Báo cáo khoa học: Piezoelectric sensors based on molecular imprinted polymers for detection of low molecular mass analytes potx

MINIREVIEW
Piezoelectric sensors based on molecular imprinted
polymers for detection of low molecular mass analytes
Yildiz Uludag
˘
1,2
, Sergey A.Piletsky
1
, Anthony P. F. Turner
1
and Matthew A. Cooper
2
1 Cranfield Health, Cranfield University, Silsoe, UK

2 Akubio Ltd, Cambridge, UK
Introduction
Biosensors are analytical devices that comprise a sam-
ple-delivery mechanism with a biological recognition
element and a suitable transducer, usually coupled to
an appropriate data-processing system (Fig. 1). The
biological recognition element is typically an enzyme,
microorganism, cell, tissue or other bioligand [1] and
the transducer is required to convert the physico-chem-
ical change resulting from the interaction of molecules
with the receptor into an electrical signal. Over the
past decade the benefits of label-free analysis have

begun to gain a foothold as a mainstream research
tool in many laboratories [2,3]. These techniques do
not require the use of detection labels (fluorescent,
radio or colorimetric) to facilitate measurement; hence
detailed information on an interaction can be obtained
during analysis while minimizing sample processing
requirements and assay run times [4]. Unlike label-
and reporter-based technologies that simply confirm
the presence of the detector molecule, label-free tech-
niques can provide direct information on analyte bind-
ing to target molecules typically in the form of mass
addition or depletion from the surface of the sensor

substrate [5].
Of the various label-free detection modalities, piezo-
electric sensing has become popular with researchers
because of the low barriers to entry, inherent simp-
licity, ease-of-use, low cost, and speed to result.
However, there are relatively few examples of small-
molecule detection using traditional immunoassay
formats: niacinamide detection has been achieved in
serum and urine [6], the endogenous cofactors NAD
+
and NADP
+

have been titrated against the enzyme
Keywords
drug; hapten; label-free; molecularly
imprinted polymer; QCRS; quantification;
quartz crystal microbalance; screening;
small molecule
Correspondence
M. A. Cooper, Akubio Ltd, 181 Cambridge
Science Park, Cambridge CB4 0GJ, UK
Fax: +44 1223 225 336
Tel: +44 1223 225 326
E-mail:

(Received 6 July 2007, accepted 24 August
2007)
doi:10.1111/j.1742-4658.2007.06079.x
Biomimetic recognition elements employed for the detection of analytes are
commonly based on proteinaceous affibodies, immunoglobulins, single-
chain and single-domain antibody fragments or aptamers. The alternative
supra-molecular approach using a molecularly imprinted polymer now has
proven utility in numerous applications ranging from liquid chromatogra-
phy to bioassays. Despite inherent advantages compared with biochemi-
cal ⁄ biological recognition (which include robustness, storage endurance
and lower costs) there are few contributions that describe quantitative ana-
lytical applications of molecularly imprinted polymers for relevant small

molecular mass compounds in real-world samples. There is, however, sig-
nificant literature describing the use of low-power, portable piezoelectric
transducers to detect analytes in environmental monitoring and other appli-
cation areas. Here we review the combination of molecularly imprinted
polymers as recognition elements with piezoelectric biosensors for quantita-
tive detection of small molecules. Analytes are classified by type and
sample matrix presentation and various molecularly imprinted polymer
synthetic fabrication strategies are also reviewed.
Abbreviations
MIP, molecularly imprinted polymer; QCM, quartz crystal microbalance.
FEBS Journal 274 (2007) 5471–5480 ª 2007 Akubio Ltd. Journal compilation ª 2007 FEBS 5471
glucose dehydrogenase and rank order binding to the

enzyme has been determined [7], and biotin has been
detected with a high-frequency microfluidic acoustic
biosensor using an anti-biotin serum [8]. Real-time
detection of 4-aminobutyrate (one of the main inhibi-
tory neurotransmitters) was achieved with an anti-
(4-aminobutyrate) serum with a minimum detection limit
of 38 lm [9]. Kurosowa et al. [10] reported a portable
dioxin sensor able to detect 2,3,7,8-tetrachlorodibenzo-
p-dioxin. Dioxin is well-known as a highly toxic com-
pound that poses a threat to the environment. The
quartz crystal microbalance (QCM) sensor surface
was prepared by immobilizing anti-(2,3,7,8-tetrachlo-

rodibenzo-p-dioxin) serum and a linear calibration
curve was created using 100–0.1 ngÆmL
)1
2,3,7,8-tetra-
chlorodibenzo-p-dioxin before detection of the com-
pound in fly ash samples.
Hence mass-sensitive acoustic immunoassays can
provide a label-free method for detecting and analysing
molecules. However, biological materials are expensive,
sensitive to harsh conditions and their shelf life on the
sensor surface can be limited. The use of molecularly
imprinted polymers (MIPs) as synthetic receptors pro-

vides an attractive alternative to biological receptors.
MIP recognition elements provide sensor surfaces that
have a long shelf life, are robust and simple to prepare,
and provide a 3D interfacial matrix layer with high
binding capacity for analytes. Such binding capacity is
crucial when the molecular mass of the analyte is
< 500 Da. This review summarizes the key approaches
to incorporating imprinted polymers as recognition ele-
ments in piezoelectric sensors for detection of small
molecules.
Acoustic biophysics
Acoustic biosensors allow the label-free detection of

molecules and the analysis of binding events. In gen-
eral, they are based on quartz crystal resonators, which
are found in electronic devices such as watches, com-
puters and televisions, with over one billion units
mass-produced each year. Quartz crystal is a piezoelec-
tric material which mechanically oscillates if an alter-
nating voltage is applied. A QCM consists of a thin
quartz disc sandwiched between a pair of electrodes.
The mode of oscillation depends on the cut and geom-
etry of the quartz crystal. If mass is applied to the sur-
face of the quartz resonator, the frequency of the
oscillation decreases. By measuring the change of fre-

quency, it is possible to determine the change in mass.
Measurement of mass using quartz crystal resonators
was first examined by Sauerbrey [11], who showed that
the frequency change of the crystal resonator is a linear
function of the mass per area m
s
, or absolute mass Dm:
Df
m
¼À
f
2

0
F
q
q
q
m
s
¼À
f
2
0
F

q
q
q
Dm
s
A
el
ð1Þ
where f
0
is the resonance frequency of the unperturbed
quartz resonator, F

q
the frequency constant of the
crystal (F
q
¼ f
0
.d
q
), d
q
the thickness, q
q

the mass den-
sity, and A
el
the electrode size of the crystal resonator.
Equation (1) is valid only for thin, solid layers depos-
ited on the resonator.
Initially, the QCM system was used for dry measure-
ments, later when suitable oscillator circuits were
developed, it was possible to carry out measurements
under liquid conditions [12]. This led to the use of
QCM systems as biosensors to detect molecular inter-
actions (Fig. 2). A new equation was derived by

Kanazawa and Gordon to explain the relationship
between density (q
l
) and viscosity (g
l
) of the liquid and
the frequency of the quartz crystal resonator:
Df ¼Àf
3=2
q
ffiffiffiffiffiffiffiffiffiffiffiffi
q

1
g
1
pq
q
l
q
r
ð2Þ
where q
q
and l

q
are the quartz density and shear mod-
ulus, respectively [13]. In a two-layer system these
frequency shifts simply add up to an overall shift:
Df ¼ Df
m
þ Df
l
¼Àf
2
0
Dm

s
F
q
q
q
A
el
þ
ffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi
g
1
q

1
f
0
pl
q
q
q
r
!
ð3Þ
In addition to the frequency shift, there also exists a
dampening of the resonator caused by the viscous

Fig. 2. Schematic representation of quartz crystal resonance
sensing.
Antibody/Protein
Enzyme
Microorganism
Cell
Recognition elements
Transducer
Optical
Piezoelectric
Electrochemical
Calorimetric

Electric
signal
Analyte
Antibody/Protein
Enzyme
Microorganism
Cell
Recognition elements
Transducer
Optical
Piezoelectric
Electrochemical

Calorimetric
Electric
signal
Analyte
Fig. 1. Schematic representation of a biosensor.
Detection of low molecular mass analytes Y. Uludag
˘
et al.
5472 FEBS Journal 274 (2007) 5471–5480 ª 2007 Akubio Ltd. Journal compilation ª 2007 FEBS
liquid layer. There are numerous examples in the liter-
ature on the applications of Eqns (2) and (3) and more
sophisticated algorithms that incorporate the measure-

ment of complex shear modulus in addition to mass,
viscosity and density, but a detailed analysis of this
approach lies beyond the scope of this review. How-
ever, it is possible to further optimize sensor sensitivity
by appropriate matching of the interfacial polymer
chemistry physical properties with the acoustic sensor
design, and a few seminal examples in this regard are
noted below.
It can been seen from the above that a quartz crystal
resonator is thus sensitive not only to mass, but also
to viscosity and density changes on the resonator sur-
face. Therefore the term quartz crystal microbalance is

not strictly accurate for all applications. The device is
also known as thickness-shear mode resonator or a
bulk acoustic wave sensor, because the bulk of the
crystal oscillates at a resonance frequency in a thick-
ness shear mode of vibration.
Surface acoustic wave sensors are also based on the
piezoelectric properties of quartz crystal. In this case
only a surface wave is generated by the electrodes and
the frequency of the surface waves is $ 100 MHz to
1GHz [14]. These frequencies are much higher than
thickness-shear mode resonators, and this is the reason
for the higher sensitivity of surface acoustic wave sen-

sors. However, higher sensitivity also means higher
response to viscosity changes and this problem causes
difficulties when surface acoustic wave sensors are used
in liquids [15]. By monitoring the change in resonant
frequency and motional resistance that occurs upon
adsorption of a ligand to the surface, quartz crystal
resonators can be used to characterize interactions
with peptides [16], proteins and immunoassay markers
[17], oligonucleotides [18], viruses [19], bacteria [20]
and cells [21]. The technology can thus be applied to
an extremely wide range of biological and chemical
entities with a molecular mass range from < 200 Da

through to an entire cell.
Application of acoustic sensors to small molecule
detection
The detection limit of many affinity biosensors is clo-
sely linked to the molecular mass of the analyte. Many
researchers prefer to immobilize the small molecule
analyte on the sensor surface and measure the binding
of a larger molecule [22] or to perform a competitive
displacement assay with a hapten–carrier conjugate
[23]. Another option is to conjugate the small molecule
to a bead or other additional mass load to increase the
molecular mass of the complex detected [24]. Other

approaches utilize a coupled assay format in which
direct binding then results in capture of an enzyme
that can convert a soluble substrate to a precipitate to
effect signal amplification. For example, organophos-
phorous and carbamate pesticides have been detected
using a two-enzyme system to produce peroxide, which
combined with peroxidase and benzidine formed an
insoluble product that absorbed to the sensor surface
[25]. Nonionic surfactants were reported to enhance
the surface deposition of suspended precipitate
enabling detection of carbaryl and dichlorvos pesti-
cides down to 1 p.p.m. This group also published an

assay for 4-aminophenol which involved precipitation
of indophenol in an amount proportional to the
4-aminophenol concentration in the sample [25].
In addition to the above strategies, the intrinsic sen-
sitivity of QCM to shear modulus, viscosity and den-
sity changes manifested at the surface interfacial layer
allows for the development of novel small-molecule
detection assays. In this case, the binding of a small
molecule that induces conformational changes in the
interfacial layer results in a modulated shear modulus
(related to rigidity and ⁄ or flexibility). Such effects can
significantly amplify the response due to mass binding

alone. For example, Carmon et al. [26] immobilized a
glucose ⁄ galactose receptor on a QCM sensor surface
and exposed the receptor to 180 Da sugars. A repro-
ducible frequency change was observed which was
ascribed to the conformational change of the receptor
upon ligand binding. Similarly conformational changes
have been invoked in the binding of ions and peptides
to calmodulin due to ion or peptide binding [27], and
the insertion of an Ad-2a model peptide onto glyco-
lipid monolayers [28]. In the latter case, the 2a-helix
structure of the peptide in the bulk solution is known
to convert to a b-structure upon association with a

lipid monolayer. This conformational change was man-
ifested as a frequency decrease for the piezoelectric
sensor.
This approach has been extended further in a
dynamic electropolymerization study in which imped-
ance data were acquired during polymerization at the
fundamental and third harmonic modes of a 10 MHz
thickness shear mode resonator [29]. At a critical
thickness, the system exhibited mechanical resonance,
a special condition in which the mechanical shear
deformation across the polymer film corresponded to
one quarter of the acoustic wavelength. At this point,

the resonant frequency and admittance data showed
dramatic changes with polymer coverage. Several
groups have also extended full impedance analysis
incorporating shear modulus modelling to protein films
[30] and phage binding [31].
Y. Uludag
˘
et al. Detection of low molecular mass analytes
FEBS Journal 274 (2007) 5471–5480 ª 2007 Akubio Ltd. Journal compilation ª 2007 FEBS 5473
For all the examples cited above, the binding
capacity of the sensor surface is critical to maximize
the sensitivity of the sensor for the detection of small

molecules. Here the synergies between robust, stable
MIPs with a mass and shear-modulus sensitive sensor
become apparent. The meso ⁄ microporous 3D matrix
structures formed by MIPs, not only increase the total
number of binding sites acoustically coupled to the
sensor, but also result in additional frequency changes
manifested in analyte- and dose-dependent modulation
of the surface-associated shear modulus of the poly-
mer layer in the polymeric structure. In other words,
binding of analyte to a MIP can result a larger
change in the total acoustic load on the sensor and
hence enable more robust detection of small mole-

cules. It is possible to model this phenomena as a
finite viscoelastic layer (of thickness h
f
, density r
f
and
shear modulus G ¼ G ¢ + jG¢¢ [32]. The latter value,
G, is the complex shear modulus where G¢ and G¢¢ are
the layer storage and layer loss moduli, respectively.
These layers are then exposed to a bulk liquid (of vis-
cosity g
L

and density q
L
). For these layered compo-
nents, it is possible to derive the surface mechanical
impedances:
Z
M
¼ jxq
s
ð4Þ
for an ideal ⁄ rigid mass layer, where r
s

is the mass per
area contributed by the interfacial layer;
Z
F
¼
ffiffiffiffiffiffiffiffi
q
f
G
q
tanhðch
f

Þð5Þ
for a viscoelastic film (MIP), where c is the shear wave
propagation constant (c ¼ j2pf
o
(r
f
⁄ G)
1 ⁄ 2
) and j ¼ Ö-1;
Z
L
¼

ffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi
2pf
0
q
L
g
L
2
r
ð1 þ jÞð6Þ
for a liquid layer (semi-infinite Newtonian liquid), and
finally;

Z ¼ j2pf
0
ffiffiffiffiffiffiffiffi
q
f
G
q
Z
L
cosh hðch
f
Þþ

ffiffiffiffiffiffiffiffi
q
f
G
p
sinhðch
f
Þ
ffiffiffiffiffiffiffiffi
q
f
G

p
cosh hðch
f
ÞþZ
L
sinhðch
f
Þ
ð7Þ
where Z is the impedance for the composite system.
Note that the impedance measured by the piezoelectric
sensor is not simply the sum of those for individual

layers, as for each layer there will be an acoustic phase
shift, which causes a transformation of the impedance
contributed by layers more distant from the resonator.
In addition, this model does not accommodate the typ-
ically inhomogeneous layers that exist in reality in
MIPs exposed to biological systems, or changes in dis-
tribution of mass induced by varying matrix conditions
and analyte binding.
When layers of a material of differing density and
viscosity to the liquid (such as a MIP) are deposited at
the surface–liquid interface, there is a contribution
from both the bulk liquid and the added material that

has displaced liquid and mechanically coupled to the
interface. In this case, the penetration depth can be
defined as:
d ¼ð1 À v
p
Þ
ffiffiffiffiffiffiffiffiffiffiffi
g
L
pf
0
qL

r
þ v
p
ffiffiffiffiffiffiffiffiffiffiffi
g
L
pf
0
q
f
s
ð8Þ

where g
p
and q
p
are the liquid viscosity and density,
respectively, and v
p
is the fraction of the volume
within the penetration depth occupied by protein. This
could be extended to encompass both a receptor layer
and analyte layer if necessary. Integrating the mole
fraction of MIP ⁄ water in combination with the defini-

tion of a composite impedance above, we can derive:
Z ¼ v
p
j2pf
0
ffiffiffiffiffiffiffiffi
q
f
G
q
ð1 À v
p

ÞZ
L
cosh hðch
f
Þþ
ffiffiffiffiffiffiffiffi
q
f
G
p
sinhðch
f

Þ
ffiffiffiffiffiffiffiffi
q
f
G
p
cosh hðch
f
Þþð1 À v
p
ÞZ
L

sinhðch
f
Þ
ð9Þ
Molecularly imprinted polymers
The history of MIPs can be traced to experiments per-
formed by Polyakov and his group in 1931. Silica gels
prepared by Polyakov’s group showed selective binding
towards one of the solvents used for the gel prepara-
tion. Later, studies by Wulf [33] and Haupt and Mos-
bach [34] helped to establish this technique in relation
to organic polymers. Initially, MIPs were used as sta-

tionary phases for chromatographic methods. Later,
the application of imprinted polymers was extended to
the biosensors area, where MIPs have been used as
recognition elements as an alternative to biological
materials such as antibodies and proteins. Similarly,
synthetic receptors formed by molecular imprinting
can be used to recognize biological or nonbiological
molecules on QCM sensors. Here the imprint of a tem-
plate molecule is formed on a synthetic polymer that
has cavities resembling the geometric shape of the tem-
plate and also has binding sites for template recogni-
tion [33]. MIPs as synthetic receptors have several

advantages over biological receptors [35]. The main
advantage of MIPs is their stability to harsh condi-
tions, in contrast to natural biomolecules that are sen-
sitive to environmental changes and can denature
easily. Because of the robust nature of MIPs, biosen-
sors that use MIP surfaces in general have a longer
shelf life than analogous biological sensors. MIPs are
simple to prepare, and their adaptation to a variety of
Detection of low molecular mass analytes Y. Uludag
˘
et al.
5474 FEBS Journal 274 (2007) 5471–5480 ª 2007 Akubio Ltd. Journal compilation ª 2007 FEBS

practical applications has been widely demonstrated.
In addition, molecular interaction studies with MIPs
can be performed in organic solvents as well as aque-
ous solvents.
Template molecules can be imprinted to a polymer
with covalent- [36], noncovalent- [37] or metal-ion-
mediated [38] interactions, followed by appropriate
cross-linking agents. Imprinting consists of three steps;
first, one or several functional monomers are mixed
with the template molecule in a solvent. Second, poly-
merization of monomers occurs in the presence of a
cross-linker. Third, the template is removed from the

polymer using basic, acidic or detergent solutions
(Fig. 3). The performance and selectivity of MIPs
towards a target molecule depend on many factors.
These include the molecular diversity of the monomer
units employed, the geometry of the imprinted cavity,
the rigidity of the cavity and associated implications
for enthalpy ⁄ entropy compensation. These factors all
govern the affinity and selectivity of the MIP recogni-
tion elements towards the analyte of interest. A discus-
sion of the importance of hydrogen bonds, van der
Waal’s forces, ionic interactions, ion dipole interac-
tions and hydrophobic interactions and other molecu-

lar phenomena between the template and the monomer
lies beyond the scope of this minireview and the reader
is referred to recent reviews and books in this area
[35,39].
To prepare MIPs with good affinity and selectivity,
polymerization conditions and the selection of mono-
mer and cross-linking agent are important parameters
for optimization (Table 1). There is no general proce-
dure for MIP preparation; therefore, depending on the
application, procedures need to be examined thor-
oughly before a decision is made. After the selection of
the polymerization procedure, the variables of the pro-

cess should be optimized. Template design, monomer
selection, solvent selection and polymerization condi-
tions all require attention. In general, the performance
of MIPs in aqueous solutions is poor, therefore, if
water-soluble templates need to be used the polymeri-
zation method needs to be carefully considered. High
nonspecific binding and heterogeneity of binding sites
needs to be addressed for successful application. If,
after polymerization, there are still embedded template
molecules remaining in the polymer, this will reduce
the capacity and invalidate analysis. Therefore extra
care needs to be taken to remove the template from

the polymer and the 3D structure of the polymer
should allow for easy regeneration of the template
from the polymer for repeated bindings. Reproducible
fabrication of MIPs is essential for gathering reliable
results from each assay.
MIP–QCM sensors
Every year many new studies are published involving
MIP–QCM sensors. In these applications, MIP synthe-
sis is performed either in situ on the sensor surface or
via preprepared MIP particles ⁄ beads that are immobi-
lized on the sensor surface using a PVC matrix. The
thickness of the imprinted polymers varies between

18 nm and 5 lm. To obtain a reproducible and reli-
able MIP–QCM sensor, it is essential to control the
thickness and properties of the polymer coating on the
sensor surface.
Monomers and target
Target & monomers complex
Polymerisation
-T arget
+ Target
Imprinted polyme
r
Monomers and target

Target & monomers complex
Polymerisation
-T arget
+ Target
Imprinted polyme
r
Fig. 3. Schematic representation of
molecular imprinting.
Table 1. Variables that need to be optimized for the preparation imprinted polymers.
Imprinting
mechanism Polymerization format Monomer selection Medium selection Polymerization conditions
Covalent Bulk polymerization Combinatorial screening Organic solvent Cross-linking agent selection

Noncovalent MIP beads Thermodynamic calculations Aqueous solvent Ratio of template ⁄ monomer ⁄ cross-linking
agent
Metal-mediated Films on bead ⁄ particles
or sensor surface
Computational methods Temperature
Pressure
Time
Y. Uludag
˘
et al. Detection of low molecular mass analytes
FEBS Journal 274 (2007) 5471–5480 ª 2007 Akubio Ltd. Journal compilation ª 2007 FEBS 5475
In situ polymerization

Two approaches to synthesize in situ imprinted poly-
mers for QCM sensing have been reported. In one
approach, the gold surface of the QCM was treated
with a thiolated molecule to create active groups on
the sensor surface and improve the adhesion of the
imprinted polymer on the gold electrode (Table 2).
Allyl mercaptan [40,41] (N-Acr-l-Cys-NHBn)
2
[42],
mercaptoethanol [43], thioctic acid-modified glycidyl
methacrylate [44], and mercaptoundecanoic acid [45]
have been used to activate the gold electrode. In the

other approach, polymer synthesis was performed
directly on to the gold surface without any activation.
MIPs could be deposited by surface grafting, spin
coating, sandwich casting, or electro-polymerization
methods.
Surface grafting method
Although it is difficult to control MIP film thickness
during polymer synthesis, in a sensing device it is
essential to reduce batch-to-batch variation. For this
purpose, Piacham et al. investigated a possible route to
prepare ultra-thin MIP films (< 50 nm) specific for
(S)-propranolol [45]. The imprinting process was per-

formed directly on to the quartz surface after coating
the gold-coated crystal surface with mercaptoundeca-
noic acid. The carboxyl groups of mercaptoundecanoic
acid were then activated with initiators, 2-ethyl-5-phen-
ylisoxazolium-3-sulfonate and 2,2-azobis(2-amidino-
propane) hydrochloride. The sensor was dipped into a
solution containing template, monomer (methacrylic
acid) and cross-linker (trimethylolpropane trimethacry-
late). UV irradiation resulted in a surface-grafted poly-
mer film on the quartz resonator. Although the
detection limit of this MIP–QCM sensor was too high
for practical application, Piacham et al. succeeded in

producing sensors that generated $ 30 Hz response on
the injection of 19 mm (S)-propranolol.
Sandwich casting method
Alternatively, a sandwich casting method can be used
either after surface activation of the QCM sensor sur-
face [40,46,47], or directly on to the sensor surface
[17,40,48,49]. In this method, a polymerization mixture
is dropped on to the quartz crystal and a microcover
glass is pressed on to the sensor while UV irradiation
is applied. The aim is to distribute the polymer
Table 2. Some examples of MIP-QCM studies. 2,2¢-azobis, (2 amidinopropane) hydrochloride; ABAH, 2-ethyl-5-phenylisoxazolium-3-sulfonate;
AIBN, azobis-(isobutyronitrile); AMVN, 2,2¢-azobis (dimethylvaleronitrile); EGDMA, ethylene glycol dimethacrylate; GMA, glycidyl methacry-

late; HEMA, 2-hydroxyethyl methacrylate; MAH, methacrylamidohistidine; MUDA, mercaptoundecanoic acid; NBAA, N-benzylacrylamide;
TRIM, trimethylolpropane trimethacrylate; 4-Vpy, 4-vinylpyridine.
Template Surface activation
Polymerization
method Initiator Cross-linker Monomer(s) Ref
Glucose Methacryloyl Surface grafted AIBN EGDMA MAH-Cu(II) ⁄ –
D-glucose [56]
Nanopeptide (N-Acr-
L-Cys-NHBn)
2
Surface grafted – – NBAA, acrylic acid,
acrylamide

[42]
(S)-Propranolol Mercaptoundecanoic
acid
Surface grafted ABAH TRIM Methacrylic acid [45]
Bilirubin Allyl mercaptan Surface grafted Benzophenone Divinylbenzene 4-Vpy [41]
Indole-3-acetic acid Allyl mercaptan and
1-butanethiol
Sandwich casting AIBN EGDMA N,N-dimethylaminoethyl
methacrylate
[40]
Sialic acid Allyl mercaptan Sandwich casting – EGDMA 4-Vpy and AMVN [40]
Dansylphenylalanine Thioctic acid-modified

GMA and thioctic
acid dodecane ester
Sandwich casting AIBN EGDMA Methacrylic acid, 4-Vpy [46]
L-Tryptophan Thioctic acid-modified
GMA
Sandwich casting – TRIM Acrylamide [44]
Sialic acid – Sandwich casting – EGDMA N,N,N-trimethylaminoethyl
methacrylate, HEMA and
AMVN
[40]
L-serine – Sandwich casting AIBN EGDMA Methacrylic acid [17]
L-menthol – Sandwich casting AIBN EGDMA Methacrylic acid [48]

Sorbitol – Electro-polymerization AIBN – m-Aminophenol [51]
Tegafur – Electro-polymerization – – m-Aminophenol [52]
Benzene, toluene
and xylene
– Spin coating AIBN Divinylbenzene Styrene [62]
Detection of low molecular mass analytes Y. Uludag
˘
et al.
5476 FEBS Journal 274 (2007) 5471–5480 ª 2007 Akubio Ltd. Journal compilation ª 2007 FEBS
solution evenly on the quartz and thus obtain a uni-
form polymer layer.
Kugimiya and Takeuchi synthesized a MIP on a

quartz crystal sensor surface using a plant hormone,
indole-3-acetic acid, as a template [40]. Platinum
coated 9 MHz quartz crystals were treated with allyl
mercaptan and 1-butanethiol to introduce vinyl groups
to the sensor surface. The monomers and the template
were dissolved in chloroform and dropped onto the
QCM sensor and held with a glass microcover. Poly-
merization was initiated by UV irradiation. Assays for
indole-3-acetic acid binding to the MIP-coated surface
were performed using three crystals and a linear rela-
tionship was obtained from 10 to 200 nmol indole-
3-acetic acid. The coefficient of variation between the

three sensors was 9.0%.
Spin coating method
The spin coating approach was used by Ling et al. [43]
to prepare MIPs for catecholamines. Ling et al. first
activated the gold-coated crystal with mercaptoethanol,
and then the homogeneous phase MIP solution was spin
coated. Using this method, dopamine-, epinephrine-
and norepinephrine-imprinted resonators were prepared
and binding assays were performed. The results show
that dopamine-specific MIP-coated resonators have bet-
ter selectivity than the other MIPs prepared (relative
binding selectivity of dopamine–MIP is 1 for dopamine

and, 0.03 for norepinephrine and 0.02 for epinephrine).
Electro-polymerization method
An electro-polymerization method was applied to pre-
pare imprinted polymers on quartz crystal sensor sur-
faces to detect sorbitol, poly(o-phenylenediamine),
tegafur and nucleotides [50–52]. Feng et al. synthesized
an o-phenylenediamine film for sorbitol on a QCM
crystal by cyclic voltammetry [51]. After MIP deposi-
tion the binding assays with sorbitol, glucose, fructose,
mannitol and glycerol were performed using a QCM
device. Glycerol could bind to the sorbitol-imprinted
surface, however, the binding of other compounds was

very limited. The detection limit of sorbitol binding
was found to be 1 mm.
Polymerization prior to sensor coating
MIPs have be prepared using a bulk polymerization
method; after grinding and sieving, the resulting parti-
cles are mixed with PVC and coated on the sensor
surface with spin coating. Imprinted polymers for
microcystin-LR, nandrolone, phenacetin, nicotine and
paracetamol were prepared using this method on
QCM sensor surfaces [17,53–55].
Application areas of MIP–QCM biosensors
The three most common application areas for MIP–

QCM sensors are clinical diagnosis, environmental
monitoring and control of enantiomeric separation.
Most studies describe detection in buffer or organic
solvents indicating the early stage in development of
these devices with respect to real applications
[17,40,43,51,52,56,57]. Although it is possible to detect
small molecules in buffered or organic solutions, it is
also important to determine the amount of a particular
drug or any other chemical in body fluids. For
instance, Tan et al. determined the amount of nicotine
and paracetamol in human serum and urine [55,58].
The detection limit of nicotine was found to be 25 nm

using an imprinted polymer-coated sensor. Wu and
colleagues fabricated bilirubin specific imprinted poly-
mers on a QCM sensor using a photo-graft poly-
merization method [41]. Bilirubin concentration is
considered an important index to identify liver dis-
eases. A bilirubin-specific sensor was challenged with
both bilirubin and its analogue biliverdin, cross-reac-
tivity of bilirubin versus biliverdin binding was 31 : 20.
Yan et al. [59] developed a MIP–QCM sensor for
daminozide (a potentially carcinogenic chemical the
detection of which is important in food safety) with a
detection limit of 50 pgÆmL

)1
. MIP–QCM sensors for
acetaldehyde, monoterpenes and bisphenol A have
been prepared for environmental pollution detection
by different groups [48,60,61]. Synthesis of enantiomer-
ically pure organic compounds is very important for
industrial production. MIP–QCM sensors capable of
enantiomeric discrimination are very useful tools for
process end-point analysis and various groups have
discriminated between R- and S-propranolol, l- and
d-tryptophan, l- and d-serine and l- and d-dansylphe-
nylalanine enantiomers [17,44,49].

Summary
The inherent robustness, ease of manufacture and
high capacity of MIPs make them a potentially useful
alternative for small molecule detection using piezo-
electric biosensors. Although the majority of applica-
tions involve the use of buffered pure solutions rather
than real clinical or environmental samples for detec-
tion, this perhaps simply reflects the early stage of
development of the technology. Selectivity is still a
significant issue for imprinted polymers and this can
hamper specific, sensitive detection of analytes in
Y. Uludag

˘
et al. Detection of low molecular mass analytes
FEBS Journal 274 (2007) 5471–5480 ª 2007 Akubio Ltd. Journal compilation ª 2007 FEBS 5477
complex fluids. Polymerization methods are critical
determinants of selectivity and overall assay perfor-
mance of MIPs and as there is no general procedure
for MIP preparation, each template requires optimiza-
tion of several parameters to fabricate reproducible,
high-performance sensor surfaces. It is expected that
the improvements to polymerization techniques should
greatly enhance the selectivity and binding capacity of
the MIP–QCM sensors.

In many ways, the stage of development of MIP
interfaces is reflected in the state of development of
robust piezoelectric biosensors compared with analo-
gous robust electrochemical and optical biosensors
that have benefited from more than two decades and
several billion dollars of research and development.
Many researchers rely on piezoelectric devices built
in-house that are generally very sensitive to artefacts
including temperature drift, humidity and pressure
effects, resulting in less reliable and less reproducible
measurements. Commercially available systems that
minimize these effects with physical compensation on

in-line referencing are appearing, but they are still far
from the idealized portable device for near patient
testing, point of care or remote environmental moni-
toring in a handheld device. However, the fundamen-
tal elements of acoustic sensors, i.e. cost of goods,
simplicity and ease of manufacture, clearly have the
potential to be exploited in an integrated portable
device for use in quantitative detection of low molec-
ular mass analytes. It is expected that recent advances
in understanding acoustic biosensing technology will
generate such devices that can be combined with the
benefits of MIPs as the recognition element to pro-

vide a new generation of sensors for use in routine
measurements outside the controlled laboratory envi-
ronment.
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