Hydrogels in Sensing Applications

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Progress in Polymer Science 37 (2012) 1678– 1719

Contents lists available at SciVerse ScienceDirect

Progress in Polymer Science

j ourna l ho me p ag e: www.elsev ier .com/ locate /ppolysc i

ydrogels in sensing applications

aniel Buengera,1, Fuat Topuza,1, Juergen Grollb,∗

DWI e.V. and Institute of Technical and Macromolecular Chemistry, RWTH Aachen University, Forckenbeckstr. 50, 52074 Aachen, GermanyDepartment and Chair of Functional Materials in Medicine and Dentistry, University of Würzburg, Pleicherwall 2, 97070 Würzburg, Germany

r t i c l e i n f o

rticle history:eceived 10 January 2012eceived in revised form 29 August 2012ccepted 7 September 2012vailable online 14 September 2012

eywords:ydrogelsensorsiosensorstimuli responsive

a b s t r a c t

Hydrogels are hydrophilic, highly water swellable polymer networks capable of convert-ing chemical energy into mechanical energy and vice versa. They can be tailored regardingtheir chemical nature and physical structure, sensitiveness to external stimuli and biocom-patibility; they can be formed in various structures and integrated into (micro-)systems.Accordingly, over the last decade, these materials have gained considerable recognition asvaluable tool for sensors and in diagnostics. This article reviews the use of hydrogels insensor development with focus on recent efforts in the application of stimuli responsivehydrogels as sensors, hydrogels as suitable matrices in which the sensing (bio-)moleculesare embedded and hydrogels for modification and protection of sensor surfaces. In the firstpart of the review, both sensors and hydrogels are defined and a basic theoretical overview

hase transitions of hydrogels and their behavior is given. Subsequent chapters focus on hydrogels in physico-chemical and biochemical sensing mechanisms with a primary emphasis on the hydrogelsas such and the applied sensing mechanism. Finally, the review is concluded by a sum-mary and discussion including an outlook on future perspectives for hydrogels in sensingapplications.

© 2012 Elsevier Ltd. All rights reserved.

ontents

1. Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 16791.1. Sensors: definition and classification . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 16791.2. Biosensors . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 16801.3. Hydrogels: general overview . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1681

1.3.1. Theory of swelling . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 16821.3.2. Mechanisms of response for stimuli sensitive hydrogels . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1683

1.4. Hydrogels for sensors . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 16841.4.1. Hydrogels as sensors . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 16841.4.2. Hydrogels for biosensors . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1685

1.5. Scope and structure of the review . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1685
2. Physicochemical sensing mechanisms . . . . . . . . . . . . . . . . . . . . . . . . . . . . .
2.1. Temperature . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

2.2. Ions, ionic strength and pH . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

∗ Corresponding author. Tel.: +49 (0) 931 201 73510;ax: +49 (0) 931 201 73500.

E-mail address: [email protected] (J. Groll).1 These authors contributed equally.

079-6700/$ – see front matter © 2012 Elsevier Ltd. All rights reserved.ttp://dx.doi.org/10.1016/j.progpolymsci.2012.09.001

. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1686. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1686. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1688

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D. Buenger et al. / Progress in Polymer Science 37 (2012) 1678– 1719 1679

2.3. Gas and humidity . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 16912.3.1. Gas . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 16912.3.2. Humidity . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1692

3. Biochemical sensing mechanisms . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 16943.1. Molecular interactions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1694

3.1.1. Enzyme–substrate interaction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 16943.1.2. Antibody–antigen . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 16983.1.3. Nucleotide, oligonucleotide and DNA . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 17013.1.4. Other binding mechanisms . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1703

3.2. Living sensors . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 17073.2.1. Bacteria . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 17083.2.2. Cells . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1708

4. Summary and perspective . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1712Acknowledgments . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1712

. . . . . . . .
References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .
1. Introduction

1.1. Sensors: definition and classification

In general, the term sensor is derived from the Latinword “sensus”, which directly translated means “sense”, orto “sense something”. If one thinks about a “sense” the firstidea that comes up are the traditional five human senses:sight, hearing, taste, smell and touch. The senses give thehuman body the ability to receive signals or stimuli fromthe environment and react or respond to them. If a sensegives the ability to receive and respond to a signal, we cantransfer this to the technical level and define a sensor asfollowed:

“A sensor is a device that receives and responds to sig-nals and stimuli from the environment”.

Furthermore, the human senses give a first hint howsensors can be classified. Sight, hearing and touch have incommon that they sense physical stimuli, namely light, inform of electro-magnetically waves, acoustic waves andpressure. Smell and taste respond to chemical stimuli,like odor and taste molecules. Accordingly, sensors canbe divided into two big classes: physical and chemicalsensors. The IUPAC defined in 1991 the terms physi-cal and chemical sensors with the following definitions[1]:

“A physical sensor is a device that provides informa-tion about a physical property of the system” and “Achemical sensor is a device that transforms chemicalinformation, ranging from the concentration of a spe-cific sample component to total composition analysis,into an analytically useful signal”.

Now that the term sensor is defined and a first classi-fication, corresponding to the sensed stimulus, is made acloser look at the general structure of sensors can be taken.Again the human senses lead the way to this structure:whenever receptor cells in the human body receive certainstimuli, e.g. an electro-magnetically wave hits a photore-
ceptor cell in the retina, a chemical reaction is triggered,and then transduced, in form of an electrical signal, to thebrain. Deductive sensors consist of two basic parts, namelya receptor and a transducer part. Technically spoken the
. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1712

receptor-part of a sensor senses a chemical or physicalstimulus and transforms it into a form of energy. The trans-ducer part is than capable of transducing this energy intoa useful analytical signal which can be processed and dis-played. Fig. 1 illustrates the general sensor structure.

On the technical level there are many ways to clas-sify sensors e.g. according to the measuring effect, givenanalyte, receptor principle, field of application, mode ofapplication and so on. The most established way is the onefollowing the principles of signal transduction, committedby the IUPAC 1991 for chemical sensors [1].

1. Optical sensors, based on absorbance, reflectance, lumi-nescence, fluorescence, index of refraction, optothermaleffect, and light scattering effects.

2. Electrochemical sensors, based on voltammetric, incl.amperometric, and potentiometric effects, chemicallysensitized field effect transistors (CHEMFET) and poten-tiometric solid electrolyte gas sensors.

3. Electrical sensors, based on metal oxide semiconductorsand organic semiconductors.

4. Mass-sensitive sensors, based on piezoelectric and sur-face acoustic wave effects.

5. Magnetic sensors, based on paramagnetic gas properties.6. Thermometric sensors, based on heat effects of specific

chemical reaction or adsorption.7. Radiation sensitive sensors, based on absorbance, or

emission of X-, �-, or �-radiation that is used for thedetermination of a chemical composition.

As the classification shows the IUPAC draws no clearline between chemical and physical sensors and countsphysical sensors that measure e.g. a chemical state, reac-tion, or composition as chemical sensors. This shows thata general approach to classify is difficult as chemical andphysical sensing almost always go hand in hand. Hence,more generally, sensors can be defined as devices thatdetermine selectively and reversibly the concentration
of analyte molecules by amplifying the signal resultingfrom interaction between analyte and receptor withoutneed of any other instrument and additional reagents[2].

1680 D. Buenger et al. / Progress in Polymer Science 37 (2012) 1678– 1719

F ith thea

1

tsimccabWbrw1

a

1

2

ig. 1. General assembly of a sensor comprising a receptor that interacts wnd display are important parts of complete sensor systems.

.2. Biosensors

If a physical sensor is a device that provides informa-ion about a physical property of a system and a chemicalensor is a device that transforms chemical informationnto an analytically useful signal, a biosensor consequently

ust work in a similar way. Actually the class of biosensorsan be counted as a cross of both, physical and chemi-al sensors. They can be devices that provide informationbout a bio-physical property of a system and/or transformio-chemical information into an analytically useful signal.hat is more important is the fact that biosensors have a

iological recognition element that enables the analysis ofelevant biological information. This concept correspondsith the definition for biosensors given by the IUPAC in

999 [3]:

“A biosensor is an integrated receptor-transducerdevice, which is capable of providing selective quanti-tative or semi-quantitative analytical information usinga biological recognition element.”

A classification of biosensors can be done for examplefter one of the following principles:

. Following the principle of signal transduction, e.g. elec-
trochemical, optical, mass sensitive and so on.
. Depending on the biological recognition element, e.g.biological ionophores, enzymes, antigens/antibodies,whole cells, membrane receptors, plant and animal

Fig. 2. Types and components of biosensors as well as their fi

analyte and a transducer. Beyond that core sensing unit, data processing

tissue, protein receptors and channels, oligonucleotides,specific ligands etc.

3. Classified after the sensed analyte, e.g. glucose, DNA,enzymes, toxins, certain drugs and so on.

From the first biosensor invention filed by Leland C.Clark in 1962 that concerned a potentiometric enzymeelectrode developed to measure the concentration ofglucose [4], numerous biosensors with different specifici-ties have been developed. During the last decade, muchwork was devoted to downsize the laboratory workspacetowards small and versatile biosensor devices. However,indispensable requirements including fast responsibility,reproducibility, reliability and stability remained the same.Crucial parameters in designing biosensors with high per-formance are (i) immobilization of biomolecules in theirnative conformation, (ii) accessibility of receptors to thetargets, and (iii) specific adsorption to the support [5–7].Especial regarding these demands, hydrogels offer specificadvantages for biosensors.

Biosensors are becoming increasingly important andpractical tools to suit a vast pool of application areasincluding point-of-care testing, home diagnostics, andenvironmental monitoring (Fig. 2) [8]. As already men-tioned, a biosensor consists of a bioelement for recognizing
a specific analyte, and a sensor element that transduces thespecific recognition into an electrical signal.
Enzymes, antibodies, living cells, or tissues mayserve as bioelements and ideally should recognize one

elds of application. Figure is partly redrawn from [8].


ensor co
Fig. 3. Overview on the four types of bioelement–s
analyte specifically while being insensitive to otherpotentially interfering molecules. These bioelements canbe coupled with sensing elements through membraneentrapment, physical adsorption, matrix entrapment, orcovalent incorporation (Fig. 3). In membrane entrapment, asemipermeable membrane is applied between the analyteand the bioelement, with the sensor attached to the bioele-ment [8]. For physical adsorption van der Waals forces,hydrophobic forces, hydrogen bonds, and/or ionic forcesmay be used to attach the bioelement to the surface ofthe sensor. In case of covalent binding, the bioelement isdirectly linked to the sensor surface by a covalent bondthrough a chemical reaction. Finally, porous entrapment isbased on forming a porous encapsulation matrix aroundthe bioelement that ensures its presence close to the sen-sor.

Biosensors can produce signals tuned by the concen-tration of analytes of interest either directly or indirectly.In a direct case, a signal that comes from the production ofheat, light or chemical products is sensed with signal inten-sity proportional to the analyte concentration. For indirectmeasurements, the analyte inhibits the production of a sec-ondary chemical product (e.g. by blocking active enzymesites) which concentration is sensed. For both ways, sen-sitivity of the sensor can be given by the change in itsresponse as a function of the corresponding change inquantity being monitored. Biosensors become convenientto use if they exhibit a linear relationship between the vari-ation in the amplitude of the output, �a, and the input,�m:

�a = s · �m (1)

where s corresponds to the sensitivity of the biosensor anddetermines the suitability of the sensor for use in a particu-lar application. It is possible that m is not an actual quantityto be measured, but a function of that quantity. This is thecase for potentiometric biosensors in which the amplitude
of the signal is proportional to the logarithm of the analyteconcentration, c, according to Nernst Law:
�a = s · �(log c) (2)

upling methods. Figure is partly redrawn from [8].

When a variation in the bioelement recognition stopsto yield an appreciable signal variation, then the detectionlimit has been reached. The detection limit usually corre-sponds to the limit of linear dependence range betweenbioelement recognition and signal intensity at low con-centrations. The linear region of a biosensor is derivedfrom a calibration curve of its response to different ana-lyte concentrations. The curve is only meaningful if thecalibration is made for increasing and decreasing seriesof analyte concentration, and it is important that thebiosensor shows no memory effect and hysteresis. A goodcalibration curve is required to indicate the stability ofthe response of the biosensor, which should neither driftnor oscillate with time. In some cases, corrective devicesare available that can go beyond the linear zone andmake it possible to exploit the non-linear responses atlow concentrations, thus lowering the detection limits ofbiosensors.

An important parameter for biosensors especiallyregarding their practicability and applicability is theresponse time of the setup. This is the time the systemneeds to reach a steady state from the instant of the varia-tion in analyte concentration. Theoretically, the responsetime is infinitely long; however, a reasonable value canbe found by fixing an interval between the instantaneousvalue of the response and the theoretical final value, as afunction of the required precision. A steady state is thenconsidered as obtained when the predetermined inter-val has been covered. The response time of the biosensorindicates how quickly it responds to a variation in concen-tration and it also defines a minimum delay, after whichdata can be collected with a given precision. It is generallyquicker to obtain the same information from the derivativeof the response curve.

1.3. Hydrogels: general overview

Hydrogels are three-dimensionally cross-linked
hydrophilic polymer networks. They swell without disso-lution up to 99% (w/w) water of their dry weights [9–11].Depending on the type of cross-linking, hydrogels canbe divided into two classes: (i) chemically cross-linked


average

hCdbbihboacIfpsstotb[ttnst

t(holdita

Fig. 4. A cross-linked hydrogel structure with the mesh size � and the

ydrogels, and (ii) physically cross-linked hydrogels.hemical hydrogels are formed by covalent networks ando not dissolve in water without breakage of covalentonds [12–15]. Physical hydrogels are, however, formedy dynamic cross-linking of synthetic or natural build-

ng blocks based on non-covalent interactions such asydrophobic and electrostatic interaction or hydrogen-ridges [16–18]. The hydrophilic and mostly inert naturef hydrogels often leads to minimized non-specific inter-ction with proteins and cells which makes hydrogels idealandidates for numerous bio-related applications [19,20].f the polymer network of a hydrogel is endowed withunctional groups, hydrogels become responsive to somehysical, chemical or biochemical stimuli [21–31]. Expo-ure to the respective stimuli causes reversible changes inwelling of the network. This response is determined byhe hydrogel composition, cross-linking type and degreef cross-linking which is the density of junctions joininghe chains into a permanent form. All these properties cane tailored to create hydrogels with adjusted properties32]. For example, a higher cross-linking degree leadso increase of the mechanical strength and decrease ofhe diffusion rate, thereby reducing the mesh size of theetwork [33]. Mesh size (�) is used to define the diffusionalpace available for the transfer of molecules or particleshrough the matrix of a hydrogel (Fig. 4) [34].

When hydrogels are used in sensing, it is usually one ofhe three following main characteristics that is exploited:i) their semiwet and inert structure that predeterminesydrogels as host-network, (ii) the amplification effectf their sensitiveness on a molecular level that is trans-ated into macroscopic effects such as a change in swelling
egree, and (iii) the ability to control the diffusion behav-
or of molecules through the polymer matrix [35–37]. Inhe following paragraph, the theory of hydrogel swellingnd the mechanism of response will be outlined in more

molecular weight between the cross-linking points MC , respectively.

detail as basis for the sensing applications described in thefollowing chapters.

1.3.1. Theory of swellingHydrophilicity of the polymers that constitute the net-

work creates an osmotic pressure within hydrogels thatleads to swelling of the matrix upon exposure to water[38]. The swelling process occurs in three steps: (i) diffu-sion of water molecules through the matrix, (ii) relaxationof polymer chains via hydration, and (iii) expansion of poly-mer network upon relaxation [39,40]. Hydrogels absorbwater to a maximum degree possible, the so-called equilib-rium water content. This degree is defined as the balancebetween osmotic pressure and the elastic retractive forcesof the polymer chains in the three-dimensional network.The stretching of polymer chains increases elastic retrac-tive forces as a counteraction for the network expansion.When these forces are balanced, the network expansionstops and comes to equilibrium. When either the osmoticpressure changes, e.g. due to deprotonation of carboxylicacid groups in the network due to a shift in pH, or thecross-linking density changes, e.g. due to degradation ofthe network and thus decreasing number of cross-linkingpoints, this balance is broken and a result of that, the hydro-gel exhibits a change in the degree of swelling.

The most widely used theory to explain the swellingin neutral hydrogels is the equilibrium swelling theory ofFlory and Rehner [41,42]. This assumes a Gaussian distri-bution of polymer chains and the average cross-links astetrafunctional branching units. Swelling of the polymernetwork is a function of elastic forces of the polymer chains,and the thermodynamical compatibility between polymer
and water molecules. Accordingly, the free energy of a neu-tral hydrogel (�G) can be expressed as:
�G = �Gel + �Gmix (3)

olymer
D. Buenger et al. / Progress in P
where �Gel represents the contribution of elastic refractiveforces, and �Gmix is the thermodynamic compatibility ofpolymer and solvent.

This equation can be rewritten in terms of chemicalpotentials. At the equilibrium conditions, the total chemicalpotential (�) has to be equal to zero:

� = �el + �mix = 0 (4)

When the total chemical potential is equal to zero, thechange in chemical potential due to mixing is balanced byelastic retractive forces. The change in chemical potentialdue to elastic forces can be expressed by using the theoryof rubber elasticity [43], while the contribution of mixingto chemical potential change is determined using the heatand the entropy of mixing [11]. With equaling these twocontributions, the molecular weight between two cross-links (MC ) in the absence of a solvent can be expressed as:

1

MC

= 2

MN

−(v̄/V1)[ln(1 − v2,s) + v2,s + �1v2

2,s]

v1/32,s −

(v2,s/2

) (5)

where MN is the average molecular weight of the polymerchains prepared in the absence of a cross-linker, v̄ and V1 arethe specific and the molar volume of water, respectively.v2,s is the polymer volume fraction in the fully swollenstate, and �1 is the polymer–solvent interaction parameter.

The presence of water changes the chemical potential,so that a new term is required for the volume fractiondensity of the polymer chains. Thus, the above originalFlory–Rehner model can be rewritten by incorporating v2,r ,a term describing the polymer fraction in the relaxed stateaccording to the theory of Peppas and Merrill [44].

1

MC

= 2

MN

−(v̄/V1)[ln(1 − v2,s) + v2,s + �1v2

2,s]

v2,r[(v2,s/v2,r)1/3 − (v2,s/2v2,r)](6)

When ionic groups are present in the network, theswelling equilibrium of the matrix becomes more compli-cated. New terms concerning the ionic strength, I and thedissociation constants, Ka, and Kb have to be added to Eq.(6). Equivalent expressions for anionic and cationic gels areformulated as shown in Eqs. (7) and (8), respectively.

V1

4I

(V2

2,s

v̄

)(Ka

10−pH − Ka

)2

= [ln(1 − v2,s) + v2,s + �1v22,s]

+(

V1

v̄Mc

)(1 − 2Mc

MN

)v2,r

[(v2,s

v2,r

)1/3

−(

v2,s

2v2,r

)] (7)

V1

4I

(V2

2,s

v̄

)(Kb

10pH−14 − Ka

)2

= [ln(1 − v2,s) + v2,s + �1v22,s]

+(

V1

vMc

)(1 − 2Mc

MN

)v2,r

[(v2,s

v2,r

)1/3

−(

v2,s

2v2,r

)] (8)

Hydrogel mesh size � (i.e. correlation length), based onvalues for the cross-linking density or molecular weightbetween cross-links can be described by the equation:

� = �−1/3(r̄2)1/2 = ˛(r̄2)

1/2(9)
2,s o o
where ̨ is the elongation of the polymer chains in any

direction, and (r̄2o )

1/2, the unperturbed end-to-end distance

of the polymer chains between the cross-linking points.

Science 37 (2012) 1678– 1719 1683

Assuming an isotropic swelling, the end-to-end distanceand the pore size of a swollen polymeric network can becalculated using the equation:

� = �−1/32,s

(2CNMC

Mr

)1/2

l (10)

where l is the length of the bond along the backbone chain,and CN, Mr are the Flory characteristic ratio and the molec-ular weight of the repeating units, respectively.

Swelling is affected by numerous physicochemical con-ditions or structural factors [45–49]. The nature of thepolymer and the cross-linking degree are crucial ones thatdetermine the swelling ratio. For instance, highly cross-linked hydrogels have a smaller mesh size and swell lessin comparison to loosely cross-linked ones. The swellingkinetics depends on the mechanism of solvent penetra-tion into the matrix, and is either diffusion-controlled (CaseI) or relaxation-controlled (Case II) [40,50,51]. In Case I,also known as Fickian type behavior, diffusion of watermolecules through hydrogel matrix occurs much fasterthan the relaxation of the polymer chains, and the swellingis controlled by a concentration gradient. However, inCase II (also called relaxation-controlled), the swelling iscontrolled by the rate of polymer relaxation. A detailedmathematical treatment of kinetics is beyond the scopeof this article and can be found in a paper by Peppas andColombo [52].

It becomes clear that analyte-induced changes in thecharacteristics of the network such as the cross-linkingdensity, content of ionic species or chemical nature of thepolymer backbone will result in macroscopic effects suchas altered swelling behavior. Such changes can result fromphysicochemical stimuli such as changes in temperature,pH or ion strength, or they may originate from biochemicalrecognition events as described below.

1.3.2. Mechanisms of response for stimuli sensitivehydrogels

Stimuli responsive hydrogels undergo a volume-phasetransitions upon exposure to the respective stimulus dueto molecular interactions resulting in abrupt changes inthe network such as swelling, collapse or solution-to-geltransitions. Thus, the stimulus as such plays key role inthe mechanism of hydrogel response. In this section, wedescribe hydrogel systems coupled with different responsemechanisms that define their behavior and application.

Hydrogels that are responsive to changes in pH are themost studied stimuli responsive hydrogels. Either acidic orbasic pendant groups on the polymer network lead to pH-induced volume-phase transitions. In the case of anionicnetworks, when the environmental pH is above the acidgroup’s characteristic pKa, ionization occurs, leading toincreases of (i) the number of fixed charges, (ii) hydrophilic-ity of the network, and (iii) electrostatic repulsion betweenthe chains [53]. Hence, the mesh size becomes sensi-tive to pH variations. However, the protonation of the
ionized acidic groups upon lowering pH of an aqueoussolution causes a decrease in the content of mobile counter-ions inside the matrix and in the strength of electrostaticrepulsions of the chain segments. So, the network gains

1 olymer

hapslwSwa

astttaaTU(isbtoitanbawmbotrbanc

tbabGnrbppabansocro

684 D. Buenger et al. / Progress in P

ydrophobic character, and the swollen form shrinks into compact state. Cationic networks (e.g. having aminoendant groups) exhibit a similar behavior but with oppo-ite trend. When pKb of the cationic pendant groups isower than the environmental pH, hydrophilicity of the net-

ork increases, and the hydrogel starts to swell [54,55].o, the hydrogels display pH-dependent swelling behaviorith opposite swelling/deswelling response to pH changes

s stated for anionic networks.Some hydrogels are responsive to temperature vari-

tions and display reversible temperature-dependentwelling behavior. These hydrogels are subdivided intohree main groups: (i) positive, (ii) negative, and (iii)hermally reversible gels [56,57]. Positive thermosensi-ive hydrogels (such as poly(acrylic acid), polyacrylamide,nd poly(acrylamide-co-butyl methacrylate)) have a char-cteristic upper critical solution temperature (UCST).hese gels shrink if the temperature sinks below theCST [58]. Negative temperature-sensitive hydrogels

like N-methylacrylamide, N,N-dimethylacrylamide or N-sopropylacrylamide) are characterized by a lower criticalolution temperature (LCST). These gels are highly swollenelow the LCST while the network collapses when theemperature increases above the LCST [59]. Several the-ries have been proposed to explain this LCST behaviorn temperature-sensitive polymers [60–64]. According tohese theories, hydrogen bonding and hydrophobic inter-ctions between non-polar components in the polymeretwork are crucial factors. Below the LCST, the hydropho-ic parts are surrounded by water molecules that form

cage around the group with only weak interactionith the non-polar center. With increasing temperature,olecules become more mobile, and this cage of immo-

ile water molecules is partially lost and the protectionf the hydrophobic groups gets weaker. At the LCST,he entropy loss of water molecules that are completelyeleased from the network is small enough to be counteralanced by the energy win of the hydrophobic inter-ction between clustering non-polar components of theetwork, so that the gel releases water and the networkollapses.

Certain hydrogels can undergo volume-phase transi-ions in the presence of biomolecules. These so-callediomolecule-sensitive gels have attracted considerablettention as intelligent materials, since they can senseiochemical changes through structural changes [65–68].lucose- and protein-sensitive hydrogels are most promi-ent examples in this family. For instance, glucoseesponsivity can be achieved by incorporating boronicased-ligands as recognition agents in the network. Com-lexation of glucose with these ligands induces a release ofrotons that can be sensed by conductance measurementsnd used for the determination of glucose content [69]. Alsooronic acid functional thermo-responsive hydrogels arelso used to sense glucose. When glucose binds to theseetworks, the hydrophilic/hydrophobic balance is changedo that the LCST of these networks shifts as a function
f glucose content what can be exploited for sensing theoncentration of glucose [70]. Alternatively, biomoleculeesponsive hydrogels can be prepared by immobilizationf specific enzymes in the hydrogel matrix. For instance,
Science 37 (2012) 1678– 1719

glucose oxidase (GOX) functional hydrogels are used tosense glucose through the glucose oxidase mediated enzy-matic oxidation from glucose to gluconic acid [71]. Antigenresponsive hydrogels are prepared either by chemicallygrafting antibodies to polymeric chains or by affinity bind-ing of antibodies to polymer-modified with correspondingantigen [72]. In the presence of free antigens, competi-tive binding occurs between free and immobilized antigensand results in volume-phase transitions. Following thesame principle, ion-responsive gels are built by endow-ing the network with appropriate ion-sensing molecules[73]. For example, calcium ions can be detected by hydro-gels endowed with calmodulin (CaM), a calcium bindingprotein capable of switching its swelling degree as func-tion of Ca2+ ion content [74]. Alternatively, a network canbe built by imprinting the sensing molecules within thematrix [75]. This approach however requires a high cross-linking density to fix the structure molecular recognitionsites (molecular cavities) in the network that are intendedfor subsequent rebinding of the target molecules with highspecificity.

These examples for different mechanisms of responsein stimuli sensitive hydrogels underline the versatility ofhydrogels that are in a permanent dialogue with their envi-ronments and thus ideal tools for (bio-)chemical sensors.High water content and minimal unspecific interactionwith biological molecules predetermine hydrogels espe-cially for sensing based on biochemical recognition. Hence,the following paragraph is dedicated to the use of hydrogelsin biosensing.

1.4. Hydrogels for sensors

Hydrogels may be prepared in aqueous solutions by UV[76] or thermo-initiated radical polymerization [77], addi-tion reaction [78], self-assembly of recognition motifs suchas coiled-coils [79,80], peptides [81–83], hydrogen-bridges[84] or DNA [12,29,85,86]. These methods can be appliedin a number of more specialized techniques for the pro-duction of peculiar hydrogel structures that are especiallyuseful for sensing. One example is the vapor-phase inducedgeneration of well-ordered arrays of hydrogel microwormswith cylindrical shape and high surface to volume ratio thatmay be used for different sensing applications, e.g. as flu-orescence sodium sensors [87–89]. Hydrogels can also beformed by in situ gelation from low-viscosity solutions andthus be injected into small voids or complex fluidic andmicrofluidic devices. This predisposition to miniaturiza-tion and integration into microsystem in desired geometrybroadens their analytical applications. Hydrogels may thusbe formed by in situ patterning on desired devices for possi-ble sensing applications [90]. All these production-relatedadvantages support the exploitation of hydrogels for ana-lytical purposes.

1.4.1. Hydrogels as sensorsStimulus-sensitive hydrogels may act as active sensing

material according to the mechanisms introduced in Sec-tion 1.3.2. Such gels are sensitive to small changes in theenvironment and give the response to physical stimuli(temperature [91], light, pressure [92], electric field [93],

olymer
ionic strength [94], and magnetic field [95]), chemicalstimuli (pH [96], ions [97]) or biological stimuli (glucose[98], enzyme [99] and antigen [100]) through volumechanges. The response rate is dependent on the hydrogelcomposition, shape and size and can be increased by sev-eral proposed techniques such as reducing the size [101],decreasing the cross-linking density [102], and increasingboth the content of ionic groups [103], and pore size [15]. Inthe presence of these stimuli, the hydrogels undergo phasetransition by the sensing molecules and simultaneouslytranslate this sense into a macroscopic event. In this way,the conversion of hydrogel swelling to an electrical outputis possible with various techniques like light transmissionmeasurement [104], conductometry [105] and pressuregenerated by the gel swelling [106].

1.4.2. Hydrogels for biosensorsThe high water content and hydrophilic nature of

hydrogel is similar to the void-filling component of theextracellular matrix, the natural environment of mam-malian cells, and renders them intrinsically biocompatible.Hydrogels are thus used for a wide range of biomed-ical applications such as soft contact lenses [107] andcontrolled drug delivery systems [108]. Hence, a straight-forward application of hydrogels in biosensors is aprotection and coating function of sensor parts to preventundesired interaction with biological molecules or cells.Their open porous structure and hydrophilic environmentallows diffusion of analytes through the hydrogel matrix.However, diffusion of larger molecules such as proteinsmay be limited and even prevented by an increased cross-linking density, and cells are usually not able to penetrateunless the gel structure is biodegradable.

Beyond simple protection, hydrogels can be used asimmobilization matrix for the biosensing elements. Essen-tial criteria in designing biosensors with high selectivityand sensitivity are receptors that have to be immobilizedin their native conformation to be able to specifically inter-act with the analyte, while support material should beresistant to unspecific adsorption. Furthermore, the qual-ity of sensing and the sensitivity level mainly dependon accessibility and activity of the immobilized sensingmolecules. Hydrogels provide excellent environments forenzymes and other biomolecules to preserve their activeand functional structure [109]. Also cells are commonlyimmobilized in hydrogel matrices for various purposes[110]. As hydrogels may easily be tailored in their prop-erties, they form an ideal platform for this purpose,and extensive work has been carried out with a vari-ety of analytes such as proteins, ions, carbohydrates, andnucleic acids [111]. Recognition in hydrogel-based sen-sors is provided by interaction between an analyte andsensing element by volume changes in response to tar-get molecules. This recognition induced volumetric changecreates a new kind of sensing system as alternative to clas-sical biosensors based on electrochemical sensing.

1.5. Scope and structure of the review

Hydrogels are a broad and extensive field ofresearch, especially for applications in cell culture,

Science 37 (2012) 1678– 1719 1685

tissue engineering, drug delivery and also in microtech-nology. Accordingly, a number of comprehensive reviewarticles have been published over the last 5 years. In 2006,Peppas et al. gave an excellent overview about hydrogels inbiology and medicine [112]. Beside a detailed descriptionof synthetic, biological and biohybrid hydrogel systems,the review intensively describes the use of hydrogels intherapeutics and also covers aspects of diagnostic devices.In 2007 Ulijn et al. published a review concerning biore-sponsive hydrogel systems that change their propertiesin response to selective biological recognition events[36]. They highlight the utilization of such hydrogels fordrug delivery, sensing and tissue engineering. In 2010,Deligkaris et al. published a comprehensive review abouthydrogel-based devices for biomedical applications [113].They survey cross-linking methods, operation principlesand transduction mechanisms. Further applications ofhydrogel devices in specific fields of interest such asfluid control, drug delivery, nerve regeneration and otherbiomedical applications constitute the main focus. Alsoin 2010, Wandera et al. published a review about stimuliresponsive membranes [114]. Their article gives someinteresting examples of hydrogel-based membrane sys-tems that react to changes in e.g. pH, temperature, ionicstrength and chemical cues. All these excellent reviewarticles are either focused on a special class of hydrogelsor a specific area of application. So far, no comprehensivereview concerns the general use of hydrogels in sensingapplications disregarding the physicochemical natureof the gel, the structure and composition of the device,the mechanism for sensing and detection and finally theanalyte.

In the last decade, remarkable efforts have beenachieved to manifold sensitivity and selectivity ofhydrogel-based sensors and biosensors. By using differentgel combinations, the systems with a variety of specifici-ties and improved sensor performance have been prepared.We have recently summarized the use of hydrogels forbiosensing [115]. This review gives an overview on theuse of hydrogels in sensing applications, not only activelyacting as sensing material themselves but also used forimmobilization of molecules that cause the sensing stim-ulus upon presence of the analyte. We have structuredthe following detailed discussion strictly according to thesensing mechanism. Chapter 2 comprises examples forphysicochemical sensing where the signal is generatedas response to changes in parameters such as temper-ature, pH, ion strength, gas and humidity. Importantly,no biochemical recognition process is involved in thesesensing mechanisms. Studies where the sensing mecha-nism is based on such processes, either through molecularinteraction or by reaction of bacteria or cells towardsthe analyte, are presented and discussed in Chapter 3.Molecular interactions may be based on enzyme–substrateinteraction, where a chemical reaction is catalyzed andcauses the effect that may be sensed, or on bind-ing/debinding events caused by presence of the analyte
that cause changes in the hydrogel network. Anti-body/analyte interaction, peptide–peptide bonding orstreptavidin/biotin are examples for such selective bindingsystems with high analyte affinity. Also chemical moieties


F h hologT 1 1) planf e (−1 1

h m temp

to3di

2

sp2hHmsd

2

ofbpfotaa

ig. 5. Structure and optical reflectance of 3D hydrogel PCs formed throughe front side of the image is the (1 1 1) plane. (b) SEM image of the (1

ractured sample. The inset image shows windows between lattices on thydrogel structure measured by Fourier transform IR spectrometry at roo

hat mimick biological ligand binding, such as the bindingf glucose to phenylboronic acids, are included in Chapter. The article is concluded with a short summary and briefiscussion on trends and future perspectives for hydrogels

n sensing applications.

. Physicochemical sensing mechanisms

An excellent and detailed overview of stimulus-ensitive hydrogels used in sensor and actuators has beenublished by van der Linden et al. in 2003 [116]. In010, the use of hydrogels in photonic crystal sensorsas been reviewed in detail by Nair and Vijaya [117].ence, this chapter will briefly introduce the underlyingechanisms that build the basis for sensing and sub-

equently focus on the advances achieved in the lastecade.

.1. Temperature

Temperature sensors are important in a wide rangef science and technology fields such as marine research,ood processing, underground geochemical studies, and iniotechnology [118,119]. Hence, sensors are needed thatrecisely detect the exact temperature in a variety of dif-erent environmental conditions. Despite the high number
f temperature sensors that are available, the sharp andunable entropy driven collapse of LCST polymeric systemst a given temperature is therefore attracting consider-ble attention [120]. Thus, thermo-responsive polymers
raphic lithography. (a) Simulated 3D holographic interferential structure.e of FCC lattices and (c) SEM image of the 408 tilted cross-section of a

0) plane. Scale bar: 500 nm. (d) Reflectance spectrum of a nitrogen-driederature. Figure partly redrawn from [123].

are widely used in temperature sensing as bulk hydrogel,in patterns, as photonic crystal gels and in other physicalforms and structures [121–126]. Accordingly, this chapterconcerns hydrogel-based temperature sensing systems inwhich hydrogels were used as 3D sensing networks formedby temperature-sensitive polymers (as bulk hydrogel form,hydrogel photonic crystals (PCs), intelligent polymerizedcrystalline colloidal arrays (IPCCA)) and as appropriatematrices for temperature sensing probes.

PCs are materials in which the dielectric constant isvarying periodically. This creates a photonic band struc-ture in which electromagnetic waves can or cannot proceeddepending on their wavelength. A number of interest-ing effects result from this such as tunable photonic stopbands that give rise to numerous sensing applications[117,127]. Sensing applications of PC usually exploit achange in the periodic structure due to external stimu-lus or presence of the analyte. With the incorporation ofa thermo-responsive material into the structure, PC canbe used for the detection of temperature. In a typicalexample, thermo-responsive hydroxyethyl methacrylate(HEMA)-based hydrogels PCs were used for temperaturesensing [123]. 3D hydrogel PCs were constructed with acombination of prism holographic lithography and hydro-gel photoresists (Fig. 5). With change in temperature,the swelling of PCs occurred resulting in morphological
changes. The changes were explored by inversion of hydro-gel PCs into silica structure. During the swelling, the latticedistance increases in the (1 1 1) direction and the swollenPCs are deformed and recovered. These changes result in


upon an
Fig. 6. Typical changes in the diffraction from IPCCA particle dispersions
shifting of the photonic stop band with proportional totemperature change (Fig. 5).

Similar to PCs, IPCCA can be used for the detection oftemperature. IPCCA consists of a crystal colloidal arrayof polymer spheres polymerized in a hydrogel matrix[128,129]. The array of colloidal spheres acts similar to pho-tonic crystal and diffracts the light giving rise to intensecolor. Following this principle, with incorporation of atemperature-sensitive polymer, IPCCA structure becomesan ideal candidate for temperature sensing. With changein temperature, the volume transition of hydrogel inducescolor changes by the means of change in diffraction proper-ties. The diffraction through PCCA particles follows Bragg’slaw, which is formalized below [130].

m = 2ndsin (11)

where d is the diffraction order, , the wavelength of lightin vacuum, and n and d refer to refractive index of the sys-tem and the diffracting phase spacing, respectively. Here, shows the Bragg glancing angle. Thus, when the hydrogelvolume increases upon temperature variation, the distancebetween the colloidal spheres increases and so the Braggpeak of the diffracted light shifts to longer wavelengths asillustrated in Fig. 6 [130].

Recently, a temperature sensing construction based onIPCCA particles was designed by cross-linking of acryl-amide (AAm) or N-isopropylacrylamide (NIPAM) withbisacrylamide (BAAm) in the presence of UV photoinitia-tor 2,2-diethoxyacetophenone (DEAP) around polystyrenecolloidal particles [131]. These particles diffract the visi-ble light because the (1 1 1) planes of the face-centeredcubic (fcc) polystyrene colloidal particle array have an∼200 nm lattice constant. This NIPAM-IPCCA is swollenin cold water, but collapses with increasing temperaturedue to the phase transition behavior, resulting in a blue
shift of the diffraction. The detection of temperature issucceeded via monitoring the diffraction wavelength ata defined angle relative to the incident light. The sys-tem works well in the temperature range of 0–60 ◦C.
alyte recognition by the hydrogel matrix. Figure is redrawn from [130].

Phase transition behavior of NIPAM gels can also becontrolled by incorporating more hydrophilic or hydropho-bic monomers. For instance, incorporation of hydrophilicAAm as a co-monomer reduces polymer hydrophobic-ity and increases the interaction between water andhydrophilic groups. Thus, the LCST is increased, since thehydrophobic interactions are compensated for up to highertemperature by the stronger polymer–water interactions[132].

Both methods described above are based on hydro-gel sensing network. However, hydrogels are also used as3D environments to prolong the lifetime of sensing ele-ments, especially in luminescence and fluorescent-basedtemperature sensing probes. These probes are sensitiveto temperature and exhibit temperature-dependent emis-sion. The advantages of this system over other temperaturesensing schemes are (i) high sensitivity, (ii) contactlessoperation, and (iii) inertness to even strong electrical fields[133]. As luminescence probes, europium(III) complexesare widely used for their strong temperature-dependentluminescence and rather narrow emission bands peakingat around 616 nm [119,134,135]. These complexes havebeen designed for wide temperature ranges and have beenimmobilized into polymer matrix to form thin films forsensing applications.

In one study, dipyrazolyltriazine tris(ˇ-diketonate)europium(III) complexes are used as luminescence tem-perature sensor [119]. Incorporation of these probes into apolyurethane hydrogel matrix provided sufficient stabilityof the probes. Temperature-sensitive microbeads areprepared by mixing of palladium(II) 5,10,15,20-tetrakis(2,3,4,5,6-pentafluorophenyl) porphyrin/poly(styreneco-acrylonitril (Pd-TFPP/ PSAN) microbeads and europium(III)tris(thenoyltrifluoroacetonate) trihydrate/poly(4-tert-butyl styrene) (Eu(tta)3L/PTBS) microbeads with a solution
of polyurethane hydrogel (type D4) in ethanol/waterfollowed by stirring for a certain time and subsequentdrying at ambient air. When the resulting beads are excitedwith a light of 405 nm, the bright luminescence exhibits


F lexes ut

ts

tctpsdlaTsoswt1[da

hf2apgbg3

2

e[mcMv

ig. 7. (a) Chemical structures of the indicators of the europium(III) comphe decay time. Figure is partly redrawn [119].

emperature-dependent decay times that can be used toense the temperature between 0 and 70 ◦C (Fig. 7).

In a similar attempt, polyurethane hydrogel stabilizederbium-tris[(2-hydroxy-bezoyl)-2-aminoethyl]amineomplexes (Tb-THBA) were studied as luminescence basedemperature sensor by Sun et al. [133]. When the probehoto-excited at 341 nm, it displays typical Tb3+ ion emis-ion bands with the strongest peak at 546 nm and a typicalecay time of 1.15 ms at 15 ◦C. The emission intensity and

ifetime decrease with increasing temperature, probably as consequence of thermal deactivation of the excited state.he proposed system is appropriate for the temperatureensing over the range from 15 to 65 ◦C and, possibly evenutside this range with further optimization. Anotherimilar luminescent-based temperature sensing systemas constructed by Borisov et al. who used the highly

emperature-dependent luminescence of ruthenium tris-,10-phenathroline embedded in a polyurethane hydrogel136]. From change in luminescence decay time, they couldetermine the exact temperature in the range between 0nd 60 ◦C.

Also fluorescence-based temperature sensing systemsave been reported, for example based on surfactant-

ree poly(vinyl alcohol)/borax/2-naphthol hydrogels [137].-naphtol acts as fluorescence indicator that exhibits

decrease of fluorescence intensity upon rising tem-erature when it is embedded in aqueous PVA/boraxel network with excitation at 365 nm. In contrast, thelue color emission intensity (photoluminescence (PL):max = 426 nm) of 2-naphthol in a basic hydrogel changesradually to strong from 30 to 80 ◦C, when excited at65 nm.

.2. Ions, ionic strength and pH

pH-sensitive hydrogels change their volume, mass andlasticity reversibly in response to a change in the pH value138]. The pH-sensitive character of hydrogels [139–145]

akes them promising materials for a broad range of appli-ations as microsensors [138–141] and microactuators inEMS devices. The principles for swelling dependent pH

alue detection can be various: changes of the holographic

sed for sensing temperature in [119] and (b) temperature dependence of

diffraction wavelength in optical Bragg grating sensors[144] shifts of the resonance frequency of a quartz crystalmicrobalance in microgravimetric sensors [146] a bendingof micromechanical bilayer cantilevers [145], as well as adeflection of silicon membranes in piezoresistive pressuresensors [143,145]. Before we give some detailed exam-ples for pH sensors based on hydrogels, we would like toadvise the reader to have a closer look at the comprehensivereview about hydrogel-based pH sensors and microsensorspublished by Richter et al. [147]. The review introducesthe physical background of the special properties of pH-responsive hydrogels and gives a comprehensive overviewabout transducers which are able to convert the changesof the physical properties of stimuli responsive hydrogelsinto electrical signals.

An example for the use of pH-sensitive hydrogels basedon an IPCCA was developed by Reese et al. [131]. Thehydrogel was constructed by dissolving of acrylamide,bisacrylamide, and 2,2-diethoxyacetophenone in a disper-sion of diffracting polystyrene colloids. This CCA mixturewas injected into a quartz cell with a 250 �m spacer andphotopolymerized. Through partial hydrolysis of an acryl-amide hydrogel with NaOH, the hydrogel network wasendowed with carboxylate groups. As described above, thishydrogel was polymerized around a face-centered cubic(fcc) array of monodisperse, ∼100-�m diameter highlynegatively charged polystyrene colloidal particles. Hence,change in pH-induced swelling/deswelling of the hydrogeldue to protonation/deprotonation of the particle surfaceand the hydrogel. The resulting color change was measuredaccording to the temperature-sensitive IPCCA describedabove. They could determine pH with a 0.05 pH unit reso-lution in deionized water.

The group of Gerald Gerlach used the strong swellingability of pH-responsive hydrogels [145] to develop a sen-sor which combines piezoresistive-responsive elements asmechano-electrical transducers and the phase transitionbehavior of hydrogels as a chemo-mechanical transducer.
They demonstrated the operational principle of chemicaland pH sensors based on the swelling behavior of hydro-gels [146]. For the realization of pH sensors poly(vinylalcohol)/poly(acrylic acid) (PVA/PAA) hydrogels with a pH


Fig. 8. Microlens sensor in which the water–oil interface forms the liquid microlens that is pinned in a construction on a hydrogel ring (a). Expansion andcontraction of the hydrogel regulates the shape of the liquid meniscus by changing the angle v of the pinned water–oil interface (b). The blue dashed linesshow the expanded state of the hydrogel ring (‘IIh’) and the corresponding divergent microlens (‘Im’) at = ˛ . The red dashed lines show the contracted

lens (‘IImure repr

state of the hydrogel ring (‘Iih’) and the corresponding convergent microfigure legend, the reader is referred to the web version of the article.) Fig

value dependent swelling behavior were used as chemo-mechanical transducers for a piezoresistive sensor setup.A similar setup but with poly(N-isopropylacrylamide) hasbeen applied and investigated for pH sensors as well.From the same group, Trinh et al. reported a simula-tion and design study of the system [148]. To optimizethe sensor design, Sorber et al. investigated the swellingeffect of the mentioned (PAA/PVA) hydrogel system byFourier transform infrared (FT-IR) attenuated total reflec-tion spectroscopic imaging under in situ conditions [149].The results of the FT-IR spectroscopic images renderedan improved chemical sensor possible and demonstratedthat in situ FT-IR imaging is a powerful method for thecharacterization of molecular processes within chemical-sensitive materials.

Lee et al. observed the transformations of pH-sensitivehydrogel-based inverse opal photonic sensors from a face-centered cubic (fcc) to a L11-like crystal structure [150].By directly imaging a Rhodamine B (RhoB)-labeled inverseopal hydrogel, using two-photon laser scanning fluores-cence microscopy, it was possible to characterize themesostructure evolution of the pH-sensitive sensors dur-ing swelling. Along with the expected swelling normal tothe substrate, they found that the template inverse opalstructure undergoes a number of unique structural trans-formations upon swelling, including a transformation fromFCC to an L11-like crystal structure.

Shin et al. [151] utilized inverse opal hydrogel struc-tures to fabricate a fast responding photonic crystal pHsensor. The sensor was fabricated by templated photo-polymerization of 2-hydroxyethyl methacrylate (HEMA)and ethylene glycol dimethacrylate (EGDM) within theinterstitial space of a self-assembled polystyrene (PS) col-loidal photonic crystal. After removal of the PS colloidalphotonic crystal structure by chloroform, pH-dependentreflectance measurements were performed using a fiber-
optic UV–vis spectrometer coupled with a reflected lightmicroscope. For calculating reflectance, a silver mirror wasassumed to have a 100% reflectance. The pH sensor exhib-ited response times of about 12 s for pH variations between
’) at = −(90◦ − ˇ).(For interpretation of the references to color in thisoduced from [152].

5 and 6 and had a lifetime of over 5 months without loss ofreproducibility or iridescent color.

A quite interesting optical sensing system was devel-oped by Dong et al. who utilized the stimuli responsiveswelling character of a hydrogel lens [152]. As centralcomponent of their system, a stimuli responsive hydro-gel was integrated into a microfluidic system that servedas the container for a liquid droplet (Fig. 8). The hydrogelsimultaneously sensed the presence of stimuli and actu-ated adjustments to the shape – and hence focal length– of the droplet. The group demonstrated two systems: apH-sensitive liquid microlens, using an acrylic acid basedhydrogel, and a temperature-sensitive liquid microlens,using a N-isopropylacrylamide hydrogel.

An easy and biologically inspired method for sensinga pH value was presented by Kim and Beebe [153]. Themethod is based on elastic instabilities of bi-polymerswelling hydrogels: two different bi-polymer gel stripeswere fixed in such a way together that their differentialresponse in swelling, induced by a designated stimulus,caused the strip to bend. Due to the fixation of the bi-polymer gel stripes the swelling energy released in anexplosive elastic instability. One example was a setupwith an acid responsive dimethylaminoethyl methacrylate(DMAEMA)/hydroxyethyl methacrylate (HEMA)-basedhydrogel on one side and an acrylic acid (AA)/HEMA-basedhydrogel on the other side. In contact with a base, orbuffer the bi-polymer stripe delaminated in a rapid motion(Fig. 9).

Sannino et al. used a quartz crystal microbalance (QCM)to recognize even very small changes in pH [154]. Theyestablished a spin coating method to coat QCM plateswith cellulose based superabsorbent hydrogels to obtainhydrogel-based fast sensors. The hydrogel network wasobtained by cross-linking hydroxyethyl cellulose (HEC) andcarboxymethyl cellulose (CMC) by di-functional divinylsul-
fone (DVS). Due to the polyelectrolyte CMC the hydrogelsdisplayed sensitivity to variations of ionic strength andpH of the water solution in which they were immersed.The change in mass resulted in a change of the oscillating


F d base

o ic force o

fa

beigpcBtt

fpiIbs(aWdia(cdmbto

ig. 9. A self-destructive hydrogel sandwich sensor composed of acid anutward bending is created by stimuli like base and buffer. Once the elast

requency of the quartz plate with a fast response and a finalccuracy on the weight variation of the order of nanograms.

A hydrogel-based method to recognize a local pH distri-ution by a change in color was developed by Maruyamat al. [155]. They found a novel and easy technique tonvestigate local pH distribution on a chip using hydro-el films made of UV sensitive resin (SU-8) patterned byhotolithography. To make the hydrogels sensitive to pHhanges they were functionalized with a pH-indicator, e.g.romocresol Green (BCG). The local pH was measured fromhe color of the hydrogel impregnated with BCG based onhe calibrated color information in YCrCb color space.

Ion sensitive hydrogels generally work in a very similarashion to pH-sensitive systems. Guenther et al. pro-osed an on-line analytical system to monitor metal

ons using gels as chemo-mechanical transducers [156].n their study, they improved a rheochemical sensorased on piezoresistive pressure sensor chips. This sen-or was constructed from two piezoresistive sensor chipschemical and viscosity sensors) that were mounted on

socket with an integrated capillary. A piezoresistiveheatstone bridge was used as mechano-electrical trans-

ucer for the transformation of the plate deflection ωnto an electrical output signal Uout. For ion-sensing,

thin layer of N-isopropylacrylamide (NIPAAm), 2-dimethyl maleimide)-N-ethylacrylamide (DMIAAm) ashromophore and poly(2-vinylpyridine) (P2VP) or N,N-imethyl acrylamide (DMAAm), respectively, as chemo-
echanical transducer, was photo cross-linked onto the
ackside of the bending plate. Immersion into an ion con-aining aqueous solution has induced a volume changef this hydrogel which was monitored by the integrated

responsive hydrogels. The layers are bond together in a way so that anvercomes the bonding force the gels delaminate in a rapid motion [153].

Wheatstone bridge inside a rectangular silicon plate. Theplate deflection caused a stress state change inside theplate, and therefore a resistivity change of the resistorsaffecting proportionally the output voltage of the sensor.An increase of this output voltage corresponded to thehydrogel swelling whereas a reduction of electrical outputvoltage corresponded to the shrinkage of the gel. The sen-sor has been tested in aqueous solutions of alcohol and saltsof different concentrations and showed proof of principlefor this sensor design.

Also the already introduced IPCCA based sensors canbe modified for ion-sensing, for example for the detectionof Pb2+ over a broad range of ion strengths ranging fromhigh ionic-strength solutions to detection in body fluids.Therefore, crown ether groups that chelate Pb2+ with highspecificity where incorporated in the hydrogel networkbetween the spheres, either by functionalization of thehydrogel with crown ether or by direct copolymerizationof vinyl-crown ether into the hydrogel matrix [157]. Whenchelation of Pb2+ with crown ether occurred, the hydro-gel started to swell driven by osmotic pressure. Due to thisswelling, the distance between the spheres increased andred-shift of the diffracted light occurred proportional to thePb2+ concentration. The detection limit of Pb2+ was lowerthan 0.5 �mol/L.

Hydrogel modified microcantilevers have been usedfor ion-sensing. Microcantilevers are small rectangularmonocrystalline silicon tips of less than 1 �m thickness
that provide an excellent dynamic response even to smallamounts of analyte by bending upon analyte adsorptionon one side of the cantilever due to change in surfacetension. CrO4
2− ions can be sensed using microcantilever


y. Inset) etche

Fig. 10. (A) Optical image of perforated diaphragm pressure sensor arra1 mm × 1 mm sensors. (B) SEM micrograph showing the pores (d = 40 �mtaken from [160].

based on self-assembled monolayer (SAM) of a triethyl-12-mercaptododecylammonium [158]. The detection limitof this sensor is as low as 10−9 M, but this sensor gradu-ally loses its sensitivity over 1 week because of instabilityof the layer. Hence, to enhance long term stability, Zhanget al. used a hydrogel layer based on acrylamide and (3-acrylamidopropyl) triethyl ammonium chloride instead ofthe SAM [159]. When this microcantilever is exposed toCrO4

2− ions, deflection occurs as a function of CrO42− con-

centration. This system shows higher specificity to CrO42−

compared to other ions like Br−, HPO42− and NO3

2− with asensitivity as low as 10−11 M.

Following a different methodology, hydrogel arrayscombined with perforated piezoresistive diaphragms wereused for sensing of ionic strength and pH. The sen-sor was constructed with three components: (i) thepiezoresistive sensors, (ii) the pH-sensitive hydrogel(composed of hydroxypropyl methacrylate (HPMA), N,N-dmethylaminoethyl methacrylate (DMA) and the cross-linker tetra-ethylene glycol dimethacrylate (TEGMA)), and(iii) a backing plate (Fig. 10) [160]. Diffusion poreswere comprised to allow analyte flow to the hydro-gel. Alternatively, the backing plate was perforated.Swelling/deswelling of the hydrogel due to changes in ionstrength induced pressure changes onto the piezoresistivesensors that are directly converted into electric signals. Thehydrogel quickly and reversibly swelled when placed envi-ronments of physiological buffer solutions (PBS) with ionicstrengths ranging from 0.025 to 0.15 M, making it ideal forproof-of-concept testing and initial characterizations withhigher sensitivity.

Very recently, a DNA hydrogel system was used for thedetection of Hg (II) ions [161]. The selective binding ofHg2+ between two thymine bases of DNA induces a hairpinstructure. Upon addition of SYBR Green I dye, green flu-orescence is observed only when this hairpin structure ispresent. In the absence of the Hg2+ ion, addition of dye leadsto yellow fluorescence. A thymine-rich DNA-functionalized
polyacrylamide hydrogel was used that allowed sensitiveand selective detection of Hg2+ via a visual fluorescencechange. The proposed system can be regenerated using asimple acid treatment to remove Hg2+ from samples. Using
shows a scanning electron microscope (SEM) micrograph of one of thed into one quarter of the 1 mm × 1 mm sensor diaphragm. The figure is

the naked eye, the detection limit in a 50 mL water sampleis 10 nM Hg2+.

2.3. Gas and humidity

2.3.1. GasIn the field of gas-sensing, hydrogels are mainly used

for three purposes. Due to their anti-adhesive and protein-repellent character, hydrogels can be used just as apassive protection coating for a sensor/electrode [162].Alternatively, hydrogels can be modified by gas sensitivemolecules, e.g. special fluorescent dyes [163] which ren-ders them sensitive for certain gases. Finally, the stimuliresponsive swelling character can be used for sensing,mainly for sensing carbon dioxide after the Severing-haus principle [164,165]. This strategy utilizes the pHchange resulting from CO2 diffusion into an electrolytesolution which can then be sensed by a pH-sensitivehydrogel.

One example for the use as protective layer is theinternal hydrogel separation layer of the planar ultrami-croelectrode nitric oxide (NO) sensor developed by Oh et al.to measure local NO surface concentrations [166]. The sen-sor array consists of a platinized working electrode and asilver paint reference electrode coated with a thin siliconerubber gas permeable membrane. The working electrodeand the gas permeable membrane were separated by aninternal hydrogel layer. They could demonstrate that theNO-selective ultramicroelectrode is an appropriate tool fordetermining accurate steady state surface NO concentra-tions.

A hydrogel gas sensor from the second sub class wasdescribed by Zguris and Pishko [167]. They developed apossible nitric oxide, hydrogel biosensor that utilized 4-amino-5-methylamino-2′,7′-difluorofluorescein (DAF-FM)entrapped in poly(ethylene glycol) (PEG) hydrogel micro-structures. By utilizing the NO-sensitive fluorescence of the
dye and bio-inert character of the hydrogel matrix it waspossible to create a disposable biosensor that is applicablefor the nitric oxide production in certain mammalian cells.The sensor has a lower detection range of 0.5 �M of nitric


f the fin

od

opaddccbrpmtpdstosabcmdcs

nctpas

Fig. 11. Exploded view of all parts and the assembly drawing o

xide in solution. However, due to the irreversibility of theye it is a non-dynamic system that may only be used once.

An example for a hydrogel sensor for the detectionf carbon dioxide based on the Severinghaus princi-le was developed by Herber et al. [168]. Therefore

pH-sensitive hydroxyethyl methacrylate (HEMA)-co-imethylaminoethyl methacrylate (DMAEMA) hydrogelisk, which swelled and de-swelled in response to pHhanges, was clamped between a silicon pressure sensorhip and a porous metal screen together with a bicar-onate solution. CO2 reacts with the bicarbonate solutionesulting in a pH change, which was converted into aressure by the enclosed hydrogel. This pressure was aeasure for the partial pressure of CO2. A more sophis-

icated hydrogel biosensor based on the Severinghausrinciple was described by ter Steege et al. [169]. Theyeveloped the prototype of a continuous hydrogel CO2ensor as an alternative and improvement to standard aironometry. The work follows up on the previous workf Herber et al. [170]. The sensor consisted of a pH-ensitive, dimethylaminoethyl methacrylate (DMAEMA)nd co-monomer 2-hydroxyethyl methacrylate (HEMA),ased hydrogel in a bicarbonate solution mounted on aatheter-tip pressure sensor. It is covered by a gas per-eable membrane (Fig. 11). The hydrogel swells/shrinks

ependent on the CO2 concentration, which leads to ahange in volume and pressure reflected by the pressureensor.

The sensor enabled continuous measurement of lumi-al CO2 and fast detection of sudden and gradualhanges in pCO2 in a clinical significant range. Due to
he hand-made assembly, the sensor lacked the tem-erature stability to meet clinical demands, but theuthors gave a positive outlook for automated assemblyystem.
al hydrogel-based CO2 sensor. Figure reproduced from [170].

A cross-over of the principles was followed byKocincova et al. [171]. They presented an interestingmethod for non-invasive, simultaneous optical mon-itoring of oxygen and pH during bacterial cultivationin 24-well microplates by using an integrated dualsensor for dissolved oxygen and pH values. The dualsensor was based on oxygen-sensitive microparticlesruthenium-tris-4,7-diphenyl-1,10-phenanthrolinedi-(trimethylsilylpropane-sulfonate) (Ru(dpp)3(TMS)2)in organosilica and methacrylate-based pH-sensitivemicrobeads (HQ-N-1-HP1), stained with the pH probe8-hydroxypyrene-1,3,6-trisulfonate (HPTS) embeddedinto a polyurethane hydrogel. The readout was based on aphase-domain fluorescence lifetime. The sensor was usedfor monitoring the growth of Pseudomonas putida bacterialcultures (Fig. 12).

2.3.2. HumidityHumidity sensing is important in numerous industries

such as food, petroleum, textiles, ceramics and manyothers [172]. Optical-based detection via color changeoffers the most attractive way for detection of humiditywithout complex system configuration and tuning incomparison to other sensing mechanisms [173]. Photoniccrystal (PC) hydrogels have been introduced in Section2.1 and are widely used in optical-based sensing devices[128,174,175]. The color change feature of PC hydrogelsdue to swelling upon contact with water predeterminesthem to act as a humidity sensor.

A humidity sensitive PC hydrogel was prepared byinfiltrating acrylamide (AAm) solution into a poly(styrene-
co-methylmethacrylate-co-acrylic acid) PC template andsubsequent photo-polymerization [174]. Due to the intrin-sic humidity sensitive property of acrylamide, a broadrange of humidity (from 20% to 100%) could be identified by


on of a w71].

Fig. 12. 24-Channel SensorDish® Reader on the shaker (a) and cross-sectidetection of the emission is performed through the bottom of the plate [1

obvious color changes. This was attributed to the humid-ity sensitivity of the sample’s stopband. The color changesand optical properties were also investigated by UV–visspectroscopy. The stopbands of the PC hydrogels shift todifferent values with regard to exposed humidity (Fig. 13).

A different photonic crystal based humidity sensingsystem was reported by Kang et al. [123]. They pre-pared a responsive photonic crystal hydrogel-based onamphiphilic poly(styrene-b-quaternized 2-vinyl pyridine)block-copolymers. This structure forms a single one-dimensional periodic lamellar structure resulting in aresponsive photonic crystal gel structure. The glassyhydrophobic layer forces expansion of the hydrophilic layeralong the layer normal, yielding extremely large opticaltunability through the changes in both layer thickness andindex of refractions. The wavelength for this system, waschanged 575% (peak = 350–1600 nm) in the position of thestopband depending on degree of humidity.

Barry and Wiltzius reported the detection of humid-ity via polyacrylamide-based inverse opal hydrogels (IOH)[176]. These smart materials have the capacity to respondto humidity changes by shifting its optical reflection
peak noticeably within the visible wavelength range dueto interlayer swelling and shrinking of the structures.With increasing humidity, equilibrium-swollen hydrogels
Fig. 13. Photographs of PAA-P(St-MMA-AA) hydrogels corresponding to relativehumidity (b). Figure redrawn from [174].

ell of a microplate with included sensor at its bottom (b). Excitation and

absorb more water from air to restore equilibrium condi-tions. This volume change results in shifting the refractivespectrum. These IOH structures exhibited reflectance spec-tra with highly differentiable peaks due to their layeredstructure and Bragg reflection properties.

Galindo et al. reported another optical-based system forhumidity detection [177]. They encapsulated fluorescentflavylium salts in a poly(2-hydroxylethyl methacrylate)(PHEMA) network matrix. Flavylium salts are stronglycolored and display a fluorescence emission that isquenched by water for flavylium salts possessing hydroxylsubstituents. This is due to the presence of efficient inter-molecular excited state proton transfer (ESPT) to water andopens the use of these molecules for humidity sensing.By this, swelling and deswelling of hydrogels at relativehumidities between 7% and 100% could be measured byfluorescence decay times of the flavylium salts.

A long period grating (LPG) coated photopolymerizedpoly (acrylic acid-co-vinylpyridine) hydrogel with N,N′-dimethyl bisacrylamide cross-linker has been used formonitoring relative humidity levels by Wang et al. [178].With increasing humidity, the response wavelength direc-
tion, and the resonance dip increases, resulting in anincrease in the coupling strength. This sensor constructionis efficient in a range from 38.9% to 100% relative humidity
humidity (a) and color/stopband changes of PC hydrogel as a function of

1 olymer

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RH) with a sensitivity of 0.2 nm/% RH and accuracy of2.3% RH.

Arregui et al. [179] used hydrogel-coated optical fibersor the detection of humidity. Four different hydrogels;oly (hydroxyethyl methacrylate), poly (acrylamide), polyN-vinyl pyrrolidinone) and agarose were deposited onptical fiber by means of direct polymerization on theptical fiber surface. All of these hydrogels were suitables humidity sensors. As a generic rule, the bigger poreize, the higher dynamic range and shorter response timesere obtained. Optical humidity sensors were based on the

efractive index changes of hydrogels with humidity. Besterformance was obtained by poly-HEMA hydrogel with aesponse time of 90 s and with a range of application from0% to 100% of RH.

. Biochemical sensing mechanisms

So far, studies and approaches were presented and dis-ussed where sensing was based on swelling or deswellingf the hydrogel network based on physical or chemicaltimuli. This paragraph concerns sensing based on moreomplex and also more specific biochemical mechanisms.t first, sensing based on molecular interactions is dis-ussed, followed by a paragraph about “living” sensorshere bacteria or cells that are immobilized in or on hydro-

els generate the signal that is used for sensing.

.1. Molecular interactions

.1.1. Enzyme–substrate interactionEnzymes are highly specific and efficient in their reac-

ions with substrates. They act as catalysts by lowering thectivation energy of chemical reactions. Therefore, in firstpproximation, the substrate is in a first step forming aomplex with the enzyme by docking into the highly selec-ive active pocket of the enzyme. This enzyme–substrateomplex represents the reactive transition state of theeaction in which the substrate is activated. After the sub-equent chemical reaction the product is released and thenzyme is restored in its functional conformation. If theoncentration of the substrate is sufficiently high, the cat-lyst is saturated with substrate and the reaction is onlyetermined by the rate of conversion from substrate toroduct. This results in a pseudo-first order kinetic of theeaction with regard to the concentration of the enzymeccording to Eq. (10) (Michaelis–Menton equation), where

is the rate of the reaction, kcat the rate constant of theatalyzed reaction, Km a constant that comprises kcat andhe rate constant for the enzyme–substrate complex for-

ation, [E0] the starting concentration of the enzyme andS] the concentration of the substrate.

= kcat · [E0] · [S]Km + [S]

(12)

In reality often more complex reaction schemes suchs multiple binding and allosteric rearrangements occur,
n which the affinity and enzymatic activity is differentor the consecutive binding events. However, the math-matical description of these processes is comparable. Inny case, this brief introduction into enzyme activity and
Science 37 (2012) 1678– 1719

enzyme kinetic shows that enzymatic activity may be usedfor precise determination of analyte concentrations. In thefollowing sections, a variety of different sensors based onenzyme substrate interaction in hydrogel matrices are pre-sented and discussed ordered by the analytes that havebeen detected.

3.1.1.1. Alcohol. Wu et al. developed an organic-phasealcohol biosensor by co-entrapping alcohol oxidase andhorseradish peroxidase (HRP) within an ionotropic poly-mer hydrogel matrix fabricated from silica gel particles[180]. By entrapping the enzymes into the hydrogel matrixit was possible to maintain a bio-catalytic reaction at highand stable rates for an on-line detection of methanol in n-hexane under flow operation mode. The analytical workingrange of the sensor was 2.3–90 mM methanol in n-hexanewith an operational biosensor-lifetime of more than 45assays and a shelf lifetime of longer than 2 weeks.

Jang and Koh [181] employed a multiplexed enzyme-based assay within a microfluidic device using shape-codedpoly(ethylene glycol) (PEG) hydrogel microparticles. Themicrofluidic device was constructed by serially connect-ing two patterning chambers and a microfilter-integrateddetection chamber through a Y-shaped microchannel.Different shapes and sizes of hydrogel microparticleswith entrapped enzymes were fabricated in the pattern-ing chamber by photolithography and collected in thedetection chamber by pressure-driven flow. As a modelexperiment for multiple sensing applications, simulta-neous detection of glucose and ethanol was investigated.Therefore glucose oxidase (GOX) and alcohol oxidase(AOX) were entrapped into differently shaped hydrogelmicroparticles. Each enzyme-catalyzed reaction was eas-ily identified and a simultaneous detection of glucose andethanol was possible by fluorescence imaging in a con-centration range from 1.0 to 10.0 mM without cross-talkby using same fluorescence indicator in a single detectionchamber (Fig. 14).

3.1.1.2. Amino acids. Castillo et al. developed a bi-enzymeredox poly(1-viylimidazole) (PVI) complexed with Os-(4,4′-diethylbpy)2Cl (termed PVI19-dmeOs)/Poly(ethyleneglycol)(400)diglycidyl ether (PEGDGE) hydrogel for detec-tion of l-glutamate [182]. The sensor was capable ofdetecting the release of this excitatory neurotransmitterfrom adherently growing HN10 and C6 cell cultures uponstimulation. Monitoring was utilized using an electrodemodified with the bi-enzyme redox hydrogel at an appliedpotential of −50 mV versus the internal Ag reference elec-trode. The detection limit, 0.5 �M, and the response timeof the sensor, 35 s, satisfied the experimental requirementsfor application in cell culture.

3.1.1.3. Ammonia. A different bi-enzyme sensor was con-structed for the amperometric determination of ammo-nium (NH4

+) by immobilizing glutamate oxidase andglutamate dehydrogenase on a clark-type oxygen electrode
by Kwan et al. [183]. To guarantee the catalytic activitythe enzymes were entrapped in a poly(carbamoyl) sul-fonate hydrogel matrix. The sensor has fast response (2 s)and short recovery times (2 min), a linear detection range


ded hyduorescense and e

Fig. 14. Simultaneous detection of glucose and ethanol using shape-cooptical and fluorescence image depending on sample composition, (c) flreacted with five different samples. (Each sample was composed of gluco

between 10 and 300 �M ammonium and with a detectionlimit of 2.06 �M.

Urea is an important parameter in clinical analysisbecause its increased concentration in blood and reducedlevel in urine is a strong indication for renal dysfunction.Due to the fact that one equivalent urea is hydrolyzedby urease to two equivalents ammonia according tothe reaction (NH2)2CO + 2H2O + H+ → 2NH4

+ + HCO3− it is

possible to determine the urea level depending on theNH4

+ concentration. Eggenstein et al. developed a doublematrix membrane (DMM) potentiometric urea biosen-sor. A NH4

+-sensitive disposable electrode was coatedby a poly(carbamoylsulfonate) (PCS) hydrogel layer withimmobilized urease and a disposable Ag/AgCl electrode asreference [184] (Fig. 15).

The group tested different polymer materials likepolyvinylchloride, polyvinylpyrrolidone, polyestersulfonicacid and photo cross-linkable prepolymers as gel matrixfor entrapping the urease. Best results concerning adhesion

rogel microparticles within microfluidic device. (a) Detection logic, (b)ce intensity of circular and square hydrogel microparticles which werethanol at different molar ratio). Figure was taken from [181].

capability, limit of detection, slope, linear range and life-time were obtained with PCS produced from a hydrophilicpolyurethane prepolymer blocked with bisulfite and cross-linked with polyethyleneimine (PEI) as gel material. Withthe biosensor setup, it was possible to detect urea con-centrations between 7.2 × 10−5 and 2.1 × 10−2 mol/L. Thedetection limit was 2 × 10−5 mol/L urea and the slope inthe linear range 52 mV/decade.

3.1.1.4. Glucose. Due to increasing numbers of diabetespatients, the demand for smart glucose-sensing systemshas grown and with it the effort to utilize hydrogels forglucose-sensing applications. Before concentrating in moredetail on hydrogel-biosensor systems that are used for glu-cose sensing we would like to draw the reader’s attention
on the review “Chemically controlled closed-loop insulindelivery” written by Ravaine et al. in 2008 [185]. Thereview focuses on glucose-responsive hydrogels and givesa great overview about hydrogel systems that are either


F H4+-sen

h de as re

maemr

coWpgTadtgbfmaabhmapmPit2

pwttopv

ig. 15. Schematic view of the disposable urea biosensor consisting a Nydrogel layer with immobilized urease and a disposable Ag/AgCl electro

odified, e.g. by glucose oxidase, lectin and phenylboroniccid moiety, release insulin, or swell/shrink in the pres-nce of glucose. Accordingly, this chapter focuses on theajor strategies for enzyme-based glucose sensing and on

ecently published studies.One of the most conventional methods to detect glu-

ose levels is the use of enzyme-systems like glucosexidase (GOX) [186,187], or glucose dehydrogenase [188].hen for example glucose oxidase is immobilized in a

H-sensitive hydrogel matrix the oxidation reactions fromlucose to gluconic acid will force the hydrogel to swell.he swelling of the hydrogel can then be sensed e.g. by

mass-sensitive magneto-elastic sensors [189]. Kim et al.escribed a micropatterning technique to immobilize pro-eins on the surface of three-dimensional poly(ethylenelycol) (PEG)-based hydrogel microstructures that coulde used for hydrogel-biosensor applications [190]. There-ore hydrogel microstructures were fabricated by replica

olding using a poly(dimethylsiloxane) (PDMS) replicas a molding insert and photolithography. Bovine serumlbumin (BSA), a model protein, was covalently immo-ilized to the otherwise protein-repellant surface of theydrogel structures. The Empty space inside hydrogelicrostructures could be occupied by different proteins

nd sequential bienzymatic reaction of glucose oxidase anderoxidase (POD) could be demonstrated in a hydrogelicrostructure by immobilizing the GOX on the surface of

OD-entrapping hydrogel microstructure. The activity ofmmobilized glucose oxidase was 16.5 U mg−1 and a detec-ion of different glucose concentrations ranged from 0.1 to0 mM was shown.

Another biosensor utilizing GOX entrapped into aoly(ethylene glycol) diacrylate (PEG-DA) hydrogel matrixas developed by Mugweru et al. [191]. In comparison

o the already mentioned swelling dependent sensors,
he group succeeded in fabricating glucose sensor arraysn gold electrodes on flexible polyimide sheets byhoto-polymerization that can be analyzed by cyclicoltammetry. The sensor principle is based on the reaction
sitive disposable electrode coated by a poly(carbamoylsulfonate) (PCS)ference [184].

of flavin adenine dinucleotide of glucose oxidase GOX(FAD)with ˇ-d-glucose, where a reduced form GOX(FADH2), glu-conic acid and hydrogen peroxide are formed inside thehydrogel. The reduced form of GOX(FADH2) is in turn oxi-dized by the electrochemically generated Os3+ form ofthe redox polymer, setting up a catalytic pathway thatproduces an enhanced oxidation peak. The electrons aretransferred from the enzyme to the redox polymer, shut-tled between the redox sites in a self-exchange reactionuntil being transferred to the electrode surface. The cat-alytic current produced is proportional to the glucoseconcentration. The sensor has a linear dependency forconcentrations of 0–360 mg/dL glucose with a calculatedsensitivity of 1.782 �A/(cm2 mM).

In 2009 Merchant et al. built a high-sensitivityreagentless amperometric double function biosensor forglucose and hydrogen peroxide by immobilizing GOXand horseradish peroxidase (HRP) in cross-linked films offerrocene-modified linear poly(ethylenimine) [192]. Fer-rocene was involved as mediator of electron transfer to theelectrode surface. Current densities ∼4 times higher thanthose obtained with other ferrocene-based redox polymerscould be achieved. This sensor generated a limiting cat-alytic density of 1.2 mA/cm2 which is the highest knowndensity ever reported for redox polymers with GOX at pH7. Due to the very high sensitivity (73 nA/cm2 �M) of thesystem it was possible to detect glucose concentrations inthe micromolar range.

In 2011 Kim et al. [193] developed a process tocreate self-assembled peptide hydrogel networks con-sisting of Fmoc-diphenylalanine nanofibers. By physicallyencapsulating enzyme bio receptors, like GOX or HRP,in combination with fluorescent reporters, like CdTe andCdSe quantum dots (QDs), the group was able to create anew optical biosensing platform. Encapsulation of QDs and
enzymes within the hydrogel matrix was achieved in situin a single step by simply mixing aqueous solutions of QDsand enzymes with the monomeric Fmoc-diphenylalaninesolution. The group successfully tested the system for the


lection o attachm

Fig. 16. Measurement setup used for fluorescence excitation and data cola fiber-optic hardware system. Various fiber-optic sensors were tested by

detection glucose and toxic phenolic compounds by usinga photoluminescence quenching of the hybridized QDs.For glucose the degree of quenching was well correlatedin the range from 1 to 10 mM of glucose concentration,which fully covers the range of 3.5–6.5 mM required forthe diagnosis of diabetes. For phenol and hydroquinone theresponse was linear up to 3 mM. Leakage of encapsulatedcompounds from the hydrogel matrix was assessed. Whilethe release of QDs was negligible and only trace amountscould be found in the supernatant, the release test for GOXshowed a release of 3% after 0.5 h, 0.06% after 1 h and 0.002%after 2 h. An initial release rate of 3% is significant and needsto be reduced, as GOX may in solution interact with the ana-lyte and interfere with correct concentration detection ofthe sensor.

In 2010 Siegrist et al. presented first promising stepstowards an enzyme-based electrochemical glucose sen-sor for in vivo operation [194]. The group reports theuse of a novel polypeptide-based fluorescent glucose-sensing system that could overcome many drawbacks ofan enzyme-based system while showing the potential forhigh accuracy, especially at hypoglycemic levels. Fluores-cently labeled glucose recognition polypeptide elementswere derived from glucose binding proteins (GBP) andgenetically modified with a unique cysteine near the ligandbinding pocket. This facilitates a conjugation reaction forsite-specific labeling with N-[2-(1-maleimidyl)ethyl]-7-(diethylamino)coumarin-3-carboxamide (MDCC), an envi-ronmentally sensitive, thiol-reactive fluorophore. The GBPswere immobilized in a polyacrylamide hydrogel matrix andplaced on the tip of an optical fiber to realize a continuousglucose-sensing device (Fig. 16).

The group performed an in vitro validation in buffered
solutions and whole blood. Performance tests in thehypoglycemic levels demonstrated that the reagentlesspolypeptide-based glucose-sensing system has an approx-imate precision of 3.6 mg/dL (0.2 mM) in a glucose range
f continuous glucose sensing by a glucose sensitive hydrogel attached toent to the permanent hardware system [194].

displayed between 90 mg/dL (5 mM) and 36 mg/dL (2 mM).It was further possible to operate the sensor in blood atphysiologic temperature (namely, 37 ◦C) for 1.5 h with aprecision of 9.5 mg/dL in a range of 36 mg/dL (2 mM) and180 mg/dL (10 mM) of glucose, without interference fromother sugars (namely, fructose).

It is important to mention the impact of oxygen concen-tration fluctuations on GOX-based sensing systems. Sinceoxidase-based devices rely on the use of oxygen as thephysiological electron acceptor, fluctuations in oxygen ten-sion and the stoichiometric limitation of oxygen, so-calledoxygen deficit, which reflects the normal oxygen concen-trations, are about 1 order of magnitude lower than thephysiological level of glucose and thus induce changes insensor response and reduced upper limit of linearity [195].Possibilities to address this problem are for example the useof mass-transport-limiting films (e.g. polyurethane or poly-carbonate) for tailoring the flux of glucose and oxygen byincreasing the oxygen/glucose permeability ratio. Alterna-tively, application of materials that act as an internal sourceof oxygen such as a fluoro-carbon pasting liquid (e.g. Kel-Foil) generates an internal flux of oxygen that can thus sup-port the enzymatic reaction, even in oxygen-free glucosesolutions [196–199]. Situations of oxygen deficit can also becounterbalanced by using glucose dehydrogenase (GDH),which does not require an oxygen cofactor, instead of GOx.

3.1.1.5. Hydrogen peroxide. Hydrogen peroxide is clinicallyrelevant as reaction product of the oxidative metabolismcatalyzed by oxidases. It is also used in many industrialprocesses as oxidizing, bleaching and sterilizing agent, sothat its detection is prominent in environmental, clini-cal and biological studies and industries [200]. Various
methods have been developed for the determination ofhydrogen peroxide via different methods like titrimetry,electrochemistry and UV-spectrophotometry [201]. Per-oxidases (POD) are generally used in these constructions

1 olymer

tidtTtasg

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o convert hydrogen peroxide into radicals. The sensitiv-ty of the determination of hydrogen peroxide via PODepends on the ability of the immobilizing matrix to retainhe functional conformation of the enzyme for a long time.hrough the last decade, an impressive number of inven-ive designs for hydrogen peroxide determination haveppeared. Here we will focus on some illustrative exampletudies of hydrogel-based hydrogen peroxide detection,enerally based on electrochemical sensing.

The detection of H2O2 based on an intracellular opti-al nanosensor utilizing poly(ethylene glycol) hydrogelpheres containing HRP was reported by Kim et al. [202].RP was encapsulated in PEG hydrogel spheres by reversemulsion photo-polymerization without losing activity.fterwards, the fluorescence emission response of theseydrogel spheres changed as a function of H2O2 concen-ration in the presence of Amplex Red, and no leachingf HRP was observed from the spheres. The HRP-loadedydrogel spheres were introduced via phagocytosis insideacrophages and were found to respond to both exoge-

ous and endogenous sources of oxidative stress. Thispproach may guide the way to use cells of the innatemmune system as real-time in vivo sensors for oxidativetress situations.

Varma and Mattiasson developed a simple amperomet-ic biosensor for the detection of hydrogen peroxide inater and organic solvents [203]. They entrapped catalase

n 30% polyacrylamide gels and created a two electrodeystem with Pt electrodes and the hydrogel in a teflonolder. By studying the kinetics of the catalase modifiedlectrode by cyclic voltammetry (CV) and coupling it to

flow injection analysis setup it was possible to create sensing system that could monitor a broad range ofydrogen peroxide concentration (0.5–100 mM) in differ-nt solvents. As mentioned above, amperometric detectionf H2O2 has also been achieved via immobilization ofRP in cross-linked films of ferrocene-modified linearoly(ethylenimine) [192]. When small amounts of H2O2ere added to the solution, oxidation peaks disappeared

nd reduction peaks increase due to increase in redox poly-er mediation of the HRP-catalyzed reduction of H2O2. The

ystem showed high sensitivity (500 nA/cm2 3 �M H2O2)nd a linear increase of the current with the H2O2 concen-ration up to 1.5 mM.

A new-type sol–gel/hydrogel composite based on sil-ca sol and a graft copolymer of poly vinyl alcohol with-vinyl pyridine was fabricated for hydrogen peroxideensing by Wang et al. [204]. The film was characterized byourier-transform infrared (FT-IR) spectroscopy and opti-um analytical performance was obtained dependent on

H and electrochemical behavior of the biosensor usingotassium hexacyanoferrate(II) as a mediator. This systemxhibited high sensitivity (15 �A mM−1) and a detectionimit of 5 × 10−7 M.

Optical-based local detection of H2O2 secreted by stim-lated macrophages was achieved through enzyme-basediosensor by Yan et al. [205]. Photolithographic pattern-

ng of a PEG-based hydrogel with incorporation of HRP waspplied to construct microstructures that served as sensinggent (Fig. 17). Amplex Red, a nonfluorescent compoundhat is oxidized to the fluorescent molecule resorufin as a

Science 37 (2012) 1678– 1719

result of the HRP-catalyzed breakdown of H2O2, was eitherimmobilized inside hydrogel elements alongside enzymemolecules or added into the cell culture media during cellactivation. The production of H2O2 after mitogenic stimula-tion of macrophages resulted in appearance of fluorescencein the HRP-containing hydrogel microstructures.

Another hydrogen peroxide sensing hydrogelmicrostructure was constructed by Lee et al. [206].They also used PEG hydrogel as matrix for the sensingenzyme, in this case hydrogen peroxidase. The hydro-gel was micropatterned on a glass substrate as shownin Fig. 18. Cells were cultured ob such substrates, andmicropatterning of co-cultures of hepatocytes and non-parenchymal cells onto non-fouling PEG structures wasachieved, thus offering a technique for precisely control-ling physical and endocrine interactions between twotypes of cells. Also here, Amplex Red served as detectionmethod for the HRP-catalyzed oxidation of hydrogenperoxide.

3.1.2. Antibody–antigenAntibody–antigen based sensors are affinity based

devices with a coupling of immunochemical reactions[207]. The general working principle of this sensor is basedon the specific immunochemical recognition of antibod-ies (or antigens) immobilized on a transducer to antigen(or antibodies), that produce signals which depend onthe concentration of the analyte. Antibody–antigen baseddetection of target molecules have gained attention in lastthree decades due to its high specificity and low detectionlimit [208].

Detection in antibody–antigen sensing system can beperformed in different ways. One of the widely used tech-niques is quartz crystal microgravimetry with dissipation(QCM-D) which provides real-time immunoassay capabil-ity. QCM is a detection method that quantifies the decreaseof the resonant frequency of the quartz crystal when themass increases, e.g. when antigen binds to immobilizedantibodies. Dissipation measurement provides additionalinformation about viscoelasticity of the surface which canincrease with antigen binding. This technique is widelyused in immunoassay applications because of high sensi-tivity.

Carrigan et al. used low molecular weightpolyethyleneimine (PEI) and carboxymethylcellulose(CMC), and covalently immobilized antibody by cross-linking between carboxymethyl groups of CMC andsuccinimide esters generated on the antibody [209]. Thishydrogel-interface has been evaluated with detectionof cytokines rhIL-1ra, rhIL-6 and rhiL-18BPa antigens.Increasing the CMC concentration leads to higher antibodyimmobilization with decreasing non-specific binding. Atoptimum conditions, this system gives a detection limitas low as 400 ng/mL for a 44 kDa antigen. Another studywas performed by the same group using specific antibodycapture ligands such as protein A, protein G, anti-IgG Fc fororiented antibody immobilization and higher frequency
QCM-D crystals to increase the detection limit [210]. Inthis case, detection limit with new technique increasedto 25 ng/mL on 5 MHz crystals. Reducing of non-specificadhesion for this method was shown with incorporation of


Fig. 17. (Up)Fluorescence images of hydrogel microstructures with embedded HRP after 5 min incubation with 5 �M H2O2 (A) and 20 �M H2O2 (B) inpatterntures (D

the presence of Amplex Red. (Down) HRP-containing PEG hydrogel micromacrophages that only attach on the glass next to sensing hydrogel strucbe sensed (E) [205].

PEG into this biointerface [211]. A different methodologyused by Lu et al. is based on periodate-oxidized dextran
layer [212]. Staphylococcus aureus protein A (SpA) wasimmobilized on poly-l-lysine modified piezoelectric crys-tal surface to improve their stability, activity, and binding
Fig. 18. Schematic diagram describing surface microfabrication procedure (Stepstromal cells such as fibroblasts (Steps 6 and 7). Hydrogel microstructures remaincellular microenvironment [206].

s are fabricated on glass substrate (C) followed by seeding and culture of). Release of H2O2 by macrophages upon mitogenic stimulation can thus

specificity with human immunoglobulin G (IgG) in flowinjection assays. The prepared sensing crystal shows best
sensitivity and reusability at a flow rate of 140 �L/minwith a detection limit as low as 0.2 nM with proposedsystem.
s 1–5) to form a micropatterned co-culture assembly of hepatozytes and free of cells and can contain enzyme molecules such as HRP for sensing

1 olymer

(owci[poprtbs[aa

fSaTdltaci(eB(dnrteta

opwigasstertpdwc

oaca


Similar in function to QCM, surface plasmon resonanceSPR) is an established method for the label-free detectionf biomacromolecular interactions. SPR uses evanescentaves to examine the surface. Analyte binding leads to

hange in refractive index at the sensing surface influenc-ng resonant angle and shifting to generate real-time signal213]. SPR allows measurements in real-time with highrecision and repeatability and provides some advantagesver other techniques. These are: (i) the amount of com-lex at equilibrium is measured in the presence of unboundeactant without disturbing the reaction equilibrium, (ii)he stability of the immobilized ligand can be monitoredy tracing the surface binding capacity and the baselinetability, and (iii) the low amount of material for analysis214]. As for QCM, the surface of SPR sensors should be anti-dhesive to unwanted compounds and interact specifically,nd hydrogels have been used for this purpose.

Guidi et al. studied the detection of insulin-like growthactor-1 (IGF-1), a suspected carcinogen, in cow milk viaPR based sensing system [215]. Hyperimmune polyclonalnti-IGF-1 antibodies were immobilized to detect IGF-1.he amount of native IGF-1 in fresh raw milk was detectedown to 4 ng/mL. A similar study was performed by Col-

ier et al. [216]. Another SPR based sensing was used forhe detection of prostate-specific antigen [217]. Anti-PSAntibodies were coupled covalently to gel sensor disc thatan detect PSA molecules with a concentration of 0.15 ngn 1 mL, which is sufficient detection of physiological<4 ng/mL) and clinical (4–10 ng/mL) PSA levels. A differ-nt study was reported for the detection of apolipoprotein100 sensitivity to 10 different monoclonal antibodiesmAbs, all IgG1). Sensitivity was evaluated in terms ofifferent structure recognition. The detection of recombi-ant human interferon-� (IFN-�) with SPR in plasma waseported by Stigter et al. [218]. To test non-specific adsorp-ion of plasma components, numbers of SPR coatings werevaluated. Non-modified dextran coated SPR disks werehe best candidate for the detection of IFN-� in plasma with

lower detection limit of 150 ng/mL.SPR based sensors are also used for the detection

f steroids. In a typical sensor, real-time detection of aair steroids, cortisol and cortisone in saliva and urineas carried out using surface plasmon resonance-based

mmunosensor [219]. A polycarboxylate based hydro-el layer was used for immobilization of the antibodiesnticortisol and anticortisone. These antibodies are verypecific for cortisol and cortisone and contribute to highpecificity of the sensor without binding other steroids likeestosterone, epitestosterone and corticosterone. The layerxhibited high degree of stability over time during repeatedegenerations. Saliva and urine sources were selected dueo presence of low amount of cortisol and cortisone in com-arison to the 100-fold higher concentration in blood. Theetection limit of both steroids in saliva and urine samplesas less than 10 �g/L which is sufficient sensitivity for both

linical and forensic use.A novel immunosensing system for determination

f human �-ferroprotein (AFP) was established using boronate immunoaffinity column via flow injectionhemoluminescence. Interaction between carbohydratesnd boronic acid was used to immobilize the AFP antigen

Science 37 (2012) 1678– 1719

through the sepharose matrix. For assessment of theusability for sensing, a mixture of the analyte AFP and HRP-labeled AFP antibody (HRP-anti-AFP) was injected into thecolumn. The antigen bound to the column which leads tosubsequent trapping of the free HRP-anti-AFP. Detectionwas performed via chemoluminescence due to the sensitiv-ity effects of HRP on the reaction of luminol with hydrogenperoxide [220].

Also three-dimensional (3D) hydrogel thin films wereused for the detection of protein binding (antigenicity) andantibody functionality in a microarray format [221]. Proteinantigenicity was evaluated using the protein toxin staphy-lococcal enterotoxin B (SEB). Cy3-labeled anti-mouse IgG(capture antibody) was microarrayed onto the hydrogelsurfaces and interrogated with the corresponding Cy5-labeled mouse IgG (antigen). Lee et al. introduced anelastic and hydrophilic composite mold, which enables thetransfer of hydrophilic polar ink such as polyelectrolytemultilayer [222]. The mold is composed of an urethane-related commercial polymer and poly(ethylene glycol)diacrylate (PEGDA) which forms a hydrogel upon UV-polymerization. The group established a simple strategyfor immobilizing antibodies exclusively inside the micro-reservoirs of the patterned composite mold, which madethe system suitable for biosensing applications. A first tryregarding the detection of Escherichia coli (E. coli) O157:H7was explored.

Electrochemical based sensing is also used inantibody–antigen sensors for the detection of IgG. Inone case, a redox hydrogel-coated carbon electrode withavidin incorporated in the hydrogel was employed [223].Redox hydrogel film was prepared by electrodeposition ofsolutions containing polyacrylamide-poly(4-vinylpyride)-[Os(bpy)2Cl]2+/3+ (PAA-PVP-Os) and phosphate bufferonto the screen-printed carbon disks (SPE) with apotential of −1.0 V (vs Ag/AgCl) for 2 min. Avidin wasconjugated with biotin labeled anti-rabbit IgG. Afterexposure the tested solution to capture of IgG, thesandwich-assay was completed by binding of HRP-labeledanti-labeled IgG. Electrical contact between the HRP andthe electrode bound hydrogel resulted in formation of anelectrocatalyst for the electroreduction of H2O2 to water.The detection limit with this method was as low as3 ng/mL. Another electrochemical based sensing wasperformed by using a three-dimensional (3D) porouschitosan network as a matrix for hepatitis B surfaceantibody (HBsAb) [224]. Porous network was preparedby electrodepositing a designer nanocomposite chitosansolution with encapsulated silica nanoparticle as hybridfilm on an ITO electrode, followed by removal of the silicananoparticles with HF solution. HBsAb were covalentlyimmobilized to the hydrogel network for the detectionof hepatitis B surface antigen. A novel label-free electro-chemical detection method of ractomine was developedby Shen and He [225]. Ractopamine-bovine thyroglobulinantigen was incorporated into agorose hydrogel filmon a glassy carbon electrode. The K3Fe(CN)6/K4Fe(CN)6
redox probe system was used to translate the bindingto an electrochemical signal with addition of polyclonalantibody against ractopamine. Using similar mechanismand same hydrogel layer with different antibody, the

olymer
detection of ferritin was performed by immobilizing theferritin antibody (FeAb) on the electrode [226]. Cyclicvoltammetry (CV) and differential pulse voltammetry(DPV) were used for the conversion of signal to electricaloutput. The detection limit of ferritin with this method was1.5 × 10−5 g/L and the linear range was 5–50 × 10−5 g/L.

3.1.3. Nucleotide, oligonucleotide and DNAAdvances in nucleic acid detection have become

increasingly important in diagnostics systems by virtue ofimportance in genetic diagnostics, forensic analysis, dis-ease diagnosis and basic research areas [227–229]. Besidethe simplicity combined with low cost, an effective DNA-sensing systems require high selectivity to eventually beable to detect single-nucleotide polymorphism (SNPs) andhigh sensitivity to enable sensing of few hybridizationevents of corresponding small amounts of target DNA frag-ments [230,231]. These high demands have led to thedevelopment of different methods, generally via electro-chemical based sensing systems due to rapid and directdetection combining high sensitivity, small dimensions,low cost, and compatibility with microfabrication tech-nology of transducers [232–236]. Generally, DNA sensorscomprise three main steps: (i) immobilizing the oligonu-cleotide probes on the substrate, (ii) hybridization withthe target sequences through base-pairing interactionsand (iii) read-out by converting DNA hybridization into auseful analytical signal by a transducer [237,238]. Fluores-cence, electrochemical [239,240], optical [241], electrical[242] and microgravimetric based systems [241] are widelyused for the detection of DNA in concentrations down to10−18 M.

Efficiency of nucleic acid fragments immobilizationon the substrate is the key factor of high-throughputDNA analysis technologies as it is crucial for accuracyand sensitivity of the detection. Most studies indicatethat the hybridization is dramatically hindered on solidsurfaces [243], whereas nucleic acid hybridization inhydrogels was reported to closely resemble hybridizationin solution [244]. Moreover, hydrogel-based sensor sys-tems result in 100-fold in signal density due to higherimmobilization capacity of sensing moieties within thematrix [245]. Due to the highly water swollen, hydrophilicnature and flexible structure, hydrogels do not disturbthe thermodynamically stable form of nucleic acid frag-ments and improve the immobilization capacity dueto the three-dimensional structure. These solution-likecharacteristics offer important advantages for capturingprobes in comparison to two-dimensional rigid substratesand self-assembled monolayers providing. In this sec-tion, we present currently used hydrogel-based techniquesfor the detection of oligonucleotides. For general DNA-sensing methods including non-hydrogel-based systems,we would like to draw the reader‘s attention on the reviews“Recent advances in DNA sensors” by Cosnier and Mail-ley [246], “Electrochemical DNA sensors” by Drummondet al. [247] and “Target-responsive structural switching for
nucleic acid-based sensors” by Li et al. [248].
One widely applied tool for DNA sensors are microcan-tilevers. As already introduced in Chapter 2.2, a cantilever isa small rectangular monocrystalline silicon tip of less than

Science 37 (2012) 1678– 1719 1701

1 �m thickness and can transduce changes of mass, tem-perature, or stress, into bending (static mode) and changein resonance frequency (dynamic mode) [249]. Accord-ingly, hydrogel-coated cantilevers combine the advantagesof hydrogels, minimal denaturation of immobilized probesand unspecific interaction, with high sensitivity and directtransduction mechanism without need of any labeling ofthe microstructure-based sensing system. The immobiliza-tion of target-selective receptors on the cantilever makes itpossible to identify analyte molecules by means of changesin surface tension. The change in surface tension will leadto a bending of the cantilever and can be sensed by opti-cal laser detection or capacitance changes [250]. In onecase, cantilevers with electrodeposited chitosan films wereused for the detection of nucleic acid hybridization [251].Electrodeposition was performed by immersing the chipin an acidic chitosan solution and then applying a negativepotential to the electrode in the solution. Subsequent nethydrogel consumption at the cathode leads to a pH gra-dient that, at pH values above 6.5, results in deposition ofchitosan on the cathode. Oligonucleotides were bound tothe chitosan and the resulting biofunctional cantilever usedfor detection of DNA hybridization by static mode (in solu-tion) and dynamic mode (resonant frequency shift in air)as illustrated in Fig. 19.

For the dynamic mode, the resonant frequency shiftwas measured in air before and after the bimoleculardetection event and in the presence of analyte. In staticmode the corresponding measurements were performedregarding bending of the cantilever in solution. Staticmeasurements in solution confirmed successful hybridiza-tion and yielded a surface tension value that, togetherwith the results of the dynamic measurements in airconfirmed surface density of oligonucleotides two ordersof magnitude larger than comparable results obtainedby probes tethered to SAMs. This direct measurementresult demonstrates the higher probe concentration in thethree-dimensional (3D) hydrogel network in comparisonto the two-dimensional (2D) self-assembled monolayersubstrates.

DNA microarrays, integrated into optofluidic hybridiza-tion systems, are also often used in the field of DNAdetection. In one case, a DNA hybridization construct wasdesigned based on microchannels in combination withUV-cross-linked hydrogels and photolithography definedweirs [252]. The PEG-based hydrogel microstructures wereprepared by UV exposure of PEG-DA through a photomaskand used for defining the DNA detection microcham-ber. Single-stranded (ss) DNA probes modified microbeadswere immobilized in the microchambers and used tocapture the complemantry target sequences. The trans-port of complementary DNA target solution was providedby an electrophoretic transport (hydrogel switch open).Subsequently, hybridized DNA targets could be isolatedvia weirs using pump-driven pressure (hydrogel switchclosed). Biotin labeled ssDNA probes were immobilized onthe microbeads and coated with strepavidin. Hybridization
reactions were conducted by moving the targets across thearray of probe-containing microchambers by electrophore-sis. The hybridization of flourescein-labeled ssDNA targetsto complementary probes was observed by fluorescence


F tos of mp

mc

pw[atcllaDtttcot

taget[riwfo

ig. 19. Cartoon illustration (a), SEM (b), and optical microscopy (c) phorofiler scan of chitosan film along dashed line in panel c (d) [251].

icroscopy. In 1 min 90% of the DNA were captured andould sequently be released and recovered.

Another strategy based on a multi-analyte sensor arraylatform consisting of analyte specific features whichere indexed by shape was followed by Meiring et al.

237]. Soft-lithography was used for array fabrication with poly(ethylene glycol) acrylate hydrogel. Fluorescentlyagged methacrylamide-modified oligonucleotides wereopolymerized into the hydrogel matrix and their cova-ent immobilization confirmed. Each target sequence wasabeled with a different fluorophore and retained thebility to hybridize selectively to the designated targets.ifferent DNA probes were incubated in the solution of

hree-different target sequence mix. After the hybridiza-ion the hydrogels were rinsed with buffer to removehe unbound samples. One bright-field and three fluores-ence micrographs were captured; each micrograph wasbtained using a specific optical filter set for one of thehree different target fluorophores (Fig. 20).

By using surface-immobilized DNA-based affinity struc-ures such as oligonucleotides, aptamers or ssDNA withppropriate redox labels, it is possible to monitor their tar-et induced structural switching. This is done by measuringlectrochemical currents which are directly associated tohe distance between the redox label and the electrode227]. In one case, an enzyme-amplified amperomet-ic DNA hybridization with ssDNA probes immobilized
n a carboxymethylated dextran layer on the electrode
as performed by Hajdukiewicz et al. [253]. Using aerrocenemethanol electron mediator, the hybridizationf ssDNA probes to biotinylated-ssDNA and addition of

icrocantilever construction after chitosan electrodeposition and contact

glucose oxidase-avidin conjugate lead to a signal current.To analyze the signal current the group use cyclic voltam-metry and achieved a detection limit of 0.2 nmol in a 500 �Lsample utilizing a graphite electrode.

DNA sensors can also be used to detect gene mutations.One example for this application is the highly sensitivedetection of a K-ras point mutation in codon 12 basedon DNA carrying hydrogel microspheres [254]. The G-G mismatch in codon 12 of the K-ras gene leads to amutation that is suspected to be a reason for pancreaticcancer. This example underlines the crucial importanceof single mismatch detection in early diagnosis of geneticdisorders. The hydrogel microspheres used in this studywere synthesized by precipitation polymerization. Todetect the mutation in the sequence a hydrogel-enhancedsandwich method of a SPR assay was used. The use of theDNA carrying microspheres enhanced the SPR response asa result of an increased dielectric constant. This methodimproved detection limit from 500 to 5 nM, correspondingto a 100-fold improved in sensitivity and 14-fold SPR angleshift compared to that of non-amplified detection.

DNA-functionalized monolithic hydrogels were suc-cessfully used for highly sensitive and selective colorimet-ric DNA detection [161]. DNA bearing methacrylamide-groups at the 5′-end were covalently incorporated intopolyacrylamide hydrogels through copolymerization dur-ing gel formation. In combination with DNA-functionalized
gold nanoparticles that selectively bind to immobilizedprobe-target pairs, high optical transparency of the hydro-gels allowed direct visual detection of down to 0.1 nMtarget DNA. By using the attached gold nanoparticles to


essful de
Fig. 20. Micrographs of the DNA detection assay demonstrating the succand negligible cross-hybridizational noise [237].
catalyze the reduction of Ag+ to metallic silver, 1 pM tar-get DNA could be detected. The gels can be regeneratedby a thermal treatment, and the regenerated gels performsimilarly to freshly prepared ones.

Gu and colleagues designed a system that combines theadvantages of quantum-dot-encoded technology, biore-sponsive hydrogels, and photonic crystal sensor [255]. DNAresponsive hydrogel bead with highly ordered 3D inverseopal structures were fabricated by template replication ofsilica colloidal crystal beads (SCCBs). Hybridization of tar-get DNA with the cross-linked ssDNA in the hydrogel gridcaused hydrogel shrinking, which could be detected as acorresponding blue shift in the Bragg diffraction peak posi-tion of the beads with a detection limit of 10−9 M.

A novel methodology for fluorometric DNA hybridiza-tion detection based on non-covalent coupling of DNA toa water soluble zwitterionic polythiophenylene derivativewas developped by Nilsson and Inganas [256]. Therefore,3-[(S)-5-amino-5-carboxyl-3-oxapentyl]-2,5-thiophenylene hydrochloride (POWT) hydrogel hasbeen cross-linked with ssDNA on a polystyrene surface.The zwitterionic side chain of the polythiophene forms acomplex with single-stranded oligonucleotide inducinga planar polymer and aggregation of the polymer chains.This change detected as a decrease of the intensity anda red-shift of the fluorescence. On addition of a comple-mentary oligonucleotide, the intensity of the emitted lightincreased and blue-shifted. This method is highly sequencespecific, and a single-nucleotide mismatch can be detectedwithin 5 min without using any denaturation steps, withan overall lower detection limit for DNA hybridization of10−11 M.

3.1.4. Other binding mechanisms3.1.4.1. Biotin–streptavidin. Responsive hydrogel-basedmicrolenses can change their optical properties uponvolume-phase transitions in the presence of stimuli. They

tection of each of the complementary target sequences with high signal

can be easily constructed in simple and rapid fashionwith high sensitivity [257]. In hydrogel microlenses,the stimulus induced phase transition changes the focallengths what causes a focusing or defocusing of theprojected image [258]. Most studies in the literatureconcerning hydrogel microlenses for biosensing havebeen performed by Lyon and co-workers. In one case,a bioresponsive hydrogel construct was developed toinvestigate specific and non-specific adsorption [259].Poly (N-isopropylacrylamide-co-acrylic acid) (pNIPAM-co-AAc) based bioresponsive microgels were preparedby incorporation of both biotin and aminobenzophenone(ABP) as a photoaffinity label via carbodiimide chemistry.Functionalization with ABP allows for photochemi-cal tethering of anti-biotin after it is associated to themicrolens through native antibody–antigen association.Hydrogel microlenses were generated by electrostaticassembly of anionic microgels on positively charged,silane modified glass substrates. Two different routes wereused for determination of the specific and unspecific bind-ing. In the first route, the hydrogel microparticles werefunctionalized with biotin with EDC coupling, whereasin the second route, these particles were functionalizedwith biotin-ABP via EDC and DCC coupling. With a sec-ondary IgG exposure, specific/ancillary binding effectswere investigated via brightfield and fluorescencencemicroscopy. Biotin and polyclonal anti-biotin sensitivehydrogel microlenses were constructed using the samemicrolens approach as shown in Fig. 21 [257].

First, pNIPAm-co-AAc hydrogel particles were pre-pared by free-radical precipitation polymerization method(Fig. 21(A)). Then, biotin was incorporated to hydrogelparticles via EDC coupling (Fig. 21(B)), followed by forma-
tion of cross-links in the hydrogel by multivalent bindingof avidin to biotin and anti-biotin antibodies to biotinprovided on the hydrogel microlenses (Fig. 21(C, D), respec-tively). Multivalent binding of both avidin and polyclonal


Fig. 21. Cartoon representation of the hydrogel microlens assay developed in [257] composed of biotinylatedmicrogels. Binding of either streptavidin orb e micro

aosimw

ascs(ioPpdibhual1

ascgirpsiifo

iotin-antibody led to volume changes that altered the focal lengths of th

nti-biotin antibody was determined by monitoring theptical properties via a brightfield optical microscopy. Theame group developed a new method label-free biosens-ng by combining antibody–antigen cross-linked hydrogel

icrolenses to use for the detection of small moleculesith a tunable sensitivity [260].

Piezoelectric biosensors have widely been used forffinity interaction [261]. In one case, hydrogel-coatedtreptavidin piezochips were used for the physical andhemical immobilization of streptavidin molecules. Sub-equent rinsing with biotinylated bovin serum albuminBSA) revealed a higher binding capacity for the covalentlymmobilized streptavidin than on the affinity-preparedne, resulting in a higher subsequent BSA binding. AnotherZ sensor was used for the detection of poly His-taggedroteins using nitrilotriacetic acid (NTA) captured oxidizedextran hydrogel layer [262]. The binding of chelating

ons, Co2+, Cu2+ and Ni2+ was investigated with regard toinding to poly His-tag proteins, and Ni2+ exhibited theighest binding capability. The system was assessed bysing glutathione-S-transferase C-terminally tagged with

6-His tag coated NTA PZ chip to determine the cel-ular glutathione level that is normally in the range of–10 mM.

A hydrogel-based sensor for monitoring proteasectivity has been constructed by combining a versatileolid-phase synthesis (SPS) with the simplicity of liquidrystal display (LCD) [263]. In this study, polyethylenelycol acrylamide (PEGA) was used as SPS support formmobilization of peptide sequences that are specificallyecognized and cleaved by proteases. In the presence ofroteases, the respective peptides were cleaved with highpecificity, resulting in their release from the hydrogel and
nteraction with the LC surface. This leads to a realignmentn the liquid crystalline layer and thus to optical changesrom bright to dark. Hence, this method offers the screeningf protease specificity by eye without need of special
lenses [257].

complex instrumentation (Fig. 22). Due to irreversibilityof the cleavage, the system can however only be usedonce.

Lee et al. created micropatterns on PEG hydrogelsthrough graft polymerization and photolithography for thedetection of streptavidin binding to an avidin modifiedhydrogel layer [264]. Surface modification of the anti-adhesive PEG hydrogel layer was performed by gratingpoly acrylic acid (PAA) via photoinduced surface-initiatedpolymerization with benzophenone as photoinitator.Biotinyl-3,6,9-dioxaoctanediamine was coupled to thePAA micropattern via EDC/NHS-coupling. Specific bindingbetween biotin and streptavidin was used for evaluation ofthis new PAA modified PEG surface and thus showed thatpotential biosensor application of these micropatternedPEG hydrogels are possible.

3.1.4.2. Calmodulin based sensing. Daunert and co-workersdeveloped a new hydrogel microlens system that canchange its optical properties via ligand exchange [265].Protein calmodulin (CaM) was used as a model hinge-motion binding protein for the chemical tunability of themicrolenses. CaM was chemically bound to the hydrogelnetwork as well as its low affinity ligand phenothiazine.The recognition between CaM and phenothiazine led toa further increase in cross-linking density. However, inthe presence of a high affinity CaM ligand like chlorprom-azine (CPZ), an exchange between phenothiazine ligandand CPZ occurs and thus leads to a decrease in cross-linking. In consequence, the hydrogel is swelling and therefractive index decreases. This ability of volume changein response to a (bio)chemical stimulus allows its uses aschemically tunable microlenses. The swelling response of
these stimulus-responsive hydrogel microlenses is quiterapid, reaching 74% of the total response within 120 s,which makes this setup interesting for rapid sensing appli-cations.


er develf the re

Fig. 22. Schematic representation of the protease sensitive sensor chambsplit upon presence of the respective protease, leading to interference oredrawn from [263].

3.1.4.3. Carbohydrate recognition by boronic acids. Besidesthe already mentioned enzyme receptors also syntheticglucose receptors like phenylboronic acids have beenintensively studied for glucose sensing [129,266–268]. Thesensing effect of phenylboronic acid is based on the fact thatthe molecule becomes negatively charged upon sugar bind-ing, which then can generate an osmotic pressure due to theresulting Donnan potential [269]. This results in a swellingof the hydrogel that can be easily sensed. The group of Asherworked intensively on this field and did some pioneeringwork by applying a crystalline colloidal readout betweenglucose and phenylboronic acid [129,270]. Due to theanalyte-induced swelling of the hydrogel described above,the separation between the colloidal spheres increases andthe Bragg peak of the diffracted light will be shifted tolonger wavelengths. Same group developed a fast respon-sive crystalline colloidal array photonic crystal glucosesensor that composed of hydrogel network embed an arrayof ∼100 nm diameter monodisperse polystyrene colloids.This network diffracts light in the visible spectral region asthe glucose concentration varies [271]. The sensor had a fullresponse rate within 90 s to the average glucose concentra-tions found in blood (5 mM) and within 300 s to the averageglucose concentrations found in tear fluid (0.15 mM). Theydemonstrated that the sensor is responsive to ∼0.15 mMglucose concentrations in artificial tear fluid solution.

Another sensing method based on phenylboronic acidwas presented by Worsley et al. in their work aboutcontinues blood glucose monitoring using a thin-filmoptical sensor [272]. The sensor was based on thin-filmpolymer hydrogel containing phenylboronic acid recep-tors. The hydrogel system is composed of acrylamidecross-linked by N,N′-methylenebisacrylamide and phenyl-boronic acid receptors (3-acrylamidophenylboronicacid,and (3-acrylamidopropyl)trimethylammonium chloride).
When the system is illuminated with ordinary whitelight, the holographic fringes within the gel collectivelydiffract a narrow band of wavelengths in a mannergoverned by the Bragg equation (max = 2ndsin �). The
oped in [263]. Fmoc-protected protease sensitive peptides are specificallylease Fmoc with the LCD and color change from white to black. Figure

end result is a characteristic spectral reflection peak,the wavelength of which can be correlated with glu-cose concentration. Utilizing the sensor it was possibleto measure glucose concentrations of 3–33 mmol/L inhuman blood plasma in static mode in vitro. In a flowcell experiment variations in the glucose concentra-tion of approximately 0.17–0.28 mmol/L min could bemonitored.

Lapeyre et al. [273] reported the synthesis of close tomonodisperse poly(N-isopropylacrylamide) submicromet-ric microgels, modified with a phenylboronic acid (PBA)derivative. The PBA-modified particles were sensitive tothe presence of glucose at pH 8.5, i.e., at pH conditions closeto the pKa of the PBA derivative (pKa = 8.2), with a swellingdegree proportional to the concentration of glucose. Dueto their characteristics the particles might serve as build-ing blocks for the design of colorimetric sensors based onthe light diffraction of colloidal crystals.

3.1.4.4. Fluorescence-based glucose sensing. Glucosesensing based on fluorescence is usually based on fluo-rescence (or Förster) resonance energy transfer (FRET).FRET involves the nonradiative energy transfer from afluorescent donor molecule to a fluorescent acceptormolecule in close proximity, and is usually brought aboutby dipole–dipole interactions. The efficiency of this pro-cess is strongly dependent on distance between donorand acceptor dye, and the signal in FRET is a decreasein fluorescence intensity and lifetime of the donor. Forglucose sensing, concanavalin A (Con A), a plant lectinwhich has four binding sites for glucose per molecule, andfluorescently labeled dextran, usually with the fluores-cence dye fluorescein isocyanate, are applied. Sensing isbased on the competitive binding of either glucose or thelabeled dextran to Con A. A more detailed description of
this process as well as some information about the useof PEG-based hydrogels for improved biocompatibility ofimplanted sensors is given in the review of Pickup et al.[274] about fluorescence-based glucose sensors.


Fig. 23. Array patterning template for the immobilization of capture antibodies and application of assay reagents. (A) When the PDMS patterning modulep appliedt e placedd

b[cgsflmplfdaQhlssrc

3Inaacribb

laced in contact with hydrogel, cross-linkers and capture antibodies areoxin and tracer antibody, the six channels of the PDMS assay module arifferent concentrations in each lane to determine detection limits [279].

One recent example for a hydrogel-based fluorescenceiosensor for glucose is the work from Chaudhary et al.275]. The group investigated the feasibility of dissolved-ore alginate-templated fluorescent microspheres aslucose biosensors in simulated interstitial fluid (SIF). Theensor worked on the principle of competitive binding anduorescence resonance energy transfer and consisted ofultilayer thin film coated alginate microspheres incor-

orating dye-labeled glucose receptor and competingigand within the partially dissolved alginate core. It wasurther possible to combine the visible sensing assay (FITC-extran-TRITC-Con A/FITC-dextran-TRITC-apo-GOx) with

near infrared sensing assay (AF-647-dextran amino-SY-21-apo-GOx), co-encapsulated inside the polymericydrogel microspheres. The incorporation of long wave-

ength NIR dyes enables more efficient excitation throughcattering tissue, which effectively improves the odds ofuccessful use in vivo. The system exhibited a maximumesponse time of 120 s and a sensitivity of 0.94%/mM glu-ose with a response in range of 0–50 mM glucose.

.1.4.5. Peptide modified hydrogels for toxin screening.mmobilization of peptides and especially proteins on pla-ar sensor substrates often results in (partial) loss ofctivity or in the case of antibodies reduced affinity tontigens [276–278]. This effect originates from interfa-ial energy driven structural deformations in the binding
egion or even the protein structure itself, with negativempact on the activity, binding kinetics, or even inhi-ition of the complete binding process. As highlightedefore, the solution-like environment of water swollen 3D
into each of the six lanes and allowed to incubate. (B) For application of parallel to the patterned capture antibodies on the hydrogel slide with

hydrogel networks can preserve the protein structurewhile also granting multidirectional access for ligand bind-ing to the active binding sites. Charles et al. made use ofthese benefits that 3D hydrogels offer when they devel-oped a galactose polyacrylate-based hydrogel scaffold forthe detection of cholera toxin (CT) and staphylococcalB (SEB) in a sandwich immunoassay format [279]. Theysucceeded in creating a hydrogel system that providedenhanced sensitivity with low non-specific binding charac-teristics. By utilizing a PDMS microchannel template theypatterned arrays of the immobilized antibodies for thedevelopment of a sandwich immunoassay (Fig. 23). Flu-orescence and 3D imaging by confocal microscopy andlaser scanning confocal microscopy of Cy3-labeled anti-CTand/or Cy3-anti-SEB tracer molecules provided qualita-tive and quantitative measurements on the efficiency ofprotein immobilization, detection sensitivity and signal-to-noise ratios. With the established system they were ableto achieve low limits of detection for SEB and cholera toxins(1.0 ng/mL).

Beside the integration of active biomolecules it is alsopossible to create hydrogels that are cross-linked by pep-tides, which dissolve in the presence of certain stimuli, likereducing agents or enzymes [280,281]. The cross-linkingsystem can be designed in such a way that the hydrogelcan act as a self-sacrificing biosensor for enzymatic activeanalytes. Such “sacrificing” hydrogels can e.g. be used for
sensing toxins that possess enzymatic activity, such asbotulinum neurotoxin type A (BoNT/A). Frisk et al. used so-called “SNAP” (synaptosome-associated protein) modifiedhydrogels in combination with microfluidics for sensing


Fig. 24. Fifteen micrograms per milliliter botulinum neurotoxin type A (BoNT/A) degrade synaptasome-associated protein (SNAP) modified hydrogel postsensors. (a) Initial 425 �m SNAP peptide post in 1× PBS, t = 0 h; (b) after 48 h exposure to BoNT/A; (c) 325 �m control post with BoNT/A insensitive cross-

–c are e
linker after 48 h toxin exposure. Corresponding images directly below ahydrogels. Scale bars = 150 �m [282].
enzymatic activity of BoNT/A [282]. They were able togenerate autonomous sensors that simply required visualreadout (Fig. 24). Introduction of ca. 20 �L toxin-laden solu-tion into the microfluidic device allowed observation of thedegradation without the necessity for detectors, pumps, orvalves, thus making microfluidic-based biosensors promis-ing as toxin-screening vehicles. Due to the fact that theanalysis of the toxin level is done by an optical comparisonof non-degraded posts to degraded posts, this system canbe used as a great indicator for toxins in a medium, but notas an accurate analytical method. To improve the readoutand analytical process it could be interesting to integratesuch a peptide cross-linked hydrogel system into a hydro-gel biosensor based on photonic crystals. In this case eventhe smallest degradation could be sensed.

3.2. Living sensors

This chapter gives an overview about the combina-tion of hydrogels with living cells and microorganismsto form living cell–polymer composites for biosensingapplication. The use of microorganisms and cells as bio-logical sensing elements in biosensors is well established.Microorganisms offer advantages of ability to detect awide range of chemical substances, amenability to geneticmodification, and broad operating pH and temperature
range, making them ideal as biological sensing mate-rials. For in-depth overviews on this field we want todraw the reader’s attention to the short but comprehen-sive reviews about “Microbial Biosensors” by Lei et al.
dge-outlined using ImageJ software for better visualization of degraded

[283] and “Whole Cell Biosensors” by Bousse [284]. Thereview of Lei et al. concentrates on the advantages of usingmicroorganisms as biosensing elements, the immobiliza-tion of microorganisms, electrochemical, amperometric,potentiometric, conductimetric, optical, bioluminescence,fluorescence and colorimetric microbial biosensors aswell as Microbial fuel cell type biosensors, fluorescenceprotein-based biosensors and O2-sensitive fluorescentmaterial-based biosensors. Luc Bousse concentrates onthe field of biosensors in which the biological componentconsists of living cells and points out some of the mostimportant reasons for using living cells, i.e. to obtain func-tional information about the effects of a stimuli on a livingsystem.

This chapter focuses on advances in the use of liv-ing cells and microorganisms in biosensors utilizinghydrogels in the last decade. The immobilization ofcells in different hydrogel types for biosensing appli-cations has been reported for polyacrylamide [285,286]polyurethane-based hydrogels [287], photo cross-linkableresins [287,288] and polyvinyl alcohol [289–293], poly-acrylonitrile membranes [294] and poly(ethylene glycol)hydrogel [295], alginate [296–298], carrageenan [286],low-melting agarose [299], chitosan [300] and proteinicmatrix [301]. One of the main benefits hydrogels have istheir high water content and the three-dimensional poly-
mer network, which allows a high gas exchange rate andtransport of nutrition’s to cells immobilized in the polymernetwork. These characteristics, in combination with theirbiocompatibility, make it possible to even cultivate cells or


F . It illusta .

bc

3

fnwobatt(5

achTcaethnmidchspdg2

ig. 25. Schematic diagram of the A. adeninivorans LS3 microbial sensorfter the addition of biodegradable pollutants. Figure redrawn from [302]

acteria inside of hydrogels and to use the cell–polymeromposites for biosensing applications.

.2.1. BacteriaBacterial cells as biological recognition elements have

or example been used in a biosensor for the rapid determi-ation of the concentration of biodegradable pollutants inastewater by Renneberg et al. [302]. The sensor consisted

f salt-tolerant yeast Arxula adeninivorans LS3 cells immo-ilized in a Poly(carbamoyl)sulfonate (PCS) hydrogel, on

Clark-type oxygen electrode. The group claimed thathe sensor made it possible to get comparable results forhe so-called standard 5-day Biochemical Oxygen DemandBOD5) method, employed as a pollution index, in just

min (Fig. 25).Another example for yeast cells in a hydrogel matrix

re the fiber-optic biosensors based on luminescent yeastells entrapped in calcium alginate, or polyvinyl alco-ol (PVA) hydrogels, constructed by Fine et al. [303].he group entrapped genetically modified Saccharomyceserevisiae cells, which contained the estrogen receptorlpha-mediated expression of the luc reporter gene andnabled a novel approach for estrogenic endocrine disrup-ing chemical (EDC) bio-detection. Due to the protectiveydrogel matrix it was possible to operate even underon-sterile conditions. Martinez et al. described a fastethod to study cell metabolic characteristics immobiliz-

ng yeast cells (S. cerevisiae) in a microbial biosensor-likeevice [304]. Main components of the sensor where S.erevisiae, e.g. entrapped in a poly(carbamoylsulphonate)ydrogel membrane, and a carbon dioxide electrode. Theensor device allowed the calculation of Michaelis–Menten
arameters related to the kinetics of transport and degra-ation of several carbohydrates (i.e., glucose, fructose,alactose, sucrose and xylose, with Km(app) of 6.0, 5.8, 0.9,.0, and 147 mM, respectively), and the study of the kinetics
rates the microbial consumption of dissolved oxygen (a) before and (b)

of expression of nonconstitutive proteins related to thetransport and degradation of galactose.

Koster et al. made use of an immobilized Staphy-lococcus type to create an amperometric, microbialrespiration-based microbiosensors for the quantification ofavailable dissolved organic carbon (ADOC) instantaneouslyrespired by microorganisms [305]. Therefore they immo-bilized aerobic seawater microorganisms, closely relatedto Staphylococcus warneri, in a polyurethane hydrogelsensing membrane. The group focused on biochemical andphysiological properties of respiration-based ADOC micro-biosensors that measure all of the dissolved organic carboncompounds that are instantaneously respired by the immo-bilized cells. The sensor had a detection limit equivalent toabout 6–10 �M glucose in a 90% response time of 1–5 min.The shelf lives of individual sensors were up to 2 weeks.

With E. coli Fesenko et al. used one of the mostpopular bacteria strains for their hydrogel microchip devel-opment [306]. They developed a cell–hydrogel systemfor biosensing and monitoring of cell populations usinga polyacrylamide-based hydrogel bacterial microchip(HBMChip). The microchip represented an array of hemi-spherical gel elements that contained live immobilizedE. coli cells. They were able to measure and monitorintracellular metabolism using vital fluorescent stains,engineered constructs encoding bioluminescent or fluores-cent reporters. It was possible to monitor the dynamics ofnucleic acids in growing E. coli cells using vital fluorescentstain SYTO 9. Further they made use of the biosens-ing opportunities the cell-chip offered and illustrated thedetection of antibiotics and sodium meta-arsenite by quan-titative analysis.

3.2.2. CellsThe same characteristics that make hydrogels prefer-

able candidates for bacterial biosensors count as well

olymer
for cell-based biosensors. Besides the already mentionedbiocompatibility and the high diffusion rate for nutri-tion’s, gas etc. hydrogels have structural similarity to themacromolecular-based cellular microenvironment in thehuman body, the so-called extracellular matrix, whichoffers applications in vivo and in vitro. To get a bet-ter overview about human cell–hydrogel composites weadvise the reader to have a closer look at the review“Hydrogels for Tissue Engineering” written by Lee andMooney [307]. Further Hynd et al. published a compre-hensive review about hydrogels in applications for neuralcell engineering [308]. The review gives a nice overviewabout acrylate-based polymers and their derivatives, usedin biomedical applications, including drug delivery, tis-sue engineering, cell encapsulation, and as templates fordirected cell growth, attachment and proliferation. Here wewant to focus on hydrogels used for cell-based biosensors.

3.2.2.1. Hydrogel-coatings and influence of the third dimen-sion. In the field of cell-based biosensors hydrogels arepreferably used in two different shapes: either as coat-ing material for recognition elements like electrodes,implantable sensors etc., or as 3D-Matrix for cells that areutilized as biological recognition element themself. Hydro-gels that are used as sensor coatings in most cases playa passive role; they e.g. support gas transportation, ion-exchange etc., or make it possible to immobilize cells onsensor surfaces. An example for the immobilization of cellson a sensor surface is the work of Ding et al. [309]. Theyprepared a nanocomposite by neutralizing a solution ofchitosan encapsulated gold nanoparticles formed by reduc-ing in situ tetrachloroauric acid in chitosan. The gel wasdesigned for immobilization and electrochemical study ofcells, monitoring their adhesion, proliferation, and apopto-sis on electrodes. The nanocomposite gel showed improvedimmobilization capacity for K562 leukemia cells, as a modelcell line, and good biocompatibility for preserving theiractivity on glassy carbon electrodes. Another examplecame from the same group when Hao et al. combinedthe biocompatibility of chitosan (CS) and conductivity ofcarbon nanofiber (CNF) [310]. By an electrodeposition ofsoluble CNF-doped CS colloidal solution they were able toform a robust CNF–CS nanocomposite film with good bio-compatibility for the immobilization and cytosensing ofK562 cells on an electrode.

The biocompatibility of hydrogels can be increased bymodifying them with certain biomolecules that supportthe immobilization of cells, or decrease the inflamma-tion of tissue in case the sensor is implanted. Nortonet al. examined the release of vascular endothelial growthfactor (VEGF) and dexamethasone (DX) from non-foulinghydrogel-based sensor coatings and their effect on the for-eign body response [311]. The hydrogels were preparedfrom 2-hydroxyethyl methacrylate, N-vinyl pyrrolidi-none, and polyethylene glycol. The study revealed thatVEGF releasing hydrogel-coated fibers increased vas-cularity and inflammation in the surrounding tissue
after 2 weeks of implantation compared to hydrogel-coated fibers. DX releasing hydrogel-coated fibers reducedinflammation compared to hydrogel-coated fibers andhad reduced capsule vascularity. Due to the fact that
Science 37 (2012) 1678– 1719 1709

after 6 weeks, there were no detectable differencesbetween drug-releasing hydrogel-coated fibers and con-trol fibers it could be concluded that the drug depot wasexhausted.

Aside of immobilizing cells on hydrogel-coatings, theycan be embedded in the coating as well. In this case thehydrogel structurally supports the cells by forming a cellfriendly 3D environment and cells can be cultivated as closeas possible to the sensor surface. Klueh et al. for examplereported the use of tissue-interactive bio-hydrogels, suchas fibrin, to enhance gene delivery at sites of sensor implan-tation [312]. Therefore cells, genetically engineered toover-express the angiogenic factor (AF) vascular endothe-lial cell growth factor (VEGF), were incorporated into anex ova chicken embryo chorioallantoic membrane (CAM)-glucose sensor model. The VEGF-producing cells weredelivered to sites of glucose sensor implantation on theCAM using a tissue-interactive fibrin bio-hydrogel as acell support and activation matrix. This VEGF–cell–fibrinsystem induced significant neovascularization surround-ing the implanted sensor, and significantly enhanced theglucose sensor function in vivo.

Desai et al. demonstrated different behavior of cellscultured on a two-dimensional coating or in a 3D net-work [313]. They used SH-SY5Y human neuroblastomacells, cultured in 3D collagen hydrogel, and applied confo-cal microscopy and immunofluorescence staining to assessthe merit of the system as a functional, cell-based biosen-sor. They were able to show the depolarization-induceddifferences in intracellular calcium release when culturedusing a 2D versus a 3D matrix. Similar results regardingthe difference in on 2D, or in 3D substrates could be shownby Wu et al. [314]. They fabricated a packed Cytodex3 microbead array as a simple 3D cell-based biosensingformat. Resting membrane potentials and voltage-gatedcalcium channel (VGCC) function of SH-SY5Y human neu-roblastoma cells cultured on the microbead array versuscollagen-coated flat (2D) substrates were evaluated byconfocal microscopy with a potentiometric dye and acalcium fluorescent indicator. The exaggerated 2D cell cal-cium dynamics, in comparison with those of 3D cells,were consistent with previous 2D/3D comparative studies.The study established the rationale and feasibility of themicrobead array format for 3D cell-based biosensing.

3.2.2.2. Cell preservation in 3D hydrogels. This paragraphhighlights some recent results regarding the influence ofentrapping cells into hydrogels and possibilities to pre-serve them over longer times in in vitro. Here we wouldlike to draw the reader’s attention to the review aboutpreservation methods to keep biosensor microorganismsalive and active, published by Bjerketorp et al. [315]. It isknown from literature that cells can survive the entrap-ment into hydrogels. Zguris et al. demonstrated that itis possible to immobilize cells in a hydrogel matrix byphoto-polymerization without a loss of activity [316]. Theydemonstrated the ability to entrap human microsomes in a
poly(ethylene) glycol hydrogel microstructures by photo-polymerization without loss of activity, as monitored byreference to the activity of cytochrome P450. They fur-ther developed two photo-polymerization methods for the


F e (PEGDp ne oxid

ftBtomtuRaTb

htfHgm

ig. 26. Schematic illustration of poly(ethylene glycol) dimethacrylatoly(allylamine hydrochloride); PAA: poly(acrylic acid); PEO: poly(ethyle

abrication of heterogeneous PEG hydrogel structures con-aining multiple phenotypes of mammalian cells [317].eside a two-step photoreaction injection molding processhey described a novel method for one-step fabricationsf stacked hydrogel microstructures using a microfluidicold. In addition to hydrogels containing fluorescent dyes

hey were able to create differing extracellular matrixes,nmodified or modified with the cell adhesion promoterGD, within one hydrogel microstructure and to gener-te patterns of hydrogels containing multiple phenotypes.he studies revealed new promising opportunities for cell-ased hydrogel biosensors.

Preservation of cells for certain time periods in aydrogel matrix has been shown by Itle et al. whenhey entrapped cells into PEG hydrogel microstructures
or biosensing applications [318]. In their work aboutepatocyte viability and protein expression within hydro-el microstructures they reported the entrapment ofammalian hepatocytes in hydrogels of varying PEG
MA) hydrogel microwell array fabrication and T-cell arraying. PAH:e) [321].

compositions. They were able to maintain cell viability for7 days in free structures and in structures contained inmicrofluidic channels. The production of protein by cellsin hydrogels decreased with increasing PEG concentra-tion, suggesting mass transfer limitations in hydrogels withhigher PEG compositions. Following up on this topic theyreported the fabrication of mammalian cell-containing PEGhydrogel microarrays by a “stage-down” freezing processfor potential applications in drug screening and pathogendetection [319]. The work revealed possibilities of fabricat-ing cell-containing microdevices and cryopreserving themfor later use. Their work resulted in the fabrication of wholemammalian cell biosensors for the optical monitoring ofcell viability and response to two model chemotoxins:sodium hypochlorite and sodium azide, and one model
biotoxin, concanavalin A [320]. Therefore the cells wereentrapped either in PEG microspheres, or in array format.It was possible to detect sodium azide in micromolar quan-tities in fewer than 30 min.


Fig. 27. (a) Representation of the array of electrode employed in [323]: i, ii and iii are working electrodes; iv, v and vi are counter electrodes; and vii isthe reference electrode. (b) Photograph and, (c) schematic representation of the well plate with and without the membrane insert positioned over the
integrated electrodes on the bottom. Figure redrawn from [323].
3.2.2.3. Pathogen sensing. One example for pathogendetection with hydrogel biosensors is the work of Kimet al. who examined the potential of a sensing strategybased on live T-cell/B-cell interactions in a biosensor chipformat [321]. They developed an approach to fabricatepatterned poly(ethylene glycol) dimethacrylate (PEGDMA)hydrogel microwells functionalized at their bases withantibodies to promote specific immobilization of lym-phocytes. The microwells were used to array single Tcells in a regular pattern on a substrate. A sensing plat-form was created by overlaying arrayed T cells witha confluent layer of antigen-capturing B cells. Utilizingthe hydrogel/lymphocyte sensing chip it was possible todetect a model peptide analyte within minutes, with dose-dependent linear signals over the concentration range0.05–5 �M (Fig. 26).

Another example for the detection of pathogens waspublished by Banerjee et al. [322]. They developed a multi-well plate-based biosensor containing B-cell hybridoma,Ped-2E9, encapsulated in type I collagen matrix, for rapiddetection of pathogenic Listeria, the toxin listeriolysin O,and the enterotoxin from Bacillus species. This sensormeasured the alkaline phosphatase release from infected
Ped-2E9 cells colorimetrically and gave an example ofa cell-based sensing system using collagen-encapsulatedmammalian cells for rapid detection of pathogenic bacteriaor toxin.
3.2.2.4. Nitric oxide/glutamate. Castillo et al. presented thesimultaneous detection of nitric oxide and glutamate usingan array of individually addressable electrodes, which weremodified with a sensitive nitric oxide sensing chemistryand a glutamate oxidase/redox hydrogel-based glutamatebiosensor [323]. The first working electrode was coveredwith a positively charged Ni porphyrin, entrapped intoan negatively charged electrodeposition paint, the sec-ond working electrode by a bi-enzyme sensor architecturebased on cross-linked redox poly(ethylene glycol) digly-cidyl ether (PEGDGE) hydrogels with entrapped peroxidaseand glutamate oxidase. Adherently growing C6-gliomacells were grown on membrane inserts and placed in closedistance to the modified sensor surfaces. After the estab-lishment of a stable background current at both electrodes,20 �L of a mixture containing 10 �L (160 nM) bradykininand 10 �L KCl solution (2 M) were carefully injected intothe buffer solution around the cells, and changes in cur-rent signals at both sensors were monitored. The currentresponses recorded at each electrode after stimulation ofglutamate and NO release by means of K+ and bradykinindemonstrated the ability of the individual electrodes in thearray to detect the analyte without interference from the
neighboring electrode or other analytes present in the testmixture (Fig. 27).
From the same group Wartelle et al. reported testresults from a compact, biocompatible, hydrogel-based,

1 olymer

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4

tittoateahgooicsstac

tUntdaplm

trErreasbtasimttbntu


lectrochemical sensing device for the detection of nitricxide released form C6-glioma cells [324]. Due to the use ofn ionically conducting kappa-carrageenan hydrogel mem-rane, in combination with the integrated electrode setup,sed by Castillo et al. as well, it was possible to create a

iving cell-based device for NO-detection.

. Summary and perspective

Especially in the last decade, hydrogels have gainedremendous interest in the field of (bio-)sensing. Manyntrinsic properties of hydrogels predetermine their use forhis application, most importantly their high water con-ent that renders them biocompatible and allows diffusionf water soluble compound through the polymer networks well as their chemical diversity and the ability to tuneheir chemical and physical properties. This allows the gen-ration and design of stimuli sensitive hydrogels that mayct as sensors themselves, for example by introduction ofydrophobic groups for temperature sensitivity or proticroups for pH sensitivity. It also enables functionalizationf the hydrogel with moieties that act as sensors. The rangef molecules that can be and has been used for this purposes very broad and comprises (fluorescence) dyes, lumines-ent compounds, enzymes, biochemical recognition sitesuch as oligonucleotides, antibodies, lectins and other verypecific recognition mechanisms. Due to accessibility ofhe complete hydrogel volume high sensitivities can bechieved what makes the use of hydrogels for sensing espe-ially interesting.

Hydrogels may be prepared in aqueous solutionshrough a variety of mechanisms. Most prominently, eitherV or temperature initiated radical polymerization tech-iques are used. The gelation from low viscous solutionso gels allows application specific preparation of the gels inesired geometries and is compatible with miniaturizationnd microfluidic applications. Moreover, hydrogels can berepared by preparation methods that are compatible with

iving cells or bacteria so that these can be encapsulated anday serve as sensors in the hydrogel network.Recently, especially multifunctionalization of the gels,

he combination of hydrogels and other intelligent mate-ials and advances in sensor design has led to advances.xamples are combinations of responsive hydrogels thateact with volume changes to stimuli and piezoelectriceadout. Also optical readout through microsphere arraysmbedded in a responsive hydrogel matrix is an elegantpproach. Double modification of hydrogels with enzymesuch as peroxidase and glucose oxidase allows sequentialienzymatic reactions in the hydrogel. Aside of applica-ion specific sensor design, microfabrication techniquesre more commonly used to miniaturize and parallelizeensing, and hydrogels offer multiple application pointsn the combination with miniaturization. Future develop-

ents of hydrogels for sensing will majorly benefit fromhe growing precision in macromolecular chemistry to con-rol size, shape, structure and functionality of polymers as
asic building blocks for hydrogels. Moreover, the combi-ation with modern biotechnological tools will allow forhe design and preparation of biohybrid-hydrogels withnsurpassed specificity and sensitivity in sensing.
Science 37 (2012) 1678– 1719

Acknowledgments

The authors gratefully acknowledge the DFG fundedSPP 1257 for financial support and stimulating discussions.F. T. thanks to Marie Curie ITN project Hierarchy (con-tract: PITN-2007-215851) for PhD fellowship. This reviewevolved from the book chapter cited in reference [115],coauthored by the authors of this review, together withD. Tanaka. This review, which presents an updated, moredetailed and broadened version of the topic in reference[115], reproduces segments from that source, with the per-mission of the publisher, Elsevier Ltd.

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