which fluoresces on binding to the antibody, was
used as the optical signal generator. This immunosensor is
highly sensitive, with a detection limit of 1 nmol L–1 for
chlordane and heptachlor and values approximately 10 to
1000 times higher for dieldrin and aldrin. The immunosensor is best suited for single measurements, because regeneration of the optical probes takes time and eliminates
the advantages of quick screening when many samples
must be analyzed.
A portable fiber-optic immunosensor has recently been
developed for detection of the herbicide methsulfuron
methyl [100]. The ovalbumin (OVA)–methsulfuron methyl
conjugate was immobilized on microscope slides and finally placed inside a detection cell at the tip of an optical
fiber bundle. Competitive immunoreactions between coating haptens and free haptens in the presence of the corresponding HRP-labeled antibodies enable determination of
the concentration of the pesticide in the sample. The portable optical device uses a 0.25-W tungsten–halogen bulb
as light source and a photosensitive diode as detector.
Typical competitive calibration curves span from 0.3 to
100 ng mL–1 methsulfuron methyl with a detection limit
of 0.1 ng mL–1, in perfect accordance with European Community legislation.
An evanescent wave optical immunosensor coupled to
a flow-injection (FI) system has been developed and used
to monitor isoproturon, a phenylurea herbicide, in a variety
of types of water sample [101]. The immunoassay is based
on the competition between an antigen derivative immobilized on clean glass slides, treated with aminodextran,
and a mixture of free and Cy5.5-labeled anti-isoproturon
antibodies. Regeneration of the optode is performed by
rinsing, first with a pepsin solution (2 mg mL–1, pH 1.9)
and then with a 50:50 mixture of acetonitrile and water. Detection limits of 0.10 µg mL–1 isoproturon can be
achieved for real water samples. The biosensor has been
validated by analysis of Aquacheck-certified water samples containing different amounts of phenylurea herbicides; correlation between certified and immunosensormeasured values was good.
Mosiello et al. [102] have constructed an evanescent
wave-based fiber-optic immunosensor for determination
of the herbicide 2,4-D and compared the performance of
two different dyes, Cy5 and FITC, as fluorescent labels.
FITC resulted in a better detection threshold but Cy5 yielded the higher signal-to-noise ratio. The authors also tested
two different assay formats, immobilizing monoclonal antibodies or haptens on the fiber. For the latter the detection
limit achieved, 60 µg L–1, was too high for environmental
applications.
Since their appearance immunosensors have become
widespread in the field of clinical analysis [103]. A fourchannel fiber-optic instrument using a competitive fluorescence immunoassay has been developed for analysis of
cocaine and its metabolites (COC) in human urine [104].
Binding of the cyanine labeled-antibenzoylecgonine monoclonal antibodies (Cy5-Ab) to the casein-benzoylecgonine-Ag immobilized on the optical fibers was inhibited
by COC. An effective benzoylecgonine (BE) concentration range of 0.75–50 ng mL–1 was measured with an assay time of 200 s, including fiber regeneration. On average each fiber was used for 11 measurements.
Toppozada et al. [105] have reported a fiber-optic biosensor for rapid detection and quantification of coca alkaloids, as cocaine equivalents, in leaf extracts. Monoclonal
antibodies against benzoylecgonine (BE), a major metabolite of cocaine, were covalently immobilized on quartz
fibers and used as the biological sensing element in the
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portable fluorimeter. When present cocaine competes for
the antibody-binding sites with fluorescein-labeled benzoylecgonine (BE-FL), reducing the fluorescence transmitted by the fiber. Fibers (mAb-coated) stored at 37 °C in
PBS solution (0.02% NaN3) gave stable responses for 14
days. The sensor can ultimately be used for plant taxonomic identification, which can prove useful in illicit drug
analysis.
By using nanotechnology Vo-Dinh et al. [106, 107]
have developed submicron fiber-optic immunosensors for
the measurement of biomarkers related to human health
which are associated with exposure to polycyclic aromatic
hydrocarbons (PAH). Optical fibers of less than 1-µm diameter at the distal end were coated with antibodies for
benzo[a]pyrene tetrol (BPT), a metabolite of the carcinogen benzo[a]pyrene. The absolute detection limit of these
immunosensors has been determined to be 300 zeptomoles (10–21 mol).
A real-time fiber-optic-based biosensor has been described for the detection of protein C (PC), one of the human body’s key anticoagulants, deficiency of which can
lead to massive thrombotic complications [108, 109]. Monoclonal antibodies against PC have been immobilized on
the surface of a quartz fiber enclosed in a 300-µL chamber. After incubation with the sample containing protein C
the fiber was probed with a fluorophore-tagged secondary
antibody, the fluorescence intensity of which can be correlated with the concentration of the protein. Regeneration
of the fibers was achieved by washing with calcium-containing PBS solution at pH 9.0. Under these conditions an
average of six consecutive assays was achieved per fiber.
The device has a linear working range of 0.2–2.0 µg mL–1,
which includes the PC range of heterozygous deficient patients.
The same optical immunosensing scheme has been employed by other authors for detection of other small biomolecules, for example D-dimer, the most specific marker
of sepsis syndrome and thrombotic disorders [110]. Dynamic ranges extended from 25 ng mL–1 to 1 mg mL–1,
which correlated perfectly with those obtained by ELISA
assays. Fiber-optic immunosensor analysis, moreover, required significantly less time (typically 11 min) and skill
than the standard ELISA method. In another example fluorescein-labeled D-dimer antibodies were immobilized on
the tip of an optical fiber by means of a sol–gel technique
[110]. When D-dimer antigens bind to the encapsulated
antibodies the fluorescence intensity decreases. The immunosensor is sensitive to the analyte in the clinically-relevant range (0.54 to 6 µg mL–1). The optode was not completely reversible, however, because changes in the spectroscopic properties of the fluorescein-tagged antibody occurred as the gel aged. The D-dimer antibodies remained
partially viable for at least four weeks when encapsulated
in the sol–gel network and stored at 4 °C in PBS solution.
Another area in which fiber-optic immunosensors have
found numerous applications in the last few years is the
analysis of food toxins [112, 113]. A increasing number of
bacteria has been identified as important food- and waterborne pathogens, and infection diseases caused worldwide
by these microorganisms account for nearly 40% of the
total 50 million annual estimated deaths [114, 115].
Salmonellosis is an infectious disease caused by Salmonella spp., dangerous food-borne pathogenic bacteria.
A miniaturized fiber-optic biosensor for detection of Salmonella spp. has been reported; it is based on a sandwich
binding format immunoassay [116]. The portable sensor
system, equipped with fiber optics and semiconductor
laser excitation, can be operated in ambient light. The sensor is rugged, cheap, stable, and can be used to detect Salmonella at concentrations as low as 104 colony-forming
units mL–1.
Staphylococcus aureus is another bacteria causing foodborne diseases – it can produce several types of enterotoxin that cause gastroenteritis. The presence of the bacterium in processed food is, therefore, a health hazard that
must be controlled. Chang et al. [117] have proposed an
optode for detection of protein A, a product secreted by
S. aureus only, based on a sandwich immunoassay. The
detection limit of the immunosensor was 1 ng mL–1 protein A. Whereas official US Food and Drug Administration methods for the analysis of the bacterium normally
take from 5 to 7 days, the assay time using this biosensor
can be reduced to 24 h.
A new portable fiber-optic evanescent wave immunosensor, commercially known as “Analyte 2000” (Research International, Woodinville, WA, USA), has been
developed and applied to the detection of Staphylococcal
enterotoxin B (SEB) [118]. A 635-nm laser diode, used as
excitation source, is connected to up to four different tapered fiber-optic probes. The measurement principle is
based on a sandwich immunoassay, using a rabbit antiSEB capture antibody, covalently attached to the optical
fibers, and a sheep anti-SEB antibody labeled with Cy5
for detection. The toxin can be monitored at levels as low
as 0.5 ng mL–1 in buffer and quantitated in other relevant
media, for example human serum, urine, and an aqueous
extract of ham. Quantitative results from four individual
fiber probes can be obtained in 45 min, including calibration; qualitative results can be obtained in 15–20 min.
Analysis of ham samples spiked with 5 and 40 µg SEB
per 100 g food resulted in recoveries of 11 and 69%, respectively. These values make the immunosensor useful
for on-site analysis of suspected food samples. King et al.
[119] have successfully applied a later version of the “Analyte 2000”, a prototype called Mantis consisting of an
automated, self-contained, portable fiber-optic immunosensor, for remote identification of airborne bacteria for
military purposes.
The same instrument has also been used to detect
E. coli O157:H7, a rare strain of E. coli considered to be one
of the most dangerous food-borne pathogens [120] and
which produces large quantities of a potent toxin that can
cause severe hemorrhagic colitis, or even hemolytic uremic syndrome, and which can lead to death, specially in
children. Because E. coli can easily contaminate ground
beef, raw milk, and chicken, careful control of this pathogen is extremely important in food production. The analysis is based on a sandwich immunoassay, using polyclonal
677
capture antibodies specific for E. coli O157:H7 immobilized on the fiber probe by biotin–avidin interactions. Cy5labeled polyclonal anti-E. coli O157:H7 antibodies are
used to generate the fluorescent signal. The immunosensor could be used for specific detection of E. coli O157:H7,
in the range 3–30 CFU mL–1, in seeded samples of ground
beef. Compared with conventional microbiological tests,
which often require several days, assays with the optode
can be completed within 20 min, leading to an almost
real-time system for E. coli O157:H7 determination.
Thompson et al. [121] have developed a fiber-optic
evanescent wave immunosensor by using a direct competitive assay for measurement of fumonisin B1 (FB1), a
mycotoxin produced by common fungal contaminants
of corn. Monoclonal antibodies were covalently bound
through an heterobifunctional silane to an etched 800-µm
core optical fiber. The sensor has a working range from 10
to 1000 ng mL–1 and a limit of detection of 10 ng mL–1
FB1. The fiber-optic biosensor is highly selective toward
fumonisins and does not cross-react with other structurally related compound. The optode can be regenerated up to
20 times before loss of sensitivity is observed. It has been
successfully used for analysis of FB1 in spiked and naturally contaminated corn, either by dilution or after affinity
column clean-up. The results obtained with the immunosensor agreed well with those from an established HPLC
method for the analysis of these compounds.
Fiber-optic surface plasmon resonance immunosensors
Surface plasmon resonance technology (SPR) is having a
substantial impact on immunosensing by addressing the
limitations of earlier methods in terms of specificity and
sensitivity. SPR has been successfully incorporated into
the development of immunosensors for simple and rapid
assay of several analytes in biotechnology, environmental
and clinical analysis and, recently, proteomics; labeling is
not required. Extended reviews about this topic can be
found in the literature [34, 122, 123, 124, 125] and on recommended web sites [126, 127, 128]. SPR has two inherent advantages over other types of biosensor [129, 130]:
1. high sensitivity – it can be used to detect subfemtomole
levels of proteins in complex fluids, and 2. the possibility
of monitoring the binding kinetics of receptor–ligand interactions in real time, without the need for fluorescent or
radioisotope-labeling of the analytes. Although many types
of SPR sensor have been developed, we focus only on
those based on the use of optical fibers. Fiber-optic SPR
are rather small, enabling use of the technique at remote
sites. Sensors based on monomode and multimode fibers
have been reported [131, 132, 133, 134]. The use of optical multimode fibers for SPR sensors was first proposed
by Jogerson et al. [131]. The cladding of the fiber was
partially removed and a gold layer was deposited symmetrically around the exposed fiber core. This type of
sensor configuration limits the interaction zone to a few
millimeters. Another drawback of using multimode fiber
SPR sensors is the difficulty of depositing a homogeneous
coating and achieving good chemical functionalization of
the sensor surface, because the modal light distribution is
affected by mechanical and surface changes. Although
several companies have developed commercially available SPR-based immunosensors [126, 127, 128], few are
based on fiber optics technology. One example is the
BIAcore probe, a fiber-optic SPR version launched by the
Swedish company BIAcore that enables dipstick-type
sampling [126].
Nucleid acid biosensors
During recent years the development of genosensor technology has attracted considerable interest, because of its
importance in DNA sequencing and the early diagnosis of
infection and genetic diseases. The decoding of the human genome has also significantly promoted this development. The favorite transduction principle employed in this
type of sensor has been the optical detection of fluorescently labeled oligonucleotides by use of a competitive
assay.
As an example, an evanescent wave fiber-optic sensor
for detection of 16-mer oligonucleotides in DNA hybridization assays has been described by Abel et al. [135]. The
optode is based on biotinylated capture probes immobilized on quartz fibers by use of the (strept)avidin–biotin
affinity system. Hybridization with fluorescein-labeled
complementary strands was monitored in real time by fluorescence detection. Efficient regeneration of the sensor
surface can be achieved by chemical treatment with a 50%
(w/w) solution of urea. Under these conditions, hundreds
of consecutive assay cycles can be performed with the
same fiber. Sixty-minute incubation times enable the detection of complementary labeled oligonucleotides at a
concentration of 2.0×10–13 mol L–1 (24 fmol) with a working range between 10–13 and 10–8 mol L–1, whereas competitive hybridization assays for the detection of unlabeled complementary oligonucleotides resulted in a detection limit of 1.1×10–9 mol L–1 (132 pmol).
In a similar approach, Kleinjung et al. [135] have reported the specific determination of DNA oligonucleotides by use of a TIRF-based fiber-optic sensor. In this
work 13-mer oligomers were immobilized on the activated surface of silica fibers, either by direct coupling or by
use of the avidin–biotin bridge. Hybridization assays have
been performed by using fluorescent double-strand DNA
ligands with a sequence complementary to that of the immobilized oligonucleotides. Detection limits of 30 fmol L–1
(3.2 attomoles) have been achieved by use of the direct immobilization method. Regeneration of the fiber
surface was possible only by heating at 90 °C, which is
not compatible with the avidin coupling. Thermal regeneration also enables more than 60 cycles to be performed
without loss of sensitivity. When stored dry at –18 °C the
fibers were stable for several months. The simultaneous
analysis of multiple DNA sequences has been tackled by
Ferguson et al. [137] using a fiber-optic biosensor array.
Seven amine-functionalized fibers (typically diameter
678
200 mm) with different cyanuric chloride-activated oligonucleotide probes immobilized on their distal ends were
assembled in a bundle. The hybridization of fluorescently
labeled oligonucleotides, complementary to one or more
probes, was monitored by measuring the increase in the
fluorescence signal upon binding. Chemical regeneration
of the DNA probes was achieved by dipping the fiber tip
in a solution of 90% formamide in TE buffer (10 mmol L–1
Tris-HCl, pH 8.3, 1 mmol L–1 EDTA) for 10 s. The device
enables rapid (<10 min) and sensitive (10 nmol L–1) detection of multiple DNA sequences, simultaneously, with
the potential for quantitative hybridization analysis. When
the fiber tips are stored at 4 °C they retain their sensing capabilities for months.
A novel technique for preparing bead-based fiber-optic
oligonucleotide arrays has recently been described [138].
It is based on the attachment of different DNA probes to
microspheres assembled in microwells created at the distal end of an optical fiber on exposure to a chemical etching reagent. By applying this procedure Lee et al. [139]
developed a fiber-optic biosensor for measurement of
thrombin, by use of a DNA aptamer receptor immobilized
on the surface of silica microspheres and distributed in
microwells at the distal end of an imaging fiber. The device could be used to detect 1 nmol L–1 thrombin and each
test could be performed in ca. 15 min, including regeneration. Degradation of the activity of the aptamer beads
was not observed during measurements for 8 h and the
beads were stable during storage for over 3 months.
A chemiluminescence-based DNA fiber-optic sensor
which can be used to detect amounts of hybridized nucleotides down to 0.1 µg mL–1 has been described [140].
The DNA probes were covalently immobilized on the distal end of an optical fiber bundle and subsequent hybridization by use of HRP-labeled complementary oligonucleotides was detected as enhanced chemiluminescence.
Kleinjung et al. [141] have fabricated a biosensor for
L-adenosine by employing high-affinity RNA as molecular-recognition element. High affinity RNA strands were
attached to the core of a multimode optical fiber by means
of avidin–biotin interactions and binding of FITC-labeled
L-adenosine was measured. Competitive inhibition assays
enabled the device to detect L-adenosine in the submicromolar range.
The DNA biosensors described above have some limitations, mainly related to the need for labeled targets, intercalation reagents, or competitive assays, which make it
difficult to perform real-time hybridization studies or to
monitor hybridization kinetics quantitatively on the sensor surface.
To overcome these problems Liu et al. [142] proposed
a novel optical fiber evanescent wave DNA biosensor
which used a newly synthesized molecular beacon DNA
probe. Molecular beacons (MB) are oligonucleotide probes
that become fluorescent on hybridization with target DNA
or RNA molecules. Biotinylated MB have been designed
and immobilized on an optical fiber surface by means of
biotin–(strept)avidin interactions. The MB-based DNA
biosensor can be used for real time detection of target
DNA and RNA molecules without the use of competitive
assays. The sensor is rapid, stable, highly selective, and
reproducible, with a detection limit of 1.1 nmol L–1 DNA,
and can be regenerated by immersion in a solution of 90%
formamide in a TE buffer for 1 min at room temperature.
Regeneration is, however, less successful after repeated
assay cycles. The optode has been applied to the analysis
of specific rat γ-actin mRNA sequences amplified by
polymerase chain reaction.
Immobilized fluorescent nucleic acid stains, acting
both as molecular recognition elements and as fluorescent
reporters, have also been used in the development of optical bacterial sensors [143, 144]. Chuang et al. [143] have
described a biosensor that uses SYTO 13, a commercially
available green fluorescent nucleic acid stain the intrinsic
fluorescence of which increases by a factor of 40 when
bound to nucleic acids (DNA or RNA). An aqueous solution of SYTO 13 (2-µL aliquots) was pipetted directly on
to the distal end of individual optical fibers. The solution
was left to dry for 20 min, resulting in a thin film of
SYTO 13 on the distal end of the optical fiber. Regeneration of the sensing surface, after each measurement, was
achieved by cleaning the tips with damp paper wipes then
sterilization in a lighter flame for 2 s. The sensor responds
to aqueous and aerosol bacterial samples in 15 and 30 min,
respectively, with a detection limit of 2.4×105 cells mL–1.
The utility of the sensor has been demonstrated by monitoring the growth of a Pseudomonas aeruginosa cell culture over a period of 50 h.
Although most DNA biosensors are based on the use
of fiber-optic probes, the development of integrated array
techniques during the last two decades, leading to so-called
bio or DNA chips [145, 146, 147], could offer a unique
combination of performance capabilities and analytical
features for the simultaneous analysis of many DNA fragments. Biochips have advantages in size, performance, fabrication, multianalysis capabilities, and production cost, because of their integrated optical-sensing microchips. We
will not discuss these devices in detail, however, because
they are beyond the scope of this review.
Biomimetic sensors
The design and synthesis of artificial receptor systems
with antibody-like recognition properties have recently attracted much interest [148, 149, 150, 151]. One technique
being increasingly adopted for the generation of artificial
macromolecular receptors is molecular imprinting of synthetic polymers (MIP). The molecular imprinting technique (Fig. 4) consists, basically, in mixing the analyte
species, the “templates”, with monomers and a large
amount of crosslinker. The template molecules or ionic species form covalent or non-covalent bonded complexes with
the functional monomers and the complexes are then
polymerized with excess cross-linking agent, in the presence of a porogenic solvent, to form a resin. On removal
of the template species (Fig. 4) cavities are formed in the
polymer matrix; the size, shape, and binding site distribu-
679
Fig. 4 Schematic representation of: (A) non-covalent and
(B) covalent molecular imprinting procedures
tion of these enables selective rebinding of the template
from a mixture of chemical species.
MIP are intrinsically stable and robust in aqueous and
organic solvents under extreme pressure and temperature
conditions; they are cheap to produce and can be stored in
the dry state. In the field of optical sensors the imprinting
technique enables the possibility of single or multicomponent analysis for in-situ monitoring and combination of
MIP with pre-organized supramolecular hollows can lead
to more sensitive and selective sensor layers even for determination of isomeric analytes [148]. Their use as recognition elements in the development of sensors, as substitutes for biomolecules, could promote the application
of these devices in areas such as industry, medicine, or environmental analysis [152, 153, 154] and the replacement
of less stable biosensors.
MIP can be prepared in different configurations depending on their final application [154]. The most common approach consists in the preparation of macroporous
block polymers which are ground and sieved to the appropriate diameter, usually of the order of micrometers. Another possibility is the grafting and/or coating of the imprinted polymer on preformed particles (e.g. silica particles). Spherical beads of controlled diameter can also be
prepared by suspension, emulsion, or dispersion polymerization. The polymerization can be performed in-situ inside a chromatographic column or in a capillary. Finally,
thin films, or membranes prepared by casting an imprinted polymer into the pores of an inert support, can also be
obtained and are specially suitable for sensor development.
Many designs are possible for optical sensors based on
MIP receptors, depending on the target analyte and its optical properties. When the analyte has optical properties
that can be used for its recognition the MIP simply acts as
an adsorbent or extractant material for preconcentration of
the analyte, which can be directly monitored. As an example, a fluorescence-based fiber-optic sensor has been
constructed for dansyl-phenylalanine [155]. The sensor
has some stereoselectivity for the L form of the analyte,
which was the original template, although the time required to furnish a stable fluorescence signal was approximately 4 h, too long for a sensor. Other authors [156]
have used polyurethanes imprinted with different polycyclic aromatic hydrocarbons (PAH) in conjunction with
fluorescence measurements in a flow system. The sensitivity of the system for PAH detection was rather high
(ng L–1 range), enabling enrichment factors up to approximately 107.
A flow-through sensing approach has recently been reported for detection of the fluorescent analyte flavonol
[157]. The analyte was enriched in a flavonol-imprinted
polymer contained in an optical detection cell; this enabled its selective detection at nanomolar concentrations.
The polymers were found to be mechanically stable after
continuous use for more than two months. The sensor has
been successfully used for determination of the flavonoid
content of olive oil samples, without the use of a pretreatment step.
When the analyte lacks a specific optical property that
can be monitored by use of a detector a competitive or
displacement assay format can be used. A labeled analyte
derivative is usually allowed to compete with the analyte
for the binding sites in the MIP. Alternatively, the labeled
analyte is allowed to bind first and is subsequently displaced on binding of the analyte. These detection schemes
can be used in combination with fluorescent or colored labels. The feasibility of displacement-sensing assays has
been demonstrated for the determination of the antibiotic chloramphenicol (CAP) by use of a chloramphenicol–
methyl red conjugate [158]. The conjugate is adsorbed in
the CAP-imprinted sites of the polymer, which is packed
into an HPLC column. On injection of the analyte some of
the conjugate is displaced, generating a colorimetric signal that can be related to the analyte concentration.
680
Another possibility for detection of analytes with no
native optical properties is to use a signal generated by the
polymer itself. One example of such a format is an optical
sensing system for chloramphenicol in which a fluorescent dye is incorporated into the MIP and its fluorescence
is quenched on binding of the analyte [159]. Although
these detection schemes have not yet been used in the development of integrated sensors, they seem to be very
promising for this type of application, because they do not
depend on a special property of the analyte.
Jenkins et al. [160] have described a very sensitive sensor for a hydrolysis product of the chemical warfare agent
Soman. The device is based on a polymer-coated fiber-optic probe; a luminescent europium complex was used for
detection. The complex of europium ligated by divinylmethyl benzoate (ligating monomer) and by the analyte
pinacoyl methylphosphonate was copolymerized with styrene; the analyte molecule was then removed by washing.
Detection limits as low as 7 ng L–1 were achieved.
Although use of MIP has great potential in the field of
optical sensors, there are few examples of the application
of this type of material to the solution of real analytical
problems. In years to come we will probably assist in the
development of useful methodologies incorporating this
type of material for the design of sensors suitable for use
under harsh conditions when biosensors might pose a
problem.
Conclusions
Biosensors have advanced substantially in recent years.
The numbers of publications and commercially available
technologies in this field have increased exponentially, at
the same time providing much information about the biology of the molecular receptors used in their design. Nowadays, biosensor research is directed toward the development of simple applications that can solve specific problems, otherwise difficult to tackle, for instance diabetes
control. Miniaturization technologies and the application
of integrated systems that include probes, samplers, and
detectors, etc., is another challenging area of research, for
instance for medical or environmental analysis. Progress
in this area will enable the design of fully automated analytical systems capable of multianalyte determination that
in combination with analytical and discriminative mathematical methods will reveal the real capabilities of optical
biosensors. More research is still needed for the development of more robust biomolecules or biomimetic receptors, that can be used as bioreceptors for sensor design.
Validation is still the workhorse in this area and is a crucial issue in the development and integration of these devices in the real world, and in achieving the confidence of
the potential users of these technologies.
Acknowledgements The authors acknowledge the Spanish government agency CICYT under contracts no. AMB98-1043-C01/C02,
PPQ2000-0778-C02/02, and the Madrid Community (contract
07M/0082/2000) for financial support. M.D.M. thanks the Spanish
Ministry of Education and Culture for a post-doctoral contract.
The authors acknowledge Dr Luigi Canavacciuolo for his assistance in the literature search.
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