Chemical Reviews: Modern Topics in Chemical Sensing issue 2

Volume 108, Number 2

Introduction: Modern Topics in Chemical Sensing Chemical Sensors is a large branch of analytical chemistry that has unique characteristics. It is very broad, ranging from solid-state physics to molecular biology. Because of this breadth, choice has to be made between the all-encompassing, but rather shallow treatment of the subject on one hand or selection of a few topics covered in depth by experts in the field. The latter approach has been chosen for this issue. It does not pretend to cover all new aspects of this active and growing field. Several rather important modern topics of chemical sensing that could have been included have not been, mainly because of the lack of available and willing authors. There is some confusion in the terminology. The label “chemical sensor” is often used to describe an analytical procedure that should be more appropriately called an “analytical assay” or “sensing system”. The main difference between the two lies in the mode of information acquisition. While a chemical sensor acquires information continuously, a sensing system obtains information in discrete steps. It does not matter that most modern analytical, specifically bioanalytical, assays are automated and can run unattended for long periods of time. The discontinuous nature of their operation still distinguishes them from true chemical sensors. The two groups of procedures are fully complementary and valuable tools of analytical chemistry. This thematic issue is limited to true chemical sensors only. A typical modern chemical sensor consists of a physical, “transducer” and a chemically selective material. Different strategies can be employed to extract maximum information about the sample. Among these are multivariate analysis of data obtained from sensing arrays, use of spatially and temporally distributed sensors, and integration of sensors with solid-state processing technology. Two examples of important sensor applications have been also included in this issue. Ion selective electrodes are one of the oldest chemical sensors. In the past decade, some significantsalthough not always positivesdevelopments have taken place, such as the lowering of detection limits by nonequilibrium operation. The important issue of internal contact in solid-state potentiometric sensors has been one of the most significant contributions of Johan Bobacka, Ari Ivaska, and Andrzej Lewenstam. Joseph R. Stetter and Jing Li focused their review of amperometric sensors on the sensing of gases. After a rather brief overview of the general amperometric sensing principles, the authors concentrate on new nanomaterials that have been recently utilized in these devices. Somewhat neglected but very interesting semiconductor

Jirˇ´ı Janata is Georgia Research Alliance Eminent Scholar in the School of Chemistry and Biochemistry, Georgia Institute of Science and Technology. Between 1991 and 1997 he was an Associate Director of Environmental Molecular Sciences Laboratory, Pacific Northwest National Laboratory, in Richland, Washington. Prior to that appointment, he was Professor of Materials Science and Professor of Bioengineering at the University of Utah for 17 years. He was born in Czechoslovakia, where he received his Ph.D. degree in analytical chemistry from the Charles University (Prague) in 1965. His current interests include interfacial chemistry, chemical sensors, and electroanalytical chemistry with particular emphasis on development of chemical sensors for environmental and security applications.

junction gas sensors are thoroughly reviewed by Karin PotjeKamloth. With the advent of organic semiconductors, the importance of this type of chemical sensor is expected to grow, as is indicated by the very large number of references included in her review. This review should also be of interest to researchers dealing with organic electronics, such as organic field-effect transistors. Four reviews are devoted to advances in optical sensing. Colette McDonagh, Conor S. Burke, and Brian D. MacCraith emphasize new hardware that has become available thanks to rapid advances in communication and signal processing technologies. Utilization of optical sensing principles in combination with biologically derived selectivity is covered in the review by Sergey M. Borisov and Otto S. Wolfbeis. It includes selectivity derived from enzymatic, immunochemical, oligonucleotide, and organismal principles. Sensors based on plasmon resonance are discussed in the review by Jirˇ´ı Homola and in the closely related review by Matthew E. Stewart, Christopher R. Anderton, Lucas B. Thompson, Joana Maria, Stephen K. Gray, John A. Rogers, and Ralph G. Nuzzo. The latter introduces the new subject of nanostructuring.

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Cantilevers are a rapidly growing subgroup of mass sensors. That growth is fueled by the wide use of microfabricated tips for various types of scanning microscopy. The subject is reviewed by Karen M. Goeders, Jonathan S. Colton, and Lawrence A. Bottomley. A review of sensors for radionuclides has been prepared by Jay W. Grate, Oleg B. Egorov, Matthew J. O’Hara, and Timothy A. DeVol and is believed to be the first of its kind. For that reason, the criterion of “continuous data acquisition”, used to define true chemical sensors, has been somewhat relaxed, allowing some crossover to sensing systems. There is much to be gained by increasing the “order” of chemical sensors, in direct analogy with hyphenated analytical techniques such as GC-MS, MS-MS, etc. Many deficiencies in the performance of single sensors can be avoided by grouping diverse sensors into higher order arrays, as shown by Andreas Hierlemann and Ricardo Gutierrez-Osuna. The benefits of increasing the sheer number of sensors to “very large sensing arrays” are demonstrated in the review by Christopher N. LaFratta and David R. Walt. This article also highlights an elegant approach to the tricky problem of addressing individual elements in such an array. Very large scale integration of electronics is an inevitable requirement in the realization of sensing arrays. That aspect is addressed in the review by Hierlemann and is also exhaustively treated by Segyeong Joo and Richard B. Brown Rather interesting coupling of the Internet with chemical sensing is discussed in a review by Dermot Diamond, Shirley Coyle, Silvia Scarmagnani, and Jer Hayes. It clearly shows that sensors are taking over the world as “wireless sensing networks”. The extraction of distributed spatial and temporal information through multiple sensing channels and the acquisition of information from the dynamic behavior of sensing arrays are demonstrated in a review by Takamichi Nakamoto and Hiroshi Ishida. Electronic noses have been in existence for

Editorial

at least two decades. The current state of this mode of multivariate sensing is discussed by Frank Ro¨ck, Nicolae Barsan, and Udo Weimar. Development of new sensing materials is essential for the advancement of modern chemical sensors. It has been discussed to some extent in all of the above reviews. However, several contributions have been devoted exclusively to this important aspect of research. A review by Jay W. Grate focuses on the development of materials for gas and vapor sensing that are based on acidic hydrogen-bonding polymers. David W. Hatchett and Mira Josowicz treat composites of intrinsically conducting polymers with various nanomaterials. These heterogeneous sensing materials include metals, carbon-based nanomaterials, and semiconductors. A broad overview of combinatorial strategies for the development of new sensing materials is presented by Radislav A. Potyrailo and Vladimir M. Mirsky. The issue concludes with two topics of perennial interest: glucose biosensors, by Joseph Wang, and application of modern sensors to critical care medicine, by Bruce A. McKinley. Chemical sensors have achieved a rather dubious distinction. They often serve as the default mode for a material or procedure that “did not quite work out” for its original purpose or to justify results for which there is no obvious “other use”. The phrase “...and it can be also used for chemical sensing” is typical and unfortunately much too common in the chemical literature. Hopefully, the reviews compiled in this issue will change this paradigm at least to some extent. Jirˇ´ı Janata Georgia Institute of Technology CR0680991

Chem. Rev. 2008, 108, 329−351

329

Potentiometric Ion Sensors Johan Bobacka,*,† Ari Ivaska,† and Andrzej Lewenstam‡,§ Åbo Akademi University, Process Chemistry Centre, c/o Laboratory of Analytical Chemistry, Biskopsgatan 8, FI-20500 Turku-Åbo, Finland; Faculty of Material Science and Ceramics, AGH-University of Science and Technology, Al. Mickiewicza 30, PL-30059 Cracow, Poland; and Åbo Akademi University, Process Chemistry Centre, c/o Center for Process Analytical Chemistry and Sensor Technology (ProSens), Biskopsgatan 8, FI-20500 Turku-Åbo, Finland Received April 25, 2007

Contents

1. Introduction

1. Introduction 2. Theory of Potentiometric Response 2.1. Total-Equilibrium Models (Classical Models) 2.1.1. Phase-Boundary-Potential Approach 2.1.2. Total-Membrane-Potential Approach 2.1.3. Semiempirical Equations 2.1.4. Comments on Model Benefits and Drawbacks of Total-Equilibrium Models 2.2. Local-Equilibrium Models (Diffusion-Layer Models) 2.2.1. Comments on Model Benefits and Drawbacks of DLM Models 2.3. Advanced Nonequilibrium Models 2.3.1. Comments on the Benefits and Drawbacks of the Advanced Model 3. Solid-Contact ISEs 3.1. Conducting Polymers 3.2. Conducting Polymers as Solid Contact 3.2.1. Polypyrroles 3.2.2. Polythiophenes 3.2.3. Polyanilines 3.3. Conducting Polymers Dissolved in the Ion-Selective Membrane 3.4. Conducting Polymers as Sensing Membranes 3.5. Comments on the Ion-to-Electron Transduction Process 4. Miniaturized ISEs 4.1. Conventional Microelectrodes 4.2. Solid-Contact Microelectrodes 4.3. Microelectrodes in Flow Analysis 4.4. Life Science and Biology Applications 5. New Wave 6. Conclusions 7. Acknowledgments 8. References

329 329 331 331 332 332 333 333 335 335 336 338 338 340 340 340 342 342 343 344 345 345 346 347 347 347 348 348 348

* Corresponding author. E-mail: [email protected]. Phone: +358 2 215 3246. Fax: +358 2 215 4479. † Åbo Akademi University, Process Chemistry Centre, c/o Laboratory of Analytical Chemistry. ‡ AGH-University of Science and Technology. § Åbo Akademi University, Process Chemistry Centre, c/o Center for Process Analytical Chemistry and Sensor Technology (ProSens).

Potentiometric ion sensors or ion-selective electrodes (ISEs) are an important subgroup of electrochemical sensors.1-3 ISEs are characterized by small size, portability, low-energy consumption, and low cost, which are attractive features concerning practical applications. ISEs based on polymeric membranes containing neutral or charged carriers (ionophores) are available for the determination of a large number of inorganic and organic ions, as described in detail, about a decade ago, in extensive reviews.4,5 However, during the past decade, the chemical sensing abilities of ISEs have been improved to such an extent that it has resulted in a “new wave of ion-selective electrodes”.6,7 This can be attributed to several important findings, such as the considerable improvement in the lower detection limit of ISEs, new membrane materials, new sensing concepts, and deeper theoretical understanding and modeling of the potentiometric response of ISEs. The aim of this review is to highlight some of these modern topics in the field of potentiometric ion sensors. This review is focused on recent achievements since the beginning of this millenium and emphasizes the results from the last 5 years (2002-2006). Section 2 gives a critical overview, placed in a historical perspective, on the theory of the potentiometric response, including classical equilibrium models as well as advanced nonequilibrium models. Section 3 deals with recent advances in the field of solidcontact ISEs, emphasizing the application of conducting polymers as ion-to-electron transducers. Recent developments in the area of miniaturized ISEs, including aplications in flow analysis, life science, and biology, are discussed in section 4. Finally, the new wave of ISEs is commented on in section 5. We hope that the issues discussed will illustrate the great possibilities offered by modern ISEs and encourage further innovations in the rapidly expanding field of chemical sensors in the years to come.

2. Theory of Potentiometric Response The response of potentiometric ion sensors, i.e., ionselective electrodes (ISEs) and/or ion-sensitive sensors (ISSs) (i.e., sensors with solid-state contact made of, e.g., conducting polymer film), is a complex time-dependent phenomenon that depends on the electroactive material (membrane/film) and the bathing solution as well as the membrane|solution interface and their composition, thermodynamic, and kinetic properties. All these features are the subjects of the theoretical modeling of the response. Modeling in the ion-sensors area serves two roles.8 One classical role is in supporting the practitioner (the user of sensors) with very basic principles

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Johan Bobacka was born in 1962 in Larsmo, a small village on the west coast of Finland. He received his M.Sc. degree in chemical engineering in 1988 and his Ph.D. degree in analytical chemistry in 1993 from Åbo Akademi University. The Ph.D. work on “Electrochemical characteristics of poly(3-octylthiophene) film electrodes” was performed under the supervision of Professor Ari Ivaska and Professor Andrzej Lewenstam. During the years 1994−1997, Johan Bobacka worked as a research scientist at Kone Instruments Corporation (now Thermo Fisher Scientific), and since 1997, he is a Docent in electroanalytical chemistry at Åbo Akademi University. He was a senior researcher (1997−2001) and Academy of Finland Research Fellow (2001−2006) at the Laboratory of Analytical Chemistry, which is part of the Åbo Akademi Process Chemistry Centre (Finnish Centre of Excellence 2000−2011). At present, Johan Bobacka is Professor of Analytical Chemistry at Åbo Akademi University (2006−2009). His current fields of interest include electroanalytical chemistry, process analytical chemistry, chemical sensors, conducting polymers, and electrochemical impedance spectroscopy, with special emphasis on the development of potentiometric ion sensors based on conducting polymers.

Bobacka et al.

Andrzej Lewenstam obtained his Ph.D. (1977, supervisor prof. A. Hulanicki) and his DSc (1987) from Warsaw University and since 1990 he has been a Professor. Currently he is Professor of Sensor Technology at Åbo Akademi University and Director of the Center for Process Analytical Chemistry and Sensor Technology ‘ProSens’ at this University, Professor in Chemistry, AGH University of Science and Technology, Cracow, and Adjunct Professor in Chemical Engineering, Laval University, Quebec, Canada. Since 2006, he is a chairman of the Working Group on Selective Electrodes and Biosensors of the International Federation of Clinical Chemistry (IFCC). His current fields of interest are chemical sensors and biosensors, clinical chemistry, membrane potential and biomimetics, conducting polymers, hydrometallurgy, and mathematical modelling, as well as methodology of chemistry and philosophy of Science. He is the author of approximately 200 papers in chemistry and philosophy of Science and 20 patents and is a member of the editorial board of 4 international journals.

Figure 1. Methodology used in models of potentiometric ionsensors response.

Ari Ivaska was born in 1946 in Kuopio, Finland. He received his M.Sc. degree in chemical engineering in 1971 and his Ph.D. in analytical chemistry in 1975 at Åbo Akademi University, Finland. The Ph.D. thesis “Potentiometric titration of weak acids and their binary mixtures” was supervised by Professor Erkki Wa¨nninen. Ari Ivaska did his post doc in 1978/79 at Chelsea College, University of London, England, and 1982/ 83 at Northwestern University Evanston, Illinois, U.S.A. He also worked as UNESCO expert at UNICAMP in Brazil, 1980/81. In 1985/86, he worked at the research centre of Neste Company in Finland and was nominated to the Chair of Analytical Chemistry at Åbo Akademi University in 1987. He has been as visiting Professor at University of Washington, Seattle, U.S.A., in 1991/92 and 1996/97 and the spring term 2003 at University of Wollongon, Australia. He is the head of the Process Analytical Group at the Åbo Akademi Process Chemistry Centre nominated to the Centre of Excellence in research by the Academy of Finland for 2000−2011. Ari Ivaska is currently the director of the Research Institute of the Åbo Akademi University Foundation. His fields of interest are electroanalytical chemistry, process analytical chemistry, electroactive materials in general, and metal ions in the paper and pulp chemistry.

of sensor response to help the application and to support quantitative measurement by simple equations. Another adVanced role is to provide a fundamental understanding of the sensor response for those interested in electrochemical

theory as the forefront of sensor technology, by helping to map electric potential and concentration changes in space and time. Related to these two roles are a few levels of modeling generality and/or idealization, as schematically shown in Figure 1. Classical models are more idealized to (intentionally) avoid mathematical, numerical, and computational difficulties stemming from solving nonlinear equations, inherent to advanced models. Classical models are easier to comprehend and to be presented and are the subject of many papers.1-7 However, the use of advanced models is the only way to achieve a fundamental understanding (often nonintuitive) of a sensor response. The main reason is that the classical models disregard migration and, therefore, do not provide adequate space and time-dependent characteristics of sensor response, whereas the advanced models do, as shown in Figure 2. Summarized below are the recent advances in the present quantitative theory of potentiometric ion sensors in contrast to more classical approaches. In all the models considered here, the potentiometric ion sensor is represented by the following scheme: sample|ion-sensitive membrane/ film|internal contact (e.g., solution, gel, solid contact). In all the models discussed, it is assumed that the potentiometric response is modeled under open-circuit conditions (under the zero-current condition). Furthermore, the sensor is made of separate, homogeneous, ionically conducting phases that form well-defined, flat interfaces; the

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glass, and plastic), they also apply to the membranes where complexation/association processes should be taken into account, e.g., the neutral carrier-based membranes.

2.1. Total-Equilibrium Models (Classical Models) In all the classical models, the electric potential (EM) of an ion sensor is represented by the sum of a boundary potential (EPB) at the sample|ion-sensitive membrane (film) boundary (PB) and by the diffusion potential inside the membrane or film (ED). The constant added includes the potential at the internal interfaces (internal contact).

EM ) EPB + ED + constant

(1)

(For simplicity in the equations below, it is arbitrarily assumed that constant ) 0.) In the classical modeling of the response of ion sensors, two possibilities are consideredsone recognizes diffusion potential, while the other disregards it. These two approaches are presented below.

2.1.1. Phase-Boundary-Potential Approach

Figure 2. Schematic presentation of differences between totalequilibrium and advanced models: (a) concentration profiles and (b) electrical potential profiles.

interfaces are unblocked for ionic charge-transfer processes (faradaic currents), which applies to both ISEs and ISSs (see section 3.5); also, the sensor’s phases are characterized by standard chemical potentials of the components and ionic mobilities invariant in space and time. It is also assumed that the only driving forces for ion-fluxes are gradients in ion concentrations and in electric potentials (the gradients perpendicular to the sensor surfaces (1D formulation)) and that the pressure and temperature in the modeled system are constant and solvent flow (osmotic effects) is ignored. It should be mentioned that, although the models presented here refer to pure ion-exchanger membranes (solid state,

The phase-boundary-potential model is based on two idealizing assumptions: (1) The phase-boundary potential at the sample|membrane interface (phase boundary) governs the membrane response, i.e., EM ) EPB. Migration effects in the membrane are ignored, which means that the kinetic parameters of all the charged species involved, i.e., all ionic mobilities, are equal. Consequently, the diffusion potential is ignored, which formally means that ED ) 0, i.e., the electroneutrality in the membrane, except of the boundary, is assumed (the electroneutrality assumption). [On some occasions, the diffusion potential is assumed to be ED * 0 ) const (the pseudoelectroneutrality assumption) and, in this way, is disregarded.] (2) Electrochemical equilibrium is assumed at the sample|membrane interface; difference in chemical potential for any ion able to transfer the interface is balanced by a difference of the inner electrical potentials EM (EM is called the equilibrium potential). Additionally, it is assumed that the electric potentials and the concentrations of ions in the phases in contact are independent of the distance (except of the phase boundaries) and of time; there are no ion concentration drops in the respective phases over distance (the total-equilibrium assumption). The two assumptions specified above provide grounds for implementing Guggenheim’s concept of the electrochemical potential, µ˜ i,9

µ˜ i ) µi + ziFφ ) µ0i + RT ln(ai) + ziFφ

(2)

where µi is the chemical potential in the phase (µ0i under standard conditions), zi is the ion valency, ai is the single free ion activity, φ is the electric (inner) potential in the phase, and R, T, and F are the universal gas constant, the absolute temperature, and the Faraday constant, respectively. Using further idealizing assumptions, namely, that only an ion “i” can transfer through the interface (the ideal permand ion-selectiVity assumptions); the ion transfer is fast and reversible (the infinite kinetics assumption); the phases in

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contact are not only of distinctly different chemical properties but are immiscible (the ideal immiscibility assumption); the convention for a single ion activity instead of a mean ionic activity (e.g., by using Debye-Hu¨ckel theory and convention) is adopted and it is assumed that the single ion activity (ai) in each phase is equal to its concentration (ci) (the ideal phase assumption); and there is no flux of solvent through the membrane/film (the solVent impermeability assumption), it is possible to employ Guggenheim’s concept to derive EPB as a function of ci. By using eq 2 for each phase, for the condition of electrochemical equilibrium between the phases (µ˜ i ) µ˜ i) and by recognizing that they are chemically distinctively different (µ0i * µj 0i ), after separation of electrical and chemical terms, one immediately arrives at a Nernstian-like equation for the electrical potential difference at the ionsensor interface,

RT RT ci EM ) EPB ) ln ki + ln ziF ziF cji

(3)

where ci denotes the concentration of an ion i in the solutions (in mol/dm3) and the barred symbol denotes the membrane/ film phase, and ki is an ion-partition constant defined as ki ) exp{(µ0i - µ j 0i )/RT}, where µ0i and µj 0i are standard chemical potentials of ion i in the respective phases. The theoretical models of the ion-sensor responses exploiting Guggenheim’s PB concept started with the paper by Nikolskii.10 The author considered the contact of a pHsensitive glass membrane with the bathing solution containing the main ions (i ) H+) and the monovalent interfering ions (j ) Na+) and obtained the equation for ion-sensor response in the form

EM ) EPB )

RT RT ci RT ln ki + ln ) ln ki + ziF ziF cji ziF

( )

cicjj cji RT ci(cji + cjj) RT RT ln ln ki + ln ) ) ziF cji(cji + cjj) ziF ziF cji + cjj RT ci + Ki,jcj RT RT ln ki + ln ln(ci + Ki,jcj) ) const + ziF ziF cji + cjj ziF (4)

(

(

)

ci +

)

where Ki,j used in the derivation above is the equilibrium constant

Ki,j )

kj cicjj ) ki cjcji

(5)

for the ion-exchange reaction,

i+ + j + a i + + j +

(6)

and cji + cjj ) const (the bar means the ions are in the membrane phase). Nikolskii was aware that including potentials arising from the different mobilities of ions in the membrane is a complex mathematical task, which was completed three decades later by Eisenman.11

He considered different membrane types, i.e., solid-state, glass, and membranes containing ion-exchangers and neutral ligands. He considered cases with and without association between ions in the membrane. Eisenman’s modeling made under classical assumptions of a total equilibrium yielded, for most cases considered, the equations in the form first obtained by Nikolskii. For instance, the equation for the fully dissociated ion-exchanger case was derived in the form

EM ) EPB + ED )

( )(

) )]

ujj RT RT ci RT ln ki + ln + ln jci + cjj ) ziF ziF cji ziF uji ujj ci + Ki,jcj RT RT ln ki + ln cji + cjj ) ziF ziF cji + cjj uji ujj RT ln ci + Ki,jcj (7) const′ + ziF uji

[(

(

)

where uji and ujj are the ion mobilities for i+ and j+ ions, respectively. If uji ) ujj, then eq 7 is identical to eq 4. As is shown above, both the Nikolskii and the Eisenman models provided strict analytical derivations of equations, possible for equal charges of the main and interfering ions (e.g., 1:1, 2:2). These models were later applied on many occasions to describe the response all kinds of ion sensors (for review, see refs 4 and 5). Analytical derivation of the above equations for unequal charges (zi * zj) in the frame of the total-equilibrium approach is impossible.12-15 To cover such cases, a caseby-case approach for each nonequal charges is used and relatively complex formalisms with implicit methods of equation solving are employed. They are based on additional ad hoc assumptions, such as ignoring the changes in concentration of the main ion in the membrane. Some ad hoc formal stratagems are employed as well, e.g., using the same mathematical equation to bind two independent variables depicting the primary ion, one representing the primary ion activity in the sample without interference from the other sample ions and the other representing the primary ion activity in the mixed sample.4,5,16 In this situation, to cover the cases of unequal charges while keeping the totalequilibrium assumption valid, the semiempirical equations, in a similar form to eqs 4 and 7, were postulated, as shown below.

2.1.3. Semiempirical Equations The semiempirical equations reflect the emphasis on the practical applications of ion-selective membrane electrodes. As a result of the arbitral decision of IUPAC, it was postulated to merge and extend the previously mentioned eqs 4 or 7 in the form of a general equation for all ionselective electrodes and ion sensors, covering all charges of ions. This equation is known today as the NikolskiiEisenman (NE) equation (for simplicity given below for one interfering ion).17

EM ) const′ +

RT pot zi/zj log(ci + Ki,j cj ) ziF

(8)

2.1.2. Total-Membrane-Potential Approach Eisenman abrogated the idealizing assumption that ED ) 0 and extended Nikolskii’s eq 4 for ED * 0 in the membrane.

This equation can be further extended to include the low detection limit (L) of the ion sensor:18

Potentiometric Ion Sensors

EM ) const′ +

Chemical Reviews, 2008, Vol. 108, No. 2 333

RT pot zi/zj log(ci + Ki,j cj + L) ziF

(9)

pot Both parameters Ki,j and L are constitutive for analytical practice and theoretical modeling. Theoretical modeling attempts to help predict their values, and, if it occurs, to show the evolution of the response for different concentrations of main vs interfering ions, over a period of time.

2.1.4. Comments on Model Benefits and Drawbacks of Total-Equilibrium Models The predominant way of modeling under total-equilibrium PB models is to ignore the diffusion potential without any formal (mathematical) justification, despite the reports on its significance even under total equilibrium.15,19,20 A typical way of ignoring ED in the PB models is formulated verbally, not mathematically, e.g., by saying that “by the end of the 1980s, it started to become clear that subtle effects of ionic mobilities in ISE membranes may be typically ignored without a significant loss in accuracy”.21 What is overlooked in the argument just cited is the fact that ignoring diffusion potential means arbitrarily placing and affixing the membrane potential at the phase boundary only andsat the same times justifying an uncritical use of the boundary-potential concept. Interestingly, this practice remains in contrast with wellbalanced criticism of the phase-boundary approach by the inventors of this concept. Namely, according to Guggenheim, “the conception of splitting the electrochemical potential µ˜ i of an ion i into the sum of a chemical term µi and an electrical term ziFφ has no physical significance; for one can assign an arbitrary value to φ for some point in each medium and this will for the ions of each type i determine µi, so as to give µ˜ i, the value which determines all the physical processes involving ions of type i”.9 Nikolskii added to Guggenheim by saying that his model “involves some difficulties, for in this case one deals with thermodynamically undefined variables, interface potential differences, diffusion potentials, and the activities of single ions”.22 The advantage of the total-equilibrium PB models lies in their simplicity. Characteristic of the phase-boundary modeling is that it considers ion complexion/association equilibria in a membrane in a similar way to solution chemistry and to coupling the membrane chemistry with that of the bathing solution via a pivotal phase boundary, eq 3. This methodology provides similar flexibility, as known from the complex formation theory, in considering many ion-equilibria and complex formation constants, the mass conservation, and the charge balance (electroneutrality) equations. Therefore, the methodology used in the total-equilibrium modeling has value: delivering the basic theoretical support and formal instrumentation needed for the practical applications of sensors, while offering an intuitive but essential understanding of the principles underlying sensor response. For these reasons, this modeling has been used extensively to provide a simple semiquantitative description of most ionselective electrode related experiments and has assisted every period of development in the ion-sensor technology. The modeling was, and still is, applied in ion-selective electrodes4,5 and conducting polymer ion sensors, including the sensors with a solid contact.23-25 The major drawbacks of the total-equilibrium models relate directly to the idealizations used; i.e., to the counterfactual assumption that all individual mobilities in the membrane are equal. This allows the migration effects (ED ) 0) to be

ignored and the assumption that (i) the electroneutrality condition is valid even in the proximity of the phase boundary, (ii) the concentrations of primary and interfering (and their complexes) as well as oppositely charged sites in the layers adjacent to the membrane/film surface are equal to those in the respective bulks, and (iii) the ion-transfer rates are infinite (see Figure 2). Primarily, assuming that the systems modeled are at total equilibrium leaves us with the sensor response interpretation that the electric potential, sensitivity, and, in particular, selectivity coefficients and detection limit are time-independent. The consideration in the total-equilibrium models that the sensor response is time-independent contradicts numerous time-dependent empirical reports in the field of potentiometric sensors, especially those delivered by the nonequilibrium potentiometry, such as the lowering of the detection limit by using transmembrane fluxes, which are timedependent.

2.2. Local-Equilibrium Models (Diffusion-Layer Models) In these models, called the diffusion-layer models (DLMs), the local equilibrium at the sensor interfaces is assumed (the local equilibrium assumption). (It means that eqs 3 and 7 as well as eqs 8 and 9 apply by substitution of the bulk concentration of ions by respective surface concentrations.) Additionally, it is assumed that the concentration of ions in the membrane phase and contacting phases are dependent on the distance but are independent of time, i.e., the sensor system is at steady state, or are dependent on time by diffusion of ions to/from membrane|solution interface(s) controlling the equilibration rate. The need to extend the time-independent thermodynamic modeling as described above was already noticed some decades ago, owing to the observations of variable and timedependent selectivity in the case of ion sensors with solidstate and plastic membranes.26-28 The problem of selectivity changes with time is still an issue of significant interest in the area of bio-ion-sensitive membranes/films.29 Sokalski et al.’s recent discovery of lowering the detection limit due to transmembrane fluxes gave an impetus to the consideration of ion fluxes30 and resulting concentration gradients in the theoretical modeling using the DLM frame. The diffusion-layer model (DLM) was first introduced by Lewenstam,31 and was continued in a number of papers,28,31-37 to model changes of selectivity coefficients with concentrations and time. This model is based on the assumption of local equilibrium at the solution|membrane interface, and consequently, a starting point in DLM are eqs 3 and 4 with surface (local) concentrations instead of bulk concentrations. The model assumes steady-state ion fluxes given by linear concentration gradients between the interface(s) and respective bulks, and it assumes constant and time-independent diffusion layers in the “local” areas. In DLM, the time-dependent response (for ions of equal charges and an ignored detection limit) is obtained by using a pivotal parameter “s(t)” called the surface coverage or site filling factor, characterizing the distance of the system under local equilibrium from total equilibrium over time (t), which is defined as29,33,35-37

s(t) )

cj0(t) ci0(t) + cj0(t)

)

Ki,jcj0(t) ci0(t) + Ki,jcj0(t)

(10)

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Here, the barred concentrations refer to ion concentrations in the membrane or film surface at time t, and the sum ci0(t) + cj0(t) is equal to cib, where cib is the concentration of the main ion in the bulk of the membrane. The concentrations without bars ci0(t) and cj0(t) are the solution ion concentrations on the membrane surface at time t. Assuming a constant diffusion-layer thickness with linear concentration changes in the diffusion layers (the Nernst approach) for any response time, and by coupling the ion fluxes (Ji,Jj) from/to the interface by virtue of mass conservation, one obtains

Jj )

ntot ds(t) cj - cj0(t) ci0(t) - ci ) Dj ) Di ) -Ji (11) 3 A dt 10 δ 103δ

where A is the electrode surface area (m2), ntot is the number of active exchangeable sites occupied by the i and j ions (mol) on the membrane surface, and δ is the diffusion-layer thickness (m). By combining eqs 5, 7, and 10 rewritten in local concentrations and eq 11, one further obtains

EM(t) ) const +

(

)

uj [1 - s(t)]Ki,j + s(t) Ki,j c + Dj c i ui Di j RT ln (12) Dj F c ib Ki,j[1 - s(t)] + s(t) Di where the const is the term including contributions independent of the concentration of the ions i and j; Di and Dj are the diffusion coefficients of the primary ion and interfering ion in the aqueous diffusion layer (m2 s-l), respectively; and uji and ujj represent the ionic mobilities of ions in the membrane phase (m2 s-1 V-1). After separation of the variables s(t) and t and integrating within time limits, t to t )+∞, and corresponding s limits, s(t) and seq, it is possible to obtain a function binding s(t) and t,

Dj ‚c Dj Di j s(t) Ki,j s(t) - Ki,j ln 1 ) Di ci + Ki,jcj seq (ci + Ki,jcj)Ct (13)

[

]

ci +

(

)

with C ) DjA/ntotδ × 10-3 and where t is a time (s), 10-3 is a unit conversion factor (dm3 to m3), and seq is given by

seq )

cjj0(t f ∞) cji0(t f ∞) + cjj0(t f ∞)

)

Ki,jcj ci + Ki,jcj

(14)

where cji0(t f ∞) and cjj0(t f ∞) represent the ion concentration in the membrane phase at the interface at steady state and/or at total equilibrium (for both cases, ds(t)/dt ) 0). In this way, in the DLM, potential is connected to time. The DLM, in contrast to total-equilibrium models, contains time as a model parameter. The time is introduced ad hoc to describe attaining of total equilibrium (i.e., equilibration) via diffusion-controlled ion transport. The DLM predicts that pot changes during the equilithe selectivity coefficient Ki,j bration process (i.e., in the measurement) as a function of s(t) and, thus, with time (t).33-37 According to DLM, the value

pot of the selectivity coefficient Ki,j can vary between two limiting valuessone for a short response time dictated by the ion-transport properties in solution,

pot Ki,j (t f 0) )

Dj ≈1 Di

(15)

and the other at steady state and/or total equilibrium, where pot Ki,j is characterized exclusively by the membrane-related parameters and expressions known from total-equilibrium models, eqs 4 and 7. In the DLM, eqs 4 and 7 are limiting cases for t f ∞: pot (t f ∞) ) Ki,j (for ED ) 0) Ki,j

ujj pot or Ki,j (t f ∞) ) Ki,j (for ED * 0) (16) uji

The above prediction given by eq 15 is of great practical importance in applications of ion sensors because it predicts that the sensor response for short readout times is characterized not by the equilibrium selectivity as given by eq 16 but by ion transport and the selectivity given by eq 15. In consequence, if it occurs, the electrode senses “equally” the main and interfering ions. This prediction resulted in important practical benefits that are also of interest today. It was, for instance, used to kinetically discriminate strong interferences by short readout times (t f 0)38,39 or, alternatively, to benefit from the measurements of strong interferents, such as heparin on chloride ISE. The latter is realized pot ≈ 1 and using the response by taking advantage of KCl,heparin to the interfering (heparin) translated “1:1” into the signal of the main (chloride) ions.40,41 DLM was also used to interpret nonmonotonic transients in ion-sensor response36 and for interpretation of long-term drifts in sensors in which a thin aqueous layer is formed between the membrane and the substrate electrode.42 DLM was used successfully to demonstrate that the poor apparent selectivities and detection limits have a common origin in the increased (vs bulk) surface concentrations of the main ion.43-46 In consequence, it was shown for the first time that, using the ion sensors in the regime of nonequilibrium response, induced as a result of the ion-complexation processes, both true (unbiased) selectivity coefficients44,46 and much lower detection limits for solid-state membranes can be achieved.45 In 1999, Sokalski et al.47 used the DLM frame to interpret the effect of lowering of the detection limit for plastic membranes by analyzing transmembrane ion fluxes in a symmetric solution/membrane/(internal) solution system. In this formulation of DLM, the diffusion potential is ignored, the fluxes of ions are treated under steady state, and the concentrations of all species are assumed to change linearly within the diffusion layers, i.e., in the adjacent solution layer and over the membrane. Ion-exchange as well as coextraction processes are considered to analyze the low detection limits. The model in this form provides a possibility to analytically find an equation for the detection limit (L) vs different model parameters, especially the concentrations of ions in the internal solution and membrane, but not vs time. An important outcome of this interpretation is reflected in the equation that allows the calculation of the steady-state surface concentration of the main ion, which dictates the low detection limit for (plastic) membranes:

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ci0(t f ∞) ) ci +

D h iδ [cj (t f ∞) - cjid(t f ∞)] Did i0

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(17)

The above equation indicates that adjusting the composition of the internal solution (i.e., by keeping low free main ion concentration due to ion complexation and optimizing vs the membrane selectivity, the interfering ion concentration in the internal solution or the solid-contact film), the membrane thickness (d), the diffusion coefficient of in the membrane (D h i), the thickness of the diffusion layer in the sample (δ), and the diffusion coefficient in the sample (Di) plays an essential role in forming an ISS response in the concentration range close to the low detection limit, as shown schematically in Figure 3. Predictions from this version of DLM are better visualized by the graphical representations of the ISS responses instead of using rather complicated algebraic equations (see Figure 4). In fact, the application of a simple commercial equation solver (e.g., Mathcad) and its graphical tools proved to be quite feasible.47 The observation of Sokalski et al. on the lowering of the ISS detection limit,30 and its mathematical modeling by DLM, heralded a new chapter in exploring the power of potentiometric sensors under local equilibrium, as proved by a number of papers (for recent review, see ref 48).

Figure 3. Schematic presentation of the concentration profile of the main ion influencing the lower detection limit.

2.2.1. Comments on Model Benefits and Drawbacks of DLM Models The DLM allows the sensor response to be theoretically modeled in a way that was not possible for total-equilibrium models; it allows modeling of the variation in the selectivity vs time and the effect of lowering the detection limit under steady state to be interpreted. From the theoretical point of view, the advantages of DLM-type modeling do not compensate for the drawbacks already known in the total-equilibrium models, as explained above. However, by extending the power of the totalequilibrium concepts, DLM, in a relatively simple way, can truly support practical advances in nonequilibrium potentiometry, since it exploits close-to-the membrane and transmembrane fluxes for comprehensive analytical benefits. One possible application of DLM is by assuming the invalidity of steady state for concentration (∂ci(x,t)/∂t * 0) and employment of the second Fick’s law for ions treated as neutral species, as shown very recently.49 The main problem of the nonequilibrium potentiometry is in finding a “niche of stability” for the reproducible analytical readouts in a variety of matrixes and in time. The models presented so far are overidealized and not able to answer the basic questions, especially important in the case of lowering the detection limit: what is the role of the membrane, its thickness and dielectric constant, in shaping the sensor response? Similarly, how are other important questions to be answered, such as the following: How do the different diffusibilities of ions and ion-exchange rates influence the signal? When can the subtle effects of ionic mobilities in ISE membranes be ignored without a significant loss in accuracy? What are the bulk-to-bulk concentration and potential changes over time? Finally, how can the signal be stabilized if the sensor works in a nonequilibrium regime in different matrixes? All these questions call for a new “upper floor” approach in theoretical modeling, namely, for the advanced nonequilibrium models, which are described below.

Figure 4. Calculated EMF functions for a series of ISEs having the same membrane but different primary ion concentrations, ci in the internal solution, from 1 to 10-15 M in the inner solution. Curves are labeled with the corresponding negative logarithm.47

2.3. Advanced Nonequilibrium Models A fundamental difference between the advanced nonequilibrium models and the (total or local) equilibrium models is in abandoning of two constitutive conditions, the electroneutrality and (total or local) equilibrium condition. It is possible to do so, owing to the admission of charge separation from the Poisson equation and the finite ion-transfer rates described by the appropriate heterogeneous ion-transfer rate constants. Modeling of the nonequilibrium potentiometric response of the ion-sensitive sensors requires employment of explicit space and time domains, which provide the platform for a set of relevant thermodynamic, extrathermodynamic, and kinetic data. This platform may be constructed by use of the Nernst-Planck and Poisson equation system (NPP)50 or by a similar system that is an appropriate and rich enough tool to encompass the ion-sensor response. The first implementation of the NPP to model nonequilibrium (non-steadystate) response of ion sensors was recently reported by Sokalski and Lewenstam51 and in the contributions that followed.52-54 The NPP equation system allows calculations of the electric potential difference and concentration profiles as a function of space and time for the conventional sensors with internal solution or for the sensors with solid internal contacts. In contrast to the total-equilibrium and diffusionlayer models, the NPP model does not require any arbitral splits of the membrane potential into phase boundaries and diffusion-potential terms (compare Figure 2). Additionally, it does not use idealizing assumptions of the total and/or local equilibrium and the electroneutrality conditions. The

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NPP model applies for different membrane and film-based ISSs and for ions of every charge, without the need for a case-by-case approach as known from the previous models, described earlier.4,13 In the NPP model, multilayers and adjacent (diffusion) layers are admitted. The membrane and the bathing solution(s), or films in contact, may have a thickness and dielectric permittivity of choice. The membrane may contain unequally mobile and/or immobile charged and uncharged sites/ligands. Ions of any charge can be considered in the charge-transfer and transport processes. Different degrees of association between extracted ions and sites and ligands in the membrane are permissible. Furthermore, the gradient of the chemical potential of the solvent across the membrane and convective flows are allowed. In the present formulations of the NPP model, all activity coefficients in the membrane are assumed to be one, ion extraction is described by the first-order rate constants, and the 1D geometry is used. The core of the NPP model is that ion fluxes in space (x) and time (t) are described by the Nernst-Planck equation,

[

fi(x,t) ) -Di

∂ci(x,t) F - zici(x,t) E(x,t) ∂x RT

( )

]

(18)

where fi(x,t) is the flux of the ith ion, ci(x,t) is the concentration of the ith ion in space point x and time t, E(x,t) is the electric field, Di is the diffusion coefficient of the ith ion, and zi is the charge of the ith ion, as in the models discussed above. In order to solve eq 18, two additional equations are used to relate fi(x,t), ci(x,t), and E(x,t). The first choice is the law of mass conservation:

∂fi(x,t) ∂ci(x,t) )∂t ∂x

(19)

The second is the Poisson equation, rewritten for the total current density (I) as

∂E(x,t)

I(t) ) F‚∑ zi‚fi(x,t) + ‚ i

(20) ∂t

where  is the dielectric permittivity. In the calculations, the Chang-Jaffe boundary conditions are used,

fi0(t) ) B k i‚ci,bL - A k i‚ci0(t) fid(t) ) - B k i‚ci,bR + A k i‚cid(t)

(21)

where fi0, fid, ci0, and cid are the fluxes and concentrations at x ) 0 and x ) d (where d is membrane thickness), respectively; B ki and A ki are the forward and backward rate constants, respectively, and their ratio is a partition coefficient; and ci,bL and ci,bR are the concentrations in the bathing solutions on the left (L) and right side (R) of the membrane, respectively. The system of nonlinear partial differential equations 18, 19, and 20 is solved to obtain the resolution in space and time by the finite-difference51-54 or finite-element method.55 Some authors stress the complexity of these procedures.21 We do not share this opinion; to the contrary, the imple-

Figure 5. Time-dependent concentration profiles for site R- ([i+] ) 10-4, [j+] ) 10-3, D h i/D h j ) 0.5, Ki,j ) 0.1, and R h TOTAL ) 10-3). Curves a-g show profiles after the following: (a) 4 × 10-4, (b) 1.64, (c) 13.1, (d) 26.2, (e) 104.8, (f) 420, and (g) 13 440 s (steady state).50

mentation and access are simple when using appropriate commercial platforms.

2.3.1. Comments on the Benefits and Drawbacks of the Advanced Model The total-equilibrium and diffusion-layer models are merely special cases (concretizations) of the NPP model. In a strict formal (mathematical) sense, all that is predicted by the PBMs and DLMs can be obtained from the NPP model by deduction (as shown in Figure 11 and discussed below). Of course, in simple, theoretical cases and routine laboratory practice, the full use of NPP is not necessary. (While admitting this, it should be emphasized that the argument repeatedly given by some authors in favor of phase-boundary models (PBMs),21 namely, that, in the NPP, the knowledge of individual mobilities of relevant ions and their transfer rates at the phase boundary is required, is not in contrast to the phase-boundary models. Actually, in the PB models, as was shown above, even stronger assumptions are used, i.e., that the rate constants are infinite and all individual mobilities are equal.) However, for modeling of time-dependent processes, in cases of suspected nonlinearities, and for someone working with sensors under nonequilibrium, as happens in the case of fast readout times or in lowering the detection limit, the NPP approach or a similar theoretical tool is a must. The reason is simple: as the model from the “upper floor”, the NPP model provides hard numerical facts instead of verbal declarations and unproven intuitions, so often made by advocates of simple modeling. The NPP approach has already provided novel insights and numerical answers to problems in ISE practice that have been intriguing for decades and which the DLM and PB models were unable to describe: the concentration and potential profiles over equilibration time and even under steady state (or equilibrium) revealed striking nonlinearities (see Figure 5), shown for R-, which in the PB and DLM models is assumed to be not dependent on distance (x) and time (t). It was presented that the contribution of the socalled diffusion potential prior to equilibrium is significant and varies with time (Figure 6), and in consequence, prediction of NPP and PB models for the same conditions can be strikingly different (Figure 7): In fact, the overall linearity of the calibration curves depends on the distance from the steady state or equilibrium in the sensor system, and even under steady state may be significantly influenced by migrational effects (Figure 8). (For simplicity, these

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Figure 6. Time-dependent and distance-dependent potential profiles, calculated as φ(x,t) ) ∫ E(x,t) dx. Total membrane potential b,R E(x,t) dx, where xb,L and xb,R are the points in the is EM(x,t) ) ∫xxb,L bulk of bathing sample solution (on the left, L) and internal solution/ film (on the right, R). Curves a-g show profiles after the following: (a) 4 × 10-4, (b) 1.64, (c) 13.1, (d) 26.2, (e) 104.8, (f) 420, and (g) 13 440 s (steady state).52

Figure 7. Comparison of the (a) steady-state NPP potential profile from Figure 6 with that presumed by (b) the phase-boundary model.52

effects are shown here schematically; for numerical details, see refs 52 and 53.) Furthermore, the NPP model shows unequivocally that the concentration profile of the species confined in the membrane phase (the oppositely charged sites) can dramatically change over the membrane distance along with highly nonlinear (and thus “nonintuitive”) profiles. For this reason, a strict electroneutrality condition, represented by the distance-independent charge-balance equation, used in PB models16,21 is expedient to allow binding of opposite charges in the membrane boundary but is clearly inadequate.53,55 Recent results from use of the NPP model allowed the analysis of concentration profiles with time for membranes bathed by strongly interfering ions (a chloride-sensitive membrane bathed with perchlorates),53 as illustrated in Figure 9, where the theoretical prediction and experimental results are compared. In the same work, selectivity changes as a function of time were examined, which had been inaccessible for the totalequilibrium model (PB model) and only partially accessible to the diffusion-layer model (DLM). Moreover, the access to spatial distributions of ions and the electric potential vs distance during equilibration time allows inspection of the underlying reasons for selectivity coefficient changes and their magnitudes.53,54 A similar methodology was used in the numerical analysis of the detection limit due to the transmembrane ion fluxes,

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Figure 8. Calibration curves in the presence of primary ([i+] ) (10-10 to 10-1 M)) and interfering ([j+] ) 10-3 M) ions calculated according to the NPP model (other data as in Figure 5). Pie charts express contributions of phase-boundary potential (EPB) (white) and inner membrane (diffusion) potential (Ed) (black) for the outermost calibration curves (Di/Dj ) 0.1 or 10) at [i+] ) 10-4 and [i+] ) 10-10 M, Ki,j ) 0.1 The curves correspond to the following diffusion coefficient rations (D h i/D h j): (a) 0.1, (b) 0.5, (c) 1, (d) 2, and (e) 10.52

to find the parameters important for its evolution and stabilization over a period of time. The NPP model allows numerical analysis of the important role of transmembrane ion fluxessas well as the appearance of the electric potential drops at one interface (sample|membrane) or at two interfaces (sample|membrane and membrane|internal contact interfaces)s in shaping the effect of the lowering of the detection limit.55,56 Consequently, for the first time, the significant influence of ion transport (diffusion), distribution and rate parameters, ion charges, dielectric constant, and thickness of the membrane are predicted. The NPP approach provides realistic concentration profiles as a function of time, which previously could only be addressed by the DLM, assuming linear concentration drops (compare Figure 3 and Figure 10a). In addition, the NPP model provides the profile of the electric potential in space and time (Figure 10b), which, according to the DLM (being a steady-state model), cannot be modeled. There is no doubt that enhanced sensitivity of measurements under nonequilibrium may be paid for by a decreased reliability of the results. Thus, an important and so far untouched question concerning the influence of the diffusion potential and its variation from sample-to-sample and over readout time, which may unfavorably and in an uncontrolled manner influence the precision of measurements with ISSs, can now be considered. Demonstrably stable and reproducible measurements have not yet been convincingly achieved. However, via the nonequilibrium model (NPP), a comprehensive analysis of this problem will soon be presented.56 The inevitable conclusion is that the NPP offers a novel tool for solving a number of, up until now, inaccessible problems in nonequilibrium potentiometry. The NPP is, by far, more general than the DLM and PB models, with the latter being special cases of the NPP, as shown in Figure 11. The relationship between these potentiometric models is characteristic of empirical sciences, for instance, relativistic mechanics and its special case, classical mechanics (even if we have ample confirmation from everyday routines in contrast with relativistic theory). For

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Figure 9. Theoretical and experimental calibration curves for ionexchanger electrode chloride ISS.53 (a) Calibration curves calculated for membrane conditioned in i- and j- for solutions of (1) j- (5 min), (2) i- (5 min), (3) i- (15 min), (4) i- (30 min), and (5) i(60 min), and (6) calibration curve obtained with a membrane conditioned in i- for a solution of i- (60 min). i is the preferred ion, and j is the discriminated ion. The inset shows the EMF time dependence for an ISE conditioned in discriminated ion for 10-4, 10-3, 10-2, and 10-1 M, in order from top to bottom. (b) Calibration curves obtained with a membrane conditioned in Cl- for solutions of ()) Cl- (5 min/conc.), (0) ClO4- (5 min/conc.), (∆) ClO4- (15 min/conc.), (3) ClO4- (30 min/conc.). Calibration curve obtained with a membrane conditioned in ClO4- for a solution of (O) ClO4(60 min/conc). ClO4- is the preferred ion, and Cl- is the discriminated ion. The inset shows the EMF time dependence for an ISE conditioned in Cl- for 10-4, 10-3, 10-2, and 10-1 M in order from top to bottom.

this reason, the theoretical value of the NPP is unquestionable. To summarize, there is currently no doubt that, according to a hierarchy of the generality of the models discussed, the NPP approach provides clear added value, with regard to both nonequilibrium response and characteristics of the steady state of potentiometric ion sensors. The NPP also allows assessment of the migrational effects, finite kinetics, and permittivity of the membrane, none of which can be accessed via more idealized models because of their limited dictionaries and formalisms, as shown in Table 1. Additionally, Table 1 shows that, although the NPP itself has its own intrinsic idealizations and limitations and shares common idealizing assumptions and parameters with the DLM and NE, it is simply more powerful. This again reflects the process of the cumulation of knowledge, which is wellknown for empirical sciences.8 A new chapter in modeling on the level of the generality dictated by the NPP has now been breached, and questions insurmountable for the simpler DLM and PB models can now be answered. There is a list of problems concerning ISE/ISSs to be addressed, as well as many technical possibilities to be investigated. In 2D and 3D, modeling of

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fluxes through nanotubes and biological channels is still under development. Assumptions such as those listed above, i.e., equating activities and concentrations, uncoupled fluxes, complex-forming processes in the membrane, higher-order rate equations, influence of the electric field on ion-transfer rates, e.g., by extending the Chang-Jaffe boundary conditions in a form of Butler-Volmer conditions, must all be considered. (The latter condition admits the influence of the electric field on rate constants; although this is formally and programwise a simple extension as will be shown shortly,54 the problem of the ion-transfer mechanisms is still controversial and open.57-59) Furthermore, the variability of the dielectric constant along the membrane distance and at the interfaces needs to be considered, and finally, increasing the library of transport parameters in the membranes and films relevant to sensor technology needs to be recognized.60 These are just some examples of the interesting problems requiring further theoretical work on the assumed “level” of approximation and experimental verification.55,56 Present applications of ion sensors in the nonequilibrium regime, and the need for proper theoretical support, show that understanding of the response mechanism of potentiometric sensors is a new and open challenge. Without ignoring the achievements and power of the earlier classical or diffusion-layer models, a call for advanced modeling is nonetheless unquestionable.

3. Solid-Contact ISEs Elimination of the internal filling solution from conventional ISEs results in solid-contact ISEs (ISSs), which are more durable and easier to miniaturize than their conventional counterparts. However, in order to obtain solid-contact ISEs with stable electrode potential, it is necessary to have sufficiently fast and reversible ion-to-electron transduction in the solid state without any contribution from parasitic side reactions.61 Research and development of solid-contact ISEs had already started in the beginning of the 1970s with the invention of the coated-wire electrode (CWE), which indeed represents a simple and robust design.62 The main drawback of the CWE is obviously the poor potential stability resulting from the blocked interface that forms between the purely electronic conductor (metal) and the purely ionic conductor (ion-selective membrane). Solid-contact ISEs with improved potential stability have, therefore, been produced by utilizing electroactive materials showing mixed electronic and ionic conductivity that serve as ion-to-electron transducers between the electronic conductor and the ion-selective membrane.61 Among the electroactive materials available today, electroactive conjugated polymers (conducting polymers) have emerged as one of the most promising ion-to-electron transducers for solid-contact ISEs.63-66 Other approaches to solid-contact ISEs involve the use of Ag/AgCl,67 Ag/AgCl/hydrogel,68 redox-active self-assembled monolayers,69,70 Prussian Blue,71 carbon-based composites,72,73 Ag-based composites,74 and Ag/AgCl/porous carbon loaded with ionophore and plasticizer that resulted in a solidcontact Pb2+-ISE with an impressive lower limit of detection below 10 pM.75

3.1. Conducting Polymers The discovery and development of conjugated polymers that can be made electronically conducting by partial oxidation (p-doping) or reduction (n-doping) has had a great

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Figure 10. Calculated by NPP, the main ion concentration (a) and the electrical potential changes in space and time (b).54 ciL ) 10-9, ciR ) 10-7, cjL ) 10-7, and cjR ) 10-1; cji ) cjR- ) 10-3 (mol/dm3); Di ) 10-9 and Dj ) 10-8 and D h i ) 10-11 and D h j ) 10-10 (m2/s); khi,l ) -7 kCi,l ) 1; khi,l ) 10 and kCi,l ) 1 (m/s), L and R denote the left and right sides of the membrane. Diffusion potential is calculated at 1 µm distance from the interfaces in the interior of the membrane.

impact on several fields of science and technology, which naturally goes far beyond the scope of this review.76-84 However, it should be mentioned here that conducting polymers have been applied in a large variety of chemical sensors and several reviews have been published on this topic.63-66,85-114 A few of these reviews focus entirely on the application of conducting polymers in potentiometric sensors.63-66,92,97,114 Although conducting polymers have been known in the field of potentiometric ion sensors since the 1980s, development is still continuing. Conducting polymers have some key features that are useful when applied as ion-to-electron transducers in solidcontact ISEs. Conducting polymers are electronically conducting materials that can form an ohmic contact to materials with a high work function, such as carbon, gold, and platinum, which ensures a proper electronic (ohmic) contact. Conducting polymers can be deposited on the electronic conductor by electropolymerization of the monomer or by solution-casting of the soluble conducting polymer, which gives some flexibility in the manufacturing process. Conducting polymers are electroactive materials with mixed

electronic and ionic conductivity, which means that they can transduce an ionic signal into an electronic one in the solid state. Furthermore, the properties of conducting polymers can be tailored via functionalization, e.g., by covalent bonding of side groups to the conjugated polymer backbone and by immobilization of functional doping ions. These are important features that make conducting polymers suitable as solid contacts in combination with conventional ionselective membranes.63 In this type of solid-contact ISE, where the conducting polymer is coated with a conventional ion-selective membrane, the ion-selectivity is determined mainly by the ionselective membrane, which allows the utilization of various ionophore-based polymeric membrane formulations that are available.4,5 Recent progress in the area of ISEs such as the lowering of the detection limit toward the picomolar level,30 enabling potentiometric trace-level analysis,6 gives an additional impetus for the development of solid-contact ISEs with improved analytical performance. Conducting polymerbased solid-contact materials are boosting “the new wave of ion-selective electrodes”.7

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Figure 11. Interrelation between models in potentiometry. (*) The condition reads as follows: there is no change in electric field strength and concentration of ions over time in the system sample/ membrane/internal contact (solution or film). (**) The condition reads as follows: there is total equilibrium over distance in the system sample/membrane/internal contact (solution) or film, i.e., electrochemical potential of each special along x is constant; additionally, electrical potential over distance in each phase, and in consequent concentrations of each species, are contact except the phase boundaries.

3.2. Conducting Polymers as Solid Contact Over the past few years, the analytical performance of ISEs where conducting polymers are used as ion-to-electron transducers (solid-contact ISEs) has been dramatically improved. Solid-contact ISEs for determination of both inorganic and organic ions were constructed by using conducting polymers as the ion-to-electron transducers. Of particular interest is the lowering of the detection limit down to the nanomolar level. The conducting polymer materials were based mainly on electropolymerized or chemically polymerized derivatives of pyrrole (Figure 12), thiophene (Figure 13), and aniline (Figure 14).

3.2.1. Polypyrroles Polypyrrole was already used as an ion-to-electron transducer in solid-contact ISEs in the beginning of the 1990s115 and is frequently still used for the same purpose today. Recent developments were focused on improving the fabrication techniques and the analytical performance of such solid-contact ISEs. Polypyrrole was found to be useful as a solid contact in planar ISEs.116,117 A critical comparison of conducting polymer- and hydrogel-based solid contacts in K+-ISEs showed that polypyrrole doped with potassium hexacyanoferrate results in solid-contact K+-ISEs with better long-term potential stability than those based on the hydrogel contact.118 Polypyrrole doped with Tiron was used as a solid contact in Ca2+-ISEs.119-121 Since Tiron complexes Ca2+, the detection limit of such ISEs was found to be as low as 10-9 M. Similar detection limits were obtained for a Pb2+-ISE using polypyrrole doped with hexacyanoferrate as a solid contact

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when the electrode was used in a flow-through cell.122 The detection limit of Cl--ISEs based on polypyrrole as the solid contact could be lowered by more than 3 orders of magnitude when applying an anodic current that compensated the Clleakage from the ion-selective membrane (due to selfdischarge of polypyrrole).123 The influence of spontaneous charge-transfer processes of polypyrrole on the linear response range and selectivity of Ca2+-ISEs using polypyrrole as a solid contact was studied.124 Galvanostatic polarization of solid-contact ISEs was found to offer some control of the ion flux through the electrode and, consequently, could be used to lower the detection limit.125,126 Interestingly, K+-ISEs based on methacrylic/acrylic membranes in combination with polypyrrole as the solid contact resulted in K+-ISEs with both improved detection limit (below 10-6 M) and excellent stability.127 These results show that polypyrrole-based solidcontact ISEs with polymeric ion-selective membranes are promising also for low-concentration measurements. Polypyrrole doped with tetraphenylborate was used as a solid contact in K+-ISEs in order to have well-defined ion transfer between the conducting polymer and the ionselective plasticized poly(vinyl chloride) (PVC)-based membrane.128 Polypyrrole doped with different anions was also employed as a solid contact in pH electrodes based on polymer membranes containing tertiary amine ionophores.129 The physicochemical properties of polypyrrole can be significantly influenced by the doping ion, which offers possibilities of improving, e.g., adhesion of polypyrrole to the ion-selective membrane. This is well-illustrated by the use of polypyrrole doped with cobalt bis(dicarbollide) ions as the solid contact in pH-, Cu2+-, and K+-selective microelectrodes.130-132 Furthermore, a composite of polypyrrole and Nafion was applied as the solid contact in pH electrodes based on glass membranes.133,134 Polypyrrole was used as an ion-to-electron transducer also in solid-contact ISEs for determination of oxytetracycline hydrochloride and methacycline hydrochloride, which are antibiotics belonging to the tetracycline family.135,136 Poly(1-hexyl-3,4-dimethylpyrrole), which is soluble in tetrahydrofuran (THF), was used as a solid contact in carbonate-selective ISEs based on a silicone rubber membrane.137,138 Solution-processable conducting polymers offer some additional flexibility in the electrode manufacturing process, when compared to electropolymerization.

3.2.2. Polythiophenes Poly(3-octylthiophene) (POT) was the first one of the polythiophenes to be used as a solid contact in ISEs.139 More recently, solution-cast films of POT on screen-printed gold substrates and on platinum (silicon-based substrates) were evaluated as solid contacts in miniature Cl--ISEs.140,141 Improved analytical performance was obtained by using an additional adhesive layer (3-aminopropyltriethoxysilane) between the screen-printed gold electrode and the POT film, while there was no significant difference between PVC and polyurethane (PUR, Tecoflex) used as the ion-selective membrane matrices.140,141 Pb2+-ISEs were constructed by using solution-cast POT as the solid contact to a Pb2+-selective membrane based on poly(methylmethacrylate)/poly(decylmethacrylate) (MMA/ DMA).142 POT is highly lipophilic, which helped to prevent the formation of an internal water layer between the solid contact and the ion-selective membrane. Interestingly, the solid-contact Pb2+-ISE showed a much faster response at low

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Table 1. Principles of the Models in Potentiometry model advanced models here, Nernst-Planck-Poisson (NPP)

parameters used (model dictionary) EM(x,t) ci(x,t) cji(x,t) ki B ki and A (including Ki,j) x (including δ, d) Di, uji  t

main assumptions

main benefits of model application in potentiometry (current state)

(1) general assumptions (see p 5); zero-current, no pressure and temperature gradients, no solvent flow, no convection (2) membrane homogeneous and flat, ionically conducting (3) one-dimension (1D) (4) infinite dilution/ideal phase (activities ) concentrations) (5) sharp border between phases (6) permittivity ) const over distance in each phasea (7) no flux couplinga

time- and space-dependent electrode response, selectivity, and low detection limit, access to membrane physicochemical properties (e.g., migrational effects, permittivity), finite charge-transfer rates

local-equilibrium models here, EM(t f ∞) ci0, cji0, diffusion-layer model (DLM) Ki,j δ, d Di, uji t (?)

(1) as in NPP (1-5) (2) electroneutrality (3) infinite rate constants (4) local equilibrium (5) steady state (limited access to time and space domains) (6) linear concentration changes over distance; electrical potential changes only at the phase boundaries

electrode response, selectivity, and low detection limit under steady state

total-equilibrium models here, Nikolskii-Eisenman (NE)

(1) as in DLM (1-3) time-independent electrode response, (2) total equilibrium selectivity, and low detection limit; (3) constant concentrations and electrical potentials access to chemical binding in each phase except of phase boundaries processes in the membrane

a

EM ci, cji Ki,j ui (?)

6 and 7 specific for NPP.

Figure 12. Pyrrole-based monomer units of polymers applied as ion-to-electron transducers in solid-state ISEs over the past few years: (1) pyrrole, (2) N-methylpyrrole, and (3) 1-hexyl-3,4dimethylpyrrole.

Figure 14. Aniline-based monomer units of polymers applied as ion-to-electron transducers and/or in solid-state ion-selective electrodes over the past few years: (8) aniline, (9) R-naphthylamine, (10) o-aminophenol, (11) o-phenylenediamine, (12) N-phenylglycine, and (13) o-anisidine.

Figure 13. Thiophene-based monomer units of polymers applied as ion-to-electron transducers in solid-state ISEs over the past few years: (4) 3-methylthiophene, (5) 3-octylthiophene, (6) 3,4ethylenedioxythiophene, and (7) 3,4-dioctyloxythiophene.

concentrations and even a slightly better detection limit (10-9.3 M) compared to the corresponding liquid-contact ISE. This has resulted in renewed interest in POT as a solidcontact material in ISEs. The use of solution-cast films of undoped POT as a solid contact together with plasticizerfree acrylate-based ion-selective membranes resulted in a number of solid-contact ISEs (Ag+, Pb2+, Ca2+, K+, I-) with detection limits close to the nanomolar (10-9 M) range.143,144 Solution-cast POT was found to be a very suitable solid-

contact material also for Ca2+-ISEs utilizing conventional PVC-based ion-selective membranes, showing, in addition, a low detection limit.145 However, the long-term potential stability of solid-contact NO3--ISEs based on POT as a solid contact was found to be somewhat inferior to conventional NO3--ISEs with internal filling solution.146 The potential stability of solid-contact K+-ISEs was found to correlate well with the bulk redox capacitance of the conducting polymer, as shown by using poly(3,4-ethylenedioxythiophene) (PEDOT) as a solid-contact material.147 This can be understood when considering that the transduction of an ionic signal into an electronic one via the redox reaction of a conducting polymer results in charging/discharging (doping/undoping) of the conducting polymer layer. From this point of view, chronopotentiometry was found to be a convenient and fast experimental method to evaluate the potential stability of solid-contact ISEs.147

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PEDOT is highly electroactive and shows good environmental stability in its oxidized (p-doped) form. Consequently, PEDOT has been studied rather intensively as a solid-contact material in recent years. PEDOT doped with poly(styrene sulfonate) (PSS), i.e., PEDOT(PSS), was applied as an ionto-electron transducer in solid-contact ISEs for various ions, including K+,147,148 Ag+,149,150 Na+,151 Cs+,151 Ca2+,152 and some aromatic cations (N-methylpyridinium, bupivacaine).153,154 The potential of solid-contact K+-ISEs based on PEDOT as a solid contact was found to be less sensitive to O2 and CO2 (pH) compared to those based on polypyrrole.147,148 A smallvolume radial flow cell for solid-contact ISEs using PEDOT(PSS) as a solid contact was also described.155 The solid contacts were prepared by electropolymerization of the 3,4ethylenedioxythiophene monomer and by solution casting of the commercially available aqueous dispersion of PEDOT(PSS) (Baytron P). Solution-cast films of PEDOT(PSS) (Baytron P) were applied as the solid contact also to screenprinted gold substrates.106 The water-solubility of the solution-cast PEDOT(PSS) film was decreased by ionic crosslinking with multicharged cations (Mg2+, Ca2+, Fe2+/3+, or Ru(NH3)62+/3+) before application of the plasticized PVCbased K+-selective membrane.156 Recently, PEDOT was applied as the solid contact in newly designed microcavitybased solid-contact ion-selective microelectrodes.157 PEDOT doped with hexacyanoferrate was used as the solid contact in Cu2+-ISEs.158 The improvement in the detection limit was attributed to spontaneous accumulation of Cu2+ in the solid contact, causing an influx of Cu2+ ions at the ionselective membrane/solution interface.158 The detection limit and stability of Pb2+-ISEs using PEDOT(PSS) as a solid contact was found to be influenced by the presence of interfering ions in the conducting polymer layer.159 Furthermore, by using soluble poly(3,4-dioctyloxythiophene) (PDOT), which is more lipophilic than PEDOT, it was possible to show the importance of the ion content of PDOT when used as a solid-contact layer.160 All-plastic disposable Ca2+-ISEs and K+-ISEs were prepared via solution-casting of PEDOT/PSS (Baytron P) and plasticized PVC-based membranes on plastic substrates.161 Here, the PEDOT/PSS worked both as the ion-to-electron transducer and as the electronic contact. The same approach was used also for Cu2+-ISEs where the leakage of primary ions from the membrane was eliminated as indicated by a super-Nernstian response for Cu2+ activities below 10-4 M.162 Poly(3-methylthiophene) (PMT) doped with BF4- and modified with EDTA as a complexing agent was used as the solid contact in Ca2+-ISEs for low-level concentration measurements.163 The Ca2+-ISEs showed a super-Nernstian response for Ca2+ activities below 10-5 M, indicating an influx of Ca2+ ions to the ion-selective membrane as a result of complexation of Ca2+ with EDTA present in the solid contact.163

3.2.3. Polyanilines Solution-cast polyaniline (PANI) was used as a solid contact in miniature Cl--ISEs.141 The analytical performance of these solid-contact Cl--ISEs was found to be similar to those based on POT as a solid contact.141 Electrosynthesized PANI was applied as a solid contact in ISEs based on plasticized PVC-based ion-selective membranes. Good overall analytical performance was obtained for such solid-contact ISEs selective to pH164 and Tl3+.165 The stability of the PANI solid contact in K+-ISEs

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based on plasticized PVC was studied by using UV-visible spectroscopy. Partial conversion of PANI from its conducting emeraldine salt form to its nonconducting emeraldine base form was found to take place during long-term measurements (1-3 months) but not during short-term measurements (4 days).166 Different derivatives of polyaniline, including poly(Rnaphtylamine), poly(o-aminophenol), poly(o-phenylenediamine), and poly(N-phenylglycine), were used as solid contacts in plasticized PVC membrane-based ISEs for determination of various organic compounds of pharmaceutical importance, including dimedrol,167 chlordiazepoxide,168 propranolol,169 papaverine,170 amiodarone,171 and dopamine.172 The potential stability of some of the solid-contact ISEs was evaluated by current reversal chronopotentiometry,168,171,172 a method used earlier for K+-ISEs based on PEDOT as the solid contact.147

3.3. Conducting Polymers Dissolved in the Ion-Selective Membrane Solid-contact ISEs, where the conducting polymer is dissolved in the ion-selective membrane, were initially called single-piece electrodes.173 Since the conducting polymer is dissolved in the ion-selective membrane, it can influence the selectivity of the electrode. Solid-contact Li+-ISEs based on plasticized PVC containing 1% (w/w) PANI protonated (doped) with bis(2-ethylhexyl) phosphoric acid were studied.174 The solid-contact Li+-ISEs showed the same dynamic response range as the conventional ISE with internal solution. However, the presence of PANI in the ion-selective membrane increased the H+ interference, due to the pH sensitivity of PANI.174 PANI protonated (doped) with phosphoric acid dihexadecyl ester was used both as an ion-to-electron transducer and as a pH-sensitive component in plasticized PVC-based membranes.175 Membranes containing 50% (w/w) PANI and 50% plasticized PVC showed the best pH sensitivity among those studied. Interestingly, the emeraldine salt-emeraldine base transition of PANI was facilitated by lipophilic cations and hindered by lipophilic anions added to the membrane. However, the analytical performance of these pH electrodes was slightly inferior to that of electrochemically synthesized PANI.175 Solid-state ISEs for determination of linear alkylbenzenesulfonates were constructed by using plasticized PVC membranes containing 5% (w/w) electrochemically synthesized polypyrrole doped with dodecylbenzenesulfonate.176 Here, polypyrrole worked both as an ion-to-electron transducer and as an anion-exchanger for alkylbenzenesulfonate anions. Solid-state K+-ISEs were prepared by using plasticized PVC membranes containing 2% (w/w) polypyrrole, PANI, or poly(o-anisidine) that were doped (protonated) with diesters of sulfosuccinic acid.177 The membranes were solution-cast on planar silver electrodes. The resulting solidcontact K+-ISEs showed comparable selectivity with the corresponding hydrogel-contact K+-ISEs. However, the presence of the conducting polymer, especially polypyrrole, in the ion-selective membrane improved the reproducibility and repeatability of the response. The signal stability of this type of solid-contact K+-ISE containing 2% (w/w) polypyrrole doped with di(2-ethylhexyl)sulfosuccinate was improved when the membrane was cast on a gold substrate instead of silver,178 which can be related to the higher work function

Potentiometric Ion Sensors

of gold compared to silver. The presence of polypyrrole in the membrane did not influence the ion-selectivity in this case.

3.4. Conducting Polymers as Sensing Membranes Solid-state ion-selective electrodes based on immobilization of ion-recognition sites in the conducting polymer membrane represent a research area of great potential. Over the past few years, the main focus has been on conducting polymers that contain ion-recognition sites in the form of immobilized doping ions.24 Covalent binding of ion-recognition sites to conducting polymers was already suggested in the 1980s.179 Covalent binding of ion-recognition sites to the conducting polymer backbone allows integration of the ionrecognition sites and the ion-to-electron transducer even within the same (macro)molecule, which may be of great importance for the construction of durable micro- and nanosized ion sensors in the future. However, the synthesis of such functionalized monomers and their polymerization is relatively demanding. A solid-state Zn2+-ISE based on electrochemically synthesized polypyrrole doped with tetraphenylborate was developed.128,180 The selectivity coefficients (log KZn,j) determined by the fixed interference method were as follows: j ) Ca2+ (-2.7), Mg2+ (-2.1), Pb2+ (-1.6), Ni2+ (-0.6), and Co2+ (-0.6).128,180 Solid-state ISEs for different cations (Ca2+, Mg2+, Cu2+, and Zn2+) based on electrochemically synthesized polypyrrole doped with metal-complexing ligands were further developed.181,182 In addition to metal-complexing groups, the ligands contain sulfonate groups that compensate for the positive charge of the oxidized (p-doped) polypyrrole backbone. The effects of chemical (soaking) and electrochemical (oxidation/reduction) conditioning on the potentiometric sensitivity and selectivity of polypyrrole doped with metal-complexing ligands were studied in detail.181,182 Polypyrrole doped with adenosine triphosphate (ATP) was found to give a near-Nernstian response to Ca2+ and Mg2+, and the response was not influenced by Na+.183 It was suggested that the polypyrrole doped with ATP can be used as artificial membranes in order to model ATP-mediated processes of real biological membranes.183 Furthermore, there was a correlation between the film topography and the potentiometric response of PEDOT doped with ATP, which was sensitive to Ca2+ and Mg2+.184 Smoother films generally showed a more stable and faster potentiometric response than the rougher ones.184 Polypyrrole and PEDOT doped with heparin were also found to give a near-Nernstian response to Ca2+ and Mg2+, and the response was not influenced by Na+ or K+.185 The response remained unchanged even after 1 year of soaking, indicating the high stability of this type of electrode.185 A solid-state pH electrode were developed by using electrochemically synthesized polypyrrole doped with cobalt bis(dicarbollide).186 The electrode showed a quasi-Nernstian response (-50 mV/pH unit) and a linear range from pH 3 to 12. Electrodeposited polymers based on various monomers containing amino groups (1,3-diaminopropane, diethylenetriamine, pyrrole, p-phenylenediamine, and aniline) were studied as pH sensors, showing a linear response in the range from pH 2 to 11.187 Polypyrrole-based pH sensors were miniaturized.188 Electrochemically synthesized PPy doped with bicarbonate was applied as a pH electrode in a Severinghaus CO2 sensor.189 Electrochemically synthesized

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polypyrrole doped with dodecylsulfate (DS-) was used for the development of a tubular solid-state ISE for determination of DS- in a flow-injection analysis system.190 A correlation between the spontaneous charging/discharging processes and the potentiometric detection limit of conducting polymers, such as polypyrole, poly(N-methylpyrrole), and PEDOT, was found.191 Furthermore, the potentiometric detection limit for electrosynthesized PEDOT(PSS) could be lowered to 10-6-10-7 M by polarization with a small cathodic current that prevented cation leakage from the polymer film into the solution. However, similar to most nonfunctionalized conducting polymers, PEDOT itself showed low selectivity.192 Solid-state Ag+-ISEs based on polypyrrole and PEDOT doped with sulfonated calixarenes (and resorcarenes) were developed.193,194 The response mechanism was studied for Ag+-ISEs based on PEDOT doped with hexabromocarborane.195 Pretreatment of these electrodes by cyclic voltammetry in KNO3 solution was found to improve the potentiometric response to Ag+.195 Electrochemically mediated doping/templating by repeated oxidation/reduction in AgNO3 solution was employed to enhance the potentiometric Ag+ response of polypyrrole that was synthesized by electropolymerization of pyrrole in the presence of Eriochrome BlueBlack B.196 Even undoped poly(3-octylthiophene) (POT) was found to give a selective potentiometric response to Ag+, indicating that Ag+ interacts with the polythiophene backbone (sulfur atoms, π-electrons).197 Electrochemically overoxidized polypyrrole (OPPy) gave a potentiometric response to alkali and alkaline earth metal cations, albeit with low selectivity.198 Overoxidation was suggested to result in oxygen-containing groups acting as “hard” Lewis bases that form complexes with hard cations, while the redox interference was simultaneously decreased because of the lower electronic conductivity of overoxidized polypyrrole.198 Novel electropolymerized films based on some functionalized polyanilines, polypyrrole,s and amino heterocyclic compounds were studied as solid-state ISEs for determination of anions, such as amino acids and ascorbic acid.199 In fact, oxidized (p-doped) conducting polymers are inherently suitable for anion sensors because of the polycationic backbone.199 A very simple procedure for manufacturing of a solid-state NO3--ISE by electrosynthesis of polypyrrole doped with NO3- on a pencil lead was described.200 Similarly, PEDOT doped with ClO4- worked very well as ClO4- sensors with similar selectivity as commercial ClO4-ISEs.201 Following the same approach, polypyrrole doped with valproate was found to give a well-functioning valproate sensor.202 Solid-state Cl--ISEs based on chemically synthesized undoped POT containing trihexadecylmethylammonium chloride (THMACl) ions were studied.203 In contrast to tridodecylmethylammonium, the more lipophilic trihexadecylmethylammonium cation required the addition of a plasticizer (2-nitrophenyloctyl ether, o-NPOE) to the POT film in order to give a functioning Cl--ISE with the following composition: 35% (w/w) POT, 23% (w/w) THMACl, and 42% (w/w) o-NPOE.203 Solid-state Cu2+-ISEs were developed by using electrosynthesized undoped polycarbazole and polyindole as sensing membranes.204 However, these electrodes showed a severely super-Nernstian response to Cu2+ at concentrations higher than 10-4 M.

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A solid-state pH nanoelectrode was constructed by electrodeposition of PANI onto an ion-beam conically etched carbon fiber with a tip diameter of ca. 100-500 nm.205 The pH electrode gave a slope of ca. -60 mV/pH unit in the linear pH range of 2.0-12.5. The selectivity coefficients (log KH,j) were around -12 with respect to K+, Na+, Ca2+, and Li+, which is comparable to conventional glass pH electrodes. A solid-state pH microelectrode based on PANI electrodeposited on a microband electrode was used for in situ pH measurement of the self-oscillating Cu(II)-lactate system.206 The pH sensitivity of PANI and its derivatives was found to depend on the substituent and the doping anion incorporated in PANI during electropolymerization.207,208 The best pH response was obtained for PANI doped with chloride, while N-substituted PANI showed no pH sensitivity, because the N-substituent hindered the emeraldine salt-emeraldine base transition.207,208 A solid-state ISE for determination of dodecylsulfate (DS-) was prepared by using electrochemically synthesized PANI doped with DS-. The DS--ISE showed a Nernstian response to DS- in the linear range from 10-9 to 10-5.5 M. The electrode showed relatively high selectivity to DS-, with the strongest interference being observed from dodecylbenzenesulfonate (DBS-) (log KDS,DBS ≈ -2.2).209 Alternatively, when PANI was electrosynthesized in the presence of DBSas the doping anion, the electrode became selective to DBS(e.g., log KDBS,DS ≈ -2.2) with a Nernstian response to DBSin the linear range from 10-5.3 to 10-2.4.210 The potentiometric response of PANI toward aniline based on the formation of charge-transfer complexes between aniline (donor) and the oxidized form of PANI (acceptor) was explained theoretically.211 The theory could also explain the potentiometric response of other similar sensor materials to aromatic amines, pyrrole, and phenols.211 A novel approach for the potentiometric detection of saccharides using poly(aniline boronic acid) was presented.212 The detection was based on pKa changes of the poly(aniline boronic acid) resulting from boronic acid-diol complexation.

3.5. Comments on the Ion-to-Electron Transduction Process Not only solid-state ISEs/ISSs but also conventional ISEs with internal filling solution (and reference electrodes as well) are asymmetric systems, because ions do not enter electronic equipment used to measure the potential. At some point, there is a transduction of the ionic signal into an electronic signal via a reversible redox reaction. In the case of a conventional Ag/AgCl electrode in contact with chloride ions, the reversible redox reaction involves the Ag/Ag+ redox couple, as follows:

Ag + Cl- h AgCl + e-

(22)

The ion-to-electron transduction is completely analogous for other electroactive materials, such as conducting polymers.63-65 Since the electrode contains a finite amount of redox-active material, the electrode has a finite redox capacitance (C).147 For a given constant current (i), there will be a change in the electrode potential (E) with time (t):

∆EC i ) ∆t C

(23)

Additionally, the electrode resistance (R), including all timeindependent resistances of the electrode, will give a constant potential shift:

∆ER ) iR

(24)

Equations 23 and 24 represent the simplest possible case that neglects diffusion/migration processes in the ion-selective membrane and the ion-to-electron transducer. However, this simple RC model was verified experimentally and is used mainly to conceptualize the ion-to-electron transduction process in ISEs.147,213 In practical potentiometric measurements, the current (i) of the high-impedance voltmeter is small enough so that (∆EC/∆t) and ∆ER can be neglected. In contrast, currents induced by electrical noise may be significant enough to influence the potential. Furthermore, miniaturization of electrodes tends to decrease C and increase R, resulting in lower potential stability. This is valid for conventional ISEs with internal filling solution as well as for solid-contact ISEs. The same reasoning is also true for ion-selective electrodes with a completely blocked interface (coated-wire electrodes) where the redox capacitance is replaced by a double-layer capacitance. Consisely, the potential stability of coated-wire electrodes should improve by increasing the contact area between the ion-selective membrane and the electronic contact, because of the increase in the double-layer capacitance. The conjugated polymers so far used as ion-to-electron transducers can be divided into two main groups of materials based on their redox capacitance, which is related to the oxidation level (degree of p-doping) of the conjugated polymers. POT represents a conjugated polymer with a relatively high oxidation potential, and therefore, it has a very low degree of p-doping under ambient conditions. This means that POT has a low electronic conductivity and a low redox capacitance under ambient conditions. On the contrary, PPy, PEDOT, and PANI represent conjugated polymers that are relatively stable in their highly oxidized (p-doped) state, showing high electronic conductivity and high redox capacitance. On the basis of the magnitude of the redox capacitance, highly p-doped conducting polymers, such as PPy, PEDOT, and PANI, should, therefore, be more suitable as ion-toelectron transducers compared to POT. However, this is not the complete account, partly because there may be electrochemical side reactions taking place in parallel with the main ion-to-electron transduction process. Such parallel reactions may be due to the presence of, e.g., O2, CO2, and H2O that can reach the solid contact, even if it is coated with a polymeric ion-selective membrane. Side reactions due to redox couples in the solution are, of course, more likely to play a role when conducting polymers are dissolved in the ion-selective membrane and especially when conducting polymers are used as sensing membranes. Electrochemical side reactions that change the redox state of the conducting polymer will not only influence the electrode potential but also cause a flux of ions into or out of the conducting polymer, which may, for example, influence the detection limit of the ISE. This is likely to happen in the case of highly p-doped conducting polymers that are electroactive in a broad potential range. PEDOT is known to be less sensitive to CO2 and pH compared to PPy and PANI. However, similar to other highly p-doped conducting polymers, PEDOT is also sensitive to redox interference, which results in long-term potential drift. Regardless, the

Potentiometric Ion Sensors

higher the redox capacitance, the smaller is the potential change for a given current and ion flux, i.e., for a given rate of an electrochemical side reaction or a given applied current that compensates for such a side reaction (chronopotentiometry). Simply expressed, this means that a high capacitance stabilizes the potential of the solid contact itself but does not eliminate the ion flux associated with side reactions. The above discussion suggests that a high capacitance is only a partial solution to the stability problem. It would be of utmost importance to have a conducting polymer (or any other electroactive solid material) with a high redox capacitance (in a narrow potential range) that would not participate in any side reactions at all. However, even in the absence of electrochemical side reactions, there may be ion-exchange processes between counterions of a highly p-doped conducting polymer and ions in the ion-selective membrane (in the case of solid-contact ISEs) or solution (in the case where the conducting polymer is used as a sensing membrane). Such ion-exchange processes depend on the selectivity of the ion-selective membrane and the solid-contact materials. Even salt formation at the interface between the conducting polymer and the ionselective membrane is possible, depending on the solubilitity of the ions present in the two phases. The presence of water at the interface leads to an unwanted situation resembling the conventional liquid-contact ISE where the conducting polymer plays the role of an internal reference electrode in contact with a very small volume of internal filling solution, the ion composition of which can easily change with time. As mentioned previously, POT has a low redox capacitance and a low electronic (and ionic) conductivity. This means that the potential of POT is more sensitive to current flow than PPy, PEDOT, and PANI. However, because of its low conductivity, POT is less electroactive and may not participate in side reactions to the same extent as the highly p-doped conducting polymers. Additionally, POT has a low content of ions and is relatively lipophilic, which prevents the accumulation of water and salt inside the solid contact. In the case of solid-contact ISEs, the compatibility between the material used as a solid contact and the plasticizer (and polymer) used in the ion-selective membrane is a key issue that has not been studied thus far. Furthermore, leakage of plasticizer, ionophore, and additives to the solution may become a serious problem in the case of miniaturized ISEs based on plasticized polymer membranes. It is interesting to note that low detection limits for several ions were achieved by using POT as a solid contact together with plasticizer-free acryl-based ion-selective membranes. This indicates that POT may be a good alternative as a solidcontact material, despite its low redox capacitance, if the main target is to reach a low detection limit. In contrast, POT is expected to be more sensitive to external electrical noise compared with the highly doped conducting polymers. The above discussion focused on ion-to-electron transducers based on conducting polymers, because these materials have been extensively studied for many years and they appear to be very promising. Conducting polymers make it possible to fabricate, e.g., all-plastic ISEs161 and pH nanoelectrodes.205 However, recent results indicate that excellent potential stability (potential drift ca. 11.7 µV/h) can be achieved by using three-dimensionally ordered macroporous carbon as the solid contact.214 Despite the relatively complicated manufacturing procedure, this latter approach is indeed very promising. In this case, the large contact area between the

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ion-selective membrane and the electronically conducting carbon material should result in a large double-layer capacitance that would stabilize the potential. Further characterization of this type of solid-contact ISE by, e.g., electrochemical impedance spectroscopy may provide important fundamental information concerning the ion-to-electron transduction process of solid-contact ISEs. Consequently, which material is the best ion-to-electron transducer? The above discussion indicates that the choice of conducting polymer or any other material used as an ionto-electron transducer must be done on a case-to-case basis. Requirements, such as analytical performance, mechanical durability, lifetime, fabrication methods, and cost of the ion sensor, have to be considered. Finding a solid material that fulfills all possible requirements of a good ion-to-electron transducer for all possible applications is really a great challenge.

4. Miniaturized ISEs 4.1. Conventional Microelectrodes Ion-selective microelectrodes are used in many different applications from life sciences to environmental control. There are several reviews, books, and sections of books dedicated to this topic, e.g., by Purves,215 Ammann,216 and Takahashi et al.217 Microelectrodes based on the concept of micropipets have already been used for a longer time, especially in life-science applications.218 Two different pipet electrodes, one as the indicator electrode and the second as the reference electrode, have been used, as well as the construction of a doublebarreled micropipet electrode. The pipets have to be silanized before applying the ion-selective cocktail into the pipet. An improved method for this procedure was described recently.219 There is also a patent for the apparatus and method to selectively induce hydrophobicity in a single barrel of a multibarreled ion-selective microelectrode.220 Scanning electrochemical microscopy (SECM) is a technique where ion-selective microelectrodes have found wide areas of application and where the advantages are the micrometer dimensions, selectivity, and low detection limit.221 pH probes based on an antimony microdisk electrode222 or a H+-sensitive liquid membrane have been used in pHmicroscopy where pH profiles have been measured in corrosion studies223,224 or in measuring biological activities.225,226 The technique has also been applied to study many other kinds of surfaces and processes.227,228 A novel type of solid-contact ammonium-selective microelectrode was recently constructed for SECM measurements.229 Doublebarreled chloride-selective microelectrode were used to map in situ Cl- ion distribution in localized corrosion systems.230 There are also microelectrodes for some “not-so-common” ions. A new Schiff’s base was synthesized and tested as the ionophore for yttrium ions in an ion-selective microelectrode.231 That sensor was tested in complexometric titration and direct determination of yttrium in dissolved yttriumaluminum alloy samples. Environmental measurement of ionic species with ionselective microelectrodes is also an interesting application area. Several ions have been measured in pore water samples,232 and a phosphate-selective microelectrode based on cobalt as the sensing material was used in studying a biological phosphorus removal process.233

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By selecting an optimum composition for the inner reference solution, the detection limit of a calcium micropipet electrode could be lowered.234 The calibration curve was found to be linear down to 10-8 M of Ca2+. Ion-selective microelectrodes were recently used in determination of extremely low levels of calcium, lead, and silver ions, 10-10 M in the presence of 10-5 or 10-6 M sodium ions.235 It was possible to detect 300 attomol of those ions at a constant ionic background.

4.2. Solid-Contact Microelectrodes During the past decade, microsensors have gained a widespread and increasing interest both by scientists, engineers, and people working with practical applications of these sensors. The collaboration between scientists and microtechnology engineers has been fruitful, leading to different manufacturing procedures for making both thin- and thickfilm microelectrodes. New technologies enabling mass production of miniaturized ion-selective electrodes were reviewed.236 The discussion concerned mainly applications in biology and medicine, but procedures to make thin- and thick-film microelectrodes were also reviewed.236 Most of the currently used microelectrodes are manufactured on solid substrates. Gold, silver, platinum, and different kinds of carbon materials are used. A recent review on advanced chemical microsensors deals with the design and performance of miniaturized chemical sensors based on Si transducers: ion-selective field effect transistors and solidstate electrodes.67 Si nitride substrate with a polymeric solid contact was used in a miniaturized sodium-selective electrode.237 Silicon substrate was also used in constructing a pH-sensitive microelectrode where the sensitivity is due to ion-selective PVC membrane on top of polypyrrole layer doped with cobalt bis(dicarbollide).238 Pulsed-laser deposition is also a practical method for making solid-state micro-ionselective electrodes. Cd2+ selective microelectrode was constructed by pulsed laser deposition (PLD) technique and using CdSAgIAs2S3 as the sensitive material.239 Screen-printing technology is a useful and practical way to produce solid-contact electrodes. A procedure to make an array of all-solid-state potassium selective electrodes on screen-printed gold substrates has been described.156 Screenprinting technology was also used to produce thick-film Ag sulfide electrodes.240 The performance of these strip sensors was comparable to those of conventional electrodes. Another way to make silver- and sulfide-selective electrodes is to etch the end of a silver wire to a sharp tip and then to insulate the wire except for the very end that then functions as the sensing part of the probe.241 Even an all-solid-state potentiometric sensor for ascorbic acid was constructed based on this screen-printing technology.73 Laser-ablation, screenprinting, and molecular-imprinting techniques were used in making nanoliter-volume vials with carbon and Ag/AgCl ring electrodes embedded in the sidewalls.242 Polypyrrole was electropolymerized on the carbon rings with nitrate as the doping ion, and the vials were used in determination of nitrate in nanoliter samples. On the basis of earlier reports on anionselectivity of polypyrrole, it is somewhat suprising how selective are the nitrate sensors that the authors have been able to produce by using nonfunctionalized polypyrrole. The main advantage of that system lies in the small sample size. Screen-printing technique is also used in making planar-form solid electrolyte modified Ag/AgCl reference electrodes.243,244 Construction and design of this type of reference electrodes

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make them suitable for measurements in physiological samples and in flow analysis. A novel method to make miniaturized solid-contact ionselective electrodes has been presented.157 The gold wire in a glass capillary was etched to form a microcavity in the electrode body. The gold disk electrode was then coated by electropolymerization with the conducting polymer poly(3,4ethylenedioxythiophene) (PEDOT) doped with poly(4-styrenesulfonate) (PSS). Potassium- and calcium-selective microelectrodes were then made by filling the cavity with the corresponding ion-selective membrane. These microelectrodes were found to have similar potentiometric characteristics as conventional electrodes. Solid-contact microelectrodes are suitable for use in harsh and remote environments. An exotic example of such an application is the use of a microelectrode array for measurement of ion concentrations in the Martian soil.245 A prototype miniature array of polymer-membrane and solid-state ionselective electrodes was developed to perform in situ analysis of soil samples for a number of ions. The array consisted of 27 microelectrodes for nine ions. Each electrode was in three replicates.68 Screen-printing technology has also been used in constructing miniaturized solid-contact electrodes for chloride ions. Either carbon or silver was used as the screen-printed substrate. The conducting polymers polypyrrole or polyaniline with an adhesive admixture of 3-aminopropyltriethoxysilane was placed on the substrate. The ion-selective PVC membrane containing the ion-exchanger for chloride (methyl-tri-n-tetradecylammonium chloride) was then placed on the conducting polymer film. This construction showed the same performance and stability as the conventional electrodes with inner reference electrode and solution.116,246 The potential stability of microfabricated potassium ionselective electrodes with two types of solid contact was studied and compared.118 Hydrogel and polypyrrole doped with potassium hexacyanoferrate(II)/(III) were used as the solid contact between the PVC-based potassium-selective membrane and the screen-printed silver or platinum substrates. The electrodes were incorporated into a flow-through cell where the measurements were conducted. The electrodes with the polypyrrole solid contact showed much higher potential stability than the electrodes with hydrogel contact.118 A microfabricated ion-selective microelectrode-array platform has been constructed and characterized.247,248 The platform contained 24 micropipets individually filled with a Ca2+ selective membrane and was developed for in vitro intracellular measurements. Different chalcogenide glasssensitive materials have also been used to make an array of miniaturized ion-selective electrodes.249 It was demonstrated in that work that the sensor array allows the problem of an insufficient selectivity of single sensors to be overcome. A microsensor array of miniaturized solid-state ion-selective electrodes to analyze sweat samples for sodium, potassium, and chloride was developed for point-of-care diagnosis of cystic fibrosis.250 Miniaturized, planar ion-selective electrodes fabricated by thick-film technology were recently reviewed.251 Different manufacturing procedures were discussed and screen printing was found to be a suitable technique because of its simplicity, low cost, high reproducibility, and efficiency in large-scale production. Thick-film technology was also demonstrated to be a proper method to produce ionophore-based ion-selective

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electrodes on inexpensive substrates. The analytical parameters of these electrodes were comparable with those of conventional electrodes. Miniaturized Ag, Ag/AgCl, or iridium oxide solid-state potentiometric sensors were used to measure the capacitance and the response time of the electrodes by an electrochemical time-of-flight method.252,253

4.3. Microelectrodes in Flow Analysis Solid-contact ion-selective microelectrodes have many advantages when used in flow channels. They are robust and rather easy to fabricate and install. A miniaturized ionselective Cu2+ electrode was developed for fluidic microsystems.254 Cu was deposited on a silicon wafer and then converted to CuS by hydrogen sulfide. The electrode performed well in the system. The authors found differences in the adhesion of the CuS layer for Si, SiO2, and Si3N4. An all-solid-state potassium microelectrode for flow analysis was developed for measurements in flowing solutions177 as well as a nitrate sensor for determination of nitrates in wastewaters, fertilizers, and pharmaceuticals by the flow-injection analysis technique.255 A small-volume flow cell was developed by incorporating all-solid-state electrodes.155 Ca2+ and pH measurements were carried out in the Lab-on-Valve instrument by using all-solid-state microelectrodes.152 Lowtemperature cofired ceramic technology was used to construct continuous-flow analytical microsystems.256 Ammonium- and nitrate-selective microelectrodes were incorporated in that microfluidic sysytem, allowing a complete on-chip integration for potentiometric detection. A microfluidic device was developed for measurement of pH as well as of Ca2+ and K+ concentrations with in-channel-incorporated all-solid-state ion-selective electrodes.257 Construction of different planar solid-state ion-selective microelectrodes as well as reference electrodes for flow analysis was recently discussed in an overview.117

4.4. Life Science and Biology Applications The use of microelectrodes to investigate transportation of inorganic ions in plants has been reviewed.258 The construction of multibarreled ion-selective microelectrodes for measurements in biological tissues was also reviewed.259 It was also demonstrated how cell volume, membrane potential, and intracellular ion concentrations can simultaneously be determined by using potentiometric measurement with multibarreled ion-selective electrodes. A special study on the cell volume regulation mechanism of nerve cells by using multibarreled ion-selective microelectrodes was made.260 Different requirements for the charge-exchange processes between a dielectric layer and the ions in the sample in ion-selective biosensors made on silicon were discussed in another study.261 It was found that standard dielectric materials normally employed in microelectronics technology can be used in making ion-sensitive field-effect transistors and ion-selective microelectrodes for biosensing applications. Use of ionselective microelectrodes and fluorescent dyes in measurement of intracellular pH has been compared and discussed.262 Ion-selective microelectrodes were used in plant physiology in studying the NH4+, H+, and NO3- fluxes around the roots of plants263 and their concentration profiles in nitrifying biofilms.264 By using these electrodes, the authors were able to determine the source of nitrogen for a particular plant. The mechanism of pH homeostasis in Listeria monocytogens subjected to, e.g., acid stress was studied by using a pH

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microelectrode,265 and ammonium transport across the plant cell membrane was measured with an ammonium-selective microelectrode.266 Ionized magnesium has been measured with an ionselective microelectrode in physiological solutions,267,268 and the release of ionized Ca from bones has been measured by a Ca-selective microelectrode.269 Ion-selective microelectrodes were also used in studying ion fluxes across cardiac membrane patches.270 Potassium, calcium, sodium, and proton fluxes were measured, and the authors were, additionally, able to quantify those fluxes. The effect of interfering ions in using ion-selective microelectrodes in biological applications has been discussed.271 The kinetics of ion fluxes in some plant leaf and root tissues was determined by using ion-selective microelectrodes,272,273 as was ion fluxes in bacteria.274 Fluxes of H+ and Ca2+ ions during the growth of a single-celled fungus were also studied by ion-selective microelectrodes.275 Salicylate selective microelectrodes were even constructed for biological applications to measure fluxes of salicylate ions.276 Ion-selective electrodes are used in many biosensors to detect, e.g., ammonia or H+ released in the enzymatic reactions involved in the detection procedure. Ion-selective microelectrodes are needed for miniaturization of such biosensors.277 Some of the main problems in using ion-selective microelectrodes are the slow response time and the susceptibility to noise because of the high electrical resistance of the convential micropipet electrodes (see section 3.5). These problems are especially crucial in biological in vivo measurements but can partially be overcome by using a concentric inner micropipet.278 Response times in the order of a few ms were obtained for pH and Ca2+ electrodes used in extracellular measurements.

5. New Wave During the past decade, the developments in the field of potentiometric ion sensors (ISEs) have been exceptionally fruitful. The discovery of the low detection limit in 1997 represents a major breakthrough in the field of ISEs.30 Much research has been devoted to the design and optimization of liquid-contact ISEs to measure low ion concentrations.234,235,279-285 The “new wave” of ISEs has unquestionably arrived.6,7 The question is: where will it lead us in the future? Solid-contact ISEs for trace-level analysis have already been demonstrated by several groups.75,120-123,125,142-144,163 When such powerful ion sensors are combined with solidstate reference electrodes,243,286-292 mass-producible miniaturized ion-sensor systems with unforeseen analytical capabilities are within reach. Deep theoretical understanding based on advanced modeling of the potentiometric response will definitely boost this development. Parallel developments of plasticizer-free membranes,293-304 covalently bound ionophores,293,296,298,302,305-309 and charged sites304 will enhance the durability of ion sensors. Other alternative membrane materials and construction principles add new tools for the future as well. Here, we would like to mention anion-sensitive epoxy resins,310,311 membranes prepared by the sol-gel method,312 polymer-supported liquidcrystal membranes,313 polymer membranes located inside an electrochemically inert porous matrix,314 monolithic capillarybased membranes,284 and fluorous membranes with exceptionally low polarity.315

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There are several other interesting developments that may have an impact on the future developments of ISEs. The socalled sandwich membrane method, which was actually introduced already in 1979, is a useful tool for studying complexation in real membranes.316,317 Polycation-sensitive electrodes can be used to detect even polycationic dendrimers.318 Highly lipophilic closo-dodecacarborane anions showing very weak ion-pair formation are promising alternatives to the commonly used tetraphenylborate derivatives as anionic additives in polymeric ion-selective membranes.319 Under certain conditions, conventional ISEs with internal filling solution can be calibrated by varying the composition of the internal solution, which may be particularly useful in applications where the electrode cannot easily be removed from the sample (e.g., process control and in vivo measurements) during calibration.320 A solvent-free method for making hotpressed ion-selective membranes is an environmentally friendly approach that avoids the use of organic solvents.321 Application of a controlled current can be used to influence transmembrane ion fluxes and is an elegant approach toward lowering the detection limit of ISEs.322,323 This method has also been applied to solid-contact ISEs using polypyrrole as the ion-to-electron transducer.126 As discussed in section 3.5, highly p-doped conducting polymers with a high redox capacitance can simultaneously work as “source” and “sink” of electrons and ions without any significant change in potential if the magnitude of the applied current is sufficiently low. Therefore, chronopotentiometry may become a useful method to control transmembrane ion fluxes and detection limits of solid-contact ISEs. Under some conditions, currentreversal chronopotentiometry allows the equilibrium potential to be determined although the electrode is significantly polarized by the applied current.213 Furthermore, a significant increase in sensitivity was achieved by imposing current pulses to the ISE, leading to the concept of so-called pulstrodes.324 Being a fascinating research area in itself, the new wave of ISEs will most probably also result in new applications of potentiometric ion sensors in the future.

6. Conclusions The response of potentiometric ion sensors, i.e., ionselective electrodes (ISEs), is a complex time-dependent phenomenon that calls for advanced theoretical modeling in addition to classical equilibrium models. This is particularly important when ISEs are used in their nonequilibrium regimes, e.g., in order to reach low detection limits, which is intensively being explored all over the world. The prerequisites for obtaining solid-contact ISEs with stable potential are well-documented, and during the past decade, conducting polymers have gained widespread acceptance as ion-to-electron transducers in solid-contact ISEs. However, in spite of extensive research in this area, obtaining solidcontact ISEs with reproducible standard potentials is still a great challenge. Miniaturized versions of both conventional and solid-contact ISEs have been used in various applications. Cost-effective miniaturized ion-sensor systems with unforeseen analytical capabilities are within reach. Advances made by numerous research groups all over the world allow us to conclude that the future of potentiometric ion sensors looks very prospective indeed.

7. Acknowledgments A.L. thanks Dr. Tomasz Sokalski and Dr. Witold Kucza for very valuable discussions and assistance during the

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preparation of this text. M.Sc. Krzysztof Szyszkiewicz, M.Sc. Peter Lingenfelter, Prof. Robert Filipek, and Prof. Marek Danielewski are acknowledged for their interest and invaluable support. The financial support (AL) by the Polish Committee for Scientific Research (KBN) project 3T09A 175 27 is acknowledged. This work is part of the activities at the Åbo Akademi Process Chemistry Centre within the Finnish Centre of Excellence Programme (2000-2011) by the Academy of Finland.

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CR068100W

Chem. Rev. 2008, 108, 352−366

352

Amperometric Gas SensorssA Review Joseph R. Stetter* SRI International 333 Ravenswood Avenue, Menlo Park, California 94025-3493

Jing Li NASA Ames Research Center, Moffett Field, California 94035-1000 Received September 4, 2007

Contents 1. Introduction 2. Characteristics of Amperometric Gas Sensors 2.1. Electrochemical Gas Sensors 2.2. Amperometry 2.3. Amperometric Gas Sensors 2.4. Theory of the Limiting Current 2.5. Structure and Geometry of an AGS 2.5.1. Gas Membrane 2.5.2. Electrolyte 2.5.3. Electrodes 2.6. Electrode Reactions 2.7. Analytes 3. Survey of Old and New Amperometric Gas Sensors 3.1. Working Electrode and Sensor Design 3.1.1. Clark Sensor 3.1.2. Diffusion Electrodes 3.2. Electrolytes 3.2.1. Nonaqueous Electrolyte 3.2.2. Solid Polymer Electrolytes 3.2.3. Ceramic Solid Electrolytes at Elevated Temperature 3.3. MEMS and Nanotechnology 3.4. Sensor Array 4. AGSs For Practical Applications 4.1. Industry 4.2. Environment 4.3. Defense and Homeland Security 4.4. Medical/Biomedical 5. Summary 6. Acknowledgment 7. References

352 352 352 354 354 355 355 355 356 356 357 357 357 357 357 359 360 360 360 361 361 362 363 364 364 364 365 365 365 365

1. Introduction The amperometric gas sensor, or AGS, belongs to a large and important class of electrochemical gas sensors, and they play an ever-increasing role in environmental monitoring, medical and health applications, industrial safety, security, surveillance, and the automotive industry. The fundamental * To whom correspondence should be addressed. E-mail: joseph.stetter@ sri.com.

principles and constituents of the AGS sensor device as well as some important events in the design and development of the AGS technology are reviewed herein. The modern AGS is based on a series of scientific discoveries, and it is has contributed much to society in its rich history. Amperometric gas sensors (AGSs) are best divided by the temperature at which they operate and can be represented by two well-known commercial gas sensors that include the ambient-temperature liquid electrolyte gas sensors used for medical and industrial hygiene applications and the hightemperature solid electrolyte O2 sensor used in automotive, process, and stack gas applications. The AGS can operate at low or high temperatures, ranging from below freezing to over 1000 °C, and the materials of construction, including the electrolyte, change significantly. The choice of materials can result in totally different designs. The room-temperature AGS uses liquid or polymer electrolytes and is the system of choice for industrial hygiene, health, safety, and medical applications, while the high-temperature AGS using solid electrolytes has been primarily developed for use in automotive exhaust and stack gas process monitoring. This review focuses more on the low-temperature AGS, but some discussions are included on the high-temperature sensors. The fact that the AGS serves many important practical applications speaks to the superb simultaneous combination of features that can be realized with this approach including: low power, low cost, small size, selectivity, stability, sensitivity, and sometimes fast response time. Given these combined characteristics, the AGS is often described as providing relatively high analytical performance at modest cost.

2. Characteristics of Amperometric Gas Sensors 2.1. Electrochemical Gas Sensors Electrochemical gas sensors can be divided into three main classes according to the operating principle: amperometric, potentiometric, and conductometric sensors. Many sensor types exist within each class. We will not cover potentiometric techniques or conductometric techniques for gas sensors but restrict discussions to amperometric sensors for the gas-phase analytes. In the AGS, the current generated by reaction of an analyte at an electrode is measured as the sensor signal and can be measured at a fixed or variable electrode potential, although it is very common at a fixed applied potential. The reaction rate, reflected by the current at the sensing electrode, occurs at a thermodynamically

10.1021/cr0681039 CCC: $71.00 © 2008 American Chemical Society Published on Web 01/18/2008

Amperometric Gas SensorssA Review

Dr. Joseph R. Stetter obtained a Ph.D. in Physical Chemistry from the University at Buffalo (SUNY) in 1975. In the 1970s, Dr. Stetter was Director of Chemical Research at the Energetics Sciences Division of Becton Dickinson and Company where he developed the first diffusion-type electrochemical CO sensors; the earliest diffusion CO dosimeters; solidstate gas sensors for CO, NOx, SO2, and other toxic gases; and an electrochemical hydrazine sensor still in use by NASA. While at the Argonne National Laboratory in Chicago, IL, in the 1980s, he led the development of the first integrated and operational “sensor-array-based” instrument with pattern recognition (now called electronic nose). In the 1980s, Dr Stetter founded TRI (Transducer Research, Inc.), where he developed portable instruments and sensors for CO and CO2, end-ofservice filter indicators, chlorinated hydrocarbon sensors, NOx sensors, personal protection instruments, and low-cost effective protection equipment for human health and the environment. In the 1990s, he sold TRI and became Professor of Chemistry at the Illinois Institute of Technology, started a sensor research group, and founded the International Center for Sensor Research and Engineering at IIT, mentoring both M.S. and Ph.D. students. He founded Transducer Technology, Inc. (TTI), a startup company focusing on nanotechnology enabled sensors and instruments in 1999. In 2007, TTI merged with KWJ Engineering, Inc., of Newark, CA, making nanosensors for health, safety, and process control applications. Recently, Dr. Stetter is Director of the Microsystems Innovation Center for SRI International (Menlo Park, CA). R&D focuses on new sensors, unique structures/materials, artificial senses, chem/biosensors, novel MEMS for drug and vaccine delivery, vacuum microelectronics including microelectron and ion sources, and micro/nanostructures and bio-MEMS. Dr. Stetter has published more than 100 refereed scientific and technical articles and has more than 25 domestic and foreign patents. He has served as Chairman of the Electrochemical Society Sensor Division and served on the boards of national and international technical societies. He has organized national and international scientific meetings and symposia in his field and serves as editor and reviewer for scientific and engineering journals. His awards include three IR-100 Awards for new product development; the Federal Laboratory Consortium Special Award for Excellence in Technology Transfer; the Argonne National Laboratory Inventor’s award; the Technology Management Association of Chicago’s 2002 “Entrepreneur-of-the-Year” award; and two NASA New Technology Awards. He is on the board of directors for several start-up companies.

determined potential for any given reaction and, when operated under appropriate diffusion-limited conditions, is simply proportional to the concentration of the analyte. The relationship between current and concentration is linear, typically over 3 orders of magnitude, and measurements with high sensitivity (ppm and ppb) are possible with excellent measurement accuracy under constant potential conditions. Faraday’s Law is often applied to the observed currents, and the design of the sensor is chosen to control many kinetic factors including mass transfer of the analyte to the electrode as well as the electrocatalytic activity of the electrode material. The geometry and dimensions of the sensor device have a profound effect on the AGS analytical performance, including the observed sensitivity, selectivity, response time, and signal stability.

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Dr. Jing Li received her Ph.D. in Materials Science in 1996 from University of Utah in Salt Lake City. She is currently a Principal Investigator in NASA Ames Nanotechnology Branch. She leads the effort and has developed a space-qualified nanochemical sensor unit for flight demonstration in a satellite that has been launched and operated successfully since March 2007. The nanotechnology-based chemical sensor platform that she developed has drawn interest from other NASA centers, many industry partners, and government Labs. She is a principle investigator (PI) on the Nano ChemSensor Unit project, which won 2007 NASA Ames Center Honor Award. She worked with a team at NASA on nanoelectronic devices that resulted in a NASA TGIR Award (2002). She is an internationally known expert in the field of chemical sensors. Her research focus is on the chemical sensor development utilizing the nanostructured materials and a sensor array with signal processing techniques for intelligent sensor systems. Prior to joining NASA Ames, she was a lead scientist at Cyrano Science, Inc., where she developed polymer-based chemical sensors for an electronic nose instrument. Dr. Jing Li worked at Pacific Northwest National Lab as a DOE postdoc fellow on the chemical sensor development for waste tank monitoring. She has extensive experience of both research and commercialization in developing chemical sensors and intelligent sensing systems, which includes the following: (1) investigating the sensing materials such as nanostructured carbon nanotube and metal, metal oxides nanowires, and their intercalation in polymers as composites; (2) studying the sensing mechanisms in scope of the charge transfer (first and secondary) interaction and diffusion/partition process between sensing materials and analytes characterized by electrical, electrochemical, and spectroscopic techniques such as UV−vis, FTIR, Raman, work function measurements with Kelvin probe, and adsorption/absorption measurement with quartz crystal microbalance and Micromeritics; and (3) developing the signal-processing techniques (i.e., pattern recognition) for a chemical sensor array to be used as an intelligent system for digitizing the chemical information. She is currently a vice chair for Sensor Division in the Electrochemical Society. Her research interests are chemical sensors and intelligent sensing systems for space applications and terrestrial applications. Table 1. Comparison of Amperometric and Potentiometric Sensors electrochemical class

sensor signal vs [analyte]

amperometric, AGS potentiometric

i ) kP E ) Ec + k ln P

principle kinetic, Faraday law thermodynamic, Nernst law

In a potentiometric electrochemical class of sensor, the open-circuit potential between two electrodes is monitored, and this potential is typically proportional to the logarithm of the concentration of analyte. The relationship is often expressed by the Nernst equation, as shown in Table 1. A potentiometric sensor can measure analyte concentrations over a very wide range, often more than 10 decades, and the sensing principle is thermodynamic, i.e., it assumes that all reactions relating to the sensing are at equilibrium. In potentiometry, the chemical and diffusion processes must be at equilibrium conditions in the sensor for a thermody-

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electrode and a counterelectrode, and the electrolyte solution in which the two electrodes are immersed. Missing from this for the moment is an interface to enable a gaseous analyte to contact the working electrode, but this is discussed later. When applying a proper potential between the two electrodes, electroactive species in the electrolyte solution will participate in electrochemical reactions on the surfaces of each of the electrodes. A reduction reaction occurs at the cathode and can be expressed as shown in eq 1, and at the same time, an oxidation reaction takes place at the anode and can be written generally as shown in eq 2:

Figure 1. Illustration of amperometric sensor with two-electrode configuration for liquids.

namically accurate signal to be observed, while the amperometric sensors rely on Faraday’s law and a dynamic reaction achieving a stready-state condition in the sensor. As may be implied, the amperometric sensor signal will get smaller with the size of the electrode and the rate of analyte reacting at the electrode surface, while the potentiometric sensor assumes a thermodynamic potential independent of size of the electrode. This situation becomes interesting in light of nanostructures wherein the thermodynamic potential is characteristic of aggregates of atoms, while the amperometric sensor reation rate is typically enhanced by the high surface areas afforded by nanostructured electrode materials.

2.2. Amperometry Amperometry is a conventional electroanalytical technique that encompasses coulometry, voltammetry, and constantpotential techniques and is widely used to identify and quantify electroactive species in the liquid or gas phase. For liquid-phase analytes, the electrodes and analytes are immersed in a common electrolyte solution. In contrast, application of amperometry to gas-phase analytes involves a unique gas/liquid/solid boundary (analyte-electrolyteelectrode) and an interfacial transport process that frequently controls the response characteristics and analytical performance of the AGS. A major consideration in the specific design of the sensor device is to choose a geometry and porosity/composition of materials to enable the gaseous analyte to effectively and efficiently transport to the interface of the electrode/electrolyte where a fast, reversible, redox reaction can occur. This interfacial redox reaction provides the charge-transfer reaction of the analyte and the analyte’s characteristic sensor signal.

2.3. Amperometric Gas Sensors Amperometric gas sensors are sometimes called voltammetric, microfuel cell, polarographic, amperostatic, potentiostat, limiting current, or other sensor names.1,2 The most descriptive name is “amperometric”, and this refers to either constant-potential amperometry or variable-potential amperometry. A simple amperometric cell with two-electrode configuration is illustrated in Figure 1 wherein the electrochemical system consists of two electrodes, i.e., a working

O1 + ne ) R1

(1)

R2 ) O2 + me

(2)

The cathode reaction takes electrons from the cathode and combines them with an oxidized species, O1, to produce a reduced species, R1. At the anode, a reduced species, R2, reacts to produce an oxidized species, O2, and electrons are given to the anode. The electrons on the anode or lack of electrons on the cathode would build up a charge on the electrode surface, except that the electrodes are conductors and are connected in an external circuit in which the electrons flow. This is an illustration of Faraday’s Law in that an exact number of electrons are generated or used per each analyte molecule and so the charge (the number of electrons or Coulombs of charge) is exactly related to the number of analyte molecules reacting in the system and the current (electron flow) at the electrode is exactly related to the rate of the electrochemical reactions (current is Coulombs/s). In the AGS, typically the current under diffusion-limited reaction conditions is directly related to the rate of electrode reaction taking place on the electrode surface and, hence, as discussed in more detail below, the rate of the electrode reaction is proportional to the concentration of reactant, i.e., the analyte. Consequently, the electrical charge or current can be used as a sensor signal and will be related to the concentration of the analyte in an AGS. The common characteristic of all AGSs is that measurements are made by recording the current in the electrochemical cell between the working electrodes and counterelectrodes over time as a function of the analyte concentration. When this experiment is performed at a fixed potential controlled by a potentiostatic circuit, the technique should be properly called “constant-potential amperometry.” The AGS produces a current when exposed to a gas/vapor containing an electroactive analyte because the analyte diffuses into the electrochemical cell, to the working electrode surface, and thereon participates in an electrochemical reaction that either produces or consumes electrons (i.e., a redox reaction). Selectivity to certain gaseous analytes is an important consideration for sensing applications. For the AGS, selectivity can be realized by optimizing the electrode material to facilitate or catalyze only selected reactions or by control of the sensing electrode potential to select an applied bias that thermodynamically favors either the oxidation reaction or the reduction reaction for a particular analyte. Noble metals such as gold and platinum, mostly in porous or high surface area designs, have been used for the working or sensing electrode in the AGS because of their excellent chemical stability in electrolyte solutions and high electrocatalytic activity toward analytes like CO, H2S, O2, Cl2, and NO. The applied potential between working and reference electrodes set by the potentiostat fixes the thermodynamic operating

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potential for the working electrode reaction and is critical for both the observed sensitivity and the selectivity of the sensor. The selection of operating potential provides an effective way to change or optimize the performance of the sensor.

2.4. Theory of the Limiting Current The observed current in the AGS can depend on many factors, and the resultant mathematical expression relating the current, potential, and concentration can be complex. There are two special operating conditions, called limiting conditions, that can be used to simplify the resulting expression for relating the rate of an electrode reaction (rr) and the rate of mass transport, e.g., by diffusion (rd), of the reactant to the electrode surface. These conditions include the situation where rr , rd (eq 1) and the situation where rr . rd (eq 2). Both cases are important for sensor applications, and in both these cases, the electrode process can be easily understood and the resulting expression can be simplified. The current obtained at the two rate-limiting conditions is called the limiting current. In the first case, where rr , rd, i.e., the rate of reaction at the electrode surface is the rate-limiting step, the limiting current is controlled by the rate of the electrode kinetics. In this case, the analyte reaches the surface much faster than it is reacted, and so the concentration at the electrode surface is the same as in the bulk of the solution and the gas surrounding the electrode and the expression for the current will take the form of eq 3,

ilim ) nFkAC exp(RnFE/RT)

(3)

where k represents the standard rate constant, F is the faraday constant, R is the gas constant, T is the Kelvin temperature, A is the area of the electrode, C is the concentration, n represent the number of electrons per molecule reacting, and R and E are the transfer coefficient and overvoltage of the electrode reaction, respectively. If all conditions are held constant, the limiting current is proportional to the concentration (C) of analyte with an exponential temperature coefficient, and the derivation of this expression can be found in many textbooks.3,4 The limiting current that is obtained in this kinetically controlled electrochemical region, however, is not always stable because of the degradation of the electrode’s catalytic activity and can be problematic for a practical AGS. This drawback can be overcome by operating the AGS in the region of diffusion control rather than the electrode’s kinetic control. Under this condition, where rr . rd, the rate of diffusion of the reactant toward the electrode surface is much slower than the rate of reaction and now the limiting current is controlled by the diffusion of analyte. In this case, the concentration of the analyte at the electrode surface is virtually zero and every molecule of analyte that reaches the surface is immediately reacted. Under these conditions, the rate of electrode reaction may be limited by the rate of diffusion through a gas membrane or a capillary that is placed somewhere between the bulk analyte sample (e.g., the air containing the anayte) and the catalyst layer of the electrode. On the basis of Fick’s diffusion law, the limiting current, ilim, should now be governed by the rate at which the analyte transports to the surface, and this can usually be written as shown by eq 4,

ilim ) k[CO]gas

(4)

where the current is directly proportional to the gaseous concentration in some convenient units like ppmv (parts per million by volume). The current can also be limited by the rate of diffusion across the diffusion layer in solution, and an expression similar to eq 4 is also obtained. Figure 2 illustrates the two limiting conditions described above wherein the concentration of the analyte is plotted vs the distance from the electrode surface. When the current is limited by the rate of the electrode reaction, the analyte supply is much faster than its consumption and the analyte concentration is almost the same in the bulk and at the electrode surface. For a mass transport (diffusion) limited electrode reaction, the analyte supply is much slower than its consumption by reaction at electrode surface, the analyte concentration falls to zero at electrode surface, and the current is now limited by the supply of analyte. This condition is often preferred by the designer of the AGS since the physical control of the analyte supply by diffusion allows one to build a more stable sensor with the square root temperature dependence of the signal. When operating the AGS using a potentiostat, the sensor can be operated at an electrochemical potential where the reaction is facilitated, and in this case, the magnitude of the sensor signal is practically independent of the electrode potential. In theory, the limiting current region can be achieved in any case when the rate-limiting step is a step prior to electron transfer. Typically, an electrode with relatively high electrocatalytic activity is sought, and the benefits of a stable long-lasting sensor are realized.

2.5. Structure and Geometry of an AGS Figure 3 illustrates the design of a typical three-electrode AGS for CO sensing having a working electrode (WE) reaction [CO + H2O ) CO2 + 2H+ + 2e-] and counterelectrode (CE) reaction [1/2O2 + 2H+ + 2e- ) H2O] and reference electrode (RE) embedded within the cell. The WE, CE, and RE are in electrolytic contact (all touch the electrolyte), and the overall cell reaction is CO + 1/2O2 ) CO2. The common feature of the typical gas sensor is an interface at the working electrode that facilitates transport of the gas to the WE/electrolyte interface. This inlet can be simple diffusion or aided by a small air pump that transports the sample to the backside of a gas porous membrane through which the analyte diffuses/permeates. The gas membrane can be used to control the gas flow into the sensor and it can aid selectivity, allowing only the analyte gas to pass as well as providing a barrier to prevent the leakage of the electrolyte from the interior of the sensor. The design and materials chosen for each part of the sensor are critical and determined, to a large part, by the application and the desired analytical performance. In the following paragraphs, we discuss alternative materials and designs that are used to provide different performance capabilities.

2.5.1. Gas Membrane Several types of porous and gas-permeable membranes exist and can be made of polymeric or inorganic materials. Most of them are commercially available very thin solid Teflon films, microporous Teflon films, or silicone membranes. Issues concerning the choice of membrane include permeability to the analyte of interest, ability to prevent electrolyte leakage, manufacturability, and the thickness and

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Figure 2. Concentration distribution of analyte (A) for different current limit conditions.

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β-alumina, and NASICON. For AGS applications, the electrolyte solution needs to support both counterelectrode and sensing electrode reactions, form a stable reference potential with the reference electrode, be compatible with the materials of construction, and be stable over long periods of time under various operating conditions. In some cases, hydrogels or an electrolyte inside a porous matrix are used to replace free liquid electrolytes to raise viscosity, lower evaporation rates, and resist leakage of the electrolyte from sensor devices. The polymers or hydrogels can prevent the evaporation of electrolyte during sensor fabrication, especially for microsensor devices where very small amounts of electrolyte are used. Solid electrolytes (polymers and hightemperature ionic materials) allow fabrication of a solid-state sensor, which avoids the danger of leakage of liquid electrolytes entirely. In order to get sufficient ionic conductivity, elevated operating temperatures may be required. High-temperature operation can be an advantage for operation in harsh environments, in an internal combustion engine control system, or for exhaust gas monitoring.

2.5.3. Electrodes

Figure 3. Illustration of amprometric gas sensor for CO with threeelectrode configuration.

durability of the membrane. For example, a semipermeable membrane composed of an acrylonitrile butadiene copolymer can be used to selectively detect the partial pressure of oxygen in blood samples by allowing oxygen transport and effectively preventing the transport of the other species present in the sample. In a number of cases, the rate of mass transport through the membrane controls the limiting current and, hence, the sensitivity of the AGS. An ideal gas membrane will have a constant permeability to the target gaseous analyte during sensor operation over a wide temperature range and possess mechanical, chemical, and environmental stability.

2.5.2. Electrolyte The role of the electrolyte is to transport charge within the sensor, contact all electrodes effectively, solubilize the reactants and products for efficient transport, and be stable chemically and physically under all conditions of sensor operation. The electrolyte is an ionically conducting medium that ionically transports charge within the cell. Examples of commonly used electrolytes include the following: aqueous electrolytes, such as sulfuric acid, sodium hydroxide, and potassium chloride; nonaqueous electrolytes, such as propylene carbonate with lithium perchlorate; and solid electrolytes, such as the polymer Nafion when wet and, for hightemperature sensors, yttria-stabilized zirconia or YSZ,

The electrode reaction, i.e., electron-transfer reaction at the WE, involves several steps,1,2 including adsorption of the analyte onto the surface, electroreaction, and desorption of products from the electrode surface. Typically, the working electrode is made from a noble metal, such as platinum or gold, which is capable of making a defined interface with the electrolyte in the cell and is porous to allow efficient diffusion of the gas phase to a large and reactive electrode/ electrolyte interface. The noble metals generally exhibit excellent stability under polarized potentials that may be corrosive to other metals. The noble metals are also excellent catalysts for many analyte reactions. Carbons, including graphite and glassy carbon, are also popular materials for working electrodes, especially for sensing bioanalytes since many forms of carbon are biocompatible. The selection of the WE material is, therefore, based on the electrochemical and electrocatalytic properties as well as its stability and manufacturability. The counterelectrode must also be stable in the electrolyte and efficiently perform the complementary half-cell reaction that is opposite of the analyte reaction.2 A Pt electrode is very often used as the CE in an AGS. In addition to the working electrodes and counterelectrodes, a reference electrode is also present when a potentiostat is used.1 The reference electrode must form a stable potential with the electrolyte, be compatible with the sensor manufacture, and, generally, not be sensitive to temperature (T), pressure (P), relative humidity (RH), or other contaminants or reactants in the sensor system. The RE must be able to maintain the WE at a constant electrochemical potential during the sensing application. An Ag/AgCl reference electrode, which shows good reversibility, is commonly used for this purpose. The other popular reference electrode is a Pt/air electrode, which is not a classical refernce electrode but is sometimes called a pseudo-reference electrode because, while it forms a stable potential, the potential is not that of a well-defined thermodynamic reaction and the electrode must be calibrated with an electrode of known potential. In a two-electrode system, a single electrode, called the auxiliary electrode, can function as both the RE and the CE for the purposes of a given analytical experiment.

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Table 2. Example Electrode Reactions for Aqueous Electrolyte Amperometric Gas Sensors target gas

electrode reaction

H2 CO O2 NO2 NO H2S SO2

H2 + 2OH- ) 2H2O + 2eCO + H2O ) CO2 + 2H+ + 2eO2 + 4H+ + 4e- ) 2H2O NO2 + 2H+ + 2e- ) NO + H2O NO + 2H2O ) NO32- + 4H+ + 3eH2S + 4H2O ) SO42- + 10H+ + 8eSO2 + 2H2O ) SO42- + 4H+ + 2e-

2.6. Electrode Reactions For an example of electrode electrochemical reactions, the typical room-temperature liquid electrolyte CO sensor is considered. The CO sensor has an H2SO4 electrolyte and three Pt electrodes, typically made with high-surface-area Pt powder. The fundamental process for the electrochemical reaction is the transfer of electrons from the reacting CO and H2O to the WE surface and the formation of some H+ to transport the charge in the cell and the formation of one additional oxygen-carbon chemical bond changing the valence of the carbon atom:

CO + H2O f CO2 + 2H+ + 2e-

(5)

In order to maintain charge neutrality in the electrochemical cell, a second reaction, the reduction of oxygen to water, will occur at the CE that uses the acid and electrons to form water:

/2O2 + 2H+ + 2e- f H2O

1

(6)

The WE current generated in the electrochemical cell is related to the rate of CO oxidation at the working electrode and is directly proportional to the rate of CO arriving at the working electrode. Upon completion of the electrochemical reaction, the products must desorb from the electrode surface. Following desorption, reaction products (CO2) diffuse away from the electrode area. If the products of the reaction are sensor poisons, the sensor lifetime or response characteristics may be severely limited. Thus, the choice of materials and methods for the electroanalytical processes are critical in the sensor design. Table 2 lists electrochemical reactions for gas analysis. However, the actual electrode reaction at the electrode surface is dependent on the nature of the electrode material, the electrolyte solution, the thermodynamic potential, the interface of the electrode-electrolyte, and, of course, the analyte. To be complete, each of these analyte reactions is balanced with a second half-cell reaction that occurs at the CE, and this was discussed in the electrolyte section. The counterreaction is important to consider because it can limit the overall performance of the resulting sensor. It is typically chosen to be a fast, reversible, and noninterfering reaction. This is not always easily achieved and has limited some successful commercialization of amperometric sensors.

2.7. Analytes Many gases can be measured with an appropriately designed AGS, such as H2, O2, CO, NO2, NO, O3, SO2, H2S, and organic vapors with electroactive functional groups likes alcohols or aldehydes. Each sensor can have a unique design and a different set of materials and geometries for mem-

branes, electrolytes, and electrodes in order to take advantage of chemical properties of a specific target analyte and survive under various operating conditions. This use of a variety of materials, unique geometries or structures, and different methods is what gives the sensors their analytical properties and allows them to serve diverse applications. It is because of the discovery of unique combinations of materials, geometries, and methods that the AGS keeps improving and the widespread use of the AGS for industrial safety, biochemistry, clinical chemistry, health and medicine, agriculture, food safety, environmental protection, automotive technology, space exploration, military threat detection, and process control continues to grow. The next section presents a historical perspective on the development of AGS technology and its path to becoming one of the most widely used gas-sensing devices.

3. Survey of Old and New Amperometric Gas Sensors In general, the AGS technology results from the basic understanding of electrochemistry arising from the works of Nernst, Faraday, and Heyrovsky as well as the understanding of mass transport and electrochemical reactions provided by Fick, Cottrell, Buttler-Volmer, and Levich. Their fundamental scientific advances allow us to make sensors and understand the many dynamic aspects of the observed AGS signals.3,4

3.1. Working Electrode and Sensor Design 3.1.1. Clark Sensor The earliest successful AGS was developed by Clark et al. in 1953 for oxygen determination.5 The original Clark sensor was first introduced to measure the oxygen content in blood samples, and the major innovation was to cover the shiny platinum button WE cathode with cellophane. Unlike previously designed O2 sensors, which employed a bare platinum electrode, the Clark sensor showed enhanced selectivity for oxygen due to the presence of the cellophane membrane, which allowed O2 to penetrate but prevented interferences from red blood cells and some other gases. The success of the Clark sensor eventually resulted in the use of cellophane-covered platinum electrodes to regulate the oxygen tension in the arterial line of the dispersion oxygenator during total bypass of the heart and lungs.5 As demonstrated by Clark, a membrane diffusion barrier for the working electrode offers several advantages over a bare electrode. For instance, the membrane eliminates interferences from ions and other substances to which it is impermeable, thus improving the selectivity for the analyte of interest. Also, the membrane serves as a diffusion layer for gaseous analytes. By appropriate selection of the membrane material and size, one can control the analytical characteristics of a sensor, permitting the analysis of several analytes over a wide range of concentrations. The earliest Teflon-bonded diffusion electrodes were prepared by Niedrach and Alford6 for use in fuel cells. Later, we will also show a method to use bare electrodes with a polymer electrolyte and thereby eliminate the membrane to improve response times.7 In 1958, Sawyer et al. investigated sensing of several gases such as oxygen, nitrogen, chlorine, bromine, sulfur dioxide, nitrogen dioxide, nitrous oxide, hydrogen, carbon monoxide,

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Figure 4. SEM micrographs of KrF laser patterned PTFE membrane electrodes: (a) PTFE-bonded platinum black layer irradiated with a fluence of 500 mJ/cm2; (b) gold film irradiated with a fluence of 60 mJ/cm2; (c) the edge of the electrode as shown in (a); and (d) timedependent limiting currents of the SO2 sensors with different pattern sizes compared to the unpatterned sensor. The concentration of the SO2 gas is 998 ppm: (a) unpatterned sensor, (b) patterned sensor with the pattern size of 100 µm, and (c) patterned sensor with the pattern size of 60 µm. Reprinted with permission from Qin, Z.; Wang, P. N.; Wang, J. Sens. Actuators, B 2005, 107, 805. Copyright 2005 Elsevier Science.

hydrogen sulfide, and hydrogen cyanide with a membranecovered platinum electrode.8 The best performance characteristics were achieved using polyethylene, as opposed to saran, Mylar, and natural rubber, as the membrane material. As shown by their results, the incorporation of the membrane into the sensor system greatly enhanced the overall practicality of the sensor, leading to many new applications. However, it remained difficult to construct practical gas sensors at that time because there was no convenient way to package the electrodes for field use. Modern Clark electrodes are fitted with a porous poly(tetrafluoroethylene) (PTFE) membrane. Because of the hydrophobicity of the material, the pores (typical diameter 10 µm) are not wetted by the aqueous solution but allow the transport of dissolved gases to the electrode-electrolyte interface.9 As the mass-transfer rate of the analyte is slow compared to electron transfer, the Faradaic current is controlled by diffusion rather than the kinetics of the electrode reaction, and this assures a linear dependence of the current on concentration of the dissolved oxygen. The arrangement can be used for the measurement of other dissolved electroactive gases, most notably for the determination of chlorine in swimming pools. The layer of electrolyte solution between membrane and electrode is kept thin in order not to compromise sensitivity and response time. This design is, in principle, suitable for measurements in the gas phase, but in practice it is found that the cell is not stable because of variations in the thickness of the thin liquid film used in the Clark design, most likely because of evaporation of the water used as a solvent for the electrolyte. For the measurement of electroactive species in the gas phase, designs derived more from fuel cell technology have been

adopted. PTFE membranes can be coated with vacuumevaporated metals for improved sensitivity.10 The potential drop across any given sensor cell includes the potentials of the working electrodes and counterelectrodes and the potential over the galvanic cell’s internal resistance. For a given material of working electrode, the sensitivity of the galvanic cell depends on the surface area on which the electrochemical reaction takes place.11-15 The assembly must provide as large a surface area as possible for the threephase boundary (TPB ) interface between the gas, the electrode, and the electrolyte), because the electrode reaction is assumed to take place primarily in this region. The particular “hair morphology” of a catalyst surface, e.g., nanostructures, can dramatically increase the three-phase boundary area, followed by an enhancement of the sensor’s sensitivity. A further improvement of the sensitivity from ppm to ppb gas concentration levels can be obtained using the new membrane-electrode assembly composed of a PTFE membrane and a nanocomposite material of carbon nanotubes and PTFE. Carbon nanotubes can greatly increase the surface area of the TPB because the carbon nanotubes have very high surface area, e.g., single-walled carbon nanotubes can achieve 1600 m2/g Brunauer-Emmett-Teller (BET) value.16,17 The catalyst is deposited electrochemically on the membrane in order to get a nanostructured surface and to increase the TPB area.18-20 Wang’s group developed a method to pattern the PTFE membrane electrodes by KrF laser ablation21 (see Figure 4). Patterning of the sensing electrode can be wellcontrolled and is expected to increase the sensitivity of the sensor because the sidewall of the patterned catalyst layer can be considered as the effective TPB. Compared with the unpatterned sensor, the sensitivity of the sensor with a pattern

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size of 60 µm was improved by 14 times, and the response time is reduced half. The dramatic increase of the limiting current and the shortening of the response time were obtained from the increase of the TPB area and the rapid gas diffusion through the membrane electrode. Other nanostructures are used for this purpose as well, such as composites of Gd0.7Ca0.3CoO3-δ and Ce0.8Gd0.2O1.9 that have potential applications as oxygenpermeable membranes in an amperometric sensor for NOx detection in exhaust gases.22 In addition, a series of manganites was investigated for membranes with the composition A0.7E0.3MnO3 (A ) Gd, Y, and Pr and E ) Ca and Sr).23 The presence of slow diffusion-coupled redox processes could be detected in all manganites at low and high oxygen activities vs an air reference electrode. The oxygen ion conductivity was considerably lower than 10-5 S cm-1 in all investigated manganites. Nevertheless, the comparison of results for glass-encapsulated and nonencapsulated microcontacts showed that Pr0.7Ca0.3MnO3 and, in particular, Pr0.7Sr0.3MnO3 are quite active as catalysts for the oxygen electrode reaction at the three-phase boundary air/zirconia/ manganite, whereas Y0.7Sr0.3MnO3 does not catalyze the oxygen electrode reaction significantly at the surface. Studies have been carried out on the selectivity, mass transfer rates, and effects of temperature on PTFE films as diffusion barriers.24 The advent of using nanostructures has made a significant impact on the design of the modern AGS as it has on other sensors.19

3.1.2. Diffusion Electrodes A major achievement in the evolution of the modern AGS occurred in 1965 with the development of the gas-diffusion electrode by Niedrach.6 The original Niedrach electrode consisted of a sintered Teflon-Pt catalyst covered with a porous Teflon film. Its performance was evaluated in hydrogen-oxygen and hydrogen-air fuel cells at ambient temperature in 6N KOH and 5N H2SO4. In both cases, it was shown that water was efficiently retained in the electrolytic solutions and did not leak into the gas fuel feed stream. The Niedrach electrode provided a model for the contemporary gas-diffusion electrode, being “semihydrophobic” and allowing an open porous electrode of the finely dispersed metal catalyst yet not allowing complete penetration by the electrolyte. These electrodes are prepared by mixing Teflon emulsion with a high-surface-area noble metal catalyst powder and then depositing the slurry on a metal wire screen or on the surface of a totally hydrophobic porous Teflon film. The electrochemical cell that is based on a gasdiffusion electrode composed of a porous PTFE membrance and a porous electrode is shown in Figure 5. The resulting gas-diffusing electrode consists of highly interlocked matrices of gas pores, electrolyte channels, electronically conducting paths, and electrocatalytic surfaces. It is porous enough to affect efficient gas permeation, has sufficient metal catalyst to be a good electronic conductor, and is hydrophilic enough to make intimate contact with the electrolyte for ionic conduction and facilitation of electrochemical reactions involving gases. In contrast to the Clark electrode, gas-diffusion electrodes based on back-side metallized porous membranes are not affected by evaporation of water because the porous electrode is directly in contact with the bulk of the electrolyte solution. The mass transfer of analyte from the sample to the working electrode can be faster, resulting in shorter response times

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Figure 5. Schematic of a cell based on a gas-diffusion electrode (GDE) composed of a porous PTFE-membrane and a porous electrode. Reprinted with permission from Knake et al., Analytica Chimica Acta, 2005, 549, 1. Copyright 2005 Elsevier Science.

and higher currents that lead to higher sensitivity. Also, the real surface area of the electrode can be several orders of magnitude higher, allowing species with relatively poor electroactivity to produce measurable currents. Therefore, they are especially valuable in those cases when the kinetics of the electrochemical reaction are very slow, such as the oxidation of hydrocarbons and CO. Examples of the earliest of these sensors reveal excellent practical sensor characteristics.25-28 A kinetic study of NO gas at the Pt/Nafion electrode with application to the amperometric NO and NO2 sensing has been carried out by Ho’s group.29,30 Sensors based on metallized porous membranes and porous catalytic electrodes have been produced from several manufacturers that are widely used in industrial applications.1,2 On the other hand, relatively few publications on this specific structure of electrode have appeared in the scientific literature in the past decade,31-33 indicating the maturity of this type device. The fuel-cell type electrode was never utilized in an AGS until, in 1969, a breakthrough in design combined it with a Pt/air reference electrode to achieve the practical ethanol and CO sensor design introduced by Oswin, Blurton, and coworkers.26 This new design was used to detect various toxic gases, such as ethanol, CO, H2S, NO2, and NO. The sample (an electrochemically oxidizable or reducable species) enters the cell through porous electrodes similar to those used in a fuel cell (see Figure 6);34 hence, the device has been incorrectly called a fuel-cell sensor, although its purpose is not the conversion of chemical to electrical energy. The relationship between the cell current and the concentration of chemical species of interest present in a more or less constant matrix can be established empirically and optimized to be linear in most cases over a fairly broad range of concentrations. In the electrochemical cell that becomes the sensor of Oswin and Blurton,26 the gas was detected from the current generated by the electrochemical reaction at a fixed potential of 0.9-1.5 V vs SHE. This avoided the generation of undesired current from reactions involving the oxygen-water redox couple within the cell. The original Oswin/Blurton cell significantly simplified sensor design and provided a model for several two- and three-electrode sensors that were eventually introduced in the 1970s including ethanol,26 CO, and other gases.28 Recently, a high-temperature solid electrolyte amperometric CO2 sensor was fabricated using a Pt/NASICON (Na+ conductor, Na3Zr2Si2PO12)/Pt cell and a porous Na2CO3-

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Figure 6. Fuel-cell-type amperometric sensor.

BaCO3 auxiliary layer.35 A platinum-loaded zeolite electrode was developed for high-temperature measurement of total NOx in combustion environment.36 A Pt-loaded Y (PtY) is a catalyst electrode for NOx equilibration with electrochemical oxidation of NO on a yttria-stablized zirconia (YSZ) electrolyte. By applying a low anodic potential of 80 mV, the NO in the NOx equilibrated mixture can be oxidized at a Pt working electrode on the YSZ electrolyte at 500 °C. A novel reference electrode for a NOx sensor was developed that did not require air, and this makes the design simpler and more robust.37 With the advent of nanotechnology, the use of nanostructured materials in fuel-cell sensors to enhance the sensor performance becomes ever more popular. The details will be discussed in the following section.

3.2. Electrolytes In addition to the progress made in the development of electrodes and sensor geometries, the 1970s witnessed improvements in the use of different electrolytes. Prior to 1970, electrolyte solutions for amperometric sensors were mainly aqueous in nature. Even nowadays the aqueous solutions are still used as the electrolytes for gas sensing, e.g., acid or halide-halate electrolytic solutions are used for acidic gas38 and for other gas-detection applications.39 As illustrated in the preceding examples, aqueous electrolytes are effective in many amperometric gas sensors. However, they are also limited in their specificity, service life, operating temperature range, and electrical potential range. As a result of the problems and limitations associated with aqueous electrolytes, sensor research began to focus on the design and development of alternative electrolytes. By the middle 1970s, nonaqueous and solid polymer electrolytes emerged for the AGS.6,40-48 As nanotechnology becomes more prevalent and many nanostructures are developed for use as working electrodes, parallel studies of nanosturctures for use as the electrolyte are also being conducted.

3.2.1. Nonaqueous Electrolyte One of the earliest amperometric sensors employing a nonaqueous electrolyte was described by Schneider in 1978.47 The novel electrochemical cell, used for the detection of chlorine, contained a nonaqueous electrolyte composed of

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lithium or sodium perchlorate as supporting electrolyte dissolved in an organic solvent selected from the group consisting of γ-butyrolactone, sulfolan, or propylene carbonate. All these solvents have a high boiling point (>200 °C) and possess a high anodic limit with the Pt electrode. This sensor was superior to previously designed chlorine sensors, which frequently failed because the electrode materials selective to chlorine often dissolved in the aqueous electrolyte medium. The new sensor was effective in detecting chlorine down to the ppm range, had a fast response time of <1 min, and had an extended useful life of >1 year. A word of caution must be added that the use of perchlorates in organic solvents always represents an explosion hazard, particularly when the assembly is, inadvertently, allowed to dry up or be exposed to high temperatures. Similarly, Venkatasetty described the invention of a carbon monoxide sensor incorporating a nonaqueous electrolyte.44 This sensor, based on an aprotic electrolyte system, was also used to amperometrically detect toxic gases such as nitrogen oxides, SO2, and H2S. Other electrochemical cells employing nonaqueous electrolytes were also introduced during the same time period. These systems were used to fabricate amperometric sensors for several gases such as methane,45 hydrazine, and carbon dioxide.47

3.2.2. Solid Polymer Electrolytes Solid polymer electrolytes became important during the mid-1970s because of the inefficiencies and maintenance requirements of liquid electrolytes. Originally, a solid polymer electrolyte (SPE) was described as a solid plastic sheet of perfluorinated sulfuric acid polymer that, when saturated with water, became an excellent ionic conductor. Nafion, a typical solid polymer electrolyte, is a hydrated copolymer of poly(tetrafluoroethylene) (PTFE) and polysulfonyl fluoride vinyl ether containing pendant sulfonic acid groups (Dupont). Nafion has good proton conductivity, high gas permeability, and chemical stability, and it has been widely used in chloralkalai, fuel-cell, and sensor applications. In the Nafion-based AGS design, Nafion film is used both as the electrolyte and as a support for the electrode structure. The ionic conductivity results from the mobility of the hydrated hydrogen ions that move through the polymer sheet by passing from one sulfuric acid group to another.19 An early amperometric gas sensor incorporating a solid polymer electrolyte was described by LaConti and later improved by Sedlak et al. The sensor, originally used to detect carbon monoxide,43 employed Nafion. Being a cation-exchange membrane, Nafion permitted the passage of cations and blocked the passage of anions. Thus, the hydrogen ions produced at the working electrode through oxidation of carbon monoxide were transported through the ion-exchange membrane to the counterelectrode, where they reduced oxygen with the addition of electrons to produce molecular water. In addition to its excellent ion-exchange capacity, the Nafion electrolyte exhibited other positive features, such as high structural stability, resistance to acids and strong oxidants, and good thermal stability. As a result of its useful characteristics, Nafion has been used to detect gases other than carbon monoxide,49,50 such as CO2 and NH3,51 NOx,52,53 hydrogen,54-56 H2S,57 SO2,58 O2,59,60 O3,61,62 and alcohol vapors.63 The only limitation for field use was the issue of the water in the Nafion freezing at low temperatures. In addition, Kuwata et al. illustrated two different amperometric oxygen sensors using a Nafion membrane opera-

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tive at room temperature.64,65 Also, Alberti and Casicola presented an amperometric three-electrode sensor for CO operating at room temperature and using thin layers of R-zirconium phosphate as a solid electrolyte.41

3.2.3. Ceramic Solid Electrolytes at Elevated Temperature The use of ceramic solid electrolytes in amperometric gas sensors at elevated temperatures allows for the detection of different gases under harsh conditions. There is an increasing need for rugged and reliable gas sensors capable of monitoring carbon monoxide, nitric oxides, sulfur oxides, and hydrocarbons from automotive exhausts. Because of such harsh environments, e.g., high temperature and high pressure, aqueous electrolytes, liquids, or polymeric materials cannot be used in sensor designs. Ceramic solid electrolyte sensors are typically designed for high-temperature applications such as stack gas, combustion process, and automotive engine exhaust. The solid electrolyte provides ionic conduction and also supports the electrodes. The interface between the solid electrolyte and the electrode plays an important role in the resulting gas sensor’s analytical characteristics. Liu and Weppner described amperometric gas sensors using solid electrolytes such as NASICON, cubic stabilized zirconium oxide, and polycrystals of tetragonal zirconium oxide to selectively detect the partial pressures of several gases.66 An AGS incorporating solid electrolytes made from rare earth elements was also designed for high temperature, and Stetter and Cao67 described an amperometric sensor for the detection of halogenated compounds such as chloropentafluorobenzene, chlorobenzene, bromobenzene, and iodobenzene. This sensor, which used a bead of sodium lanthanum fluoride silicate as the solid electrolyte, was shown to operate around 550 °C.67 Similarly, Coillard et al. discussed a sensor employing commercial zirconia lambda gauge for the amperometric detection of NO in automobile exhausts at high temperature.68 As a result of their versatility, solid electrolytes have become important components of amperometric gas sensors. An amperometric NO sensor using a LaGaO3 solid electrolyte was investigated, and the response was based on the difference of catalytic activity of the electrodes.69 It was found that the oxygen pumping current increased upon exposure to NO when Sr0.6La0.4Mn0.8Ni0.2O3 and La0.5Sr0.5MnO3 (LSM-55) were used as active and inactive electrodes for NO oxidation, respectively. Compared with various common electrolytes such as YSZ or Sm-doped CeO2, it was clearly demonstrated that much higher sensitivity in the amperometric mode is achieved by using LaGaO3-based fast oxide-ion conductors. The fuel-cell configuration was used for the investigation of three types of metallic electrodes for NO detection in the exhaust gases.70 Their results show that the cell defined as RhPtAu/YSZ/ RhPtAu gives a relatively high NO sensitivity of ∼14 nA/ppm and a low O2 sensitivity of 0.02 nA/ppm. The availability of electrode materials with selectivity opens up possibilities for the development of highly sensitive amperometric NO gas sensors with a single working electrode, i.e., the NO concentration in O2 and NO containing gas mixtures can be directly measured without oxygen pumping.

3.3. MEMS and Nanotechnology Microelectrodes (MEMS), especially micro-working electrodes, with a very small electrode surface area, have been used for amperometric gas sensors. In recent years, MEMS

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technologies have increasingly been employed to fabricate microelectrode structures. The microelectrodes have shown several advantages of high sensitivities, fast response speeds, small sample amounts, and well-defined and reproducible geometries. In contrast to the development of sensor parts, such as membranes, electrodes, and electrolytes, which occurred during the 1960s and 1970s, amperometric gas sensor research of the 1980s focused on the production of novel sensor designs. One of the highlights of the 1980s in relation to analytical chemistry was the introduction of microsensors. Microfabrication of electrochemical devices has numerous advantages over standard fabrication procedures. These advantages include precise reduced sensor size, reduced cost, smaller sample size, faster response time, higher concentration sensitivity, well-defined geometrical features, and potential for mass production. These advantages, however, must be obtained without degradation of the signal-to-noise ratio as the sensor size is reduced. Progress in the development of microamperometric sensors was slow when compared to the production of micropotentiometric sensors. Microamperometric sensors of the early 1980s consisted only of microfabricated electrodes on a suitable substrate. These electrodes were then coated with the liquid electrolyte solution that also carried the sample to the electrode area for analysis. An example of an early microamperometric sensor was presented by Sleszynski and Osteryoung in 1984.71 The given sensor, which used very small electrodes constructed from nonconductive epoxy and reticulated vitreous carbon (RVC), was shown to yield desirable results as compared to standard fabricated sensors. However, the goal of constructing complete, practical, and commercially successful amperometric gas sensors using the microfabrication approach has still not been achieved. During the late 1980s, microfabrication technology for the construction of amperometric sensors was investigated with the introduction of novel sensor electrode designs. In 1988, Maclay and co-workers72,73 introduced a series of Nafionbased microfabricated amperometric gas sensors using gold sensing electrodes in the shape of a square grid. The newly designed sensors were evaluated by comparing their analytical characteristics with those of conventional sensors. The response time of the miniaturized sensors was >1 order of magnitude faster than conventional sensors, although they had a lower sensitivity. This work also elucidated the fact that the sensitivity of the device depended not only on the chemical nature of the electrode surface but also on the specific structure of the electrocatalytic surface and the interface created by gas/electrode/electrolyte. In an effort to improve the sensitivity of microamperometric sensors, Buttner et al.73 constructed devices with an integrated design in which the working electrodes, counterelectrodes, and reference electrodes were photolithographically etched onto a gold-coated silicon oxide surface of a silicon wafer and spin-coated with a thin film of a solution of Nafion. The working electrode of the sensor was an ultrafine square grid with evenly spaced regions of gold and holes every 50 µm. The signal observed for the sensor following exposure to gases such as H2S and NO greatly exceeded the response that would be expected on the basis of simple geometric considerations of microelectrodes.73 The analytical characteristics of microamperometric sensors can also be improved by varying the sensor components, such as electrolyte composition. For example, Hauser et al.31

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recently described the preparation of a carbon dioxide sensor using mixed organic solvents as the electrolyte. As shown by the results of this study, a mixed acetonitrile/dimethyl sulfoxide solvent enhanced the overall performance of the sensor by producing a large potential window over which the CO2 reduction currents can be measured.31 As illustrated by these examples, advances in microtechnology during the 1980s have resulted in the production of newly designed amperometric gas sensors with characteristics superior to those of conventional sensors. More recently, gold microelectrodes are used with and without membranes in nonaqueous media for oxygen, carbon dioxide, and nitrous oxide detection simultaneously over a wide concentration range.74,75 It is believed that, by using sufficiently small microelectrodes, it should be possible to effectively “outrun” the reaction kinetics by the much faster loss of species from the electrode/electrolyte interface by diffusion, thereby giving rise to independent oxygen and carbon dioxide signals. Hahn and co-workers also conducted the microelectrode studies of isoflourane vapor, an inhalation anesthetic agent, and oxygen mixtures in dimethyl sulfoxide.76 However, such sensors never demonstrated the long lifetimes required by field applications nor the low cost and, thus, have not been commercialized for popular use. Very recently, nanosize working electrodes using carbon nanotubes have been reported. The advantages of such microsensors are small in size, low cost, low power, high signal-to-noise ratio, and low detecting limit. For the discussion of using the nanostructures for making the microsensors, a sensor array consisted of various highly sensitive carbon nanotube-based sensors developed by Li’s group are described here, such as sensors using pristine carbon nanotubes,77 carbon nanotubes loaded with metal clusters,78,79 and carbon nanotubes coated with polymers.80 An interdigitated electrode (IDE) was utilized as a transducer, with many of them on a chip that was fabricated on a silicon wafer.81 This electrochemical sensor depends on the transfer of charge from one electrode to another electrode. The gases and vapors introduced to this sensor array are NO2, HCN, HCl, Cl2, acetone, and benzene in parts per million (ppm) concentration levels in dry air. These are toxic gases and vapors that are of interest to both military and civilian personnel in defense applications, as well as in industry process and environmental monitoring. The detection limit of carbon nanotube sensors can achieve the ppb concentration level.82 Carbon nanotubes can be coated with metals to prepare hydrogen sensors83 or with Nafion to make RH sensors.20 For many years, it was generally reported that potentiometric sensors were more easily microfabricated because the signal size did not depend at all on the electrode size while amperometric sensor currents were directly proportional to electrode size. An ideal use of nanostructures to create large surface area in small-sized devices is found here. With the use of nanostructures, the electrode area of the amperometric sensor can be reduced by orders of magnitude while preserving the sensor signal magnitude, and this is a perfect example of the importance of nanostructures in the design and construction of the AGS. Several other designs for amperometric gas sensors were also introduced during the 1990s. For example, Wallgren and Sotiropoulos reported84 the construction of an oxygen sensor based on the planar sensor design. The design consisted of both the working electrodes and the counter-

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electrodes being vacuum-deposited as nonporous Au layers on the same face of a Nafion electrolyte and in contact with the gas sample. Potential advantages of this design include faster response times and higher sensitivities per useful electrode area due to the high mass transport rates of the gas to the sensing electrode since no membrane or significant liquid film barrier is present. Also, smaller quantities of precious metal catalyst could be used for the working electrode, and the labor-intensive stacking steps required for the fabrication of conventional-type designs are avoided. As shown by the results of this research, oxygen reduction leads to an exponential current rise over a wide potential range, indicating very high mass transport rates and implying that the electroactive gas reacts at the line formed by the gas/ solid electrolyte/metal layer interface.84 However, the practicality of the design may be questioned because, in practice, if the water content of the membrane changes, so do the dimensions, and this can lead to stresses that will fracture the electrodes and alter the response. Berger and Edelman also described the production of a planar sensor used to detect the partial pressure of oxygen in blood samples.85 In this sensor, a semipermeable membrane, comprising an acrylonitrile butadiene copolymer or an acrylate-based copolymer, was provided to act as an impermeable barrier for ions and blood constituents other than oxygen.85 In 2004, Meyerhoff and co-workers reported a planar amperometric nitric oxide sensor based on platinized platinum anode.86 A microporous poly(tetrafluoroethylene) gaspermeable membrane was used. Platinization of the working platinum electrode surface dramatically improves the analytical performance of the sensor by providing ∼10-fold higher sensitivity (0.8-1.3 pA/nM), ∼10-fold lower detection limit (e1 nM), and extended (at least 3-fold) stability (>3 d) compared to sensors prepared with bare Pt electrodes. In this article, both experimental results and theory of NO measurements as a function of sensor diameter and distance from a point source were discussed in the context of an investigation of monitoring NO release from diazeniumdiolate-doped polymeric films. By modifying the above-discussed NO amperometric gas sensor with thin hydrophilic polyurethane films containing catalytic Cu(II)/(I) sites, the direct amperometric detection of S-nitrosothiol species (RSNOs) is realized.87 Catalytic Cu(II)/(I)-mediated decomposition of S-nitrosothiols generates NO(g) in the thin polymeric film at the distal tip of the NO sensor. Another type of planar thick-film sensor with centrosymmetric diffusion geometry for detection of hydrocarbons in oxygen, nitrogen, and hydrocarbon gas mixtures has been developed for monitoring in exhaust gases.88 In this work, propene (C3H6) is chosen for sensor tests as a model hydrocarbon. The solid-state electrochemical cell based on yttria-stabilized zirconia (YSZ) operates in the amperometric mode. Oxygen is pumped out at the Au/YSZ electrode, and C3H6 is oxidized at the Pt/YSZ electrode (see Figure 7). At a gas temperature of 600 °C, the sensitivity is 2 nA/ppm C3H6 for this high-temperature amperometric gas sensor design.

3.4. Sensor Array The 1980s also witnessed the microcomputer revolution, which, coupled with the advent of chemometrics and available AGSs, encouraged the introduction of arrays of gas sensor systems.89-93 While the earlier research occurred in

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Figure 7. Schematic of the sensor design: (A) one side of the sensor showing the counterelectrodes upon the YSZ substrate, (1) Au/YSZ ring electrode (RE) and (2) Pt/YSZ point electrode (PE); (B) back side of the sensor, showing the diffusion barrier covering the working electrodes; and (C) propene characteristics of the Pt/YSZ point electrode at various oxygen concentrations. Reprinted with permission from Schmidt-Zhang, P.; Guth, U. Sens. Actuators, B 2004, 99, 258. Copyright 2004 Elsevier Science.

the early 1980s93,94 in conjunction with a K-nearest neighbor or KNN algorithm for pattern recognition, a very recent development features an amperometric sensor array for a gas mixture of NOx, CO, O2, and SO2 detection using Nafion/Pt (or Au)/ceramic plate microelectrodes prepared by microfabrication technologies as the sensing electrodes and a backpropagation neural network (BPN) as the patternrecognition algorithm.95,96 The limiting current potentials for anodic oxidation of NO on Au, anodic oxidation of SO2 and CO on Pt, and cathodic reduction of NO2 on Au and O2 on Pt were applied to each of the sensors in the array, respectively. The system was optimized for identifying these gases with their concentrations in the mixture. Sensor array technology has been utilized in chemical detection of gases and is sometimes known as an electronic nose97,98 and, when used in liquids, is sometimes know as the electronic tongue.99 However, the instrument using a sensor array as a detector has not been fast to get into the commercial analytical market and remains in the stage of a new technology in various application areas. One of the main reasons for the slow introduction of arrays into the market is the difficulty in calibration and in the isolation of the analyte response from the matrix response. Recent conferences, reviews, and books have chronicled progress in sensor arrays incorporating all types of sensors.90 A recent renaissance has occurred and provided new interest in sensor-array-based analysis. Much of this interest has to do with important applications like the detection of lung cancer, recently shown to be possible with a gas sensor array.100 While the relationship of the sensor array signals to lung cancer is undisputed, there is something unsettling about an analytical technique for which the molecular basis for the response is clearly not well-understood if understood at all. Statistically, while we expect the analysis to be somewhat robust and to apply to unknowns, the statistical basis for this can be calculated and the molecular basis is only implied. Thus, the sensor array work often blurs the line between functional dependence and association of a signal and complex endpoint. It will take progress in multivariate calibration, the construction of multivariate models like the net analytical signal recently used for NIR determination of glucose, or the development of another robust

method in order to gain widespread analytical acceptance for routine analytical determinations. What the sensor array has always been good at is the comparison of two complex mixtures, one a standard and one sample or a mixture of standard and a sample to the standard. In analysis, the sensor array can be a highly valuable tool for determination or comparison of complex nonmolecular endpoints like taste, odor, or flavor, and the AGS was one of the earliest contributors.93,94 More comparisons of different analytical myths and operational aspects of the electronic nose can be found89 for practical applications.

4. AGSs For Practical Applications The modern AGS can be considered one of the primary methods for field detection of CO in air for industrial hygiene applications since the 1970s, when it was introduced to replace the hopcalite catalyst method.101 The 1970s through the 1990s also brought about many additional applications of amperometric sensors in field analytical use.102 The AGS provides ppm level sensitivity, stability, low cost, low power, good selectivity, and fast response all at the same time, and this makes for a powerful and versatile sensor. Also, because the analysis is “on the spot”, samples do not have to be preserved or returned to a lab for a slow and costly analysis. By manipulating the standard three-electrode potentiostat arrangement of amperometric gas sensors, one can design an amperometric detector for various target analytes and conditions in field analysis. Hauser et al. recently described a simplified amperometric detector for capillary electrophoresis. This system employed only two electrodes in total to analyze catecholamines, ascorbic acid, carbohydrates, and heavy metals. Detection was carried out with a single working electrode by using an electrophoretic counterelectrode as a polarized pseudo-reference electrode with an appropriate electronic circuitry. As shown by the results of this study, the amperometric detector achieved similar performance characteristics as compared to the traditional amperometric gas sensor.31 The amperometric gas sensor has a long history of development and some of them have produced mature for field applications. With the advent of microfabrication

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Figure 8. (A) Schematic cross section of a typical amperometric oxygen sensor with a channel-type diffusion barrier and (B) principle of an amperometric two-electrode cell for the simultaneous detection of oxygen and NO. Reprinted with permission from Reinhardt, G.; Mayer, R.; Rosch, M. Solid State Ionics 2002, 150, 79. Copyright 2002 Elsevier.

technology, computational power in a PC, and nanotechnology, improved sensors with better working electrodes, different types of electrolytes, and miniature design have evolved, and this has expanded the spectrum of real-world applications.103 The challenges are common to all applications including the following: inexpensive fabrication steps, sensor robustness, reliability, reproducibility, low power consumption, and easy system integration.

4.1. Industry AGSs possess high sensitivity, small size, and low power consumption, which can be used to quickly verify incoming raw materials at the delivery point. The AGS technology can significantly reduce the amount of time and money spent analyzing materials in a lab, as well as reduce the amount of material handling. Most changes in chemical processes can be reflected in the changing composition of the vapor phase surrounding or contained within the process. Thus, vapor-phase sensors can enable the quick assessment of the chemical status of most industrial processes. Examples are found across many sectors including food processing (coffee roasting and fermentation), petrochemical (plastics manufacture and gasoline blending), and consumer products (detergents and deodorants). Much like vision inspection is used to assess the visual integrity (color, shape, and size) of products, olfactory inspection assesses the chemical integrity (consistency and presence of contaminants) of products.1,2,89,91,109,110,111

4.2. Environment Increasing awareness of new regulations for safety and emission control makes environmental monitoring one of the most important applications for reliable gas sensors. Current methods for air quality analysis typically require the use of very costly and bulky analytical equipment. For applications in environmental monitoring or indoor air quality measurement, sensors that are able to selectively detect various gases at concentration levels ranging from a few ppb to hundreds of ppm are valuable. New developments in MEMS sensors are trying to address these difficult measurement tasks.103 With unique combined advantages of high sensitivity, small size, and low power consumption, the amperometric gas sensor is uniquely suited for environmental applications, but it is often challenged with sensitivity or selectivity demands not encountered in other applications. For example, the amperometric sensors based on stabilized zirconia have been developed for the detection of oxygen-containing gases like NOx or combustible gases such as CO, H2, and hydrocarbons;104 for NOx detection;105 for lean combustion gas control;106 and for selective detection of propane.107 Multielectrode amperometric sensors are able to detect O2 and NO or O2 and combustibles at the same time, and such a

multielectrode cell is illustrated in Figure 8. The basic idea is to adjust geometric and working parameters in such a way that, at each electrode, essentially one-electrode reaction takes place. Selectively reacting electrodes are desired to improve sensor performance. Recently, SrTi0.65Fe0.35O3-δ (STF-35) has received considerable attention for use as a resistive oxygen sensor because of its unusual temperature-independent conductivity above 700 °C and pO2 > 1 Pa.108 This material has the potential for automotive oxygen sensor applications. Combined with physical sensors for temperature, anemometry, pressure, and humidity monitoring, an AGS system can be built to get comprehensive information for an environment simultaneously. The Indoor air quality application is just now beginning to emerge as exposure to environmental pollutants in the air is being suspected for many human health issues. The Exposure Biology program at the National Institutes of Health addresses many of these concerns. The application to measure trace levels of pollutant exposure of children and the population at large will require new sensors for measurements as well as new modeling tools (http://www.gei.nih.gov/ exposurebiology/index.asp). This need for personal measurement sensors at ppb levels for specific pollutants represents a new challenge that is ideally met with the AGS since it has all of the important logistical features of low power, light weight, small size, and low cost. However, the challenge is obtaining increases in sensitivity and selectivity and often response time and stability in the AGS for analytes such as ozone, CO, NOx, and others.

4.3. Defense and Homeland Security Chemical sensors like the AGS are suitable for security and defense applications because of their portability and low power consumption. AGS potentially can offer higher sensitivity, specificity, and lower power consumption for toxic gas detection. Some examples include monitoring filter breakthrough (e.g., toluene concentration inside the mask), personnel badge detectors (e.g., toxins accumulation), embedded suit leak-detection sensors (e.g., soldiers’ suits for upcoming warfare agents), and other applications. Additionally, a wireless capability with the sensor can be used for networked mobile and fixed-site detection and early warning systems for military bases, work facilities, and battlefield areas. Explosives detection at the airport is recently a priority application for Department of Homeland Security (DHS), and sensors or detectors will require quick detection and a low false-alarm rate. One type of the liquid explosive hazard that is easily detected by amperometry is hydrogen peroxide. The challenges will be to obtain the stability, reproducibility, and low maintenance of the fielded sensors to achieve quick screening and early warning at the airport. The detection of mustard gas and HCN are two specific agents that can be

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addressed with the AGS, as well as TNT, when the electrochemical sensors are used in arrays.109-116

4.4. Medical/Biomedical Chemical sensors will provide physicians with a quicker and more accurate diagnostic tool. Applications will include breath analysis that can provide objective information on the identity of certain chemical compounds such as in exhaled air and in the excreted urine or other body fluids that are directly related to specific metabolic conditions, certain skin diseases, or bacterial infections, such as those common to burn wounds. Additionally, the chemical sensors may also provide more accurate, real-time patient monitoring during anesthesia administration. An amperometric enzyme/gas-diffusion electrode is developed that combines a porous air-breathing gas-diffusion electrode with an enzyme immobilized on its catalytic layer. The oxygen needed for the enzymatic reaction is supplied to the reaction zone of the electrode in the gas phase through the gas-supplying layer of a gas-diffusion electrode. The performance of a gas-diffusion/enzyme electrode operating in oxygen-containing solutions is practically independent of the concentration of the dissolved oxygen.

5. Summary Amperometric gas sensors have a long and rich history. The present review shows that AGSs are an important member of the electrochemical class of chemical sensors. Many research and development efforts have resulted in practical life-saving sensors, such as the AGS sfor CO, H2S, NO2, O2, or other electrochemically active gases that are now routinely used in industrial hygiene and safety monitoring of people and workplaces. There is an ever-expanding list of analyte targets for the AGS, and its use for electroactive materials like TNT will challenge the sensitivity of the amperometric approach. With the advent of arrays, the AGS is contributing to the detection of a diverse set of complex endpoints including toxicity, cancer detection, and off-odor of materials. New applications such as homeland security and defense against threats have arisen and will challenge the sensitivity and selectivity of the electrochemical technique in general and the amperometric approach specifically. As we demand ever-lower detection levels and since the sensor signal in amperometry is directly proportional to the electrode area, the challenge is to maintain and improve the signal/ noise ratio for the analyte. However, the challenge of amperometry is to utilize the ability provided by electronics to detect smaller and smaller amounts of charge/current and to use nanostructures to continue to increase the electrode reactive surface area even with less material used in the electrode. The advancement of MEMS technology and the fast pace of nanotechnology will no doubt enable new AGS designs and materials such that new applications are found for large gas molecules and weakly electroactive gas molecules in applications that will analyze for chemical threats, toxins, food flavors, and fragrances. Much of the modern AGS work focuses on new designs that incorporate microfabrication and nanofabrication to achieve smaller size, low power, lower cost, and portable sensors and sensor arrays with intelligence. New nanomaterials development changes the building blocks that will provide well-organized nanostructures with high surface area, high chemical reactivity at lower temperature,

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good mechanical strength, and better thermal stability, leading to new catalysts for selectivity, new electrolytes for higher-temperature operation, multiple working electrodes for self-amplifying sensors, and combinations with bioanalytical approaches for biosensors and enzyme-based sensors. MEMS technology and nanotechnology combined with new computational power brighten the future of the AGS and its use within analytical chemistry and especially in field analytical measurements. We now begin a new era of AGS development using new materials, new time-resolved and spatially resolved sensor array approaches, and smaller sensor devices. It is clear that the utility of amperometry will continue and will preserve its place in gas-sensor applications now and in the future37 and will continue to serve for the betterment of human health, safety, and the environment.

6. Acknowledgment The authors would like to thank Sheng Yao for help with some of the illustrations used herein and to thank our colleagues because, without their work, this review could not have been written. Also, we express regret if some work has not been included as it is only possible to include a fraction of the work in this field.

7. References (1) Stetter, J. R.; Butner, W. J.; Cao, Z. Electroanalysis 1992, 4. (2) Stetter, J. R.; Chang, S. C.; Cha, C. S. Talanta 1993, 40. (3) Bard, A. J.; Faulkner, L. R. Electrochemical Methods: Fundamentals and Applications, 2nd ed.; John Wiley & Sons: New York, 2001. (4) Kissinger, P. T.; Heinemann, W. Laboratory Techniques in Electroanalytical Chemistry, 2nd ed.; Marcel Dekker Inc.: New York, 1996. (5) Clark, L.; Wolf, R.; Granger, D.; Taylor, Z. J. Appl. Physiol. 1953, 6. (6) Niedrach, L. W.; Alford, H. R. Chem. Abstr. 1965, 62, 11416c. (7) LaConti, A. B. U.S. Patent 4,227,984, 1979. (8) Sawyer, D.; George, R. S.; Rhodes, R. C. Anal. Chem. 1959, 31. (9) Knake, R.; Jacquinot, P.; Hodgson, A. W. E.; Houser, P. C. Anal. Chim. Acta 2005, 549, 1. (10) Wienecke, M.; Bunescu, M. C.; Pietrzak, M.; Deistung, K.; Fedtke, P. Synth. Met. 2003, 138, 165. (11) Parsonage, E. E.; Debe, M. K. U.S. Patent 5,333,8430, 1994. (12) Thomas, J. H.; Ziegler, R. J.; Debe, M. K.; Haugen, G. M.; Steinbach, A. J. U.S. Patent 5,879,827, 1999. (13) Lauer, J. M.; Akmal, N. Patent EP0777121, 1997. (14) Ziegler, P.; Demisch, U.; Pahlke, J. U.S. Patent 5,733,436, 1998. (15) Lee, S. B.; Maclay, J. G.; Keyvani, D. U.S. Patent 5,716,506, 1998. (16) Cinke, M.; Li, J.; Chen, B.; Cassell, A.; Delzeit, L.; Han, J.; Meyyappan, M. Chem. Phys. Lett. 2002, 365, 69. (17) Stetter, J. R.; Hesketh, P. J.; Hunter, G. W. Interface 2006, 15 (1), 66. (18) Hong, Y. J.; Oh, S. M. Sens. Actuators, B 1996, 32, 7. (19) Stetter, J. R.; Maclay, G. J. AdVanced Micro and Nano Systems; Baltes, H.; Brand, O.; Fedder, G. K.; Hierold, C.; Korvink, J. G.; Tabata, O., Eds.; Wiley-VCH Verlag-GmbH and Co.: Weinheim, Germany, 2004; Vol. 1, p 357. (20) Star, A.; Han, T. R.; Joshi, V.; Stetter, J. R. Electroanalysis 2004, 16 (1-2), 108. (21) Qin, Z.; Wang, P. N.; Wang, J. Sens. Actuators, B 2005, 107, 805. (22) Nigge, U.; Wiemhofer, H. D.; Romer, E. W. J.; Bouwmeester, H. J. M.; Schulte, R. R. Solid State Ionics 2002, 146, 163. (23) Wiemhofer, H. D.; Bremes, H. G.; Nigge, U.; Zipprich, W. Solid State Ionics 2002, 150, 63. (24) Do, J. S.; Wu, K. J.; Tsai, M. L. Sens. Actuators, B 2002, 86, 98. (25) Oswin, H. Chem. Abstr. 1970, 72, 106582t. (26) Bay, H. W.; Blurton, K. F.; Lieb, H. C.; Oswin, H. G. Nature 1972, 240. (27) Sedlak, J. M.; Bay, H. W.; Valentine, A. M. Anal. Chem. 1974, 46. (28) Stetter, J. R.; Blurton, K. F. J. Chromatogr. 1978, 155. (29) Ho, K. C.; Liao, J. Y.; Yang, C. C. Sens. Actuators, B 2005, 108, 820. (30) Ho, K. C.; Hung, W. T. Sens. Actuators, B 2001, 79, 11. (31) Hauser, P.; Jordan, L.; Dawson, G. A. Electroanalysis 1997, 9. (32) Hauser, P.; Kappes, T. Analyst 1999, 124.

366 Chemical Reviews, 2008, Vol. 108, No. 2 (33) Hauser, P.; Dawson, G. Electroanalysis 2000, 12. (34) Janata, J. Principles of Chemical Sensors; Plenum Press: New York, 1989. (35) Lee, J. S.; Lee, J. H.; Hong, S. H. Sens. Acuators, B 2003, 96, 663. (36) Yang, J. C.; Dutta, P. K. Sens. Actuators, B 2007, 123 (2), 929. (37) Yao, S.; Stetter, J. R. J. Electrochem. Soc. 2004, 151 (4), H75. (38) Imaya, H.; Ishiji, T.; Takahashi, K. Sens. Actuators, B 2005, 108, 803. (39) Ho, K. D.; Hung, W. T. Sens. Actuators, B 2001, 79, 11. (40) Lee, S. B.; Maclay, J. G.; Keyvani, D. U.S. Patent 5,716,506, 1998. (41) Alberti, G.; Casicola, M. Solid State Ionics 1993, 61. (42) Campbell, D. N.; Davis, R. C.; Schmidt, J. C. U.S. Patent 4,744,954, 1986. (43) LaConti, A. B. Chem. Abstr. 1979, 90, 60474a. (44) Venkatasetty, H. V. U.S. Patent 4,522,690, 1985. (45) Otagawa, T.; Zaromb, S.; Stetter, J. R. J. Electrochem. Soc. 1985, 132. (46) Schmidt, J. C.; Campbell, D. N.; Clay, S. B. U.S. Patent 4,595,486, 1985. (47) Scheider, A. A. U.S. Patent 4,184,937, 1978. (48) Albery, W. J. U.S. Patent 4,377,446, 1982. (49) Xing, X. K.; Liu, C. C. Electroanalysis 1991, 3, 111. (50) Yan, H.; Liu, C. C. Sens. Actuators, B 1994, 17, 165. (51) Mayo, N.; Harth, R.; Mor, U.; Marouani, D.; Hayon, J.; Bettelheim, A. Anal. Chim. Acta 1995, 310, 139. (52) Opekar, F. Electroanalysis 1992, 1, 133. (53) Do, J. S.; Shieh, R. Y. Sens. Actuators, B 1997, 37, 19. (54) Sakthivel, M.; Weppner, W. Sens. Acutators, B 2006, 113, 998. (55) Samec, Z.; Opekar, F.; Crijns, J. E. F. Electroanalysis 1995, 7, 1054. (56) Chao, Y.; Buttner, W. J.; Yao, S.; Stetter, J. R. Sens. Actuators, B 2005, 106 (2), 784. (57) Schiavon, G.; Zotti, G.; Toniolo, R.; Bontempelli, G. Anal. Chem. 1995, 67, 318. (58) Schiavon, G.; Zotti, G.; Toniolo, R.; Bontempelli, G. Analyst 1991, 116, 797. (59) Hwang, B. J.; Liu, Y. C.; Hsu, W. C. J. New Mater. Electrochem. Syst. 1999, 2, 33. (60) Hwang, B. J.; Liu, Y. C.; Hsu, W. C. J. Solid-State Electrochem. 1998, 2, 378. (61) Schiavon, G.; Zotti, G.; Bontempelli, G.; Farnia, G.; Sandona, G. Anal. Chem. 1990, 62, 293. (62) Knake, R.; Hauser, P. C. Anal. Chim. Acta 2002, 459, 199. (63) Schiavon, G.; Comisso, N.; Toniolo, R.; Bontempelli, G. Electroanalysis 1996, 8, 544. (64) Kuwata, S.; Miura, N.; Yamazoe, N. Chem. Lett. 1988, 7. (65) Kuwata, S.; Miura, N.; Yamazoe, N. Nippon Kagaku Kaishi 1989, 10. (66) Liu, J.; Weppner, W. U.S. Patent 5,322,601, 1994. (67) Stetter, J. R.; Cao, Z. U.S. Patent 5,226,09, 1993. (68) Coillard, V.; Juste, L.; Lucat, C. Meas. Sci. Technol. 2000, 11. (69) Dutta, A.; Ishihara, T. Sens. Actuators, B 2005, 108, 309. (70) Dchmidt-Zhang, P.; Zhang, W.; Gerlach, F.; Ahlborn, K.; Guth, U. Sens. Actuators, B 2005, 108, 797. (71) Sleszynski, N.; Osteryoung, J. Anal. Chem. 1984, 56. (72) Maclay, G. J.; Buttner, W. J.; Stetter, J. R. IEEE Trans. Electron DeVices 1998, 35. (73) Buttner, W. J.; Maclay, G. J.; Stetter, J. R. Sens. Mater. 1990, 2. (74) Floate, S.; Hahn, C. E. W. J. Electroanal. Chem. 2005, 583, 203. (75) Wadhawan, J. D.; Welford, P. J.; McPeak, H. B.; Hahn, C. E. W.; Compton, R. G. Sens. Actuators, B 2003, 88, 40. (76) Floate, S.; Farmery, A. D.; Hahn, C. E. W. Sens. Actuators, B 2005, 109, 200. (77) Li, J.; Lu, Y.; Ye, Q.; Cinke, M.; Han, J.; Meyyappan, M. Nanoletters 2003, 3 (7), 929. (78) Lu, Y.; Li, J.; Han, J.; Ng, H.-T., Binder, C.; Partridge, C.; Meyyappan, M. Chem. Phys. Lett. 2004, 391, 344. (79) Young, P.; Lu, Y.; Terrill, R.; Li, J. J. Nanosci. Nanotechnol. 2005, 5 (9), 1509. (80) Li, J.; Lu, Y.; Meyyappan, M. IEEE Sens. J. 2006, 6 (5), 1047. (81) Li, J.; Lu, Y.; Ye, Q.; Delzeit, L.; Meyyappan, M. Electrochem. Solid State Lett. 2005, 8 (11), H100. (82) Qi, P.; Vermesh, O.; Grecu, M.; Javey, A.; Wang, Q.; Dai, H.; Peng, S.; Cho, K. J. Nano Lett. 2003, 3 (3), 347. (83) Stetter, J. R.; Bradley, K.; Cumings, J.; Gabriel, J.-C. P.; Gru¨ner, G.; Star, A. Nanotech 2003, 3, 313. (84) Wallgren, K.; Sotiropoulos, S. Sens. Actuators, B 1999, 60. (85) Berger, J.; Edelman, P. U.S. Patent 6,068,748, 2000. (86) Lee, Y.; Oh, B. K.; Meyerhoff, M. E. Anal. Chem. 2004, 76, 536.

Stetter and Li (87) Cha, W.; Lee, Y.; Oh, B. K.; Meyerhoff, M. E. Anal. Chem. 2005, 77, 3516. (88) Schmidt-Zhang, P.; Guth, U. Sens. Actuators, B 2004, 99, 258. (89) Stetter, J. R.; Penrose, W. R. Understanding Chemical Sensors and Chemical Sensor Arrays (Electronic Noses): Past, Present, and Future. In Sensors Update; Wiley-VCH: Weinheim, Germany, 2002; Vol. 10, Chapter 2.3, p 189. (90) Stetter, J. R., Penrose, W. R., Eds. Artificial Chemical Sensing: Proc. Eighth International Symposium on Olfaction and the Electronic Nose; ISOEN8-ISOEN 2001; The Electrochemical Society: Pennington, NJ, 2001; ISBN 1-56677-321-0. (91) Stetter, J. R.; Penrose, W. R. The Electrochemistry Encyclopedia; http://electrochem.cwru.edu/ed/encycl/art-n01-nose.htm (accessed 2001). (92) Zaromb, S.; Stetter, J. R. Sens. Actuators 1985, 6, 225. (93) Stetter, J. R.; Zaromb, S.; Penrose, W. R.; Findlay, M. W.; Otagawa, T.; Sincali, A. J. Hazard. Mater. Spills Conf. Proc., PreV., BehaV., Control Cleanup Spills Waste Sites (Ludwigson, John, Ed.) 1984, 183. (94) Stetter, J. R.; Jurs, P. C.; Rose, S. L. Anal. Chem. 1986, 58, 860. (95) Do, J. S.; Chen, P. J. Sens. Actuators, B 2007, 122, 165. (96) Pardo, M.; Dwong, L. G.; Sberveglieri, G.; Brubaker, K.; Schneider, J. F.; Penrose, W. R.; Stetter, J. R. Sens. Actuators, B 2005, 106, 136. (97) Li, J. Sensors 2000, 17 (8), 56. (98) Gardner, J.; Bartlett, P. N. Electronic Nose: Principles and Applications; Oxford University Press: New York, 1999. (99) Nanto, H.; Stetter, J. R. Introduction to Chemosensors. In Handbook of Machine Olfaction, Electronic Nose Technology; Pearce, T. C., Ed.; Wiley-VCH: Weinheim, Germany, 2003; Chapter 4. (100) Mazzone, P.; Mekhail, T.; Czich, C. Am. J. Respir. Crit. Care Med. 2006, A777. (101) Stetter, J. R. Ann. Am. Conf. GoV. Ind. Hyg. 1984, 11, 225. (102) Miller, M.; David, F.; Wilson, E.; Kling, R. J. Air Pollut. Control Assoc. 1971, 21. (103) Bryzek, J. S.; Roundy, S.; Bircumshaw, B.; Chung, C.; Castellano, K.; Vestel, M.; Stetter, J. R. IEEE Circuits DeVices, 2006, 22 (2), 8 (ISSN 8755-3996). (104) Reinhardt, G.; Mayer, R.; Rosch, M. Solid State Ionics 2002, 150, 79. (105) Nakamura, T.; Sakamoto, Y.; Saji, K.; Sakata, J. Sens. Actuators, B 2003, 93, 214. (106) Ivers-Tiffee, E.; Hardtl, K. H.; Menesklou, W.; Riegel, J. Electrochim. Acta 2001, 47, 807. (107) Ueda, T.; Plashnitsa, V. V.; Nakatou, M.; Miura, N. Electrochem. Commun. 2007, 9, 197. (108) Litzelman, S. J.; Rothschild, A.; Tuller, H. L. Sens. Actuators, B 2005, 108, 231. (109) Stetter, J. R. Amperometric Electrochemical Gas Sensors: Description and Applications. In NIST Workshop on Gas Sensors: Strategies for Future Technologies, Sept 8-9, 1993; NIST Pub. #865; National Institute of Standards and Technology: Gaithersburg, MD, 1993; pp 61-64. (110) Findlay, M. W.; Penrose, W. R.; Stetter, J. R. Quality Classification of Grain Using a Sensor Array and Pattern Recognition. Anal. Chim. Acta 1993, 284, 1. (111) Stetter, J. R. Instrumentation to Monitor Chemical Exposure in the Synfuel Industry. Ann. Am. Conf. GoV. Ind. Hyg. 1984, 11, 225. (112) Strathmann, S.; Penrose, W. R.; Stetter, J. R.; Gopel, W. Detection of TNT with chemical sensors. Proc. International Symposium on Olfaction and the Electronic Nose, (ISOEN 99), Tuebingen, Germany, Sept 20-22, 1999. (113) Wang, J. Electrochemical Sensing of Explosives. In Counterterrorist Detection Techniques of ExplosiVes; Yinon, J., Ed.; Elsevier: Amsterdam, The Netherlands, 2007; Chapter 4. (114) Wang, J. Electrochemical Sensing of Explosives.Electroanalysis 2007, 4, 415. (115) Wang, J.; Pumera, M.; Collins, G. E.; Mulchandani, A. Measurements of Chemical Warfare Agent Degradation Products Using an Electrophoresis Microchip with Contactless Conductivity Detector. Anal. Chem. 2002, 74, 6121. (116) Wang, J.; Pumera, M.; Chatrathi, M. P.; Escarpa, A.; Musameh, M.; Collins, G.; Mulchandani, A.; Lin, Y.; Olsen, K. Single-Channel Microchip for Fast Screening and Detailed Identification of Nitroaromatic Explosives or Organophosphate Nerve Agents. Anal. Chem. 2002, 74, 1187.

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367

Semiconductor Junction Gas Sensors Karin Potje-Kamloth* Institut fuer Mikrotechnik Mainz, Carl-Zeiss-Strasse 18-20, 55129 Mainz, Germany Received July 11, 2007

Contents 1. Introduction 2. Properties of Organic Semiconductors 2.1. Electronic Properties of Organic Semiconductors 2.1.1. Molecular Semiconductors 2.1.2. Conducting Polymers 2.2. Doping in Organic Semiconductors and Self-localized States 2.2.1. Molecular Semiconductors 2.2.2. Conducting Polymers 2.3. Transport Mechanism in Organic Semiconductors 2.3.1. Molecular Semiconductors 2.3.2. Conducting Polymers 3. Metal/Semiconductor Junctions and Their Use in Gas Sensing 3.1. Introduction 3.2. Interfacial Electronic StructuresIdeal Junction Characteristics 3.2.1. Formation of a Schottky Barrier 3.2.2. Charge Carrier Transport Mechanism across the Junction Potential Barrier 3.2.3. Ohmic Contact and Space Charge Limited Current 3.2.4. Tunneling through the Barrier 3.3. Interfacial Electronic StructuresDeviations from the Ideal Case 3.3.1. Introduction of the Ideality Factor 3.3.2. Schottky Barrier Lowering 3.3.3. Effect of Interface States 3.3.4. Metal/Organic Semiconductor Junctions with an Interfacial Layer 3.4. Gas/Solid Interactions in Schottky Barrier Junction Devices 3.4.1. Bulk Effect upon Gas DopingsModulation of the Electron Work Function 3.4.2. Formation of a Dipole and an Interfacial Layer 3.5. Extraction of Schottky Diode Parameters for Sensor Applications 3.5.1. Introduction 3.5.2. Current−Voltage Characteristic 3.5.3. Capacitance−Voltage Characteristic 3.5.4. Impedance Measurement 4. Schottky Diodes and Their Sensor Applications 4.1. Schottky Diodes Based on Inorganic Semiconductors * E-mail: [email protected].

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4.1.1. Silicon Based Schottky Diodes 4.1.2. GaAs Based Schottky Diodes 4.1.3. InP Based Schottky Diodes 4.1.4. GaN and AlGaN Based Schottky Diodes 4.1.5. SiC Based Schottky Diodes 4.1.6. CdSxSe1.x Based Schottky Diodes 4.2. Schottky Diodes Based on Organic Semiconductors 4.2.1. Phthalocyanines 4.2.2. Polythiophene/Poly(3-alkylthiophenes) 4.2.3. Polyanilines 4.2.4. Polypyrrole 5. Summary and Perspectives 6. List of Symbols and Abbreviations 7. References

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1. Introduction Over the past three decades, a number of studies have been undertaken to characterize the interactions of gases and chemical species with semiconductors as well as to develop useful sensing devices based on these effects.1 Examples include gas detectors using metal-oxide semiconductor fieldeffect transistors,2-4 metal-oxide semiconductor capacitors,5 as well as Schottky barrier diodes.6 Among those, Schottky barrier diode sensors are devices which are extremely simple to fabricate, obviating the need for photolithography or hightemperature diffusion/oxidation steps. These sensors may be broadly classified into two categories: (1) interfacecontrolled deVices, in which the species of interest are adsorbed at the metal surface and thus affect interfacial polarization, and (2) bulk-controlled deVices, which are dependent on the change in bulk properties of the semiconductor caused by absorption of and interaction with the diffusing species. Schottky barrier based gas sensors have been fabricated with a number of inorganic semiconductors using catalytic metals as the metal contact,6-11 mostly for detection of hydrogen. Due to various hydrogen-induced changes in catalytic metals, the change of the interface properties of the Schottky barrier diode readily lends itself to applications in interface-controlled sensors devices.6 The metal/inorganic semiconductor Schottky diodes can exhibit a very high sensitivity and low detection limit to hydrogen. The success of these devices translates into a huge number of papers published on the subject of hydrogen sensing. Unfortunately, the range of applicability is limited to the detection of hydrogen and hydrogen-producing gases or vapors, which comes from the exclusive solubility and permeability of atomic hydrogen in such materials. This limitation does not apply to the class of organic semiconductors, which includes organic molecules as well

10.1021/cr0681086 CCC: $71.00 © 2008 American Chemical Society Published on Web 01/30/2008

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Karin Potje-Kamloth studied chemistry at Ludwigs Maximilians Universita¨t Mu¨nchen and received her Ph.D. and Habilitation in physical chemistry. Currently, she is a senior scientist at Institut fuer Mikrotechnik Mainz GmbH. Her main research areas of interest are low cost chemical sensors, thin film technology of advanced polymer materials, energy harvesting, and chemical sensing platforms for microsystems

as conjugated polymers. They show almost unlimited permeability for gases and vapors, and the concomitant ability to form charge-transfer complexes with the matrix. Moreover, they offer the possibility to tune the sensitivity and selectivity toward gases and vapors by chemical alternation of the polymer backbone, addition of side groups, alternation of the lengths of the polymer chain, and alternation of the nature of the dopant. Variations in the chemical and physical properties of the organic semiconductors are manifold and represent a further advantage for organic semiconductors in this application area. For this reason, the focus of this review is on the progress in metal/organic semiconductor based Schottky barrier junction sensors. The use of organic semiconducting materials has received considerable interest as the active component in electronic device structures,12-18 such as organic light emitting diodes (OLEDs),19,20 solar cells,21,22 and organic field-effect transistors,23,24 Schottky diodes,25 and gas sensors.26-30 In the case of field-effect transistors (FETs), the organic semiconductor can either be the active component, i.e., the current carrying material, of the organic FET (OFET) or the selective layer used as the gate electrode in a chemical sensitive FET (CHEMFET). The interaction of gases or vapors with the bulk of organic semiconductors affects the work function of the organic semiconductor, which modifies the electrical properties of the electronic device. This leads to the application of organic semiconductors in bulk-controlled devices. The metal/semiconductor interface is the obvious component in any semiconductor device and controls the Schottky barrier device characteristic. The theoretical aspects of interfaces in devices based on organic semiconductors closely resemble those of inorganic ones. This allows in general the use of the extensive knowledge accumulated during many years of studies of the electronic properties of inorganic semiconductors. The purpose of this review is to introduce the reader to the field of metal/semiconductor junctions, in particular metal/organic semiconductor junctions, by providing insights into some of the theoretical and experimental approaches that have been employed thus far, although it is not meant to be an exhaustive study of all of the research work and techniques employed in the study of interface formation in semiconductor devices. This topic is too broad and is beyond the scope of this review.

Potje-Kamloth

The second section gives an introduction to the physics of organic semiconductors, which are relevant to the understanding of the peculiar properties of interfaces in organic semiconductor devices. The third section deals with some theoretical aspects concerning metal/semiconductor junctions and their use in chemical sensing. The fourth section is a review of this area of semiconductor junction gas sensors and deals with the electrical and gas sensing properties of Schottky barrier diodes. Both sections cover inorganic and organic semiconductor Schottky barrier diodes. Furthermore, section 4 shows how experimental and structural parameters of semiconductors can influence the Schottky barrier junction and hence the diode characteristics. The reader who is familiar with organic semiconductor polymers and the electrical properties of semiconductor contacts can skip directly to section 4, and those interested in an overview of this class of materials and of these types of devices may find sections 2 and 3 to be useful.

2. Properties of Organic Semiconductors The word “organic semiconductors” includes polymers and low-molecular-weight organic materials. The former are usually characterized by ill-defined chain lengths while the latter show both a well-defined composition and length. The use of the name “organic semiconductor” is based on the extrinsic semiconducting properties of organic systems, i.e., the capacity to transport charge generated by light, injected by electrodes, or provided by chemical dopants. Due to the weak electronic coupling between organic molecules, disorder plays a central role in the understanding of these materials. Disorder manifests itself in varying the energy levels of the electronic states of different molecules and the structure of the solid state. Both strongly influence the transport mechanism of charge carriers and charge injection at metal/semiconductor interfaces.

2.1. Electronic Properties of Organic Semiconductors 2.1.1. Molecular Semiconductors In inorganic semiconductor crystals such as silicon or germanium, the strong coupling between the atoms and the long-range order lead to the delocalization of the electronic states and the formation of allowed valence and conduction bands, separated by a forbidden energy gap. By thermal activation or photoexcitation, free electrons are generated in the conduction band, leaving behind positively charged holes in the valence band. The transport of these free charge carriers is described in the quantum mechanical language of Bloch functions, k-space, and dispersion relations familiar to solid-state physicists.31 In organic solids, intramolecular interactions are mainly covalent, but intermolecular interactions are due to much weaker van der Waals and London forces. As a result, the transport bands in organic crystals are much narrower than those of their inorganic counterparts and the band structure is easily disrupted by introducing disorder in the system. Thus, even in molecular crystals, the concept of allowed energy bands is of limited validity and excitations as well as interactions localized on individual molecules play a predominant role. The common electronic feature of molecular semiconductors is the π-conjugated system, which is formed by the overlap of carbon pz-orbitals. In Figure 1 the

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Figure 1. Molecular structure of metallophthalocyanine tetrasulfonate (MPcTS) salt.

molecular structure of a typical representative of this class of semiconductors, metallophthalocyanine tetrasulfonate (MPcTS), is shown. The orbital system of the ring comprises 42 π-electrons. Due to the orbital overlap, the π-electrons are delocalized within the molecule and the energy gap between the highest occupied molecular orbital (HOMO) and the lowest unoccupied molecular orbital (LUMO) is small, i.e., exhibiting transition frequencies within the visible range. Activation of free electrons by photoexcitation or chemical doping can lead to a tremendous increase of the electric conductivity from about 10-12 S cm-1 to values in the range of 102 S cm-1, which indicates the semiconducting character of the material.

2.1.2. Conducting Polymers Polymers are typically associated with flexible, processible materials having electrically insulating properties. Although this is true of most polymers, a special class of these materials called conjugated polymers has the electrical and optical properties traditionally associated with metals and semiconductors, yet they retain the mechanical properties and the processability of plastics. Conjugated polymers are characterized by repeated units in which atomic valence is not satisfied by covalent bonds. The leftover valence electrons of adjacent carbon atoms overlap and form double bonds that give rise to π-bonds, in which the π-electrons are delocalized over large segments of the polymer chain. They are forming the so-called π-conjugated system, which is responsible for the electronic properties of the conjugated polymers. The chemical structures of the most thoroughly studied conducting polymers are shown in Figure 2. Polymers can consist of on the order of 105 to 106 (or even a greater number) monomers strung together to form a macromolecular chain. Typical conjugated polymers include straight chain units, five or six-membered rings, and all combinations of these. The addition of heteroatoms (atoms other than carbon and hydrogen) and side chains allows for an even larger variety in this class. Their electronic properties can be tailored by the synthesis, and their electrical conductivity can be varied from about 10-12 S cm-1 to values in excess of 105 S cm-1. In order to understand the operation of polymer devices better, it is useful to first examine some basic properties of organic semiconductors based on conjugated polymers, because they differ qualitatively from those of crystalline inorganic semiconductors in several important respects. In crystalline materials, the properties are determined by the three-dimensional arrangement of the atoms and their resulting interactions. Conjugated polymers are treated as a quasi-one-dimensional system, wherein the polymer chains

Figure 2. Schematic representation of the chemical structures of the most commonly studied conjugated polymers, as well as their corresponding nomenclature and energy band gap Eg.

are assumed to behave independently of one another, and their physical and chemical properties depend on interactions within the single chains. The π-bonding scheme of conjugated polymers decreases the gap between occupied (HOMO) and unoccupied (LUMO) states. The valence band and conduction band of conducting polymers are generally derived from such π-bonds. The band gap of these polymers tends to lie between 1.5 and 3 eV (see Figure 2), in the same range as that of inorganic semiconductors. Since the bonds formed by σ-orbitals are stronger than the π-bonds, the polymers do not break apart when excited states are created in the π-electron levels. Also, in most polymers, there is only a weak overlap between the π-orbitals between neighboring polymer chains, so electrons and holes tend to be delocalized on individual polymer chains, although they can hop between chains. The π-electron delocalization along the chain and the weak interchain bonding give conjugated polymers a quasi-one-dimensional nature and give rise to strong anisotropies when the macromolecules are chain extended and chain aligned. Due to the electron-phonon-coupling of the π-electrons, which leads to a lattice distortion (Peierls distortion), a gap is generated at the Fermi level. If an electron is added or removed to/from the polymer chain, a self-localized electronic state within the previously forbidden semiconductor band gap is formed as a result of the easy deformability of the polymer lattice without chain scissoring. The creation of electronic defects allows charge transport within a single chain. This property of conjugated polymers demonstrates that many ideas are used to explain why inorganic semiconductors do not directly carry over to conjugated polymers. Early on, most of the research effort was concentrated on the conducting properties of the doped conjugated polymers,32 which were often referred to as ‘‘conducting polymers’’. In this review, this notation is used throughout the text to describe this class of polymers when they are in both pristine and doped states.

2.2. Doping in Organic Semiconductors and Self-localized States Doping of a disordered organic semiconductor by charged moieties has two counteracting effects: (1) the increase of concentrations of charge carriers and, thus, the change of the Fermi level;33,34 (2) the increase of energetic disorder by formation of additional deep Coulomb traps of the opposite polarity. Therefore, the average hopping rate is controlled

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by the charge carrier release from the Coulomb traps.35 The former effect facilitates conductivity; the latter strongly suppresses the charge carrier hopping rate. This process determines, together with the intrinsic order, the temperature dependence of the mobility. Arkhipov et al. showed that the doping efficiency on charge carrier hopping strongly depends on the energy disorder and the external field.35 Upon doping by ionized species, charge neutrality must be maintained. There are two ways to accomplish this. One is electrochemical doping. If the ionization (or reduction) potential of the electrode more or less matches the HOMO of the organic semiconductor, a negatively charged majority carrier can be injected, provided that the electrolyte supplies appropriate counterions that can diffuse into the semiconductor. For p-conducting organic semiconductors, positively charged majority carriers are injected, as the oxidation potential of the electrode matches the LUMO. An example is the oxidation of polyhexylthiophene by injection of holes from a solution containing tetraethylammonium perchlorate. Charge injection is compensated for by concomitant “doping” with perchlorate anions. The alternative method is doping by a neutral entity, whose electron affinity is different enough to allow for charge transfer from the semiconductor to the dopant. In both cases, excess mobile majority carriers and immobile countercharges (i.e., dopants) are generated that roughen the energy landscape, in which the charge carriers migrate, but in the “neutral” doping case, in addition, charge redistribution can and does occur. The majority of charge carriers will actually form metastable geminate pairs, whose dissociation is facilitated by the ambient phonon bath and the external electric field. A high charge carrier concentration can be reached without introducing counterions due to either field effect or high level of monopolar charge injection across a contact. Under these circumstances, the Coulomb interaction between charge carriers can strongly change the effective potential landscape. The interaction between charges of the same sign is repulsive and, therefore, cannot create Coulomb traps. It gives rise to transiently fluctuating potential barriers that affect the mobility.

2.2.1. Molecular Semiconductors Doping of organic molecular films has been investigated relatively little compared to the doping of inorganic semiconductors or of conjugated polymers. The main reason is that, unlike inorganic semiconductors, traditional n- and p-doping has not been a requirement for achieving bipolar transport in the most common molecular devices, i.e., OLEDs and OFETs. The ability to stack electron and hole transport layers, to build organic/organic or metal/organic junctions, alleviates the need to “dope” the organic materials in order to inject electrons or holes into the active layer(s) of the organic device. However, the performance of organic devices is now reaching a level at which electrical doping seems to be attractive as a means to further improve efficiency. In this respect, considerable work remains to be done, however, to understand and control doping in materials which exhibit fundamental differences with standard inorganic semiconductors. The weak intermolecular bonds, large energy gaps, and small dielectric constants of these materials are not particularly conducive to low dopant ionization energies. Several groups have started to investigate doping mechanisms and their effect on the electronic structure of the host organic molecular semiconductors.36-39 Particularly important are the relative energies of the dopant and the host molecular levels,

Potje-Kamloth

which determine the ‘‘ionization energy’’ and doping efficiency. Doping of metal/organic contacts with inorganic donors, such as lithium in aluminum tris(8-hydroxychinoline) (Alq3),40 and with inorganic acceptors, such as antimony pentachloride in N,N′-diphenyl-N,N′-bis(3-methylphenyl)1,1′-biphenyl-4,4′-diamine (TPD),41 has shown the potential for significant improvements in current injection. Similar results could be obtained by doping with molecular acceptors, e.g., tetrafluorotetracyanoquinodimethane (F4-TCNQ),37,42 and molecular donors, e.g., bis(ethylenedithio)tetrathiafulvalene (BEDT-TTF).43

2.2.2. Conducting Polymers When an electron is added to (or withdrawn from) a rigid band semiconductor, it must go into the conduction band (or come from the valence band) because the lower energy valence band is completely filled. However, when an electron is added to (or withdrawn from) a conducting polymer, a chain deformation takes place around the charge, which costs elastic energy and puts the charge in a lower electronic energy state. The competition between elastic deformation energy and electronic energy determines the size of the lattice deformation, which can be on the order of 20 polymer units long.44 This localized charged particle together with the simultaneous chain deformation is known as a polaron and is defined in the semiconductor literature as an electron that is dressed by a phonon cloud. The lattice distortion by the charges is also called self-trapping. In Figure 3A, the doping process of p-type polypyrrole (PPy) is depicted, which represents a typical conducting polymer. The corresponding band structures and the allowed electronic transitions are shown in Figure 3B. Bipolarons are similar to polarons, but they are double charged. Instead of being a single charge that distorts the chain, there are two charges, which are bound together in and by the same chain deformation. Although these charges repel each other via the Coulomb interaction, they remain bound together by their common chain deformation; the energy increase to form two distinct chain deformations rather than a single one is greater than the Coulomb energy gained by their separation. Associated with each one of these excitations are energy states located in the energy gap of the p-doped polymer (Figure 3B). The polaron has two subgap states, of which the lower one is singly occupied and the higher one is empty for a positive polaron (i.e., an electron and a hole). The doubly charged bipolaron has its subgap states completely empty for a positive bipolaron. A significant difference between polarons and bipolarons is their spin signatures: polarons have charge (1 and spin 1/2, the same spin and charge as a free electron. Bipolarons have a zero spin and are doubly charged, mainly because of the pairing of the electrons. The charge injection into the macromolecular chain, i.e., the creation of any of the above-mentioned defects (polaron, bipolaron) in the conjugated backbone of the polymer, is called doping. Besides this charge injection, doping in polymers also implies the insertion or repulsion of a counterion (referred to as “dopants” or “doping ions”) to maintain charge neutrality. In order to match conventions between physicists and chemists, a polymeric cation-rich material is called p-doped, and a polymeric anion-rich material is called n-doped. Since every monomer is a potential redox site, conducting polymers can be doped to a high density of charge carriers. The doping level is up to 5

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Figure 3. (A) Neutral (undoped) state of polypyrrole and doped (polaron and bipolaron) structures of lightly and heavily doped polypyrrole, respectively. (B) Corresponding schematic energy band structures and the allowed electronic transitions.

orders of magnitude greater than that in common inorganic semiconductors. The standard value of the doping levels lies in the range of 0.27-0.5 positive charge per monomer (2750% by mol), depending on the anion and the type of conducting polymer. A slight decrease in the doping level is observed with an increase of the molecular weight and the negative charge of the anion. For all conjugated polymers, except for polyacetylene, polarons and bipolarons are the entities through which charge transport is accomplished in conducting polymers.45 They can travel along the chain as an entity, with the atoms in the path changing their positions so that the deformation travels with the electron or hole.46 Doping in polymers can be accomplished chemically, electrochemically, and photochemically, as well as by charge injection at the metal/insulator/semiconductor (MIS) interface.47 In the case of electrochemical and/or chemical doping, the induced electrical conductivity is permanent, until the charge carriers are purposely removed by dedoping, i.e., by reversing the electrochemical reactions, or until the charge carriers are chemically compensated. Initially, conducting polymers have been regarded as onedimensional semiconductors, because the π-conjugated systems extend over the whole polymer chains, thus allowing for delocalized states in one direction. In their pioneering work, Su, Schrieffer, and Heeger48 studied the interaction of an excitation with an ideal one-dimensional lattice, neglecting Coulomb interactions and disorder effects. Today, conducting polymers are discussed either as one-dimensional semiconductors with strong electron/phonon interactions or as disordered molecular solids with strong Coulomb interactions. The discussion of the precise nature and dynamics of excited states in conducting polymers is still very much alive. The one-dimensional semiconductor nature of conducting polymers might play a role in experiments on highly ordered, stretch-oriented materials or on isolated single chains. Furthermore, in conducting polymers, the influence of chemical dopants on the morphology of the polymer layer and on the charge-transporting states is still a matter of

controversy. The (quasi) one-dimensional nature of the conducting states may explain the exceptionally high conductivities in doped conducting polymer systems.49

2.3. Transport Mechanism in Organic Semiconductors A feature common to all amorphous and disordered systems, among them organic semiconductors, is the frequency dependent conductivity that increases approximately linearly with frequency, at least in the range 101-107 Hz, i.e., σ(ω) ∝ σωs, where the frequency exponent s e 1.50 The origin of this obviously universal behavior of the dispersive component σac of the conductivity has been ascribed either to the relaxation process or to a transfer process. The former is caused by the motion of charge carriers by thermally assisted quantum mechanical tunneling between localized states lying deep within the band gap. The latter is called hopping and involves classical thermal activation over the barrier height, separating two sites with an energy difference ∆E. The frequency exponent s is predicted to have a temperature dependence, and the magnitude of s at any temperature is determined by the binding energy of the charge carrier in its localized site.51 The temperature dependence of the dc conductivity term provides first-hand information on the possible processes involved in the conduction in amorphous materials. One of the theoretical models most successfully used to predict the probable transport mechanism in disordered materials is the Mott’s variable-range hopping (VRH).52 It describes a phonon-assisted quantum mechanical transport process, in which a balance is obtained between the thermodynamics constraint on a charge carrier moving to a nearby localized state of different energy and the quantum mechanical constraint53 on a charge carrier moving to a localized state of similar energy, but spatially far away. The VRH model is equally applicable to charge carriers such as electrons, polarons, or bipolarons, provided that the appropriate wave

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function is incorporated. This model and its field of applications have been extensively reviewed by various authors.52,54 According to the phonon-assisted quantum mechanical transport process, the temperature dependence of the dc conductivity provides information about the conduction mechanism and the dimensionality of the charge transport mechanism, which ranges from a three-dimensional to a onedimensional transport process. Considering a one-dimensional system, the most commonly invoked descriptions are charging-energy-limited tunneling (CELT),55 fluctuation-induced tunneling (FIT),53 and quasi-one-dimensional variable-range hopping.52,56

2.3.1. Molecular Semiconductors Charge carrier hopping within a positionally random and energetically disordered system of localized states was shown to be an adequate model for the description of both the equilibrium and the nonequilibrium conductivity in noncrystalline organic semiconductors.57-60 Ba¨ssler et al. discussed the effect of doping on the charge carrier hopping in disordered organic semiconductors.58,61 Doping such a system by charged moieties will create a random distribution of dopant ions that will coulombically interact with charge carriers localized in randomly located intrinsic hopping sites62 and, thus, broaden the effective density of states distribution. This effect is especially important for molecular semiconductors, because the dielectric permittivity is low, and the range of the Coulomb potential is large in organic solids.

2.3.2. Conducting Polymers Much work has focused on the nature of charges in doped conducting polymers. Even though there is a high density of conduction electrons at the Fermi level for the highly doped polymers, the charge carriers may be spatially delocalized, so that they cannot participate in the electronic transport except through hopping. The prime source of localization in conducting polymers is structural disorder. Most conducting polymers are noncrystalline, i.e., amorphous materials, although some of these systems possess a molecular low crystallinity, in which regions of the material are more ordered and other regions are more disordered. The percentage of crystallinity may vary from near zero to 11% for polypyrroles,63 20-30% for polyanilines,64 and approximately 80% for highly doped polyacetylenes.65 The charge carrier mobilities in these low-conducting amorphous solids are frequently estimated to be very low, µ , 1 cm2 V-1 s-1.66 These values bear witness against a band transport mechanism. However, detailed knowledge of their charge transport is usually quite limited. Nevertheless, the basic principles underlying electronic states and electronic transport in this class of solid materials are those of amorphous solids. It should be emphasized that the great variety of polymer constituents and structures yields a diversity of interesting and perhaps specific electronic properties. The dimensionality of the system can play a major role in determining the nature of both electronic states and electronic transport. The intrinsic contributions to charge transport are strongly intermixed, and it is often not trivial to quantify the individual contributions from various parameters (e.g., interchain interactions, anisotropic diffusion of charge carriers, the role of dopant ions, the extent of disorder, etc.). The theoretical modeling of transport properties in conducting polymers is a challenging problem because of the extreme complexity of these systems.

Potje-Kamloth

For a transport mechanism to be described in a certain material, charge carriers must be present. In conducting polymers, upon injection of an electron or a hole, a defect in the polymer chain is created, i.e., a polaron or a bipolaron, as already mentioned. These charged defects are the real charge carriers in nondegenerated conducting polymers. Charge transport in conducting polymers consists generally of two components: intrachain and interchain transport.67,68 Intrachain charge transport occurs along the polymer backbone and requires less energy than interchain charge transport, which involves the hopping of the charge to neighboring chains. The dopant compounds play a very important role in the hopping process during interchain charge transport. The charge carrier mobility may be thought of as a measure of the ease with which the charge carriers move through the material. It is sensitive to the level of structural order present in the polymer. Thus, structural defects decrease the conductivity by lowering the mobility. It is interesting that the theoretical conductivity of perfectly aligned defectfree polyacetylene is believed to be greater than 106 S cm-1.67 As a general consideration, the conjugation length appears to be the most significant structural parameter controlling the charge mobility. It has been recognized that the effective conjugation length of most conducting polymers is finite, because of the imperfect planarity of the conjugated system. It is supposed that the chain architecture of a conducting polymer consists of rigid blocks of a conjugation length of several polymer units separated by flexible nonconjugated polymer segments. This kind of macromolecular chain architecture should be kept in mind when one deals with the structure-property relationship of conducting polymers. Currently, the description of the conductivity in these kinds of disordered materials remains open, and the understanding of the influence of the disorder, which is present at every scale of the structure (defects on the chains, heterogeneities in the doping distribution, etc.) is far from being complete. Moreover, several studies have been devoted to the theoretical search for microscopic pictures leading to different new descriptions, such as Fermi glass,69 Coulomb glass,70 superlocalization,71 multifractal localization,72 etc. Such a great variety of theoretical approaches reflects the wide range of phenomena involved in electronically conducting polymers. Furthermore, the transport process may differ from one polymer to another, as well as from one sample to another, principally according to the preparation method.

3. Metal/Semiconductor Junctions and Their Use in Gas Sensing 3.1. Introduction In the following sections, the behavior of basic metal contacts with semiconducting materials and their application in chemical gas sensing are described. The elucidation of the interfacial electronic structure forms the basis for understanding and improving the performance of electronic devices. A key theoretical prediction is that the electric field at the semiconductor/metal interface should, in principle, respond to changes in the Fermi level of the contacting metal.73,74 The behavior of such contacts has been discussed thoroughly in several books and reviews articles,75-77 although the interfacial chemistry of these systems is still controversial.

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Figure 4. Formation of a barrier between a metal and a semiconductor: (A) neutral and isolated; (B) electrically connected; (C) in perfect n-type rectifying contact (φm >φs); (D) n-type ohmic contact (φm < φs); (E) p-type ohmic contact (φm > φs); (F) p-type rectifying contact (φm < φs). o denotes electron in conduction band; + denotes donor ion.

Metal/semiconductor contacts are components of any semiconductor device. At the same time, such contacts cannot be assumed to have a resistance as low as that of two connected metals. In particular, a large mismatch between the Fermi energy of the metal and the semiconductor can result in a high-resistance rectifying contact. A proper choice of materials can provide a low-resistance ohmic contact. However, for a lot of semiconductors, there is no appropriate metal available. Instead, one then creates a tunnel contact. Such a contact consists of a thin barrier, obtained by heavily doping the semiconductor, through which charge carriers can readily tunnel. Thin interfacial layers also affect the formation, which is discussed for metal/organic semiconductors contacts.

3.2. Interfacial Electronic StructuresIdeal Junction Characteristics 3.2.1. Formation of a Schottky Barrier Experiments on rectifying contacts started in 1874 with the pioneering work of Braun, who observed asymmetries in the transport of electrical current across metal/semiconductor interfaces.78 The following decades brought a variety of technical applications, but it took more than 60 years until Schottky,79 and independently Mott,80 used microscopic concepts to describe these so-called Schottky contacts. Since then, metal/inorganic semiconductor contacts have been extensively studied and have become rather well understood. The subject of interfacial electronic structure can be roughly divided into two aspects: (1) the energy level alignment at the interface and (2) the band bending in the space-charge region.81 The former is important for charge carrier injection; the latter is essential for charge carrier separation. A complete

overview of the works on Schottky contacts is beyond the scope of this review. Several excellent textbooks report on their electronic properties and the development of the field.81-83 Figure 4 illustrates the band diagram of Schottky’s Gedankenexperiment for the formation of a Schottky barrier. The metal and the semiconductor are supposed to be electrically neutral and to be separated from each other for an n-type semiconductor with a work function φs smaller than that of the metal φm (Figure 4A). When the metal and the semiconductor come in electrical contact, the two Fermi levels are forced to coincide and electrons pass from the semiconductor into the metal. The result is an excess of negative charge on the metal surface and the formation of a positive charge depletion zone with a thickness WS in the semiconductor near its surface (Figure 4B). These excess charges form an interface dipole and produce an electric field, directed from the semiconductor to the metal. By bringing the metal and the semiconductor closer together, the gap between the two materials vanishes and the electric field corresponds now to a gradient of the electron potential in the depletion layer, resulting in the well-known band-bending regime (Figure 4C). The Mott-Schottky model leads to the Schottky barrier height φb. For an ideal contact between a metal and an n-type semiconductor, the height of the barrier φb measured relative to the Fermi level is given by

φb ) φm - χS

(1)

φb ) eVd + ξ ) φbi + ξ

(2)

or

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where χs is the electron affinity of the semiconductor, defined as the difference in energy between an electron at rest outside the surface and an electron at the bottom of the conduction band just inside the surface, e is the electronic charge, ξ is the difference between the bottom of the conduction band EC and the Fermi level of the semiconductor, eVd is the band bending at zero bias voltage, and eVd, the diffusion potential at zero bias voltage, usually called the built-in potential φbi or the contact potential formed between the metal and the semiconductor, is given by82

eNDWS2 φbi ) eVd ) 2S

(3) J ) A**T2 exp

where ND is the donor concentration, s is the permittivity of the semiconductor, and WS is the depletion layer width. Equation 2 has been referred to as the Mott-Schottky limit. In obtaining it, a number of important assumptions have to be made, namely that (1) the surface dipole contributions to φm and χS do not change when the metal and semiconductor are brought into contact, (2) there are no localized states on the surface of the semiconductor, and (3) there is a perfect contact between the semiconductor and the metal (i.e., there is no interfacial layer). In the case that the work function of the n-type semiconductor φs is larger than that of the metal φm, the contact is biased so that electrons flow from the semiconductor to the metal. They encounter no barrier (Figure 4D). If a bias voltage is applied such that electrons flow in the reverse direction, the comparatively high concentration of electrons in the region where the semiconductor bands are bent downward (usually referred to as the accumulation region) behaves like a cathode, which is easily capable of providing a copious supply of electrons. The current is then determined by the bulk resistance of the semiconductor and the applied voltage. Such a contact is termed ohmic contact. A p-type semiconductor/metal junction, for which φm exceeds φS, represents an ohmic contact (Figure 4E). The case of a p-type semiconductor, for which φS exceeds φm, is shown in Figure 4F. Bearing in mind that holes have difficulty in going underneath a barrier, one sees that Figure 4F is the p-type analogue of Figure 4C and gives rise to rectification, which is the case of Schottky contacts based on p-type organic semiconductors. The barrier height φb for an ideal contact between a metal and a p-type semiconductor is given by

φb ) Eg - (φm - χs)

metal/semiconductor interface, contribute to the current flow. Quantum mechanical tunneling through the barrier takes into account the wave-nature of the electrons, allowing them to penetrate through thin barriers. In a given junction, combinations of all three mechanisms exist. However, typically one finds that only one charge carrier mechanism dominates. Thermionic Emission Theory. Thermionic emission theory assumes that the barrier height φb is much larger than kT and that electrons, with an energy larger than the top of the barrier, will cross the barrier, provided they move toward the barrier. The actual shape of the barrier is hereby ignored. The current density, J, can be expressed as

(4)

3.2.2. Charge Carrier Transport Mechanism across the Junction Potential Barrier The models initially developed for charge carrier transport across potential barriers in crystalline materials have also been used for noncrystalline systems. The current across a metal/semiconductor junction is mainly due to majority carriers. Three distinctly different mechanisms exist: diffusion of charge carriers from the semiconductor into the metal, thermionic emission of charge carriers across the Schottky barrier, and quantum mechanical tunneling through the barrier. The diffusion theory assumes that the driving force is distributed over the length of the depletion layer. The thermionic emission theory, on the other hand, postulates that only energetic charge carriers, which have an energy equal to or larger than the conduction band energy at the

( )[ ( ) ] φb eV exp -1 kT kT

(5)

or

[ (eVkT) - 1]

J ) J0 exp

(6)

where A** is the effective Richardson constant for thermionic emission, neglecting the effects of optical phonon scattering and quantum mechanical reflection (equal to 120 A cm-2 K-2 for free electrons), T is the absolute temperature, k is the Boltzmann constant, V is the bias voltage, and J0 is the saturation current density, given as

( )

J0 ) A**T2 exp -

φb kT

(7)

Diffusion Theory. For semiconductors with low charge carrier mobility (µ < 10-4 cm2 V-1 s-1), it has been shown that the dominant barrier transport mechanism is diffusion.81 The diffusion theory by Schottky81 is derived from the assumption that the depletion layer is large compared to the mean free path, so that the concepts of drift and diffusion are valid. The resulting current density equals

( )[ ( ) ]

J ) eNcµeE¨ max exp -

φb eV exp -1 kT kT

(8)

eV -1 kT

(9)

or

[

J ) JSD exp

]

where NC is the effective density of states in the conduction band, µe is the electron mobility, and E¨ max is the maximum field strength at the metal/semiconductor interface given by E¨ max ) eNDW/s. It can be seen that eq 9 is almost, but not quite, of the form of the ideal rectifier characteristic for thermionic emission. The difference arises because E¨ max is not independent of bias voltage but is proportional to (V0 V)1/2. For large values of reverse bias, the current does not saturate but increases roughly with |V|1/2.

3.2.3. Ohmic Contact and Space Charge Limited Current A metal/semiconductor junction results in an ohmic contact (i.e., a contact with voltage independent resistance), if the Schottky barrier height φb is zero or negative. In such cases, the charge carriers are free to flow in and out of the semiconductor so that there is a minimal resistance across the contact (see Figure 4D and E). For an n-type semiconductor, the work function of the metal must be close to or smaller

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than the electron affinity of the semiconductor. For a p-type semiconductor, it requires that the work function of the metal must be close to or larger than the sum of the electron affinity and the band gap energy. Since the work function of most metals is less than 5 V and a typical electron affinity is about 4 V, it can be problematic to find a metal that provides an ohmic contact to p-type semiconductors with a large band gap such as GaN, SiC, or organic semiconductors. The values of the work function of organic semiconductors strongly depends on the preparation method and the doping level. The values range between 4.0 and 5.3 and can even differ for the same type of polymer and doping because of differences in the treatment of the material. Therefore, the reader is referred to the literature to obtain information on the work function values of the organic semiconductors of interest. Effect of Background Doping-Bulk Limited Transport. In J-V characteristics, a slope of the ln J versus ln V plot of approximately unity indicates an ohmic conduction. Assuming conduction is via holes, the current flow may be expressed in the form84

(VL)

J ) eh0µh

(10)

where h0 is the concentration of thermally generated holes in the valence band, µh is the hole mobility, and L is the thickness of the semiconductor. The concentration of holes at thermal equilibrium is given by

[

h0 ) NV exp -

]

(EF - EV) kT

(11)

where NV is the effective density of states in the valence band and (EF - EV) is the separation of the Fermi level from the valence band edge. The current density in the ohmic region becomes

() [

J ) eµhNV

]

(EF - EV) V exp L kT

(12)

By plotting ln(J/V) against 1000/T, the values eµhNV and (EF - EV) can be calculated from the slope and the intercept at 1/T ) 0. When holes are injected into the p-type bulk of a metal/ p-type semiconductor Schottky diode, the total hole density NV(x) at the point x measured from the ohmic (metal) contact consists of two parts: (1) the density of the existing holes h0 induced by the background acceptors and (2) the density [NV(x) - h0] of the injected holes. The total density of the holes determines the current. The space charge is determined by the distribution of injected holes, provided that the charge of the existing holes h0 is compensated by the ionized fixed dopant ions. As long as the total hole density NV(x) is equal to the doping density over most of the sample thickness, there is no space charge present, and it is expected that Ohm’s law will be obeyed. The electric field in the doped samples becomes constant over practically the whole length of the sample. This assumption would be equivalent to performing an electrochemical experiment in which the resistance of the solution dominates and no electrochemical activity at the surface electrode (i.e., contact) can be detected. In effect, such a case corresponds to a chemiresistor in which the useful information is obtained from the changes of bulk resistivity of the sensing layer. This is not the case for Schottky barrier junctions.

If current is passing through a forward biased junction with a low doped semiconductor or a semiconductor exhibiting low mobility, a depletion layer will be formed in front of the ohmic contact of the Schottky barrier diode, as discussed by Chen et al.85 It can be described in analogy to the depletion polarization effect in an electrochemical experiment by a limited charge carrier transport process inside the semiconductor toward the ohmic contact. The width of the mass-transport limited contact resistance can be modulated by the external field. The time dependence of the polarization resistance is controlled by the charge carrier mobility and by the geometry of the contact. In most Schottky barrier devices, the ohmic contact resembles an electrode with a rectangular geometry, for which the polarization resistance increases with 1/xt at constant applied voltage. The current increases roughly with |V|1/2 (see also section 3.2.2). Moreover, a second bulk effect has to be taken into account in the presence of an intrinsic or low doped semiconductor forming the Schottky barrier junction. If the semiconductor has no free or low charge carrier density due to doping or due to intrinsic thermal ionization, the theory of space charge limited current (SCLC) applies and the current density J is given by the V2 (Mott-Guerney) Law81,86

9 V2 J ) s µ 3 8 L

(13)

where s is the permittivity of the semiconductor and L is the thickness of the sample. With increasing applied voltage, the charge carrier density of the undoped sample increases continuously with x. The injected charge carrier density becomes considerably larger than the doping induced density in most of the volume of the sample and the electric field increases with xx. The behavior of the current density at high currents, and therefore at high voltages, is quite different. The curves for high and low doped semiconductors become identical. One would therefore expect that the current would obey the SCLC V2 law at high voltages in the doped material as well. The transition from Ohm’s law to SCLC takes place at the point at which the Ohm’s law straight line and the SCLC line (corresponding to the V2 law) meet. Therefore, an expression for the voltage Vtr, at which the transition occurs, can be derived by equating the ohmic current and the SCLC (eqs 10 and 13):

8eh0L2 Vtr ) 9s

(14)

This equation provides a method for determining the background doping concentration. The transition voltage Vtr can be experimentally determined. Effect of Trapping and Field Dependent Mobility. A modification of the V2 law occurs when trapping and the dependence of the mobility on the electric field are taken into account. The effects of trapping and field dependent mobility on the J-V curves have been discussed by several authors.86-89 Most workers have assumed an exponential distribution of traps Ntr given by

( )

Ntr(E) ) NV exp -

E kTtr

(15)

where NV is the density of states for a p-type semiconductor

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semiconductor that there is only a thin barrier separating the metal from the semiconductor. For heavily doped semiconductors (doping level: 1019 cm-3 or higher) exhibiting a depletion region width at the metal/semiconductor interface on the order of 3 nm or less, or for operation at low temperature, it may be possible for electrons with energies below the top of the barrier to readily tunnel across such barriers by quantum mechanical tunneling.82 If the tunneling current dominates the current flow, the current-voltage relationship is of the form

( )[

Figure 5. Dark J-V characteristic of an Al/doped PPy Schottky diode. The dashed line indicates the fit of ln J vs V according to thermionic emission (see section 3.2.2).

and Ttr is the characteristic temperature of the distribution. The ratio of free to trapped charge carrier density θ for a p-type semiconductor is given by

θ)

[

]

NV -(Et - EV) exp Ntr kT

(16)

where (Et - EV) is the activation energy of hole traps. For the distribution of traps, the current density J is given by the J ) ktrVm power law, where ktr is a constant and the value of the exponent m depends on the characteristic temperature Ttr. The effect of field dependent mobility has also been discussed by several authors.89,90 Most authors have used the following equation for the mobility:

µ ) µ0 exp(βxE)

(17)

Variations of this equation are discussed by Campbell et al.89 It is shown that both SCLC with traps and field dependent mobility can give rise to the equation J ) ktrVm. However, experimental currents measured at different temperatures could not be described with one value of m. At a given temperature and in a small voltage range, the Vm law with constant m has been experimentally verified.89 Figure 5 shows a typical current density vs applied voltage curve of an Al/PPy Schottky diode. The diode is forward biased at positive voltages, implying that the Al contact is negative. The current density in the low bias region between 0.1 and 1.5 V varies exponentially with the bias voltage and fits closely with the thermionic emission theory (eq 6). At a bias voltage > 1.5 V, the current starts to become linear, indicating a dominant contribution of the bulk resistance, which is in series with the resistance of the depletion layer of the diode. As the voltage increases to more than 5 V, the current increases again. This indicates the transition from the contribution of the ohmic current to the contribution of the SCLC, according to eq 14. The current density J varies with Vm. The determination of the exponent m by plotting J as a function of Vm gives information if the SCLC regime is trap free and the mobility is field independent (m ) 2) or controlled by trapping or by the effect of the electric field on the mobility (m > 2).

3.2.4. Tunneling through the Barrier An alternate and more practical contact is a tunnel contact, if no appropriate metal is available to form an ohmic contact. It can be particularly problematic to find a metal that provides an ohmic contact to doped organic semiconductors that usually show work functions of about 5 V or higher. Tunnel contacts do have a positive barrier at the metal/semiconductor interface. They also have a high enough doping level in the

J ) J0 exp -

φb eV exp -1 E00 kT

( )

]

(18)

where E00 ) (h/4π)(ND/m*s)1/2/eV, where h is Planck’s constant and m* is the effective mass of the electron. The equation indicates that the tunneling will increase exponentially with xND. A possible mechanism for the charge transfer from the metal to a conductive polymer based on tunneling through the barrier is given by Gustafsson et al.91 Mobile charges in the conducting polymer exist in the form of polarons and bipolarons which occupy states in the band gap (see section 2.2.2). They are different from the valence band holes normally encountered in inorganic semiconductors. Injection of a hole by tunneling or by thermionic emission into the polymer should be similar to the photogeneration of charge in conducting polymers, where, after its formation, a rapid relaxation to polaronic/bipolaron states occurs.92 Thus, the majority carrier transfer from the metal to the polymer is based on (1) tunneling through the barrier to the valence band, followed by (2) an immediate relaxation to a polaron or bipolaron defect state, whereas, in the opposite direction, the charge transfer occurs by direct tunneling from the defect state to the metal. This gives rise to different tunneling paths at different energy levels in the forward and reverse bias voltage directions, which probably can affect the rectification ratio of the device.

3.3. Interfacial Electronic StructuresDeviations from the Ideal Case 3.3.1. Introduction of the Ideality Factor Real Schottky diodes do not always follow the expressions derived for the ideal case of a charge carrier transport mechanism across a potential barrier. Deviations from ideal behavior arise from imperfections in fabrication, or factors, which are not included in the relatively simple theories taken for derivation of eqs 5-9 and eq 18. A few of the major limitations are given below. In practice, diodes never satisfy eq 5 exactly but can be more closely described by the modified eq 19.82

J ) J0 exp

eV (nkT )[1 - exp eVkT]

(19)

or

J ) J0 exp

eV (nkT )

for V .

3kT e

(20)

where n is the ideality factor, which can be estimated from the slope of the straight part of the ln J vs V plot and is usually greater than unity. The ideality factor n is introduced in eq 19 to account for the failure of the simple, single-

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Figure 6. Schematic band diagram of the band bending interface in the presence of surface states for a metal/semiconductor: (A) before contact; (B) after contact formation. The Fermi level is pinned by a high density of surface states of the semiconductor.

carrier theories of barrier behavior, e.g., when image lowering of the potential barrier occurs. There are many other possible reasons why n should exceed unity, e.g., series resistance of the bulk semiconductor, generation-recombination within the depletion region, tunneling through the barrier, interface states, and the presence of an insulating interfacial layer.82,93 In the case of polymeric diodes, bulk resistance and interfacial layer effects are likely to be the most important.

3.3.2. Schottky Barrier Lowering It has been assumed that the Schottky barrier height remains constant under all conditions of applied voltage. However, the barrier height varies with applied voltage because conduction electrons experience a force from their image charges in the metal. This force attracts the electrons toward the metal surface, effectively lowering the barrier and allowing voltage-dependent deviations from ideal behavior. In theory, this “image force” should give the reverse current a fourth-power dependence of bias upon voltage, rather than the constant value implied by eq 7. This effect is usually not observed, because charge carrier generation in the depletion region at high reverse bias and tunneling effects dominate the reverse leakage. At forward biases above approximately 0.1 V, the effect causes a slight deviation of the ideality factor n from unity. For a diode that is assumed to be ideal except for barrier lowering, the ideality factor is

(

n) 1-

)

dφb dV

-1

(21)

where dφb/dV is the derivative of the barrier height with respect to the applied voltage. The image force lowering of the barrier ∆φbi is given as

∆φbi )

( ) eE¨ max 4π0

1/2

(22)

E¨ max is the maximum field at the interface, and 0 ) 8.85 × 10-14 C cm-1 V-1 is the permittivity of free space.

3.3.3. Effect of Interface States Very often, a deviation from the ideal Mott-Schottky behavior is observed. It can be measured in terms of the slope parameter

S)

dφb dφm

(23)

describing the dependence of the Schottky barrier height on the metal work function. According to the Mott-Schottky

model, S should be equal to unity, while S is smaller than unity for any deviation. An important limitation of the Mott-Schottky model is the neglect of intrinsic surface states with energy levels located in the semiconductor band gap which are pinning the metal Fermi level as shown in Figure 6. In this extreme case, the Schottky barrier height does not depend at all on the metal work function, i.e., S ) 0. Considering a continuum of interface states, a phenomenological formula for the Schottky barrier height can be formulated76

φb ) S(φm - χs) + (1 - S)φ0

(24)

where S is the slope factor and φ0 is the barrier height when the charge neutrality level of the continuum of interface states coincides with the Fermi level at the interface. The barrier height becomes independent of the metal work function and is determined entirely by the doping and surface properties of the semiconductor. In real situations, the constant S lies between 0 and 1. “Fermi level pinning” indicates that the Fermi level position at the surface of the semiconductor measured relative to the vacuum level does not vary when either the work function or the Fermi level or the contacting phase is varied. Under strong Fermi level pinning, a constant electric potential is dropped across the semiconductor regardless of the nature of the metal. The remaining potential is dropped across a thin dielectric layer near the metal/ semiconductor junction. Besides the interface states, other reasons have been proposed for inorganic semiconductor/metal junctions as explanations for the Fermi level pinning behavior, i.e., chemical reactions such as alloy formation,94 or other stoichiometry changes,95 and/or quantum mechanically induced interface states.96 Although these theories were primarily developed to explain the behavior of inorganic semiconductor/metal contacts, many of them also make testable predictions regarding the behavior of contacts comprising organic semiconductors in Schottky barrier junction sensors.

3.3.4. Metal/Organic Semiconductor Junctions with an Interfacial Layer The elucidation of the interfacial electronic structure, particularly the energy-level alignment at organic/substrate interfaces, forms the basis for understanding and improving the performance of organic electronic devices.97,98 Despite the advances in device application, there is still only a limited understanding of the interface between organic materials and metals. A number of recent studies have helped to understand the basic properties of the interface; see, e.g., the review by

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Figure 7. Energy diagram of the organic/metal interface with (A) band bending leading to Fermi level alignment and (B) without band bending. Evacuum denotes the vacuum level, EF is the Fermi level, eVi is the shift of the vacuum level at the interface due to the dipole layer formation, and ∆ is the interfacial layer thickness.

Seki at al.99 and references therein. In particular, ultraviolet and X-ray photoemission spectroscopy (UPS, XPS) have been used to investigate metal/organic100,101 and organic/ organic102-104 interfaces. The organic layers employed in such studies were nominally undoped, although the purity of the organic layers is usually insufficient to achieve intrinsic behavior. The simplest model used to discuss the energy level offset at these interfaces is the “vacuum-level alignment” model (i.e., Mott-Schottky limit, eq 2). The band of an organic semiconductor in contact with a metal is bent to achieve the alignment of the Fermi level of the metal with the Fermi level of the organic layer fixed somewhere in the band gap, simply by analogy to an interface between a doped inorganic semiconductor and a metal. As shown for several interfaces between an undoped organic semiconductor and a metal layer prepared in UHV, this is not always true.105 If the Fermi level fixing and its alignment with the metal Fermi level occur by atmospheric doping for an organic semiconductor, this assumption may become valid for practical devices prepared and used under atmospheric conditions such as ambient air. The doping process is expected to cause a characteristic band bending at organic semiconductor/metal interfaces. Several problems have remained in understanding whether or not this simple model is applicable to typical organic/inorganic and organic/ organic interfaces. Among these, the two main problems are (1) little is confirmed experimentally about the effects of the molecular orientation and/or packing structure on electronic structures near organic/organic and organic/inorganic interfaces and (2) it is not easy to clearly distinguish the electronic structure in the “contact interface” from that in the bulk region.106 Changes in the molecular orientation during film growth may cause unintentional dipole-layer formation in the film107 and unintentional changes in the ionization potential of the film.108 In contrast, several groups have reported the electronic structure of a well-defined organic/substrate contact interface.109-111 The organic/substrate interface can be divided into two adsorption systems:99 (1) a chemisorbed system, in which the interface dipole is formed by chemical interactions and electron transfer, which gives rise to pinning of the energy levels of the organic semiconductor to the substrate Fermi level, independent of the initial semiconductor work function, and (2) a physisorbed system, in which the interface dipole is formed by the push-back effect and electron transfer. The important aspects can be discussed using Figure 7. When a metal and an organic solid come into contact, a dipole layer may be formed right at the interface, due to

rearrangement of electronic charge. Several mechanisms have been identified to explain this type of interface dipole.38,112 Possible origins of the rearrangements are, e.g., charge transfer across the interface, redistribution of the electron cloud spilled out of the metal surface, mirror force, formation of an interfacial electronic state, interfacial chemical reaction, and aligned permanent dipoles of adsorbed molecules.99 Such interfacial dipole layer formation results in an abrupt shift of the potential across the dipole layer, leading to a shift of hypothetical vacuum levels (eVi) at the interface, as shown in Figure 7. If the organic semiconductor is doped, for example by an acceptor, as shown in Figure 7A, the dopant can be ionized to form a fixed space charge, and the energies of the occupied and unoccupied levels are shifted with increasing distance from the interface, indicating the space-charge region at the metal/semiconductor interface. However, most organic semiconductors are used, at least nominally, undoped and the electronic structure of the semiconductor is represented without surface “band bending”, as shown in Figure 7B. The presence of a space-charge region at the interface between a doped semiconductor and a metal, with a significant molecular level bending in the doped layer, has been amply demonstrated.39,113 The magnitude of the spacecharge region has also been shown to be larger in wide band gap semiconductors with low intrinsic charge carrier concentrations and low conductivity,114 which are common characteristics of organic semiconducting materials. Nishi et al. discussed the effect of bulk oxygen doping on the interfacial electronic structure of titanyl phthalocyanine films (TiOPc) deposited on various metal substrates.105 Films prepared under UHV did not show the alignment of the metal Fermi level with the Fermi level of the organic semiconductor at fixed energy within the band gap, as shown in Figure 7B, possibly because of insufficient charge density in the TiOPc films. On the other hand, there is a conversion from n- to p-type semiconductor seen, when the film is exposed to O2, as well as a clear alignment of the Fermi level together with the formation of a space-charge region and an upward molecular level bending (Figure 7A). Heeger et al. presented a model to understand the interactions at doped polymer/metal interface48,115 which was refined by Brazovskii et al.116 The model considers specifically the charge carriers in degenerated conjugated polymers. The nondegenerate continuum model of Davids et al.117 extended the work toward conjugated polymers with a nondegenerate ground state. Based on the relative energies between the Fermi level of the metal (EF) and the formation energy of polarons and bipolarons of the polymer, Eform,

Semiconductor Junction Gas Sensors

conclusions about the extent of charge transfer during interface formation can be made. If EF of the contact is higher in energy than the Eform of the negative bipolaron/polaron, the transfer of a large negative charge density into the polymer occurs, whereas a large positive charge density is transferred if EF of the contact is lower than Eform of the positive bipolaron/polaron. In addition, Davids et al. suggested that the presence of traps in the polymer will accommodate charges as well.117 After values are assigned to the polaron and bipolaron formation energies, as well as to the trap energies and densities, it is possible to model the potential and charge density at the metal/polymer interface for a given Fermi energy with respect to the center of the energy gap by solving the Poisson equation. As a result, the transferred charge remains close to the metal/polymer interface. By considering the metal/polymer interactions from the microscopic point of view of quantum interactions, a molecular modeling was carried out, which took into account the quasi-one-dimensionality of the polymer system.118-120 As a result of the difference in the Fermi levels of the metal and the conducting polymer, it is expected that charge transfer may occur across the interface. Furthermore, the quantum calculation approach considers specific chemical reactions or even charge donation between the metal and the polymer as a result of bringing metal atoms in the vicinity of a conjugated system and observing the changes in the electronic structure of the resulting system. Based on such calculations, it has been postulated that the metal plays an important role in determining the extent of charge transfer that occurs in metal/conducting polymer interfaces: Ca and Na appear to transfer charge at the interface without significantly altering its chemistry.118,119 Al, on the other hand, disrupts the chemical structure of the surface of the conducting polymer.119 The results obtained in the case of Ca and Na agree with the general behavior postulated by Davids et al.,117 whereas Al deposition introduces an additional effect of surface chemical reactions, which are not considered by the continuum model.

3.4. Gas/Solid Interactions in Schottky Barrier Junction Devices The key difference between organic and inorganic Schottky barrier junction devices is the ability of many gases and vapors to penetrate through the organic semiconductor either to change the Schottky junction resistance or to interact with the semiconductor itself, which causes a work function change of the material, as shown in section 3.4.1. In the case of an inorganic Schottky barrier junction device, gas permeation toward the metal/semiconductor interface works only for hydrogen or hydrogen-producing compounds and then gives rise to formation of a dipole layer, as discussed in section 3.4.2. Adsorption at the catalytic metal surface has only a limited secondary effect, i.e., the dissociation of molecular hydrogen to atomic hydrogen.

3.4.1. Bulk Effect upon Gas DopingsModulation of the Electron Work Function The rectifying behavior of the Schottky diodes relates directly to the Fermi level of the layer of the organic semiconductor and can be influenced not only by the structure, the dopant type, and the doping level,121 but also by the interaction of the semiconductor with gases or vapors. By exposure of the organic semiconductors to certain gaseous

Chemical Reviews, 2008, Vol. 108, No. 2 379

species exhibiting electron donor or acceptor behavior, the Fermi level position is changed27,30,122 by either a decrease or an increase in the doping level. In this case, a gas-induced change of the Schottky barrier characteristics can be observed, which can be exploited for gas sensing measurements. In 1980, Van Ewyk et al. proposed the formation of a charge-transfer complex due to an electron donator/acceptor interaction during the absorption of strong electron donating or accepting gases on organic p-type semiconductors.123 The absorption causes a change in the doping level of the semiconductor and thus a change in its conductivity. The transfer of a partial amount of electron density without complete ionization of the reacting species is generally known for adsorption processes of molecules on conducting surfaces and has been treated comprehensively, from both the catalysis124 and chemical125 points of view. It was assumed that this type of reaction is also responsible for the chemical modulation of the electron work function of the polymer layer due to gas absorption.126 This type of interaction process is often described as a secondary doping of the sensitive semiconducting layer. The primary doping process is carried out during formation of the semiconducting sensing layer by incorporating doping ions. In the case of an organic semiconductor, primary doping is carried out during the chemical or electrochemical preparation process.67,127 Janata presented a model that describes the potential concentration relationship based on the formation of a chargetransfer complex between the secondary dopant (gas or vapor) and the matrix (sensitive phase), which is combined with a fractional charge transfer δ.128 The dependence of the Fermi level position of the sensitive material on the partial pressure, pgas, of the dopant gas or vapor is given by

EF ) E/0 +

kT ln(pgas + const) 2δ

(25)

where in E/0 the donor/acceptor level of the organic semiconductor and the equilibrium coefficients of all relevant reactions are combined. For pgas ) 1, EF equals E/0, denoted as the value of the Fermi level under standard conditions. A change of the Fermi level is reflected in the change of the work function EWF of the semiconductor

∆EF ) ∆EWF

(26)

Equation 25 has the familiar form of the Nernst equation for ion and electron transfer across the interface of two condensed phases, except that they account for the fractional charge transfer δ. The magnitude and the polarity of the response depend on the value of the charge-transfer coefficient δ, and with 2|δ| < 1, a fractional value of the slope is possible. It depends on the ability of the entering gas or vapor molecules to exchange charge density with the semiconductor matrix either by oxidizing (2δ < 0) or reducing (2δ > 0) the active sites of the matrix. The gas/vapor behaves like an electron acceptor and donor, respectively. In the absence of interfacial layer and Fermi level pinning, the gas-induced change of the Fermi level causes a change in the Schottky barrier characteristics in the low bias region. The diode current is then exponentially related to the change of the Fermi level, i.e., the work function of the semiconductor and the applied voltage, according to eq 5.

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adsorbed species. In the steady state, the coverage is

()

c1 θ ) d1 (1 - θ)

Figure 8. Band diagrams of a catalytic metal/semiconductor Schottky diode without and with hydrogen adsorption, respectively.

Figure 9. Cross-sectional view of a typical geometrical arrangement of a Schottky barrier diode used for chemical sensing applications.

It is known that a number of metals, e.g., palladium or platinum, adsorb and dissolve hydrogen and that the adsorbed atoms change the work function of the metal surfaces. By using a thin metal film as the catalytic metal electrode of the Schottky diode, hydrogen and other gases that react with the catalytic metal can be detected. Hydrogen molecules are adsorbed and dissociated on the outer metal surface. Then, the atomic hydrogen permeates through the bulk lattice toward the metal/semiconductor, causing a perturbation at this interface that gives rise to a change in the sensor output signal. Two hypothetical steps were proposed to explain the detection mechanism:129 (1) the hydrogen atom is polarized at the metal/semiconductor interface, which gives rise to a dipole layer, or (2) an excess of charge states at the metal/semiconductor interface is created in the presence of hydrogen and reduces the Schottky barrier height. Figure 8 illustrates the band diagrams without and with hydrogen adsorption, respectively. The hydrogen-polarized layer, arising from the intrinsic electric field of the diode, redistributes the charges in the depletion region and abates the degree of band bending. Hence, the corresponding current is correlated with the number of hydrogen atoms adsorbed at the metal/semiconductor interface. When atomic hydrogen has been formed on the outer metal surface, which is exposed to the ambient (see Figure 9), an equilibrium between the hydrogen concentration at this metal surface and that at the metal/semiconductor interface is reached. Based on this model, the change in the J-V characteristics and the decrease in the Schottky barrier height are strongly related to the hydrogen concentration. Assuming that the adsorption is not affected by conditions outside the metal surface, the metal/semiconductor interface has similar adsorption properties as the free metal surface. The fractional coverage θ is equal at the outer surface and the inner interface in the steady state. At the outer surface, the following reaction occurs in pure hydrogen: c1

H2(g) 79 8 2Ha d

(27)

1

where (g) refers to gas-phase species and (a) refers to

pH2(g)1/2

(28)

where θ is the fractional coverage of hydrogen and pH2(g) is the hydrogen gas pressure. Equation 28 is valid, for both the metal/semiconductor interface and the outer metal surface. The change in the build-in voltage can be written as130

∆Vbi ) -µNadsθ/S

(29)

where µ is the dipole moment of an adsorbed hydrogen atom, Nads is the density of adsorption sites, and S is the permittivity of the semiconductor. Moreover, the increase of the interfacial charge density causes the change in the ideality factor. The analytical relation between the ideality factor and the interfacial charge density is given by131

n)1+

3.4.2. Formation of a Dipole and an Interfacial Layer

1/2

(

)

∆ s + eNssb i W

(30)

where ∆ is the interfacial layer thickness (see Figure 7), i and s are the permittivities of the interfacial layer and the semiconductor substrate, respectively, W is the depletion layer width, e is the electronic charge, and Nssb (cm-2 eV-1)131 is the density of the interface states that are in equilibrium with the semiconductor.

3.5. Extraction of Schottky Diode Parameters for Sensor Applications 3.5.1. Introduction In Figure 9, a cross section of a typical chemical sensor based on a Schottky junction is shown. Each Schottky barrier diode comprises two contacts, or junction areas, by definition: (1) between metal I and the semiconductor forming the Schottky barrier junction, which is the origin of the sensor signal, and (2) between metal II and the semiconductor forming the ohmic contact, which has to be inactive and, hence, should not contribute to the sensor response. The type of contact is defined by the work function of the materials used for fabrication (see section 3.2.1 and Figure 4). It should be noted that the area of the ohmic contact of the sandwich structure shown in Figure 9 is much larger than that of the Schottky junction, in analogy to auxiliary and working electrodes in electrochemical experiments. In Figure 10, the schematic band diagram and the equivalent circuit of a chemical sensor based on an ideal Schottky junction device, i.e., without an appreciable interfacial layer or interface states, is shown. The equivalent circuit of the Schottky barrier diode in Figure 10 shows those parts which can be influenced by the chemical interaction of gases with the chemical sensing material: (1) the barrier region (of thickness WJ) formed by metal I and the semiconductor having a capacitance CJ in parallel with the junction resistance RJ, and (2) the neutral bulk (of thickness WB) of the semiconductor with a bulk capacitance CB in parallel with the bulk resistance RB. Both metal I and the semiconductor, forming the Schottky barrier junction, can be used as the chemical sensing components interacting with gases and vapors. Their choice and combination defines the working principle as well as

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Figure 11. Typical change in the J-V characteristics of a Schottky barrier diode due to interaction with a gas or vapor of interest. ∆V indicates the gas or vapor induced voltage shift at constant current density.

Figure 10. Schematic band diagram and equivalent circuit of an ideal Schottky barrier diode.

the sensitivity and the selectivity of the sensor device. The generation of the “sensor response”, i.e., sensor signal, of the Schottky barrier diode is originated in the change of the junction characteristics (i.e., J-V or C-V characteristics) by alternating the Schottky barrier height or the built-in voltage. The response can be due either to adsorption of the species of interest at the metal surface affecting interfacial polarization by formation of a dipole layer or to absorption of gases or vapors of interest by the semiconductor and their interaction with the semiconductor, which changes its work function and, hence, the contact potential or built-in voltage of the diode. The former is mostly related to Schottky barrier diodes based on inorganic semiconductors and is discussed in section 4.1, whereas the latter is related to Schottky barrier diodes based on organic semiconductors discussed in section 4.2. It should be noted that the interaction of gases or vapors with the semiconductor, which causes a change of the doping level, leads not only to a change of the work function of the semiconductor, i.e., a change of the junction resistance RJ and the junction capacitance CJ of the Schottky barrier junction, but also to a change of the bulk resistance RB of the semiconductor (see Figure 10). In this case, the sensor device works as both a junction controlled device, with a diode current in the low bias voltage region exponentially related to the change of the built-in voltage, and a bulk controlled device in the high bias region, with a diode current linearly related to the change in conductivity. The latter represents a simple chemiresistor device. These compensatory effects are discussed in section 4.2.4.3. In general, the gas sensitivity of diodes is represented by the so-called gas-induced voltage shift, ∆V, extracted from current voltage characteristics and is defined as29,132

∆V ) Vgas - Vair at J ) constant

(31)

where Vgas and Vair are the applied voltages at a constant current density in gas and in air, respectively. In Figure 11 a typical change in the J-V characteristic of a Schottky barrier diode due to interaction with a gas or vapor of interest is shown. ∆V indicates the gas or vapor induced voltage shift

at constant current density. It can be either junction (i.e., ∆VJ) or bulk (i.e., ∆VB) controlled, as discussed above. The extraction of the sensing information from Schottky barrier diodes can be done by operation in constant current mode. Hence, the sensor response is the output voltage. Figure 20 shows a typical example of an experimentally obtained transient response of the output voltage of a Schottky barrier diode operated in the constant current mode.

3.5.2. Current−Voltage Characteristic Taking into account the image force between an electron and the surface of the metal, eq 7 can be rewritten as82

[

J0 ) A**T2 exp -

]

φb - ∆φbi kT

(32)

For metal/semiconductor contacts, the value of J0 can be readily correlated with a thermodynamic property of the interface, the barrier height of the junction, φb. Provided that the current-voltage characteristics are sufficiently close to ideal characteristics to allow a reliable value of J0 to be determined, the effective barrier height can be deduced in two ways: (1) If A** is known, the value of J0, found by extrapolating the straight part of the ln{J/[1 - exp(-eV/kT)]} vs V plot to V ) 0 (eq 19), immediately gives the effective barrier height φe ) φb - ∆φbi. A** is often not known with any great precision, but because an error of a factor of 3 in A** gives rise to an error of about kT/e in φe, an imprecise value can usually be used unless a very accurate measurement of φe is required. The determination of the zero-bias barrier height from current-voltage characteristics is only reliable if the semilog plot of J vs V is linear with a low value of n (1 < n < 1.1). For large values of n or a nonlinear plot of ln J as a function of V, the Schottky barrier is far from ideal. This is probably due to the presence of a thick interfacial layer and/or to recombination in the depletion region and/or to tunneling conduction. (2) If A** is not known, a plot of ln(J0/ T2) against T-1 should give a straight line of slope -φe/k and intercept on the vertical axis equal to ln A**. This is the most common case for the Schottky diodes based on organic semiconductors, because of the lack of the effective mass of charge carriers. The barrier height is generally a decreasing function of temperature, because the expansion of the lattice causes a change in the band gap. To a first approximation, one can write φe(T) ) φe(0) - bT, in which case the method gives the barrier height at absolute zero.

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A problem arises if the diode has a large series, i.e., bulk resistance (see Figure 10), or the semiconductor is low doped. This is typical for most polymeric Schottky diodes.133 The resistance is mainly due to the resistivity of the polymer film. In this case, the region over which the plot of ln J vs V is linear may be small, and accurate extrapolation to zero voltage may be difficult. Moreover, the polymer surface cannot be exposed to any chemical etching process; therefore, it can be expected that irregularities will be brought into contact with the metal and the effective area of the junction will be somewhat uncertain. To circumvent these difficulties, the use of a Chot plot,134 which is a modified Norde plot, has been suggested by Ingana¨s135 to determine the effective barrier height.

3.5.3. Capacitance−Voltage Characteristic An independent way to determine the barrier height involves the measurement of the differential capacitance of the Schottky barrier contact. The procedure is known as the Mott-Schottky or capacitance-voltage method. According to the Schottky theory, the depletion layer capacitance (/F‚ cm-2) of the metal/semiconductor contact

C)

δQ δVr

(33)

can be expressed as81

C

-2

)

(

2 Vbi - V -

kT e

)

eSN

(34)

where Q is the charge associated with the band bending, Vr is the reverse bias voltage, V is the applied voltage, and N is the free charge carrier concentration. If φb is independent of V (i.e., if there is no appreciable interfacial layer), a plot of C-2 against V (Mott-Schottky plot) should give a straight line with an intercept -VI on the horizontal axis equal to -(Vbi - kT/e). The barrier height is then given by

φb ) eVbi + ξ ) eVI + kT + ξ

(35)

The free charge carrier concentration, N, can be determined from the slope of the C -2 vs V plot. The depletion layer width, W, is given by

W)

( ) 2sφbi eN

1/2

( )[ ] -2 dV 2 C es d

tion, which disappears at low frequency. A compromise must be found to minimize these two contributions. This method of barrier height determination is generally straightforward for metal/semiconductor contacts, yielding the flat-band barrier height, which does not include the image force lowering present at the junction. However, when some of the applied voltage is dropped across an interfacial dielectric layer, or when traps are present in an interfacial layer, the capacitance-voltage data will not generally adhere to the simple relationships of eq 33 and eq 34, respectively. Hence, the capacitance method gives results which can differ significantly from those obtained by the current-voltage methods.

3.5.4. Impedance Measurement To gain more insight into the nature of junctions and to obtain the equivalent circuit models for diodes, the complex impedance is usually measured as a function of frequency at different bias voltages.137 This methodology is analogous to impedance analysis used in electrochemical experiments. It has been proposed that an ideal Schottky diode (i.e., without an appreciable interfacial layer) can be divided into two regions: (1) the barrier region (of thickness WJ), with a capacitance CJ in parallel with the junction resistance RJ, and (2) the neutral bulk (of thickness WB), with a bulk capacitance CB in parallel with the bulk resistance RB (see Figure 12A).133 In most practical metal/semiconductor contacts, the conditions RJ . RB, CJ . CB, and ωmaxτB , 1 are satisfied.133,138 The plane plot of the complex impedance is shown in Figure 12B. From Figure 12B, the bulk resistance RB and the junction resistance RJ can be estimated by the intercepts of the high- and low-frequency extrapolations of the semicircle with the horizontal axis, respectively. Using these values of RJ and RB, the junction capacitance (for each frequency) and thus the depletion width can be fitted according to

(36)

If N is not constant, the plot is no longer linear. In this case, the differential capacitance method can be used to determine the doping profile. The derivative of eq 35 gives81

N)

Figure 12. (A) Common equivalent circuit of a Schottky diode junction. (B) Impedance plane plot. Arrows indicate the direction of increasing frequency. The dotted lines imply that the bulk capacitance CB is often not appreciably distributed.

Zr )

RJ 1 + ω Rj Cj

Ri ) (37)

In practice, the junction is reverse polarized, and small oscillations δV are superposed. If no trap exists, the variation of the charge associated with the band bending δQ corresponds to the variation of the free charge carrier concentration at the edge of the depletion zone. If traps exist, they contain part of the charge δQ. To get rid of them, one has to perform measurements at high frequency (>100 kHz).136 For such a frequency, only the semiconductor depletion charge changes. Another problem then arises: the bulk resistance contribu-

2

2

2

+

ωRJ2CJ

RB 1 + ω2RB2CB2

1 + ω2RJ2CJ

2

ωRB2CB 1 + ω2RB2CB2

(38)

(39)

In most practical metal/semiconductor contacts, the ideal situation shown in Figure 12A is never reached because there is usually a thin insulating layer with a capacitance CI in parallel with the resistance RI (see Figure 12B). The insulating layer has three effects:82 (1) because of the potential drops in the layer, the zero-bias barrier height is lower than it would be in an ideal diode, (2) the electrons have to tunnel through the barrier presented by the insulator so that the current for a given bias is reduced in a manner equivalent to a reduction in the Richardson constant A**,

140

1.7 × 1.33/-0.02 (500 ppm) 0.59/-0.005 (800 ppm) -/-0.01428 (500 °C) -/-0.00547 (500 °C) 1.73/-0.05 13.5

RT

synth air

1.7 × 10-4/2.1 × 10-4

h h ) homogeneous. a

sat. current/gas induced sat. current/A

Pt-Pd/(porous)Si

From the rapid progress in semiconductor materials and devices, solid-state hydrogen sensors based on the Schottky diode structure with a catalytic palladium metal have attracted intensive interest.10,140 Silicon based Schottky diodes are inexpensive, and they are compatible with monolithic silicon integrated circuit technology. The measured diode current

h h h h h

4.1.1. Silicon Based Schottky Diodes

Table 1. Electrical Parameters of Schottky Barrier Diodes Based on SiC and Si

In the field of solid-state chemical sensors, Schottky diodes and related devices (MOSFET, MIS structures, etc.) incorporating catalytic metals have been widely studied for their application in the detection of hydrogen and hydrogenproducing compounds.6 In hydrogen sensors based on Schottky diodes, the key role is played by the Schottky metallization, mostly belonging to the platinum group metals, which ensures the catalytic dissociation of molecular hydrogen into hydrogen atoms (eq 27). It should be noted that the semiconducting substrate is only used to provide sufficiently high Schottky barrier heights, necessary for good sensing performance, but is not the gas-sensitive part of the sensor. According to the dipole model, originally developed by Lundstro¨m,139 the hydrogen sensitivity of catalytic gate devices is based on the dissociation of hydrogen molecules on the catalytic metal surface and the diffusion through the metal film to form a polarized layer at the metal/insulator interface (see section 3.4.2). The polarized layer gives rise to a shift of the capacitance-voltage (C-V) curve of a MOS capacitor or the current-voltage (J-V) curve of a Schottky diode. Catalytic gate devices also respond to hydrogencontaining molecules such as hydrocarbons, provided that the molecules are also dissociated on the catalytic metal surface. Oxygen atoms are also dissociated on the catalytic metal surface. Water formation with oxygen atoms from oxygen-containing molecules consumes hydrogen and, therefore, decreases the sensor response. In other words, catalytic metal gates have a direct response to hydrogen and hydrocarbons as well as an indirect response to oxygen molecules, whose effect is to decrease the direct response. Besides Si, compound semiconductors such as GaAs, InP, SiC, and GaN have been alternatively employed as substrate materials for Schottky diode type hydrogen sensors. Tables 1-3 give an overview of the electrical and gas sensing parameters of the Schottky barrier diodes mentioned above.

ideality factor/gas induced shift, n/Dn

4.1. Schottky Diodes Based on Inorganic Semiconductors

synth air synth air synth air synth air synth air

4. Schottky Diodes and Their Sensor Applications

202-400 500 400-600 400-600 RT

(41)

Pt/n-SiC Pt/n-SiC Pd/-SiC Pt/-SiC Pt-Pd/(polished)Si

e δV kT δ(ln J)

background atmos

n)

barrier height/gas induced shift, φb/∆φb (eV)

in a manner which can be described in terms of an ideality factor n (greater than unity) defined by eq 41

NO/10-500 ppm H2/800 ppm H2/800 ppm CH4/4800 ppm NO2/5-20 ppm NO/50-250 ppm NO2/5-20 ppm

charge carrier density, NA, ND (cm-3)

(40)

operating temp (°C)

kT e

gas/conc

for V .

metal layera

eV (nkT )

structure metal/semiconductor

J ) J0 exp

ref

and (3) when a bias is applied, part of the bias voltage is dropped across the insulating layer, so that the barrier height φb is a function of the bias voltage. This effect changes the shape of the current voltage characteristics of the tunneling current given by81

170 11 165 165 140

Chemical Reviews, 2008, Vol. 108, No. 2 383

1016

Semiconductor Junction Gas Sensors

structure metal/semiconductor Pt/n-GaAs Pt/n-GaAs Pt/n-GaAs Pd/n-GaAs Pd/n-GaAs Pt/n-GaAs Ir/n-GaAs Pt/Al0.3Ga0.9As Pd/Al0.3Ga0.9As Pd/n-GaAs Au/n-GaAs Au/n-GaAs Pd/n-GaAs Pd/n-GaAs (MOS)

metal layera h h h h h p p h h h p p h h

operating temp (°C)

gas/conc H2 /1000 ppm H2 /23-1350 ppm H2 /23-1350 ppm H2 /23-1350 ppm H2 /23-1350 ppm NH3 /6-95 ppm NH3 /23-1350 ppm H2 /49-9090 ppm H2 /49-9090 ppm H2 /1-5% CO /1-5% NO2 /1-5% H2 /500-1500 ppm H2 /5-5000 ppm

50-200 °C 150 °C 50 °C 150 °C 50 °C 150 °C 150 °C 30-120 °C 30-120 °C RT RT RT-250 °C 80 °C/110 °C

background atmos

sat. current/gas induced sat. current/A

synth air synth air/N2 synth air/N2 synth air/N2 synth air/N2 synth air/N2 synth air/N2 synth air synth air synth air N2 N2 synth air N2

6 × 10-10/9 × 10-9 A

ideality factor/gas induced shift, n/∆n

barrier height/gas induced shift, φb/∆φb (eV)

1.57/+0.16 (500 ppm) 1.083/-0.009 (1010 ppm)

charge carrier density, NA, ND (cm-3)

-/-0.19, -0.15 -/-0.02 (100 ppm, N2) -/-0.001 (100 ppm, N2) -/-0.075 (100 ppm, N2) -/-0.025 (100 ppm, N2)

4 × 10

-/-0.074 (9090 ppm) -/-0.07 (9090 ppm) -/-0.28 (5%)

2 × 1017 2 × 1017 2 × 1017 1 × 1015 1 × 1015 1 × 1015 1.5 × 1017

0.95(RT)/-0.08 (500 ppm) 0.786/-0.046 (1010 ppm)

ref

14

7 8 8 8 8 8 8 148 148 147 151 151 150 150

384 Chemical Reviews, 2008, Vol. 108, No. 2

Table 2. Electrical Parameters of Schottky Barrier Diodes Based on GaAs and AlGaAs

Table 3. Electrical Parameters of Schottky Barrier Diodes Based on InP and GaN structure metal metal/semiconductor layerb

gas/conc

Pd/InP (MOS) Pd/n-InP Pd/n-InP (MOS) Pd/n-InP Pd-Ge-Pd/p-InP (pseudojunction) Pd*/n-InPa Pt/AlGaN/GaN

h h h h h

H2 /5-5000 ppm H2 /100 ppm H2 /15-1010 ppm H2 /15-1010 ppm NO2/10 ppm

h h

Pt/AlGaN/GaN IrPt/AlGaN/GaN PdAg/AlGaN/GaN Pt/GaN

h h h h

H2/5 vol % H2/5 vol % H2/5 vol % H2/500-2000 ppm

80 °C/110 °C RT RT-110 °C RT-110 °C RT-100 °C

background atmos

sat. current/gas induced sat. current/A

ideality factor/gas induced shift, n/∆n

N2 synth air synth air synth air synth air

1.072/-0.047 (1010 ppm) 7.5 × 10-8/4.3 × 10-6 1.8/-0.88

H2 /15 ppm-1 vol % RT-118 °C CO/8-160 ppm 250-300 °C

synth air synth air

/3.2 × 10-6

1.12/+2.55

200-800 °C 200-800 °C 200-800 °C 400 °C

N2 N2 N2 synth air

3.5 × 10-12

1.5 1.4 1.6 1.37

Pd* indicates electroless plated. b h ) homogeneous.

1.06/+0.14

barrier height/gas induced shift, φb/∆φb (eV)

charge carrier density, NA, ND (cm-3)

0.677/-0.21 (1010 ppm) 0.72/-0.19 (100 ppm) 0.63/-0.15 (1010 ppm) 0.42/-0.02 (1010 ppm) 0.906/-0.039 (10 ppm)

1.5 × 1016 1.5 × 1017 1.5 × 1017 5 × 1016

150 10 9 9 155

1017

152/153 163

0.588/-0.177 (5000 ppm) 1.01/-0.06 (250 °C, 8 ppm) 1.12/-0.03 (300 °C, 8 ppm) -/-0.0114 -/-0.0246 -/-0.0941 1.15/-

1017

ref

162 162 162 161

Potje-Kamloth

a

operating temp (°C)

Semiconductor Junction Gas Sensors

varies exponentially with the lowering of the hydrogeninduced Schottky barrier height.141 Therefore, these Schottky diodes often exhibit excellent hydrogen detection sensitivities. The formation of palladium silicide at the Pd/Si Schottky interface leads to the degradation of detection capacity.142 An ultrathin (∼30 Å) oxide layer introduced into the Pd/Si interface prevents the formation of palladium silicide,142,143 forming a conducting MIS (metal-insulator-semiconductor) diode. However, the sensitivity toward hydrogen is significantly decreased, if the oxide layer is not properly defined. Therefore, the high quality of the oxide layer is a key factor for the device performance. This certainly increases the fabrication complexity and the cost. In the MIS configuration, the Schottky barrier height of the Pd/Si system is found to depend quite strongly on the Pd effective work function and hence on the hydrogen ambient. Zhang at al. investigated Pt-Pd/Si-Al Schottky diode structures for detection of NO2 from 6 to 22 ppm gas concentrations and of NO from 50 to 250 ppm gas concentrations, respectively. The sensor operates at room temperature. The response is roughly proportional to the logarithm of the gas concentration.140 Mesoscopic polysilicon wires coated with a thin film of palladium have been used as hydrogen sensors. The wires show an increase in their electrical resistance, when hydrogen reacts with the sensitive layer.144 Porous silicon obtained from p+-type silicon wafers were impregnated with Pd nanoparticles using an electroless process. It has been shown that the distribution of nanoparticles over the porous media creates a higher gas sensitivity toward hydrogen.145

4.1.2. GaAs Based Schottky Diodes In a metal/semiconductor junction, the height of the Schottky barrier depends on the work function difference between the metal and the semiconductor according to eq 4. However, as for many III-V semiconductors, the observed value of the barrier height of a metal/GaAs contact does not follow the simple theory in terms of the metal work function and the electron affinity of GaAs due to the high density of surface states, which causes a pinning of the Fermi level. Many models have been developed to explain this discrepancy. Among them, the effective work function model by Freeouf et al.146 emphasizes the role of the intermediate layer formed by the reaction between the metal and GaAs to determine the barrier height. However, some authors have found evidence of unpinning of the Fermi level in metal/ GaAs interfaces fabricated on clean (100) GaAs surfaces. This offered a way to fabricate metal/GaAs Schottky diodes with high sensing capability.147,148 Lechuga et al. reported a Pt/GaAs Schottky diode reaching a low detection limit for hydrogen of 6 ppm in nitrogen and 200 ppm in air.7,8 Liu et al. proposed the fabrication of a planar-type Pd/GaAs Schottky diode as a hydrogen sensor exhibiting an interfacial oxide, which weakens the Fermi level pinning and, hence, improves the barrier height.149 The use of a porous GaAs semiconductor has been investigated to improve the performance of Schottky diode hydrogen sensors, both in terms of sensitivity and response. Pd/porousGa/As and Au/porous-GaAs Schottky junctions have been fabricated, which show sensitivity toward hydrogen and polar gases, e.g., CO, NO, respectively.150,151 Pd/porous-GaAs Schottky diode gas sensors exhibited large sensitivity of more

Chemical Reviews, 2008, Vol. 108, No. 2 385

than three times higher than that of Pd/GaAs contacts with a response time of 1 s toward 500 ppm hydrogen gas. It has been found that this promoting effect could be related to the formation of pore structures (effective adsorption centers) in the Pd contact induced by the porosity of the GaAs wafer. A urea sensor has been presented using an ammonia gas sensor based on a Pt/GaAs Schottky diode and a ureaseimmobilized membrane. The authors used discontinuous platinum films which proved to be effective and reliable in solid-state chemical sensors for the detection of NH3 gas.8 The Schottky diodes have been coated with an organic layer in order to protect the devices during measurements in solution, as well as to use it as a support for biological material.

4.1.3. InP Based Schottky Diodes Pd/InP Schottky diodes have been regarded as promising devices for hydrogen sensing.149,152 At room temperature, it has been found that the Pd/InP Schottky diode exhibits superior sensing performances for the detection of hydrogen. Yousuf et al. demonstrated essentially high hydrogen response in the J-V characteristics of a Pd/InP Schottky diode.10 Although enormous current variations are observed, the low barrier height associated with the high defect state density at the Pd/InP interface severely restricts the allowable variation in the barrier height. Recently, electroplating techniques have also been proposed to fabricate the Pd/InP Schottky diodes.153 Although uniformity and adhesion related problems must still be solved compared with traditional vacuum deposition techniques, electrochemical techniques do exhibit the advantages of easy operation, simple equipment, and low cost.154 Particularly, because of the specificity the low-energy process, the electrochemical techniques can eliminate the Fermi-level pinning effect and lead to well behaved Schottky contact properties, which can remarkably improve the sensing performances of Pd/InP Schottky diodes. Talazac et al. have presented a different type of Schottky based gas sensor.155 The detection mechanism is based on the electron exchange of the gaseous molecules with a thin semiconducting layer between the Schottky contact and the semiconducting substrate forming a pseudo-Schottky barrier diode. High sensitivities in the ppb range toward NO2 and O3 could be found. However, the sensors suffer from longterm aging effects.

4.1.4. GaN and AlGaN Based Schottky Diodes GaN and AlGaN are attractive as materials for the fabrication of various chemical and biochemical sensors:156 the bulk and surface properties are chemically stable due to large band gap energies, the materials allow sensing operation at high temperatures, and the materials are environment friendly. The Schottky barrier heights are much more dependent on the metal work function157 than those for other III-V materials, indicating a weaker Fermi level pinning. However, Schottky diodes formed on GaN and AlGaN materials exhibit excess reverse leakage currents, which are many orders of magnitude larger than the prediction of the standard thermionic emission model (see section 3.2.2). Another issue related to the metal/semiconductor interface is a poor ohmic contact. Large leakage currents in Schottky diodes are particularly problematic for constructing high-performance sensors.158 As a model which explains large leakage currents in nitride-based Schottky barriers, Hasegawa et al. recently

386 Chemical Reviews, 2008, Vol. 108, No. 2

proposed a thin surface barrier model159 involving donors near the surface. Due to the presence of ionized high-density donors, the width of the Schottky barrier is greatly reduced, and the electrons tunnel through this barrier by the thermionic field emission or the field emission mechanism. On free surfaces of GaN and AlGaN, high-density surface states exist, which cause charge-discharge transients leading to performance instability, such as current collapse, and poor long-term reliability. Recently, experimental and theoretical investigations on the role of oxygen impurities in GaN and AlGaN have indicated that oxygen acts as a shallow donor close to the surface with an activation energy of about 30 meV.160 Therefore, the elimination of oxygen from AlGaN and GaN layers is highly desirable prior to the sensor fabrication in order to realize high-performance sensors, e.g., by using an ultrathin Al layer as a getter material for oxygen.157 Schottky diodes on AlGaN/GaN heterostructures with Pt, IrPt, and PdAg catalytic metals have been fabricated and characterized from 200 to 800 °C for hydrogen sensing operation.161 The hydrogen sensitivity of Pt and IrPt diodes improves with the increase in temperature due to a more effective hydrogen dissociation. The sensitivity of PdAg diodes degrades with increase in temperature due to thermal instability of PdAg.162 Schottky gas sensors for CO were fabricated using a PtAlGaN/GaN Schottky diode. The CO sensors show good sensitivity in the temperature range 250-300 °C.163

4.1.5. SiC Based Schottky Diodes The operating temperature of gas sensors using silicon substrates is limited below 250 °C due to the small band gap of silicon. This restricts their use in specific environments such as automotive, aeronautical, and environmental areas. Semiconductor substrates with a large energy band gap such as silicon carbide and diamond can be used for sensors operating at high temperatures. The use of catalytic metal gate SiC devices as gas sensors for hydrogen and hydrocarbons has been extensively investigated by several groups. Janson et al. reviewed the effect of hydrogen on the wide band gap semiconductor SiC and its influence on catalytic metal/SiC Schottky diodes.164 Kim et al. investigated hydrogen and methane gas sensors using Pt-SiC and Pd/SiC Schottky diodes operating at temperatures in the range 300-500 °C.11,165 Annealing effects on Pd/6H-SiC Schottky diodes have also been investigated.166 The studies indicated that the response of the diode to hydrogen degraded after annealing at 425 °C due to interfacial diffusion of silicon into the Pd film region after dissociation from SiC. Palladium silicides are formed as the major interfacial products. The use of a Pd0.9Cr0.1 alloy film deposited on a 6H-SiC epilayer results in Schottky diodes exhibiting a stable catalytic surface and a significant improvement of the device sensitivity.167 Some reports deal with the response of catalytic metal gate SiC devices to nitric oxides (NOx). Zubkans et al.168 reported the direct response to NOx, which could be increased after a treatment with ammonia. Katsube et al. showed that thin catalytic metal gate Schottky diodes and p/n heterojunctions on Si substrates can be used for direct NOx detection at room temperature.169 They extended their investigations using SiC substrates for high-temperature operation170 for direct detection of NOx at elevated temperatures up to 400 °C at

Potje-Kamloth

minimum concentrations down to 10 ppm. The Langmuir adsorption model can describe the response of these devices.

4.1.6. CdSxSe1.x Based Schottky Diodes It has been shown that the bulk photoluminescence of semiconductors can be used to probe the Schottky barrier characteristics of both semiconductor/metal diodes171 and photoelectrochemical cells.172 In particular, the electric field in the semiconductor can be estimated from the PL intensity using a dead-layer model: electron-hole (e--h+) pairs, formed within a distance from the interface in the order of the depletion width, do not contribute to the bulk photoluminescence. Schottky diodes have been fabricated with Pd because of the known sensitivity of the current-voltage characteristic to hydrogen. Carpenter et al.173 showed that, besides the decrease of the Pd work function on exposure to hydrogen, the variation in Schottky barrier height with hydrogen strongly influences the bulk photoluminescence intensity of the Pd/CdS Schottky diode. The bulk photoluminescence intensity is enhanced upon exposure to hydrogen. In air the bulk photoluminescence intensity of the Schottky diode returns to its original value. The spectral changes are consistent with a reduction in Schottky barrier height resulting from the dissolution of hydrogen in Pd. Qualitatively, the bulk photoluminescence is expected to rise because the smaller electric field in the semiconductor allows a larger fraction of e--h+ pairs to radiatively recombine. By regarding the region supporting the electric field as being completely nonemissive, a quantitative expression for the relative intensity of the bulk photoluminescence can be obtained:

φair/φH2 ) exp(-R∆D)

(42)

where φair and φH2 are the radiative quantum yields in air and in hydrogen, respectively, ∆D is the difference in deadlayer thickness between the two media, and R is the sum of the solid’s absorptivities for the exciting and emitted light, respectively. ∆D can be used to calculate the reduction of the Schottky barrier height. By substitution of the depletion width W by ∆D, eq 36 gives the change in barrier height upon exposure to hydrogen.

4.2. Schottky Diodes Based on Organic Semiconductors Electrical conductivities of organic semiconductors can be varied over the full range from insulator to semiconductor through p-doping or n-doping (see section 2.2). They offer a viable alternative to conventional inorganic semiconductors in many applications because of their unusual electrical properties, diversity, ease of fabrication, large area, and potentially low cost. Much research has been carried out on the use of organic semiconductors as active materials in electronic devices. Fabrication of electronic devices, such as OFETs, OLEDs, and devices based on Schottky barriers is the most important application of these materials. The use of different semiconducting organic films for the detection of gas and vapors has been reported by a number of groups.27-30,174 Whereas the chemical stability in a given temperature range and in ambient environment is a basic requirement of organic semiconductors to be used as active material in electronic devices, the modulation of physical properties of the organic semiconductors as a function of gas or vapor concentration is a prerequisite to be used as

Semiconductor Junction Gas Sensors

Chemical Reviews, 2008, Vol. 108, No. 2 387

Table 4. Parameters of Molecular Semiconductor Schottky Barrier Diodes Based on Phthalocyanines (Pc) doping

barrier height φb /eV

charge carrier density Ns /cm-3

oxygen as deposited oxygen

1.03 1.09 0.65

oxygen as deposited oxygen

1.2

1.7 × 1019 7.5 × 1018 2.2 × 1017 5.1 × 1018 1.6 × 1018 3.4 × 1017 1.1 × 1016

device config Au/AlPcCl/Al ITO/AlPcCl/Al Au/FePcCl/Au Au/TiOPc/HOPG Au/NiPc/Al

p-type p-type p-type p-type p-type n-type

depletion width W/ nm

169

charge carrier mobility µ /m2 V-1 s-1

ref

1.73 × 10-8 1.58 × 10-8 1.2 × 10-6

194

5.8 × 10-6

176

196 193 105

Table 5. Electrical Parameters of Polymer Schottky Barrier Diodes Based on Polythiophene (PT), Poly(3-alkylthiophenes) (P3AT)a, Polyanilines (PANI), and Polypyrroles (PPy)a structure Al/PT/Au Al/P3DT/Au-Sn Al/P3HT/Au Al/P3HT/ITO Al/P3MeT/Pt Al/P3MeT/Au Al/P3OT/ITO Al/P3OT/Au

doping agent ClO4 undoped undoped PF6 reduced BF4 reduced PF6 undoped FeCl3 undoped

ideality factor n 13.0

contact or built-in potential φbi/eV PolythiophenessPT 0.38

barrier height φb/eV 0.66 0.4

3.9 3.0 4.5 2.5

0.35 1.2 0.6

6.1 4.9 1.2

0.52 0.7

0.34 0.75 1.5 0.8 0.45 1.01 0.72

charge carrier density NA/cm-3 3.0 × 1014 1.0 × 1017 3.5 × 1015 6.4 × 1019 4.6 × 1016 2.9 × 1017 7.0 × 1016

ref 211 217 91 205 216 239 239 205 205 203 204

PolyanilinessPani Al/Pani/Au

Al/Panib/Au Al/Pani/ITO In/Pani/Pt Al/PPy/Au Al/PPy-NMPy/Au In/PPy-NMPy/Au Au/PPy/Pt

HCl TOS PAA undoped HCl HCl DBS/HCl

1.24 1.3 1.7 1.95 1.46-1.64 4.2 3.12

BF4 BF4 ClO4 CuPcTS PbPcTS NiPcTS CoPcTS

4-7 1.4-2 1.2 3.6 3.8 4.0 4.2

0.72 0.82 0.9 0.65 0.69 PolypyrrolessPPy ∼0.7 0.78-0.82 0.23 0.26 0.24 0.23 0.2

0.38 0.78 0.87 1.03 0.33-0.41 0.76 0.803 0.7-0.76 0.7 0.81 0.63

2.8 × 1018 1.4 × 1018 1.8 × 1017 3.1 × 1017 3.06 × 1017 1.0 × 1022 1016-1017 7.0 × 1017 1.2 × 1020 4.6 × 1019 3.7 × 1019 1.9 × 1019

222 220 222 228 221 185 185 182 186

a ClO ) perchlorate; CuPcTS ) copper phthalocyanine tetrasulfonate; PbPcTS ) lead phthalocyanine tetrasulfonate; NiPcTS ) nickel 4 phthalocyanine tetrasulfonate; CoPcTS ) cobalt phthalocyanine tetrasulfonate; DBS ) dodecylbenzenesulfonic acid; D ) dodecyl; O ) octyl; M ) methyl; H ) hexyl; NMPy ) N-methylpyrrole. b Modified with metal-2.5-diaminobenzenesulfonic acid.

the sensing element in chemical sensors. However, the details of the gas response mechanism have not been completely elucidated. From a practical point of view, the gas sensing characteristic is strongly influenced by a number of processing parameters: for example, the growth rate, the thermal annealing,29 the polymerization temperature,175 and the nature of the dopants.132 By exposure of the polymers to certain gaseous species exhibiting electron donor or acceptor behavior,176,177 the Fermi level position is changed by either increasing or decreasing the doping level according to section 3.4.1. In this case, a gas-induced change of the Schottky diode barrier characteristic can be observed. Furthermore, the increase or decrease in the ideality factor during gas/ polymer interactions, which is due to a generation or reduction of the interface state density, can influence the gas sensitivity of the junction as well. Despite these advances in the use of chemical sensing applications, the detailed properties and roles of interfaces in organic semiconductor devices remain elusive. Parker suggests that charge injection into organic materials proceeds via tunneling across interfacial barriers.178 Improvement of device performance through interface modification has also

been reported.179 Up to now, most Schottky diodes based on organic semiconductors were fabricated and characterized using metals with low work function, such as Al, In, or Ca.27,180-182,217 These diodes show unstable electrical and chemical sensing properties, which have been attributed to the high chemical reactivity of the metal used.183 The nature of the interfacial layer between these metals and the semiconducting polymers is commonly difficult to define.181 In Tables 4-7, diode parameters of organic Schottky diodes are given. The data show that, unlike the case of conventional metal/semiconductor contacts, the electrical properties of contacts based on organic semiconductors can be manipulated not only through the choice of materials, methods, and conditions of preparation but also through the type of dopants incorporated into the organic semiconductor (see section 2.2). This can be exploited to tune the diode characteristics in a definite way. On the other hand, a careful control of the preparation parameters is necessary to obtain reproducible results. The choice of the metal forming the ohmic contact depends on the work function of the organic material used. For the commonly used polymer-based diodes, consisting

388 Chemical Reviews, 2008, Vol. 108, No. 2

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Table 6. Electrical Parameters of PPy-MPcTS Schottky Barrier Diodes Extracted from J-V and C-V Characteristics, Measured in Air and Exposed to 11 pm NO2 saturation current density J0/A cm-2

ideality factor n

rectification ratio (at |V| ) 1 V)

built-in voltage Vbi/V

dopant

air

NOx

air

NOx

air

NOx

air

NOx

CuPcTS PbPcTS NiPcTS CoPcTS TOS

8.1 × 1.5 × 10-3 6.5 × 10-3 1.0 × 10-2 2.5 × 10-2

7.0 × 7.9 × 10-4 2.4 × 10-3 4.8 × 10-3 1.1 × 10-2

3.6 3.8 4.0 4.2 4.9

3.2 3.5 3.7 3.8 4.3

7.2 5.2 3.7 3.2 2.1

13.5 8.7 5.1 4.7 2.9

0.26 0.24 0.23 0.20 0.15

0.29 0.27 0.26 0.22 0.17

10-4

10-5

Table 7. Electrical Parameters of PPy-MPcTS Schottky Barrier Diodes Extracted from J-V and C-V Characteristics: Change of NO2 Sensitivity for Different Polymerization Temperatures of PPy, Measured in Air and in 11 pm NO2 polymerization temp/K

rectification ratio at |V| ) 0.9 V

saturation current density J0/A cm-2

ideality factor n

charge carrier conc NA/1019 cm-3

278 283 288 293 303 313 323 333

2.1 3 5 6.3 27.3 156 307 68

3.1 × 10-2 6.3 × 10-3 1.9 × 10-3 7.9 × 10-4 1.2 × 10-4 7.6 × 10-5 5 × 10-5 1 × 10-4

4 3.7 3.4 3.6 3.3 3.1 2.5 2.9

97.5 52.6 37.1 12 2.1 0.82 0.37 1.4

of poly(3-octylthiophene),27 poly(3-methylthiophene),184 and PPy,29,185,186 gold was used to form an ohmic contact. The work function of gold has been assumed to be about 5.1 eV. Several studies show that the work function of π-conjugated polymer films is about 5 eV, depending upon the dopant type and the doping level.64,126 In this case, the barrier between the p-type polymer and gold is relatively small but is predicted to be negative, representing a quasi-ohmic contact, with a symmetric current-voltage dependence. Nevertheless, the experimental contacts are sometimes rectifying,133 indicating a higher work function of the polymer than that of gold. Hence, the junction characteristics of the ohmic gold/polymer barrier are strongly influenced by the gas sensing behavior of the polymer layer, exhibiting either symmetric or a non-ohmic, i.e., asymmetric, current-voltage characteristics. To overcome this uncertainty, platinum is recommended as metal to form an ohmic contact with the doped p-type conducting polymers.

4.2.1. Phthalocyanines Phthalocyanines (Pc) are a class of organic compounds of high thermal and chemical stability, which are classified as p-type semiconductors characterized by a low mobility and low charge carrier concentration.187 They have potential advantages for use as active layers in electroluminescent,188 photovoltaic,189 and gas sensor devices,190 because they are easily processable in low cost and large area device fabrication.191 However, it is known that oxygen profoundly influences the electrical properties of phthalocyanines. It has been demonstrated that the dark rectification ratio of Au/Pc/M (M ) Au, Cu, Cr, Al) is related to the presence of O2.192,193 When the devices are made and studied under vacuum, no rectifying effect is observed. However, by exposure to air, a strong rectifying effect has been found. Acceptor levels are generated within the band gap of these materials in the presence of O2, and hence, their thermal activation energy is lowered and the conductivity is enhanced.192,194

Besides O2, any other electron acceptors, such as NOx and Cl2, induce acceptor levels, while electron donors, such as H2 or NH3, remove them. Both change the Schottky diode characteristics. With most of the phthalocyanines, gold is found to form the ohmic contact,14,15 and aluminum is found to form the blocking contact.192,195 Under forward bias, oxygen-doped phthalocyanine thin films show ohmic conduction at low voltages and space charge limited conduction controlled by an exponential distribution of traps above the valence band edge at higher voltages.192,194,196 As already discussed above, almost all phthalocyanines show excellent gas-sensing properties for both electron acceptors, such as NOx,197-199 HCl,197 Cl2200 and electron donors, such as H2 or NH3.105 The change in the doping level induces a Fermi level shift and a change in the Schottky barrier height. It is supposed that a charge-transfer complex128 between the organic semiconductor and the gas or vapor is formed during the interaction. Spectroscopic investigations show the accumulation of the gas species in the bulk of the sensitive film and the formation of organic radical cations.197,201 Studies of phthalocyanines indicated that the sensitivity and selectivity of the sensor device are strongly influenced by the nature of the central metal ion and also by the crystal structure and/or crystallinity. Nieuwenhiuzen et al.202 reported that the sensitivities for detection of NOx decrease in the order Co > Cu > Pb > H > Fe > Mg > Ni, whereas the sensitivities toward NH3 decrease in the order Pb . Fe > Cu > Ni > Mg. Pc films either prepared by spin-coating or by vacuum deposition show a detection range of 0.3-200 ppm for NO2197,202 and one of 1-200 ppm for NH3202 without further annealing after film formation. Independent of the central metal ion, phthalocyanines show no sensitivity toward CO, CO2, CH4, C6H14, C7H8, SO2, H2O, and H2S.202 High stability and reproducibility are observed in the sensing behavior of all phthalocyanines, with the exception of PbPc and FePc. Both show a deformation during exposure to NH3 and NO2, respectively. Additionally, the crystal structure and size have been shown to vary with operating time.198,202 The presence of disordered phases between the PbPc particles reduces the response and recovery times during the gas interaction. It was concluded by Sadaoka et al. that the gassensing properties of PbPc can be improved by annealing the Pc films at 300 °C in order to form a homogeneous layer of triclinic crystals with a mean diameter less than 0.2 µm.198

4.2.2. Polythiophene/Poly(3-alkylthiophenes) For more than one decade, polythiophene and its derivates have been the subject of intensive experimental and theoretical work.203-205 Polythiophene and its 3-methyl derivative exhibit a poor processability due to their insolubility. The available materials have not been of the high electronic

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quality necessary to use them in electronic devices. In 1986, poly(3-alkylthiophenes) with an alkyl group of more than four carbon atoms became available.206 They are soluble in several organic solvents and even fusible, which makes them attractive for use in electronic devices ranging from fieldeffect transistors207 and light emitting diodes208 to nonlinear optoelectronic components.209 Sundberg et al. investigated the manufacture of metal/ polymer contacts.210 The comparison of the current-voltage and capacitance-voltage characteristics of melt-processed and sputtering-processed devices shows no significant difference between these structures. Both processes influence the polymer surface. The sputtering involves some relatively high-energy particles impinging on the polymer surface, whereas melt-processing can give rise to surface modifications and, hence, to a stronger enhancement of the doping concentration close to the interface than obtained for sputtering. For both devices, C-V characteristics indicate an additional constant capacitance in series with the variable space-charge capacitance of the polymer, which is assigned to an insulating oxide layer at the metal/polymer junction. In the melt-processed device, this layer is probably thicker because the aluminum is exposed to air before making the contact. Similar results were obtained when the rectifying contact was formed by pressing aluminum on the polymer surface.211 In this context, it should be taken into account that reactive metals such as Al and In might form a thin insulating interfacial oxide layer and probably react with the dopants.212 This would lead to a lower accumulation of the dopant species at the interface and a wider barrier. Lous et al. investigated Schottky contacts on highly doped organic semiconductors based on thiophene oligomer by current-voltage and capacitance-voltage measurements.213 The Schottky diodes obtain their rectifying behavior from a partly undoped p- layer at the metal/oligomer interface with respect to the higher doped p+ bulk layer. This low conductive p- layer implies a large electric field at the interface which can cause a substantial lowering of the Schottky barrier. Schottky diodes with poly(3-octylthiophene) as the semiconductor and aluminum- and indium-doped tin oxide as the rectifying and ohmic contact, respectively, show a rectification ratio up to 3-4 orders of magnitude. The shape of the J-V characteristics, the high value of the ideality factor, and the observation of series resistance indicate an interfacial layer at the aluminum/polymer interface and the effects of trapping in a region adjacent to the metal.214 The effect of oxygen is reversible on structures based on poly(3-methylthiophene).215,216 The effect of various atmospheres on the electrical characteristics of Mo/poly(3methylthiophene) Schottky barrier diodes was investigated by Tagmouti et al.216 The electrical characteristics were mainly affected by the presence of water vapor and oxygen. The capacitance shows a frequency dispersion, which is characteristic for the presence of deep traps induced by oxygen forming acceptor states in the band gap. The authors concluded that oxygen fills the traps and dopes the polymer. The effect of water vapor is manifest by reduction in dopant density and a filling of deep traps, which is irreversible. Gas-sensitive characteristics toward water vapor, chloroform, and ethanol were observed in poly(3-dodecylthiophene) Schottky diodes.217 The two types of gases have opposite effects on the diode characteristics; one type (water vapor) increases and the other type (chloroform and ethanol)

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Figure 13. Chemical structure of polyaniline, showing three different oxidation states.

decreases the reverse bias current. The results were explained by the acceptor and donor behavior, respectively, which results in a change in the Schottky barrier height by doping or undoping of impurity levels in the band gap of the p-type semiconductor.

4.2.3. Polyanilines The advantages of polyaniline compared with other conducting polymers are the easy and cheap oxidative synthesis, the thermal and environmental stability, and the simple nonredox doping in protic acids.218 Polyaniline is a p-type organic semiconductor, and it can exist in several oxidation states, ranging from fully reduced leucoemeraldine over intermediate and moderately oxidized (doped) emeraldine to the fully oxidized pernigraniline, as named by Green and Woodhead219 (Figure 13). Acid-doping of the soluble emeraldine base, using organic and inorganic acids, leads to the insoluble highly conducting emeraldine salt. Schottky diodes have been prepared using either casting220,221 or electrochemical techniques.222,223,228 The polymers investigated involve undoped and acid-doped polyanilines, which behave as p-type semiconductors. The diodes, in which aluminum was used as low work function metal to form a Schottky junction, exhibit a moderate rectifying behavior and low leakage current. The values of ideality factor n lie between 1 and 2, indicating either a recombination process in the barrier or a thin insulating layer at the interface. The acid-doping of polyaniline, e.g., with toluene-4-sulfonic acid, can cause a higher rectifying effect and photovoltaic conversion efficiency. Calderone et al. investigated the metal/polymer interface theoretically, focusing on the interaction between aluminum and the fully reduced form of polyaniline, i.e., leucoemeraldine.224 The study obtained for the model system Al/pphenylenediamine indicated that an electron charge is transferred from the metal to the organic system, and the π-electronic levels are strongly perturbed. These results are considered as the basis of the Al/leucoemeraldine interaction. Chemical sensitivities toward gases and vapors, such as ammonia,225 hydrogen cyanide,226 hydrogen,227 and methane,228 have been shown. Campos used a sandwich type device based on Al/polyaniline as a sensor for the detection of methane gas. HCl-doped polyaniline was electrochemi-

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cally deposited on an indium tin oxide coated glass electrode, which serves as the ohmic contact metal in the Schottky diode. Methane decreased the forward and reverse bias current, indicating an increase in the barrier height.

4.2.4. Polypyrrole 4.2.4.1. General Aspects. Among various conducting polymers, polypyrrole (PPy) has been extensively investigated because of properties that are attractive from the practical point of view: that is, relatively good environmental stability, high conductivity, and ease of preparation either by chemical or electrochemical methods, as well as the possibility of forming homopolymers or composites with optimal properties. A general difficulty of the reproducible PPy preparation arises from its complexity. The structure and hence the properties of the resulting PPy are strongly influenced by a number of variables, e.g., the polymerization potential, the monomer concentration, and the preparation temperature,229 which are not perfectly controllable. Therefore, the results on PPy vary widely. Two basic methods are used for the preparation of PPy: chemical and electrochemical synthesis. Anions, present in the polymerization solution, are simultaneously incorporated into the polymer as dopants. The final polymer chain bears a charge of unity for every three or four pyrrole rings, which is counterbalanced by the dopant anion.229 A lot of work has been carried out in order to improve both the chemical and electrochemical synthesis of pyrrole, as reviewed by Rodrı´guez et al.229 There exists a huge number of publications dealing with parameters influencing the final properties of PPy samples during the synthesis. Only some important parameters are briefly discussed below. The counterion is indispensable for the charge compensation of PPy; at the same time the size, shape and structure of the counterion are expected to influence the electrical conductivity. There are a large number of reports comparing the conductivity values of PPy containing aromatic sulfonate anion, such as p-toluene sulfonates, aromatic carboxylates, etc., and inorganic anions, such as BF4-, Cl-, ClO4-, SO42-, etc., and it was found that conductivity was always higher in the case of aromatic dopants.230 Few studies have dealt with PPy containing anionic transition-metal complexes, e.g., tetrasulfonated metallophthalocyanines (MPcTS).231,232 Saunders et al.231 have studied the physical and spectroscopic properties of PPy-MPcTS (M ) Co, Cu, Fe, and Ni) prepared using both aqueous and nonaqueous solution. The conductivities of these materials were found to rapidly decrease upon exposure to air. As a reason for this, it was suggested67 that the large size of the MPcTS counteranions produced packing inconsistencies within PPy-MPcTS, which facilitates the chemical attack of the polymer backbone by atmospheric oxygen. Walton et al.232 have studied a competitive doping of PPy during polymerization in an aqueous and acetonitrile solution of ClO4- and copper phthalocyanine tetrasulfonate. The effect of temperature on the polymerization process has been studied. It has been shown that the temperature has a significant effect on the conductivity of PPy depending on the combination of solvent and dopant. Toluensulfonateand perchlorate-doped PPy films, deposited at a lower temperature (280 K) in water, exhibit higher conductivities than films prepared at 293 or 313 K.229 The authors found that an increase in the temperature favors the polymer growth, which has been related to an overpotential decrease.

Potje-Kamloth

4.2.4.2. Metal/Polymer Interface. The polypyrrole family of polymers is a promising candidate for device fabrication because of its relatively high stability in ambient air. Oxidized (p-type) PPy is reported to have a work function close to 5 eV.122 From theoretical considerations, therefore, the junctions between PPy and metals with low work functions are expected to form a Schottky barrier. Several reports are available on metal/PPy junctions.182,233,235 Almost all of the results were explained on the basis of the Richardson equation, eq 6. The high charge carrier concentration found in doped PPy (NA ∼ 1022 cm-3 assuming that all dopant charges give rise to one hole polaron or half a bipolaron) results in the formation of a very thin Schottky barrier with a narrow depletion width (∼ 10 Å), which may be the reason for the bad rectifying behavior of the polymer Schottky junction due to the dominant tunneling process of the charge carrier transport mechanism (see section 3.2.4). It has been suggested that the introduction of N-substituted pyrroles to PPy lowers the conductivity of the homopolymer.234 The relationship between the density of dopants and the conductivity has led to investigations of the rectifying characteristics of various copolymers having pyrrole and N-methylpyrrole units. The introduction of N-methylpyrrole would less influence the delocalization of π-electrons along the polymer chain than that of other N-substituted pyrroles. Schottky junction devices from a pyrrole/N-methylpyrrole copolymer have been fabricated exhibiting a relatively low charge carrier concentration of NA ∼ 1017 cm-3 and a conductivity of 10-210-4 S cm-1.182,185 The values of the ideality factor (n ) 1.20-1.38) are closer to that of an ideal diode, and the rectification ratio is relatively high. In or Al have mostly been used as low work function metals to from a non-ohmic Schottky barrier with PPy.137,181-183,233 Gupta et al. reported leaky Schottky barriers of reactive metals, such as Al and In, with the polymer.233 Some authors concluded that the formation of metal salts at the metal/polymer junction as a result of reaction between the dopant anion and the low work function metal leads to insulating interphases.137 Other reports show that the ambient (i.e., oxygen) influences the current-voltage characteristics of contacts between these metals and PPy and that the observed hindrance of charge transport was due to the formation of insulating metal oxides.181,183,235,236 The thin insulating layer at the interface between the metal and the polymer was assumed to be the reason for the anomalous J-V characteristics with a small plateau at low forward bias.217,237,238 The forward characteristics of this type are comparable with nonequilibrium effects in MIS diodes.237 Taylor et al. noted that this effect could be further produced when the polymer Schottky diode comprising an aluminum rectifying contact was given a thermal post-metal annealing in air.239 On the other hand, it has been shown by Nguyen et al. that a metal with a high work function (e.g., gold) can be used to form a non-ohmic contact with PPy when metal phthalocyanine tetrasulfonates (MPcTS) are used as dopants.132 These contacts show an exponential and asymmetrical behavior of the J-V curves (Figure 14). As can be seen in Table 4 and Figure 14, the junction parameters of Au/PPy-MPcTs diodes are strongly influenced by the type of dopant. The junction between PPy-TOS and Au shows a low rectifying ratio and a relatively high saturation current density. This junction seems to be practi-

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The capacitance-bias voltage curve shown in Figure 17 indicates the formation of a Schottky junction135 and allows the use of the simple depletion layer theory as already discussed in section 3.5.3. Nguyen et al.132 determined a value of 0.26 V for the built-in voltage Vbi and a charge carrier concentration NA (concentration of polarons and/or bipolarons) of 1.2 × 1020 cm-3 for the CuPcTS-doped PPy layer of 120 nm in thickness. This is in good agreement with NA ) 1.5 × 1020 cm-3, reported by Gupta233 for electrochemically prepared PPy of 3 µm in thickness, doped with tetrafluoroborate anion. Figure 14. Dark current density vs applied voltage (J-V) characteristics of Au/PPy Schottky barrier diodes doped with different dopants.

Figure 17. Mott-Schottky plot (C-2 vs bias voltage) for a Au/ PPy-CuPcTS Schottky barrier diode. Figure 15. Dark current density vs applied voltage (J-V) characteristics of a Au/PPy-CuPcTS Schottky barrier diode.

cally quasi-ohmic. The most optimal diode parameters, such as the rectifying ratio and the ideality factor, were found for the Au/PPy-CuPcTS junction (Figure 15). The rectification ratios (forward bias current/reverse bias current at 0.6 V) were observed to be between 7.7 and 3.2. The values are slightly larger than that of the Schottky diode based on trans-polyacetylene and Pb (n ) 1.5) reported by Kanicki,133 and are in good agreement with that for Ti/PPy (n ) 10) reported by Gupta et al.240 and for Al/PPy Schottky diodes (n ) 6) reported by Bantikassegn et al.137 Impedance and capacitance measurements give insight into the nature of the diode junctions and the physical properties of PPy as an organic semiconductor.132 Figure 16 shows the plane plot of the complex impedance of a Au/PPyCuPcTS Schottky barrier diode, measured at zero bias. The device shows only a single semicircular arc. The experimental curve can be fitted well by applying the equivalent circuit of the diode shown in Figure 16. This indicates the formation of a negligible small insulating interfacial layer (see Figure 7A).

The charge carrier concentration can also be estimated from the doping level, determined by electrochemical quartz microbalance (EQCM) measurements.241 A doping level of 30.2% is deduced, and a charge carrier concentration of about 3.3 × 1021 cm-3 is expected. This value is in accordance with that reported by Buhks et al.242 (about 3 × 1021 cm-3). It can be seen that the acceptor concentration calculated from capacitance measurements is far different from that estimated from the doping level obtained from the EQCM measurements. This may be due to the large concentration of defects present in PPy that strongly influence the conjugation lengths of the polymer chain231 (see section 4.2.6.1). 4.2.4.3. Chemical Sensing Properties. Influence of Dopants. Several authors,64,121,243 have reported that the sensitivity of PPy to certain vapors or gases changes with the doping anion. Sensitivities toward organic vapors, such as alcohols, chlorinated hydrocarbons,121,244 aromatic hydrocarbons,245 nitrogen oxide gases, and organophosphorous compounds,121 have been reported. The influence of different metal phthalocyanines incorporated in PPy on the NOx sensing properties of the Au/PPy diodes has been studied.121 NOx exhibits a strong electron acceptor effect on PPy when doped with metal phthalocyanines.27,29,64 Nguyen at al.

Figure 16. Complex impedance plane plots of a Au/PPy-CuPcTS Schottky barrier diode at zero bias, and a complex plane model.

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Figure 18. Mott-Schottky plot (C-2 vs bias voltage) for a Au/ PPy-CuPcTS Schottky barrier diode measured in air and in 11 ppm NOx.

Figure 19. Change in J-V characteristics of Au/PPy-CuPcTS measured in air and in 11 ppm NOx. ∆V indicates the NOx induced voltage shift at constant current density.

Figure 20. Response curve of a Au/PPy Schottky barrier sensor toward exposure to NOx by applying a constant current density (J ) 2 × 10-3 A cm-2).

showed that the exposure of a Schottky diode with PPy as the semiconducting organic layer to NOx causes an increase in the doping level of PPy. This leads to a positive shift of its Fermi level and, therefore, increases the built-in potential at the Au/PPy junction.132 This effect could be confirmed by capacitance voltage measurements (Figure 18).

Potje-Kamloth

The change in the J-V characteristics is shown in Figure 19 for the same measurement. The gas sensitivity of the diodes is represented by the so-called gas-induced voltage shift, ∆V, at a constant current density in gas and in air, respectively, which was extracted from current voltage characteristics as discussed in section 3.5.1. Figure 20 shows the transient response of the output voltage of a Au/PPy-CuPcTS Schottky barrier diode to exposure to NOx, measured at constant current density. As already discussed in section 3.5.1, an ideal Schottky barrier diode (i.e., without an appreciable interfacial layer) can be divided into two regions: (1) the barrier region (of thickness WJ), having a capacitance CJ in parallel with the junction resistance RJ, and (2) the neutral bulk (of thickness WB), with a bulk capacitance CB in parallel with the bulk resistance RB. The NOx exposure can lead to an increase in the built-in voltage Vbi and the depletion width WJ; hence, the resistance RJ of the barrier region can increase as well. In contrast, NOx exposure may cause a decrease in the bulk resistance RB of the PPy layer because of the strong acceptor behavior of NOx. These effects have been confirmed by impedance measurements.132 Figure 21 shows the plane plot of the complex impedance of the Au/PPy-CuPcTS Schottky barrier diode, measured in air and in NOx after different time intervals. It can be clearly seen that when the diode is exposed to NOx, the bulk resistance decreases and the junction resistance simultaneously increases. A combination of those compensatory effect can decrease the gas sensitivity of the Schottky barrier diodes. In the low bias region, the junction voltage increases slightly due to NOx-exposure, while in the high bias region it increases largely. This suggests a large change in the interface layer resistance upon NOx-exposure, which has been assumed to play an important role in the charge carrier transport across the Schottky barrier.82 The change of the diode parameters extracted from current voltage measurements of PPy-MPcTS Schottky diodes exposed to NOx is listed in Table 6. For comparison, the sensitivity of a PPy-TOS Schottky diode is included. The data listed in Table 6 show that the NOx sensitivity is strongly influenced by the nature of the dopants. A lower saturation current density is accompanied by a higher rectification ratio and by a lower value of the ideality factor, which are related to a larger value of the NOx-induced voltage shift measured in the low bias range (Vbias < 0.3 V). The ideality factor shows a tendency to decrease when the diode is exposed to NOx. It has been reported by Bardeen246 that the discrepancy between the real Schottky diode and the ideal may be due

Figure 21. Plane plots of complex impedance of a Au/PPy-CuPcTS Schottky barrier diode (1) in air, (2) and (3) in 11 ppm NOx.

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Figure 23. Doping level (O) and built-in voltage (0) as a function of PPy film thickness of Au/PPy-CuPcTS Schottky barrier diodes.

Figure 22. Dependence of the diode parameter and sensor response of Au/PPy-MPcTS Schottky barrier diodes on the nature of dopants.

to the effect of surface states and that the increase or decrease in the ideality factor may be due to a generation or reduction of the interface state density. Hence, the tendency of the ideality factor to decrease when the diode is exposed to NOx may be attributed to the compensation of the interface states by NOx molecules.132 In the low bias range, the diode current is primarily controlled by the depletion region and the gas response is usually explained in terms of the change of the work function (i.e., Fermi level) in the polymer layer. Capacitance-voltage (C-V) measurements show that PPy-CuPcTS exhibits the highest charge carrier density of all layers studied and hence the most positive work function. Hence, PPy-CuPcTS leads to the highest built-in voltage of all polymer diodes studied (Figure 22A). The interaction of PPy-CuPcTS with NO2 causes the highest change in work function, which is related to the highest change in the charge carrier concentration, as C-V measurements have shown (Figure 22B). Both effects are enhancing the sensor signal of the Schottky barrier diode when compared with work function measurements using Kelvin Probe or POSFET.247 The charge carrier concentration of the PPy layer decreases from CuPcTS to TOS as dopant. The junction between PPy-TOS and Au shows a very low rectifying ratio due to the low built-in voltage and is practically quasi-ohmic. In this case, the bulk resistance cannot be neglected any more even in the low bias range. Hence, the tendency of the NO2-induced junction voltage shift to decrease when the built-in voltage of the diode decreases can be explained in terms of the enhancement of the compensatory effect in the PPy bulk. This can be seen in Figure 22C along the series CuPcTS, PbPcTS, NiPcTS, CoPcTS, and TOS. This effect is consistent with results obtained by Cabala et al.121 They observed (1) an increase of the work function

of doped PPy during exposure to NOx, which was assumed to be due to an increase of the doping level, and (2) a decrease of the sensitivity toward NOx from PPy-CuPcTS to PPy-PbPcTS and PPy-TOS. According to the model given by Janata,128 describing the chemical modulation of the work function of a sensitive polymer layer upon exposure to an interacting gas or vapor (see section 3.4.1), the magnitude and the polarity of the response depend on the value of the charge-transfer coefficient δ, with |2δ| < 1 (eq 25). The charge-transfer coefficient δ can be estimated from quartz crystal microbalance (QCM) measurements during NOx exposure. From the obtained data, an absorbed NOx gas concentration of about 1.14 × 1020 cm-3 is deduced when the PPy layer is exposed to 11 ppm NOx. Assuming that most absorbed NOx molecules participate in the formation of a charge-transfer complex, a value of 0.36 can be deduced for the charge-transfer coefficient δ, taking into account the change of the charge carrier concentration in the PPy layer extracted from C-V measurements (NA ) 4.2 × 1019 cm- 3). Evaluating the change in built-in voltage with NOx gas concentration, a value of 0.41 is deduced for the charge-transfer coefficient δ. This value is slightly larger than that estimated from QCM measurements, and it indicates that the change in the Fermi level of PPy does not follow exactly the change in the charge carrier concentration according to eq 25 and eq 26, respectively. This fact may be attributed to the pinning of the Fermi level due to the presence of interface states, as discussed in section 3.3.3 Influence of Doping LeVelsFilm Thickness. It is wellknown that the doping level of MPcTS-doped PPy layers prepared by electropolymerization is strongly influenced by the PPy layer thickness. Rosenthal et al. have shown that the doping level of a MPcTS-doped PPy layer electrochemically formed at a constant potential decreases sharply when the PPy layer thickness increases.248 The parasitic oxidation of the MPcTS dopant gives rise to a depletion layer near the electrode surface and, hence, to a decrease in the concentration of the dopant anions incorporated into the polymer backbone. The influence of the film thickness on the doping level of PPy has been monitored in situ using EQCM measurements.241,249 J-V and C-V measurements of Au/PPyCuPcTS Schottky barrier diodes comprising PPy layers of different thicknesses showed that a decrease in the doping level causes a decrease in the Fermi level of PPy. As shown in Figure 23, the built-in voltage formed between the PPy layer and Au and, in turn, the depletion width of the junctions decrease according to eq 3.

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Figure 24. Change in junction voltage ∆VJ as a function of PPy film thickness of Au/PPy-CuPcTS Schottky barrier diodes.

The two compensatory effects of NOx on the gas sensitivity of the Schottky barrier diodes in the bulk of PPy and at the interface region can be studied more extensively with the variation of the polymer film thickness, which strongly influences the doping level of the electropolymerized MPcTsdoped PPy layers.248 They can be expressed as the changes in the junction and bulk voltages, which can be determined from the J-V characteristics according to the following equations:

( )

(43)

VB ) Vbias - VJ

(44)

Figure 25. Dependence of the electrical parameter and sensor signal of a Au/PPy-CuPcTS Schottky barrier diode on the preparation temperature.

The influence of the polymer film thickness of PPyCuPcTS/Au diodes on chemical sensing properties, expressed as the change in the junction voltage ∆VJ upon exposure to NOx, is shown in Figure 24. The NOx-induced voltage shift ∆VJ, which was measured in the low bias range, decreases when the PPy film thickness increases. The increase in the bulk resistance due to the decrease of the doping level causes a stronger influence of the neutral bulk effect. A decrease in the diode current exposed to NOx is observed in the low bias voltage region, because the junction resistance RJ is dominant relative to the bulk resistance RB. In the higher forward bias region, where the flat-band condition is reached, the depletion width WJ decreases to zero, the junction resistance reaches a minimum value, and the bulk resistance of the PPy layer may become dominant. Therefore, the gas response of the current-bias behavior is dominated by the gas-induced change of the PPy layer resistance. Influence of Doping LeVelsFilm Preparation Temperature. The effects of substrate temperature during film formation, of the annealing process on the film morphology and the crystalline structure, and the influence on the gassensing properties for NO2 have been studied. Several authors have reported that, by manipulation of the film morphology, it is possible to control the sensitivity and response characteristics of the gas sensors.198,250,251 Masui et al. found that the gas sensitivity of the amorphous copper phthalocyanine (CuPc) film is higher than that of the polycrystalline CuPc film because the amorphous film has a looser stacking of CuPc molecules and, hence, more adsorption sites for NOx.251 Schottky barrier diodes based on PPy layers prepared at room temperature (295 K) do not show very large rectification ratios (Table 4). This may be due to a very thin barrier formed at Au/PPy junctions and may be an effect of the tunneling current.137 With an acceptor concentration on the

order of 1019 cm-3 and assuming the depletion width to scale as the square root of the acceptor density (eq 37), a depletion width of about 1-10 nm is expected for these devices. Hence, the tunneling current component cannot be neglected at room temperature (section 3.2.4). This is associated with the relatively large values of the ideality factor (see Table 4). In general, the rectification ratio of diodes can be improved by the enhancement of the depletion width. According to eq 3, this may be reached either by increasing the built-in potential (i.e., contact potential) or by decreasing the active charge carrier concentration. It was found that the charge carrier concentration calculated from 1/C2 versus V curves is far different from that estimated from the doping level obtained from EQCM measurements.241,249 (∼1021 cm-3), assuming that all dopant charges give rise to one hole polaron (or half a bipolaron). This may be due to the existence of a large concentration of defects in amorphous semiconductor materials. Due to these structural defects, the majority (about 80 to 90%) of the self-localized electronic defects (polarons or bipolarons) are trapped, such that they cannot contribute any more to the charge carrier transport through the polymer layer. The concentration of structural defects increases with increasing preparation temperature, because the polymerization rate rises simultaneously.252 In Figure 25 and Table 7 the electrical parameters of Au/ PPy-CuPcTS Schottky barrier diodes can be found as a function of the polymerization temperature of PPy.249 The rectification ratio increases rapidly with the increase of the preparation temperature, and simultaneously, the saturation current density decreases largely. This may be due to the enhancement of the defect concentration in PPy with the increase in the preparation temperature, which decreases the so-called active acceptor concentration and, hence, enhances the depletion width of the Schottky barrier according to eq 37.

VJ =

nkT J ln e J0

Semiconductor Junction Gas Sensors

Figure 26. Transient response of a Au/PPy-CuPcTS Schottky barrier diode electrochemically prepared at different temperatures, exposed to 11 ppm NOx in ambient air.

Therefore, the concentration of the active charge carriers is dependent on the preparation temperature and decreases with increasing preparation temperature (Figure 25A). Simultaneously, the size of the space charge region increases and therefore the junction resistance is enhanced with respect to the bulk resistance. Hence, the NO2-induced shift of junction voltage increases for the same concentration (Figure 25D), enhancing the sensitivity of the Schottky barrier diode. Furthermore, with increasing temperature, the crystallinity decreases and is zero when reaching the glass transition temperature (which is between 315 and 328 K for PPyCuPcTS64). The gas or vapor can only be absorbed by the amorphous part of the polymer, which increases with the preparation temperature. Hence, NO2 interacts more strongly with layers prepared at higher temperatures, leading to a higher relative change in the active charge carrier concentration (Figure 25B) and a higher change in the built-in voltage for the same NO2 concentration (Figure 25C). Both effects cause a further increase of the gas-induced junction voltage shift with the preparation temperature. The diffusion of gas or vapor in conducting polymers and their adsorption on polymer surfaces are the fundamental processes that control the chemical sensing response. It is supposed that diffusion processes in the crystalline polymer phase are much slower compared to those in the amorphous phase.253 Hence, the polymerization temperature is an important parameter influencing the diffusion processes of polymeric sensor transducers. Figure 26 shows the response curves of diodes with PPy films exposed to 11 ppm NO2. The diode based on the PPy film prepared at 323 K exhibits the response time and the recovery time, which are much faster than those of diodes based on PPy films prepared at 283 and 303 K. The decrease in response and recovery time can be explained by an increase in the amount of the amorphous phase compared to that of the crystalline phase in PPy when the preparation temperature increases.

5. Summary and Perspectives Schottky barrier devices are promising candidates for sensor applications. The devices are simple to fabricate, obviating the need for expensive processes used in microtechnology. Both the metal and the semiconductor forming the Schottky barrier junction can be used as the chemical sensing components interacting with the gases or vapors. The material choices and combinations define the working principle, the sensitivity, and the selectivity of the sensor devices.

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In Schottky barrier devices, particular attention must be paid to the formation of the junction barrier between the metal and the semiconductor. In the case of organic semiconductors, there is still only limited understanding of the interface between the organic material and the metal, which makes data interpretation complicated. Experimental data show that the junction barrier turns out to be not ideal because of the presence of interface states as well as the undesired formation of an interfacial layer caused either by instability of the components in air or by chemical reaction between the metals and semiconductors. The deviation from ideality can influence the sensor characteristics up to a complete breakdown of the ability to detect gases or vapors. Therefore, a high quality of the interface is the key factor affecting the sensing performances of the Schottky barrier devices. This certainly can increase the fabrication complexity and cost. The key difference between organic and inorganic Schottky barrier junction devices is the ability of many gases and vapors to penetrate through the organic semiconductor to the interface either to change the Schottky junction resistance or to interact with the bulk of the semiconductor, which causes a work function change of the material. In the case of inorganic Schottky barrier junction devices, gas permeation toward the metal/semiconductor interface works only for hydrogen or hydrogen-producing compounds and causes formation of a dipole layer. Organic semiconductors offer a viable alternative to conventional inorganic semiconductors for application in chemical sensors. They show sensitivities toward many gases or vapors, ranging from organic solvents to inorganic gases. Moreover, their porosity enables an ease of penetration of gases or vapors. The mechanical flexibility, environmental stability, and solution processability offer an enormous potential for applications within the field of microsensors. The macromolecular character and the high degree of flexibility in preparation make various physical and chemical properties realizable. The tunability of the sensing properties by the nature of the dopants as well as by the preparation procedures is an important benefit. Individual modification of each sensor is possible in only one step. This allows for the inexpensive fabrication of multisensing arrays, an aspect that makes sensors based on organic semiconductors suitable for commercialization. They are inherently compatible with solid-state integrated chemical sensors, because they can be readily incorporated into microfabricated structures. Additional benefits are the low device prices, the small dimensions, and low power consumption.

6. List of Symbols and Abbreviations ∆ δ 0 s i φb φe ∆φbi φbi µp,n µ θ τ τtr

interfacial layer thickness, cm2 fractional charge-transfer coefficient upon gas interaction permittivity of free space, A s V-1 cm-1 permittivity of the semiconductor, A s V-1 cm-1 permittivity of the interfacial layer, A s V-1 cm-1 Schottky barrier height, eV effective barrier height, eV image force lowering of the barrier height, eV contact potential or built-in potential, eV charge carrier mobility, cm2 V-1 s-1 dipole moment, A s cm-1 ratio of free to trapped charge carriers dielectric relaxation time, s electron lifetime within the depletion layer, s

396 Chemical Reviews, 2008, Vol. 108, No. 2 χ ω ξ A A** C E EC EF EWF Eg Emax

electron affinity, eV angular frequency, Hz energy difference between EC and EF, eV area or junction area, cm2 effective Richardson constant, A cm-2 K-2 capacitance or capacity per area, F cm-2 energy, eV conduction band edge, eV Fermi level, eV electron work function, eV energy band gap, eV maximum field strength at the metal/semiconductor interface, V cm-1 EV valence band edge, eV eV0 band bending at zero applied voltage ()φbi), eV e electronic charge, A s h Planck constant, J s h0 concentration of thermally generated holes in the valence band, cm-3 J0 saturation current density, A cm-2 J current density, A cm-2 k Boltzmann constant, J K-1 L layer thickness, cm m* reduced mass of electron, kg Nads density of adsorption sites, cm-3 Nssb density of interface states, cm-2 eV-1 NC effective density of states in the conduction band, cm-3 ND donor concentration, cm-3 Ntr effective density of traps, cm-3 NV effective density of state in the valence band, cm-3 n ideality factor ni intrinsic electron concentration, cm-3 pgas partial gas pressure, N m-2 S slope parameter s frequency exponent T absolute temperature, K V voltage or bias voltage at Schottky contact, V Vbi built-in voltage, V Vd diffusion voltage at zero bias voltage, V Vr reverse bias voltage, V W charge depletion zone or depletion layer width, cm index “s” semiconductor index “m” metal index “h” hole index “e” electron index “b” bulk index “j” junction Alq3 aluminum tris(8-hydroxychinoline) BEDT-TTF bis(ethylenedithio)tetrathiafulvalene BN benzonitrile CELT charging-energy-limited tunneling CHEMFET chemical sensitive field-effect transistor DCM dichloromethane EQCM electrochemical quartz crystal microbalance F4-TCNQ tetrafluorotetracyanoquinodimethane FIT fluctuation-induced tunneling HOMO highest occupied molecular orbital LUMO lowest unoccupied molecular orbital MIS metal-insulator-semiconductor MOSFET metal-oxide-semiconductor field-effect transistor MPcTS metallophthalocyanine tetrasulfonate OLED organic light emitting diode OFET organic field-effect transistor Pc phthalocyanine QCM quartz crystal microbalance THF tetrahydrofuran TPD N,N′-diphenyl-N,N′-bis(3-methylphenyl)-1,1′-biphenyl-4,4′-diamine VRH variable-range hopping

Potje-Kamloth

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Chemical Reviews, 2008, Vol. 108, No. 2 399 (243) Nagase, H.; Wakabayashi, K.; Imanka, T. Sens. Actuators, B 1993, 13-14, 596. (244) Topart, P.; Josowicz, M. J. Phys. Chem. 1992, 96, 7824. (245) Barisci, J. N.; Wallace, G. G.; Andrews, M. K.; Partridge, A. C.; Harris, P. D. Sens. Actuators, B 2002, 84, 252. (246) Bardeen, J. Phys. ReV. 1947, 71, 717. (247) Potje-Kamloth, K. Crit. ReV. Anal. Chem. 2002, 32, 121. (248) Rosenthal, M. V.; Skotheim, T. A.; Linkous, C. A. Synth. Met. 1986, 15, 219. (249) Nguyen, V. C.; Potje-Kamloth, K. J. Phys. D: Appl. Phys. 2000, 33, 1. (250) Lee, Y.-L.; Tsai, W.-C.; Maa, J.-R. Appl. Surf. Sci. 2001, 173, 352. (251) Masui, M.; Sasahara, M.; Wada, T.; Takeuchi, M. Appl. Surf. Sci. 1996, 92, 643. (252) Novak, P. Electrochim. Acta 1992, 37, 1227. (253) Hsieh, J. C.; Liu, C. J.; Ju, Y. H. Thin Solid Films 1998, 322, 98.

CR0681086

Chem. Rev. 2008, 108, 400−422

400

Optical Chemical Sensors Colette McDonagh,† Conor S. Burke,‡ and Brian D. MacCraith*,† Biomedical Diagnostics Institute, Dublin City University, Glasnevin, Dublin 9, Ireland and Optical Sensors Laboratory, National Centre for Sensor Research, Dublin City University, Glasnevin, Dublin 9, Ireland Received July 12, 2007

Contents

1. Introduction

1. Introduction 2. Sensing Platforms 2.1. Introduction 2.2. Fiber Optic Sensor Platforms 2.2.1. Passive FOCS 2.2.2. Active FOCS 2.3. Planar Waveguide-Based Sensor Platforms 2.3.1. Fluorescence-Based PWCS 2.3.2. Absorption-Based PWCS 2.3.3. Refractometric PWCS 2.3.4. Light Coupling Strategies for PWCS 2.4. Summary 3. Direct Sensors 3.1. Introduction 3.2. Direct Spectroscopic Sensing 3.2.1. Absorption-Based Sensors 3.2.2. Direct Fluorescence Sensing 3.2.3. Raman and SERS Sensing 3.3. Summary 4. Reagent-Mediated Sensors 4.1. Introduction 4.2. Reagents 4.2.1. Reagents for Colorimetric Sensing 4.2.2. Reagents for Luminescence Sensing 4.2.3. Summary 4.3. Immobilization Matrices 4.3.1. Introduction 4.3.2. Sol−Gel Matrices 4.3.3. Polymer Matrices 4.3.4. Interaction of Reagent and Support Matrix 4.4. Recent Developments in Absorption-Based Sensors 4.5. Recent Advances in Luminescence-Based Sensors 4.5.1. Introduction 4.5.2. Intensity-Based Sensing 4.5.3. Lifetime-Based Sensing 5. Key Trends and Future Perspectives 6. References

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* To whom correspondence should be addressed. Phone: +353 (0)1 7005299. E-mail: [email protected]. † Biomedical Diagnostics Institute, Dublin City University. ‡ National Centre for Sensor Research, Dublin City University.

The field of optical chemical sensors has been a growing research area over the last three decades. A wide range of books and review articles has been published by experts in the field who have highlighted the advantages of optical sensing over other transduction methods.1-5 An appropriate definition of a chemical sensor is the so-called “Cambridge definition”:1,6 Chemical sensors are miniaturised deVices that can deliVer real time and on-line information on the presence of specific compounds or ions in eVen complex samples. Figure 1 shows a schematic of a sensor system, illustrating the three main elements, the sample (or analyte), transduction platform, and signal-processing step. Optical chemical sensors employ optical transduction techniques to yield analyte information. The most widely used techniques employed in optical chemical sensors are optical absorption and luminescence, but sensors based on other spectroscopies as well as on optical parameters, such as refractive index and reflectivity, have also been developed. This review will deal mainly with the transduction stage in Figure 1. There will be only limited discussion of sample preparation/enrichment and signal-processing techniques. Recent developments in the field have been driven by such factors as the availability of low-cost, miniature optoelectronic light sources and detectors, the need for multianalyte array-based sensors particularly in the area of biosensing, advances in microfluidics and imaging technology, and the trend toward sensor networks. In the case of luminescencebased sensors, direct intensity detection has been replaced in many applications by lifetime-based sensing, often using sophisticated phase-based techniques. This review will focus on developments in optical chemical sensing over the last 10 years with major emphasis placed on the literature from 2000 to the present day. Recent novel developments will be highlighted, and future trends will be discussed. While the optical principles used in chemical sensing have not changed substantially over the years, in many cases the transduction platforms have changed considerably, yielding sensors with vastly improved performance, the most relevant performance parameters being sensitivity, stability, selectivity, and robustness. The structure adopted for this review is shown in Figure 2. Because of the important role played by the sensor platform in current sensor developments, this review begins with a section on different platforms and platform technologies. This section includes a description of waveguide sensor technology, including fiber-based and planar waveguide systems. Refractometric platforms are also dealt with here as are interferometric sensors.

10.1021/cr068102g CCC: $71.00 © 2008 American Chemical Society Published on Web 01/30/2008

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Colette McDonagh studied undergraduate physics at the National University of Ireland in Galway and was awarded her Ph.D. degree in Physics from Trinity College, Dublin, in 1980. After postdoctoral work at Trinity College and the Department of Applied Science at the University of California, Davis, she took up an academic position in the School of Physical Sciences at Dublin City University. She currently holds the position of Associate Professor, and her research interests include development of sol−gelbased optical sensors for environmental monitoring, luminescence-based optical biosensors, and development of strategies for luminescence enhancement in biochips including metal-enhanced luminescence and highbrightness nanoparticles.

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Brian MacCraith is Director of the Biomedical Diagnostics Institute (BDI) at Dublin City University. The BDI is a Science Foundation Ireland (SFI) Centre for Science, Engineering & Technology (CSET) focused on developing the underpinning science leading to next-generation biomedical diagnostics. Established in October 2005, the BDI is an academic−industry partnership involving six industrial and four academic partners and has been funded for 5 years in the first instance. The funding awarded to the BDI includes over 6 million euros from its industry partners and 16.5 million euros from SFI. He was founding Director of the National Centre for Sensor Research at Dublin City University and held this position from its establishment in October 1999 until the establishment of the BDI. Currently, the NCSR comprises over 200 full-time researchers working on the fundamental science and applications of chemical sensors and biosensors. With a strong track record and international reputation in the field of optical chemical sensors and biosensors, he has published widely (over 150 publications) on these topics as well as developing significant Intellectual Property.

Conor S. Burke received his B.Sc. (Hons.) degree in Physics with French from Dublin City University in 1999. In the same year, he started his Ph.D. studies with the Optical Sensors Laboratory in the School of Physical Sciences at Dublin City University under the supervision of Professor Brian MacCraith. He was awarded his Ph.D. degree in 2004 with a thesis entitled “Development of microfabricated optical chemical sensor platforms using polymer processing technology”. His work involved development of both colorimetric and fluorometric chemical sensors based on novel, enhanced optical platforms using polymer microprocessing technology. Since then, he has worked as a postdoctoral researcher at the National Centre for Sensor Research, Dublin City University, where his primary research interests include development of novel optical sensors for breath monitoring and environmental applications.

In general, optical chemical sensors may be categorized under the headings of direct sensors and reagent-mediated sensors. In a direct optical sensor, the analyte is detected directly via some intrinsic optical property such as, for example, absorption or luminescence. In reagent-mediated sensing systems, a change in the optical response of an intermediate agent, usually an analyte-sensitive dye molecule, is used to monitor analyte concentration. This latter technique is useful particularly in the case where the analyte has no convenient intrinsic optical property, which is the case for many analytes. Section 3 of this review deals with direct optical sensing using a range of optical parameters. This section deals exclusively with direct spectroscopic sensing,

Figure 1. Principal stages in the operation of a sensor.

including infrared (IR) and ultraviolet (UV) absorption techniques, direct fluorescence measurements, as well as Raman and surface-enhanced Raman spectroscopy (SERS). The section on infrared absorption includes a brief discussion of sensing based on the relatively newly developed quantum cascade lasers (QCL). Direct refractometric sensing is

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Figure 2. Structure of review.

discussed in section 2. Section 4 on reagent-mediated sensors includes a discussion of analyte-sensitive reagents and associated immobilization approaches, especially those based on polymer and sol-gel matrices, and is followed by a comprehensive overview of recent developments in absorption- and luminescence-based sensors. This review does not contain formal sections on optical biosensors or surfaceplasmon-resonance (SPR) based sensors as these topics are dealt with in separate review articles in this issue. However, throughout this review, examples of biosensor applications are given where appropriate, which is a reflection of the current predominance of optical biosensors in sensor research and the overlap between chemical sensing and biosensing.

2. Sensing Platforms 2.1. Introduction The successful development of an optical chemical sensor is intrinsically linked to the nature of the physical platform on which it is based. When careful consideration of the transduction mechanism dictates the design of the sensor platform, this leads to development of highly efficient integrated optical sensors that combine light delivery/ collection with intrinsic sensing functionality. This section will present an overview of optical chemical sensor platforms with an emphasis on innovations in platform design during the past decade. In order to lend clarity to the discussion, the various types of sensor platform are grouped according to the underlying waveguide geometry on which the sensor is based, e.g., optical fiber based or planar, with further differentiation achieved according to the transduction mechanism employed by the sensor.

2.2. Fiber Optic Sensor Platforms The optical fiber is arguably the most exploited platform in the development of optical chemical sensors, a fact that is reinforced by the sheer volume of publications reporting such systems (over 2500 in the past 10 years). Indeed, this area of sensor research alone merits the publication of regular reviews of the field.5 Given the scope of fiber optic chemical sensor (FOCS) development, it is difficult for any review of the field to be completely exhaustive. This section will cover the principal fiber optic configurations and sensing schemes that are currently driving sensor research in this area. The scope of this discussion will be limited to sensing strategies

that are not the subject of other reviews in this issue. With this in mind, fiber optic sensors that utilize surface plasmon resonance (SPR) techniques and quantum-cascade laser (QCL) spectroscopy or those used to produce high-density sensor arrays will not be described in any great detail, except where such sensors are illustrative of a more general sensing configuration or strategy. While there exists a multitude of FOCS, the physical configurations employed are relatively few in number, with most sensors falling under one of the fiber categories illustrated in Figure 3, e.g., standard (unmodified) fiber, declad, active cladding, fiber bundle, bifurcated fiber bundle, U-bend, or tip based. In some cases, these configurations are combined in a single sensor, i.e., a standard fiber with a modified tip or a U-bend fiber with an active cladding. A significant fraction of the sensors reviewed here utilize one or more of these configurations/modifications. In this section, FOCS are classified according to the role played by the fiber in the operation of the sensor, i.e., passive or active. The fiber’s role is considered to be passive if the sensor response is not linked in any way to an intrinsic change in the optical properties of the fiber, which acts merely to transport the optical signal to and from the sensing environment. An active FOCS utilizes a fiber that has been modified so as to impart intrinsic analyte sensitivity to the fiber, for example, by doping the fiber cladding with an analyte-sensitive indicator. In this manner, the optical properties of the fiber are in some way modulated by the presence of the analyte. Passive and active systems are treated separately in the following subsections.

2.2.1. Passive FOCS A wide variety of passive FOCS have been developed over the past two decades, with perhaps the most prevalent example being that of the fiber-coupled spectrometer. However, only those systems developed within the past decade will be reviewed here. A passive FOCS commonly takes the form of a separate sensor element that is interrogated in some fashion by the fiber optic assembly. Several configurations are typically used to achieve this. The reflectance-based configuration is common in the development of colorimetric or absorption-based FOCS. A typical embodiment of such a sensor comprises an analytesensitive material (typically deposited on a planar support) that changes color upon interaction with the analyte of

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Figure 3. Commonly used configurations for FOCS: (a) unmodified; (b) declad; (c) active or doped cladding; (d) fiber bundle; (e) bifurcated fiber bundle; (f) U-bend (shown declad); (g-i) tip based ((g) tip with active cladding, (h) etched tip, (i) modified end-face).

interest and a fiber bundle (as shown in Figure 3(d)) that is used to deliver light to this material and collect the reflected light. In normal operation, the light is delivered to the analyte-sensitive component via the central fiber in the bundle and the reflected light is collected by the outer fibers, which are arranged around the central fiber. Changes in the color (i.e., the absorption coefficient at a particular wavelength) of the sensor material are detected as changes in the intensity of the reflected light. Such sensor configurations benefit from the commercial availability of fiber bundles (as standard reflectance probes) in addition to the relatively simple optical setup. A reflectance-based system has been used in the development of a sensor for amine vapors based on silica microspheres coated with the pH indicator Bromocresol green.7 A bifurcated fiber bundle (see Figure 3(e)) has been used in a similar fashion for detection of sulfur dioxide by an organopalladium complex encapsulated in a PVC membrane.8 However, the use of a reflectance-based configuration is not limited to detection of analytes by colorimetric means. The same fiber bundles can equally be applied to the excitation and detection of fluorescence. In the case of the fiber bundle depicted in Figure 3(d), the central fiber delivers the excitation light to the external sensor element and the outer fibers are used to collect and transport the emitted fluorescence to a suitable detector. This configuration has been applied to detection of dissolved oxygen.9 The bifurcated fiber bundle can be used to deliver excitation light with one fiber and collect the excited fluorescence with the second. This configuration has been used extensively for the fluorescence-based detection of a

variety of analytes including oxygen,10 iodine,11 2,6-dinitrophenol,12 pH,13 and chloride.14 Transmission-based configurations are also commonly employed in the development of FOCS, but in recent years, research has moved away from the conventional configuration employed in many spectrometers, which sees two fibers aligned on opposing sides of a solid sensor element or optical cuvette. While current work still involves the use of fibers that are aligned in such a manner, the sensor element that is interrogated has seen some changes. A subnanoliter spectroscopic gas sensor has been developed. This consists of input and output fibers that are inserted into a fused silica capillary, which acts as a transmission cell for gaseous samples (in this case acetylene) that can be detected using an infrared source.15 A similar configuration was employed by Eom et al. in the development of a dual-wavelength measurement system for refractive index and turbidity compensation in liquid flow systems.16 In that case, a length of Teflon tubing played the role of transmission cell and the input/output fibers were used to transport visible light signals from two LED sources, one blue (λpeak ) 468 nm) the other red (λpeak ) 621 nm), to a photodiode detector. In this case, the red LED was used to provide a reference signal. Such configurations are not limited to cylindrical sample cell geometries. A variety of planar (passive) fiber-coupled sensor platforms have arisen in recent years due to the emergence of microfluidic systems as technology platforms and, in particular, the drive to develop opto-fluidic sensor chips that provide integrated, on-chip optical coupling and detection. One example of this recent trend is a PDMS-based microflow cytometer that incorporates grooves for several

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optical fibers positioned at different angles on either side of a flow channel that can be used to excite and detect fluorescence from particles within the channel in addition to scattered light detection.17 A second PDMS-based device has been developed that utilizes fiber alignment grooves with integrated 2D lenses on either side of a microfluidic channel for fluorescence detection.18 While these two examples exploit fluorescence detection, the optical configurations used are also inherently suited to transmission-based sensing strategies. Many opto-fluidic microsystems make use of fiber coupling to fulfill either the role of light delivery or light collection but not necessarily both. Yang et al. reported a LED-based capillary electrophoresis system that employs fiber optic detection of the LED-induced fluorescence from within a fused silica capillary.19 A PDMS microchip for the fluorescence-based detection of human serum albumin has also been described.20 This system employs an organic LED (OLED) as an excitation source and uses an integrated optical fiber to transport the fluorescence signal to a spectrometer. Chabinyc et al. described development of an integrated, PDMS-based microfluidic system for fluorescence detection that uses a fiber-coupled LED to excite fluorescence from within the microchannels of the system while employing a microavalanche photodiode, encapsulated within the chip, as a photodetector.21 These examples serve to highlight the importance of optical fibers in rendering these increasingly important platforms optically addressable at a localized level.

2.2.2. Active FOCS In this section, active FOCS are further categorized under one of three headings, which reflect the transduction mechanism employed: (1) fluorescence based, (2) absorption based (colorimetric and spectroscopic), and (3) refractometric. (1) Fluorescence-Based Active FOCS. In order to transform a standard optical fiber into an intrinsically fluorescence-based optical chemical sensor, it is necessary to impart analyte-sensitive fluorescence to the fiber in some manner. This is done through addition of fluorescent indicator molecules to the fiber platform. This may be achieved by replacing the cladding of the fiber over a portion of its length with a solid matrix that contains the fluorescent compound, a process that involves removing a portion of the original cladding of the fiber and coating the declad region with a liquid sensor material, which is subsequently cured to form a solid, fluorescent cladding (Figure 3c). A slight variation of this involves coating the declad, distal tip of the fiber with the sensor material (Figure 3g). In some instances, the distal tip may be chemically etched and the fluorescent material deposited within the etched cavity (Figure 3h), while yet another variation involves doping the fiber with the fluorescent material during the fiber fabrication process. However, the latter approach is rarely employed due to its adverse effect on the guiding properties of the fiber. A variation of the fluorescent cladding configuration was employed by Ahmad et al., who described a system based on two “positive” fibers attached to either end of a “negative” fiber.22 The system actually consisted of a single fiber, which was declad for a portion of its length and then coated along that portion with a solution of rhodamine 6G in glycerol. The higher refractive index of the glycerol relative to the silica core of the fiber caused any light guided within the fiber to leak out, making this the negative fiber component of the system, while the clad portions on either side of this

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section constituted the positive (i.e., guiding) fibers. The excitation light that leaked out from the negative fiber induced fluorescence from the rhodamine 6G in the glycerol film, which was coupled into the output positive fiber for detection. Park et al. described development of fiber optic sensors for detection of inter- and intracellular dissolved oxygen (DO) that utilize “pulled tip” fibers.23 These are multimode fibers, the tips of which have been pulled down to submicrometer dimensions and then dip coated with a liquid PVC matrix incorporating two luminescent dyes, one that is oxygen sensitive and the other acting as a reference. A similar configuration was employed by Preejith et al. in the realization of a tapered fluorescent fiber optic evanescent wave-based sensor for serum protein.24 In that work, a fiber was declad over a length of 5 cm from its end and the declad region was then contoured to produce a tapered fiber tip. The tip was then coated with a sol-gel layer in which the fluorescent complex NanoOrange was encapsulated. Evanescent wave excitation of fluorescence was employed for detection of serum albumin. Both this system and the previous one described by Park et al. differ from that produced by Ahmad et al. in that the latter system does not employ evanescent wave interrogation of the fluorescent cladding, but all three systems demonstrate the use of such a cladding or sensing layer in the production of an active FOCS. Another example of an active, fluorescence-based FOCS is the platform of Walt et al., which centers on development of high-density fiber bundles for fluorescence-based sensing.25,26 While this is a high-density array-based platform, it also demonstrates the use of chemical etching to transform the distal tip of a fiber into a well, which can be filled with a fluorescent material. In this case, the fluorescent probe takes the form of suitably functionalized microspheres that are distributed into individual microwells (which correspond to the etched distal tips of individual fibers in a high-density fiber bundle) and can bind specific, fluorescently labeled biomolecules for fluorescence detection in a highly parallel format. (2) Absorption-Based Active FOCS. Absorption-based optical sensors can be colorimetric or spectroscopic in nature. Colorimetric sensors, as the name would suggest, are based upon detection of an analyte-induced color change in the sensor material, while spectroscopic absorption-based sensors rely on detection of the analyte by probing its intrinsic molecular absorption. FOCS that utilize these techniques have been modified in order to facilitate the interaction of the light guided within the fiber with either the colorchanging indicator or the analyte itself, depending on the nature of the sensor. Both colorimetric and spectroscopic platforms will be discussed in this section, but it will become clear that some of the fiber configurations described for fluorescence-based sensors are equally useful for development of both colorimetric and spectroscopic sensors. Indeed, the modified cladding configuration described above has been used extensively in a number of modes for development of absorption-based sensors, and it should be stressed that only some of the more recent representative examples of such work are reported here. Recent examples include development of sensors for ammonia (concentration range 0.03-1%),27 pH,28,29 and ethanol.30 In some cases the sensor consisted of a dye-doped

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section of fiber cladding that was interrogated by the evanescent field of the guided light,27 while in other cases a U-bend configuration was also used in conjunction with the dye-doped cladding in order to increase the interaction of the evanescent field with the analyte-sensitive reagent, thereby enhancing device sensitivity.28 The U-bend configuration was also used in conjunction with a declad section of fiber for spectroscopic monitoring of ethanol.30 In this case, a dye-doped cladding was not required as the sensor was being used to probe the intrinsic absorption of ethanol molecules in solution and it was necessary merely to remove the cladding from the section of the fiber forming the U-bend in order to yield a sensor element. This work actually described a distributed sensor network and utilized optical time domain reflectometry (OTDR) to interrogate a series of U-bend probes along a 1 km length of fiber. In a recent review of fiber optic sensors, James and Tatam described the development of a Langmuir-Blodgett (LB) film-based sensor for pH.29 This was fabricated by deposition of a LB film onto a side-polished optical fiber and is another example of a modified cladding-type sensor. With this platform, it was possible to monitor pH-induced changes in the absorption of the LB film at 725 nm, thereby providing the basis for the pH sensor. Refractometric FOCS. A range of refractometric FOCS has been developed in the past decade that involves addition of refractive index-sensitive optical structures to the optical fiber. Such structures include fiber Bragg gratings (FBG’s), Fabry-Perot cavities, and metal films for SPR measurements. FBG’s are written into photosensitive optical fibers (typically germanosilicate fibers) using a suitable laser light source. An interference pattern is produced by passing the laser output through a phase mask or splitting and recombining the laser beam. The fiber is exposed to this pattern, which causes a local modulation of the refractive index of the fiber that matches the pitch of the interference pattern. This refractive index modulation is known as a FBG, and its sensing capabilities derive from the dependence of the Bragg wavelength (and, therefore, the transmission of the grating) on the period of the grating and the effective refractive index, Neff, of the medium. FOCS based on FBG’s detect changes in Neff and require that the fiber section bearing the FBG be declad in order to impart sensitivity toward the refractive index of the environment in which the fiber is placed. A twin FBG-based sensor, developed by Sang et al., made use of the second grating in order to provide temperature compensation for the sensor.31 The device was used to detect different concentrations of sugar and propylene glycol solutions. Long period fiber gratings (LPFG’s) are similar structures that are fabricated in the same manner as FBG’s but have periods in the range 100-1000 µm. These structures are also refractive index sensitive and have been used for detection of analytes such as sodium chloride,32 ethylene glycol,32 and antibodies.33 Analyte specificity can be imparted to such sensors through deposition of an overlay29 that possesses its own refractive index response to particular analytes. A distal tip-based sensor for dissolved ammonia was recently reported by Pisco et al.34 The sensing element consisted of a tin dioxide (SnO2) thin film that was deposited onto the distal tip of a fiber by electrostatic spray pyrolysis (ESP). This is an example of a Fabry-Perot interferometric (FPI) FOCS, where changes in the thickness and refractive

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index of the SnO2 film at the tip of the fiber and/or changes in the refractive index of the sensing environment cause changes in the reflectivity of the fiber film interface. A zeolite thin-film-based FPI sensor was similarly employed by Liu et al. for detection of organic solvents in water.35 The optical configuration employed was similar to that of the ammonia sensor mentioned above with the zeolite thin film replacing the SnO2 layer. An optical humidity sensor for breathmonitoring applications that was also based on a fiber tip FPI configuration was described by Kang et al.36 In that case, the Fabry-Perot cavity was fabricated through layer-by-layer self-assembly of polyelectrolytes. Other examples of refractometric FOCS include a sensor for methane developed by Benounis et al., which exploited a cryptophane-doped section of fiber cladding coupled with evanescent wave interrogation.37 The specific absorption of methane by the cryptophane molecules caused a change in the cladding refractive index, which was indicated by a change in the transmitted optical power of the system. Jung et al. reported a refractive index sensor based on a threesegment optical fiber platform that consisted of a coreless silica fiber (CSF), sandwiched between two standard multimode fibers.38 The core modes from the first multimode fiber couple to cladding modes in the CSF, which then couple to core modes in the second multimode fiber. The coupling conditions are dependent on the refractive index difference between the CSF and the surrounding medium, thereby providing the system with refractive index detection capabilities.

2.3. Planar Waveguide-Based Sensor Platforms Compared with FOCS, planar waveguide chemical sensors (PWCS) are a relatively recent innovation in the field of optical chemical sensors. While most FOCS developed to date have adopted one of a small number of relatively conventional configurations (as demonstrated in the previous section), PWCS platforms research is exhibiting rapid growth in terms of innovative design and integration of multiple functionalities onto a single sensor chip. This is largely due to the compatibility of the planar geometry with a range of advanced microfabrication technologies and the ease with which such a sensor geometry can be integrated with microfluidic, lab-on-a-chip systems. Yet another advantage of such a configuration is its robust nature, which, when compared with optical fiber-based systems, is an attractive characteristic when contemplating development of practical sensor devices intended for deployment outside a laboratory environment. These attributes have made planar waveguides an ideal platform for development of integrated optical sensors. As for the previous section on FOCS, PWCS are categorized here according to the three principal transduction mechanisms employed in this field, i.e., fluorescence, absorption, and refractometry. Before examining each of these areas, it is useful to define the general configuration of a PWCS, the concept of which is illustrated in Figure 4. At its core, a PWCS comprises a planar substrate (e.g., glass, plastic, or silicon) that forms the basis of the sensor chip. In some cases, this substrate acts as the waveguide, while in others an additional waveguide layer is deposited onto the substrate. With regard to the waveguide layer, several configurations have been employed that impart various optical functionalities to the sensor platform, and some of these are reported in the following sections. In many cases,

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Figure 4. Generalized configuration of a PWCS. In many cases, evanescent wave interactions with the analyte form the basis for the sensor. In some cases, a sensing layer may be employed to facilitate transduction by imparting colorimetric or fluorometric properties to the sensor.

the light that propagates within the waveguide facilitates the operation of the platform as a sensor through the interaction of its evanescent field with the sensing environment above the waveguide. However, this is not always the case, and examples of sensors employing interrogation principles other than evanescent field-based techniques are also given in the sections to follow.

2.3.1. Fluorescence-Based PWCS PWCS based on detection of fluorescence have been developed for a broad range of analytes including oxygen,39-42 pH,41 carbon dioxide,43 and a variety of biological species.44-46 Many of these systems employ the technique of evanescent wave excitation of fluorescence (also referred to as total internal reflection fluorescence). The sensor configuration is relatively straightforward, consisting of a planar waveguide to which light is coupled using one of a variety of techniques, such as prism coupling, grating coupling, or end-fire coupling. The evanescent field of the guided light extends into the superstrate (i.e., the sensing environment) and induces fluorescence in susceptible molecules located within a distance of typically 100-200 nm from the waveguide surface. Such a technique is particularly well suited to optical biosensing applications due to the high degree of surface selectivity afforded through the use of the evanescent field for excitation purposes. Chronis and Lee reported development of a total internal reflection-based biochip platform that excites fluorescence from within a microfluidic network using a single reflection,47 while the platforms described by Rowe-Taitt et al.45 and Duveneck et al.46 (the Zeptosens platform) are examples of multiple reflection (i.e., waveguiding based) evanescent wave excitation systems, which are well suited to development of array-type optical sensor chips. In some cases, the excited fluorescence is detected either above or below the sensor platform, but it is also possible to use the waveguide itself to capture fluorescence for detection at the output end face of the waveguide.39,42,48 With such a configuration, it is possible to utilize direct excitation of the fluorescence as opposed to evanescent wave excitation, which results in higher signal levels and improved sensor performance, while improving the inherent ability of the platform to discriminate between the excitation light and the fluorescence.39,48 Zourob et al. reported an alternative configuration, based on a metal-clad leaky waveguide platform, that also affords enhanced interrogation of the sensing medium, and this system can equally be applied to absorption-based or refractometric sensing schemes49(Figure 5a,b).

2.3.2. Absorption-Based PWCS One of the most prevalent examples of absorption-based PWCS platforms is that based on evanescent-wave absorp-

Figure 5. (A) Mode profile of a metal-clad leaky waveguide covered by water at a wavelength of 620 nm using a 1 mm glass substrate. (Reprinted with permission from ref 49. Copyright 2003 Elsevier Science B.V.) (B) Instrumental set up for optical sensor based on a metal-clad leaky waveguide using a (a) 473 nm solidstate laser or 610 nm LED, (b) collimating lens, (c) filter, (d) cylindrical lens, (e) polarizer, (f) grating, (g) sensor chip, (h) linear CCD detector or photomultiplier detector. (Reprinted with permission from ref 49. Copyright 2003 Elsevier Science B.V.)

tion. The configuration is similar to that described for evanescent-wave excitation-based fluorescence sensors. In this case, the sensing functionality comes about due to changes in the absorption coefficient of the sensing environment. Such changes result in more or less absorption of the evanescent field intensity by the environment, which is reflected as a change in the detected output intensity of the sensor. Sensing layers that are doped with colorimetric, analyte-sensitive indicators can be deposited onto the upper surface of the waveguide, and any analyte-induced color changes can be probed by the evanescent field of a suitable light source (i.e., one that is spectrally matched to the indicator used). In cases where the analyte is detected using direct spectroscopic or refractometric techniques, such a sensing layer may not be necessary, although transparent enrichment layers with high permeability coefficients for the analyte(s) of interest are sometimes employed. PWCS that utilize thin-film-based sensing layers have been employed for detection of analytes such as gaseous ammonia,50,51 pH,52 iodine,53 and water vapor.54 However, in recent years, a number of platforms have been developed that demonstrate a shift away from what might be seen as conventional evanescent wave absorption-based platforms. These platforms typically incorporate design features intended to enhance the interaction of the interrogating light with the sensing environment, thereby improving platform sensitivity. The integrated waveguide absorbance optode (IWAO) developed by Puyol et al. incorporates a PVC-based sensor membrane that is located between two

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Figure 7. Schematic of a Mach-Zehnder interferometer.

Figure 6. Integrated waveguide absorbance optode (IWAO) configuration developed by Puyol et al. (Reprinted with permission from ref 55. Copyright 1999 American Chemical Society.)

antiresonant reflecting optical waveguides (ARROWs)55 (see Figure 6). The sensor membrane also serves as a waveguide and transports the interrogating radiation from the input ARROW to the output waveguide. The increased optical path length within the sensing layer provides enhanced sensitivity, and the device was applied to detection of potassium chloride solutions of varying strength. This configuration was later improved upon to incorporate curved ARROW waveguides and achieved improved sensitivity and response time.56 Hisamoto et al. reported development of ion-selective optodes57 that employ a similar strategy in that the sensing layer also serves as the waveguide (referred to as an “active waveguide”). This system is based primarily on a prism-coupled glass platform, which incorporates the PVC-based sensor membrane and has been applied to detection of a variety of analytes including potassium, sodium, and calcium. Qi et al. reported a composite optical waveguide (COWG) platform that consists of a potassium-ion-exchanged (PIE) waveguide onto which a tapered film of bromothymol blue (BTB)sa colorimetric pH indicatorsis deposited.58 The BTB film acts as a singlemode waveguide, and the sensor achieved a limit of detection (L.O.D.) of 1 ppb ammonia. Other systems that exploit enhanced interrogation of the sensing medium include the previously mentioned metal-clad leaky waveguide platform developed by Zourob et al.,49 while Burke et al.59 reported an injection-moulded waveguide platform that facilitated enhanced interrogation of a thin absorbing sensing layer through correct choice of the incident angle of interrogation, which was dictated by the design of integrated refractive optical elements.

2.3.3. Refractometric PWCS PWCS designed to measure refractive index changes represent one of the most extensively developed types of planar waveguide platforms. These systems exploit techniques such as interferometry, surface plasmon resonance, and light-coupling strategies to transduce refractive index changes. Comprehensive reviews of integrated optical chemical sensors based on such devices have recently been published by Lambeck60 and Gauglitz.61 A representative cross-section of cutting-edge sensor technology in this field will be delivered here, but readers should refer to these reviews for further reading on this sector of sensor technology. Sensors based on surface plasmon resonance (SPR) are perhaps the most widely known examples of refractometric

optical sensor platforms. However, this family of sensors will not be discussed here as it is the subject of a separate review in this issue. The first branch of refractometric PWCS to be discussed here is that of interferometric sensor platforms. A significant number of refractometric PWCS are based on exploitation of interferometric techniques, and one of the most commonly employed optical configurations is the Mach-Zehnder interferometer (MZI), shown in Figure 7. The MZI consists of a single-input waveguide whose transported optical power is split equally between two parallel waveguide branches using a Y-splitter. A second Y-splitter is used to recombine the optical signals from both branches into a single-output waveguide, at the end of which the output power is measured using a suitable photodetector. For the purposes of chemical sensing applications, one of the waveguide branches is either coated with a sensor membrane or exposed to the sensing environment while the other unmodified branch serves as a reference waveguide. Changes in the refractive index of the sensing layer/environment influence the effective refractive index, Neff, in the sensing channel, which induces a phase shift in the optical signal that propagates through this channel. Upon recombination, interference of the two optical signals (i.e., sensor and reference) occurs and the measured output power changes depending on the phase shift between these two signals. In the recent review by Lambeck, a comprehensive treatment of MZI interferometric platforms has been given, including a discussion of a range of strategies intended to enhance sensitivity and general sensor performance. This involves optimizing parameters such as operating wavelength, interaction length, and, most importantly, the waveguide composition. The MZI configuration has recently been exploited by a number of groups for detection of a variety of chemical and biochemical species.62-66 Improved performance can be achieved by implementing serrodyne modulation, which involves inserting electro-optical modulators into each branch.67 Other interferometry-based PWCS include the Young interferometer,68-70 the Michelson interferometer,71 and the difference interferometer (also referred to as a polarimeter).72-74 These platforms can be described as derivatives of the MZI and operate in an analogous fashion (see ref 53 for further details). Some noteworthy examples are the Young-type interferometer that has been commercialized by Farfield Sensors [www.farfield-scientific.com] and the Zeeman interferometer,75 which is a variant of the difference interferometer that is capable of resolving refractive index changes on the order of 10-8. In addition, Kribich et al. describe development of a multimode interference (MMI) coupler with tunable sensitivity that has been used to detect changes in relative humidity.76 This platform consists of single-mode input and output waveguides that are coupled to a central

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Figure 8. RifS configurations and measurement principle. (Reprinted with permission from ref 61. Copyright 2005 Springer-Verlag.)

multimode waveguide. The field profile within the multimode region is dependent on the refractive index of the superstrate, and changes in this are detected as changes in the output light intensity. Lavers et al. described a related coupled-waveguide-type refractometer, which consists of a buried, single-mode potassium ion-exchanged (PIE) waveguide in a BK7 glass substrate, onto which a multimode polymer waveguide is deposited by photolithographic means.51 While this device also exploits the effects of coupling between dissimilar waveguides, it is more properly an example of a resonanttype refractometric PWCS, other examples of which include the resonant mirror platform,77 metal-clad leaky waveguides,49,78 and some grating-based devices.79-81 Other grating-based devices have been developed that rely on the refractive-index- and thickness-dependent nature of diffraction to yield a sensor response that manifests itself as a change in the intensity or spatial distribution of the diffraction pattern.82-86 Platforms based on this phenomenon are under development by Axela Biosensors [www.axelabiosensors. com], although in that particular case the sensor response is primarily due to changes in the thickness of the receptor layer that defines the diffraction grating as opposed to alterations in refractive index, and the platform is included here in order to illustrate the perceived utility and commercial potential of such a sensor configuration. Several other noteworthy PWCS platforms have been developed that exploit thickness-sensitive measurement principles, as opposed to refractometry. Reflectometric interference spectroscopy (RIfS),87 developed by Gauglitz, is one such technique. The principle of RIfS is illustrated in Figure 8 and, in brief, centers on detection of changes in the interference pattern generated by light that is reflected from the interfaces of a thin film deposited on a substrate. Changes in the thickness of the thin film (e.g., due to swelling upon interaction with the analyte of interest) result in a shift in the observed interference pattern, and the technique is capable of resolving a thickness change of 1 pm. Nikitin et al.88 exploited the interference pattern generated by light reflected from the upper and lower surfaces of a

glass cover slip in the development of the Picoscope. This system measures the correlation signal of two interferometerss the cover slip and a scanned Fabry-Perot interferometers to monitor changes in the thickness of the cover slip (due, for example, to biomolecular binding events) with a reported resolution of 3 pm. Lambeck et al. proposed three theoretical strip waveguide configurations implemented in SiON technologyssegmented, absorptive, and directional coupler basedsthat are designed to transduce thickness changes through changes in the field profile of the waveguide modes with a calculated resolution of 3 × 10-5 nm.89

2.3.4. Light Coupling Strategies for PWCS An important design consideration in the development of any PWCS is the method by which light is coupled to the sensor platform. The most commonly employed techniques are prism coupling,49,57,58,77 grating coupling,45,46,50,69,70,78 and end-fire coupling,42,55,62,63,65,66,68,72,74-76 with the choice of technique typically involving a tradeoff between fabrication costs, practicality, and sensor performance. The use of prism couplers provides relatively high coupling efficiency but detracts from the cost effectiveness, planarity, and overall robustness of the sensor platform, while end-fire coupling requires extremely precise alignment optics in order to be effective, which is not an attractive feature in a field deployable sensor but is an obstacle that can be overcome by employing fiber-pigtailing strategies (although this makes easy exchange of the sensor element problematic). The use of grating couplers can be seen as a viable alternative as this is a low-cost option that preserves the robustness of the sensor, but the reduced coupling efficiency may adversely affect sensor performance. Some efforts have been made to integrate the coupling functionality into a single chip using polymer microprocessing techniques such as injection molding59 or hot embossing.41 These techniques are compatible with the production of integrated grating couplers and refractive/prismatic optical elements that provide adequate coupling efficiency, do not reduce the robustness of the sensor, and also emphasize the trend toward integration of multiple functionalities on a single sensor platform.

Optical Chemical Sensors

2.4. Summary A range of recently developed optical chemical sensor platforms have been categorized and discussed according to the underpinning waveguide geometry employed. Further categorization was achieved according to the transduction mechanism upon which the sensors were based, and development of a number of highly innovative sensor platforms has been highlighted. This is particularly true of research in the field of planar waveguide-based devices where this robust, microfabrication-compatible geometry has been shown to lend itself well to development of optical chemical sensors having a range of integrated functionalities. However, integration of multiple functionalities (e.g., light delivery and analyte detection capability) is not limited to a planar geometry, as evidenced by the volume of work on active fiber-based sensors. The broad range of exciting developments in the field of optical chemical sensor platform development is indicative of the increasing interest in these devices and an acknowledgment of their effectiveness in providing sensitive, practical sensor solutions for a variety of application areas.

3. Direct Sensors 3.1. Introduction Direct optical detection involves the measurement of some intrinsic optical property of the analyte. Many direct sensors, for example, gas sensors, measure the intrinsic optical absorption, usually in the infrared (IR), using a variety of techniques including Fourier transform infrared (FTIR) spectroscopy or correlation spectroscopy. A range of optical configurations has been used, from free-space sensors to fiber-based configurations to optical waveguide sensors. Fiber optic and planar waveguide sensing platforms have been discussed in section 2 of this review as well as refractometric platforms such as the Mach Zehnder interferometer. Developments in IR sensing have been driven largely by the emergence of new diode laser sources, for example, QCL lasers. This section includes a comprehensive discussion of direct IR absorption sensing including FTIR-based systems. While a review of QCL-based sensing appears elsewhere in this issue, a short section on this topic is included here under the heading of diode laser-based sensing. A brief section on UV absorption is included, mainly concerned with environmental sensing of organic pollutants. Similarly, direct fluorescence sensing is dealt with mainly in the context of biomedical applications. Recent developments in SERSbased sensing are reported, which reflect the huge growth in this area in the last 10 years.

3.2. Direct Spectroscopic Sensing 3.2.1. Absorption-Based Sensors IR Absorption. For many applications, spectroscopic detection has been a reliable method of detecting chemical species. IR spectroscopy has been widely used for detection of gases, for example, CO2, CO, NO2, NH3, and CH4. In its simplest form, the technique involves confining a sample of the gas in an optical absorption cell and measuring the absorption at specific IR wavelengths, which are characteristic of the vibrational modes of the molecule. The system components usually include an IR source, optical filters to select specific absorption wavelengths, and a detector, which

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is sensitive at the wavelength of interest. Referencing can also be employed using a wavelength that is not absorbed by the analyte molecule. The fundamental vibrational transitions of many molecules of interest occur in the near- and mid-IR, spanning wavelengths from 1 to >10 µm. Gas analysis is becoming increasingly important in control of various industrial processes, for example, in combustion processes and monitoring industrial emissions. The need for environmental control of emissions has also led to development of analyzers to measure the concentration of a variety of gases. Much of this instrumentation is based on near- and mid-infrared absorption spectroscopy. While conventional IR sensing uses broadband IR sources, recent developments in diode laser technology, including the relatively recent availability of novel QCL sources, have led to development of highly sensitive and selective sensor systems. Fourier transform infrared (FTIR) spectroscopy is also a very powerful analytical tool. Nondispersive infrared (NDIR) sensors are widely used for gas analysis, and many commercially available gas analyzers are based on this principle. NDIR Sensors. NDIR sensors use band-pass filtering in order to select the specific analyte absorption wavelength as opposed to dispersion by a prism or a grating as employed in a spectrometer. NDIR-based sensors have been developed to monitor CO2 concentrations in the ocean in order to monitor the effects of increased atmospheric CO2 on the global carbon cycle. Total inorganic carbon measurement involves acidification of the seawater sample and quantitative measurement of the evolved CO2. NDIR detection has been found to yield a precision of 0.11-0.25% and an accuracy of ∼0.1% for total inorganic carbon detection.90,91 Kaltin et al.92 employed NDIR to measure the extracted CO2, and the system continually compares the signal to a certified reference material, which constitutes the baseline signal. A precision of ∼0.05% was achieved with an accuracy of ∼0.2%, which was comparable to the performance of the standard colorimetric technique, and the sampling time of 5 min was 3 times faster than the colorimetric method. In order to produce the speed of measurement desired to enable highspeed continuous analysis, Bandstra et al.93 employed a gaspermeable hydrophobic membrane contactor, which continuously strips the CO2 out from a flowing stream of seawater. The system continuously quantifies the CO2 concentration using an NDIR analyzer, which uses a lead selenide detector and broadband blackbody source, yielding an accuracy and precision better than (0.1% and a response time of 6s. FTIR-Based Analyzers. FTIR-based sensing is an alternative to NDIR-based analyzers and has advantages for many applications. Current FTIR spectrometers offer precise quantification of a wide range of analytes for concentrations down to single-digit ppb levels and are now available as integrated and relatively compact units compared to the complex and bulky units offered in the past. An FTIR spectrometer consists of an interferometer, usually a Michelson type, which generates an interferogram from the IR emission of the sample and then performs a Fourier transform to obtain the spectrum. The ability of FTIR to measure multiple analytes simultaneously differentiates the technique from NDIR sensors. Castro et al.94 recently reported the use of FTIR to characterize low-temperature combustion gases in biomass fuels. The spectrometer, operating in absorption mode, was placed in front of an IR source and mounted such

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that the optical line of sight was just above the sample, allowing the absorption of the emitted gases to be measured. A high acquisition speed was necessary in order to evaluate temporal evolution of gas concentrations and improve the signal-to-noise ratio. A spectral resolution of 1 cm-1 was selected as a compromise between required spectral resolution and acquisition time. The concentration of three gases, CO2, CO, and CH4, which are the most important carbonrelated products of biomass combustion (shrub species in this case), was measured in situ. These results can contribute to improved modeling of pyrolysis for predicting forest fire behavior. FTIR has also been used for analysis of vehicular emissions; for example, Reyes et al.95 monitored a range of pollutants including CO, NO, SO2, and NH3 under different driving conditions using a hybrid car. However, no detailed performance parameters were given. FTIR imaging has been used for high-throughput analysis of processes such as epoxy curing,96 gaseous or solid-phase reactions,97 and monitoring of pharmaceutical formulations under different conditions.98 Chan et al.98 applied FTIR imaging to the study of the behavior of drug formulations such as ibuprofen as a function of different environmental parameters. They combined a microdrop device with an FTIR imaging system, which consisted of a step scan spectrometer coupled to a macrochamber extension and a 64 × 64 focal plane array (FPA) detector. Spectra were measured with a 16 cm-1 resolution and a spectral range of 3950-900 cm-1. They demonstrated for the first time the simultaneous analysis of 100 samples. The system allows fast screening of many formulations under controlled environments. FTIR spectroscopy has been used for on-line monitoring of formaldehyde-based resin synthesis.99 The system used a FTNIR spectrometer fiber optic data acquisition, and spectra were collected at 8 cm-1 optical resolution. Chemometrics was employed to provide fast and reliable process information on an industrial scale. No sample manipulation was required, and data was acquired in less than a minute. Diode Laser Sensing Systems. While FTIR is well suited to multicomponent analysis of gases and other chemical species for a range of processes, laser spectroscopy is the method of choice for trace gas analysis because of its higher sensitivity and specificity, which arises when a spectrally narrow laser source probes a narrow absorption feature of the analyte. Werle100 and Allen101 produced comprehensive reviews of diode laser-based gas sensor systems for applications such as automated control of industrial processes, environmental monitoring, and combustion process monitoring. The lasers used for gas analysis span the near- and midIR spectral regions with wavelengths ranging from 1 to >2µm for conventional diode lasers and from 2 to >4µm for recently developed QCL lasers. Fiber-coupled multiwavelength diode laser sensor systems have been designed which use modulation spectroscopy and multipass absorption cells to increase sensitivity. Wavelength modulation spectroscopy has been used to monitor gas emission in industrial processes.102 Tunable diode laser absorption spectroscopy (TDLAS) systems, using lead salt lasers which are based on IV-VI semiconductors, have been used to achieve parts per billion detection levels of atmospheric gases, for example, methane.103 However, these systems require laser cooling and are not suited to unattended routine industrial applications. While lead salt lasers operated at cryogenic temperatures cover the fundamental gas absorp-

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tion bands required for ultrasensitive gas analysis, nearinfrared diode lasers, operated at room temperature, probe mainly the weaker overtone and combination bands of the analyte molecule. Therefore, there is a trade off between sensitivity and convenience of operation. Sensitivity problems can be overcome using appropriate signal processing and double-modulation techniques.104 QCL lasers are semiconductor lasers based on transitions in a multiple quantum well heterostructure. The emission wavelength depends mainly on the quantum well thickness rather than on the size of the band gap, as is the case with conventional diode lasers. QCLs operate at wavelengths in the MIR starting at about 3 µm, which match well with the fundamental vibrational absorption bands of many gases and other chemical species in comparison with conventional diode sources where the laser emission generally matches the weaker overtone bands. QCLs operate at near room temperature, produce milliwatts of radiation, and offer the possibility of tailoring the emission wavelength within a broad range of frequencies.105,106 The most technologically developed QCLs are based on InGaAs-InAlAs and GaAs-AlGaAs heterostructures.107 They are usually fabricated as single-mode lasers and often in a distributed feedback (DFB) configuration. QC-DFB lasers are of particular interest for gas sensing as they produce highly monochromatic radiation that is suitable for highresolution spectroscopy.108 The first reports of gas sensing using a CW QC-DFB laser was by Sharpe et al.,109 where they presented absorption data on NO and NH3 at low pressure. The first QCL-based system for determination of CO2 in aqueous solution was reported by Schaden et al.110 A QCL gas sensor system for measuring trace gases in air was reported by Kosterev.111 Gases such as CH4 and N2O were detected down to ppb levels. A QCDFB laser system was used to measure ppb levels of NO in vehicle emissions,112 while high-precision measurements of atmospheric N2O and CH4 were reported by Nelson et al.113 Sensing in liquids has also been reported. A Fabry-Perot QC laser system, consisting of two lasers, one tuned to a water vibration and the second to a water window, was used in a flow injection system equipped with a fiber optic flow cell with an adjustable path length. Adenine and xanthosine were used as test analytes, and the results demonstrated the advantages of QCL lasers in overcoming solvent interference for liquid sensing.114 UV Absorption. Direct UV absorption sensing has been used mainly in environmental applications to monitor pollutants such as heavy metals, hydrocarbons, and volatile organic compounds (VOC) in air and water. A sensor to detect Cr(VI) in water that is based on a flexible fused silica capillary and intrinsic evanescent-wave UV absorption of Cr ions in a water sample in the capillary has been reported.115 A Cr(VI) absorption band peaking at 374 nm was monitored, and a limit of detection of 31 parts in 109 was achieved using an absorption path length of 30 m of capillary. The concentrations of ozone and NO2 in the atmosphere were measured by UV-differential optical absorption spectroscopy (UV-DOAS), and the results were used to evaluate the contribution of NO2 in limiting ozone formation.116 The UV-DOAS system used a broad-band xenon source that was placed at a level of 12 m on top of a building, and the detector was placed 150 m away. A premeasured clean-gas reference spectrum was used to generate the differential absorption spectrum. UV-DOAS was

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also used to monitor VOCs and aromatic hydrocarbons in ambient air in the vicinity of a refinery.117 The system detected high levels of benzene and toluene and also monitored seasonal variations in VOC concentrations. UVDOAS was one of four VOC monitoring techniques used in a field study in Mexico to identify VOCs that were primarily due to vehicle emissions.118 The data were analyzed to understand the spatial distributions, diurnal patterns, and reactivity of various VOCs. One of the key results of this study was a better understanding of the interaction between vehicular activity and meteorological processes.

3.2.2. Direct Fluorescence Sensing Direct fluorescence sensing is widely used in biomedical applications. Autofluorescence spectroscopy is a useful tool for noninvasive detection of precancerous development of the epithelium, where most human cancers originate. This fluorescence arises mainly from two intrinsic fluorophores, reduced nicotinamide adenine dinucleotide (NADH) and flavin adenine dinucleotide (FAD). The fluorescence of these species can be used to monitor cell activity. For example, the ratio of free and bound NADH, differentiated by their different lifetimes, is a good indicator of the cell metabolic state. Wu et al.119 used a time-resolved confocal fluorescence spectroscopy system to study cell metabolism and were able to differentiate between cancerous and normal cells by analyzing the dual-exponential decay of NADH. Depthresolved fluorescence spectroscopy, using a confocal microscope and multiple excitation wavelengths, was used to study different layers of the epithelium.120 By exciting at 355 and 405 nm, the system was able to isolate the fluorescence from different sublayers. In particular, the ratio of NADH to FAD fluorescence can be used to indicate the presence of precancerous cells. Autofluorescence patterns have also been used to follow changes in cervical tissue121,122 and oral tissue.123 The intrinsic fluorescence of tryptophan has been used to study protein folding and unfolding. Tryptophan fluorescence is very sensitive to the environment, whereby the peak emission wavelength shifts from 308 to 350 nm for the range of protein conformational changes.124 This spectral behavior has been used to study protein stability125 and protein aggregation.126 Fluorescence can be used to characterize different oils as the fluorescence is dependent on chemical composition. For example, crude petroleum oils have been analyzed using synchronous fluorescence, where both the excitation and emission monochromators are scanned simultaneously, and lifetime measurements were carried out.127,128 Recently, biocrude oils, derived from seeds, have been analyzed.129 Synchronous fluorescence has also been used to quantitatively discriminate between virgin olive oil and sunflower oil,130 and fluorescence spectroscopy combined with artificial neural networks has been used to classify a range of edible oils.131

3.2.3. Raman and SERS Sensing Like IR-based absorption spectroscopy, Raman spectroscopy probes the vibrational energy levels of molecules and therefore can be used as a highly selective technique which will distinguish between similar molecules. The Raman effect occurs when a photon interacts with the vibrational energy levels of a molecule and is scattered.132 Raman spectroscopy normally requires laser excitation and relies on good optical filtering to separate the scattered photon from the intense

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incident beam. Unfortunately, Raman scattering crosssections are very small, many orders of magnitude smaller than fluorescence cross-sections. However, it is still a useful technique for sensing of gaseous and liquid analytes. It has a particular advantage over IR absorption for liquid sensing in that there is no interference from the vibrational spectrum of water. For this reason, Raman is a useful technique for life science applications, including biomedical diagnostics, for example, the study of living cells. Lasers with large spot sizes (>10 µm) can be used to study whole cells and tissues. Raman imaging using diffraction-limited spot-size lasers can be used to map the chemical distribution in a cell.133 Uzunbajakava et al.134 used a laser with high lateral resolution to map the DNA and RNA distribution in GeLa cells. Principal components analysis of Raman spectra has been used to distinguish between different cancer cells.135,136 A Raman imaging system, based on fiber optic probes, has been used to image breast tissue.137 Different tissue features were identified from differences in the Raman spectra. Toxins such as sulfur mustard and ricin have been identified in cells with an identification accuracy of 88.6% and 71.4%, respectively, using Raman spectroscopy.138 The disadvantages of conventional Raman, such as the low scattering cross-sections, can be overcome to a large extent using the SERS technique. The strength of Raman scattering from a molecule in close proximity to a nanostructured metal surface can be enhanced by as much as a factor of 108, compared to conventional Raman. This surface-enhanced effect is strongest for silver and due to the interaction of the localized surface plasmon, generated at the metal surface, with the vibrational levels of the molecule. The large electromagnetic field generated induces a dipole in the molecule that is adsorbed on the metal surface, hence producing the hugely enhanced Raman signal.139 The two principal conditions required for SERS to be observed are the presence of a suitable SERS-active nanostructured metal surface and the requirement that the sample under investigation be immobilized on or in close proximity to the surface. In the past, routine application of the technique has been hampered by a poor understanding of the theoretical background and lack of reproducibility of SERS substrate materials. Recently, with advances in materials fabrication and better understanding of the details of the plasmonic interaction, SERS is being increasingly used in diverse fields such as biomedicine and environmental analysis. A number of review articles have been published recently with particular emphasis on biomedical applications of SERS.140-142 The use of SERS to obtain quantitative in-vivo glucose measurements has been reported,143 where a suitable SERS surface was achieved using a silver-coated self-assembled monolayer (SAM) of polystyrene nanospheres. The SERS probe was implanted in a rat, and the data obtained agreed well with simultaneous data monitored using a conventional electrochemical glucose sensor. The same substrate has also been used to detect the presence of the chemical warfare agent half-mustard.139 A combination of near-field scanning optical microscopy and SERS has been used to detect dyelabeled DNA with 100 nm resolution.144 SERS has been used to detect specific dyes in works of art. Anthraquinone145 and alizarin dyes have been identified.146 For the alizarin investigation, silver nanoparticles were deposited directly onto a sample of the painting. The advantage of SERS over conventional Raman was highlighted in this work as strong fluorescence of the dye completely masked the Raman signal,

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while in the SERS measurement, the fluorescence was quenched due to the presence of the metal substrate, enabling detection of the enhanced Raman signal. In environmental sensing, a SERS optode has been used to detect polycyclic aromatic hydrocarbons (PAHs) in seawater.147

3.3. Summary This section on direct sensing has focused exclusively on spectroscopic techniques. A short review of the range of IR absorption techniques and configurations used in gas analysis and sensing of organic compounds was given. NDIR- and FTIR-based sensing are both used extensively, FTIR being more appropriate for high-throughput, multianalyte applications. Laser-based systems are more complex and expensive but offer very high sensitivity, allowing detection down to ppb levels. The availability of QCL sources has been a novel development in this area and enables relatively compact, robust systems for use in the field. The recent growth of SERS-based sensing was discussed, and examples of diverse applications, including some biosensing applications, were given. In summary, it has to be said that, currently, the most widely used direct optical sensing technique is that of IR absorption-based sensing, mainly for gas analysis. For example, of the total commercial gas analyzer market in the United States, NDIR sensors had about a 50% share in 2005. It will be interesting to observe whether this technique still dominates the optical sensing market over the next decade. It will also be interesting to see if the current growth in SERS sensing will be sustained and whether it will lead to commercial sensor systems.

4. Reagent-Mediated Sensors 4.1. Introduction When an analyte does not exhibit a convenient spectroscopic optical response such as absorption or luminescence, sensing can be achieved by monitoring the optical response of an intermediate species or reagent, whose response is modulated in some way by the presence of the analyte. Reagent-mediated sensing is illustrated very well by the principle of optical oxygen sensing, whereby the luminescence intensity or decay time of an oxygen-sensitive luminescent complex, for example, a ruthenium polypyridyl or a porphyrin complex, is quenched in the presence of oxygen. This enables the oxygen partial pressure to be measured as a function of the intensity or luminescence decay time of the complex. Indirect colorimetric sensors also require an intermediate reagent, for example, many optical pH sensors have been based on monitoring the change in optical absorption of pH indicators such as bromocresol purple. This indirect sensing technique requires the reagent to be immobilized, either in the liquid or in the solid phase, to facilitate interaction with the analyte. In recent years, reagentbased optical sensing has been based mainly on solid-phase immobilization matrices, where the reagent dye can be adsorbed, covalently or ionically attached, or simply encapsulated in a solid matrix that is permeable to the analyte. If the immobilization matrix has the capability of being coated on a substrate in liquid form, as is the case for sol-gel glasses or polymer coatings, a wide range of sensor configurations is enabled including, for example, fiber optic, planar waveguide and array-based sensors. These configura-

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tions have already been described in section 2. Section 4.2 of this review will report on the wide range of reagents used in recent years in optical absorption- and luminescence-based chemical sensors, while section 4.3 describes two widely used sensor immobilization matrices, namely, sol-gel and polymer materials. Subsequent sections review the state-ofthe-art in absorption- and luminescence-based systems, dealing mainly with work that has been published in this decade.

4.2. Reagents 4.2.1. Reagents for Colorimetric Sensing Indirect colorimetric pH sensing uses organic pH indicators, the absorbance of which is modified by the pH of the environment. The pKa of these indicators indicates the center of the measurable pH range, for example, cresol red, bromophenol blue, and bromocresol purple respond to acidic pH (pH < 7), while cresol red, naphtholbenzene, and phenolphthalein respond at basic pH (pH > 7).148 The relatively narrow pH response range of most of the above dyes has been addressed whereby several indicators with different pKa values have been combined in one sensor in order to produce a linear pH response over a wide range.149 Makedonski et al.150 synthesized new reactive azo dyes for use in pH sensing where the dyes were covalently bound to a polymer matrix for increased sensor stability. Recently, phenolphthalein has been used to monitor the behavior of corrosion-resistant polymers in alkaline solutions in chemical plants.151 Since colorimetric CO2 sensing is normally achieved by measuring the pH change of an indicator in response to carbonic acid generated by the acidic CO2 gas, many of the same reagents that are used for pH sensing are also used in absorption-based CO2 sensing. Thymol blue immobilized in a sol-gel matrix has been used for gaseous CO2 sensing,152 while bromothymol blue in an ionic liquid matrix has been used for both gas-phase and dissolved CO2 sensing.153 Cresol red has been used in an optical fiber configuration in order to measure in vivo gastric CO2.154 Recent reviews by Mohr155,156 reported the development of new indicator dyes for neutral and ionic analytes. Many of these use reversible covalent bond formation to detect analytes such as amines, cyanide, formaldehyde, nitrites, and peptides. Selective colorimetric optical probes based on pyrylium dyes have been developed recently for detection of aliphatic amines.157,158 Novel, highly selective probes for formaldehyde detection have been developed by Suzuki et al.,159 while a formaldehyde sensor, based on the reagent 4-amino dydrazine-5mercapto-1,2,4-triazole, has been reported recently.160 A highly selective colorimetric Hg2+ sensor, which uses a squaraine reporter dye, has been reported.161 This highly toxic heavy-metal ion has also been detected optically by immobilizing dithizone on a triacetylcellulose membrane.162 Squaraine derivatives have also been used for colorimetric detection of cyanide ions.163 A fluoride-selective sensor based on the complexation of fluoride ion with an aluminum octaethylporphyrin ionophore dye has been reported.164 The sensor exploits the fact that aluminum(III) forms very stable complexes with fluoride in aqueous solutions. A breath sensor for acetone measurement, which detects the colored product that is formed when acetone reacts with alkaline salicylaldehyde, has been reported by Teshima et al.165

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4.2.2. Reagents for Luminescence Sensing Analytes such as pH, CO2, ammonia, O2, and many cations and anions can be measured using luminescent probes. Luminescence is intrinsically more sensitive than absorption as a sensing technique, so, for many applications, luminescence-based sensing offers higher sensitivity than absorption sensors. As in the case of absorption-based sensors, the principle of luminescence measurement of pH, CO2, and ammonia relies on the change in luminescence of a pH indicator. For these analytes, the literature is dominated by two pHdependent luminescent probes, namely, fluorescein and 8-hydroxypyrene-1,3,6-trisulfonic acid (HPTS). Both dyes absorb at convenient LED-emission wavelengths in the blue, while emission occurs just above 500 nm. Fluorescein has been used recently in an optical fiber imaging sensor based on drop-on-demand inkjet printing.166 Fluorescein-related dyes have been used in order to facilitate covalent immobilization to the sensor matrix to eliminate dye leaching. For example, pH sensing has been realized using a fluoresceinamine isomer II complex that was covalently bound to a sol-gel matrix.167 Ratiometric sensing is a self-referencing technique where either an analyte-insensitive excitation or a luminescence band of the dye is ratioed with the analytedependent band. Excitation ratiometric sensing has been achieved using methacryloyl-modified HPTS that facilitates covalent bonding to a polymer matrix and has a sensing range of pH 6-9.168 Ratiometric pH sensing has also been demonstrated using mercurochrome, another fluorescein-related dye, in a solgel matrix.169 A newly synthesized boron-dipyrromethene derivative has been used in a single-excitation dual-emission ratiometric scheme to measure pH in human gastric juices.170 Naphthalimides have also been investigated as luminescent probes,171,172 while a pH sensor using covalently immobilized piperazinyl-1,8-napthalimide has been published recently.173 Luminescent transition-metal complexes have larger Stokes shifts and longer lifetimes than, for example, fluorescein or HPTS. This allows more flexibility in excitation/detection optoelectronics as well as enabling lifetime-based sensing schemes. pH sensing has been demonstrated recently using a range of ruthenium and rhenium complexes based on the pH-sensitive ligand 5-carboxyl-1,10-phenanthroline.174 In this work, the complexes have been tailored for maximum dynamic range. Aminofluorescein has been used for optical ammonia sensing where an enhancement has been achieved compared to fluorescein-based sensing due to the reaction of the ammonia with the amine group on the fluorescent dye.175 Ruthenium(II) tris(4,7-dephenyl-1,10-phenthroline) (Ru(dpp)3) has been ion paired with the pH indicator bromophenol blue.176 On exposure to ammonia, the increase in the deprotonated band of the pH indicator gives rise to energy transfer whereby the luminescence and lifetime of the Ru(dpp)3 complex decreases. In a novel, multianalyte approach, Nivens et al.177 used HPTS to detect pH, CO2, and NH3 by designing multilayer sol-gel structures whereby, for CO2 and NH3 sensing, a hydrophobic outer layer is used to eliminate cross-reactivity to pH. HPTS has been widely used for CO2 sensing, where, as above, the matrix has been tailored to prevent ingress of protons. HPTS is stable and water soluble, and the pKa is ideal for CO2 detection.178 Lifetime-based sensing based on HPTS in polymer matrices has been reported.179,180 A HPTS-

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based ratiometric sensor has been published recently for monitoring CO2 in marine sediments.181 The luminophores used for optical oxygen sensing have been based mainly on organometallic complexes of ruthenium, in particular ruthenium polypyridyl complexes, and also metalloporphyrin complexes. However, a number of papers have reported on a range of polycyclic hydrocarbons (PAHs) that are efficiently quenched by oxygen.182-184 Ruthenium polypyridyl complexes are characterized by high quantum efficiency and convenient absorption and emission maxima located in the visible spectrum with a large Stokes shift and relatively long lifetime (∼1 µs). These complexes are efficiently quenched in the presence of oxygen, giving rise to a decrease in both luminescence intensity and lifetime. The complex Ru(dpp)3, referred to above in the context of pH sensing, has one of the longest unquenched lifetimes (∼5 µs) and been widely reported for oxygen-sensing applications, both in polymer and sol-gel matrices.185-187 Porphyrin complexes have longer lifetimes (∼100 µs), emit at the red end of the visible spectrum, and are generally considered to have inferior photostability properties compared to ruthenium complexes. However, fluorinated porphyrins have been shown to be highly stable against photo-oxidation and photoreduction.188 The most widely reported complexes for oxygen-sensing applications are the platinum and palladium octaethylporphyrin (PtOEP and PdOEP).189,190 Platinum and palladium tetrakis(pentafluorophenyl)porphyrin (PtTFPP and PdTFPP) have also been reported.191 Ruthenium complexes have also been used for luminescence-based relative humidity sensing based on the quenching of a phenazine ligand by protons.192-194 There are extensive literature reports on the use of luminescent probes to measure anionic and cationic species as well as organic chemicals. Some of this work has been referenced in the recent review by Wolfbeis.5 Probes used for chloride sensing include the chloride-quenchable lucigenin dye14 and the halide-selective ionophore9 mercuracarborand-3.195 Aluminum(III) concentration has been measured using the dye 1,4-dihydroxyanthraquinone (quinizarin). Sensing is based on the highly fluorescent complex which is formed when Al3+ complexes with quinizarin.196 Molecularly printed polymers (MIPs) have also been used where Al3+ is used as the template and 8-hydroxy-quinoline sulfonic acid is the luminescent tag which complexes to the Al3+ ion.197 Badagu et al.198 developed a range of cyanide-sensitive probes that are based on the binding of the cyanide ion to a boronic acid functional group. Both intensity and lifetimebased sensing have been demonstrated. Luminescence-based nitrate sensing has been achieved using the cationic potentialsensitive dye 4-(4-(dihexadecylamino)styryl-N-methylpyridinium iodide), which acts as an anion-exchange catalyst that extracts nitrate from aqueous solution to form a luminescent complex.199

4.2.3. Summary Clearly, a comprehensive account of the large number and variety of absorption and luminescence probes reported in recent years is beyond the scope of this general review of optical chemical sensors. The most widely used reagents in areas such as pH and oxygen sensing have been highlighted. In the case of ion sensing, a selection of recently published work has been cited, which deals with some environmentally topical analytes such as chloride and aluminum ions.

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4.3. Immobilization Matrices 4.3.1. Introduction In most reagent-based optical sensors, the reagent is immobilized in a solid matrix usually in the form of a monolith or a thin film. The matrix serves to encapsulate the reagent such that it is accessible to the analyte while being impervious to leaching effects. In this section, two commonly used immobilization matrices will be described, namely, sol-gel and polymer materials, and a selection of work published in recent years will be reported for sol-gel and polymer-based absorption and luminescence sensors. Since the immobilization matrix very often influences the sensor response, a brief discussion of reagent-matrix interactions will be given.

4.3.2. Sol−Gel Matrices The sol-gel process provides a relatively benign support matrix for the immobilization of analyte-sensitive reagents and dyes. The basic process involves the hydrolysis and polycondensation of the appropriate metal alkoxide solution to produce a porous glass matrix in which the reagent is encapsulated in a nanometer-scale cage-like structure and into which the analyte molecules can diffuse. The versatility of the process facilitates tailoring of the physicochemical properties of the material in order to optimize sensor performance. For example, key sol-gel process parameters such as sol pH, precursor type and concentration, water content, and curing temperature can be adjusted to produce materials of desired porosity and polarity.200 This tailorability was demonstrated for oxygen-sensitive films doped with the Ru(dpp)3 complex and were fabricated using different ratios of tetraethoxyorthosilane (TEOS) and methyltriethoxysilane (MTEOS) precursors.201 In a later study, the effect of the alkyl (R) chain length on the oxygen sensitivity was measured.202 Monoalkoxysilanes of the form (CnH2n+1)-Si-(OR)3, where n ranges from 1 to 12, have been investigated and exhibit a linear oxygen response which increases up to C8 and then begins to decrease. The decrease in response at longer alkyl chain lengths was correlated with a decrease in oxygen diffusion coefficient in the matrix for these films. Very high oxygen sensitivities have been achieved using fluorinated sol-gel precursors. For example, oxygen sensitivities have been achieved using precursors such as 3,3,3-trifluoropropyl-trimethoxysilane (TFP-TMOS), which are ∼10 times greater than that achieved for MTEOSbased sensors.203,204 This high sensitivity is attributed to the highly hydrophobic and nonpolar nature of the fluoro films. These materials also exhibit very high stability, with stabilities of up to 2 years being achieved for some formulations. Novel phenyl-substituted ORMOSILs were used as matrices for oxygen sensors based on both the Ru(dpp)3 complex and a platinum(II)-octaethylporphyrin complex. A highly stable and sensitive oxygen response was obtained, and the films were sterilizable.205 As an alternative to physical entrapment, the analytesensitive dye can also be covalently bound to the sol-gel matrix. This can increase stability and eliminate dye-leaching effects in aqueous environments. Some examples have already been referred to in section 4.2.167,168 A triethoxysilane-modified ruthenium tris(bipyridyl) (Ru(bpy)32+) complex has been anchored to a range of polysiloxane matrices,206 and an oxygen sensor based on the same complex has been recently reported where the complex was covalently grafted

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to 3-aminopropyltriethoxysilane.207 The sensor exhibited no leaching and was used for dissolved oxygen measurement. Sol-gel matrices have also been used for other analytes such as in pH and CO2 sensing. As discussed in section 4.2, for pH sensing a hydrophilic matrix which allows ingress of protons is required, whereas for CO2 sensing a more complex matrix is required which retains the water required for conversion of the gas to carbonic acid in the films but has sufficient hydrophobicity to be pH insensitive. These conditions have been achieved for both sol-gel and polymer matrices using an ion-pairing approach. This will be described in more detail in the next section as the majority of the literature on optical CO2 sensors is based on polymer matrices. However, sol-gel-based optical CO2 sensors for both gas and dissolved phase have been reported.208,209,43 Standard sol-gel materials have a nonordered, amorphous structure where diffusion of analytes can be limited by the random microporosity of the structure. Mesostructured porous films, on the other hand, have large open porosity, which can offer enhanced diffusion and accessibility for analytes. Mesoporous sol-gel sensor films have been realized via an evaporation-induced self-assembly (EISA) approach.210,211 These films have a highly structured mesoporosity and excellent optical quality. The films are based on the surfactant cetyltrimethylammonium bromide (CTAB). The CTAB/Si and probe/Si ratio is optimized to give the required structure. A sensor to detect metallic cations was reported which was based on the chelating properties of dibenzoylmethane.211 An oxygen sensor based on a mesostructured sol-gel silica matrix was also reported.212 A ruthenium complex was grafted to the mesostructured network that was based on the precursor 3-(triethoxysilyl)propyl isocyanate and CTAB as surfactant. It was established that the stability, homogeneity, and sensitivity of the matrix was superior to a nonmesoporous matrix where the dye was just physically entrapped. A novel sol-gel-based sensor application has been reported whereby an optical fiber core is fabricated from reagent-doped porous sol-gel. Hence, the core acts as an active fiber core optical sensor (AFCOS). The fabrication and optical properties of the fiber were investigated, and a humidity sensor was constructed as proof of principle.213 Oxygen sensing has been reported using calcined mesoporous silica spheres which have been postdoped with the Ru(bpy)32+ complex.214 The oxygen sensitivity was dependent on the pore morphology in the sphere, and the response times (2-4 min) were longer than for thin films and most likely related to the relatively large sphere diameter of 0.1-0.2 mm.

4.3.3. Polymer Matrices Polymers have been widely used as support materials for a broad range of optical sensors. They have many desirable features and compare well with sol-gel matrices for most applications. While polymers may not be as photochemically stable or printable as sol-gels, some polymers are more suitable than sol-gels for high-temperature applications such as autoclavation. The most widely used materials include polystyrene (PS), polyvinyl chloride (PVC), polymethyl methacrylate (PMMA), polydimethyl siloxanes (PDMS), and polytetrafluoroethylenes (PTFE) and cellulose derivatives such as ethyl cellulose. As discussed in the previous section, hydrophobic matrices, for example, PMMA and PDMS, are selected as the optimum matrix for optical oxygen sensing,

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while more hydrophilic matrices such as ethyl cellulose have been widely used for pH sensing. Much has been published on polymer-based oxygen sensors. Mills215 carried out a comprehensive investigation of the effect of various polymers and plasticizers on the oxygen sensitivity for mainly ruthenium-complex-doped materials. It was clear from this work that the value of the oxygen diffusion coefficient in the material was the main contributing factor in the oxygen sensitivity. Of all materials studied, PDMS yielded the highest diffusion coefficient, which is associated with the opening and closing of void volumes in the polymer. It is to be noted that similarly high oxygen sensitivities were achieved using fluoro-sol-gel materials by Bukowski.204 This has been attributed to the high O2 diffusivity in fluorinated materials.216 Polymer-based oxygen optrodes have been reported which withstand autoclavation. Polysulfone and polyetherimide materials were found to withstand steam sterilization at 135 °C.217 It is thought that these polymers have a different water uptake compared with other materials such as polyvinylnaphthalene, which are destroyed under these conditions. The dissolved oxygen sensors based on these polymers, once calibrated, do not need recalibration in between sterilization cycles. Polysulfone was also used successfully to monitor cell viability by measuring the oxygen consumption.218 A novel oxygen sensor membrane has been reported by Holmes-Smith et al.219 where an electropolymerized diphenyl-di(4-aminophenyl)porphyrin (Pt(II)DAPP) has been characterized and tested as an optical oxygen sensor. The sensitivity obtained was comparable to other Pt-porphyrin systems reported in the literature, but the main advantages of this sensor were the versatility of deposition via the electropolymerization technique as well as the lack of leaching as the porphyrin is part of the polymer. Optical CO2 sensors are generally based on the Severinghaus principle, which relies on conversion of the gas to carbonic acid. In the presence of bicarbonate, it has been shown that the partial pressure of CO2 is directly related to the pH change of a colorimetric or luminescent pH indicator.178 In solid optical CO2 sensors, a phase-transfer agent such as tetraoctylammonium hydroxide (TOAOH), a quaternary ammonium hydroxide, is used to solubilize the pH indicator into a hydrophobic polymer which acts as a proton barrier and reduces cross-sensitivity to pH. This gives rise to formation of an ion pair which facilitates encapsulation of the dye in the gas-permeable polymer. The quaternary ammonium hydroxide is normally associated with sufficient water molecules to facilitate formation of carbonic acid. A fiber optic microsensor based on ion-paired HPTS in an ethyl cellulose matrix has been reported for highresolution pCO2 monitoring in the marine environment.220 A gas-permeable but ion-impermeable coating was used as a protective layer to eliminate interferences such as chloride and pH, and the limit of detection achieved was 60 ppb. A highly stable, autoclavable sensor, based on HPTS ion paired to cetyltrimethylamonium hydroxide (CTMAOH) in a twocomponent silicone film, has been investigated for use in biofermentation.221 PDMS and vinyl-terminated dimethylsiloxane copolymers were used to make the sensor film using a three-step process. The sensor used ratiometric detection, was stable for several months and sterilizable, and had an LOD of 0.03% CO2. It has been observed that CO2 sensors based on a quaternary ammonium hydroxide base can have a limited shelf life in ambient air due to the ingress of other acidic

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gases such as NOx and SO2, which cause irreversible protonation of the indicator. It has been suggested that this is due to a process called Hofmann degradation.222 The feasibility of replacing the quaternary base with a neutral phosphazene base which would not be susceptible to degradation has been investigated by Schroder et al.223 Ethyl cellulose was the polymer used, and p-xylenol blue was the colorimetric pH indicator. The CO2 response of the new phosphazene base system was generally comparable to that of the quaternary base system. However, the new base could only be used in dissolved phase due to its volatility, and the sensor film exhibited cross-sensitivity to relative humidity. Molecularly imprinted polymers (MIPs) are being increasingly used as sensor supports. MIPs are synthesized by copolymerizing functional and cross-linking monomers in the presence of a target analyte molecule which acts as a template or imprint. When the template molecule is removed, the resulting cavity has a “memory” of the target analyte which imparts a high degree of selectivity to the MIP material.224 A polyurethane-based MIP sensor imprinted with anthracene was reported where a theoretical model of the sensor response was developed in order to provide an optimization strategy for sensor design.225 MIP sensors for detection of PAHs have been synthesized where the selectivity was tuned using different polystyrenes. Selectivity was tuned down to differences in size as little as one methyl group.226 Because of their selectivity, MIPs are now being developed for biosensor applications where antibodies and enzymes are used as templates. Reviews dealing with MIPs for biosensor and other applications include those by AlKindy et al.227 and Kindschy et al.228 Apart from their high degree of selectivity, MIP sensor materials are generally very robust, stable, and resistant to interferences such as pH and humidity.225

4.3.4. Interaction of Reagent and Support Matrix In order to be efficiently encapsulated in the support matrix, the reagent must be soluble in the support material. To enhance solubility, an ion-pairing approach can be used as discussed above for CO2-based sensing. In the case of sol-gel matrices, for example, in order to encapsulate oxygen-sensitive ruthenium complexes, the counterion and sol-gel solvent can be chosen to facilitate homogeneous distribution of the dye in the matrix. In the case of optical oxygen sensors, the linearity of the Stern-Volmer (SV) sensitivity plot is strongly dependent on the matrix. For most sol-gel and polymer materials, a downward curved SV plot is obtained due to the inhomogeneous nature of the amorphous matrix where each dye molecule experiences a slightly different microenvironment. However, linear SV plots have been obtained, for example, where the indicator dye is soluble in the matrix and homogeneously entrapped in a defined polymer microdomain.229 Linear SV plots have also been obtained for sol-gel materials which use precursors with long alkyl chain groups.202,204 The optical properties of some reagents are very sensitive to the environment; thus, often these reagents will experience spectroscopic shifts when encapsulated in a solid matrix. Of all the oxygen-sensitive ruthenium complexes, the spectroscopic properties of Ru(dpp)3 are relatively insensitive to the sensor matrix.230 In fact, these complexes exhibit enhanced fluorescence and longer decay times in solid matrices compared to in solution due to the reduction in nonradiative decay pathways. A substantial improvement in

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selectivity and sensitivity has been observed in the case of the metal-ion indicator pyrocatechol violet (PV) when encapsulated in a plasticized PVC membrane. A shift in the working wavelength was accompanied by enhanced sensitivity to Cu(II) ions. This is possibly due to a reduction in the complexation ability of the dye and the different permeability selectivity imparted by the PVC matrix compared to solution behavior.231 Many optical sensor reagents photodegrade under conditions of high-intensity illumination, and in many cases the matrix has on influence on the degree of photodegradation. In general, reagents entrapped in rigid supports exhibit increased photostability due to reduced ligand photodegradation compared to a solution environment.232 It has been shown that reagents encapsulated in sol-gel matrices have higher photostability than in polymer matrices. It has been suggested that the more organic polymer environment provides less stability than the more inorganic sol-gel matrix. In the case of oxygen-sensitive dyes, as mentioned in section 4.1, transition-metal complexes, for example, the ruthenium complexes, have higher stability than porphyrin complexes in general, but fluorinated porphyrin complexes exhibit high photostability.182 It has been established that, for oxygen sensors, singlet oxygen, which is a byproduct of the quenching process, causes ligand dissociation in oxygensensitive dyes.233 This degradation process is minimized in fluorinated matrices, either polymer or sol-gel based, which provide highly stable environments due to the high electronegativity of the C-F bond.182,203 In general, the behavior of the reagent can be dependent on the matrix as well as the analyte concentration.

4.4. Recent Developments in Absorption-Based Sensors Reagent-mediated absorption or colorimetric sensing has already been dealt with in section 4.2 on reagents, where numerous examples were given of colorimetric sensing for a wide range of analytes including pH, ammonia, and some ion-sensing applications. This brief section will highlight some recent examples of colorimetric sensors including some novel sensing techniques. A generic sensing platform has been reported which is based on a flexible fused silica capillary.234 The inner surface of a flexible light-guiding capillary was coated with a reagent-doped film, either polymer or sol-gel, to produce a long-path evanescent-wave-based sensor. The authors suggest that the flexibility and robustness of the fused silica capillary and the increased sensitivity provided by the increased length offer advantages over conventional optical fiber-based sensors as well as over previously reported capillary-based sensors. Sensing has been demonstrated for ammonia, toluene, and Cu(II) using this system. Cu(II) was detected down to 2.5 ppb, while ammonia was detected down to 15 ppm in the as-yet unoptimized system. A dynamic technique for measuring pH over a very wide range has been reported, where the limited measurable pH range of the indicator in normal steady state measurements has been overcome using a flow system and optimizing the diffusion of the analyte in the sensor membrane.235 In this work, the pH indicator was immobilized in a triacetyl cellulose membrane. The optical response to pH was measured before equilibrium was reached. At a fixed time before equilibrium, a linear plot is obtained for absorbance versus pH at any flow rate. The time and flow rate are

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optimized for maximum linear dynamic range. pH was determined over a range of up to 11 pH units. A combination of a colorimetric and fluorometric reagent has been used to measure CO2 where the luminescence change of a tetraphenylporphyrin dye due to the CO2dependent change in absorbance of a co-immoblized R-naphtholphthalein dye was measured.236 The technique relied on the overlapping absorption and luminescence of the colorimetric and fluorometric reagents, respectively, and the fluorophore which was insensitive to CO2 acted as an internal reference. The sensor operated in both gas and dissolved phase, was immune to photobleaching, and exhibited short response times of <6 s in the gas phase. The same authors used a thymol blue-europium complex combination for gasphase CO2 sensing.237 Mesoporous sensor matrices have already been discussed in section 4.3. A fiber-based colorimetric pH sensor based on bromothymol blue in a mesostructured silica film has been reported.238 The sensor responds over the pH range 2-5. Dye leaching is overcome by tailoring the mesopore dimensions to entrap the dye molecule, while the hydrophilic nature of the pore wall allows adsorption of the dye via hydrogen bonds. A CO2 sensor based on ion-paired bromothymol blue, which has been dissolved in an ionic liquid matrix, has been reported.239 The solubility of CO2 in water- miscible ionic liquids, such as 1-methyl-3-butylimidazolium tetrafluoroborate used in this work, is 10-20 times that observed in conventional solvent and polymer matrices. Sensing of CO2 was demonstrated both in gas and dissolved phase. However, regeneration of the reagent phase using nitrogen was necessary, and pH interference studies have not been reported in this preliminary report. The same authors reported a luminescence-based CO2 sensor based on HPTS dissolved in similar ionic liquid matrices.240

4.5. Recent Advances in Luminescence-Based Sensors 4.5.1. Introduction Reagent-based luminescence sensing can be based on monitoring the intensity of the luminescence, often using a ratiometric approach as mentioned in section 4.2.2, or the reagent lifetime can be monitored via a lifetime-dependent phase measurement for example. Simply monitoring the intensity, in the absence of an analyte-independent reference, can give rise to drift and instability due to light source and detector drift, changes in optical path, and drift due to degradation or leaching of the dye. These problems can be overcome to a large extent by employing ratiometric sensing which, as explained in section 4.2.2, involves taking the ratio between an analyte-dependent luminescence or excitation band and a second band which is analyte independent. Referenced sensing can also be achieved using techniques such as phase fluorometry and dual-luminophore referencing (DLR), which will be explored in more detail in section 4.5.3. This section will be divided into a review of recently reported intensity-based sensing and a discussion of lifetime-based sensing, including some recent examples.

4.5.2. Intensity-Based Sensing Fiber optic oxygen sensors based on two different Pt(II) complexes encapsulated in a fluorinated sol-gel matrix have been reported.241 A composite matrix of n-propyltrimethoxy-

Optical Chemical Sensors

silane and 3,3,3-trifluoropropyltrimethoxysilane was doped with either platinum tetrakis(pentafluorophenyl)porphine (PtTFPP) or platinum octaethylporphine (PtOEP). The higher oxygen sensitivity of metalloporphyrin complexes compared to ruthenium complexes has already been discussed in section 4.3, as also has the stability and enhanced oxygen permeability offered by fluoro sol-gel matrices. Both sensor materials demonstrated very high oxygen sensitivity, linear SV response, and short response times of ∼5 s. The disadvantages of intensity sensing have been addressed by the work of Tang et al.,242 where tailored solgel sensor arrays were employed to yield improved accuracy and precision in O2 detection with the use of artificial neural networks (ANN). Pin-printed arrays, where each element has a different O2 response profile, are produced. The tailored oxygen response is achieved by co-doping the sol-gel with different ratios of two ruthenium complexes which have very different oxygen sensitivities. The ANN was “trained” to identify the CCD images from the array. Overall, a 5-10fold improvement in accuracy and precision compared to a single-element sensor was achieved for quantifying O2 in unknown samples. Transparent nanostructured metal oxide oxygen sensor matrices were fabricated by embedding oxides such as aluminum oxide, silicon oxide, and zirconium oxide in PVA.243 Organometallic complexes of ruthenium or iridium were used as oxygen-sensitive dyes, which were incorporated in the nanostructured matrix. SV constants for these matrices were ∼100 times larger than for the same dyes incorporated in conventional polymer matrices. Furthermore, the materials were sterilizable by autoclavation and γ radiation, exhibited response times of <1 s, and were stable for up to 9 months. The superior oxygen sensor performance was attributed mainly to the nanoporosity and high total pore volume of the nanostructured matrices. The majority of both colorimetric and fluorometric pH indicators reported in the literature and in this review operate in the acidic or neutral pH range. Recently, luminescent pH indicators based on Schiff bases have been developed which respond in the pH range 7-12.244 The indicators chlorophenyliminopropenylaniline (CPIPA) and nitrophenyliminopropenylaniline (NPIPA), with pKa values of 10.30 and 8.80, respectively, have been immobilized in PVC. Photostability was superior to fluorescein for both dyes, and CPIPA in PVC matrix exhibited an enhanced quantum yield and reversible pH response over the basic pH range. Both dyes have convenient absorption and emission bands in the visible spectrum. A Li+ sensor for clinical applications, which is based on a polymer derivative of a novel lithium fluoroionophore, has been reported.245 The fluoroionophore is based on a tetramethyl blocking subunit of a 14-crown-4 as a Li+-selective binding site and 4-methylcoumarin as a luminophore. Ratiometric detection is enabled by the quenching of the main emission band by Li+ and the simultaneous appearance of a blue-shifted shoulder. The sensor exhibits good reversibility and immunity from interference for pH and the major biological interfering cations, K+, Ca2+, and Mg2+. Highly selective K+ sensing has been demonstrated using a fluoroionophore based on aza-18-crown-6 as K+ chelator and the fluorophore BODIPY (from Molecular Probes Inc.). The cavity size of the crown ether is chosen to be selective for K+ and discriminate against Na+ ions. By using novel substitution reactions of the BODIPY complex, the system

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was tailored such that the emission band shifts from ∼529 to 504 nm in the presence of K+, which facilitates ratiometric optical detection.246 Advances in array-based biosensing have been summarized in a recent review by Ligler.247 The main focus of this review is development of a portable, multianalyte luminescencebased array biosensor by the Naval Research Laboratories (NRL). The device uses planar waveguide optical coupling and PDMS microfluidics to allow simultaneous detection and quantification of multiple target analytes and multiple samples.248 The array-based biosensor has been used for antibody detection, for example, to detect food borne contaminants such as staphylococcal enterotoxin B and other toxins249 and the toxin deoxynivalenol, which is a common food contaminant. LODs of <1 ng/mL were achieved in each case. The system has also been used to detect bioterrorism agents such as anthrax.250

4.5.3. Lifetime-Based Sensing A primary disadvantage associated with the use of fluorescence intensity detection techniques is the susceptibility of sensors based on such techniques to signal fluctuations caused by changes in excitation source intensity, photobleaching of the fluorescent indicator complex, or changes in the concentration of the indicator within the sensing environment (due, for example, to leaching). A more robust fluorescence detection strategy involves measuring analyteinduced changes in the lifetime of the indicator compound, as this parameter is unaffected by the factors mentioned above. Sensors based on detection of fluorescence lifetime may be broadly divided into those that directly measure the lifetime and those that employ phase measurement techniques to indirectly monitor the lifetime. Systems that employ direct lifetime measurement techniques typically include a pulsed laser source and gated, or delayed, detection with a suitable photodetector, e.g., a photon-counting photomultiplier tube or intensified CCD camera. Such systems have been applied to detection of water in organic solvents,192 pH,251-253 and oxygen204,254 in recent years. However, the cost and complexity of systems capable of direct fluorescence lifetime measurements render them less attractive for development of low-cost, mass-producible systems. This has been an important driver in optical chemical sensor research in recent years, and while some efforts have been made to develop low-cost lifetime measurement systems,255 more effort has been applied to development of phase measurement instrumentation, which is well suited to development of low-cost sensor platforms. Perhaps the most common example of a luminescence-based sensor that utilizes phase measurement techniques is that of a phase fluorometric oxygen sensor.10,39,40,256-260 The concept of phase fluorometry is illustrated in Figure 9. The excitation source is modulated at a particular frequency, and the emitted fluorescence is similarly modulated but phase shifted relative to the excitation signal. Phase measurement electronics are used to measure this phase shift, which is related to the luminescence lifetime by eq 1

tan φ ) 2πfτ

(1)

where φ is the measured phase angle, f is the modulation frequency, and τ is the lifetime of the luminescent complex. This measurement concept can be implemented using lowcost phase detection electronics in conjunction with LED

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Figure 9. Principle of phase fluorometry.

sources and photodiode detectors, making it an extremely attractive candidate for development of commercial oxygen sensors by companies such as PreSens Precision Sensing GmbH [www.presens.de], YSI environmental [www.ysi.com], Gas Sensor Solutions [www.gss.ie], Ocean Optics [www.oceanoptics.com], and Interlab [www.interlab.es]. The technique has also recently been applied to development of optical humidity sensors.193,194 An important element in development of a phase fluorometric sensor is the use of an indicator complex that is sufficiently long lived so as to be accessible to the modulation frequency range provided by the low-cost sensor electronics, which typically requires that the lifetime be at least in the single microsecond regime. In the case of phase fluorometric oxygen sensors, indicators based on ruthenium or porphyrin dye complexes adequately fulfill this role, but for analytes such as carbon dioxide, pH, or chloride, the fluorescent indicators employed are not sufficiently long lived to be directly accessible to phase fluorometric detection strategies. In response to this, a number of strategies have been developed with the goal of rendering the intrinsically intensity-based measurements required for detection of such analytes compatible with phase fluorometry instrumentation. This would then allow for development of multianalyte sensor technology requiring only a single instrumentation platform. Fluorescence resonance energy transfer (FRET) has been exploited in the past to bring intensity measurements into the phase domain.208,261 However, the most predominant example of a technique that fulfils this goal is dual-lifetime referencing (DLR), also referred to as dual-luminophore referencing, which was first reported by Huber et al. in 2000262 and has been applied to detection of analytes such as CO2,195, 20943, 263 pH,264,265 nitrate,266 copper(II) ions,267 and chloride.199 As the name suggests, the DLR technique makes use of two luminescent indicators that are typically coimmobilized within a solid matrix. One of the indicators is analyte sensitive and possesses a short luminescence lifetime; the other is analyte insensitive and long lived. Changes in the fluorescence intensity ratio of the two compounds are reflected as changes in the measured phase angle (see Figure 10), thereby facilitating the use of phase detection instrumentation in the realization of any sensor that exploits this technique. Additionally, Borisov et al. recently reported the use of a modified DLR (m-DLR) technique that facilitates the simultaneous detection of two analytes.268 A related technique is that of gated phase-modulation (GPM) fluorometry,269 which is similar to DLR in that it makes use of two

Figure 10. Principle of dual-lifetime referencing (DLR). Changes in the amplitude of the indicator signal that are caused by changes in analyte concentration result in a modification of the measured phase angle, φm. This is shown above for two different indicator signal amplitudes.

luminophores with long and short lifetimes, but unlike DLR, square-wave modulation is employed in conjunction with gated detection in order to develop sensors for analytes such as pH.

5. Key Trends and Future Perspectives While the basic sensor transduction principles employed in optical chemical sensors are now well established and remain largely unchanged over the years, there have been significant recent advances in both the design of sensor platforms and strategies for performance enhancement. These developments, which are likely to drive future trends in the research and development of optical chemical sensors, include advances in nanomaterials, microfluidics, wireless networks, as well as optical materials and components. As in many other areas of scientific research, many developments in optical chemical sensors (OCS) are to be found at the micro-nano interface. The massive investment in nanotechnology across the world is leading to new materials and structures, the unique properties of which can have significant implications for the design and implementation of OCS. For example, the area of plasmonics, which is based primarily on the novel optical properties associated with localized surface plasmon resonance on metal nanoparticles and nanostructured surfaces, is attracting considerable research attention with conferences and a new journal dedicated to the topic. The key implications of plasmonics for optical chemical sensing are in nanoscale sensors (e.g., refractometric) and enhancements of spectroscopic sensors. For example, surface-enhanced Raman scattering (SERS) is an area that has been the subject of intensive research over a relatively short time frame for development of ultrasensitive diagnostic devices. A number of novel geometries/platforms based on metal nanoparticles/nanostructures have been developed recently,134,135 and these have resulted in improved sensitivity and sensor stability. Research in this area is likely to accelerate further in coming years. Similarly, the area of metal-enhanced fluorescence270-273 (MEF) is attracting considerable attention and likely to play a key role in sensitivity enhancement strategies over the coming years as the underlying principles of this phenomenon become better understood. Another area where the impact of nanotechnology is likely to be seen is in the use of nanowires; the use

Optical Chemical Sensors

of subwavelength nanowires as evanescent wave-based optical sensors was recently reported.274 Nanomaterials must be supported on a suitable platform which must often provide additional functionalities such as sample delivery and signal collection. Such platforms typically employ features in micrometer dimensions, and this highlights the increasingly important role that microsystems and microfluidics will play in the future development of OCS. More generally, OCS employing microfluidics (labon-a-chip; micro-total analysis systems (µTAS); Bio-MEMS) will see significant growth especially in the context of (bio)sensor arrays and the need for miniaturized systems. It is important to note here that such microfluidic systems may also enable the use of chemical sensors based on renewable reagents (using irreversible indicators which would otherwise not be suitable for sensors) by virtue of the low volumes involved. Such miniaturized systems are also critical in the field of wireless sensor networks, which is the subject of a separate review in this issue. This area is likely to have a profound impact on our society as a whole, potentially providing real-time information on environmental and healthrelated parameters in a spatially distributed context. Another area where developments in microfluidics are likely to see major application is in the fabrication of low-cost diagnostic platforms for widespread use, e.g., in the home but more significantly for application in low resource environments (global diagnostics). Advances in optical materials and platforms are likely to play an important role in future developments of OCS. For example, refractometric sensor platforms based on geometries such as microsphere275,276 or microring277 resonators and 2D photonic crystals278 show promise for development of compact, array-based chemical sensors, while a number of groups have emphasized and demonstrated the importance of integration of all aspects of the sensor platform (i.e., source, sample, detector) onto a single chip using, for example, CMOS technology.279-281 The use of quantum dots for optical sensing is a rapidly developing area. In particular, quantum dots are increasingly being used as high-brightness fluorescent labels in optical biosensing. These nanoscale colloidal semiconductors have many advantages over conventional dye labels, for example, very high quantum efficiencies, narrow, tunable emission bands, and enhanced stability. It is likely that ongoing improvements in synthesis and functionalization of quantum dots will lead to a new generation of luminescence-based optical sensors.282-285 Another recent innovation in fluorescence-based sensor research is the exploitation of fluorescence emission anisotropy in the design of highly efficient, ultrasensitive sensor platforms based on a range of optical elements and platforms.285-288 Such platforms can enhance sensitivity by a factor of up to 100 and will have significant implications for development of efficient, low-cost diagnostic platforms. In conclusion, the field of OCS, which is based on solid fundamental principles, has an exciting future as the need for rapid supply of measurement information increases and the convergence of disparate technologies provides significant opportunities for performance enhancement.

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Chem. Rev. 2008, 108, 423−461

423

Optical Biosensors Sergey M. Borisov† and Otto S. Wolfbeis* Institute of Analytical Chemistry, Chemo- and Biosensors, University of Regensburg, D-93040 Regensburg, Germany Received April 2, 2007

Contents 1. Introduction and Scope 2. General Remarks 2.1. Definition of Biosensors 2.2. Classification of Biosensors 2.2.1. Catalytic Biosensors 2.2.2. Affinity Biosensors 2.3. General Aspects of Signal Generation, Immobilization of Biomolecules, and Sample Handling 2.4. Frequently Used Spectroscopies and Internal Referencing 3. Enzymatic Biosensors 3.1. General Considerations 3.2. Enzymatic Glucose Biosensors 3.3. Other Enzyme-Based Biosensors 4. Immunosensors 4.1. General Remarks 4.2. Immunosensor Formats 4.2.1. Direct Immunosensors 4.2.2. Competitive Immunosensors 4.2.3. Sandwich Immunosensors 4.2.4. Displacement Immunosensors 4.2.5. Immunosensors Based on Binding Inhibition 4.2.6. Comparative Study on Immunosensor Formats 4.3. Preferred Optical Readout Formats in Immunosensing 4.3.1. Conventional Readout Formats 4.3.2. Evanescent Wave, Capillary, and Other Readouts 4.4. Immobilization of Antibodies on Sensor Surfaces and Nonspecific Protein Binding 4.5. Specific Examples of Immunosensors 4.5.1. Biosensors for Proteins and Antibodies 4.5.2. Biosensors for Toxins 4.5.3. Biosensors for Drugs 4.5.4. Biosensors for Bacteria Cells 4.5.5. Biosensors for Pesticides 4.5.6. Multianalyte Biosensors 5. Biosensors Based on Ligand−Receptor Interactions 5.1. Receptor-Based Biosensors for Saccharides and Glycoproteins

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* Corresponding author. E-mail: [email protected]. Tel.: (+49) (941) 943-4066. Fax: (+49) (941) 943-4064. † New address: Institute of Analytical Chemistry and Radiochemistry, University of Technology of Graz, Stremayrgasse 16, A-8010 Graz, Austria.

5.2. Receptor-Based Biosensors for Inorganic Ions 5.3. Receptor-Based Biosensors for Gaseous Species 5.4. Receptor-Based Biosensors for Toxins 6. Nucleic Acid Biosensors 6.1. Single DNA Sensors on Solid Supports and on Fiber-Optics 6.2. DNA Arrays 6.3. Molecular Beacons in DNA Sensors 6.4. Liposome-Based DNA Assays 6.5. Aptamer-Based DNA Sensing 7. Whole-Cell Biosensors 7.1. Catalytic Whole-Cell Biosensors 7.2. External Stimuli-Based Cellular Biosensors 7.3. Genetically Engineered Whole-Cell Biosensors 8. Solid Supports for Use in Optical Biosensors, and Other Methods of Immobilization 9. Outlook 10. List of Abbreviations and Acronyms 11. References

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1. Introduction and Scope Numerous kinds of biosensors do exist, but this chapter is confined to sensors and systems where the information is gathered by the measurement of photons (rather than electrons as in the case of electrodes). More specifically, it relates to sensors based on the measurement of absorbance, reflectance, or fluorescence emissions that occur in the ultraviolet (UV), visible, or near-infrared (NIR). It does not cover sensors based on infrared or Raman spectroscopy, nor those based on surface plasmon resonance. Molecular imprintsswhile very promisingsare not covered either since they do not match the definition of a biosensor (which asks for a biological recognition element to be at work). Fluorescence is by far the method most often applied and comes in a variety of schemes. Parameters that are being measured in such sensors include intensity, decay time, anisotropy, quenching efficiency, luminescence energy transfer, and the like. Optical layouts include plain sensor foils (stripes) and also waveguide optical systems, capillary sensors, and arrays. Chemical sensors and biosensors do not have separation capabilities unless coupled to respective additional devices that, however, make the system more complex, require larger instrumental effort (and power consumption!), and prevent sensing to be combined with imaging. Hence, specificity can only be based on selective (bio)molecular recognition. To achieve this goal, use is made of more or less specific biorecognition elements such as enzymes, antibodies, oligonucleotides, or even whole cells and tissues. The variety

10.1021/cr068105t CCC: $71.00 © 2008 American Chemical Society Published on Web 01/30/2008

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Sergey M. Borisov, born in 1978, at present is a postdoctoral fellow at the Institute of Analytical Chemistry and Radiochemistry at the University of Technology in Graz, Austria. He recived a Ph.D. degree in chemistry from the Herzen State Pedagogical University (St. Petersburg, Russia) in 2003 and worked as a postdoctoral fellow at the University of Regensburg before he assumed his present position in Nov. 2006. His research interest is in the chemistry of porphyrins, in the application of fluorescent probes to biosensing schemes, and in the use of (fluorescent) micro- and nanomaterials in bioanalytical methods. He has authored several articles in recent years, mainly in analytical and material science journals.

Borisov and Wolfbeis

solid-phase-based diagnostic devices such as those for glucose, pregnancy markers, or cardiac markers. They are better referred to as test strips. Others (mainly bioorganic chemists) often refer to molecular bioprobes as biosensors. However, true biosensors are solid state, not certain molecules, give a reading after having been contacted with the sample to be analyzed, and do not require the addition of reagent(s). It is noted at this early stage of the review that the world still does not have a fully reversible glucose sensor for in vivo use over >1 months, which would be a great relief to the 4-5% of the population suffering from various forms of diabetes and that would enable the construction of an artificial pancreas. Definitions of biosensors have been given but are diverse.28-30 However, all include the use of a biological component such as an enzyme, an antibody, a polynucleic acid, or even whole cells or tissue slices. In other words, a pH electrode capable of sensing the pH of blood is not a biosensor by all current definitions, as is xenon gas that can be used to probe the structure and dynamics of a protein.31 Certain authors confuse the terms sensor and probe; we are referring such authors to the homepage of the world’s largest manufacturer of bioprobes (www.probes.com), which never would refer to its many bioprobes as “sensors”. Over time, this has led to the undesirable situation that electronic searches for literature on biosensors result in two sets of data. The first (still larger one) is on true sensors of all kind (electrochemical, fluorescence, piezo, thermal, surface plasmon resonance (SPR), reflectometric, chemo/ bio-luminescent, IR, and the like). The second set of data is on (mainly optical) molecular probes whichsin a wrong fashionsare referred to as biosensors.

2.2. Classification of Biosensors

Otto S. Wolfbeis, born in 1947, is a Full Professor of Analytical and Interface Chemistry at the University of Regensburg, Germany. His research interests are in optical chemical sensing and biosensing, in the design of novel schemes in analytical fluorescence spectroscopy, in fluorescent probes, beads, and labels, in biosensors based on thin gold films and molecular imprints, and in the design of advanced materials for use in (bio)chemical sensing. He has authored numerous papers on optical (fiber) chemical sensors and fluorescent probes, has edited a book on Fiber Optic Chemical Sensors and Biosensors, acts as the editor of the Springer Series on Methods and Applications of Fluorescence, is the Editor-inChief of Microchimica Acta, is a member of the board of various journals including Angewandte Chemie, and acts as the chairman of the Permanent Steering Committee of the biannual conference on Methods and Applications of Fluorescence (MAF) since 1989. For details, see www.wolfbeis.de.

of approaches that have been made in the past makes it rather difficult to sort biosensors systematically. Excellent books and reviews cover the first few decades of research on biosensors,1-27 but none of those describes the state of the art as comprehensively as the one presented here, since it covers the work from the early stages to the state as of early 2007.

2. General Remarks 2.1. Definition of Biosensors Unfortunately, the terminology on biosensors is not systematic. Medical doctors tend to refer to “biosensors” as

The biosensors described in this review can be divided into two kinds of groups, viz. biosensors of the catalytic type and biosensors of the affinity type. Their features are briefly discussed in the following.

2.2.1. Catalytic Biosensors These make use of biocomponents capable of recognizing (bio)chemical species and transforming them into a product through a chemical reaction. This type of biosensor is represented mostly by enzymatic biosensors, which make use of specific enzymes or their combinations. Many wholecell biosensors also rely on biocatalytic reactions. More recently, catalytically active polynucleotides (DNAzymes) have been used as well. This type of biosensor also includes biosensors based on measurement of the rate of inhibition of a catalytic reaction by an inhibitor such as a heavy metal ion or a pesticide. Catalytic whole-cell sensors often are employed to sense sum parameters such as toxicity, antibiotic activity, or cell viability.

2.2.2. Affinity Biosensors These make use of the specific capabilities of an analyte to bind to a biorecognition element. This group can be further divided into immunosensors (which rely on specific interactions between an antibody and an antigen), nucleic acid biosensors (which make use of the affinity between complementary oligonucleotides), and biosensors based on interactions between an analyte (ligand) and a biological receptor.

Optical Biosensors

Some whole-cell biosensors act as recognition elements responding to (trigger) substances by expressing a specific gene.

2.3. General Aspects of Signal Generation, Immobilization of Biomolecules, and Sample Handling The most usual format of a biosensor is that of a biological compound immobilized on the surface of a transducer. The function of the latter is to gather the analytical information when in contact with the sample and to convert it into an electrical signal. Optical transducers respond to an analyte by undergoing a change in their optical properties, such as absorption, reflectance, emission, or a change in an interferometric pattern. Signal changes are recorded by a photodetector and, thus, transformed into an electrical signal. The development of appropriate (and stable) materials probably is more of a challenge in biosensor development than the development of appropriate spectroscopies. In all kinds of biosensors, recognition is accomplished by a biomolecule. In the overwhelming majority of biosensors, this biomolecule is immobilized on the surface of the sensor. Immobilization serves one or more of the following purposes: (a) It enables continuous sensing of analytes in flowing systems such as blood, bioreactor fluids, or water samples. (b) The biomolecule is “added” to the sample in welldefined quantity. (c) The biosensor becomes reusable or regenerable. The stability of immobilized biomolecules is a serious issue. It is noted that many articles on biosensors do not consider aspects of long-term stability in a proper way. Much more often than chemical sensors, biosensors have been combined with (micro)fluidic devices such as (micro)flow injection analyzers or lab-on-a-chip devices. Optical biosensors are particularly useful in the case of the latter, where voltages of up to several kV are applied that may disturb (or make impossible) electrochemical detection. Valcarcel and Luque de Castro32 have reviewed the state of flow-through “biosensors”, which, however, often are based on flow-injection and corresponding detectors.

2.4. Frequently Used Spectroscopies and Internal Referencing There are two main types of optical biosensors: The first exploits any changes that can occur in the intrinsic optical property of the biomolecule as a result of its interaction with the target analyte. Such changes can occur in absorbance, emission, polarization, or luminescence decay time of a receptor. Such sensors are not numerous because their sensitivity is usually low, and because many effects occur in the deep UV where spectroscopy has academic merits but is difficult to implement when it comes to analyzing complex (such as environmental or clinical) samples. An additional challenge when using intrinsic biosensors consists of the separation of the shortwave signal from background fluorescence (or absorbance). Enzymes using FAD as a coenzyme are examples of more longwave absorbing receptors that undergo intrinsic spectral changes on binding a ligand (during catalytic conversion), as are some cytochromes and hemoglobin. The second type of biosensor is making use of optical labels and probes of various kinds. This requires the

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biomolecule to be covalently labeled (an extra step) but enables the analytical wavelength(s) to be shifted to almost any desired value. Moreover, luminescence decay times and anisotropy can be adjusted to desired values, and effects such as dynamic or static quenching can be exploited in a more systematic manner. Not surprisingly, luminescent labels are widely used for this purpose. In recent years, the use of luminescent nanoparticles has strongly increased. Absorptiometry and reflectometry are still the most widely used methods, both in solution assays and in test strips. Absorptiometry is well-established (hardly any lab does not have a photometer) and has the unique merit of being selfreferenced (i.e., the intensity of the signal measured is always referenced to the intensity of the incoming light beam in the case of two-path photometers as used in cuvette assays or microtiter plates). Surprisingly enough, fluorescence intensity is by far the most often used analytical parameter when it comes to biosensing. According to Parker,33 luminescence intensity (I) is proportional to the concentration of an analyte present,

I ) I0clφfk where I0 is the intensity of the exciting (laser) beam,  is the molar absorption coefficient (molar absorbance) of the fluorescent probe or label, c is its concentration, l is the length of penetration, φf is its quantum yield, and k is a geometrical factor that accounts for the geometry of the optical system. This linear relationship between measured fluorescence (phosphorescence) intensity and I0 is valid only for solutions of low molar absorbance. Fluorescence intensity at a single wavelength is not referenced and obviously depends on numerous variables (and can be compromised by drifts in the photodetection system). Ratiometric (two-wavelength) measurements have, therefore, become quite widespread.34-40 This either requires addition of an inert reference fluorophore or the application of a FRET system (i.e., a donor dye and an acceptor dye). FRET systems have often been employed in immunosensors,41-44 nucleic acid sensors,40,45-47 and those based on ligand-receptor interactions.48-56 Another method for self-referencing consists in the measurement of luminescence decay time.52,57 Since the measurement of decay times in the order of a few nanoseconds (or even picoseconds) so far has required complex and expensive instrumentation (this has changed in recent years, though), labels and indicators were employed with decay times in the order of µs and ms. Decay time-based sensing is widely employed in optical oxygen sensing and in enzyme sensing based on oxygen transduction.58-60 Measurement of decay time also was reported for a fluorescent hydrogen peroxide transducer (a europium(III) complex) for use in a glucose biosensor.61 A final self-referenced method is based on measurement of fluorescence polarization,62-64 which also is independent of various variables (such as the degree of labeling, photobleaching, quenching, and solvent effects) that make other methods prone to errors. Table 1 summarizes the more important (self-referenced) methods of read-out in luminescence and their respective merits. Measurement of intensity is most common because they are easily performed, routine instrumentation is available, and one label is required only. Ratiometric (2-λ) methods are more reliable but require the availability of appropriate probes ands labels. Dual-lifetime referencing (DLR) is quite powerful, too, but requires the presence of a reference dye

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Table 1. Fluorescent Schemes and Their Ability to Compensate for Interferences of Various Origin. (++): Well Compensated for; (+) Partially Compensated for; (-) Not Compensated for. These Are General Statements Only That May Be Different in Particular Situations parameter interference resulting from

intensity

2-λa

ref dye

FRET

lifetime

optical components (e.g., filters) instrumental drift (light source, detector) optical misalignment background fluorescence of sample light scatter by sensor material/ sample intrinsic color of sample dye leaching/bleaching temperature inhomogeneous dye loading

-

+ ++ + + +e ++

+b ++ + + -

+ -c + +e 0

++ ++ ++ +d ++ ++ + ++

anisotropy + ++ + ++ ++ ++

DLR + ++ ++ ++ + +e +

a Using a dual-wavelength label or probe. b Only compensated for if detected with the same set of optical components. c Not used in fiber-optic sensors. d Well-compensated only in case of time domain for probes having decay times in the µs or ms range. e Assuming both dyes to display the same temperature-dependence of their spectral properties.

with long decay time, so it is more complicated. DLR has not been applied to biosensing so far. The measurement of decay time (“lifetime”) is superior to measurement of intensity in many respects. In the case of affinity sensors, probes are needed whose decay time (that usually is not strongly affected by binding) changes upon biomolecular interactions. It has been demonstrated, for example, that the fluorescence lifetime of certain fluorescent labels (supposed to be inert in terms of changes of decay time) is a useful parameter to detect affinity binding between biotin and streptavidin and between biotinylated bovine serum albumin and streptavidin.65 Lifetime also can be determined in FRET, preferably if a long-lived donor dye is used. Refractometry also is self-referenced and has been used in immunosensors.66-68 Less common spectroscopic techniques such as reflectometric interference spectroscopy,69 optical waveguide lightmode spectroscopy, 70 supercritical angle fluorescence,71 and light scattering72 also shall only be mentioned here.

3. Enzymatic Biosensors 3.1. General Considerations Determination of such analytes as glucose, lactate, urea, ethanol, phenols, pesticides, and many others is of high significance in clinical medicine, food and environmental analysis, and bioprocess monitoring. The lack of indicators that give changes in color or luminescence at room temperature without addition of (aggressive) reagents and at nearneutral pH, in reasonably short time and in a fully reversible way, has made researchers look for alternatives. Enzymes catalyze reactions with a high degree of specificity, and the products of these reactions (or of reactants consumed) are detected directly if colored or luminescent, or by using optical transducers. The steady-state concentration of detectable species is, thus, related to the concentration of the analyte. Some enzymatic reactions require the presence of other specific reactants called coenzymes, e.g., nicotinamide adenine dinucleotide or flavine mononucleotide, which change their optical properties during the reaction. A cross section of the typical enzymatic biosensor is shown in Figure 1. An indicator layer (often sensitive to oxygen or pH) is spread over a transparent inert support, usually a polyester film. An indicator dye is either directly dissolved in a polymer matrix or, alternatively, covalently immobilized or physically adsorbed on a surface of microbeads, which

Figure 1. Cross section of a fiber-optic enzymatic biosensor. The analyte (substrate) enters the enzyme layer where it is converted into products. The indicator (sensing) layer consists of an indicator dye in a polymer layer and registers the formation of reaction products or the consumption of coreactants such as oxygen. The transparent support is inert and used only to facilitate manufacturing. It may as well be omitted. Exc and Em symbolize the paths of exciting and emitted light, respectively.

then are dispersed in the matrix polymer. The indicator layer is responsible for sensing of either cosubstrates consumed or of products produced during the enzymatic reaction. Enzyme(s) can be chemically immobilized onto the surface of a polymer membrane (e.g., cellulose, nylon, or inorganic porous glass) or physically entrapped into a polymer network, e.g., sol-gels, hydrogels, or Langmuir-Blodgett films. To avoid leaching of the enzyme, it is often cross-linked to bovine serum albumin via glutaraldehyde linkers. Alternatively, preactivated membranes may be used. When the analyte (the substrate) diffuses into the enzyme layer, it is converted into products. The indicator layer registers the formation of reaction products or the consumption of coreactants such as oxygen. In Figure 1, the sensor “sandwich” is mounted on the tip of an optical fiber that transmits excitation light from a light source to the sensor foil and emitted (reflected) light back to a photodetector. However, the majority of biosensors are not based on fiberoptics. Optical sensors that exploit chemi- and bioluminescent reactions are usually simpler because no indicator layer is required. The chemical species generated during an enzymatic process are involved in subsequent reactions that result in the production of light. In this case, other substrates (“reagents”) are needed along with the sample solution. Most chemi- and bioluminescent reactions are catalyzed by enzymes that have to be co-immobilized in the enzyme layer. Biosensors that make use of the intrinsic optical properties of the enzyme do not require optical transducers and, thus, usually include an enzyme layer placed

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on a planar support or at the tip of an optical fiber, preferably in a hydrogel.

3.2. Enzymatic Glucose Biosensors Not surprisingly, this is by far the most often investigated type of biosensor. Those based on the use of glucose oxidase (GOx) function on the basis of the following reactions: GOx

D-glucose + O2 98 D-gluconolactone + H2O2 (1) D-gluconolactone + H2O 98 D-gluconate + H+ (2) The concentration of glucose thus can be related to (a) the amount of oxygen consumed,35,58,59,73-88 (b) the amount of hydrogen peroxide produced,39,61,89-102 or (c) the decrease in pH due to the conversion of D-gluconolactone to Dgluconic acid.103-105 The above equations indicate that the response of such a sensor depends on a number of variables (notwithstanding the effects of temperature and diffusion). The first is pH. If pH transduction is used, the initial pH and the buffer capacity of the sample will govern the shape and the relative signal change. pH also affects enzyme activity. The second is oxygen. Depending on how its concentration is related to that of glucose, different shapes of the response curve and different signal changes will be observed, as can be seen in Figure 2. If the sample is anaerobic (i.e., does not contain any oxygen), no signal change will be detectable. If oxygen is present in large excess, the concentration of oxygen is low, and diffusional processes are fast, hardly any signal changes will be detectable once the steady-state equilibrium is reached. It also needs to be reminded that the shapes are quite different for standing samples, stirred samples, and flowing samples. Finally, the quantity (more precisely, the activity) of immobilized GOx will strongly affect the signal change and the response time. This is shown in Figure 2 for flowing samples. The sensor is first equilibrated with a buffer solution saturated with air. The flow of buffer is then replaced by a flowing sample at time t1. At time t2, the sample is replaced by a flow of buffer again. Various curves (A-E) are obtained depending on the levels of oxygen and glucose in the sample: (A) No oxygen and no glucose in the sample; the shape of the response is mainly determined by the rate of the diffusion of oxygen out of the sensor membrane into the sample flow. The same signal level is reached (even faster) if the sample contains no oxygen but a relatively large concentration of glucose. (B) Response to a sample where [O2] , [glucose]; all oxygen in the sensor is quickly consumed as a result of enzymatic oxidation and of diffusion. (C) Air-saturated sample where [O2] > [glucose]; the shape is mainly determined by the rate of the enzyme-catalyzed oxidation of glucose. (D) Sample where [O2] . [glucose]; the steady-state signal is smaller than that in (C). (E) Sample without glucose where the pO2 is lower than at air saturation. These are exemplary plots; the shapes and steady-state intensities also depend on the activity of the enzyme, the flow rates, and the thicknesses of the various layers (and, thus, on the oxygen storage capacity). Note that the shapes for flowing samples are quite different from those obtained

Figure 2. Typical signal shapes that can be obtained if a glucose sensor based on immobilized glucose oxidase and using an oxygen sensor as the transducer (in contact with air-saturated buffer) is exposed to flowing samples containing various levels of oxygen and glucose, respectively, and then again to air-saturated buffer. See the text for an explanation of shapes and signal changes.

with standing samples and that even these can differ depending on whether they are stirred or not. Detection of glucose via the quantity of hydrogen peroxide formed appears to be the most attractive approach since it works at virtually zero background, even though it also is affected by the initial pO2 in the sample. Optical continuous sensors for H2O2 are scarce, however. The group of Luebbers106 probably were the first to describe a glucose sensor based on transduction via oxygen, which acts as a dynamic quencher of the luminescence of certain indicator dyes. The sensors consisted of an oxygen sensor (using pyrenebutyric acid as the oxygen probe) onto which GOx was deposited as a thin layer. The sensor reported glucose in physiological concentrations. The temperature dependence of the biosensor was studied in some detail. Temperature is known to exert an effect on various parameters including the rate of diffusion, the activity of the enzyme, the efficiency of quenching of the indicator by oxygen, and the quantum yield of the fluorophore used. In essence, a reduced analytical range and a steeper slope of the response curve toward glucose is observed. Diphenylanthracene (DPA) was used as a probe for oxygen in a sol-gel based glucose biosensor. The sensing material was obtained107 by controlled hydrolytic polycondensation of tetraethoxysilane (TEOS) to give a fairly inert inorganic glassy matrix whose porosity and size of pore network can be varied by polymerization conditions. Both DPA and GOx were entrapped into the sol-gel. Because the material has no absorption in the near UV and visible, it is well-suited for fabrication of optical sensor membranes. Enzymes in sol-gels can be substantially stabilized by addition of polycations.108 Other biosensors based on oxygen transduction made use of polyaromatic hydrocarbons such as pyrene, decacyclene, and their derivatives, which were dissolved in silicone.73-75,77 Following their discovery as probes for oxygen in 1986,109 ruthenium(II) complexes with ligands such as bipyridyl (Rubipy), 1,10-phenanthroline (Ru-phen), and 4,7-diphenyl1,10-phenanthroline (Ru-dpp) rapidly replaced the polycyclic aromatic hydrocarbons. They possess visible absorption, relatively long decay times (0.6-6 µs), and good photostability and, therefore, are widely used oxygen probes.35,76,78,80-88 The probe Ru-dpp is a preferred indicator because of its good brightness (Bs; defined as the product of quantum yield and the molar absorption coefficient at the excitation wavelength), which is 10 500 M-1 cm-1 at 465 nm excitation.110 In being cationic, ruthenium probes can be adsorbed

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onto silica gel beads (which are negatively charged at pH 7) and then be dispersed in silicone, which results in an good sensitivity to oxygen. In addition, this material is highly scattering, which increases the efficiency of collection of fluorescence. Platinum(II) and palladium(II) porphyrins represent another group of viable luminescent oxygen indicators because of their high chemical and photochemical stability, large Stokes’ shifts, good brightness, and long luminescence lifetimes. They also are often used in pressure-sensitive paints. Papkovsky58 used a phosphorescent platinum(II) complex with octaethylporphyrin (PtOEP) dissolved in polystyrene as the oxygen transducer for the glucose sensor. Luminescence intensity and decay times were measured. The sensor was applied to the determination of 0.05-1.2 mM of glucose with a limit of detection of 0.05 mM. Recently, glucose was dually sensed by immobilizing (a) a europium probe acting as a reporter for hydrogen peroxide; (b) the iridium-trisbipyridine complex as a reporter for oxygen; and (c) glucose oxidase in a hydrogel membrane.111 This sensor measures the hydrogen peroxide formed (without any background) and can compensate for variations in oxygen partial pressure in the sample, which has a strong effect on the shape of the response function (see Figure 2). Miniaturized glucose sensors are particularly attractive for a number of clinical applications, including measurements of glucose in extremely small volumes or monitoring of localized events where high spatial resolution is desired. Microsensors also are attractive because they produce less injury to patients. Rosenzweig and Kopelman82,83 designed a fiber-optic glucose microsensor in which a ruthenium oxygen probe and GOx were incorporated into an acrylamide polymer covalently attached to the surface of an optical fiber (of an outer diameter between 2 and 100 µm). The analytical range of the sensor was rather high (0.7-10 mM), but the detectable quantity of glucose was very small because of the small sample volumes needed. The group of Klimant reported on a fiber-optic flow-through biosensor for online monitoring of glucose.112 A microdialysis membrane in a Tygon tubing contained a fiber-optic sensor composed of immobilized glucose oxidase and an oxygen transducer layer, and a reference oxygen sensor was used to compensate for interfering effects. The authors have also demonstrated outstanding selectivity of the sensor, which makes use of an oxygen optode as a transducer.113 No interference was observed from ascorbic acid, acetylsalicylate, uric acid, mannitol, and dopamine in concentrations exceeding physiological levels by several folds. Measurement of glucose in humans via a sensing catheter was demonstrated. Xu et al.35 prepared luminescent probes that were encapsulated into nanoparticles to give so-called PEBBLE sensors designed for intracellular glucose imaging. The polyacrylamide nanoparticles of 45 nm diameter incorporate GOx, the oxygen indicator (a sulfonated Ru-dpp derivative), and an oxygen-insensitive fluorescent dye, Oregon Green 488dextrane, that is used as a reference for the purpose of ratiometric intensity measurements. The small size and inert matrix of these sensors allows them to be inserted into living cells with minimal physical and chemical perturbations of their biological functions. Because glucose biosensors based on oxygen transducers measure the consumption of oxygen during the enzymatic reaction, the response of such sensors to glucose is actually dependent on concentration of oxygen in the analyzed

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medium. This is shown in Figure 2 for the two extremes of ratios of concentrations of glucose and oxygen. To overcome problems associated with variable oxygen supply, oxygen should be present in large excess (compared to the amount consumed) or its concentration should be kept constant (which is not always easy to achieve) or known. Dual biosensors have, therefore, been developed and represent one possible solution to the above problem.80,85 Such sensors do contain both an oxygen-sensitive and a glucose-sensitive element located in the proximity, e.g., on a distal end of an imaging fiber.80 The glucose-sensitive element is prepared by covering the oxygen sensor with an enzyme layer. Wolfbeis et al.85 tested three different combinations of oxygen transducer and sol-gel immobilized GOx. In the first, GOx was sandwiched between a sol-gel layer doped with Ru-dpp and a second sol-gel layer composed of pure sol-gel. Such configuration provided the highest enzyme activity and the largest dynamic range (0.1-15 mM) but suffered from a distinct decrease in sensitivity upon prolonged use. In the second, which provided the fastest response time (t90 ) 50 s), a sol-gel layer doped with Ru-dpp was covered with sol-gel-entrapped GOx. In the third sensor type, both the oxygen-sensitive sol-gel powder and the sol-gel powder containing GOx were incorporated into a single sol-gel phase. Such a sensor type provided the best operational lifetime. The authors also have derived equations that describe how the effect of varying oxygen supply can be compensated for by making use of two sensors, one sensitive to oxygen only and the other sensitive to both oxygen and glucose. Since both the decay time of the luminescence of oxygen indicators (especially for ruthenium(II) complexes)114,115 and the quenching by oxygen itself116 are highly temperaturedependent, the performance of all biosensors based on oxygen transducers also is influenced by temperature. The temperature dependence of such sensors may be compensated for by making use of dual sensors for oxygen and temperature,116,117 but so far this problem has not been addressed for glucose sensors, which, therefore, need to be thermostatted. This option has not been applied to glucose sensors. The concentration of glucose may also be related to the amount of protons produced in reaction 2; however, only few optical glucose biosensors made use of pH transducers. The fluorescence of the pH probe 8-hydroxypyrene-1,3,6trisulfonate (having a pKa ≈ 7.3) contained in a protonpermeable hydrogel served as a signal to monitor pH changes during enzymatic reaction.103 The limit of detection (LOD) for glucose was 0.1 mM. Polyaniline was found to exhibit pH-sensitive spectra and, thus, was used itself as a pH transducer.105 The enzymatic reaction can be monitored at 550-650 nm (where the absorbance decreases) or at 700900 nm (where the absorbance increases). The LOD for glucose was 1 mM. An interesting approach was made by McCurley.104 A pHinsensitive fluorophore linked to cadaverine was incorporated, along with GOx, into a cross-linked acrylamide-based hydrogel placed at the end of an optical fiber. The amine moiety of cadaverine is responsible for the pH-dependent swelling of the hydrogel. When the volume of the hydrogel increases, a decrease in fluorescent intensity is observed because the total quantity of fluorophore remains constant. The sensor was operative in the range from 0 to 1.6 mM of glucose.

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Finally, the amount of hydrogen peroxide produced during the enzymatic reaction can be related to the concentration of glucose. Most “sensors” rely on irreversible chromogenic reactions of hydrogen peroxide. The H2O2 transducer can be considered a biosensor itself because it makes use of the oxidation of luminol (5-amino-2,3-dihydrophthalazine-1,4dione) catalyzed by horseradish peroxidase (POx) as shown in eq 3: POx

luminol + 2H2O2 + OH- 98 3-aminophthalate + N2 + 3H2O + hν (3) The intensity of chemiluminescence (peaking at 430 nm) is proportional to the concentration of hydrogen peroxide and, in the case of the glucose biosensor, to the concentration of glucose. The method was pioneered by Freeman and Seitz,118 who immobilized POx in a polyacrylamide gel to monitor H2O2 in concentrations as low as 1 µM. The chemiluminescent reaction was widely used for determination of hydrogen peroxide93,119-121 and glucose.39,90,91,93 The sensors typically operate in the flow injection mode where GOx and POx are immobilized in a polymer membrane immersed into a solution containing luminol and sample. The problem to overcome is a mismatch between the optimal pH needed for enzymatic oxidation of glucose (neutral pH) and for enzymatic oxidation of luminol (pH ≈ 9). While an intermediate pH may be used, other possible solutions include the use of cetyltrimethylammonium bromide micelles to incorporate luminol and HPOx90 or the use of an internal solution of POx and luminol, located close to a membrane containing immobilized GOx.93 Similar to the oxidation of luminol, POx also catalyzes the reaction of H2O2 with other substrates such as homovanillic acid122 and Amplex Red.123 The products of oxidation are highly fluorescent species whose intensity can be monitored. Unlike in chemiluminescent “sensors”, the signal is not transient in these cases. Heo and Crooks102 used the POx-Amplex Red system for simultaneous determination of glucose and galactose in a microfluidic array biosensor. The enzymes (GOx and POx or galactose oxidase and POx) were entrapped in hydrogel micropatches where they show good storage stability. Amplex Red was added to the analyte solution, which was pumped over the surface of the sensor. The fluorescence of resorufin, the product of the oxidation of Amplex Red, was imaged via a conventional charge-coupled device camera. Glucose was determined in the range of 1-5 mM. By using specific enzymes located in different micropatches, several analytes can be determined simultaneously, as was demonstrated for the sensing of glucose and galactose mixtures. Production of resorufin also was monitored in a biosensor for superoxide ion, which makes use of superoxide dismutase and POx.124 Luminol also may be electrochemically oxidized by hydrogen peroxide, a reaction that does not require the enzyme POx and gives strong electrochemiluminescence (ECL). In this case, the polymer membrane containing immobilized GOx is placed on a carbon electrode,92,94,96 and the intensity of ECL is monitored from the other side. A sol-gel containing the enzyme also was coated on the surface of an electrode,99 and ECL was measured. Alternatively, the enzyme may be immobilized in a ceramic-carbon composite material.100 This graphite-containing sol-gel material was placed in a glass tube and served as an electrode to generate the ECL of luminol.

Several other kinds of H2O2 transducers were reported for use in glucose biosensors. Thus, a mixture of titanium(IV) ion and a pyridylazophenol dye was found to produce a reddish-purple product.89 Formation of a colored adduct with a dinuclear iron(III) complex was used to quantify H2O2 and glucose.95 The colored form of Prussian Blue was formed from the colorless one (Prussian White) upon oxidation by hydrogen peroxide.97,98 Wolfbeis et al.61,101 introduced a novel hydrogen peroxide transducer, which is based on the luminescent europium(III) tetracycline complex (EuTc).125 The probe is excitable by visible light and responds to H2O2 by an ∼15-fold increase in luminescence intensity. Unlike in previous methods, the determination of H2O2 does not require the addition of POx, and the reaction is fully reversible although rather slow in both directions (∼10 min). Moreover, the transducer operates at neutral pH. The large Stokes’ shift of ∼200 nm and the long-lived emission (with decay times in the microsecond time domain) enable the time-resolved suppression of fluorescent species. The probe immobilized into a hydrogel was successfully used for sensing101 and time-resolved imaging61 of glucose. Trettnak and Wolfbeis126 exploited the intrinsic fluorescence of GOx, which contains the fluorophore flavine adenine dinucleotide (FAD). FAD displays visible absorption (∼380-450 nm) and weak emission at ∼530 nm. The optical properties of the enzyme are slightly different for the reduced form (FADH2) that is produced during reaction with glucose and, therefore, could be used for analytical purposes. The sensor shows full reversibility (FADH2 is back-oxidized by molecular oxygen), but the analytical range is narrow (0.50.8 mM). The same authors used the intrinsic fluorescence of lactate mono-oxygenase for determination of lactate.127 Similarly, Chudobova et al.128 used the intrinsic absorption of GOx to obtain the sensor with a wider analytical range (1-10 mM) and limit of detection (LOD) of 2 mM. As expected, the LOD was significantly lower (0.5 mM) when the sample was deoxygenated. Sierra et al. reported on the use of the intrinisc fluorescence of GOx for the determination of glucose in serum.129 The green fluorescence of FAD is strongly quenched by serum proteins so that high concentrations of enzyme are required. Therefore, the authors suggested to exploit the UV fluorescence of the protein part of GOx, which peaks at 334 nm at λexc ) 224 nm. Even though such a biosensor is suitable, in principle, for the determination of glucose in the range from 0.5 to 20 mM, there are substantial drawbacks that include the lack of affordable (semiconductor-based) excitation light sources for 224 nm and interferences by other luminescent species. De Marcos et al.130 have investigated sensors based on intrinsic fluorescence of GOx immobilized on different polymer supports and in polymer matrixes, with respect to sensitivity, leaching of the enzyme, and sensor shelf life. The best results were achieved when GOx was immobilized on photopolymerized polyacrylamide. The sensor polymer films had a lifetime of over 2 months and adequate analytical characteristics. The linear range was between 1.67 and 11 mM of glucose. In an attempt to shift analytical wavelengths into the visible, De Marcos et al.131 have labeled GOx with fluorescein and found an increase in fluorescence intensity in the presence of glucose, probably a result of an inner filter effect. In fact, the absorption spectrum of GOx-bound FAD (but not of FADH2) overlaps that of fluorescein. Consequently,

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when FAD is reduced to FADH2 during enzymatic action, fluorescence is enhanced (λexc ) 492 nm). No such enhancement is observed if GOx is labeled with Cy-5 or Texas Red. For fluorescein-labeled GOx, the linear response is from 0.55 to 5.5 mM of glucose. The labeled GOx was entrapped into polyacrylamide gel, and the sensor was used in the flow injection mode.132 Another optical biosensor for glucose makes use of the intrinsic absorption properties of horseradish peroxidase (POx), which undergo spectral changes upon binding of H2O2.133 Both GOx and POx were entrapped in a polyacrylamide gel matrix. When glucose is present, H2O2 is produced and reversibly bound by POx. The intermediate species produced during enzymatic activity display different absorption spectra between 400 and 450 nm. The sensor has a linear response between 1.5 and 300 µM of glucose. The intrinsic optical properties of enzymes also were used for determination of nitrate (absorbance of nitrate reductase)134 and nitrite ions (absorbance of cytochrome cd1 nitrite reductase)135 as well as for sensing ethanol,136 pyruvate,137 and lactate,138 using intrinsic fluorescence of alcohol dehydrogenase, lactate dehydrogenase, and lactate oxidase, respectively. The characteristics of various glucose biosensors based on the use of GOx are summarized in Table 2 along with those for other enzyme-based sensors.

3.3. Other Enzyme-Based Biosensors A number of biosensors were designed by analogy to GOxbased glucose sensors by making use of other oxidases. Enzymatic oxidation of glutamate, for example, can be described by eq 4, L-glutamate

GlOx

+ H2O + O2 + H+ 98 2-oxoglutarate + NH4+ + H2O2 (4)

where GlOx stands for glutamate oxidase. Thus, biosensors for glutamate can be based on an oxygen transducer (e.g., decacyclene in silicone)77 or on detection of H2O2 by chemiluminescence (via the GlOx/POx system).139,140 It was found that peroxidase from Arthromyces ramosus produced a 100 times stronger luminescence signal than the commonly used POx from horseradish. A glutamate biosensor based on an ammonia transducer also was reported.141 The enzyme glutaminase (GLase) was used in a glutamine biosensor in which glutamine was converted into glutamate according to eq 5, GLase

glutamin + H2O 98 glutamate + NH3

oxidized by molecular oxygen. The enzymes were immobilized on an eggshell membrane and showed a remarkable long-term stability there. When stored at room temperature over a period of 6 months, the sensor retained >95% of its initial activity. When the enzymes were immobilized in plasticized poly(vinyl chloride), the sensor lost ∼45% of its activity in 5 days. The group of Lu¨bbers also reported on biosensors for xanthine, lactate, and cholesterol using oxygen transduction. Pyrene butyric acid acted as a fluorescent probe for oxygen and was covered with a layer containing the appropriate oxidase.155 Similar to the aspartame biosensor, a dual enzymatic system was used for determination of choline-containing phospholipids (e.g., lecithin).156 Phospholipids were hydrolyzed by phospholipase D to produce choline, which was subsequently oxidized by oxygen in the presence of choline oxidase. A similar bienzymatic system was used by Kotsira and Clonis,157 however, in combination with a pH optical transducer. Production of betaine during oxidation of choline results in a change of pH and protonation of the indicator bromothymol blue. Lactate monooxygenase is more stable than lactate oxidase and was used in a lactate biosensor.158 Lactate was monitored by measurement of the oxygen consumption via the fluorescence of decacyclene and can be determined with an LOD of 0.3 mM. Its stability in sol-gel matrix can be improved by addition of polycations.108 Detection of hydrogen peroxide via chemi- and electrochemiluminescence of luminol was used in biosensors for lactate,94,96,159 ethanol,160 choline,161 lysine,119,162 sulfite,163 xanthine and hypoxanthine,140,164 and choline.96,165,166 Uric acid and D-amino acids were detected by bienzymatic systems including uricase and POx,167 and D-amino acid oxidase and POx,168 respectively. Here, thiamine was oxidized by hydrogen peroxide to give fluorescent thiochrome, whose fluorescence was detected.

(5)

Glutamate produced in the first reaction was subsequently oxidized to 2-oxoglutarate (reaction 4). Hydrogen peroxide was detected by the chemiluminescence resulting from the oxidation of luminol that was catalyzed by either hexacyanoferrate(III) ion142,143 or POx.139 To avoid interference by glutamate, which can be present in samples along with glutamine, an ion-exchange resin was used142,143 to remove glutamate, which is an anion at pH 7. Sensors that make use of an oxidase-type enzyme and an oxygen transducer also were designed for lactate,142 ethanol and methanol,145-147 cholesterol,148-150 sulfite,151 bilirubin,152 and phenol.153 An aspartame biosensor incorporated both R-chymotrypsin and alcohol oxidase.154 Aspartame is hydrolyzed by R-chymotrypsin to produce methanol, which is

Figure 3. Schematic of an electrochemiluminescent multifunctional biosensing chip: GCE, glassy carbon electrode; Pt, platinum pseudo-reference electrode; S, silicone spacer; W, plexiglass window. A solution containing different analytes is injected in each channel, a working potential of +850 mV is then applied, and the emitted light (integrated over 3 min) is detected by a CCD camera. Reprinted with permission from Marquette, C. A.; Degiuli, A.; Blum, L. J. Biosens. Bioelectron. 2003, 19, 433. Copyright 2003 Elsevier.

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Table 2. Overview of Optical Enzymatic Biosensors: Abs, Absorbance; BL, Bioluminescence Intensity; CL, Chemiluminescence Intensity; EL, Electroluminescence Intensity; LI, Luminescence Intensity; LL, Luminescence Lifetime analyte acetylcholine acetylcholine acetylcholine acetylcholine ATP ATP ATP ATP ADP AMP L-alanine alcohols ethanol ethanol ethanol ethanol ethanol ethanol ethanol ethanol methanol D-amino acid aspartame captan Captan + ORP (paraoxon) bilirubin cholesterol cholesterol cholesterol choline choline choline with phospholipids fructose H2O2 H2O2 H2O2 H2O2 H2O2 H2O2 H2O2 hypoxanthine hypoxanthin glucose glucose glucose glucose glucose glucose glucose glucose glucose glucose glucose glucose glucose glucose glucose glucose glucose glucose glucose glucose glucose glucose glucose glucose glucose glucose

enzyme AChE AChE AChE AChE f. luciferase f. luciferase f. luciferase f. luciferase hexokinase + pyruvate kinase + glucose 6-P dehydrogenase adenylate kinase + creatine kinase + luciferase L-alanineDH ADH ADH + aldehyde-DH ADH AOx AOx AOx ADH + OR + b. luciferase ADH ADH AOx + HPOx D-amino acid oxidase + HPOx R-chymotrypsin + AOx GST GST + AChE bilirubin oxidase cholesterol oxidase cholesterol oxidase cholesterol oxidase ChOx + HRP ChOx phospholipase-D + ChOx GFOR HPOx HPOx HPOx HPOx HPOx HPOx HPOx XOx + HPOx XOx + POx GOx GOx GOx GOx GOx GOx GOx GOx GOx GOx GOx GOx GOx GOx GOx GOx GOx GOx GOx GOx GOx GOx GOx GOx + HPOx GOx + HPOx GOx + HPOx

analytical range

transducer pH pH pH pH

oxygen H2O2 oxygen

oxygen H2O2 oxygen

oxygen oxygen oxygen oxygen H 2O 2 H 2O 2 oxygen

H 2O 2 H 2O 2

oxygen oxygen oxygen oxygen oxygen oxygen oxygen oxygen oxygen oxygen oxygen oxygen oxygen oxygen oxygen oxygen oxygen pH pH pH H 2O 2 H 2O 2 H 2O 2

LOD

spectroscopy

indicator or substrate

ref

0.5-20 mM 0-20 µM 2-13 mM ? 0.1 nM-1 µM 0.1 nmol-1 µM ? ? 0.1-20 µM

0.5 mM ? 1 mM 50 µM 0.05 nM 0.1 nmol 0.1 pmol 10 pmol µm

LI ratio of LIs LI LI BL BL BL BL LI

?

25 pmol

BL

0.45-4.5 mM ? 1-100 mM 10-1000 mM 50-500 mM 3-750 µM 0.5-9 mM 0.4-70 µM 0.043-1 mM 0-1.1 mM 80 µM-60 mM 0-10 mM 0.056-3.07 mM 0-2.0 ppm 0-2.0 mM

? 0.9 mM ? ? 10 mM 3 µM 0.5 mM 0.4 µM ? ? 80 µM 0.3 µg/mL 32 µM ? ?

LI. LI. LI. LI. LI. CL LI. BL LI. LI. LI. LI. LI. Abs. Abs.

223 219 183 220 145 160 146 233 PEG-NAD+ 221 PEG-NAD 222 Ru-dpp 147 thiamine 168 Ru-dpp 154 CDNB + GSH 202 synthestic substrate 203

0.1-300 µM 0.2-3 mM 0.15-3.0 mM 0.07-18 mM 3-150 µM 10 pmol-30 nmol 0.08-3.00 g/L

0.1 µM 0.2 mM 0.15 mM 0.05 mM 3.0 µM 10 pM 0.08 g/L

LI. LI. LI. LI. CL EL. Lum. I.

Ru-dpp decacyclene Ru-dpp Ru-dpp luminol luminol Ru-dpp

0.278-331 mM ? 0.01-1 mM 0.05-1.2 mM 0.1-3 mM 17-117 µM 1-130 µM 0.5-250 µM 1-320 µM 0.5 µM-1 mM 0.5-0.8 mM 1.7-11 mM 1-10 mM 0.1-20 mM 0.1-500 mM 0.01- 2 mM 0.06-1 mM 0.05-1 mM 0.03-1.2 mM 0.1-8.3 mM 0-2.5 mM 0-20 mM 0.5-15 mM 0.7-10 mM 0.06-30 mM 0.1-15 mM 0.30-2.0 mM 0.3-5 mM 9.0-200 µM 0.1-0.8 mM 0.1-2 mM 0-1.7 mM 1-30 mM 0.25-250 nmol 0.3-300 µM 1-5 mM

0.278 mM 1 µM 1 µM 0.025 mM 0.67 mM 16.7 µM 1µM ? 0.55 µM ? 0.5 mM ? 0.5-2 mM 0.05 mM 0.1 mM 0.01 mM 0.06 mM ? 0.05 mM 0.1mM 80 µM 0.6 mM ? 0.75 mM 6 µM 0.1 mM 0.3 mM 0.3 mM 9.0 µM

intrinsic LI. CL CL CL CL CL LI. LI. CL CL intrinsic LI. intrinsic LI. intrincic Abs. LI. LI. LI. LI. LI. LI.,LL LI. LI. LI. LI. LI. LI. LI. LI. ratio of LIs LI. LI. LI. LI. Abs. CL CL CL

0.1 mM ? 1 mM 0.25 nmol 0.1 µM 0.43 mM

FITC SNARF HPTS FITC

206 34 208 207 228 229 234 235 NAD+ (coenzyme) 236 235 PEG-NAD+ NAD+ (coenzyme) NAD+ (coenzyme) NAD+ (coenzyme) Ru-bipy luminol a Ru(II) complex

152 148 149 150 165 96, 166 156

224 118 119 93 120 121 122 123 164 140 126 130 128 decacyclene 73 decacyclene 74 decacyclene 75 Ru-phen 76 decacyclene 77 PtOEP 58, 59 Ru-dpp 78 Al-ferron complex 79 Ru-ligand complex 80 Rudpp 81 Ru-phen 82 Ru-dpp 84 Ru-dpp 85 Ru-dpp 86 sulfonated Ru-dpp 35 a Ru complex 87 Ru-dpp 88 HPTS 103 Rhodamine (inert) 104 polyaniline 105 luminol 91 luminol 90 luminol 93

luminol luminol luminol luminol luminol homovanilic acid amplex red luminol luminol

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Table 2 (Continued) analyte

enzyme

analytical range

LOD

spectroscopy

0.01-0.8 mM 1-5 mM 0-10 mM 60 pmol-5 µmol 60 pmol-0.6 µmol 50 µM-10 mM 0.01-10 mM 17 µM-15 mM 30-200 µM 0.05-2.0 mM 0.1-5 mM 0.1-2 mM 1.1-11 mM 0-0.55 mM 0.055-55.5 mM 1-12 µM 0.1-60 µM 1 µM-1 mM ? 0-18 µM 1-100 µM 1 µM-2.5 mM 0.5-1 mM 0.3-6.0 mM 0.02-0.5 mM 3-200 µM 0.1-1 µmol 30-300 pmol 2-50 µM 0.2-1.0 mM 11-150 mg/L 10-250 IU/L 20-120 µM 10 µM-1 mM 5 µM-10 mM 0-0.1 g/L 0.04-0.12 µM 1 nM-10 µM 1 nM-3 µM 0.3 nM-3 µM 2 pmol-1 nmol 5.5 nM-1 µM 5-500 nM 1 nM-1µM 10-500 pmol 0-1.5 µM 0.07-1.25 µM

80 µM 0.8 mM 0.3 mM 60 pmol 60 pmol 26 µM 8.16 µM 17 µM 10 µM ? 0.2 mM 0.05 mM 0.6 mM ? 55 µM 0.1 µM 0.1 µM 1 µM 0.13 µM 0.2 µM 1 µM 1 µM ? 0.3 mM ? 3 µM ? 30 pmol 2 µM ? ? 10 IU/l 20 µM 10 µM 5 µM ? 0.04 µM 1 nM 0.3 nM ? 2 pmol 1 nM 5 nM 1 nM ? 0.125 µM 0.075 µM

LI. LI. EL. EL. EL. EL. EL. Abs. Abs. Abs. LI. LL. LI. LI. intrinsic LI. LI. CL CL LI. LI. CL CL intrinsic, LI. LI. LI. LI. CL EL. LI. Abs. & LI. LI. BL Abs. CL CL LI. Abs. Refl. BL BL BL BL BL BL BL intrinsic abs. intrinsic abs.

10-380 µM

?

Abs.

p-nitrophenyl phosphate alkaline phospatase

0-40 µM

?

LI.

ORP (paraoxon) ORP carbamate pesticides ORP (paraoxon) ORP ORP (Carbaryl) ORP (paraoxon) ORP (paraoxon) ORP (paraoxon) ORP (DFP) oxaloacetate

0.5-16 µM ? 0.8-3.0 mg/L ? ? 0.1-8.0 mg/L 0.01-0.48 mM 0.8-15 µM 20-140 µM 2-400 µM 3 nM-2 µM

0.2 µM 0.1-58 mg/L 25 ng 27 ppb 2 ppm 108 µg/l 2 µM 0.8 µM 20 µM 0.05 µM 1 nM

Abs. LI. Refl. LI. Abs. Abs. Abs. ratio of LI ratio of LI ratio of LI BL

0-10 mM 0.25-10 mM 0.1-10 mM 0.3-10 mM 0.1-25mM 1-10 mM 0.5-8 mM 0.25-10 mM 0.08-40 mM

? 75 µM 0.1 mM ? 0.1 mM 1 mM ? 0.25 mM 0.08 mM

Abs. LI. LI. Abs. LI. LI. Abs. & Refl. LI. LI.

GOx + HPOx GOx + HPOx GOx GOx GOx GOx GOx GOx GOx GOx GOx GOx GDH GDH GFOR GlOx GlOx + POx GlOx + POx GlDH + GPT GlDH GAH + GlOx GAH + GlOx + POx lactate monooxygenase lactate monooxygenase LOx LOx + HPOx LOx + HPOx LOx LDH LaDH LDH + GPT OR + b. luciferase phospholipase-D + ChOx LyOx + HPOx LyOx + POx mannitol-DH urease urease OR + b. luciferase OR + b. luciferase OR + b. luciferase OR + b. luciferase OR + b. luciferase OR + b. luciferase OR + b. luciferase nitrate reductase cytochrome cd1 nitrite reductase p-nitrophenyl phosphate alkaline phospatase

glucose glucose glucose glucose glucose glucose glucose glucose glucose glucose glucose glucose glucose glucose glucose glutamate glutamate glutamate glutamate glutamate glutamine glutamine lactate lactate lactate lactate lactate lactate lactate/pyruvate lactate lactate LDH lecitin lysine lysine mannitol Hg2+ Hg2+ NADH NADH NADH NADH NADH NADH NADH nitrate ion nitrite ion

penicillin penicillin penicillin penicillin penicillin penicillin penicillin penicillin phenol

AChE AChE AChE AChE AChE AChE OPH OPH OPH OPH malate DH + OR + b. luciferase penicillinase penicillinase penicillinase penicillinase penicillinase penicillinase penicillinase penicilinase tyrosinase

transducer H2O2 H2O2 H2O2 H2O2 H2O2 H2O2 H2O2 H2O2 H2O2 H2O2 H2O2 H2O2

NH3 H2O2 H2O2 H2O2 H2O2 oxygen oxygen H2O2 H2O2 H2O2

pH H2O2 H2O2 NH3 pH

pH pH pH pH pH pH pH pH pH pH pH pH pH oxygen

indicator or substrate Amplex Red Amplex Red luminol luminol luminol luminol luminol Ti(IV) reagent Fe(III) complex Prussian White EuTc EuTc NAD+ (coenzyme) PEG-NAD+ carboxyfluorescein luminol luminol NAD+ (coenzyme) NAD+ (coenzyme) luminol luminol decacyclene decacyclene homovanillic acid luminol luminol NAD+ (coenzyme) NAD+ (coenzyme) PEG-NAD+ BTB luminol luminol PEG-NAD+ Nile Blue pH ind. strip

p-nitrophenyl phosphate umbelliferyl phosphate AMPT indoxyl acetate chlorophenol red FITC o-nitrophenol bromcresol purple paraoxon carboxy SNARF-1 DDAO phosphate carboxy SNARF-1 bromocresol green acrylofluorescein FITC azo dye FITC FITC phenol red acryloylfluorescein Ru-dpp

ref 39 102 92 94 96 99 100 89 95 97, 98 101 61 213 222 224 141 139 140 217 218 142, 143 139 127 158 77, 144 122 159 96 214 216 221 232 157 162 119 222 210 211 225 226 227 228 230, 231 229 159 134 135 197 198 199 200 204 206 201 205 209 36 37 38 233 172 175 176 173 177 178 174 179 153

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Table 2 (Continued) analyte

enzyme

L-phenylalanineDH phenylpyruvateDH LOx + LDH sorbitolDH + OR + b. luciferase sulfite sulfite oxidase sulfite sulfite oxidase + HPOx superoxide radical superoxide dismutase + HPOx urea urease urea urease urea urease urea urease urea urease urea urease urea urease urea urease urea urease urea urease urea urease urea urease urea urease urea urease urea urease urea urease urea urease urea urease urea urease urea urease uric acid uricase + HPOx xanthine XOx + HPOx

transducer

analytical range

LOD

spectroscopy

indicator or substrate +

ref

0.6-6 mM 0-0.7 mM 0-0.1 mM 20 nM-10 µM

? ? 8.4 µM 20 nM

LI. LI. LI. BL

PEG-NAD PEG-NAD+ NAD+ (coenzyme)

223 222 215 233

oxygen H2O2 H2O2

0-100 ppm 1-100 µM ?

? 0.5 µM 20 nM

LI. CL LI.

perylene luminol Amplex Red

151 163 124

pH pH pH pH pH pH pH pH pH pH pH pH NH3 NH3 NH3 NH3 NH3 NH3 NH3 NH3 H2O2 H2O2

0-40 mM 0-2 mM 40-250 µM ? 0-1 mM 0-4 mM 0-100 mM 0.2-100 mM 0.06-1 M 2.0-12.0 mM 0.001-10 mM 0-400 µM 0.05-2.5 mM 0.25-8 mM 0.1-5 mM 0.01-1 mM 0.1-100 mM 10 µM-100 mM 0.1-10 mM 0.1 mM-0.1 M 3-30 mM 3.1-320 µM

? ? ? ? 20 µM ? ? 0.1 mM 0.06 M ? 2.5 µM ? ? 0.25 mM ? 0.03 mM 0.1 mM ? ? 0.1 mM 0.9 mM 2.2 µM

Abs. Refl. Refl. LI. LI. Refl. Abs. Abs. Abs. Abs. ratio of LI ratio of LI LI. Abs. LI. LI. Abs. Abs. Abs. LI. LI. CL

BTB BTB phenol red FITC NBD-PE FITC PVP Prussian Blue polypyrrole Prussian Blue FITC SNARF CF BTB HPTS Nile Blue acridine orange ETH 5350 brilliant yellow octadecyl dichlorofluorescein thiamine luminol

172 180 181 182 186 183 187 185 184 97 39 34 188 189 190 191 192 193 194, 195 196 167 164

L-phenylalanine phenylpyruvate pyruvate sorbitol

A microarray biochip for simultaneous electrochemiluminescent detection of several analytes was reported169,170 that contains the following six enzymes: GOx, glutamate oxidase, choline oxidase, lactate oxidase, lysine oxidase, and uricase. They were noncovalently immobilized (along with luminol) on anion-exchanger beads consisting of diethylaminoethyl sepharose, and the resulting beads were dispersed along with the luminol beads into poly(vinyl alcohol) bearing styrylpyridinium groups. This “cocktail” was spotted on the surface of a glassy carbon electrode, giving spots of 0.8 mm in diameter (Figure 3). The spots were allowed to polymerize under UV light. The electrochemiluminescence from the sixchannel, six-parameter biosensor was read by a CCD camera. Simultaneous measurements of glucose, glutamate, choline, lactate, lysine, and uric acid could be performed in the ranges 20 µM-2 mM, 1 µM-0.5 mM, 2 µM-0.2 mM, 2 µM-0.2 mM, 1 µM-0.5 mM, and 1 µM-25 µM, respectively. A biosensor for acetate also was developed.170 Acetate kinase (pre-immobilized on sepharose beads) and pyruvate kinase were immobilized in a layer brought into contact with the sample solution, while pyruvate oxidase was entrapped in a layer placed between the kinase layer and the glassy carbon electrode. Acetate could be measured in the range from 10 µM to 100 mM. A related microarray biochip of nine screen-printed graphite electrodes was prepared for determination for glucose and lactate.171 A reproducibility of within 4.4% was found at an optimum luminol oxidation potential of +650 mV. The LODs for simultaneous determination of glucose and lactate were 10 and 3 µM, respectively. Like in glucose biosensors, pH transducers were also used in biosensors for penicillin and urea. Penicilloate and protons are produced from penicillin in the enzymatic hydrolysis catalyzed by penicillinase. The decrease in pH is monitored by changes of absorption (reflectance)172-174 or emission

intensity175-179 of pH indicators. Fluorescein-derived indicators were used almost exclusively in order to monitor pH changes in neutral media. The pH indicator is usually contained in a hydrogel (most often a polyacrylamide gel), which is permeable for protons formed during the enzymatic reaction. A photopolymerization process was used to obtain a pH/penicillin array biosensor with spot diameters of ∼27 µm located on the surface of a 350 µm thick optical fiber.179 Penicillin and pH could be measured simultaneously (via imaging of fluorescence intensity with a CCD camera), and effects of changing pH, which often are large in complex fermentation media, could thus be compensated for. Biosensors based on pH transduction suffer from the fact that pH changes depend on the buffer capacity of the sample medium, which often is unknown and can hardly be compensated for. Urease-catalyzed hydrolysis of urea leads to formation of ammonium ions (eq 6) but also results in an increase in pH: urease

urea 98 2NH4+ + HCO3- + OH-

(6)

Consequently, two types of urea biosensors can be developed. Those based on pH transducers are designed analogously to penicillin optical sensors and make use of absorptionbased172,180,181 or fluorescent34,39,182,183 pH indicators. In contrast to most optical biosensors that rely on measurements of fluorescence intensity only, the one designed by Tsai and Doong39 employs a ratiometric scheme of self-referencing. Here, the intensity of the indicator (a fluorescein isothiocyanate-dextrane conjugate) is referenced to the pH-independent intensity of tetramethylrhodamine isothiocyanatedextrane. Such referencing makes it possible to overcome drawbacks of intensity-based measurements (for example, drifts in the intensity of the light source) but also seems to

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increase the sensitivity of measurements (LOD ) 2.5 µM, compared to 20-100 µM for other urea biosensors). The ratiometric approach also was used by Yadavalli et al.,34 who prepared sensor arrays composed of poly(ethylene glycol) hydrogel microspots with a diameter of ∼200 µm containing urease labeled with a seminaphthofluorescein indicator (SNAFL-1). The ratio of the intensities for the acidic (green emission) and basic (red emission) forms of the indicator was determined using a microscope with two different sets of excitation and emission filters. Polypyrrole was found to exhibit a pH-dependent intrinsic absorption with a peak at 650 nm. It was used in an urea biosensor where it acts as a support and a pH indicator simultaneously.184 In other work, Prussian Blue (also having a pH-dependent near-IR absorption) along with the enzyme was chemically incorporated into polypyrrole films.97,185 Brennan et al.186 co-immobilized the fluorescent probe nitrobenzoxadiazole dipalmitoylphosphatidylethanolamine (NBD-PE) and urease on the surface of an optical fiber whose surface was modified with alkylamine monolayers. Alterations of pH during the enzymatic reaction result in a change of physical and electrostatic structure of the membranes, which in turn alters the emission intensity of the NBD-PE. Fluorescence intensity decreases with the degree of ionization of head groups, consistent with an increase in self-quenching. By analogy to the approach made by McCurley,104 a pHdependent swelling of a polymer also was used for sensing urea.187 A layer of poly(vinyl pyrrolidone) cross-linked with sulfonated bisazidostilbenes was coated onto a metal mirror. Protons and certain other ions cause swelling of the material, and the changes in reflectance can be monitored. The analytical range of this fully reversible sensor is from 0 to 100 mM of urea. A number of urea biosensors are based on the determination of ammonia gas produced during hydrolysis of urea.188-196 Two types of ammonia transducers were used. In the first, a pH indicator is contained in a buffer solution positioned behind a gas-permeable membrane, in which the enzyme is immobilized. Gaseous ammonia diffuses through the membrane and dissolves in the buffer. This results in an increase of pH of the internal solution and in deprotonation of the indicator. Changes in absorbance or fluorescence intensity of the indicator are related to the ammonia concentration present in the external solution and, thus, to the level of urea. This sensor type was introduced by Rhines and Arnold188 and often used in later work.189,190,194,195 A completely different scheme is utilized in the second type of ammonia transducers. Such transducers contain the NH4+-selective neutral ionophore nonactin, a proton-selective neutral chromoionophore (a pH indicator), and a lipophilic anionic counterion dissolved in a plasticized poly(vinyl chloride) (PVC) matrix. The sensor layer is covered with a gas-permeable membrane to warrant selectivity for ammonia gas by inhibiting a direct ion-exchange reaction between the sample solution and the sensor membrane. A membrane with immobilized urease is mounted on top of the sensor. Ammonia gas diffuses through the gas-permeable membrane and reaches the PVC layer, where the reaction outlined in eq 7 occurs:

IndH+ + NH3 + IP f Ind + IP-NH4+

(7)

Here, Ind and IP are the neutral chromoionophore and the

neutral ionophore, respectively. Again, both absorptionbased192,193 and fluorescent191,196 pH indicators came to use. Kawabata et al.192 manufactured a 140 µM thick urea microsensor based on this principle. Certain enzymatic reactions do not require optical transduction via a chemical sensor because optically detectable species are generated or consumed during the reaction. Such sensors usually consist of a membrane that contains the immobilized enzyme. Chromogenic or fluorogenic substrates and any cosubstrates are added to the sample into which the sensor is submerged. Because of the absence of a transducer, such biosensors are often referred to as direct optical biosensors. One type of a direct optical biosensor is based on hydrolysis of a substrate catalyzed by a hydrolase-type enzyme. The principle was first demonstrated by Arnold,197 who used immobilized alkaline phosphatase to catalyze hydrolysis of p-nitrophenyl phosphate, which results in the formation of yellow p-nitrophenolate. A linear dependence of the change in absorbance on the concentration of the substrate was observed. The method was further developed by Freeman and Bachas,198 who introduced a sensor that makes use of a competition between two substrates (4-methylumbelliferyl phosphate and p-nitrophenyl phosphate) for the active state of the model enzyme alkaline phosphatase. If the sensor was placed in a solution containing 4-methylumbelliferyl phosphate, the highly fluorescent anion of 4-methylumbelliferone was produced upon hydrolysis. In the presence of the analyte (p-nitrophenyl phosphate), the rate of fluorescence change caused by production of 4-methylumbelliferone was decreased. The method also was demonstrated to work for the determination of adenosine monophosphate, another substrate of alkaline phosphatase. A sensing scheme for determination of organophosphorous pesticides such as paraoxon199 is based on the inhibitor action of acetylcholine esterase (AChE). The biosensor makes use of a synthetic yellow substrate that is converted into a blue product by AChE. Inhibition of the reaction by pesticides is monitored spectroscopically. The LOD for paraoxon was 200 nM. Although the immobilized enzyme showed a very good long-term stability, that of the synthetic substrate was rather low in that the half-lifetime was ∼2 weeks only at room temperature. Others200 have used indoxyl acetate as a substrate for AChE. The fluorescence intensity of indoxyl (λmax ) 470 nm) was related to the concentration of the inhibiting pesticide. Similar to the work of Freeman and Bachas,198 enzymatic hydrolysis of substrate o-nitrophenyl acetate was used also for determination of organophosphates.201 Choi et al.202 developed a biosensor for captans, a group of systemic organophosphorus fungicides and pesticides. The enzyme glutathione-S-transferase (GST) converts the substrates, 1-chloro-2,4-dinitrobenzene and glutathione, into yellow S-(2,4-dinitrophenyl) glutathione. In the presence of captans, GST is inhibited and the amount of the product is decreased. A dual enzymatic system consisting of GST and AChE was shown to be suitable for simultaneous determination of both paraoxon and captan.203 The absorbance of S-(2,4-dinitrophenyl)glutathione (the product of the reaction catalyzed by GST) and R-naphthol (the product of the reaction catalyzed by AChE) was detected at 400 and 500 nm, respectively. It was observed that AChE was inhibited by both captan and organophosphorus compounds, while GST was inhibited by captan only. Thus, simultaneous detection of both analytes becomes possible.

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Inhibitors of AChE can also be detected using pH optical transducers. The hydrolysis of acetylcholine chloride according to eq 8 results in the formation of acetic acid and, therefore, in a decrease in pH: AChE

acetylcholine + H2O 98 choline + acetate + H+ (8) Different absorption-based204,205 and fluorescent34,206,207 indicators were used for determination of organophosphorous and carbamate pesticides as well as of acetylcholine itself.208 All biosensors that make use of the inhibition of AChE show good sensitivity but are severely limited in specificity because AChE is the target of a wide variety of toxic inhibitors. These range from heavy metal ions to chemical warfare agents. Both organophosphate- and carbamate-based pesticides inhibit AChE. Additionally, most sensors using AChE inhibition have lengthy response times because of long incubation periods, inhibition is often irreversible, and subsequent reactivation of AChE sometimes is impossible. More recently, interest has been directed to organophosphorus hydrolase (OPH), which is not susceptible to nonspecific inhibition and offers much better specificity than AChE. OPH hydrolyzes a range of organophosphate esters, including pesticides such as paraoxon and chemical warfare agents such as soman or sarin. Catalytic hydrolysis of these compounds is accompanied by a release of protons, which makes possible determination of organophosphorus pesticides (ORP) using pH transducers.36,38 Hydrolysis of some ORP also produces detectable chromophoric products.209 An interesting approach was introduced by Simonian et al.37 The sensing scheme is far from a conventional enzymatic biosensor but resembles a competitive immunoassay. The enzyme OPH is covalently attached to the surface of a gold nanoparticle. The fluorophore 7-hydroxy-9H-(1,3dichloro-9,9-dimethylacridin-2-one)phosphate binds weakly to the active site of the enzyme. The fluorescence of the bound fluorophore is enhanced because of the proximity to the gold surface. An inhibitory substrate is added that has a much higher affinity for the active site of OPH and, thus, replaces the fluorophore. When far away from the gold surface, no enhancement of fluorescence is observed any longer. Biosensors for heavy metal ions have been reported that exploit the inhibition of urease by heavy metal ions.210,211 Urease is chosen as the enzyme since it is particularly sensitive to ions such as Pb(II), Cd(II), Ag(I), and Hg(II). In contrast to the catalytic biosensors that monitor inhibition of enzymatic activity, certain biosensors use a different mechanism of signal transduction. Walker and Asher212 designed an ultasensitive biosensor for the pesticide parathion. The sensor utilizes an array of colloidal polymer particles (which diffract light in the visible spectral region) emulgated into a polyacrylamide-based hydrogel. AChE is covalently attached to the hydrogel backbone where it irreversibly binds parathion, which in turn results in the formation of a charged product. This induces swelling of the hydrogel network and results in a shift of the wavelength of the diffracted light that is proportional to the concentration of the analyte. The sensor is capable of sensing parathion in the fM to pM concentration range. The LOD is 4.3 fM of parathion, which is several orders of magnitude lower than those for other sensors for organophosphates. As expected, ionic strength severely influences swelling properties and, thus, the performance of the sensor; therefore, 30 min of

washing with deionized water is necessary after incubation of the sensor with parathion (30 min). A number of enzymatic redox reactions require the presence of coenzymes. Most systems involve nicotinamide adenine dinucleotide (NAD+), to which a hydrogen atom and an electron can be transferred, while the substrate is oxidized according to eq 9: dehydrogenase

substrate + NAD+ 98 product + NADH (9) The formation of the reduced form of nicotinamide adenine dinucleotide (NADH) can be monitored via its characteristic absorption at λmax 350 nm and emission peaking at λmax 450 nm, and this enables the optical quantification of substrate concentrations. In early work, the dehydrogenases were immobilized on (or in) a hydrophobic polymer membrane, and the “sensors” were operated in the flow-injection mode, where the analyzed substrate and NAD+ are passed over the sensor layer. Such sensors were developed for glucose,213 lactate and pyruvate,214-216 glutamate,217,218 and alcohols183,219 using glucose dehydrogenase, lactate dehydrogenase, glutamate dehydrogenase, and alcohol dehydrogenase, respectively. Enzymatic oxidation of lactate by NAD+ results in the formation of pyruvate and NADH. The reaction thus was used not only for determination of lactate but also in the reverse direction for pyruvate,215 with NADH being supplied instead of NAD+. Evidently, the main disadvantage of such sensors relies on the fact that the coenzyme needs to be added to the sample solution. Therefore, some effort was focused on designing a self-contained biosensor, i.e., a sensor that does not require the supply of any additional components. For example, NAD+ was immobilized together with alcohol dehydrogenase in a sol-gel monolith.220 However, leaching of the coenzyme into the solution was not completely excluded, and the response of the sensor was rather slow (∼30 min). A novel approach was proposed by Scheper and Buckmann,221-223 who used a poly(ethylene glycol) molecular weight-enlarged NAD+ (PEG-NAD+) instead of NAD+. A pair of dehydrogenase-type enzymes (for substrate detection and for regeneration of the coenzyme) and PEG-NAD+ were enclosed in the sensing compartment between the ultrafiltration membrane and the fiber-optic tip. The analytes and products were allowed to diffuse freely through the ultrafiltration membrane. In contrast to previous sensors (where NAD+ was supplied in solution), this sensor type allows for the regeneration of PEG-NAD+ in a subsequent reaction such as the one in eq 10: LDH

pyruvate + NADH + H+ 98 lactate + NAD+ (10) The scheme was used for determination of glucose, lactate, ethanol, pyruvate, mannitol, formate, L-alanine, and L-phenylalanine. The unique enzyme glucose-fructose oxidoreductase (GFOR) is capable of dehydrogenating glucose to gluconolactone and of simultaneously reducing fructose to sorbitol in a ping-pong mechanism according to GFOR

glucose + fructose 98 gluconolactone + sorbitol (11) Both the intrinsic absorbance and fluorescence of NADH can be measured and enable optical detection of both substrates. In the GFOR-based biosensor,224 the enzyme was

436 Chemical Reviews, 2008, Vol. 108, No. 2

cross-linked with glutaraldehyde and placed between an optical fiber and a dialysis membrane. Glucose was sensed via the increase in fluorescence of the enzyme due to formation of NADH, and fructose was sensed via the decrease in fluorescence due to consumption of NADH. The system can be regenerated by passing fructose or glucose solutions, respectively, over it. Apart from measurements of its intrinsic absorbance or fluorescence, NADH can be detected with much higher sensitivity via reactions 12 and 13, which are catalyzed by bacterial enzymes and result in blue-green bioluminescence:

Borisov and Wolfbeis

entrapped in a collagen membrane, as well as that of adenosine monophosphate (AMP) and adenosine diphosphate (ADP) by entrapping two additional enzymes, adenylate kinase and creatine kinase, responsible for conversion of AMP and ADP into ATP. The sensitivity for ATP was significantly lower and the limit of detection was significantly higher (10 pmol) than for sensors using luciferin in solution. ADP was also determined via the fluorescence of NADH that is formed in the following sequence of reactions236 (eqs 15-17) that are catalyzed by the enzymes pyruvate kinase (PyKin), hexokinase (HexKin), and glucose-6-phosphate dehydrogenase (Glu-6P-DH):

oxidoreductase

NADH + H+ + FMN 98 NAD+ + FMNH2 (12) bacterial luciferase

FMNH2 + R-CHO + O2 98 FMN + R-COOH + H2O + hν (13) Here, FMN is flavine mononucleotide and R-CHO is a longchained aldehyde, e.g., decanal. The emission of bioluminescence peaks at 490 nm. Initially, biosensors for NADH only made use of enzymes immobilized onto a polymer support (a preactivated polyamide membrane), while the cofactor FMN and the longchained aldehyde were added to the solution to be analyzed.225-229 Attempts were made to design a self-contained biosensor that would not require the addition of coreactants and, therefore, would operate in a reagentless mode.230,231 Hence, the flavine cofactor was noncovalently entrapped in a matrix of poly(vinyl alcohol), which allowed its controlled release in the vicinity of the immobilized enzymes. The method works well but the aldehyde needs to be continuously supplied to the reaction medium. Oxidation of a substrate by a dehydrogenase-type of enzyme coupled to bioluminescent detection of NADH also was used for determination of the activity of lactate dehydrogenase232 (lactate was supplied together with FMN and the aldehyde) and for analysis of sorbitol, ethanol, and oxaloacetate.233 Compared to the ethanol biosensors based on direct detection of NADH and those using alcohol oxidase, the biosensor with coupled bioluminescent detection of NADH proved to be 1-2 orders of magnitude more sensitive, with a typical LOD being 0.4 µM. At the same time, such a system is more complicated because it makes use of three enzymes and requires cosubstrates such as FMN and an aldehyde to be added. The luciferin/luciferase bioluminescent system with its λmax of 560 nm was adapted to the determination of adenosine triphosphate (ATP). Oxidation of luciferin is catalyzed by firefly luciferase according to eq 14, firefly luciferase

ATP + luciferin + O2 98 AMP + oxyluciferin + P2O74- + CO2 + hν (14) and results in green luminescence. As in the case of bioluminescent determination of NADH, biosensors for ATP are extremely sensitive (LODs are <1 pmol). In earlier systems, luciferin again had to be added to the sample solution.228,229,234 To overcome this inconvenience, a reagentless biosensor was designed.235 Here, luciferin was incorporated into acrylic (Eudragit) microspheres entrapped in a film of poly(vinyl alcohol). Such a controlled-release system allowed the determination of ATP via firefly luciferase

PyKin

ADP + phosphoenol pyruvate 98 ATP + pyruvate (15) HexKin

ATP + glucose 98 ADP + glucose-6-phosphate (16) Glu-6P-DH

glucose-6-phosphate + NAD+ 98 6-phosphogluconate + NADH + H+ (17) ADP could be determined in concentrations as low as 0.1 µM. It was discovered recently that the europium tetracycline (EuTC) 1:1 complex can act as a luminescent probe for nucleoside phosphates including AMP, ADP, and ATP. The probe can be excited with the 405 nm laser diode and is nonspecific, but the response to the various phosphates is different. It has been applied to the determination of the activity of soluble kinases (which are important in highthroughput screening for new drugs).237 The same group has used EuTC to monitor the activity of alkaline phosphatase or the efficiency of an inhibitor by determining the amount of phosphate released by the enzyme from phenyl phosphate.238 An overview on enzyme-based biosensors is given in Table 2 along with typical data of merit. In conclusion, it can be stated that most enzymatic biosensors (a) are fairly easy to fabricate; (b) do not require labeling but a transducer capable of detecting reaction products or coreactants; (c) are suitable for continuous analyte monitoring; (d) have moderate sensitivity and limits of detection; (e) are prone to poisoning; and (f) are thermally labile (with few exceptions) and frostsensitive in aqueous solutions.

4. Immunosensors 4.1. General Remarks Affinity biosensors make use of specific interactions between an antibody (Ab) and an antigen (Ag) or a hapten. Antibodies are large Y-shaped proteins (∼150 kD) used by the immunosystem to identify and neutralize alien objects like bacteria and viruses. The affinity of Ag’s to Ab’s is very strong (Ka 1012-1014) but of strictly noncovalent nature. Binding of an antibody to its specific antigen can cause precipitation of the Ab-Ag complex, result in blocking of viral receptors, or mark the Ag for digestion by phagocytes. Smaller molecules such as pesticides or hormones often cause immune response only when attached to a large carrier such as a protein and usually are referred to as haptens. Antibodies to the hapten-carrier adduct produced by the body are able, however, to bind the hapten. It should be stressed that most

Optical Biosensors

components of physiological pathways are not immunogenic, e.g., glucose, citrate, fatty acids, amino acids, and the like. Immunosensors are mainly used for determination of concentration of antigens or haptens or, alternatively, for sensing antibodies because their presence can indicate an infection. We differentiate between immunoassays (performed in solution and not treated here except for certain examples) and immunosensors (on solid supports). The latter are treated here but actually are not sensors in their strictest definition because they are not capable of continuously and reversibly recording a parameter. Solid-phase immunoassays make use of a recognition element (Ab or Ag) immobilized on the surface of an inert support which, however, also may act as an optical fiber or a planar waveguide. Despite the fact that a binding event between an antibody and an antigen is reversible and noncovalent, most immunoreactions are irreversible in practice because of very large association constants and very slow dissociation rates. As a result, practically all immuno“sensors” are suitable for a single measurement only. This makes calibration difficult and requires an enormous reproducibility in manufacturing. Given this, attention has been paid to the regeneration of sensors (e.g., by washing with solutions of high osmolarity, high ionic strength, or low pH), which allows for multiple measurements with a single sensor. However, regeneration procedures do not always result in full recovery of the activity. In recent years, a number of devices were developed that made possible simultaneous detection of several analytes performed automatically.

4.2. Immunosensor Formats The most widely used formats are illustrated in Figure 4, where the upper panel represents the situation before immunobinding has occurred and the lower panel represents the situation after it. A selection of immunosensors for various analytes is presented in Table 3.

4.2.1. Direct Immunosensors These sensors are fairly straightforward but have been reported for a limited number of analytes only.43,44,66-68,72,239-248 The sensing format is schematically shown in Figure 4. An unlabeled antigen binds to an unlabeled antibody. Interferometric readout is common since it has the advantage of not requiring a label.66-68 The change of refractive index, however, is much smaller than in a fluorescent or radiolabel sandwich format (Figure 3c) because antigens and particularly haptens possess relatively low molecular mass. The intrinsic fluorescence of benzo[a]pyrene tetraol (BTP) was used as analytical information; the anti-BTP antibodies were immobilized onto silica microbeads.239,240 The optical signal of such single-shot probes is directly proportional to the amount of BTB captured. The LOD is 0.5 nM. Another example is represented by a biosensor for human serum albumin (HSA).241 When HSA binds to dansyl-labeled antibody attached to the surface of an optical fiber, an increase in fluorescence is observed because the antigen shielded the label from quenching water molecules. Engstro¨m et al.249 observed an enhancement of the intrinsic UV fluorescence of tryptophan of monoclonal mouse antibodies immobilized on the surface of a quartz slide upon binding maltose and panose (a rather rare triglucoside). The low affinity of the antibodies for the saccharides enabled a virtually reversible sensing, with no need for sensor regeneration. The analytical range was from 0 to 8 mM of the

Chemical Reviews, 2008, Vol. 108, No. 2 437

Figure 4. Typical formats of heterogeneous optical immunoassay; situations before (upper part) and after equilibration (lower part). The enzyme-linked immunosorbent assay (ELISA) in practically all cases is a modification of the sandwich method as it makes use of an enzyme as a label. Thus, it requires a subsequent enzymatic reaction to produce a colored or fluorescent product whose concentration can be determined, usually in solution and not on the surface of the sensor.

respective sacharides, and the LODs were 5 and 15 µM, respectively, for panose and mannose. Reck et al.250 reported a homogeneous immunoassay for thyroxine hormone. Quenching of the intrinsic fluorescence of the thyroxine-binding globulin was observed upon binding the thyroxine. Although the initial response was achieved after 5 min of incubation, almost 2 h were needed until the system reached saturation. The main drawback of this approach is its low sensitivity, since the LOD was found to be ∼100-fold higher than the concencentration of free thyroxine in serum. Other fluorescent immunosensors used in the direct format242-245 can only serve as model systems, since labeling of the analyte is necessary in the case of the assay, which is difficult (or even impossible) for real samples.

4.2.2. Competitive Immunosensors In this format (see Figure 4b),41,42,49,246,251-261 an unlabeled antigen (the analyte) and its labeled form compete for a limited number of binding sites of the immobilized antibody. Fluorescence intensity is inversely proportional to the amount of the analyte concentration. The application of the methods requires a labeled antigen to be available. The method can be inversed to enable the detection and assay of antibodies via the competitive binding of labeled and unlabeled antibodies, respectively, to an immobilized antigen.262

4.2.3. Sandwich Immunosensors Such assays (Figure 4c)60,243,261,263-272 are widely used and require relatively large antigens that contain at least two epitopes (the site of a macromolecule that is recognized by an antibody) for the antigen to be bound to the immobilized capture antibody and to the labeled second antibody. Fluorescence intensity is proportional to the concentration of the fluorescently labeled antibody, which, in turn, is related to the concentration of the antigen. Two different protocols are usual. In the “stepwise” protocol, the antigen and fluorescently labeled second antibody are added sequentially to the biosensor. In the “premixed” protocol, the antigen and antibody are premixed before injection into the biosensor. The stepwise protocol is said to give a significantly higher

analyte

assay format

analytical range

LOD

scheme

label

direct

0-9 mg/L

?

FI

dansyl/anti-HSA attached to fiber

direct direct direct

0.011-0.11 g/L 1-100 nM 10-250 nM

0.011 g/L 0.5 nM 20 nM

FI intrinsic FI FI

FITC/anti-rabbit IgG

anti-goat IgG IgG trinitrobenzene PSA LDH hemoproteins

direct direct direct direct direct direct

0.3-10 mg/L 0-5 mg/L 1-8 µg/L 1-100 µg/L 0.125-5.0 mg/L 10 nM-10µM

0.3 mg/L ? 1 µg/L 0.5 mg/L 0.03 mg/L ?

FI FI FI FI FI FI & FRET

Salmonella typhimurium

direct

?

1.03 × 105 cfu/L

ratio of FI

anti-rabbit IgG phenytoin

competitive competitive homogeneous competitive homogeneous competitive competitive competitive competitive competitive competitive competitive competitive competitive competitive

0-150 nM 1-20 µM

8 nM 1 µM

FI FI & FRET

FITC/anti-rat IgG Q-dot/protein A on optical fiber Cy-5/trinitrobenzol allophycocyanin/anti-PSA antibody FITC/LDH FITC labeled anti-hemoprotein IgG on LB monolayer complex of Alexa Fluor 546/ anti-Salmonella IgG and Alexa Fluor 594/protein G on optical fiber; FRET FITC/anti-rabbit IgG Texas red/anti-phenytoin IgG

0-300 µM

?

FI & FRET

Texas red/anti-theophylline IgG

1 nM-1 mM 0.5-200 nM 10-4-10-1 g/L 0.01-1 µM 0.1 nM-1 µM 10-1000 µg/L 1-1000 µg/L 20-200 µg/L 1-100 µg/L 50 µg/L-10 mg/L

1 nM 0.5 nM ? 5 µg/L ? 10 µg/L 5 µg/L 20 µg/L 2.5 µg/l 10 ppb

FI FI FI FI FI FI FI FI FI FI

aminofluorescein /inazethapyr fluorescein/atrazine FITC/rabbit IgG fluorescein and benzoylecgonine fluorescein/CCA Cy-5-labeled TNT sulfonate Cy-5/EDTA-TNB Cy-5/trinitrophenyl Cy-5/EDTA-RDX fluorescein/TCPB

competitive sandwich sandwich

1-50 mg/L 0-50 nM 0.03-1.2 nM

? ? 30 pM

FI FI FI

sandwich sandwich

1.5-200 mg/L 10-100 µg/L

2 mg/L ?

FI FI

Cy-5/theophylline FITC/anti-hCG IgG TRITC/ anti-botulinium toxin A IgG FITC/anti-AFS IgG NIR dye1/goat anti-human IgG

sandwich sandwich sandwich

40-300 ng/L 0.1-10 mg/L 3 × 107-3 × 1011 cfu/L 0.1-250 µg/L 0-1 mg/L 30-400 ng/L 1-1000 µg/L 5-30 µg/L 0.5-25 µg/l 2.5-10 µg/L 10-50 µg/L

40 ng/L 0.03 mg/L 107 cfu/L

FI FI FI

100 ng/L 10 µg/L 30 ng/L 1 µg/L 5 µg/L 0.5 µg/L 2.5 µg/L 10 µg/L

FI FI FI FI FI LL LL FI

theophylline inazethapyr atrazine human IgG cocaine CCA TNT TNT TNT RDX TCPB theophylline hCG Clostr. botulinium toxin A ASF protein human IgG mouse IgG LDH Salmonellax typhimurium ricin SEB SEB hCG TNT mouse IgG LDH TNT

sandwich sandwich sandwich sandwich sandwich sandwich sandwich displacement

material

assay time

ref

30 min

241

? 45 min ?

244 239, 240 245

30 min 10 min 5 min ? 4 min 30 min

242 44 246 247 243 248

5 min

43

20 min 15 min

262 41

15 min

42

2 min 10 min 15 min 15 min 20 min 4 min 5 min 5 min 5 min 20 min

252, 253 254 255 256 258 257 259 246 259 260

5 min 2 min 2 min

261 263 264, 265

40 min 30 min

266 267

Cy-5/anti-mouse IgG FITC/anti-LDH IgG Cy-5/anti-Salmonella IgG

anti-AFS IgG on Immobilon membrane goat anti-human IgG on PMMA droplet on fiber anti-mouse IgG on capillary anti-LDH IgG on fiber anti-Salmonella IgG on fiber

20 min 4 min 60 min

271 243 268

Cy-5/anti-ricin IgG Cy-5/anti-SEB IgG Cy-5/sheep anti-SEB IgG Cy-5/anti-hCG IgG Cy-5/anti-TNT IgG GOx/anti-mouse IgG GOx/anti-LDH Cy-5/trinitrophenyl

anti-ricin IgG on fiber anti-SEB IgG on PS waveguide sheep anti-SEB IgG on a capillary surface anti-hCG IgG on waveguide anti-TNT IgG on waveguide mouse IgG on PtOEPK/PS layer LDH on PtOEPK/PS layer anti-TNT IgG on waveguide

15 min 10 min 20 min 5 min 15 min 1h 1h 5 min

269 270 271 261 246 60 60 246

FITC/protein A

rabbit IgG on silica beads anti-BTP IgG on silica beds (7 µm) human IgG on an ion exchange waveguide goat IgG on a patterned waveguide anti-TNT IgG on a waveguide anti-LDH IgG on fiber

rabbit IgG on fiber phycoerythrin-phenytoin; semipermeable membrane phycoerythrin-theophylline in a well with semipermeable membrane sheep anti-inazethapyr IgG on fiber anti-atrazine IgG on fiber anti-human IgG on waveguide anti- benzoylecgonine IgG on fiber anti-CCA IgG on fiber anti-TNB IgG on fiber anti-TNT IgG on fiber anti TNT IgG on waveguide anti-RDX IgG on fiber anti- polychlorinated biphenyls IgG on fiber anti-theophylline IgG on waveguide anti-hCG IgG on waveguide anti-botulinium toxin A IgG on fiber

Borisov and Wolfbeis

human serum albumin (HSA) anti-rabbit IgG BTP protein A

438 Chemical Reviews, 2008, Vol. 108, No. 2

Table 3. Overview of Immunosensors (Acronyms Used for Schemes: FI, Fluorescence Intensity; FRET, Fluorescence Resonance Energy Transfer; LL, Luminescence Lifetime; RI, Refractive Index)

2 min 1h 30 min 10 min 20 min 15 min 15 min 15 min 15 min 15 min 15 min 12 min

ref assay time material

anti-paclitaxel on a capillary aminohexylatrazine on fiber aminohexylatrazine on fiber atrazine on the interference layer atrazine derivative on waveguide atrazine derivate on waveguide isoproturon derivate on waveguide estrone derivate on waveguide 2,4-D on fiber dichloroanilineglutaric acid on waveguide paraquat derivative on waveguide testosterone derivative on waveguide

label

Rhodamine/paclitaxel FITC/anti-triazine IgG anti-triazine IgG mouse anti-atrazine IgG Cy-5.5/anti-atrazine IgG Cy-5.5/anti-atrazine IgG Cy-5.5/anti-isoproturon IgG Cy 5.5/anti-estrone IgG Cy-5/anti-2,4-D Cy-5/anti-2,4-D Cy-5.5/anti-paraquat IgG Cy-5.5/anti-testosterone IgG

scheme

FI FI RI RI FI FI FI FI FI FI FI FI 1 µg/L 0.1 mg/L 15 mg/L 0.25 ppb ? 0.16 µg/L 0.05µg/L 0.08µg/L ? 0.03 µg/L 10 ng/L 0.2 ng/L

LOD analytical range

1-100 µg/L 0-10 mg/L 20-200 mg/L 0.1-1 mg/L 0.1-1000 µg/L 0.35-1.47 µg/L 0.11-2.83µg/L 0.17-10.7µg/L 0.3-10 µM 0.07-1.8 µg/L 0.01-100 µg/L 0.2 ng/L-100 µg/L

assay format

displacement binding inhibition binding inhibition binding inhibition binding inhibition binding inhibition binding inhibition binding inhibition binding inhibition binding inhibition binding inhibition binding inhibition

analyte

paclitaxel terbutryn terbutryn atrazine atrazine atrazine isoproturon esterone 2,4-D 2,4-D paraquat testosteron

Table 3 (Continued)

273 274 275 69 276 277 277 277 279 280 281 282

Optical Biosensors

Chemical Reviews, 2008, Vol. 108, No. 2 439

response than the premix protocol. Typical examples of fluorescent sandwich assays are listed in Table 3. Gold nanoparticles may be used to enhance the fluorescence signal of a labeled antibody. Hong et al.272 have designed a biosensor operating in a sandwich format by immobilizing protein C antibody, protein C, and secondary antibodies that were labeled with Cy-5 and Alexa Fluor 647. Their signal was typically inhanced by 10-200% when gold nanobeads were added. Self-assembled nanolayers (SAMs) can be used to control the distance between the nanobeads and the labeled antibodies. Maximal enhancement of the signal was achieved with SAM thicknesses of 2 nm. It was also found that using ethanol instead of water resulted in an up to 10-fold enhancement of fluorescence intensity.

4.2.4. Displacement Immunosensors This less common format (Figure 4d)246,273 requires an initial saturation of all the antibody binding sites with a fluorescently labeled antigen. Upon introduction of the unlabeled antigen, displacement of the labeled antigen occurs and is measured in this sensor as a decrease in the fluorescence intensity. It needs to be kept in mind that the “sensor” only works if the reaction rates of displacement are adequately fast. In fact, such sensors usually are quite slow.

4.2.5. Immunosensors Based on Binding Inhibition In contrast to other formats, such sensors (Figure 4e)69,70,274-282 require immobilization of an unlabeled analyte derivative on the surface of a waveguide. In the absence of the antigen, the labeled antibody can bind to the surface. Binding is inhibited, however, in the presence of the analyte because it blocks the binding sites (paratopes) of the antibody. For the binding inhibition assay to be quantitative, the number of high-affinity binding sites on the surface has to be much higher than the number of labeled antibodies in the bulk solution. If an antigen or an antibody is immobilized on a solid surface, the flexibility of the system becomes limited. More flexible sensors can be prepared by immobilization of an arbitrarily chosen ligand (antigen) on the surface.276 Different analyte derivatives that are conjugated to an antibody can be then attached to the ligand for assay operation in the binding-inhibition format.

4.2.6. Comparative Study on Immunosensor Formats Snapsford et al.246 compared the performance of four immunoassay formats (direct, competitive, sandwich, and displacement) that can be used for determination of 2,4,6trinitrotoluene (TNT). Anti-TNT was attached to waveguide surface via avidin-biotin linkage. Cy-5-labeled trinitrobenzene was used in direct, competitive, and displacement assay formats, and Cy-5-labeled antibody was used in a sandwich format. While TNT itself is unsuitable for detection using the sandwich assay format, it was conjugated to ovalbumin (OVA), so that the OVA-TNT complex can be bound not only to the antibody attached to the waveguide but also to the tracer antibody. Each assay format resulted in different LODs and dynamic ranges. The LOD was the lowest in the direct assay (1 µg/L); however, it also had the narrowest dynamic range (1-8 µg/L). On the other hand, the dynamic range was the widest in the competitive assay (20-200 µg/ L), but it had the highest LOD (20 µg/L). The LODs in the displacement and sandwich assay formats were 10 and 5 µg/ L, respectively. While 15 min was required to perform the

440 Chemical Reviews, 2008, Vol. 108, No. 2

Figure 5. Principle of the total internal reflection fluorescence in an optical fiber waveguide. On reflection at dielectric interface, light penetrates into the second phase that has a lower refractive index than that of the core. Intensity decreases exponentially over the penetration depth dp (which typically is about as long as the wavelength of the light employed). Any labeled antibodies located in the declad zone within dp are excited to produce fluorescence, while those located outside this distance will not.

two-step sandwich assay, the other assays required only 5 min. It was also shown that complete regeneration of the sensor was possible within 2 min by passing a regeneration buffer containing 50% ethanol over the sensor layer. No loss of activity was observed after 10 regeneration cycles.

4.3. Preferred Optical Readout Formats in Immunosensing 4.3.1. Conventional Readout Formats In the most simple version, immunosensor spots on glass, metal, or plastic supports are read out by either absorption, fluorescence, interferometry, various methods of polarization spectroscopy, or surface plasmon resonance (treated elsewhere).283 They can be combined with methods of optical spectroscopy. Fluorescence intensity serves as the analytical parameter in most immunosensors (see Table 3) and is mostly read at a single wavelength, but this may cause difficulties in measuring reproducible data. According to Parker’s law (see section 2.4), luminescence intensity depends not only on the concentration of the fluorophore but also on other variables such as the intensity of the exciting light and the geometry of the experimental arrangment. Self-referenced methods, where the latter parameters are being referenced out, are, therefore, preferred. Measurement of intensity at two wavelengths (e.g., after addition of a reference dye or by making use of fluorescence resonance energy transfer from a donor fluorophore to an acceptor fluorophore) is one common self-referenced method. Solution immunoassays often are performed by measuring polarization, but less often in solid-state devices for obvious reasons. The measurement of fluorescence decay time represents another, albeit less common, self-referenced method.

4.3.2. Evanescent Wave, Capillary, and Other Readouts It was recognized rather early that the solid supports required in biosensors also may act as optical components. Evanescent wave spectroscopy (EWS) has become particularly useful and is often applied in immunosensors (which contrasts the situation in the case of enzymatic biosensors). EWS can be performed in various ways, but total internal reflection fluorescence (TIRF) is, by far, the most often applied. A schematic of how TIRF works is shown in Figure 5. Light transported by a waveguide (here, an optical fiber) excites the fluorescence of a label on its surface only within the evanescent field. This has several advantages: (a) the unbound labeled species in solution remain unexcited and, thus, do not form a background signal; (b) measurements can be performed in absorbing or turbid media such as most biological solutions; and (c) background fluorescence of the serum can also be largely reduced.

Borisov and Wolfbeis

Figure 6. Capillary flow sensor. The capillary acts (a) as a sample compartment (or flow-through cell); (b) as an optical waveguide; and (c) as the solid support for immobilized antibodies. Light is coupled into and out of the capillary through grating couplers. Antibodies are deposited on the inner surface of the capillary. The fluorescence of labeled antibodies or antigens is interrogated by the evanescent wave mode. Reprinted with permission from Weigl, B. H.; Wolfbeis, O. S. Anal. Chem. 1994, 66, 332. Copyright 1994 American Chemical Society.

In contrast to surface plasmon resonance and interferometric sensors, the response of TIRF immunosensors does not depend on the mass of an analyte, which makes possible detection of even small haptens. If planar waveguides are used, the fluorescence is typically collected perpendicularly to the surface of the waveguide. In most fiber-optic biosensors, the emission is, however, collected at the distal end of the optical waveguide, as shown in Figure 5. The use of capillaries284,285 for optical immunosensing is very attractive because a capillary can not only guide fluid but also light in its wall (see Figure 6). Usually, the excitation light is introduced at the end of the capillary and propagates on the inner surface as an evanescent wave. When the evanescent excitation generates a signal from an antigenantibody-fluorophore complex, the emitted light is coupled into and propagates along the capillary. At the distal end, a grating can be used to couple the light out of the waveguide into a photodetector. When the active surface area is increased inside the capillaries, the fluorescence signal integrates over their length, but the electronic background noise remains constant. Sensitivity of the sensor, thus, is significantly improved. The group of Seeger71 has designed an immunosensor that exploits changes in the supercritical angle of the fluorescence (SAF) of molecules bound to a glass surface. Because the detection volume in the aqueous sample is significantly reduced in this technique, bulk (i.e., background) fluorescence from solution is strongly suppressed. The SAF signal can be captured by a parabolic glass lens, thus leading to high spatial collection efficiency and detection sensitivity. As little as 2 pM concentrations of labeled rabbit IgG could be detected in a direct immunoassay format. The sensitivity could be further improved by using tight focusing and smaller excitation spots. In this case, however, rapid photobleaching is a serious limitation. On the other side, spot diameters of 60 µm allow for up to 200 measurements with photodegradation not exceeding 1%. This is more than adequate to obtain a smooth response plot. The group also reported on a confocal reader for biochip screening and fluorescence microscopy.286

4.4. Immobilization of Antibodies on Sensor Surfaces, and Nonspecific Protein Binding There are several widely used methods for immobilization of large proteins such as antibodies on solid support. One is

Optical Biosensors

based on the creation of a covalent linkage between the support and the protein, often via a spacer group. The surface of a support is rendered reactive with groups such as -COOH, maleinimide, -NH2, or, less often, iodoacetamide groups, isothiocyanate groups, or boronic acid. These can be reacted with amino groups, thiol groups, carboxy groups, or saccharide groups of antibodies to form covalent chemical bonds. Glass surfaces and metal oxide nanoparticles are often derivatized using silyl reagents (aminopropyltrimethoxysilane being a typical example), while gold surfaces are derivatized with thiolated reagents of the type HS-(CH2)n-X, where X is one of the reactive groups given before. Such thiols readily bind to gold to form self-assembled monolayers on its surface. Quantum dots based on metal (Zn, Cd) sulfides and selenides also are surface-modified first by making use of an appropriate thiol chemistry. Plastic materials are more difficult to derivatize unless they contain intrinsic chemical groups such as in the poly(acrylonitrile-co-acrylic acid) copolymers.287 Another important method is based on the strong affinity of biotin to (strept)avidin.67, 68, 246, 270,288 Neutravidin also has been used recently.289-291 They all have four binding sites for biotin. Typically, the surface of a biosensor is modified (as described in the previous paragraph) by introducing biotin groups, in the overwhelming majority by using di- or tri(ethylene glycol)-modified biotin of the chemical structure

biotin-CONH-(CH2-CH2-O)n-CH2-CH2-NH2 (where n ) 2 or 3). Any protein that has been modified with (strept)avidin will strongly bind to such a surface. The affinity (binding) constants of the resulting noncovalently linked conjugates are in the order of 1012 to 1014 depending on the protein and the surface (or particle) used. One may wonder why the rather affordable avidin (a glycoprotein) is used so much less often than the rather expensive streptavidin (not a glycoprotein). On the other side, streptavidin has an isoelectric point (pI) of 5 and is less prone to nonspecific binding as compared to avidin with its pI of 10.5. Proteins also can be immobilized via a polyhistidine tag. The method is based on an amino acid motif in proteins that consists of at least six histidine (His) residues. It also is known as hexahistidine tagging, or 6xHis-tag, or by the tradename His-tag. The polyhistidine tag can be used for the immobilization of proteins on a nickel- or cobalt-coated microtiter plate, on glass, or on another protein. The most simple way to add a poly-His unit to a protein is to insert a protein DNA in a vector encoding a His-tag so that it will be automatically attached to one of its ends. The other technique is to perform a PCR with primers that have repetitive histidine codons next to the start or stop codon in addition to several (16 or more) bases encoding specifically to the protein to be tagged. Another widely used method (with particular applicability to the immobilization of antibodies) is based on the use of protein A. This is a 40-60 kD surface protein that strongly binds to immunoglobulins from many mammalian species. Specifically, it binds to the Fc region through interaction with the heavy chain and, thus, does not strongly compromise its affinity to the respective antibody. Shriver-Lake et al.292 investigated different heterobifunctional cross-linkers for covalent attachment of antibodies through thiol-terminated silanes onto glass, silica, silicone, and other surfaces. A variety of cross-linkers were found to be suitable for effective immobilization of antibodies. The

Chemical Reviews, 2008, Vol. 108, No. 2 441

use of carbohydrate-reactive cross-linkers resulted in immobilized antibodies having higher activities than when using reactive succinimide residue but required a more complex procedure, which implied the risk of denaturation of some antibodies. Preininger et al.293 have investigated three different types of polymer supports used for immobilization of antigens and antibodies with respect to specific binding and nonspecific binding (NSB, better referred to as nonspecific adsorption) and regeneration of the sensor. Interaction of rhodaminelabeled anti-human IgG with immobilized human IgG was used in a direct assay format for investigation of specific binding. The degree of NSB was determined using antihuman IgG and was found to be quite high (∼80% of the level of specific binding) for human IgG immobilized onto a sol-gel support. Carboxylated poly(vinyl chloride) (PVC) support showed ∼50% of nonspecific absorption, while the NSB to polystyrene was negligible. However, only 35% of initial signal was regenerated when polystyrene support was used, while the regeneration level of 50% was possible for carboxy-modified PVC. Liu et al.294 have shown that, when using polystyrene fibers, the extent of NSB can be significantly reduced by introducing a spacer between the polystyrene surface and the photoimmobilized antibody. Poly(ethylene glycol) crosslinkers with five ethylene glycol units decrease the degree of NSB dramatically, and additional treatment of the surface with BSA eliminates it completely. The authors also showed that the “stepwise” protocol of sandwich assay formats resulted in a much higher sensitivity than when using the more convenient “premix” protocol. NSB also can be significantly reduced by using a dextrane linker.295 Carboxymethylated dextrane was attached to the surface of a fiber-optic waveguide whose surface was treated with aminopropyltriethoxysilane; this was followed by carboxamide formation using activation with EDC and formation of a reactive NHS ester. The same method was applied to covalently immobilize an antibody to dextrane. The amount of NSB was shown to be only 2% of the level obtained for the untreated glass chip. Different immobilization techniques for glass fibers were investigated by Tedeschi et al.296 Immobilization of antibodies via glycidyloxypropyl-trimethoxysilane-dextrane resulted in the highest density of active sites. The Langmuir-Blodgett (LB) technique represents an alternative for immobilization of antibodies.255 Protein A has a specific affinity for a specific section of IgG and can be prepared as a stable monolayer by the LB film technique. Such a monolayer was immobilized on an alkylsilanized hydrophobic synthetic quartz plate. Anti-human IgG antibody was self-assembled on the protein A film. Rabbit IgG labeled with fluorescein isothiocyanate was used in a competitive assay for determination of human IgG over the analytical range from 10-4 to 10-1 g/L. Anderson et al.297 showed that the sensitivity of fluorescent immunoassays for determination of antigens was similar when the antibody was covalently attached to the support or via protein A.

4.5. Specific Examples of Immunosensors 4.5.1. Biosensors for Proteins and Antibodies Barnard and Walt251 developed a kind of reversible immunosensor for continuous measurements of IgG over a prolonged period of time using a controlled-release system. The fluorescein-labeled antibody and the Texas Red-labeled

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antigen were separately incorporated into an poly(ethyleneco-vinyl acetate) matrix, allowing for controlled release of the components. The analyzed IgG from the media and the two released biocomponents diffuse into the reaction chamber, where the competitive immunoreaction occurs. FRET is observed in the absence of IgG but is suppressed once the complex between fluorescein-labeled antibody and unlabeled IgG has been formed. Fluorescence intensity is monitored with the help of an optical fiber located above the reaction chamber. The steady-state rate of release is achieved after 2 days, allowing continuous monitoring of 0-500 mg/L of IgG over a period of 1 month. The approach, thus, can be applied to situations where continuous monitoring of analyte is required over prolonged periods of time and where rapid response (limited here by the diffusion of the analyte from the bulk solution or the release rate) is not an issue. Lepesheva et al.248 developed a FRET assay for hemoproteins in the analytical range from 10 nM to 10 µM. In the direct biosensor, fluorescein-labeled anti-hemoprotein IgG was absorbed onto a LB monolayer contained on a quartz support. Fluorescence intensity decreased in the presence of hemoprotein as the result of the quenching of the fluorescence of the label by heme. Luminescent quantum dots are viable optical markers44 and have been used in a direct assay for IgG. Protein A was labeled with CdSe/Zn Q-dots with a fluorescence λmax of 655 nm and then was immobilized at the tip of an optical fiber. Once the immunoreaction with IgG occurs, a decrease in fluorescence intensity is observed as a result of FRET from the Q-dot to the bound protein. Enzyme-linked immunosorbent assays employ enzymes as labels. Several attempts were made to design biosensors based on the use of enzyme labels. Papkovsky et al.60 demonstrated the feasibility of an enzyme-linked biosensor with the IgG-anti-IgG model system. Mouse IgG antigen was absorbed onto a surface of a glass fiber membrane combined with an optical oxygen transducer (a phosphorescent metalloporphyrin contained in polystyrene). Mouse IgG was detected in a sandwich ELISA using anti-mouse IgG antibodies labeled with glucose oxidase (GOx) as the secondary antibody. The amount of the antigen was quantified by measurement of the consumption of oxygen that results from the enzymatic reaction in the presence of glucose. The (rather long) luminescence decay time of the oxygen probe was monitored. This is in contrast to immunosensors based on the measurement of fluorescence intensity. A glass cover was used to limit oxygen access, and this significantly improved sensitivity, the LOD being 0.5 µg/L. Lactate dehydrogenase (LDH) was detected analogously using anti-LDH antibodies labeled with GOx. As little as 2.5 µg/L ()10 pM) of LDH could be sensed. The assay time was 1 h. Kim et al.298 made use of microbeads (made from a modified acrylamide) of ∼1 µm diameter that were assembled onto an amino-functionalized glass surface. A network of biotin and anti-biotin couples was attached to the beads (that also act as microlenses) via photopolymerization with aminobenzophenone. In the absence of the analyte, the immobilized antigens and antibodies interact with each other, which results in microspheres that are in the “on”state (Figure 7). The interactions between the attached antigens and antibodies are disrupted when a sample containing antigen (biocytin) is introduced. The microlenses are transformed from the “on”-state to the “off”-state as a result

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Figure 7. Microscope images of a microlens-based optical immunosensor array. Its sensitivity can be tuned by changing the concentration of the antibody used in the cross-linking stage. The concentrations of antibiotin were (a) 6.7 µM, (b) 2 µM, (c) 1 µM, and (d) 0.6 µM. With decreasing concentrations of antibiotin, the microlenses are inversed at lower concentration of biocytin. Reprinted with permission from Kim, J.; Singh, N.; Lyon, L. A. Angew. Chem., Int. Ed. 2006, 45, 1446. Copyright 2006 Wiley.

of gel swelling once the concentration of the analyte exceeds a certain critical value. The changes can be monitored microscopically. The sensitivity can be tuned by changing the concentration of the antibody used in the photochemical cross-linking stage. When the hydrogel microlenses are prepared with an excess of binding pairs above the critical point, they swell only after a suitably large number of displacement events have occurred. However, if the number of cross-linked units is just slightly above this critical point, only a few displacement events will result in gel swelling. The system was demonstrated to be fully reversible as the microlenses return to the initial “on” state when the antigen is removed by washing the sensor with phosphate-buffered saline.

4.5.2. Biosensors for Toxins In the integrating capillary biosensor described by the group of Ligler et al.,271 antibodies (anti-mouse IgG or sheep anti-staphylococcal enterotoxin B (SEB) were coated onto the entire inner surface of the capillary. Immunosensing of mouse IgG and SEB was accomplished in a sandwich format using antibodies labeled with Cy-5. Compared to conventional fiber-optic biosensors and planar waveguide-based about ∼2 orders of magnitude (40 and 30 ng/L for IgG and SEB, respectively). The analytical range of the sensor is from 40 to 300 and from 30 to 400 ng/L of IgG and SEB, respectively. Moreover, multianalyte detection can be attained by passing the sample through multiple capillaries, each coated with a different antibody, either sequentially or in parallel, depending on the amount of sample available. An ELISA type of biosensor for antibodies against cholera toxin B (CTB) was developed by Konry et al.299 An electroconductive surface was created on a fiber-optic waveguide by coating it with indium tin oxide to allow surface electropolymerization of biotin-pyrrole monomers. Biotin-conjugated CTB was attached to the surface using avidin. Anti-CTB was quantified via a competitive assay format in which the sensor was incubated first with the analyzed antibodies and then with goat anti-rabbit IgG labeled with horseradish peroxidase. Fibers were then immersed into a solution of luminol and oxidizing agent (a standard kit), and chemiluminescence was monitored. The

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LOD of the sensor is said to be 1 × 106 “titers”. The total assay time was 65 min. In another biosensor for anti-CTB, the antigen was attached to the surface of the optical fiber via a photoreactive benzophenone derivative,300 and the same LOD was achieved. An analogous system was developed for detection of antibodies against anti-West Nile virus IgG,301 and the LOD was similar. A 2-fold increase in sensitivity was observed compared to a chemiluminescent ELISA, presumably because light emission occurs near the optical fiber, which enhances the efficiency of light collection.

4.5.3. Biosensors for Drugs A capillary displacement-type immunosensor for the drug paclitaxel was designed by Sheikh and Mulchandani.273 The anti-paclitaxel antibodies were attached via protein A to the silanized inner walls of a glass capillary, and the binding sites were saturated with rhodamine-labeled paclitaxel. Although fluorescence intensity of the displaced labeled antibodies was monitored in another chamber, thus taking no advantage of the light-guiding properties of the capillary itself, the sensor still proved to be very sensitive. In fact, the detection limits were 1 and 4 ng/mL at flow rates 0.1 and 1 mL/min, respectively. The assay time ranged from 2 min at 1 mL/min to 8 min at 0.1 mL/min. Moreover, the regeneration of the capillary column was possible without affecting the performance of the biosensor. Anderson and Miller41 developed a homogeneous immunoassay for the anticonvulsant drug phenytoin where Bphycoerythrin-labeled phenytoin and Texas Red-labeled antibody were contained in a 200 µm cellulose dialysis tube connected to the distal end of an optical fiber (see Figure 8). The two species form a complex in which fluorescence resonance energy transfer (FRET) occurs from phycoerythrin to Texas Red. The interaction thus results in the quenching of the fluorescence of phycoerythrin. Phenytoin is small enough to diffuse through the dialysis membrane, where it displaces some of the phycoerythrin-phenytoin conjugates in the complex. The increase in the fluorescence of phycoerythrin is, thus, proportional to the concentration of phenytoin. This biosensor, notably operating in a fully reversible way(!), was suitable for determination of phenytoin with an LOD of 5 µM and a measurement time of 15 min. Later, the system was optimized to detect phenytoin in concentrations as low as 1 µM.49 A similar system was used for determination of theophylline.42 This biosensor design is of wider interest because full reversibility is achieved. However, the system is applicable only to antibodies having high reverse rate constants, which is not usually the case (see Figure 8).

Figure 8. Schematic representation of the reversible competitive immunosensor for phenytoin. Phenytoin and the phycoerythrinphenytoin conjugates competitively bind to the antiphenytoin-TR complex. Redrawn with permission from Anderson, F. P.; Miller, W. G. Clin. Chem. 1988, 34, 1417. Copyright 1988 American Association for Clinical Chemistry.

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4.5.4. Biosensors for Bacteria Cells Ko and Grant developed a FRET-based immunosensor for determination of Salmonella typhimurium.43 S. typhimurium antibody was labeled with a donor dye, while protein G was labeled with an acceptor. Both were immobilized on the surface of an optical fiber. In the absence of the antigen, very little fluorescence is observed from the acceptor. FRET occurs as a result of conformational changes when the antigen binds to the antibody. The ratio of the fluorescences of donor and acceptor serves as the analytical information. Concentrations of the bacterium as low as 1 × 105 cfu/L can be detected in 5 min. Antimicrobial peptides (AMPs) were recently shown to be suitable as recognition elements for microbial cells. AMPs are produced by many organisms for protection against invasion of harmful microbes. AMPs recognize microbes by interacting with their membranes, which then are destroyed. Several immunosensors reported recently302,303 rely on this recognition ability of AMPs. A direct immunosensor302 for Escherichia coli and Salmonella typhimurium was demonstrated to work with the AMP magainin I, which was either covalently attached to the patterned microscope slide surface or immobilized via biotin-avidin chemistry. Cy-5-labeled cells could be detected with detection limits comparable to analogous antibody-based assays. The LODs for E. coli and S. typhimurium were 1.6 × 105 and 6.5 × 104 cfu/mL, respectively, in the case of covalently immobilized magainin, and 6.8 × 105 and 5.6 × 105 cfu/mL, respectively, for the AMP immobilized via biotin-avidin. The assay time was 70 min. AMPs also were applied in the more-practical sandwich types of assays.303 Here, the immobilized AMPs were used to capture the unlabeled targets, while detection of bound cells was accomplished using fluorescently labeled antibodies. A significant degree of nonspecific binding was found in the case of tracer antibodies labeled with Cy-5 and Alexa Fluor 647 dyes. Replacement of the marker to Cy-3 was found to significantly reduce the amount of nonspecific binding. It was also found that high peptide density was necessary for optimal results. Limits of detection for E. coli and S. typhimurium were 5 × 104 and 1 × 105 cfu/mL, respectively, when magainin I was used, and 2-10-fold higher with other peptides. Martinez et al.304 reported on a biosensor for the protective antigen (PA) and for cellular components of Bacillus anthracis using SiONx waveguides. The sensor can detect 83 µg/L (i.e., 1 picomolar concentrations) of PA in a complex fluid within 10 min when operated in a sandwich assay format, but it possibly can become even more sensitive if interferences by background fluorescence and nonspecific binding can be further reduced. When whole cells are monitored immunologically by optical methods, the use of ultrasonic standing waves significantly improves the sensitivity of a biosensor. Zourob et al.72 showed that ultrasonication for 3 min enhanced sensitivity of detection of Bacillus subtilis cells by 2 orders of magnitude because the diffusion-limited capture rate is replaced by much faster cell movement. Rabbit anti-B. subtilis antibodies were immobilized on the surface of a metal-clad leaky waveguide. Evanescent light-induced scattering was detected by a CCD camera. Obviously, fluorescent labeling was not required, and the analytical range of the sensor operated in the direct assay format was from 1 × 106 to 1 × 1012 cfu/L.

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4.5.5. Biosensors for Pesticides A biosensor for the pesticide atrazine made use of polystyrene nanobeads dyed with a luminescent europium(III) chelate.278 Beads containing carboxy groups were covalently coupled to atrazine antibodies, which, in turn, were used in a binding inhibition assay format. An indium tin oxide (ITO) waveguide was used for immobilization of the capture analyte derivative to minimize nonspecific binding of the beads. Particles with diameters of 107, 304, and 396 nm were investigated in some detail. A decrease in their size resulted in faster binding but did not increase the assay sensitivity, which was comparable with the sensitivity of a known ELISA for atrazine. The use of such labels (having decay times of 0.1-1 ms) enables an almost complete elimination of background fluorescence by applying time-resolved measurements on the ITO waveguide.

4.5.6. Multianalyte Biosensors Lately, significant effort was devoted to the development of automatted biosensor devices capable of simultaneous immunosensing of several analytes of importance in environmental monitoring but, in particular, for the determination of (bio)chemical warfare agents and explosives. A compact, portable, multichannel fiber-optic instrument named MANTIS (of 5.5 kg weight)270 was reported to be capable of automatically transporting sample, buffer, and labeled antibodies to fibers and to perform fluorescence measurements. It enables four fluorescence immunoassays to be performed simultaneously on the surface of miniaturized polystyrene fiber-optic probes. The device was demonstrated to work for determination of staphylococcal enterotoxin B (SEB). Antibodies to SEB were immobilized on the surface of a polystyrene waveguide through avidin-biotin bridging. Cy5-labeled SEB antibodies were used in the sandwich assay format. The analytical range of the biosensor was from 10 µg/L to 1 mg/L, and the time required was 10 min. The improved version of the MANTIS device (termed RAPTOR) was shown to be able to determine 105 cfu/mL of Bacillus globigii, 107 cfu/mL of Erwinia herbicola, and 109 pfu/mL of MS2 coli phages simultaneously.305 Analogously, the RAPTOR system can determine staphylococcal enterotoxin B, ricin, Francisella tularensis, and Bacillus globigii simultaneously.305 In other work,306 plastic capillaries with immobilized antibodies were used for simultaneous determination of hormones prolactin, follitropin, and human chorionic gonadotropin (hCG). Fluorescein-labeled antibodies were used in a sandwich assay format. The detection limits were 1.3 µg/L, 2.3 IU/L, and 3.6 IU/L for prolactin, follitropin, and hCG, respectively. Feldstein and others at the Naval Research Laboratory307 developed an automatted array biosensor for determination of biological warfare agents. The central element of the sensor is a planar optical waveguide (Figure 9) used for direct excitation of antibodies that are bound to the waveguide surface within the penetration depth of the evanescent field. A physically isolated patterning method has been developed to manufacture an array of recognition elements (each ∼1 mm2 in size). In this technique, the multichannel patterning cell is placed on the prefunctionalized surface of the waveguide. Then, the recognition species (e.g., antibodies) are introduced into appropriate channels and are patterned on the waveguide surface during incubation, after which the cells are removed. The sensor is then used in a sandwich format.

Figure 9. Optical system of an array biosensor. Light from a diode laser is coupled into the waveguide. Fluorescence from the waveguide surface is focused by a graded index lens array (GRIN) through optical filters onto a Peltier-cooled CCD imaging array. Reprinted with permission from Feldstein, M. J.; Golden, J. P.; Rowe, C. A.; MacCraith, B. D.; Ligler, F. S. J. Biomed. MicrodeV. 1999, 1, 139. Copyright 1999 Springer Science and Business Media.

Simultaneous determination of the explosives 2,4,6trinitrotoluene (TNT) and hexahydro-1,3,5-trinitro-1,3,5triazine (RDX) was performed by a system named “Analyte 2000” developed at the Naval Research Laboratory.259 Two probes for determination of TNT were coupled with two probes for RDX. Cy-5-labeled analyte derivatives were used in a competitive immunoassay format to determine as low as 5 µg/L of TNT and 2.5 µg/L of RDX. Only a minimal cross-reactivity for the two haptens was observed in the multianalyte immunosensor, which was, therefore, capable of analyzing samples containing mixtures of the two compounds. The same array biosensor307 was used (a) for simultaneous screening of human serum for antibodies against bacterial and viral antigens including Staphylococcus aureus enterotoxin B, tetanus toxin, diphtheria toxin, and hepatitis B, with detection limits from 0.2 to 3 µg/mL and an LOD as low as ∼100 fg;308 (b) for the mycotoxin deoxynivalenol (LOD ) 0.2 ng/mL in buffer);309 (c) for simultaneous determination of large food pathogens such as Campylobacter jejuni (LOD ) 500 cfu/mL) and of small toxins such as aflatoxin B1 (LOD ) 0.3 ng/mL);310 (d) for staphylococcal enterotoxin B and botulinum toxin A (with LODs of 0.1 and 20 ng/mL, respectively);311 (e) for ochratoxin A in cereals and beverages (LODs ranging from 4 to 100 ng/g);312 (f) for Salmonella typhimurium (LOD ) 8 × 104 cfu/mL within 15 min and 8 × 104 cfu/mL within 1 h);313 (g) for the aggressive Escherichia coli species O157/H7 in food samples (with LOD of 5 × 103 cfu/mL in buffer and 1-5 × 104 cfu/mL in spiked food matrixes, and an assay time of 30 min only);314 and (h) for Campylobacter and Shigella species in food (LOD ) 9.7 × 102 and 4.9 × 104 cfu/mL, respectively).315 These articles reveal that the standard 6 × 6 array sensor can be used to analyze six samples for up to six different analytes. Taitt et al. showed316 that the same format is suitable for analyzing a single sample for 36 different analytes by using complementary mixtures of capture and tracer antibodies. Mixtures were optimized to allow detection of closely related species without significant cross-sensitivity. The approach was demonstrated to work when analyzing a sample for 9 targets with a simple 3 × 3 array. The only limitation of the approach is that the quantity of reagents needed increases significantly. Several other array biosensors suitable for simultaneous immunosensing were developed. The RIANA (river analyzer)280 array biosensor is based on a TIRF arrangement that was combined with a flow injection technique for

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automated and reproducible fluid handling. Herbicides can be detected by a heterogeneous binding inhibition immunoassay. Analyte derivatives were attached via an aminodextrane linkage to defined detection spots (L 3 mm) located on a glass waveguide, with a 2.5 mm distance between the spots. Analyte-specific antibodies labeled with a fluorescent marker were preincubated with analyte present in the sample, and the remaining (nonblocked) antibodies were then selectively attached to the spots. Each test cycle includes regeneration with hydrochloric acid of pH 1.9 and washing with an acetonitrile/water/propionic acid (49/50/1) mixture and is finished within 15 min. Because the glass surface included three sensitive spots, simultaneous detection of three analytes was possible. The following pesticides were detected: atrazine, simazine, isoproturon, alachior, 2,4-dinitrophenoxyacetic acid, and pentachlorophenol. The respective LODs were 0.03, 0.03, 0.11, 0.07, 0.07, and 4.23 mg/L. In related work,277 the pollutants atrazine and isoproturon and the hormone and endocrine disruptor estrone were determined simultaneously with LODs of 0.155, 0.046, and 0.084 µg/L using the binding inhibition assay format. No interference in the analysis of target compounds was observed upon simultaneous quantification. It was shown later317 that much lower concentrations of the pollutants can be detected by the RIANA biosensor. The highest standard deviation observed at very low analyte concentration results from the inaccuracy of the dilution procedures that require up to 11 steps. Much lower standard deviations were observed when stock solutions were used for each concentration. The errors resulting from dilutions can be compensated for by using a statistical method. The LODs for atrazine, estrone, and isoproturon were 0.002, 0.019, and 0.016 µg/L, which is ∼1 order of magnitude lower than if the statistical method was not applied. The RIANA array biosensor also was employed for monitoring testosterone in water samples (LOD 0.2 ng/L),282 progesterone in water (LOD 0.37 ng/L318 and 0.2 ng/L319), progesterone in milk (LOD 45 ng/L),319 the pesticide propanil in aqueous samples (LOD 0.6 ng/L),320 and other antibiotics, hormones, and endocrine-disrupting chemicals with similar LODs.321 A completely different approach toward array immunosensors was developed by Rissin and Walt.322 The underlying scheme of this type of microsensor is similar to the one used in DNA array sensors that were developed by the same group. In a typical example, antibodies to lactoferrin and IgA were covalently immobilized on the surface of 3 µm sized poly(methylstyrene) microspheres. These, in turn, were positioned onto the array of ∼50 000 individual optical fibers. A luminescent europium(III) chelate in two different concentrations was applied for encoding purpose, i.e., to establish the position of the two types of microbeads. Lactoferrin and IgA were determined in a sandwich assay format, with the secondary antibodies being labeled with Alexa Fluor 546. IgA can be determined in concentrations between 700 pM and 100 nM, while for lactoferrin the range is between 385 pM and 10 nM. While simultaneous determination of only two analytes was demonstrated, the approach is likely to be suitable for multianalyte sensing. In conclusion, it can be stated that immunosensors (a) are versatile because they enable the determination of highly different species that range from haptens to viruses and cells; (b) display excellent sensitivity and have very low limits of detection; (c) are highly specific; (d) act irreversibly and,

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therefore, are not suitable for continuous sensing in most cases; (e) are less prone to poisoning than enzyme-based biosensors; (f) are thermally labile and frost-sensitive in aqueous solutions; and (g) can be multiplexed.

5. Biosensors Based on Ligand−Receptor Interactions This type of affinity biosensor makes use of specific interactions between a biological receptor and an analyte. Reports on receptor-based biosensors are less numerous than those on immunosensors for the following reasons. This is mainly due to the fact that working with receptors is limited by the facts that (a) their isolation often is tedious ands unlike the production of antibodiessrequires individual protocols for each receptor; (b) they need the natural (cellular) environment to function best; (c) they are not stable over time; and (d) their molecular diversity requires an individual labeling protocol for each single species.

5.1. Receptor-Based Biosensors for Saccharides and Glycoproteins The first optical receptor-based biosensors were developed for sensing glucose and made use of its specific interaction with the glucose-binding protein concanavalin A (ConA).323-325 The approach (see Figure 8 for a closely related scheme) is similar to competitive immunoassays, with the notable exception that it acts reversibly. ConA is immobilized on the inner wall of a mm-sized hollow dialysis chamber via a 1,6-hexanediimine glutaraldehyde spacer. The chamber is placed at the distal end of an optical fiber. Fluoresceinlabeled dextrane is contained in the solution filling the chamber. In the absence of glucose, it binds to ConA. Unlike the large ConA, glucose can freely diffuse through the membrane and bind to ConA, and this results in the displacement of dextrane. The released fluorescein-labeled dextrane is distributed within the chamber and “seen” by the fiber if located in the cone determined by the numerical aperture. Because the walls of the chamber are located out of the aperture of the fiber, no fluorescence is registered in the absence of glucose. The sensor operates in the 0-50 mM glucose range and has a response time of ∼7 min. A FRET sensor based on the same principle was designed later.48 Here, the interaction of fluorescein-labeled dextrane with rhodamine B-labeled ConA resulted in a decrease in fluorescence intensity because of the more efficient FRET. Aggregation of ConA is prevented by chemical modification of the protein with succinic anhydride followed by labeling with the rhodamine and results in more stable calibration plots.326 Concentrations of glucose as high as 0.08 M could be analyzed. Russell et al.327 provided another solution for the FRET system. Concanavalin A was labeled with the FRET acceptor dye tetramethylrhodamine isothiocyanate (to give TRITCConA) and covalently immobilized in photopolymerized poly(ethylene glycol) hydrogel spheres with an average diameter of ∼2 µm. The FRET donor dye fluorescein isothiocyanate dextrane (FITC dextrane) was physically entrapped in the hydrogel. It can bind TRITC-Con A in the absence of glucose, while in its presence, FITC-dextrane is liberated. An increase in its fluorescence is observed as a result. The dynamic range of the sensor was from 0 to 44 mM of glucose, and the response time is 10 min for a glucose concentration step from 0 to 11 mM.

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Figure 10. Schematic of an affinity glucose sensor (redrawn with permission from Ballerstadt, R.; Schultz, J. S. Anal. Chem. 2000, 72, 4185; Copyright 2000 American Chemical Society). (Left) The Alexa-488-ConA conjugate is bound to dextrane and, therefore, confined in the porous microbeads, which do not allow the excitation light to penetrate. (Right) As glucose diffuses through the dialysis membrane, it liberates Alexa-488-ConA conjugate so that fluorescence is enhanced a result of better exposure to the exciting light beam and because fluorescence emission is no longer screened off.

Ballerstadt and Schultz328 have further developed the ConA-based system, thereby providing a most elegant solution. The system is based on an inner filter effect. Highly porous Sephadex beads were colored with two red dyes (safranine and para-rosaniline), which were selected to block the excitation and emission spectra of the fluorescent Alexa 488-ConA conjugate (see Figure 10). The latter is bound to dextrane inside the beads but is liberated when glucose is present. Once located outside the beads, the conjugate is fully exposed to the excitation light and fluorescence intensity, therefore, increases. The sensor has a dynamic range from 0.2 to 30 mM of glucose, and the total signal change was much higher than that for the FRET-based system.327 Moreover, faster response times were accomplished (4-5 min). In other work,329 the IR-dye Alexa-647 was conjugated to ConA in order to make the sensor work in the IR, which is less prone to spectral interferences by the intrinsic fluorescence of serum samples. The authors also carefully investigated the long-term stability of the biosensors by monitoring their performance over a 4 month period. The sensors displayed an initial increase in fluorescence over the first 3-4 weeks, which later on gradually decreased with an approximately linear drop of 25% per month. The decrease in fluorescence was not due to denaturation of the ConA but rather due to leakage of the fluorescently labeled ConA through the interface between the outer sealant and the membrane. If this problem can be coped with, the sensors potentially are suitable for continuous usage for up to 1 year. The main drawback of all sensors based on the use of ConA is their poor selectivity because they also respond to many other carbohydrates. For example, binding of fructose is ∼3 times stronger than that of glucose. This is in distinct contrast to the high specificity of enzymatic glucose sensors. On the other side, they have lower limits of detection. Other schemes for affinity sensing of glucose make use of the E. coli glucose-binding protein (GBP) that can be gene engineered. In contrast to ConA, GBP binds glucose with very high selectivity. In fact, the affinity to other saccharides is 100-1000-fold weaker than that for glucose or galactose. These proteins are rather stable and can be stored several months without degradation of activity.330 Since GBP

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undergoes significant conformational changes upon binding of glucose, a polarity-sensitive fluorescent indicator introduced at specific positions can act as an optical transducer. The proof of principle for the case of glucose was demonstrated by Marvin and Hellinga,331 who used GBP labeled with acrylodan. A homogeneous assay for glucose was developed that operates in the micromolar concentration range (0-2 µM). However, it was much earlier shown that a similar scheme could be applicable to sensing of maltose using a fluorescently labeled maltose-binding protein.332 A 60-180% increase of fluorescence intensity was observed in the presence of maltose. Concentrations of maltose from 0 to 200 µM could be determined. Ge et al.330 showed that the sensitivity of such assays strongly depends on the brightness (Bs) of the label used (Bs ) ‚QY). In fact, GBP labeled with the thiol label 2-(4′iodoacetamidoanilino)naphthalene-6-sulfonate resulted in a working range from 0.3 to 10 µM of glucose. For comparison, the assay based on the glutamine binding protein (GlnBP) labeled with a much brighter acrylodan was performed in the analytical range from 0.03 to 3.2 µM. The response of the proteins is very fast in solution (<1 min) but, however, much slower in dialysis cassettes (5-12 min). The recovery in the dialysis cassettes (which have a volume of 1 mL) was unacceptably slow in taking several hours, but this can be possibly accelerated by using smaller volumes. A ratiometric assay also was reported333 where GBP was labeled with both an environmentally sensitive fluorophore acrylodan and a reference luminophore, a ruthenium(II) polypyridyl complex. Sensing of glutamine using GlnBP was also demonstrated by Dattelbaum and Lakowicz.63 The E. coli GlnBP was covalently modified with acrylodan and other environmentally sensitive probes and used in a homogeneous assay format to detect glutamine from 0.05 to 6.4 mM. Time-resolved and polarization-based sensing of glutamine also was demonstrated. No interference by glutamate was demonstrated, which is a common drawback of the enzymatic glutamine biosensors based on the enzyme glutaminase. Ye and Schultz53 engineered a novel glucose indicator protein (GIP) that makes use of FRET from the donor (green fluorescent protein, GFP) to the acceptor (yellow fluorescent protein, YFP). In the absence of glucose, the two fluorescent proteins are in close proximity so that FRET occurs. When excited at 395 nm (corresponding to the absorbance maximum of GFP), the emission from the YFP (peaking at 527 nm) is observed. The spatial separation between the two moieties increases when GIP binds glucose, and FRET is reduced. In the biosensor, a solution of GIP was brought into a hollow cellulose dialysis fiber (L 190 µm), which was placed into a microcuvette, and fluorescence intensity was monitored at two wavelengths and the ratio was determined. The sensor responded reversibly to glucose with response and recovery times of ∼100 s, although some drift in the baseline occurred. Glucose could be determined in concentrations between 0 and 10 µM. Chinnayelka and McShane54 have used an inactive form of glucose oxidase as a selective glucose-binding protein. The apoenzyme labeled with tetramethylrhodamine isothiocyanate was placed inside nanoengineered polymeric microcapsules together with fluorescein-labeled dextrane. The FRET that was observed in the absence of glucose was reduced when glucose diffused into the microcapsules and replaced labeled dextrane in a competitive way. The ratio of fluorescence intensities of the two labels was used as the

Optical Biosensors

analytical information. The sensor reversibly responded to glucose concentrations in the range from 0 to 40 mM, was fast (1-2 min), and was rather selective (a 5-10-fold higher sensitivity for glucose compared to other saccharides). The sensitivity of the system can be easily tuned by varying the concentration of the GOx-dextrane complex. Ogasawa et al.334 designed an affinity sensor for the riboflavin-binding protein (RBP). It exploits the intrinsic green fluorescence of riboflavin. Hydrophobic 3-octylriboflavin is adsorbed on an optical fiber whose surface was made lipophilic by modification with octadecyl groups. On exposure to a solution of RBP, it binds to the surface-immobilized octylriboflavine. The sensor can be renewed with good reproducibility by removing the RBP-riboflavine complex and loading the surface again with octylriboflavine. In fact, 15 individual sensors for RBP were prepared on a single optical fiber, which varied in performance by <5%. When RBP binds to the immobilized riboflavin, quenching of fluorescence is observed. As little as 0.73 µM of the RBP can be detected in 5 min. Medintz et al.55 prepared two kinds of biosensors for maltose, which make use of quantum dots (QDs) and a maltose-binding protein (MBP) from E. coli. The first type of the sugar-sensing nanoassemblies consisted of 560-nm donor QDs conjugated to 10 molecules of MBP (in average). The displaceable dextrane was labeled with a fluorescent acceptor dye. No FRET was observed when maltose replaced dextrane in the sugar-binding site. The second maltose biosensor consists of 530-nm donor QDs loaded with 10 cyanine-labeled MBP molecules. The protein-bound label served as a bridging acceptor/donor for ultimate energy transfer to maltose-displaceable labeled dextrane. Both sensors respond to maltose concentrations from 10 nM to 1 mM. Another nanobiosensor for maltose56 was reported that is based on the use of thiolhexanoate-capped CdSe nanoparticles (L 3.0-3.5 nm) conjugated to the MBP, which was labeled with a luminescent ruthenium(II) complex. Little fluorescence is detectable in the absence of maltose because of the electron transfer from the ruthenium complex to the nanobead. Fluorescence is enhanced as a result of conformational changes that occur upon binding of maltose. Glucose and lactose were shown not to interfere, while maltose could be sensed from 10 nM to 10 µM. An unusual approach toward receptor-based sensing of glucose in blood was proposed by Sanz et al.335 It is based on the fact that oxidation of hemoglobin (Hb, contained in the sample) by H2O2 (generated after addition of glucose oxidase GOx according to eq 1) results in distinct changes in the absorption spectrum of Hb. To make the assay operative, blood samples prepared without pretreatment and reactions of H2O2 with other blood components such as catalase need to be blocked (which is achieved by addition of azide). The activity of GOx should be high enough for the “chemistry” to work, given the fact that the activity of GOx is inhibited by azide by ∼30%, and low enough to avoid its interaction with hemoglobin. The linear response range of the system was from 0.1 to 30 mM of glucose.

5.2. Receptor-Based Biosensors for Inorganic Ions Biosensors for cations exploit the need of enzymes for certain ions in order to function. This results in two kinds of sensing schemes. The first is the effect of the catalytic ion on the actiVity of the enzyme (which is treated in section

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3). The second is the effect of the catalytic ion on the conformation of the enzyme (regardless of its activity). Respective sensors are treated in this section. In such sensors, an enzyme serves as a kind of chelator for certain metal ions, but with much higher selectivity than classical chelators such as EDTA. The ions usually are required as enzyme cofactors. For example, Thompson and Jones336 reported a highly sensitive zinc(II) biosensor, which makes use of the enzyme carbonic anhydrase (CA). The Zn2+ ion is its natural cofactor and is bound by CA with excellent selectivity. The apoenzyme prepared by dialysis was contained in a chamber together with the fluorescent probe dansylamide. The ionpermeable chamber was attached to an optical fiber. The probe does not bind to the apoenzyme and, thus, remains weakly fluorescent in water. However, when (practically irreversibly) bound to CA in the presence of Zn2+, the fluorescence intensity increases significantly. The analytical range of the sensor was from 40 to 1000 nM. The main drawback of the system is its moderate brightness and the limited applicability to real samples because UV excitation is required at ∼330 nm where background fluorescence of most samples is very strong. A more flexible system for detection of Zn2+ is based on fluorescence quenching.52 In the presence of the metal cation, the active site of CA labeled with the fluorescent donor fluorescein permits binding of the colored quencher azosulfamide. The fluorescence decay time (measured by phase fluorometry) was shown to decrease with increasing concentrations of zinc cation, which could be determined in the concentration range from 1 to 100 µM. A cobalt biosensor was reported51 that is based on the finding that the d-d absorption of Co2+ coordinated to CA labeled with a cyanine dye promoted radiationless FRET since the absorption of the Co2+ ion overlaps the emission of the label. The labeled apoenzyme was entrapped in a polyacrylamide gel positioned at the distal end of an optical fiber. The sensor was capable of sensing Co2+ in the concentration range from 0 to 20 µM with response times of a few minutes. The fact that certain variants of CA exhibit different selectivity to metal ions was used by Zeng et al.57 to design a sensor for Cu2+. Two variants of human CA II were labeled (with Oregon Green and Alexa Fluor 660, respectively) and immobilized at the distal end of an optical fiber. An ∼85% drop in fluorescence intensity and decay time (as measured by phase fluorometry) was observed upon binding of Cu2+. The analytical range of the system was from 0.1 to 100 pM. Ions such as Co2+ and Ni2+, and even Zn2+, were shown to interfere only if present in much higher concentration, because the affinity of the CAs to these ions is much lower than that for Cu2+. It was demonstrated that the sensor was suitable for real-time analysis of Cu2+ in seawater. The phosphate-binding protein isolated from Escherichia coli and labeled with the fluorophore acrylodan is potentially suitable for biosensing purposes337 by showing a 50% increase of fluorescence intensity in the presence of micromolar concentrations of phosphate. In another detection scheme for phosphate,338 two cysteine mutations were introduced into the phosphate-binding protein, allowing it to be labeled with two rhodamine fluorophores. When close to each other (in the absence of phosphate), the rhodamine molecules form a noncovalent and nonfluorescent dimer. A linear correlation between fluorescence intensity and the concentration of phosphate was observed in the range from 0 to 6 µM of the analyte. At saturation (concentrations of

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phosphate of 6 µM or higher), an 18-fold increase (at average) in fluorescence intensity was observed for the mutants investigated. A homogeneous assay for sulfate was reported that is based on fluorescently labeled sulfate binding protein.339 The fluorescence of the labels was quenched upon binding sulfate. The LODs varied from 30 to 200 nM depending on the label used. Sumner et al.340 have discovered that the wild form of red fluorescent protein (RFP) can reversibly bind Cu+ and Cu2+ ions with high selectivity and sensitivity. In fact, the quenching of fluorescence of the RFP by the heavy metal ions occurred even at 10 nM concentrations of the analytes. The RFP was found to be ∼107-fold more selective for copper over Mg2+ and Ca2+, with its fluorescence being virtually unaffected even by high concentrations of those. Fluorescence was recovered on addition of EDTA, albeit not to the initial level. A nanobiosensor for Cu2+ also was designed on the basis of RFP.341 The protein was immobilized in 80nm polyacrylamide nanobeads together with an inert reference dye. The ratio of the intensities under 488-nm excitation served as the analytical parameter. The signal change caused by Cu2+ returned to 95% of its original value within 3 min when EDTA was added, thus showing an almost complete reversibility of the sensor. As in the case of the homogeneous assay in solution, no interference by other metal ions was observed. The sensitivity to Cu2+ was independent of pH in the range from 6 to 8.5. However, the nanobiosensor was found to be much less responsive to Cu2+ at low pH. The analytical range for Cu2+ is from 0.2 to 50 µM, but sensitivity can be fine-tuned by varying the concentration of the nanobeads, thus generating a larger signal change per nanoparticle at the same concentration of the analyte.

5.3. Receptor-Based Biosensors for Gaseous Species A number of optical affinity biosensors make use of heme proteins, natural compounds that can weakly bind gases such as oxygen and carbon dioxide, and much more strongly bind carbon monoxide and nitrogen monoxide. Blyth et al.342 showed that the heme proteins cytochrome c, myoglobin, and hemoglobin (Hb) enable semiquantitative detection of CO and NO in aqueous medium. The heme proteins immobilized into a sol-gel matrix exhibited a distinct change of their absorption spectra upon coordination of NO and CO. Although the effect was reversible, desorption of gases took up to 2 h. However, fast regeneration was accomplished by using other reagents. For example, cytochrome c could be regenerated from its complex with NO by reduction with sodium dithionite, washing with buffer, and addition of potassium ferricyanide. Cytochrome c embedded in a sol-gel was demonstrated to sense NO in a gas phase, with a response time of 200 s and full recovery within 300 s.343 The sensor operated in the dynamic range from 1 to 25 ppm of NO. Reversible and fast micro- and nanosensors for NO were developed by Barker et al.344 Cytochrome c′ was immobilized on gold nanobeads (L 50 nm). Two ways of optical interrogation were reported. The first is to measure the intrinsic fluorescence of cytochrome c′, which changes in the presence of NO. The second is to measure the increase in the efficiency of FRET from the label (Oregon Green) to cytochrome (which is enhanced when the latter binds NO because of better spectral overlap between the emission spectrum of Oregon Green and the excitation spectrum of cytochrome). The response to NO is linear in the concentra-

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tion range from 20 µM to 1 mM. The sensors operate in a fully reversible way and have response and recovery times of <1 s. Although the LOD of the sensor was relatively high (20 µM), the limits of quantification are very low because only small volumes are required. Cytochrome c′ also was entrapped in polyacrylamide; however, binding of NO was found to be irreversible. The performance of such microsensors was further improved by using a reference fluorescent indicator and employing ratiometric measurements.345 The microsensor incorporated labeled with Oregon Green cytochrome c′ along with polystyrene nanobeads (L 40 nm) labeled with a fluorescent reference dye whose red emission (λmax ) 685 nm) allowed for ratiometric (two-wavelength) measurements. Compared to the previous work,344 the LOD of the sensor was improved to as low as 8 µM. Immobilization of the reference spheres was, however, not reproducible, and the ratio of fluorescence intensities and calibration plots, therefore, varied from sensor to sensor. The ratiometric sensors were employed to measure extracellular NO released by macrophages. Blyth et al.346 observed that cytochrome c′, when immobilized into a sol-gel, undergoes irreversible conformational changes (in 2-4 days), which lowers binding affinity of the metalloprotein. However, after these changes have occurred, the protein remains selective for NO and the calibration plots are well-reproducible. Analogous biosensors were prepared with the heme domain of soluble guanylate cyclase, the only protein receptor known for signal transduction involving in vivo produced NO and having many similarities to cytochrome c′, including a very low affinity for oxygen and a high affinity for nitric oxide.347 The LOD of the sensor was 1 µM. Reversible binding of oxygen to hemoglobin (Hb) was exploited by Zhujun and Seitz348 to design an oxygen biosensor. They showed that Hb immobilized on a Sephadex ion-exchange resin can sense oxygen in the dynamic range from 0 to 160 Torr. The shelf life of this reflectance-based biosensor is very short due to fast irreversible degradation of the immobilized Hb (within 2 days at room temperature or within 7 days at 4 °C). The sensor is, thus, hardly an alternative to conventional optical oxygen sensors, which are based on stable quenchable luminescent indicators.

5.4. Receptor-Based Biosensors for Toxins An evanescent-wave biosensor for R-bungarotoxin was designed by immobilizing the nicotinic acetylcholine receptor on an optical fiber.349 Fluorescein-labeled R-bungarotoxin was used in the competitive assay format. As little as 1 nM of the toxin could be detected within 5 min. Although the nonspecific binding was totally eliminated by addition of bovine serum albumin, the sensor was inhibited by agonists such as acetylcholine, nicotine, and carbamylcholine and by antagonists such as pancuronium and D-tubocurarine. No regeneration of the biosensor was possible. Song and Swanson50 developed biosensors for cholera toxin (CT). The bioreceptor ganglioside GM1 was incorporated into a biomimetic membrane surface (composed, e.g., of 9-octadecenoyl phosphatidylcholine), which, in turn, was spread onto glass microbeads. The labeled receptor molecules are homogeneously distributed in the lipid bilayer but aggregate in the presence of CT, which has five binding sites for GM1. As a result, fluorescence self-quenching is observed. Alternatively, the receptor molecules are labeled

Optical Biosensors

with a fluorescent donor, and a fluorescent acceptor dye, FRET, is observed in the presence of CT. Generally, the sensors respond to the concentration of toxin from 0 to 10 nM. Sensitivity and dynamic range can be tuned by varying the total concentration of the labeled GM1 in the membrane. Limits of detection as low as 0.05 nM and small dynamic range are associated with samples having low concentration of the receptor, while lower sensitivity and large dynamic range are found for samples with high concentration of the receptor. Boradipyrroles also were found to be viable labels in both types of the biosensor. No interference by albumin was observed in this case, whereas a significant nonspecific drop in fluorescence was observed for GM1 labeled with fluorescein. Many bacterial toxins, viruses, and bacteria target carbohydrate moieties on the surface of a cell so as to attach and gain entry into the cell. Ngundi et al.350 designed a monosaccharide-based array biosensor for detection of protein toxins. Arrays of N-acetyl galactosamine (GalNAc) and N-acetylneuraminic acid (Neu5Ac) derivatives were immobilized on the surface of a planar waveguide (similar to ref 307) and were used as receptors for protein toxins. These arrays were probed with fluorescently labeled bacterial cells and protein toxins. While Salmonella typhimurium, Listeria monocytogenes, Escherichia coli, and staphylococcal enterotoxin B did not bind to either of the monosaccharides, both cholera toxin and tetanus toxin bound to GalNAc and Neu5Ac and could be detected at concentrations down to 100 ng/mL. In conclusion, it can be stated that biosensors based on ligand-receptor interactions are (a) often highly specific; (b) sensitive in giving rather low limits of detection; (c) characterized by virtually irreversible response; (d) sensitive to environmental effects; (e) prone to poisoning; and (f) tedious to fabricate, in particular in terms of genetic modification and isolation of proteinic receptors.

6. Nucleic Acid Biosensors Such sensors (also referred to as DNA biosensors) represent an attractive alternative approach to immunological sensing of species such as bacteria. They take advantage of the exceptional long-term stability of nucleic acids and the high selectivity of the interaction of complementary chains of polynucleotides. Nucleotides and their polymers also can be synthesized easily. Typical examples of DNA biosensors are described in the following (see Figure 11).

Figure 11. Two fundamental forms of nucleic acid-based sensors: (a) conventional DNA sensor using a fluorescently labeled counter strand or a fluorogenic intercalator; (b) molecular beacon DNA sensor.

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6.1. Single DNA Sensors on Solid Supports and on Fiber-Optics The feasibility of optical nucleic acid biosensing was demonstrated by Graham et al.351 16- and 20-base oligonucleotides, but also long (204-base) oligonucleotide chains, were attached to the surface of an optical fiber as shown schematically in Figure 11a. Fluorescein-labeled complementary chains were shown to interact with complementary chains within ∼1 min as determined by TIRF spectroscopy (see section 4.3.2). Regeneration of the biosensor was accomplished by raising the temperature from 65 to 80 °C, resulting in complete dissociation of the bound duplex within 10-15 min. The analytical range of the sensor was from 0 to 200 nM of nucleotides. Piunno et al.352,353 immobilized a DNA sequence on an optical fiber by first activating the surface with a long-chain aliphatic spacer arm terminated with a nucleoside to which a longer chain was attached through automated step-by-step DNA synthesis. Detection of hybridized DNA at the fiber surface was achieved by treating it with a solution of the intercalator ethidium bromide. The sensor was regenerated by exposing it to a 85 °C hybridization buffer for 5 min. Total analysis time was <1 h, and the LOD was 86 ng/mL. The sensor showed reproducible results within 3 months of storage. Watts et al.354 immobilized biotinylated oligonucleotide sequences on a solid surface via streptavidin and used a resonant mirror technique (see section 4.3.2) for direct and rapid detection of hybridization. The lowest detectable concentration of the target 40-base nucleotide was 9.2 nM. Hybridization at the sensor surface was followed for 15 min, although a positive response was obvious within 30-60 s. Abel et al.355 compared the performances of DNA biosensors operating in the direct and competitive assay formats, respectively. A biotinylated capture probe was immobilized on a glass surface via avidin or streptavidin. A complementary fluorescein-labeled 16-base oligonucleotide could then be determined with an LOD of 0.2 pM. A competitive assay (using labeled and unlabeled nucleotides) resulted in a much higher LOD (1.1 nM). The use of poly(acrylic acid) sodium salt and Tween 20 reduced the nonspecific binding to 1-2% of the amount of specific binding. The signal loss during long-time measurements, i.e., after consecutive hybridization assays, can be described by a single-exponential function and, thus, compensated for. After 200 cycles, the net signal had decreased by 50%, corresponding to a signal variation of only 2.4% after correction for this signal loss. By using a 50% (w/w) aqueous urea solution for regeneration of the biosensor, the duration of an assay cycle was reduced to 3 min. Pilevar et al.356 used a near-IR cyanine dye (λexc ) 787 nm, λem ) 807 nm) as a label for an oligonucleotide sequence in order to make measurements outside the background fluorescence from natural compounds, which is substantial when using fluorescein labels. The feasibility of detecting bacterial cells using rRNA as the target was demonstrated in a solid-phase sandwich-type of assay where Helicobacter pylori rRNA was used along with IR dye-labeled detector oligonucleotide probe. The result indicates that this biosensor is capable of detecting H. pylori RNA at picomolar concentrations. A biosensor for detection of L-adenosine was developed by Kleinjung et al.357 An L-adenosine specific RNA was attached to an optical fiber via an avidin-biotin link.

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Fluorescein-labeled L-adenosine was used in a competitive assay format. The sensor responds to concentrations from 1 nM to 100 µM. Bagby et al.358 found that an intercalating thiazole orange derivative (TOMEHE) gives a 10-fold larger signal change over the commonly used ethidium bromide, thus providing higher sensitivity for hybridization events. TOMEHE, however, also showed a significant response to single-strained DNA and concentration-dependent phenomena at high loading with the dye. This limits the dynamic range over which TOMEHE can be used. A capillary sensor (see section 4.3.2.) for DNA also was reported.359 A capture DNA was attached to the inner walls of a capillary via biotin-streptavidin chemistry. The complimentary DNA sequence labeled with Alexa Fluor 532 can be sensed with a detection limit of 30 pg/mL.

6.2. DNA Arrays Ferguson et al.360 created a fiber array biosensor capable of simultaneous measurements of 7 DNA sequences. The optical fibers (each 200 µm in L) carrying the immobilized oligonucleotide probes were bundled to form a multiplex DNA sensor. The fluorescence intensity of all the fibers was imaged with a CCD camera. Up to 7 DNA sequences could be detected within 10 min with an LOD of 10 nM. As in the case of individual sensors, the array can be stored for prolonged periods (months) without loss of activity. Walt and co-workers demonstrated later361 that the actual detection limits of the array microsensor can be much smaller if microspheres are being used. Small sample volumes (10 µL for a 500 µm array) and higher local concentrations of the DNA enable further amplification. Fewer sensor numbers in the array also increase the signal because more target molecules hybridize per microsphere. By examining multiple identical sensors simultaneously, the signal-to-noise ratio can be improved by allowing incoherent noise to be signal averaged. The authors demonstrated that as few as 600 target DNA molecules (10-21 mol) can be detected. Although DNA at higher concentration can be assayed within 10-30 min, 17 h were necessary to complete hybridization and to achieve the lowest detection limits. The same group developed a method for encoding a set of randomly ordered functionalized microbeads (each bearing alkaline phospatase, avidin, or biotin) using luminescent dyes.362 This method formed the basis for a smart array system suitable for simultaneous detection of numerous DNA sequences.363 Polymer microbeads (L 3.1 µm) were dyed with various fluorophores such as Cy-5 and europium(III) complexes and functionalized with different oligonucleotide probes. Because microbeads with different concentrations of a single fluorophore are optically distinguishable, a total of 100 different beads could be prepared. A mixture of the beads was distributed over a distal end of a fiber bundle (L 500 µm, 6 000 individual fibers) so that each microbead occupied a single well (Figure 12). The position of the recognition elements was decoded by imaging with a CCD chip because each type of the beads has a characteristic emission wavelength and luminescence intensity. The signal was monitored after hybridization to fluorescein-labeled complementary oligonucleotides. Only 10 min are needed to determine 100 pM of oligonucleotides, but up to 17 h are needed for the lowest concentration (10 fM). In continuation of this methodology, Walt and co-workers have designed array biosensors for simultaneous determi-

Borisov and Wolfbeis

Figure 12. Overview of a fiber-optic array system. The functionalized beads occupy the micrometer-sized wells (1 bead/well) located on the tip of the optical fiber. The position of the beads is decoded by imaging its color. When immersed into a sample solution containing labeled DNA, a signal is observed only on the beads bearing the DNA probe complementary to the target in solution. Reproduced with permission from Ferguson, J. A.; Steemers, F. J.; Walt, D. R. Anal. Chem. 2000, 72, 5618. Copyright 2000 American Chemical Society.

nation of Bacillus anthracis, Yersinia pestis, Francisella tularensis, Brucella melitensis, Clostridium botulinum, Vaccinia virus, and a biological warfare agent simulant named Bacillus thuringiensis kurstaki.364 The replacement of the 20mer probes by 50-mer probes allowed for a high specificity of the array. The authors report LODs of “10 fM” (10 femtomolar concentrations) for B. anthracis, Y. pestis, Vaccinia virus, and B. thuringiensis kurstaki, and of 100 fM for B. mellitensis and C. botulinum. This is difficult to interpret since a bacillus does not have a molecular weight. It was also found that overlapping target sequences are partially complementary to the probe sequences. This can result in a nonspecific response, unfortunately. The use of multiple probes (at least two for each analyte) minimizes the potential possibility of false identification. The assay time was 30 min. The above DNA optical fiber array subsequently was coupled to a microfluidic system365 operated at a flow rate of 1 µL/min. This resulted in faster hybridization (15 min, compared to 30 min required for static measurements) and in ∼100-fold lower LODs (10 aM, compared to 1 pM as achieved in static measurements), which makes this approach highly advantageous. The systems described in refs 360367 are quite successful in commercial terms. The target rRNA of an algal bloom species can be determined with a microarray biosensor operating in a kind of “sandwich” assay format.366 The RNA to be analyzed interacts with a long capture probe, and the labeled tracer oligonucleotides then interacts with the residual free end. Fluorescently labeled oligonucleotides (acting as tracers that can capture nucleotide sequences) were coupled to the surface of microbeads positioned in the wells at the tips of optical fibers in an array. As few as 5 cells could be detected within 45 min, and the LOD of the rRNA is 4 × 104 molecules. In similar work,367 the microarray biosensor was used for detection of different Salmonella strains with LODs of 103-104 cfu/mL in pure samples and of 104-105 cfu/mL in the presence of interfering organisms. All investigated Salmonella strains were detectable, albeit with different sensitivities. Other common food pathogens were shown not to interfere at concentrations of 108 cfu/mL. The assay time was 1 h. Another approach of addressing and specifically depositing DNA was demonstrated by Swanson et al.,368 who designed a semiconductor biochip containing a microelectrode array. In order to immobilize a specific capture DNA probe at a

Optical Biosensors

certain location, an electric field was applied that causes the attraction of DNA fragments and their deposition at predetermined electrodes. The target DNA may be guided to the specific locations as well. The assay is performed in a sandwich format by exposing the array to the tracer DNA bearing a fluorescent label. If coupled to an integrated fluidic system, the following steps of the DNA analysis of E. coli can be automatically performed:369 (a) di-electrophoretic collection of bacteria; (b) DNA amplification and electronic DNA hybridization; and (c) fluorescence readout with the help of a CCD camera. The whole assay can be performed in 2.5 h. Rissin and Walt have demonstrated recently370 that the sensitivity of array biosensors can be significantly improved by employing an enzymatic amplification step similar to that known from ELISAs. To prove its feasibility, a biotinylated femtoliter array was incubated with a solution of streptavidin-β-galactosidase (SβG) conjugate, and then with a solution of the enzyme substrate, which generated fluorescent resorufin upon hydrolysis. Since the volume of the wells is very small, the limit of quantification for SβG was found to be 2.6 amol. No signal was measured in control experiments, in which the enzyme was bearing no streptavidin, or, alternatively, the surface of the fiber was not biotinylated.

6.3. Molecular Beacons in DNA Sensors Described first by Tyagi and Kramer,45 molecular beacons (MBs) have become an important tool for studies in genetics, disease mechanisms, and molecular interactions. MBs represent single-stranded types of oligonucleotide probes that possess a stem-and-loop structure (see Figure 11b). The stem is formed by the two ends of an MB containing complimentary nucleotides. A fluorophore attached to one end of the stem and a quencher attached to the other are in close proximity, and little or no fluorescence is observed. The loop portion of the molecule is responsible for reporting the specific complimentary oligonucleotide. Hybridization of a matching oligonucleotide to the loop portion results in conformational reorganization that brings the stem apart so that fluorescence is enhanced. This smart technique has the advantage that no labeling of other species is required (compared to, e.g., the competitive assay format). Following the work of Tyagi and Kramer,45 who used MBs in homogeneous solution, Fang et al.46 immobilized an MB onto the surface of a silica plate via avidin linkage to design a solid-state biosensor. A biotinylated MB was prepared that had a total of 28 bases, including 18 bases complementary to the sequence of interest and 5 base pairs for the stem. Tetramethyl rhodamine was selected as a fluorophore, and a modified azobenzene Dabcyl was selected as a quencher. A significant increase in fluorescence was observed upon addition of the complementary DNA both for the MB contained in homogeneous solution and in the immobilized form. In the control experiment, no effect was observed on addition of the noncomplementary DNA. The results indicated that the immobilized MB could be used to detect target DNA molecules in the subnanomolar range. Liu and Tan47 have investigated DNA sensing in more detail using a similar MB (also labeled with TMR and Dabcyl). The biotinylated MB was immobilized onto the surface of an optical fiber via a streptavidin bridge. A spacer group between the MB and the fiber substantially reduces steric hindrance and increases its mobility. A strong increase in fluorescence intensity was observed only upon hybridiza-

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tion with a complementary DNA, while the increase was much lower for the oligonucleotide strand having a single mismatch. The MB could be regenerated by immersing it into 90% formamide plus 10% tris/EDTA buffer mixture for 1 min. While a single regeneration cycle completely recovered the sensing properties of the MB, a few repeated regeneration cycles resulted in a significant drift. A 50% aqueous urea solution can be used for regeneration of DNA sensors as well.371 The authors also showed that higher ionic strength (IS) favored hybridization by decreasing electrostatic repulsion between the loop chain of an MB and the target DNA. Moreover, the initial reaction rate in the presence of divalent cations was ∼20 times higher than in the presence of monovalent ones at the same IS. The best results were obtained by using a high IS created by the divalent cation Mg(II). Under these conditions, 1-100 nM of DNA could be sensed in 10 min. Ueberfeld and Walt40 have designed an MB capable of virtually reversible binding of the target nucleotide. This can be achieved if the free energy of the duplex formation is of the same magnitude as the free energy of the stem formation. To do so, 6 oligonucleotides in the fully complementary 20mer loop were replaced by adenine moieties. The reversibility was indeed achieved, but only when working at a carefully adjusted temperature (38.8 °C), while working at 34 °C resulted in irreversible binding, and working at 43.8 °C resulted in the melting of the loop-target hybrid. In order to measure the efficiency of FRET, the stem of the MB was labeled with a fluorescent donor, while the acceptor dye was positioned between the stem and the loop sequence where it neither hinders stem formation nor inhibits target-loop hybridization. A two-wavelength ratiometric approach was made use of. This has the advantage that the ratio of the fluorescence intensities of the donor and the acceptor, respectively, is independent of the concentration of the labels. Du et al.372,373 attached a fluorescently labeled oligonucleotide to a thin gold surface to create an MB that requires no quencher on the second stem. In the absence of a complementary DNA, the label is in close proximity to the gold surface and its fluorescence is quenched. The intensity increased about 100-fold upon hybridization with a complementary DNA, which was detectable at concentrations from 0.2 to 3 µM. An 8-fold lower sensitivity was observed for a singly mismatched target. The sensor was, however, not suitable for multiple measurements in showing an ∼40% degradation of fluorescence intensity after each regeneration cycle.

6.4. Liposome-Based DNA Assays Some biosensors for nucleic acids make use of liposomes containing thousands of dye molecules and thus generating strong signal for even low nucleic acid concentrations so that quantitative reflectance (or even qualitative visual) measurements are possible. A typical biosensor of this type includes a capture oligonucleotide attached to the surface of a liposome. Biotin-streptavidin binding is employed to attach a reporter oligonucleotide to a liposome loaded with sulforhodamine B. In the presence of the target sequence, a sandwich is formed and the detection zone becomes colored. RNA from B. anthracis spores in concentrations from 0.1 pg/L to 1 ng/L could be detected in 90 min.374 As little as one B. anthracis spore is detected in 12 h. Ba¨umner et al.375 used a similar system for detection of RNA sequences from B. anthracis, C. parVum, and E. coli,

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but with the difference that the fluorescein-labeled capture oligonucleotides were immobilized on a polyethersulfone membrane via anti-fluorescein antibodies. Quantification of the RNA was possible between 10 nM and 1 µM by simultaneously incubating the RNA with streptavidin-labeled liposomes, biotinylated capture oligonucleotide, and the membrane containing a target sequence. The assay time was 20-30 min. An analogous RNA biosensor was used for detection of Dengue virus in blood samples.376

6.5. Aptamer-Based DNA Sensing Aptamers are nucleic acid species that have been evolutionary engineered through in vitro selection to bind various molecular targets such as haptens, proteins, nucleic acids, and even cells, tissues, and organisms. Lee and Walt377 have designed an aptamer biosensor for thrombin. The aptamer was immobilized onto nanoporous silica beads that were placed on the distal end of a fiber array. Fluorescein-labeled thrombin was used in a competitive assay format. Although fluorescence signals from the individual aptamer beads showed significant variability, the average signals of 100 beads provided much more precise values. The fiber-optic microarray system had a detection limit of 1 nM for nonlabeled thrombin, and each test could be performed in ∼15 min including the regeneration time. Rupcich et al.378 reported on the immobilization of an aptamer-based system in a sol-gel matrix. Fluorescein was covalently attached to the aptamer, and the quencher (dabcyl) was attached to the complementary oligonucleotide and, thus, was in close proximity of the fluorescein upon hybridization. Binding of the target molecule (ATP) to the aptamer results in a conformational change in the aptamer and formation of a stem-and-loop structure. The quencher-labeled oligonucleotide (QDNA) strand is displaced, and a large increase in fluorescence is observed. A tripartite aptamer complex also was prepared, where the fluorophore was attached to an oligonucleotide sequence complimentary to the stem part of the aptamer. The aptamer was attached to streptavidin to provide a lager molecular mass so to reduce leaching from the sol-gel matrix. Aminopropyltriethoxysilane was added to the sol-gel to promote electrostatic retention of the DNA because it introduces amino groups into the sol-gel, which are positively charged at pH’s below 8. The leaching rate was ∼12%/h. The authors also showed that a QDNA composed of 11 nucleotides is the best compromise between sensitivity and response time. The analytical range of the sensor was from 0.01 to 3 mM ATP. The sensor retained full signaling capability for 1 month but showed no response after a 3 month storage in buffer solution, presumably because of irreversible aging of the sol-gel matrix. To summarize this chapter, one can state that DNA sensors (a) have exceptional stability; (b) possess very low limits of detection; (c) are highly specific; (d) can be easily produced using automated procedures; (e) are self-contained in case of using molecular beacon; (f) are relatively tolerant to heat and frost; (g) are rather sensitive to effects of ionic strength; and (h) have very wide applicability and large potential in the case of aptamer DNA or DNAzyme sensors, thus allowing sensing even of haptens and proteins.

7. Whole-Cell Biosensors This type of biosensor makes use of living cells such as individual microorganisms or tissue, rather than relying on

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using specific biomolecules such as enzymes, proteins, or DNA. Whole-cell biosensors generally exhibit longer shelf lifetimes compared to, e.g., enzymatic biosensors because the active components are contained and produced in the “natural” environment and not in a polymer matrix on the surface of a sensor, which has limited biocompatibility. Whole-cell biosensors often are less costly than the corresponding enzymatic biosensors because some microorganisms can be cultivated and isolated rather easily, which is not the case for many enzymes. On the other side, they often lack specificity for the respective analytes. Whole-cell biosensors mostly are self-contained, do not require the addition of cofactors, and are the biorecognition elements of choice when the total amount of hazardous substances or pollutants is to be determined. Biosensors for determination of biological oxygen demand (BOD) represent a typical example. In contrast to such biosensors, those making use of genetically modified microorganisms can be highly specific. Another disadvantage of whole-cell biosensors is the relatively slow response (tens of minutes to hours) because the analytes need to diffuse through a cell membrane. Such response times are, however, adequate in certain cases. The standard BOD5 test, for example, requires 5 days.

7.1. Catalytic Whole-Cell Biosensors Biosensors for determination of biochemical oxygen demand (BOD) make use of oxygen transducers. As in the case of many enzymatic biosensors, the consumption of oxygen is monitored optically over time or at the endpoint. Trichosporon cutaneum379 and Bacillus subtilis380 bacteria were immobilized into poly(vinyl alcohol) (PVA) and solgel PVA-PVP [(poly(4-vinylpyridine)] networks. A combination of Bacillus licheniformis with Dietzia maris and Marinobacter marinus contained in a sol-gel-PVA matrix381 was used to obtain even lower selectivity and, thus, monitor much more possible pollutants. Compared to the sensor that makes use of B. Licheniformis only, a decrease in LOD from 0.9 to 0.2 mg/L and in response time from 30 to 3.2 min was observed for the multibacteria sensor. In all BOD sensors, the microorganism layer was placed on an oxygensensitive layer of various layouts and materials. These include a quenchable ruthenium(II) complex contained in plasticized PVC,379 in ormosil,381 or in silicone.380 Analytical ranges of the sensors were reported to be from 0 to 110,379 0 to 25,380 and 0.2 to 40 mg/L,381 respectively (expressed as equivalents of a glucose/glutamate BOD standard solution). The first sensor379 possessed moderate stability, and a 30% drop in sensor response was observed after 1 month of storage. The shelf life of the sensors reported later was much higher (a 12% decrease in activity in 1 month for the B. subtilis biosensor380 and only a 5% decrease in 6 months for the multibacteria sensor).381 A biosensor for the organophosphorous pesticide methyl parathion382 was prepared by analogy to the enzyme-type biosensors that made use of organophosphorous hydrolase (OPH). Whole cells of FlaVobacterium sp. containing OPH were immobilized on a glass fiber filter. Hydrolysis of methyl parathion (catalyzed by OHP) results in the formation of p-nitrophenol, which is readily detected by absorbance. The LOD (0.3 µM) and analytical range (4-80 µM) are comparable to the properties of OPH-based enzymatic biosensors. Recombinant E. coli cells immobilized in agarose were placed on a nylon membrane and used for determination of

Optical Biosensors

organophosphorous pesticides.383 These cells are capable of expressing OPH directly on the cell surface, which improves response times because of the low resistance to mass transport of the analytes and products across the cell membrane. The sensor detected as little as 3 µM of paraoxon and 5 µM of coumaphos and could be stored >1 month without a decrease in activity. Evidently, the main advantage of the whole-cell biosensors for pesticides over the enzymatic ones is that no isolation of the enzyme OPH is necessary. This reduces the costs significantly and improves the longterm stability. Arain et al.410 have studied the inhibitory effect of toxic metal ions on the cellular activity of E. coli and P. putida. Microtiter plates were prepared with integrated, fluorescencebased sensors for pH and oxygen, and bacterial respiratorial activity was monitored via the decrease in oxygen partial pressure of the closed system and also via the decrease in pH value. Other applications of such biosensors include respirometry and general toxicity assays.

7.2. External Stimuli-Based Cellular Biosensors Whole-cell biosensors were developed for the determination of various toxicants. For example, Bains384 immobilized Escherichia coli into an agarose membrane and monitored its UV absorbance at the unusual wavelength of 200 nm. In the presence of toxicants (e.g., sodium azide), the cells were stressed so that their metabolism was reduced and a drop in optical density was observed within 15 s. The sensitivity of the system was, however, poor with respect to the requirements of environmental monitoring. Frense et al.385 used algae cells (Scenesdesmus subspicatus) for the determination of environmentally harmful impurities in water. These chlorophyll-containing cells were immobilized on a filter paper that was covered with a thin alginate layer. Pollutants such as atrazine, endrine, and many other pesticides inhibit the electron-transport occurring during photosynthesis. This results in the increase of fluorescence of chlorophyll. The increase of fluorescence is well-related to the concentration of the pesticides, which can be measured at levels of several parts/billion (ppb). The sensor showed comparatively fast response (∼10 min) and good long-term stability in that storage at 4 °C within 6 months did not alter the sensor properties significantly. A similar approach was used by Naessens et al.386 who used algal cells from Chlorella Vulgaris (immobilized on a glass microfiber) for determination of atrazine, simazine, diuron, and other pesticides with high sensitivity (e.g., as little as 5 nM of atrazine when using the sensor in a flow mode). The biosensor showed good storage stability only for a limited period of time (7 days), and significant loss of activity was observed during longer storage. Different bacteria and mutants were found to respond to different pesticide classes387 because they can selectively modify the activity of photosystem II. The microorganisms were immobilized in a BSA-glutaraldehyde network deposited on a porous septum filter that was placed in a flow-through cell. Several flow-through cells were combined as an array to enable simultaneous sensing of several pesticides. The selectivity of the individual sensors, however, remained low enough, because each type of bacteria was sensitive to several classes of pesticides. The long-term stability of the biosensor was rather poor (half-life from 12 to 54 h).

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7.3. Genetically Engineered Whole-Cell Biosensors Progress in genetic engineering made possible the creation of a new type of microbial biosensor. It relies on the ability of a cell to respond to an environmentally harmful toxin by expressing specific genes. As a result, the toxin is transported out of the cell. To monitor the process, a reporter gene is fused to the induced gene. In the presence of the toxin, both genes are activated and expressed. A reporter gene usually is responsible for production of species that could be monitored optically, e.g., the green fluorescent protein. In a certain sense, all genetically engineered biosensors are external stimuli biosensors. A variety of biosensors makes use of the lux gene coding for the enzyme luciferase. In contrast to enzymatic bioluminescent sensors, the whole-cell biosensors often are selfcontained because luciferase and other reagents such as the cofactor, flavine mononucleotide, and substrate (a long-chain aldehyde) are produced in vivo. Typical examples of such biosensors are described below. Virta et al.388 developed a mercury biosensor that makes use of Escherichia coli containing the lux gene fused to the mer operon. The latter encodes for resistance to mercury, which is a nonessential and toxic metal for bacteria. The bioluminescence was triggered in the presence of Hg2+. Concentrations as low as 0.1 fM are detectable. A linear dependence is observed up to 0.1 µM of Hg2+. At higher concentrations, the luminescence rapidly drops to zero because of the toxicity of Hg2+ ions. The assay exhibits high selectivity, and no interference by other metals ions (except Cd2+) is observed. Sensitivity to Cd2+ is, however, ∼107fold lower than that to Hg2+ and, therefore, does not really compromise the performance of the assay. Another metal ion biosensor389 makes use of Escherichia coli containing the znt A gene fused to the reporter lac Z gene. While the first is responsible for transporting heavy metal ions out of the cell, the second produces the enzyme β-galactosidase, which cleaves the added substrate fluorescein di-β-D-galactopyranoside. Hence, fluorescence is increased in the presence of heavy metal ions. Individual cells of E. coli were immobilized on an array of 50 000 fibers (L of 2.5 µm) so that each bacterium occupied one individual fiber. Averaging the response from multiple identical sensors improved the signal-to-noise ratio. The LOD for Hg2+ was 100 nM. Unfortunately, no information is given about conceivable interferences by other heavy metal ions. Leth et al.390 have developed a biosensor for copper(II) ion based on a genetically engineered strain of Alcaligenes eutrophus into which was inserted a lux operon from Vibrio fischeri under the control of a copper-induced promoter. As a result, copper ions induce bioluminescence whose intensity is proportional to the concentration of the triggering ion. The cells were immobilized into calcium alginate and agarose gels, which were positioned in a home-made flow-cell, and luminescence intensity was monitored by means of a photodetector. The biosensors based on the two gels showed similar performance, which was highly influenced by the growth medium used. In fact, the analytical range of the sensor for both alginate and agarose was from 0 to 250 µM of Cu2+ when using the Luria-Broth (LODs of 50 and 25 µM, respectively). The use of a modified mineral reconstruction medium resulted in an LOD of 1 µM and an analytical range from 0 to 25 µM for alginate. The authors also showed

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that the performance of the biosensor is influenced by the concentration of riboflavin, suggesting the necessity of using a standard composition of nutrient medium. Unfortunately, the activity of the immobilized species was shown to decrease dramatically with time. In fact, a 2 week storage resulted in the complete loss of the activity of the bacteria for both alginate and agarose biogels. After the first 6 days of storage, the activity was, however, almost unchanged for alginate, while it dropped by 7-fold in agarose. Aromatic hydrocarbons are widespread and harmful pollutants that can be successfully detected by whole-cell biosensors. Here, the content of individual hydrocarbons is of less significance than the determination of overall toxicity. For example, Heitzer et al.391 used Pseudomonas fluorescens, which carried the nah G gene fused to a lux reporter gene to design a naphthalene biosensor. The bacteria were physically immobilized in alginate gel that was hardened at elevated temperature in the presence of SrCl2. Exposure of this sensor layer to both naphthalene and its degradation intermediate, salicylate, results in an increase of gene expression and, consequently, an enhancement of bioluminescence. The sensor demonstrated complete reversibility for at least 20 h, but the lowest detectable concentration of naphthalene was 0.12 mM, which is rather high. Ikariyama et al.392 used Pseudomonas putida bearing a xyl R gene (which recognizes benzene and its derivates) fused to the reporter lux gene from firefly. Because firefly luciferase catalyzes a different bioluminescent reaction than bacterial luciferase, addition of the substrate luciferin was necessary. The luminescent signal had a good linear relationship to the concentration of xylenes, which ranged from 0.05 to 1 mM. The response of the biosensor (where P. putida was immobilized onto a polycarbonate membrane) was shown to be much slower than the respective assay in solution and required hours of incubation time to achieve an adequate intensity of bioluminescence. In contrast to the sensors described above, biosensors for toxic chemicals are based on measurement of the reduction of intensity of the bioluminescent reaction when cells experience toxic or lethal conditions. Gil et al.393 immobilized a recombinant E. coli species bearing a lux reporter gene in a solid agar gel located in proximity of the distal end of an optical fiber. The biosensor was used for detection of toxic gases and vapors. As little as 48 ppm of benzene vapor could be detected. The sensor had a response time of ∼10 min and could be stored up to 20 days without degradation of activity. The sensitivity can be improved by increasing the surface that is exposed to vapors and by enhancing the diffusion of vapors, which can be accomplished by addition of glass beads.394 The sensitivity of other strains of bacteria bearing a lux reporter gene also was investigated.395 Shetty et al.396 developed a bioassay for determination of L-arabinose. The binding of the monosaccharide to the ara C regulatory protein was linked to the production of green fluorescent protein (GFP) by the reporter gene in an E. coli strain. The amount of GFP expressed is, thus, directly related to the concentration of L-arabinose. The dynamic range of the assay is from 0.5 µM to 5 mM. The sensitivity to other monosaccharides was ∼10 times lower than to arabinose. A biosensor also was developed where the bacteria were entrapped behind a membrane at the tip of a bifurcated fiber bundle. Although operated similarly to the bioassay, an ∼10fold increase in the LOD was reported.

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Recently, several bioassays were developed for detection of endocrine-disrupting chemicals (EDCs), compounds that affect human health by irregularly modulating endocrine functions. Michelini et al.397 used recombinant Saccharomyces cereVisiae cells that were modified to express the human androgen receptor together with Photinus pyralis luciferase. The assay responds to testosterone in the concentration range from 50 pM to 1 µM. Luciferin needs to be added to the assay solution. Compared to other available methods, the assay is rather fast (150 min of incubation needed for optimal performance against 24 h for other methods). A biosensor also was developed for estrogenic EDCs.398 Genetically modified Saccharomyces cereVisiae cells containing the estrogen receptor were entrapped in hydrogel matrices based on calcium alginate or poly(vinyl alcohol). The LODs for the two EDCs investigated were 0.08 and 0.64 µg/L. The immobilized cells retained their activity for 1 month, however only if stored at -80 °C. A lactate biosensor was reported399 that utilizes a bacterial cytoplasmic membrane isolated from E. coli, which was genetically modified to express its lactate-oxidizing activity. The sensor properties can be tuned by varying the conditions of cultivation. The cytoplasmic membranes were adsorbed on a cellulose disk placed above an oxygen transducersa ruthenium(II)-derived luminescent probe in a silicone matrix. The sensing scheme is based on the consumption of oxygen during oxidation of lactate and is the same as that used in the enzyme based on the use of lactate oxidase. The analytical range of the biosensor is from 0.05 to 5 mM of lactate. In some optical biosensors, even whole tissues have been used as recognition elements. For example, Lundstro¨m et al.400 used fish scales from Labrus ossifagus containing melanophores. Their cells contain pigment granules that are either dispersed or aggregated. Aggregation of the granules is promoted by noradrenaline, which can be monitored optically via the increase in the transmittance of the scale. Addition of the noradrenaline antagonist yohimbine recovers the initial signal. Nanomol quantities of noradrenaline could be measured, and the response time of the sensor was ∼10 min. In summary, it can be stated that cellular biosensors (a) are relatively easy to manufacture; (b) are rather unspecific in the case of catalytic biosensors but fairly specific in the case of external-stimuli biosensors and gene-modified biosensors; (c) possess rather slow response; (d) are selfcontained; (e) are more stable, in general, than enzymatic biosensors but are sensitive to heat and, less so, to frost; and (f) vary over a wide range in terms of sensitivity.

8. Solid Supports for Use in Optical Biosensors, and Other Methods of Immobilization The success of (bio)sensor research and development dependssmore often than anticipatedson the availability of adequate materials. One may differentiate between materials for mechanical sensor supports and materials for use as matrices or membranes that contain the biologically active species, or indicators in the case of catalytic biosensors. These shall be discussed here briefly in addition to the specific examples of immobilization given in section 4.4. There are three kinds of “supports” for optical biosensors. The first one is of the completely inert type. Its only purpose is to serve as a mechanical support to facilitate the handling of planar sensors. The second is of the optical waveguide type and, thus, can serve as an essential compo-

Optical Biosensors

nent in the process of optical interrogation of the sensor material. The third (and most recent) group of supports are the “active” supports such as fluorescent nanoparticles (Qdots), metal films, beads of noble metals, or inorganic or organic micro- and nanoparticles. These can act as microlight sources or quenchers, for example, and thus can actively take part in the spectroscopic scheme. Any of these supports can be preactivated (i.e., made bioreactive) form to enable covalent immobilization of the biocomponent or of other species. Inert supports come in various forms and include films of poly(ethylene terephthalate), which is readily available at low costs and also easy to handle. Other supports include poly(methyl methacrylates) and polycarbonates, with their excellent optical transparency, and polystyrene, which is widely used in microtiterplates (MTPs). In most cases, the chemically responsive material (the sensor “cocktail”) is deposited, or printed, or stamped on such a support, in a groove of this material, or in the wells of a (plastic) MTP. The material, after having been deposited as a thin film on the support, is punched into sensor spots, and these are being placed in disposable sensor devices. The sensor layer is then interrogated by guiding the light beam onto the sensor layer, and reflectivity or fluorescence is measured or interference is measured. It is obvious that the mechanical supports are expected to create no background signal. The support also may act as a waveguide material. Planar waveguides, fiber-optics, and, less often, capillaries have been applied. There are two ways to guide the exciting light to the sensor material. The first is by direct illumination and by collecting luminescence via the waveguide. The second is to use the waveguide for both the exciting beam and for collection of emitted luminescence. Both geometries have their merits (see section 4.3). Waveguide-based sensors are most elegant and, therefore, have found widespread application. Among the third kind of supports, the nanobeads exploit the fact that, because of their intrinsic luminescence, they can act as a donor in FRET assays.55 Metal particles and films, in turn, can act as quenchers or enhancers of luminescence (depending inter alia on the spatial distance between metal and fluorophore and on the kind of metal). A most interesting class of micro- and nanoparticles is represented by the so-called upconverting phosphors (UCPs). They are capable of converting near-infrared light (from lowcost diode lasers) into visible light with fair to high efficiency. Upconversion is not related to 2-photon excitation, which occurs at strong light intensities only. UCPs (mostly oxides, sulfides, andspreferablysfluorides of trivalent lanthanide ions) enable complete elimination of autofluorescence, which commonly impairs the performance of fluorescence-based assays. UCPs are ideal donors for fluorescence resonance energy transfer (FRET)-based assays. UCP-based FRETs have been applied in immunoassays401,402 and in nucleic acid hybridization assays.403,404 Arguably, these methods are at the borderline between solid-phase-based biosensors of the conventional type and of classical solution assays. With respect to materials for use as a bulk matrix for sensors, it is important to remind that the design of such sensors depends on the size of the analyte. Enzyme-based sensors usually digest or metabolize substrates of low or medium molecular weight. Hydrogel matrices, for example, are useful in the case of low-molecular-weight analytes only,

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since these can penetrate the bulk matrix, where, in the case of enzyme sensors or whole-cell sensors, metabolism can occur. Immunosensors almost never are designed as bulk sensors even though they could be, at least for small analytes such as atrazine. Immunosensors for large analytes require the antibody to be immobilized on the surface of a support, usually as a nanometer thick coating or film. The situation is similar for DNA sensors where the size and diffusion of the analyte is critical. Whole-cell biosensors (in the majority of cases) have been used for low-molecular-weight analytes and, therefore, have been incorporated into analyte-permeable gels such as from alginates. The surface of a biosensor layer is either covered with a polymer/enzyme matrix (like in most enzymatic biosensors) or directly with biorecognition elements such as antibodies or oligonucleotides. In this case, the surface needs to be made reactive first in order to allow immobilization of a biomolecule. Covalent attachment of a biomolecule to a support is much more commonly used than physical absorption. Numerous cross-linkers and spacers can be used.20,292,296 Proteins such as bovine serum albumin often are also deposited on the surface, so to saturate remaining binding sites. Immobilization of antibodies and oligonucleotides via the (noncovalent) biotin-avidin46,67,68,246,288,299 or biotin-streptavidin47,354,355,371,374 couple is widespread. A support modified with (strept)avidin can be used, in principle, for immobilization of any biotinylated molecule. Alternatively, biotinylated recognition elements can be immobilized via streptavidin onto a surface modified with biotin. Enzymes are often covalently immobilized onto preactivated transparent polyamide or poly(vinylidenedifluoride) membrane supports such as Immunodyne,73,90,140,163,227,228,230,232-234 Biodyne,177,191,225 and Immobilon.164,194,195 Less common supports include eggshell membranes86,154 and swim-bladder membranes88 and were reported to be highly biocompatible. Hydrophilic polymer matrixes are widely used for immobilization of biocomponents (such as enzymes and bioreceptors) and of optical indicators. Sol-gels (whose polarity can vary over a wide range by introducing organic groups to end up with ormosils)401,402 have been often used for immobilization of enzymes,44,84,87,120,131,134,135,380,206,216 bioreceptors,343 and even whole cells.380,381 One major reason for the popularity of sol-gels results from the fact that the activity of biocomponents is retained over a long time. Hydrogels have also become popular36,80,102,118,147,150,212,398 because they do not require a modification of the biological component. Enzymes sometimes are cross-linked with glutaraldehyde and BSA to form a polymer network located on a support or directly in a hydrophilic polymer (e.g., PVA) layer.75,175,176,188,215,217,387 Koncki et al.97 have designed optical biosensors based on the use of thin films of Prussian Blue incorporated into polypyrrole. Other semiconducting organic materials may be used as well. The composite film is deposited on a nonconducting support and used as an optical transducer for flow-through biosensors based on hydrolases and oxidases. Immobilization of glucose oxidase resulted in a glucose biosensor where the film responds to both pH and hydrogen peroxide by a change in its color. Millimolar concentrations can be determined. The biosensor is said to be quite stable owing to the presence of a poly(pyrrolylbenzoic acid) network in the composite material. This organic polymer plays a dual role as a binding agent for inorganic material

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and as a functionalized support for strong covalent immobilization of enzyme molecules. Polymer matrixes also can act as a support for the immobilization of indicators used in cellular or enzymatic sensors. Ruthenium(II) polypyridyl complexes are common when oxygen transducers are used (see Table 2). To achieve the desired oxygen sensitivity, they can be adsorbed onto silica beads that are dispersed in a layer of a highly oxygenpermeable silicone layer. Absorption-based and fluorescent pH indicators were employed in enzymatic biosensors using pH and NH3 transducers. Permeability to protons is mandatory in these cases, and the indicators, therefore, often are contained in hydrogels. To prevent leaching of indicators to the sample media, they are sometimes absorbed onto the surface of microbeads.103,194 Dialysis (ultrafiltration) membranes are used in certain sensor types.41,42,48,49,221,222,224 They allow small analytes to diffuse freely in and out of a chamber and interact there with a biorecognition element. The latter is either large enough itself or is conjugated with a large molecule in order to not leach out of the chamber. In contrast to indicators, labels do not respond to substrates or reaction products but render a biomolecule detectable. Fluorescent labels are the most common ones. Fluoresceins, rhodamines, cyanine dyes, and numerous others are commercially available403 and widely used (see Table 3). Ideally, a label should absorb in visible light to reduce background fluorescence, be bright, and be inert. Brightness (defined as the product of molar absorbance and quantum yield) is particularly significant and should exceed 30 000 M-1‚cm-1). Since fluoresceins are viable pH indicators, thorough control of pH is essential for these labels. Luminescent colloidal semiconductor nanocrystals (quantum dots, QDs) also represent viable labels,44,55,56 despite their cell toxicity and difficult surface chemistry, and can largely expand the range of useful fluorophores for biosensors. The group of Seeger have found404 that biotin-functionalized cellulose monolayers can act as a new kind of support and have used it for the fluorescent detection of single molecules via laser-induced confocal single-molecule spectroscopy in glass-bottom microplates. Gold nanobeads can be used to increase the brightness of fluorescent biosensors.37,272,345,405

9. Outlook Optical biosensing has experienced a substantial growth despite the usual critical comments of certain “experts” that expect new technologies to make a breakthrough (mainly in commercial terms) within a few years and despite the overoptimistic presentations of certain researchers, which often does more harm to a new technology than supporting it. Optical biosensor technology is not a matter of spectroscopy only, or of material sciences, or any other single discipline, but rather requires various kinds of scientists to cooperate in order to end up with a viable biosensor scheme and, ideally, commercial products. Optical biosensors have numerous applications, and not each scheme will be applicable to any given analyte. Moreover, methods that may work for a specific analyte in a certain matrix may not even work for the same analyte in another matrix. This fact is but one of the reasons why biosensors, in a commercial sense, are not as successful as was expected initially. The trend toward multianalyte sensing and toward biosensor arrays is obvious, even though certain single sensor

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spots in an array may not be needed in any conceivable application. DNA arrays are technically the most advanced, not the least because DNA is built from 4 nucleotides only, which makes synthetic and surface chemistry comparably more simply than in the case of protein arrays. Enzyme arrays are rather established and have found application in clinical analyzers for glucose, urea, cholesterol, and lactate. Protein arrays are most versatile but also can be most complex, and this has hampered their technical realization. Proteins not only come in a variety of sequences (of 20 different amino acids!) but also in a variety of tertiary and quaternary structures, which makes labeling and immobilization an experience in each single case. Unfortunately, the current terror hype has directed research in protein arrays away from health and environment into other areas, which implies a massive redirection of tax money and appears not to really represent a useful long-term investment.

10. List of Abbreviations and Acronyms Abs. AChE ADH AOx AMPT ASF BL BSA BTB BTP b. luciferase CCA ChOx CLum. CF CFU CDNB Cy-5 2,4-D DDAO DFP f. luciferase EDC ELum. EuTc FITC FRET GAH GDH GFOR GLOx GlDH GOx GPT GSH GST hCG HPOx HPTS IgG I. LDH LOx Lum. LyOx NHS OPH

absorption acetylcholine esterase alcohol dehydrogenase alcohol oxidase 2-(2-acetoxy-5-methylphenylazo)-N-methyl-1,3thiazolium methosulfate African swine fever bioluminescence bovine serum albumine bromothymol blue benzo[a]pyrene tetraol bacterial luciferase chlorendic capronic acid choline oxidase chemiluminescence 5(and 6)-carboxyfluorescein colony-forming units 1-chloro-2,4-dinitrobenzene carboxymethylindocyanine succunimidyl ester 2,4-dinitrophenoxyacetic acid 7-hydroxy-9H-1,3-dichloro-9,9-dimethylacridin-2-one diisopropyl phosphorofluoridate firefly luciferase ethyl-3-[1-dimethylaminopropyl]carbodiimide electroluminescence europium(III) tetracycline fluorescein isothiocyanate fluorescence resonance energy transfer glutaminase glucose dehydrogenase glucose-fructose oxidoreductase glutamate oxidase glutamate dehydrogenase glucose oxidase glutamic-pyruvic transmitase glutathione glutathione-S-transferase chorionic gonadotrophin horseradish peroxidase 8-hydroxypyrene-1,3,6-trisulfonate immunoglobulin intensity lactate dehydrogenase lactate oxidase luminescence lysine oxidase N-hydroxysuccinimide organophosphate hydrolase

Optical Biosensors OR ORP PEG PMMA POx PSA PtOEP RDX Refl. Ru-bipy Ru-phen Ru-dpp

oxidoreductase organophosphorous pesticides poly(ethylene glycol) polymethylmethacrylate peroxidase prostate-specific antigen platinum(II) octaethylporphyrin hexaxydro-1,3,5-trinitro-1,3,5-triazine reflectance ruthenium(II) trisbipyridyl ruthenium(II) tris(1,10-phenanthroline) ruthenium(II) tris(4,7-diphenyl-1,10-phenanthroline) SEB staphylococcal Enterotoxin B SNARF seminaphthofluorescein TCPB 2,4,6-trichlorophenoxybutyrate Ti(IV) reagent titanium(IV) + 24(5-bromopyridyi)azo)5-(Npropyl-N-sulfopropylamino)phenol TMR tetramethylrhodamine TNB trinitrobenzene TNT trinitrotoluene TRITC tetramethylrhodamine-5-isothiocyanate triazine derivative 4-chloro-6-(isopropylamine)-1,3,5-triazine-2-(6aminohexane carboxylic acid) TSH thyroid stimulating hormone XOx xanthine oxidase

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CR068105T

Chem. Rev. 2008, 108, 462−493

462

Surface Plasmon Resonance Sensors for Detection of Chemical and Biological Species Jirˇ´ı Homola† Institute of Photonics and Electronics ASCR, Chaberska´ 57, 182 51 Prague 8, Czech Republic Received May 31, 2007

Contents 1. Introduction 2. Surface Plasmons on Planar Structures 2.1. Surface Plasmons on Metal−Dielectric Interface 2.2. Long-Range and Short-Range Surface Plasmons 2.3. Optical Excitation of Surface Plasmons 3. Fundamentals of SPR Sensors 3.1. SPR Sensors 3.2. SPR Affinity Biosensors 3.3. Performance Considerations 3.3.1. Sensitivity 3.3.2. Resolution 3.3.3. Limit of Detection and Minimum Resolvable Surface Coverage 4. Advances in SPR Biosensor Technology 4.1. Optical Platforms Used in SPR Sensors 4.1.1. SPR Sensors Based on Prism Couplers 4.1.2. SPR Sensors Based on Grating Couplers 4.1.3. SPR Sensor Based on Waveguide Couplers 4.2. Biorecognition Elements and Their Immobilization 4.2.1. Biorecognition Elements 4.2.2. Immobilization of Biorecognition Elements 4.2.3. Nonfouling Surfaces 4.3. Summary 5. Applications of SPR Sensors for Detection of Chemical and Biological Species 5.1. Detection Formats 5.2. Food Quality and Safety Analysis 5.2.1. Pathogens 5.2.2. Toxins 5.2.3. Veterinary Drugs 5.2.4. Vitamins 5.2.5. Hormones 5.2.6. Diagnostic Antibodies 5.2.7. Allergens 5.2.8. Proteins 5.2.9. Chemical Contaminants 5.3. Medical Diagnostics 5.3.1. Cancer Markers 5.3.2. Antibodies against Viral Pathogens

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† Telephone +420 266 773 448; fax +420 284 681 534; e-mail [email protected].

5.3.3. Drugs and Drug-Induced Antibodies 5.3.4. Hormones 5.3.5. Allergy Markers 5.3.6. Heart Attack Markers 5.3.7. Other Molecular Biomarkers 5.4. Environmental Monitoring 5.4.1. Pesticides 5.4.2. 2,4,6-Trinitrotoluene (TNT) 5.4.3. Aromatic Hydrocarbons 5.4.4. Heavy Metals 5.4.5. Phenols 5.4.6. Polychlorinated Biphenyls 5.4.7. Dioxins 5.5. Summary 6. Conclusions 7. Abbreviations 8. Acknowledgment 9. References

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1. Introduction Optical sensors based on excitation of surface plasmons, commonly referred to as surface plasmon resonance (SPR) sensors, belong to the group of refractometric sensing devices including the resonant mirror sensor,1,2 the grating coupler sensor,3-5 the integrated optical Mach-Zehnder interferometer,6,7 the integrated Young interferometer,8,9 and the white light interferometer,10,11 which measure changes in the refractive index occurring in the field of an electromagnetic wave supported by the optical structure of the sensor. Since the first demonstration of surface plasmon resonance for the study of processes at the surfaces of metals12 and sensing of gases13 in the early 1980s, SPR sensors have made vast advances in terms of both development of the technology and its applications. SPR biosensors have become a central tool for characterizing and quantifying biomolecular interactions. Moreover, development of SPR sensors for detection of chemical and biological species has gained considerable momentum, and the number of publications reporting applications of SPR biosensors for detection of analytes related to medical diagnostics, environmental monitoring, and food safety and security has been rapidly growing. Over the past decade, thousands of research papers on SPR biosensors have been published. SPR biosensors, as one of the main optical biosensor technologies, have been also extensively featured in biosensor books14-17 and reviews.18-26 Review articles27-30 and books31 focused on SPR biosensor technology have also been published.

10.1021/cr068107d CCC: $71.00 © 2008 American Chemical Society Published on Web 01/30/2008

SPR Sensors

Jirˇ´ı Homola received his M.S. degree in physical engineering in 1988 from the Czech Technical University, Prague, Czech Republic, and his Ph.D. degree from the Academy of Sciences of the Czech Republic in 1993. Since then, he has been with the Institute of Photonics and Electronics, Prague, Czech Republic. From 1997 to 2002, he was with the Department of Electrical Engineering, University of Washington, Seattle, since 2001 as a Research Associate Professor. He is currently Head of Photonics Division and Chairman of Department of Optical Sensors at the Institute of Photonics and Electronics, Prague, and Affiliate Associate Professor at the University of Washington, Seattle. Dr. Homola is a Member of the Editorial Board of Sensors and Actuators B (Elsevier) and Associate Editor of the Journal of Sensors (Hindawi). He has served on technical committees of numerous scientific conferences and is a Member of the Permanent Steering Committees of Advanced Study Course on Optical Chemical Sensors and Europt(r)ode Conference Series. He is also a Member of the NATO CBP Advisory Panel. He received the Otto Wichterle Prize, Czech Republic, in 2003, and the Roche Diagnostics Prize for Sensor Technology, Germany, in 2006. Dr. Homola has edited 2 books, 3 book chapters, and 60 research papers in scientific journals and presented over 100 research papers at international conferences. His research interests are in biophotonics, guided-wave photonics, optical sensors, and biosensors.

This paper reviews advances in SPR sensor technology and its applications since the year 2000. The review focuses on SPR sensors employing conventional (unlocalized) surface plasmons propagating along planar structures and their applications for detection of chemical and biological species. Aspects related to applications of SPR method for the study of biomolecules and their interactions are outside the scope of this review.

2. Surface Plasmons on Planar Structures The first observation of surface plasmons was made in 1902 by Wood, who reported anomalies in the spectrum of light diffracted on a metallic diffraction grating.32 Fano has proven that these anomalies are associated with the excitation of electromagnetic surface waves on the surface of the diffraction grating.33 In 1968 Otto demonstrated that the drop in the reflectivity in the attenuated total reflection (ATR) method is due to the excitation of surface plasmons.34 In the same year, Kretschmann and Raether observed excitation of surface plasmons in another configuration of the attenuated total reflection method.35 These pioneering works of Otto, Kretschmann, and Raether established a convenient method for the excitation of surface plasmons and their investigation and ushered surface plasmons into modern optics. Fundamentals of surface plasmons can be found in numerous review articles and books.36-39

2.1. Surface Plasmons on Metal−Dielectric Interface The simplest geometry in which a surface plasmon can exist consists of a semi-infinite metal with a complex

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Figure 1. Spatial distribution of the magnetic intensity for a surface plasmon at the interface between gold and a dielectric (nd ) 1.328) in the direction perpendicular to the interface, λ ) 850 nm.

permittivity, m ) ′m + i′′m, and a semi-infinite dielectric with permittivity, d ) ′d + i′′d, where ′j and ′′j are real and imaginary parts of j (j is m or d). Analysis of Maxwell’s equations with appropriate boundary conditions suggests that this structure can support only a single guided mode of electromagnetic fieldsa surface plasmon. Surface plasmon is a transversally magnetic (TM) mode, and therefore its vector of intensity of magnetic field lies in the plane of metal-dielectric interface and is perpendicular to the direction of propagation. If we use such a Cartesian system of coordinates that the metal occupies the region z < 0 and the surface plasmon propagates along the x-axis, the vector of magnetic intensity H B of the surface plasmon can be written as

H B j ) (0,Hy,0)j ) (0,1,0)A exp[-Rjz + i(βx - ωt)]

(1)

where ω is the angular frequency, t is time, β is the propagation constant, Rj ) xβ2-(ω/c)2j, where j is either m for metal or d for dielectric, and i ) x-1. The vector of the electric field is perpendicular to the vector of the magnetic intensity and can be calculated from Maxwell’s equations and eq 1. A typical profile of the magnetic field of a surface plasmon is shown in Figure 1. The intensity of the magnetic field reaches its maximum at the metal-dielectric interface and decays into both the metal and dielectric. The field decay in the direction perpendicular to the metal-dielectric interface is characterized by the penetration depth, which is defined as the distance from the interface at which the amplitude of the field decreases by a factor of e (where e is the base of the natural logarithm). The penetration depth depends on the wavelength and permittivities of the materials involved. Penetration depth into the dielectric for a surface plasmon propagating along the interface of gold and a dielectric with nd ) 1.32 increases with wavelength and ranges from 100 to 600 nm in the wavelength region from 600 to 1000 nm.31 The propagation constant of a surface plasmon at a metaldielectric interface can be expressed as

βSP )

x

ω c

x

dm 2π )  d + m λ

dm  d + m

(2)

where c is the speed of light in a vacuum and λ is the wavelength in a vacuum.36,37 If the structure is lossless (′′m ) ′′d ) 0), eq 2 represents a guided mode only if the

464 Chemical Reviews, 2008, Vol. 108, No. 2

Homola

Figure 2. Spatial distribution of the magnetic intensity of symmetric and antisymmetric surface plasmons propagating along a thin gold film (m ) -30.5 + i 1.6, thickness ) 20 nm) embedded between two identical dielectrics (nd ) 1.328), λ ) 850 nm.

permittivities ′m and ′d are of opposite signs and ′m < -′d. As the permittivity of dielectric materials is usually positive, this requires that the real part of the permittivity of the metal is negative. Metals such as gold, silver, and aluminum exhibit a negative real part of permittivity in the visible and nearinfrared regions of the spectrum. These metals also exhibit a considerable imaginary part of the permittivity, which causes the propagation constant of a surface plasmon to have a nonzero imaginary part. The imaginary part of the propagation constant is associated with the attenuation of the surface plasmon in the direction of propagation.36,37 The propagation constant is related to the effective index nef and attenuation b as

c 0.2 Im{βSP} nef ) Re{βSP}, b ) ω ln 10

(3)

where Re{ } and Im{ } denote the real and imaginary parts of a complex number, respectively; the attenuation b is in dBcm-1 if β is given in m-1.

2.2. Long-Range and Short-Range Surface Plasmons A planar structure consisting of a thick metal film sandwiched between two semi-infinite dielectrics supports two independent surface plasmons at the opposite boundaries of the metal film. If the metal film is thin, coupling between the surface plasmons at opposite boundaries of the metal film can occur, giving rise to mixed modes of electromagnetic fieldssymmetric and antisymmetric surface plasmons. Characteristics of these surface plasmons can be found from Maxwell’s equations and appropriate boundary conditions.40-42 The symmetric surface plasmon exhibits a propagation constant and attenuation, which both increase with increasing metal film thickness.31 The propagation constant and attenuation of the antisymmetric surface plasmon decrease with increasing thickness of the metal film. The symmetric surface plasmon exhibits a lower attenuation than its antisymmetric counterpart, and therefore it is referred to as a long-range surface plasmon, whereas the antisymmetric mode is referred to as a short-range surface plasmon.40,41 Figure 2 shows the distribution of magnetic intensity of the symmetric and antisymmetric surface plasmons propagating along a thin gold film surrounded by two identical dielectrics. The profiles of magnetic intensity of symmetric and antisymmetric plasmons are symmetric or antisymmetric

Figure 3. Coupling of light to a surface plasmon via (A) prism coupler, (B) waveguide coupler, and (C) grating coupler.

with respect to the center of the metal. The field of the symmetric surface plasmon penetrates much more deeply into the dielectric medium than the field of the antisymmetric surface plasmon or the field of a conventional surface plasmon at a single metal-dielectric interface (Figure 1).

2.3. Optical Excitation of Surface Plasmons A light wave can couple to a surface plasmon at a metaldielectric interface if the component of light’s wavevector that is parallel to the interface matches the propagation constant of the surface plasmon. As the propagation constant of a surface plasmon at a metal-dielectric interface is larger than the wavenumber of the light wave in the dielectric, surface plasmons cannot be excited directly by light incident onto a smooth metal surface. The wavevector of light can be increased to match that of the surface plasmon by the attenuated total reflection or diffraction. This enhancement and subsequently the coupling between light and a surface plasmon are performed in a coupling device (coupler). The most common couplers used in SPR sensors include a prism coupler, a waveguide coupler, and a grating coupler (Figure 3). Prism couplers represent the most frequently used method for optical excitation of surface plasmons.35-37 In the Kretschmann configuration of the attenuated total reflection (ATR) method (Figure 3A),35 a light wave passes through a high refractive index prism and is totally reflected at the base of the prism, generating an evanescent wave penetrating a

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thin metal film. The evanescent wave propagates along the interface with the propagation constant, which can be adjusted to match that of the surface plasmon by controlling the angle of incidence. Thus, the matching condition

2π n sin (θ) ) Re{βSP} λ p

(4)

can be fulfilled, allowing the evanescent wave to be coupled to the surface plasmon. θ denotes the angle of incidence, np denotes the refractive index of the prism (np > nd), and βSP denotes the propagation constant of the surface plasmon. Surface plasmons can be also excited by a light wave guided in an optical waveguide. This approach is illustrated in Figure 3B. Light propagates in a waveguide in the form of guided modes. The electromagnetic field of a guided mode is concentrated in the waveguiding layer, and a portion of the field propagates, as an evanescent wave, in the low-refractive index medium surrounding the waveguiding layer. When light enters the region of the waveguide containing a metal layer, the evanescent wave excites a surface plasmon at the outer boundary of the metal layer. The coupling condition for the guided mode and the surface plasmon is fulfilled when the propagation constants of the two waves are equal

βmode ) Re{βSP}

(5)

3.2. SPR Affinity Biosensors

where βmode denotes the propagation constant of the waveguide mode. Another approach to optical excitation of surface plasmons is based on the diffraction of light in a grating coupler (Figure 3C). In this method, a light wave is incident from the dielectric medium on a metallic grating. The diffracted light can couple to a surface plasmon if the momentum of diffracted light parallel to the grating surface is equal to the propagation constant of the surface plasmon

2π 2π n sin θ + m ) (Re{βSP} λ d Λ

are classified as sensors with angular, wavelength, intensity, or phase modulation.31 In SPR sensors with angular modulation, a monochromatic light wave is used to excite a surface plasmon. The strength of coupling between the incident wave and the surface plasmon is observed at multiple angles of incidence, typically by employing a convergent light beam. The excitation of surface plasmons is observed as a dip in the angular spectrum of reflected light. The angle of incidence yielding the strongest coupling is measured and used as a sensor output.44 In SPR sensors with wavelength modulation, a surface plasmon is excited by a collimated light wave containing multiple wavelengths, typically a beam of polychromatic light. The excitation of surface plasmons is observed as a dip in the wavelength spectrum of reflected light. The wavelength yielding the strongest coupling is measured and used as a sensor output.45 SPR sensors with intensity modulation are based on measuring the strength of the coupling between the light wave and the surface plasmon at a single angle of incidence and wavelength, and the intensity of light wave serves as a sensor output.13 In SPR sensors with phase modulation the shift in phase of the light wave coupled to the surface plasmon is measured at a single angle of incidence and wavelength of the light wave and used as a sensor output.46

(6)

where m is an integer and denotes the diffraction order and λ is the grating period.43 In the process of optical excitation of surface plasmon, a portion of the energy of the light wave is transferred into the energy of a surface plasmon and dissipated in the metal film, which results in a drop of intensity of the light wave. In addition to the change in the intensity, the light wave exciting a surface plasmon undergoes a change in phase.31

3. Fundamentals of SPR Sensors 3.1. SPR Sensors In principle, SPR sensors are thin-film refractometers that measure changes in the refractive index occurring at the surface of a metal film supporting a surface plasmon. A surface plasmon excited by a light wave propagates along the metal film, and its evanescent field probes the medium (sample) in contact with the metal film. A change in the refractive index of the dielectric gives rise to a change in the propagation constant of the surface plasmon, which through the coupling condition (eqs 4-6) alters the characteristics of the light wave coupled to the surface plasmon (e.g., coupling angle, coupling wavelength, intensity, phase). On the basis of which characteristic of the light wave modulated by a surface plasmon is measured, SPR sensors

SPR affinity biosensors are sensing devices which consist of a biorecognition element that recognizes and is able to interact with a selected analyte and an SPR transducer, which translates the binding event into an output signal. The biorecognition elements are immobilized in the proximity of the surface of a metal film supporting a surface plasmon. Analyte molecules in a liquid sample in contact with the SPR sensor bind to the biorecognition elements, producing an increase in the refractive index at the sensor surface, which is optically measured (section 3.1). The change in the refractive index produced by the capture of biomolecules depends on the concentration of analyte molecules at the sensor surface and the properties of the molecules. If the binding occurs within a thin layer at the sensor surface of thickness h, the sensor response is proportional to the binding-induced refractive index change, which can be expressed as

∆n )

(dndc)Γh

(7)

where (dn/dc) denotes the refractive index increment of the analyte molecules (typically 0.1-0.3 mL/g47,48) and Γ denotes the surface concentration in mass/area.49

3.3. Performance Considerations The main performance characteristics of SPR (bio)sensors include sensitivity, linearity, resolution, accuracy, reproducibility, dynamic range, and limit of detection.31

3.3.1. Sensitivity Sensor sensitivity is the ratio of the change in sensor output to the change in the quantity to be measured (e.g., concentration of analyte). The sensitivity of an SPR affinity biosensor depends on two factorsssensitivity of the sensor output (e.g., resonant angle or wavelength) to the refractive index and efficiency of the conversion of the binding to a change in

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Homola

the refractive index (section 3.2).31 The sensitivity of an SPR sensor to a refractive index SRI can be expressed as a product of two terms

SRI )

δY δnef δnef δnd

(8)

where Y denotes sensor output. The first term describes the sensitivity of sensor output to the effective index of a surface plasmon and depends on the method of excitation of surface plasmons and the used modulation approach. The second term describes the sensitivity of the effective index of a surface plasmon to refractive index and is independent of the modulation method and the method of excitation.31 In general, the sensitivity of the effective index of a surface plasmon to refractive index depends on the distribution of the refractive index change. Homola and Piliarik used the perturbation theory50 to calculate the sensitivity of the effective index to the refractive index for two limiting cases: (i) the change in the refractive index that occurs within the whole sample and (ii) the change in the refractive index that occurs only within a very short distance from the sensor surface.31 They showed that the sensitivity of the effective index of a surface plasmon to bulk refractive index change can be expressed as

( ) δnef δnd

B

)

nef3 nd3

>1

(9)

and therefore is always larger than the sensitivity of a free space plane wave in the infinite dielectric medium, which is equal to 1.31 The sensitivity of the effective index of a surface plasmon to surface refractive index change occurring within a layer with a thickness h can be, assuming h , Lpd ) 1/Re{Rd} and |′m| . ′′m, written as

( ) ( ) δnef δnd

δnef h )2 δnd BLpd S

SPR sensors of all the modulation approaches need to measure the intensity of the light wave coupled to a surface plasmon to determine the sensor output. Therefore, their resolution is limited by the noise in the intensity of the detected light. Dominant sources of noise are the fluctuations in the light intensity emitted by the light source, shot noise associated with photon statistics, and noise in conversion of the light intensity into electric signal by the detector.31 To reduce the noise, light intensities are averaged. The averaging involves either averaging of time series of intensity from the same detector (time averaging) or averaging of intensities from multiple detectors (e.g., of a two-dimensional array) measured at a single time (spatial averaging). The time averaging reduces the noise in the intensity of light by a factor of xM, where M is the number of averaged intensities. The spatial averaging used in spectroscopic SPR sensors (averaged spectra are measured in several rows of a 2D detector51,52) or intensity-modulated sensors (averaged area of a 2D detector forms a signal of one measuring channel53-55) is less efficient, as the light fluctuations affect all of the measured intensities in the same way and therefore cannot be eliminated by the spatial averaging. The noise in the light intensity is translated to sensor output noise by a data processing algorithm used to generate the sensor output. Although various methods for processing data from spectroscopic SPR sensors have been developed (centroid method,56,57 polynomial fitting,58,59 and optimal linear data analysis60), the noise in angular or wavelength spectra was found to transform to the noise in the sensor output in a similar fashion.61 Piliarik and Homola investigated the propagation of noise through the centroid data processing algorithm and demonstrated that the noise of the centroid method can be expressed as

1 σth w σRI ) K xN d SRI

(11)

(10)

This suggests that the surface refractive index sensitivity is proportional to the bulk refractive index sensitivity and the ratio of the thickness of the layer within which the surface refractive index change occurs and the penetration depth of the surface plasmon, Lpd. As the penetration depth of a surface plasmon on gold increases with increasing wavelength, the surface refractive index sensitivity of the effective index decreases with the wavelength more quickly than the bulk refractive index sensitivity.31

3.3.2. Resolution Resolution is a key performance characteristic of an SPR sensor and ultimately limits another important performance characteristic of an SPR affinity biosensorsthe limit of detection (LOD). The resolution of an SPR sensor is defined as the smallest change in the bulk refractive index that produces a detectable change in the sensor output. The magnitude of sensor output change that can be detected depends on the level of uncertainty of the sensor outputs the output noise. The resolution of an SPR sensor, rRI, is typically expressed in terms of the standard deviation of noise of the sensor output, σso, translated to the refractive index of bulk medium, rRI ) σso/SRI, where SRI is the bulk refractive index sensitivity.

where N is the number of intensities used for the calculation of the centroid, σth is the total intensity noise at the threshold, d is the difference of intensities at the SPR dip minimum and at the threshold, w is the width of the dip, SRI is the bulk refractive index sensitivity of the sensor, and K is a factor depending on the relative contributions of the sources of noise.31 As follows from eq 11, the noise in the sensor output is mainly determined by the ratio of the noise in the light intensity at the threshold and the depth of the SPR dip. The ratio w/SRI depends only weakly on the choice of coupler and modulation and therefore has only a minor effect on the sensor resolution.31 Although the analysis was performed for the spectroscopic sensors (N > 1), it can be extended to intensity-modulated SPR sensors (N ) 1). Equation 11 also explains why spectroscopic SPR sensors typically exhibit better resolution than their intensity-based counterpartssthe N values in spectroscopic SPR sensors are typically of the order of 100,62-64 which improves resolution by an order of magnitude. Another important conclusion is that in the intensity-modulated SPR sensor σRI is proportional to σth/I, which for most kinds of noise decreases with increasing intensity of light. Ran and Lipson performed theoretical and experimental comparisons of resolution of intensity and phase modulationbased SPR sensors.65 They demonstrated that under identical noise conditions, the performances of SPR sensors based on

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intensity and phase modulations are comparable. This is due to the facts that in phase modulation-based sensors not the phase but the intensity of a light beam (produced by interference) is measured and that the configurations providing higher phase sensitivity to the refractive index exhibit highest absorption and consequently worse signal-to-noise ratios. They also demonstrated that the refractive index resolution can be improved for both the modulation approaches if the intensity of light coupled to the surface plasmon and received by the detector is increased,65 which is consistent with the results of Piliarik and Homola.31

3.3.3. Limit of Detection and Minimum Resolvable Surface Coverage In the field of SPR sensors, the term resolution usually refers to a bulk refractive index resolution. On the other hand, the limit of detection (LOD) is usually defined as the concentration of analyte that produces sensor output corresponding to 3 standard deviations of sensor output measured for a blank sample.66 The ultimate LOD can be predicted only when the parameters of the interaction between the analyte and biorecognition element and mass transport to the sensor surface are known. However, the smallest detectable surface concentration (minimum resolvable surface coverage) can be determined independently of these factors. The minimum resolvable change of molecular mass captured by the biorecognition elements σΓ depends on the sensor sensitivity and the noise in the sensor output:

σΓ )

σSO h Sh ∂n ∂c vol

( )

(12)

(∂n/∂c)vol denotes the volume refractive index increment of the molecular concentration, and Sh denotes the refractive index sensitivity of sensor output to a refractive index change within the sensitive layer of a thickness h. For thicknesses much smaller than the penetration depth of the surface plasmon (see eq 10), the following relationship between the bulk refractive index resolution and the resolution of surface coverage can be written:

Lpd σΓ ) σRI ∂n 2 ∂c vol

( )

(13)

For an SPR sensor operating at the wavelength of 760 nm (Lpd ) 320 nm) and a typical analyte with a refractive index increment (∂n/∂c)vol ) 0.18cm3/g (DNA or BSA48), eq 13 suggests that a refractive index resolution of σRI ) 10-6 RIU corresponds to a surface coverage resolution of σΓ ) 0.91 pg/mm2.

4. Advances in SPR Biosensor Technology An SPR affinity biosensor consists of a biorecognition element and an SPR transducer. The core of the transducer is an optical platform in which a surface plasmon is optically excited and interrogated and the binding between a biorecognition element (e.g., antibody) immobilized on the surface of the transducer and target analyte in a liquid sample is measured. An SPR biosensor also incorporates a fluidic system that usually consists of a flow cell or cuvette

confining the sample at the sensing surface and a samplehandling system for sample collection and preparation. In the following sections, recent advances in the two most critical elements of the SPR biosensor technologysoptical platforms and biorecognition elements and their immobilizationsare reviewed.

4.1. Optical Platforms Used in SPR Sensors 4.1.1. SPR Sensors Based on Prism Couplers Most of the SPR sensors developed to date, including the first reported SPR sensor,13 use a prism coupler to couple light to a surface plasmon. Prism coupling is convenient and can be realized with simple and conventional optical elements. Moreover, it can be readily combined with any type of modulation. Sensors Based on Intensity Modulation. Research into SPR sensors with intensity modulation focuses mainly on the two important aspectssimproving performance (sensitivity, resolution) and increasing throughput. To increase the sensitivity of intensity-modulated SPR sensors, Lechuga’s group proposed an approach based on combination of the magneto-optic activity of magnetic materials and a surface plasmon resonance in a special multilayer structure.67 They demonstrated an improvement in sensitivity by a factor of 3 compared to a conventional intensity-modulated SPR sensor and a refractive index resolution of 5 × 10-6 RIU.67 A typical example of a high-throughput SPR sensor is the SPR imaging.68,69 In a typical SPR imaging configuration, a beam of monochromatic light passes through a prism coupler and is made incident on a thin metal film at an angle of incidence close to the coupling angle. The intensity of reflected light depends on the strength of the coupling between the incident light and the surface plasmon and therefore can be correlated with the distribution of the refractive index at the surface of the metal film.68,69 Corn’s group has researched SPR imaging for over a decade. In their earlier works, they employed a HeNe laser as a source of illumination.69 However, a highly coherent light source generated images with parasitic interference patterns that were disturbing SPR measurements. In 1997 they improved their SPR imaging instrument by introducing an incoherent light source and a NIR narrow band-pass filter.70 Using this approach, they detected hybridization of short (18-base) oligonucleotides at concentrations as low as 10 nM71 (this was estimated to correspond to a refractive index resolution in the 10-5 RIU range). The use of a white light source and a bandpass filter was also advocated by Yager’s group.72 They demonstrated that by tilting the interference filter, an operating wavelength of the SPR imaging sensor can be tuned.72 Later they demonstrated that their SPR imaging instrument operating at a wavelength of 853 nm can provide a refractive index resolution of 3 × 10-5 RIU.55 In 2005 Corn’s group reported SPR imaging with a special multilayer structure supporting long-range surface plasmons; however, the use of long-range surface plasmons led only to minor sensitivity improvements of 20% (experiment) and 40% (theory) compared to the conventional SPR imaging.73 A dual-wavelength SPR imaging system was reported by Zybin et al.74 In their SPR sensor, they used two sequentially switched-on laser diodes, and the intensities of the reflected light at the two different wavelengths were measured and

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Figure 4. Concept of SPR biosensing.

Figure 5. Typical image obtained with an SPR imaging sensor with a polarization control. Bright rectangles correspond to areas of an SPR chip (300 × 300 µm) covered with a monolayer of albumin molecules formed on the surface of gold by microspotting.

the sensor output was defined as the difference of these two signals. A refractive index resolution of 2 × 10-6 RIU was achieved when the signal was averaged over a large beam diameter (6 mm2). Campbell’s group reported an SPR imaging system with a controllable angle of incidence.75,76 This feature allows SPR images to be acquired sequentially at different angles of incidence and selection of the optimum angle of incidence for the SPR measurements. With a HeNe laser as a source of light, their sensor was able to measure simultaneously in 120 sensing channels with a refractive index resolution of 2 × 10-5 RIU. Recently, they claimed an improvement in sensor resolution down to 5 × 10-6 RIU.77 Recently, Piliarik et al. investigated SPR imaging with an elliptically polarized light78 and concluded that a change in the polarization of light induced by the coupling of light to a surface plasmon can be exploited to significantly improve the sensitivity and operating range of SPR imaging sensors. In addition, this approach, as illustrated in Figure 5, provides high-contrast SPR images (with a low background), which are well suited for automated image analysis. Homola’s group developed an SPR imaging approach based on polarization contrast and excitation of surface plasmons on spatially patterned multilayers.53 In this configuration a prism coupler with an SPR chip containing a spatially patterned multilayer structure was placed between two crossed polarizers. The output polarizer blocked all of the light reflected from the (inactive) areas outside the sensing areas, generating high-contrast images. Two types of SPR multilayers with opposite sensitivities to refractive index were employed, and the output signal was defined as a ratio of the intensities generated from the two neighboring

Homola

multilayers. This sensor was shown to be able to detect refractive index changes down to 2 × 10-6 RIU and to detect short oligonucleotides (23-mers) at concentrations as low as 100 pM.79 Currently, commercial SPR imaging instruments are available from GWC Technologies, Inc.80 (Madison, WI), Lumera81 (Bothell, WA),29 IBIS Technologies (Hengelo, The Netherlands),82 and SPRi-Array from GenOptics (Orsay, France).83 Sensors Based on Spectroscopy of Surface Plasmons. In sensors based on spectroscopy of surface plasmons, the angular or wavelength spectrum of a light wave coupled to a surface plasmon is measured and sensor output is related to a change in the angular or wavelength position of the SPR dip. In the early 1990s, an angular modulation-based SPR sensor consisting of a light-emitting diode (LED), a glass prism, and a detector array with imaging optics was introduced.58,84,85 A divergent beam produced by the LED was collimated and focused by means of a cylindrical lens to produce a wedge-shaped beam of light, which was used to illuminate a thin gold film on the back of a glass prism containing several sensing areas (channels). The imaging optics consisted of one imaging and one cylindrical lens ordered in such a way that the angular spectrum of each sensor channel was projected on separate rows of the array detector.86-88 This design has been adopted by Biacore and resulted in a family of commercial SPR sensors with high performance (resolution down to 1 × 10-7RIU) and multiple sensing channels (up to four). In 2004, Thirstrup et al. integrated several optical elements into a single sensor chip.63 In this approach, the cylindrical focusing optics utilized to create a beam of a desired angular span was replaced by a diffraction grating of a special design incorporated into the sensing element.89,90 A wide parallel light beam was diffracted by the focusing grating and focused into a small spot on the SPR measuring surface. The reflected light followed a similar path, producing a parallel beam with an angular spectrum superimposed across the beam. A twodimensional photodetector was used to measure the angular spectrum of the reflected light for several parallel channels. This design offered a compact SPR platform with a resolution of about 5 × 10-7 RIU.89 An SPR sensor with wavelength modulation and parallel channel architecture was reported by Homola’s group.91 In this sensor, a polychromatic light from a halogen lamp was collimated into a large-diameter parallel beam, which was launched in a prism coupler. The light reflected from different sensing channels was collected by different output collimators coupled and transmitted to different inputs of a spectrograph. The SPR sensor of this design was demonstrated to be able to resolve refractive index changes down to 2 × 10-7 RIU.62 An SPR sensor with wavelength division multiplexing (WDM) of sensing channels was proposed by Homola et al.92 In this approach, signals from multiple surface plasmons excited in different areas of a sensing surface are encoded into different regions of the spectrum of the light wave. Two configurations of WDMSPR sensors have been developed.93,94 In the first configuration, a wide parallel beam of polychromatic light is made incident onto a sensing surface consisting of a thin gold film, a part of which is coated with a thin dielectric film. As the presence of the thin dielectric film shifts the coupling wavelength to a longer wavelength

SPR Sensors

(compared to the bare gold), the reflected light exhibits two dips associated with the excitation of surface plasmons in the area with and without the overlayer.93 The second configuration of WDMSPR sensor employs a special prism coupler in which a polychromatic light is sequentially incident on different areas of the sensing surface at different angles of incidence. Due to the different angles of incidence, the surface plasmons in different regions are excited with different wavelengths of the incident light.94 Therefore, the spectrum of transmitted light contains multiple dips associated with surface plasmons in different areas of the sensing surface. The WDMSPR approach was combined with the parallel architecture, yielding an eight-channel SPR sensor with a resolution around 1 × 10-6 RIU.94 An optical sensor based on spectroscopy of long-range surface plasmons was reported by Nenninger et al.95 In that work, a long-range surface plasmon was excited on a special multilayer structure consisting of a glass substrate, a Teflon AF layer, and a thin gold layer. A resolution as low as 2 × 10-7 RIU was achieved.95 Most recently, Homola’s group demonstrated an improved configuration for excitation of long-range surface plasmons and demonstrated an SPR sensor with a resolution as low as 3 × 10-8 RIU.96 Development of portable/mobile SPR sensor platforms suitable for deployment in the field presents an important direction in SPR sensor research. Several miniaturized SPR optical platforms based on spectroscopy of surface plasmons have been developed. A concept of the miniature SPR sensor based on integration of all electro-optical components in a monolithic platform developed by Texas Instruments in the mid-1990s97 was further advanced by researchers at Texas Instruments and the University of Washington. The Spreeta 2000 SPR sensor (Texas Instruments, USA) consists of a plastic prism molded onto a microelectronic platform containing an infrared LED and a linear diode array detector. The LED emits a diverging beam that passes through a polarizer and strikes the sensor surface at a range of angles. The angle at which light is reflected from this surface toward a detector varies with the location on the surface. The initial version of this platform exhibited a refractive index resolution of 5 × 10-6 RIU.98 Baseline noise and smoothness of response of this sensor were investigated by Chinowsky et al.,99 who showed that the baseline noise established under constant conditions was <2 × 10-7 RIU; however, the sensor response to a gradual change in the refractive index revealed departures from the expected sensor output of about 8 × 10-5 RIU. A portable SPR instrument based on the Spreeta 2000 design was reported by Naimushin et al.100 Their instrument incorporated temperature stabilization and was demonstrated to provide a refractive index resolution of 3 × 10-6 RIU. Recently, an SPR instrument based on two Texas Instruments Spreeta devices was adopted for deployment in a surrogate unmanned aerial vehicle and applied for detection of airborne analytes.101 Another compact, portable SPR sensor platform was developed by Kawazumi et al.52 Their system was also based on angular modulation and Kretschamnn geometry. A line-shape light beam from an LED was focused on the sensing surface using a cylindrical lens. Two-dimensional Fourier transform images of the reflected light were measured with a compact CCD camera. SPR measurements were carried out at a fixed angle, and the resonance angles were obtained by analyzing the twodimensional images. The system provided four independent

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sensing channels and a refractive index resolution of 10-4 RIU.52 Nowadays numerous SPR sensors based on spectroscopy of surface plasmons are commercially available. They include Biacore systems from Biacore (now part of GE Healthcare, USA),88 Spreeta sensor from Texas Instruments (Dallas, TX),102 Multiskop system from Optrel (Kleinmachnow, Germany),103 SR 7000 platforms from Reichert Analytical Instruments (Depew, NY),104 Plasmonic from Hofmann Sensorsysteme (Wallenfels, Germany),105 Autolab Esprit and Springle SPR systems from Eco Chemie (Utrecht, The Netherlands),106 SPR-20 from DKK-TOA Corp. (Tokyo, Japan),107 BIOSUPLAR 6 from Analytical µ-Systems (Sinzing, Germany),108 and Sensia β-SPR Research Platform (Madrid, Spain).109 SPR Sensors Based on Phase Modulation. The research group of Nikitin demonstrated two SPR sensor platforms based on interferometry.110,111 The first approach was based on the interference of the TM-polarized signal beam with the TE-polarized reference beam,110 whereas the second method was based on a Mach-Zehnder interferometer combining TM-polarized signal and reference beams.111 This configuration was demonstrated in two modes: (a) phase contrast (Zernike phase contrast) increasing the sensor sensitivity and (b) “fringe mode”, in which there was a definite angle between the interfering beams and a pattern of interference fringes was superimposed on the image of the surface. Local variations in the phase of the signal beam resulted in bending and moving of the interference fringes. A refractive index resolution was on the order of 10-7 RIU.111 At the same time as Nikitin et al. published their work,111 a similar configuration of SPR sensor based on the MachZehnder interferometer was reported by Notcovich et al.112 They used their system for measurement of the refractive index of gases and demonstrated a refractive index resolution on the order of 10-6 RIU.112 Wu et al. proposed a phase-modulation SPR sensor based on common-path, heterodyne interferometry.113 Two acoustooptic modulators were used to split the incoming laser light from a HeNe laser into two linearly orthogonally polarized beams with a frequency difference of 60 kHz. These two light beams were merged into one beam by a polarization beam splitter. One portion of the beam was directed to a detector while the other was coupled into an SPR prism coupler. The TE and TM components of light reflected from a thin layer of gold on the base of the prism were recombined using a polarizer, and the output beam was received by a detector. SPR-induced phase shift was determined by an electronic phase meter (lock-in amplifier). The refractive index resolution of this design was estimated to be 2 × 10-7 RIU.113 Alieva and Konopsky developed an SPR sensor based on interference between a surface plasmon supported on a metal film and a bulk wave propagating at grazing angle in the flow cell just above the surface of metal.114 This approach suppresses the sensitivity of the SPR method to variations in the refractive index of a liquid sample, which in SPR biosensors interfere with binding measurements. Naraoka and Kajikawa reported a phase-modulation SPR sensor based on a rotating analyzer method.115 In their approach, a linearly polarized light from a semiconductor laser was coupled to an SPR prism coupler and reflectivity was detected while the rotational angle of the analyzer was scanned. The phase difference between the TE and TM

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components of the reflected light were determined from the dependence of the reflectivity on the angle of analyzer. The refractive index resolution of their system was estimated to be below 2 × 10-7 RIU.115 In recent years, SPR sensors with phase modulation have been extensively studied by the researchers at the Chinese University of Hong Kong and City University of Hong Kong. In 2002 Ho et al. reported an SPR sensor based on the Mach-Zehnder interferometer.116 In that work, an optical beam from an unpolarized HeNe laser passed through a prism coupler and the TM-polarized component of the beam excited surface plasmons in two parallel sensing channels, one filled with a sample and the other with a reference solution. TE and TM polarization components of the output beam were split by a polarizing beam splitter. The optical path for TE polarization was modulated by means of a piezoelectric actuator. Finally, TM polarization was converted to TE polarization by a half-waveplate, and the two beams were recombined. The shift between the interference patterns for the measuring and reference sensing channels was measured. Resolution of this sensor was estimated to be about 3 × 10-6 RIU.116 An alternative configuration of an SPR sensor with phase modulation was reported by Wu et al. in 2004.117 In this configuration, one Mach-Zehnder interferometer performed independent interference of TE and TM polarized components of a signal beam emerging from a prism coupler and a reference beam. Subsequently, the output TE and TM beams were separated in a Wollaston prism and directed to two separate detectors. A piezoelectric actuator modulated the optical path in a reference arm of the interferometer, producing a periodic intensity modulation in both TE and TM polarizations. The interference patterns for TE and TM polarizations were processed to reduce the noise and compensate for instabilities in the setup,118 and the sensor output was determined as a mutual shift of the two patterns. Resolution of the sensor was estimated to be 5.5 × 10-8 RIU.117 Another SPR sensor based on measuring the phase difference between TE and TM polarization components of light beam was reported by Ho et al.119 They used a single beam and a photoelastic phase modulator to introduce a carrier frequency so that the phase can be determined by measuring the relative amplitude of the first harmonic signal. Resolution of the sensor was determined to be 1.2 × 10-6 RIU.119 In 2007, a phase modulation-based SPR sensor employing a Michelson interferometer was reported by Yuan et al.120 They used a Michelson interferometer with an SPR prism coupler inserted in one arm of the interferometer. This arrangement allows the TM-polarized component of the light beam to incur a 2-fold phase shift compared with that in the Mach-Zehnder interferometer. Therefore, the sensitivity of the Michelson interferometer-based SPR sensor was twice as high as that of the SPR sensor with the Mach-Zehnder interferometer. This improvement was demonstrated in a single experimental system incorporating both Michelson and Mach-Zehnder interferometers in which the refractive index resolutions were established to be 7.7 × 10-7 and 1.5 × 10-6 RIU, respectively.120

4.1.2. SPR Sensors Based on Grating Couplers Grating couplers have not been used in SPR sensors as widely as the prism couplers. However, their compatibility with mass production (in particular, replication into plastic) makes a grating coupler an attractive approach for fabrication of low-cost SPR sensing structures.

Homola

In 2001 Brockman and Fernandez presented an SPR imaging device based on grating coupling.121 In this approach, a collimated monochromatic light beam (wavelength ) 860 nm) was made incident onto a plastic chip with a gold-coated diffraction grating. An array of 400 sensing channels (spot diameter ) 250 µm) was prepared on the chip by means of spatially resolved functionalization. Upon reflection from the chip, the light was projected onto a twodimensional CCD array.121 This concept was further developed by HTS Biosystems (Hopkinton, MA).122 In 2005, Biacore International AB acquired the FLEXChip technology from HTS Biosystems. Another high-throughput SPR sensor based on grating coupling was reported by Homola’s group.64 This approach was based on angular spectroscopy of surface plasmons on an array of diffraction gratings. A collimated beam of monochromatic light was focused with a cylindrical lens on a row of gold-coated diffraction gratings and reflected under nearly normal incidence. The angular spectra were transformed back to a collimated beam by means of a focusing lens and projected onto a two-dimensional CCD detector. Different rows of gratings were read sequentially by moving the beam splitter and cylindrical lens with respect to the sensor chip. A refractive index resolution of 5 × 10-6 RIU was achieved for simultaneous measurements in over 200 sensing channels.64 Homola’s group reported an SPR sensor based on simultaneous spectroscopy of multiple surface plasmons on a multidiffractive grating.123 A polychromatic light beam was made incident onto a special metallic grating with a grating profile composed of multiple harmonics. The reflected light contained multiple SPR dips, one for each grating period. The use of multiple surface plasmons of different field profiles can provide more detailed information about the refractive index distribution at the sensor surface. This approach was illustrated in a model experiment in which three-surface plasmon spectroscopy was used to determine background refractive index variations and changes in the thickness and refractive index of a bovine serum albumin (BSA) multilayer.123 Recently, an SPR sensor based on twoplasmon spectroscopy on a bi-diffractive grating was investigated in terms of its ability to distinguish contributions to sensor response due to refractive index changes at the sensor surface (i.e., binding) and due to refractive index changes in the whole sample. Theoretical analysis yielding an estimate of an error of such decomposition was reported.124 Recently, Homola’s group reported an SPR biosensor using both longrange and short-range surface plasmons excited simultaneously on a diffraction grating of a special design.125 This approach offers several interesting features such as extended probe depth of the long-range surface plasmon and ability to distinguish sensor response caused by bulk and surface refractive index changes. The sensor was demonstrated to be able to detect changes in the refractive index as small as 3.5 × 10-6 RIU.125 An SPR sensor based on angular spectroscopy of surface plasmons was reported by Unfricht et al. in 2005.126 In their configuration, an LED source was moved by an angle encoder along an arc centered on the chip in order to change the incident angle. The light reflected from the chip surface was detected using a CCD camera that captured sequential images across the range of interrogated angles, and the coupling angle was measured.126

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Figure 6. Concept of SPR sensor based on simultaneous excitation of surface plasmons by a polychromatic light and the dispersion of light on a special grating coupler.127

A new approach to the development of sensors based on spectroscopy of surface plasmons on diffractive gratings was reported by Telezhnikova and Homola.127 A collimated beam of polychromatic light was made incident on a special diffraction grating. A portion of incident light was coupled to a surface plasmon at the metal-dielectric interface via the second order of diffraction. Simultaneously, the light diffracted into the first diffraction order was dispersed and the light components of different wavelengths were directed to different areas of a position-sensitive detector (Figure 6). The coupling of light into a surface plasmon resulted in a drop in the intensity of diffracted light, which was observed as a narrow dip in the spectrum of diffracted light.127 A refractive index resolution of this sensor was established at 3 × 10-7 RIU. In 2007, Chien et al. reported an SPR sensor in which light was coupled into a metal-dielectric waveguide by a subwavelength grating.128 In this configuration, white light was made incident on the waveguiding layer through the grating structure and was coupled to a surface plasmon supported by a thin metal film. An SPR dip was observed in the spectrum of reflected light. A refractive index resolution of this sensor was established at 1 × 10-6 RIU.128

4.1.3. SPR Sensor Based on Waveguide Couplers Fiber optic SPR sensors present the highest degree of miniaturization of SPR sensors. The first fiber optic SPR sensors were reported in the early 1990s.129-131 In SPR sensors based on side-polished single-mode optical fibers, a fundamental mode of the fiber couples to a surface plasmon at the outer surface of a metal layer deposited on a sidepolished region of the fiber. The coupling results in attenuation of a transmitted light at a fixed wavelength (intensitymodulated SPR sensor) or a characteristic dip in the spectrum of transmitted light (wavelength-modulated SPR sensor). The main challenge for obtaining a stable performance is the sensitivity of polarization of light guided in the fiber to deformations of the fiber. The deformations generate changes in the strength of coupling between the light and a surface plasmon and thus interfere with SPR measurements. In 1999 Homola’s group reported an intensity-modulated fiber optic SPR sensor with a resolution better than 2 × 10-5 RIU.132 Later, they reported a wavelength-modulated version of this sensor with resolutions of 5 × 10-7 RIU (no deformations) and 3 × 10-5 RIU (under moderate deformations).133 In 2003 the sensitivity of the sensor to deformation of the fiber was

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dramatically reduced by the introduction of polarizationmaintaining fibers. An SPR sensor based on a side-polished polarization-maintaining fiber was demonstrated to exhibit a refractive index resolution of 2 × 10-6 RIU.134 Chiu et al. reported a fiber optic SPR sensor based on a D-shape optical fiber and heterodyne interferometry.135 Their sensor measured refractive index changes down to 2 × 10-6 RIU. In 2007 Lin et al. extended this approach to multimode fibers and reported an SPR sensor with wavelength modulation based on a side-polished multimode optical fiber.136 The sensor was shown to be able to resolve refractive index changes as small as 3 × 10-6 RIU. An integrated optical SPR sensor with intensity modulation and one sensing channel and one reference channel was reported by Mouvet et al.137 The signal from the sensing channel was normalized to the signal from the reference channel, resulting in an increased stability and a refractive index resolution of 5 × 10-5 RIU.138 A wavelength modulation-based integrated optical SPR sensor was reported by Homola’s group. The sensor was demonstrated to provide a refractive index resolution as low as 1 × 10-6 RIU.139 An SPR sensor based on a strip waveguide consisting of a germanium-doped silicon dioxide waveguiding layer on a silicon substrate and wavelength modulation was reported by Huang et al.140 Their sensor exhibited a sensitivity similar to that reported in ref 139. In conventional waveguide-based SPR sensors, the resonant coupling between a surface plasmon and a waveguide mode occurs for refractive indices of sample considerably higher than the refractive index of a typical aqueous sample. Various approaches to control the operating range of waveguide-based SPR sensor, so that it includes aqueous environments, were proposed. They include an integrated optical waveguide fabricated in low refractive index glass,141 a buffer layer,142 or a high refractive index overlayer.143 Skorobogatiy and Kabashin proposed to overcome this limitation of conventional waveguides by employing a photonic-crystal waveguide.144 In their paper, they theoretically demonstrated the feasibility of a photonic crystal waveguide-based SPR sensor using a single-mode photonic crystal waveguide in which the effective index of a mode confined in the waveguiding layer can be made considerably smaller than the refractive index of the waveguiding layer material, enabling phase matching with a surface plasmon at any wavelength.144 Debackere et al. proposed an interferometric SPR sensor consisting of a silicon-on-insulator (SOI) waveguide and a thin metal layer.145 In their design, a mode guided by a thin silicon film excited two surface plasmons on opposite sides of a thin metal film. These two modes propagated side by side over a short distance (10 µm) and were recombined in a waveguide mode at the end of the metal film. The transmitted intensity depended on the mutual phase delay of the two interfering surface plasmons. As the propagation constant of the surface plasmon propagating at the outer boundary was sensitive to the refractive index of the adjacent medium (sample), changes in the refractive index of sample could be measured by measuring changes in the intensity of transmitted light. As this sensor is based on interference rather than phase-matching, the operating range of the sensor can be conveniently controlled by the geometry of the device.145 Wang et al. reported an alternative approach to integrated optical SPR sensing based on electro-optical modulation.146

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Their sensor consisted of titanium-diffused channel waveguide in a lithium niobate substrate in which the propagation constant of a mode was modulated by voltage applied between the electrodes on two sides of the waveguiding layer. In the amplitude modulation mode, the intensity of transmitted light was measured as the electrode voltage was scanned and the slope of the intensity-voltage dependence was correlated with the amount of analyte captured at the sensor surface. The same approach was demonstrated for the wavelength modulation mode in which the resonant wavelength was measured.146

4.2. Biorecognition Elements and Their Immobilization In SPR affinity biosensors, one of the interacting molecules (biorecognition element or target molecule) is immobilized on the solid surface of the SPR sensor and the other is contained in a liquid sample. Which of the molecules is immobilized depends on the used detection format (section 5.1)sin direct, sandwich, and competitive detection formats, the molecule that needs to be immobilized is a biorecognition element; in the inhibition detection format, the immobilized molecules is the target molecule or its derivative. The choice of appropriate biorecognition elements and immobilization method is of critical importance with direct impact on key performance characteristics of the sensor such as sensitivity, specificity, and LOD.

4.2.1. Biorecognition Elements Various kinds of biorecognition elements have been employed in affinity SPR biosensors. Antibodies remain by far the most frequently used biorecognition element. They offer high affinity and specificity against target analyte. Moreover, antibodies against numerous target molecules are now commercially available. Development of high-quality antibodies is, however, a rather expensive and laborious process.147 Recently, single-chain antibody fragments (scFvs) have been also used as biorecognition elements.148 Biotinylated scFv fragments expressed in yeast can be spotted on streptavidin-coated sensor surfaces directly from cell supernatant without the need of purification.149 Another type of biorecognition element that has been employed in SPR sensors are peptides. In comparison with antibodies, peptides, in general, are inexpensive, more stable, and easier to manipulate. However, peptides sometimes lack high affinity and specificity against the target. In SPR biosensors, peptides have been applied mainly for the detection of antibodies, for example, antibodies against hepatitis G,150 herpes simplex virus type 1 and type 2,151 and Epstein-Barr virus,152 and for the detection of heavy metals.153 Recently, aptamers emerged as another promising type of biomolecular recognition element for SPR biosensors.154,155 DNA or RNA aptamers are single-stranded oligonucleotide sequences, which can be produced to bind to various molecular targets such as small molecules, proteins, nucleic acids, and even cells, tissues, and organisms.155,156 Moreover, the synthesis of aptamers is straightforward and reproducible.

4.2.2. Immobilization of Biorecognition Elements In SPR biosensors, one of the interacting molecules (mostly biorecognition element) is immobilized on the sensor surface. The surface chemistry has to be designed in such a way that it enables immobilization of a sufficient number

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of biorecognition elements on the sensing surface while minimizing the nonspecific binding to the surface. In addition, biorecognition elements need to be immobilized on the sensor surface without affecting their biological activity. In principle, the molecules can be immobilized either on the surface or in a three-dimensional matrix. Although immobilization on surfaces is more straightforward to perform, the number of accessible biorecognition elements is limited by the capacity of the surface (too high density of immobilized biorecognition elements can lead to lower response due to sterical hindrance). Immobilization in a threedimensional matrix typically provides more binding sites than immobilization on the surface and a better environment for the preservation of immobilized molecules during prolonged storage.157 The most widely used three-dimensional matrix for immobilization of molecules in a structured environment is the carboxymethylated dextran.158 For two-dimensional (surface) immobilization of biorecognition elements on the sensing (gold) surface, selfassembled monolayers (SAMs) of alkanethiolates or disulfides have been widely used.159 To provide a desired surface concentration of biomolecular recognition elements and a nonfouling background, mixed SAMs of long-chained (n ) 12 and higher) alkanethiolates terminated with a functional group for further attachment of biomolecular recognition elements and oligo(ethylene glycol)-terminated shorterchained alkanethiolates for a nonfouling background have been developed.160,161 The main approaches to immobilization of molecules to the surface of SPR sensors are based on physical absorption and hydrophobic and electrostatic interactions,162 covalent coupling,158 and attachment of tagged molecules by a sitespecific non-covalent interaction between the tag and an immobilized capture molecule via biotin-avidin163 or histidine-chelated metal ion164 interaction or DNA hybridization.165 A more detailed account of immobilization methods is given in refs 31 and 166. As proteins represent the class of molecules that are most frequently used as biorecognition elements, the following section is dedicated to the immobilization of proteins. The most commonly used approach to the immobilization of proteins is via a covalent bond formed between the nucleophilic functional groups supplied by amino acids (e.g., amino groups, lysine; thiol groups, cysteine) of the protein and electrophilic groups (e.g., activated carboxyls, aldehydes) on the sensor surface.158,167 The covalent immobilization is stable; however, as proteins typically contain many functional groups, immobilization via these functional groups results in random orientation of immobilized proteins. Moreover, simultaneous immobilization via multiple functional groups may restrict conformational flexibility of the protein and impair its function. Another approach to the immobilization of proteins is based on biochemical affinity reaction. The most common example of this approach is the immobilization based on avidin-biotin chemistry. In this immobilization method, protein avidin (or a closely related streptavidin) is immobilized on the sensing surface (covalently or via preimmobilized biotin) and provides binding sites for subsequent attachment of a biotin-conjugated protein. The protein can be biotinylated by various methods targeting different groups on the protein. Orientation of the immobilized proteins depends on the orientation of avidin/streptavidin molecules, the biotinylation method used, and the properties of the protein. Alternatively, antibodies can be immobilized via

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interaction between the Fc region of the antibody and protein A or protein G. This method provides good access to the binding site of the antibody, located on the Fab variable region; however, to control orientation of the antibody, the orientation of protein A itself needs to be controlled.168 Recently, DNA-directed immobilization of antibodies has been described.165 This approach takes advantage of DNA chip technology, which provides an exceptionally stable pattern of single-stranded DNA (ssDNA) sequences and uses it for the immobilization of proteins conjugated with complementary ssDNA sequences via DNA hybridization. This approach provides an elegant and flexible platform both for SPR sensors with a limited number of sensing channels and for high-throughput screening SPR systems. The drawback of this method is that it requires that the protein is conjugated with ssDNA.165 Another bioaffinity immobilization method is based on the interaction between histidinetagged protein and chelated metal ions, for example, nitrilotriacetic acid (NTA) immobilized on the surface and loaded with bivalent metal cations.164 NTA can be attached to the sensing surface covalently via acetic group using EDC/ NHS chemistry.169 Recombinant proteins with affinity tags can be produced by genetic engineering. As the tag can be placed at a defined position on the protein, this approach allows site-specific and thus highly ordered protein immobilization. Moreover, the binding between chelated metal ions and histidine is reversible and the immobilized protein can be released by introduction of a competing ligand (e.g., imidazole) or a chelation agent (e.g., ethylenediaminetetraacetic acid, EDTA). Whereas reusability of the surface is one of the main advantages of this immobilization method, the main drawback is the low affinity of the His-tag to an individual chelator. In 2005 Piehler’s group expanded this immobilization technique by designing supramolecular entities binding to oligohistidine tags with high affinity and stability.170 They designed supramolecular multivalent chelator heads (MCH) containing multiple NTA moieties and investigated their binding with hexahistidine (H6)- and decahistidine (H10)-tagged molecules. It was demonstrated that the binding stability of the complex increases with an increasing number of NTA moieties. An improvement of the stability of the chelator-oligohistidine complex by 4 orders of magnitude compared to that of mono-NTA was achieved.170 Tinazli et al. developed multivalent metal-chelating thiols for attachment of histidine-tagged proteins to the surface of SPR sensors via SAMs.171 Dramatically improved stability of protein binding by these multivalent chelator surfaces was observed compared to mono-NTA/His(6) tag interaction. Regenerability of the surface (removal of the protein) using EDTA was also demonstrated.171 Immobilization and arraying of histidine-tagged proteins by combining molecular and surface multivalency was demonstrated by Valiokas et al.172 They employed SAMs formed by triethylene glycolterminated alkyl thiols171 functionalized with either a single NTA moiety (mono-NTA) or a chelator head group containing two NTA moieties (bis-NTA). The process of immobilization of proteins was observed with SPR imaging.172 Immobilization of peptides can follow strategies similar to those developed for proteins, including electrostatic attraction and amine- or thiol-based covalent coupling. The most straightforward approach to the immobilization of oligonucleotides is based on the use of biotinylated derivatives.173 Small molecules with functional groups (aliphatic amines, thiols, aldehydes, or carboxylic groups) can be

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covalently linked to suitable corresponding groups on the sensor surface. Small molecules without suitable functional groups need to be derivatized.174 To deliver molecular recognition elements to different areas of the SPR sensor surface, the immobilization chemistry needs to be spatially controlled. For instance, Campbell’s group demonstrated the microspotting of double-stranded DNA on gold for SPR microscopy using two approaches.175 Both methods use streptavidin and biotinylated oligonucleotides. In the first method, the robotic microspotter was used to deliver nanoliter droplets of dsDNAs onto a uniform layer of streptavidin. In the second method, a streptavidin layer was also microspotted on a mixed-alkanethiolate SAM and, subsequently, microspots of dsDNA were added using microspotting. Homola’s group compared the microspotting technique with the conventional flow-through in situ functionalization.79 In their study, the spatially resolved functionalization based on microspotting applied to immobilization of short oligonucleotides was shown to provide a surface concentration of oligonucleotide probes of about 2.2 × 1012 oligonucleotides per cm2, which was higher by 80% than the surface coverage provided by the flow-through functionalization method.79

4.2.3. Nonfouling Surfaces As the adsorption of proteins is of major concern in numerous important biomedical and biological applications (biocompatible materials for prostheses, tissue engineering, cell culturing, implantable devices, microarrays, etc.), the adsorption of proteins to synthetic surfaces has been the subject of extensive research worldwide. Nonspecific adsorption of proteins to sensing surface presents a key challenge also for affinity biosensors.176,177 This problem is more severe when complex samples such as blood or cell lysate are to be analyzed.178,179 Although the molecular mechanism of protein resistance has not been fully understood, research into protein-resistant surfaces has made significant advances. Various surface coatings displaying low fouling or even nonfouling properties (i.e., exhibiting complete resistance to protein binding and cell colonization) have been proposed. For affinity biosensors it is especially important to create surfaces with low fouling background providing also abundant binding sites for immobilization of biomolecular recognition elements. Hydrophilic polymers such as poly(ethylene glycol) (PEG) and its derivatives have been successfully employed in the design of protein-resistant coatings for SPR sensors. The key factors that influence nonfouling properties of PEG molecules have been considered to be steric-entropy barrier characteristics and a high degree of hydration.180 Recent measurements of interfacial forces have shown that the protein resistance of PEGylated surfaces correlates with a net repulsive force versus distance curve.181 An approach utilizing a poly(Llysine) grafted with poly(ethylene glycol) (PLL-g-PEG) have been used by Pasche et al. to minimize the nonspecific adsorption of proteins.182,183 In their work, they varied the ratio of the number of lysine monomers and the PEG side chains and, with the optimized surface composition, they observed adsorption from blood serum below 2 ng/cm2.182 It was demonstrated that the immobilization of biorecognition elements to PLL-g-PEG surfaces can be performed by introducing biotin to the surface by assembling mixed (PLLg-PEG/PEGbiotin + PLL-g-PEG) from the corresponding mixed solutions. The resulting biotinylated surfaces have been shown to be highly resistant to nonspecific adsorption

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from serum while allowing binding of linkage proteins (e.g., streptavidin or avidin) and subsequent attachment of biotinylated biorecognition elements.184-186 PEG-containing molecules were also successfully employed as the secondary blocking agents on the surface with covalently immobilized proteins.187 Oligo(ethylene glycol) (OEG)-terminated alkanethiolates are widely used to form SAMs on gold surfaces of SPR sensors, typically also in combination with alkanethiolates containing various functional groups.188,189 Several heterobifunctional OEG or PEG surfaces have been employed in affinity biosensors. 190Although OEG-terminated SAMs have shown high resistance to nonspecific protein adsorption, the real biomedical applicability of these surfaces has been limited, mainly due to the limited oxidative stability of thiolates and the difficulty of integrating them into biomedical devices.191,192 It has been reported that PEG molecules autoxidize relatively rapidly, especially in the presence of oxygen and transition metal ions.193 Whitesides and co-workers explored alkanethiolates in SAMs with various functional groups and concluded that inertness of the surface is a general property of a group of surfaces rather than a specific property of ethylene glycol groups. They have also investigated the resistance of SAMs terminated with various groups to bacteria and mammalian cells and found that there was very little correlation between the resistance to the adsorption of protein and the adhesion of cells.191 On the basis of experimental investigation of SAMs with different functional groups, Whitesides’s group concluded that important requirements for protein resistance are (1) hydrophilicity, (2) ability to accept a hydrogen bond, (3) inability to donate to a hydrogen bond, and (4) a net neutral charge.194 Subsequently, they hypothesized that a zwitterionic SAM combining positively charged and negatively charged groups might offer a new type of protein-resistant surface.194 Kitano et al. formed a SAM of zwitterionic telomers on a metal surface and demonstrated its ability to reduce the nonspecific adsorption of proteins.195 Subsequently, Jiang’s group extended this material into a dual-functional zwitterionic poly(carboxybetaine methacrylate) (polyCBMA) polymer using reactive carboxyl groups for protein immobilization.196 They demonstrated a polyCBMA polymer with immobilized antibodies against human chorionic gonadotropin (hCG) that, when exposed to high concentrations of lysozyme and fibrinogen, exhibited nonspecific adsorption of <0.3 ng/cm2 (Figure 7).196 Most recently, Jiang’s group has investigated bacterial adhesion to the zwitterionic poly(sulfobetaine methacrylate) (polySBMA) and demonstrated that polySBMA surfaces dramatically reduce bacterial adhesion.197 Chen et el. showed that oligo(phosphorylcholine) SAMs exhibit strong resistance to protein adsorption, specifically to high concentrations of fibrinogen, lysozyme, and bovine serum albumin.198 In the past several few years, numerous SPR biosensors for detection in complex matrices have been reported. The most frequently targeted complex medium is blood serum, which is a key medium for medical diagnostics applications. Most of the reported SPR biosensors were designed to operate in serum diluted by buffer to serum concentrations from 1 to 25%. The detection of antibodies against EpsteinBarr virus (anti-EBNA) in 1% serum was reported by Homola’s group.199 A synthetic peptide, which was used as a biorecognition element, was immobilized on the surface via hydrophobic and electrostatic interactions. Nonspecific

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Figure 7. Adsorption of 1 mg/mL fibrinogen, 1 mg/mL lysozyme, and 20 µg/mL hCG from PBS on (a) polyCBMA-grafted surfaces and (b) polyCBMA-grafted surfaces with immobilized hCG antibodies. Reprinted with permission from ref 196. Copyright 2006 American Chemical Society.

adsorption from 1% serum was found to be negligible, and the LOD for anti-EBNA was 0.2 ng/mL.199 Ayela et al. reported an SPR sensor for the detection of IA-2 autoantibodies also in 1% human serum.200 In that work various types of mixed SAMs were evaluated in terms of specific and nonspecific binding. It was observed that the nonspecific adsorption from serum to surface coated with EG6SAM100%COOH was about an order of magnitude lower than the adsorption to a SAM100%COOH-coated surface. Using an EG6-SAM25%COOH-coated surface, their SPR sensor was able to detect antibody at a concentration of 0.2 nM.200 Cao et al. detected prostate-specific antigen-1antichymotrypsin (PSA-ACTcomplex) in 10% serum. Their sensor employed a mixed SAM of alkanethiolates terminated with EG6-COOH and EG3-OH groups. The COOH group was biotinylated and used for subsequent immobilization of streptavidin and antibody against the PSA-ACTcomplex. Nonspecific adsorption of albumin, IgG, and fibrinogen on the sensing surface was found to be negligible, and the sensor was demonstrated to be able to detect the PSA-ACT complex at concentrations below 50 ng/mL.201 Chung et al. demonstrated detection of antibody against human hepatitis B virus (hHBV) in serum dilutions from 5 to 60%.202 Their sensor employed a thiol monolayer prepared using 11mercaptoundecanoic acid to which hHBV antigen was coupled using the EDAC/NHS coupling chemistry. The level of nonspecific binding at different concentrations of serum

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was measured; serum concentrations of 5, 10, and 20% produced a nonspecific sensor response corresponding to protein adsorption of 20, 26, and 68 ng/cm2, respectively.202 Miura’s group reported an SPR sensor for the detection of insulin in 10 and 25% serum.203 The sensing surface of their sensor was constructed using a heterobifunctional oligo(ethylene glycol)-dithiocarboxylic acid derivative (OEGDCA) containing dithiol and carboxyl end groups to which insulin was covalently bound. The results observed with OEG-DCA SAMs were further compared to those reported using a SAM of monothiol tethered oligo(ethylene glycol)carboxylic acid (OEG-COOH), and it was concluded that the resistance of the present bare dithiol-tethered OEG-DCA SAM is comparable to or better than that of the monothioltethered OEG-COOH SAM.204 The sensor was able to detect insulin in serum at insulin concentrations down to 6 ng/mL.203 Although these developments present clear progress toward SPR biosensing in serum, an SPR sensor capable of detecting relevant concentrations of analyte in whole serum has not been demonstrated yet.

4.3. Summary The past decade has witnessed development of numerous SPR sensors based on excitation of surface plasmons via prism coupling, waveguide coupling, or diffraction coupling and angular, wavelength, intensity, or phase modulation. SPR sensor platforms based on prism coupling remain by far the most common. SPR sensing platforms providing the highest resolutions are typically based on angular88,127 or wavelength62 spectroscopy of surface plasmons or phase modulation.113,115,117 The best SPR sensor platforms with a limited number of sensing channels (<10) provide a refractive index resolution around 10-7 RIU. High-throughput SPR sensors with a large number of sensing channels (>100) are usually based on intensity modulation (SPR imaging) and offer an order of magnitude worse performance.79 One of the prospective approaches to further improving the resolution of SPR sensors involves long-range surface plasmons. SPR sensors based on wavelength spectroscopy of long-range surface plasmons were demonstrated to be able to deliver resolution as low as 3 × 10-8 RIU.96 However, as the field of long-range surface plasmons extends much farther from the sensing surface than that of conventional surface plasmons, this improvement can be fully harnessed only when large analytes (e.g., bacterial pathogens) are targeted or biorecognition elements are immobilized in a extended coupling matrix. Various types of biorecognition elements and immobilization methods are available to allow the SPR sensing platforms to be tailored for specific detection of chemical and biological substances. Proteins (e.g., antibodies) and peptides are most frequently immobilized via covalent bonds formed between amino groups of the protein and activated carboxyls on a SAM of alkanethiolates or within a dextran matrix. Oligonucleotides can be efficiently immobilized via interaction between avidin or streptavidin immobilized on the sensing surface and biotinylated oligonucleotide. Small molecules are usually conjugated with a larger protein (BSA), which is subsequently (covalently) immobilized on the sensor surface. High-throughput SPR sensors demand immobilization methods capable of accurate spatially controlled delivery of different biorecognition elements to different areas of the sensing surface. This can be achieved by combining the streptavidin-coated surface with a spatially controlled de-

Figure 8. Main detection formats used in SPR biosensors: (A) direct detection; (B) sandwich detection format; (C) competitive detection format; (D) inhibition detection format.

livery of biotinylated biorecognition elements by microspotting.79 An interesting alternative approach is based on the use of a conventional DNA chip and its conversion to a protein chip by incubating the chip with a mixture of proteins conjugated with complementary DNA sequences.165

5. Applications of SPR Sensors for Detection of Chemical and Biological Species SPR biosensors have been applied in numerous important fields including medical diagnostics, environmental monitoring, and food safety and security.

5.1. Detection Formats Various formats for the detection of chemical and biological analytes have been applied in SPR sensors.205,206 The format of detection is chosen on the basis of the size of target analyte molecules, binding characteristics of available biomolecular recognition element, range of concentrations of analyte to be measured, and sample matrix.206 The most frequently used detection formats include (a) direct detection, (b) sandwich detection format, (c) competitive detection format, and (d) inhibition detection format (Figure 8). In the direct detection mode (Figure 8A), the biorecognition element (e.g., antibody) is immobilized on the SPR

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sensor surface. Analyte in solution binds to the antibody, producing a refractive index change detected by the SPR sensor. Direct detection is usually preferred in applications, where direct binding of analyte of concentrations of interest produces a sufficient response. The specificity and LOD can be improved by using the sandwich detection format (Figure 8B), in which the sensor surface with captured analyte is incubated with a second antibody. Smaller analytes (molecular weight < 5000) often do not generate a sufficient change in the refractive index and therefore are measured using either competitive or inhibition detection format. Figure 8C shows an example of the competitive detection format, in which the sensing surface is coated with an antibody interacting with the analyte; when a conjugated analyte is added to the sample, the analyte and its conjugated analogue compete for a limited number of binding sites on the surface. The binding response is inversely proportional to the analyte concentration. In the inhibition detection format (Figure 8D) a fixed concentration of an antibody with affinity to analyte is mixed with a sample containing an unknown concentration of analyte. Then, the mixture is injected in the flow cell of the SPR sensor and passed over a sensor surface to which analyte or its analogue is immobilized. Noncomplexed antibodies are measured as they bind to the analyte molecules immobilized on the sensor surface. The binding response is inversely proportional to the concentration of analyte. In recent years, various modifications and extensions of these basic detection formats have been developed in order to expand and improve detection capabilities of SPR biosensors.207 Several detection formats for the detection of multiple analytes have been reported. Chung et al. modified the sandwich assay approach to allow detection of multiple analytes in a single sensing channel.208 They immobilized antibodies against two different analytes on the same area of the sensor surface. After incubation of sample with the sensor surface, solutions containing respective antibody were sequentially injected. Sensor response to each antibody was proportional to a concentration of the respective analyte.208 The same concept was adopted for inhibition detection format by Lechuga’s group.209 They immobilized three analyte derivatives on the sensor surface, and the sensor response to each analyte was determined by incubation of the sensor with a solution containing a respective antibody.209 Enhancement of sensor sensitivity through the “labeling” of a secondary antibody in the sandwich detection format by latex particles210 or gold nanoparticles211 was demonstrated in the 1990s. In 2005 Mitchell et al. used the labeling approach in the inhibition detection format.212 They used gold nanoparticles to improve the sensitivity of the SPR sensor for detection of progesterone. In their study, the (primary) antibody mixed with a sample was conjugated with biotin. Upon incubation of the sample with a sensor surface coated with progesterone, streptavidin conjugated with a gold nanoparticle (10 or 20 nm in diameter) was injected. The LOD for progesterone in buffer was established at 23 pg/ mL; this corresponds to an improvement by a factor of 17 compared to the inhibition format.212 In 2003 Sato et al. described amplification of SPR response to DNA based on the use of DNA-carrying hydrogel microspheres.213 Acrylamide-based microspheres with carboxyl groups were conjugated with DNA probes. Binding of DNA-carrying acrylamide-based microspheres with target DNAs at the sensor surface resulted in a 100-fold increase in sensitivity

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compared to the sensitivity of nonamplified DNA target hybridization. In 2004 Okamura et al. reported the enhancement of an SPR sensor response by means of hydrogel nanospheres and the same protocol.214 Their hydrogel nanospheres were prepared by precipitation polymerization of acrylamide, methylenebisacrylamide, and methacrylic acid.214 Styrene-glycidyl methacrylate (SG) microspheres prepared by soap-free emulsion copolymerization were proposed by Sato et al.215 They used DNA conjugated with SG microspheres to enhance the signal from hybridization of DNA to a complementary DNA strand immobilized on the SPR sensor surface. They immobilized short ssDNA onto the sensor surface to capture (a longer) target DNA. Subsequently, DNA-carrying microspheres were injected to bind to the free portion of the target DNA.215 In 2006 Komatsu et al. proposed an amplification method suitable for intensitymodulated SPR sensors based on dye-doped polymer particles.216 The dye-doped polymer particles can enhance the sensitivity of intensity-modulated SPR sensors by two mechanismssby the shift in the resonant coupling condition due to the refractive index change induced by the presence of the particles and by absorption of light in the dye-doped particles. In a model experiment, the authors compared sensor responses due to the binding of BSA or BSA conjugated with dye-doped polymer particles to anti-BSA immobilized on the sensor surface. The use of particles was demonstrated to provide a 100-fold improvement in sensor sensitivity.216 Recently, several methods for enhancing the sensitivity of SPR biosensors based on enzymatic amplification were developed by Corn’s group. Goodrich et al. reported an enzymatic amplification method for the detection of DNA molecules that utilizes RNA microarrays in conjunction with the enzyme RNase H.217 In this method, a single-stranded RNA microarray is exposed to a solution containing both the complementary DNA and RNase H. The DNA binds to its RNA complement on the surface and forms an RNADNA heteroduplex. RNase H then binds to this heteroduplex, selectively hydrolyzes the RNA probe, and releases the DNA complement back into solution. The released DNA molecule binds to another RNA probe on the surface, so that a single DNA molecule can initiate the destruction of many surfacebound RNA probes. Eventually, all of the RNA probe molecules are destroyed and removed from the surface. The loss of RNA probe molecules from the surface is detected by the SPR method. Using this approach, DNA solutions were detected at levels down to 10 fM.217 Another enzymatic amplification approach for the detection of DNA was demonstrated by Lee et al.218 In the first step of this approach, an ssDNA array is exposed to a solution containing target DNA and enzyme exonuclease III (ExoIII); the target DNA hybridizes to its complementary ssDNA array elements, and ExoIII binds to the dsDNA. ExoIII selectively hydrolyzes the probe DNA strand from the duplex, releasing the target DNA strand back into solution. The released target DNA is then free to bind to another surface-bound ssDNA probe. This cyclic reaction progresses until all ssDNA probes on the surface are destroyed by ExoIII.218 Using this approach, a 16-mer ssDNA was detected down to 10-100 pM, which presents a 102-103 -fold improvement.218 Fang et al. described an amplification method for the detection of RNA based on poly(A) enzyme chemistry and nanoparticle enhancement.219 In this method the target RNA is adsorbed from solution onto a single-stranded LNA microarray. Subsequently, poly(A) tails are introduced to the surface-

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bound RNAs via the poly(A) polymerase surface reaction. Finally, poly(A) tails are hybridized with T30-DNA-coated Au nanoparticles. This approach was demonstrated to allow detection of RNAs down to 10 fM.219 Recently, Li et al. reported another enzymatic amplification approach.220 In the first step of this method, the target protein binds to the aptamer immobilized on the sensor surface. Then, a horseradish peroxidase (HRP)-conjugated antibody to the target protein is introduced to create an aptamer-protein-antibody sandwich, which is subsequently exposed to the substrate 3,3′,5,5′-tetramethylbenzidine (TMB). Reaction of TMB with HRP gives rise to a dark blue precipitate. This methodology was applied in an SPR sensor for detection of human thrombin. The LOD was demonstrated at a concentration of 500 fM, which corresponds to an enhancement factor of ∼104.220

5.2. Food Quality and Safety Analysis As the acceptance of SPR biosensor technology in food analysis continues to increase, the number of publications on SPR biosensors for the detection of analytes related to food quality and safety increases.221-225 The targeted analytes include pathogens, toxins, drug residues, vitamins, hormones, antibodies, chemical contaminants, allergens, and proteins.

5.2.1. Pathogens In recent years, various pathogens have been targeted by SPR biosensors.226 In particular, they include bacteria, protozoa, fungi, and parasites. Escherichia coli O157:H7 was first detected by SPR by Fratamico et al. in 1998.227 Since then, numerous SPR biosensors for the detection of E. coli O157:H7 have been reported. Choi’s group used the commercial SPR sensor Multiskop (Optrel, Germany) and monoclonal antibodies immobilized on a protein G-coated sensor surface. The sensor was demonstrated to be able to directly detect E. coli O157: H7 at concentrations as low as 104 cells/mL.228 Subsequently, they demonstrated that in conjunction with the immobilization of antibodies via a mixed SAM of alkanethiolates, the same SPR instrument can detect E. coli O157:H7 down to 102 cells/mL.229 Taylor et al. detected E. coli O157:H7 using a custom-built SPR sensor with wavelength modulation and examined the effect of various treatment methods on sensor performance.230 A monoclonal antibody was immobilized on a mixed -COOH- and -OH-terminated SAM of alkanethiolates via amine coupling chemistry. Detection of E. coli O157:H7 was performed in the sandwich detection format using a secondary polyclonal antibody. Detection limits for detergent-lysed bacteria, heat-killed bacteria, and untreated bacteria were determined to be 104, 105, and 106 colonyforming units (cfu)/mL, respectively.230 Detection of E. coli O157:H7 using a commercially available Spreeta sensor (Texas Instruments Co.) was reported by Meeusen et al.231 Biotinylated polyclonal antibody against E. coli O157:H7 was immobilized on the avidinated gold surface. The SPR biosensor was shown to be capable of detecting E. coli O157: H7 in cultures at levels down to 8.7 × 106 cfu/mL in 35 min.231 Another SPR sensor for the detection of E. coli O157: H7 based on the Spreeta sensor was reported by Su and Li.232 In that work, polyclonal E. coli O157:H7 antibodies were immobilized via protein A adsorbed on the sensor surface. The sensor was demonstrated to be able to detect E. coli O157:H7 in an aqueous environment at levels down to 106 cells/mL.232 Subramanian et al. demonstrated an SPR bio-

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sensor for the detection of E. coli O157:H7. They used the commercial SPR sensor SR 7000 (Reichert Analytical Instruments) and attachment of E. coli O157:H7 polyclonal antibodies via alkanethiolate SAM and amine coupling chemistry. Detection of E. coli O157:H7 was performed in sandwich format. The detection limit for E. coli O157:H7 was established at 103 cfu/mL.233 Taylor et al. reported SPRbased detection of E. coli O157:H7 in apple juice using a custom-built multichannel SPR sensor with wavelength modulation and sandwich detection format.234 Biotinylated polyclonal antibodies against E. coli O157:H7 were immobilized via streptavidin attached to a mixed SAM of oligo(ethylene glycol) alkanethiolate and biotinylated alkanethiolate. Detection of heat-killed E. coli O157:H7 was performed in buffer, in a mixture of four bacterial species, and in apple juice. The effect of the pH of the apple juice on the sensor response was investigated, and SPR responses were higher for bacteria in apple juice at pH 7.4 than in apple juice at pH 3.7. The LOD was 1.4 × 104 cfu/mL in buffer and about 105 cfu/mL in apple juice with an adjusted pH of 7.4.234 Waswa et al. used two commercial SPR sensorsslaboratory instrument Biacore 2000235 and a miniature SPR sensor Spreeta236sto detect E. coli O157:H7. Immobilization of E. coli O157:H7 antibody for the laboratory SPR system was performed by first attaching protein A using a carboxymethylated dextran layer and amine coupling chemistry and subsequent attachment of the antibody.235 The LOD for E. coli O157:H7 in pasteurized milk was determined to be 25 cfu/mL.235 In the Spreeta sensor, the biotinylated E. coli O157:H7 antibody was attached to a layer of neutravidin molecules adsorbed on the gold surface.236 The detection limit for E. coli O157:H7 in milk, apple juice, and ground beef was estimated to be in the range of 102-103 cfu/mL.236 Salmonella enteritidis was detected using a custom-built SPR sensor with wavelength modulation by Koubova´ et al.162 In that work, a double layer of antibodies was physisorbed on a bare gold surface and cross-linked with gluteraldehyde. Direct detection of heat-killed, ethanol-soaked S. enteritidis at a concentration as low as 106 cfu/mL was demonstrated.162 Bokken et al. demonstrated detection of Salmonella groups A, B, D, and E using the commercial SPR sensor Biacore 3000.237 Antibodies were immobilized in a carboxymethylated dextran layer via amine coupling chemistry, and detection of Salmonella serotypes was performed using the sandwich format. Salmonella serotypes were detectable at a concentration of 1.7 × 105 cfu/mL even in the presence of other bacteria at 108 cfu/mL levels.237 Choi’s group demonstrated the detection of Salmonella typhimurium using the commercial SPR sensor Multiskop and monoclonal antibodies immobilized via protein G attached to an alkanethiolate SAM on the sensor surface. The LOD was 102 cfu/mL.238 An SPR sensor for the detection of Salmonella paratyphi was demonstrated by the same group.239 They used the same SPR instrumentsa Multiskopsand a similar method for the attachment of monoclonal antibodies via protein G. Detection of S. paratyphi was shown down to concentrations of 102 cfu/mL.239 In 2006 detection of Salmonella in food matrices was reported by two groups. Taylor et al. reported SPRbased detection of Salmonella choleraesuis serotype typhimurium in apple juice using a custom-built multichannel SPR sensor with wavelength modulation and sandwich detection format.234 Biotinylated polyclonal antibodies against Salmonella were immobilized via streptavidin attached to a mixed SAM of oligo(ethylene glycol) alkanethiolate and

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biotinylated alkanethiolate. Detection of Salmonella was performed in buffer, in a mixture of four bacterial species, and in apple juice. The effect of the pH of the apple juice on the sensor response was investigated, and SPR responses were higher for bacteria in apple juice at pH 7.4 than in apple juice at pH 3.7. The LOD for S. choleraesuis was 4.4 × 104 cfu/mL in buffer and about 104 cfu/mL in apple juice with an adjusted pH of 7.4.234 Waswa et al. detected S. enterica serovar Enteritidis in milk using the commercial SPR sensor Biacore 2000.235 In that work, polyclonal S. enterica antibody was immobilized by first attaching protein A using a carboxymethylated dextran layer and amine coupling chemistry and subsequent attachment of the antibody to protein A.235 The LOD for Salmonella in pasteurized milk was determined to be 23 cfu/mL.235 In 2007 Mazumdar et al. also reported the detection of Salmonella in milk.240 They used the commercial SPR sensor Plasmonic and sandwich detection format. Polyclonal capture antibody was immobilized by self-assembly on the hydrophobic sensing surface formed by alkylsilanes. Milk spiked with S. typhimurium cells, killed by thimerosal (1%, w/w), was incubated with the sensing surface for 15 min and then switched with a solution containing the second antibody. The LOD for S. typhimurium cells in milk was at 105 cells/mL.240 Listeria monocytogenes was detected by Koubova´ et al.162 They used a custom-built SPR sensor with wavelength modulation and a double layer of antibodies adsorbed on a bare gold surface and cross-linked with glutaraldehyde. Heatkilled Listeria bacteria were detected at levels down to 107 cfu/mL.162 Leonard et al. used the commercial SPR sensor Biacore 3000 and competitive format to detect the L. monocytogenes.241 A polyclonal anti-goat antibody was immobilized in a carboxymethylated dextran layer using amine coupling chemistry. Solutions of known concentrations of L. monocytogenes were incubated with rabbit anti-Listeria antibodies. Cells and bound antibodies were then centrifuged out of solution, and the unbound antibodies remaining in solution were detected by the SPR sensor. The LOD was determined to be 105 cells/mL.241 Detection of L. monocytogenes in apple juice was reported by Taylor et al.234 They used a custom-built multichannel SPR sensor with wavelength modulation and sandwich detection format. Biotinylated polyclonal antibodies against L. monocytogenes were immobilized via streptavidin attached to a mixed SAM of oligo(ethylene glycol) alkanethiolate and biotinylated alkanethiolate. Detection of L. monocytogenes was performed in buffer, in a mixture of four bacterial species, and in apple juice. In the considered range of concentrations of L. monocytogenes in apple juice, the sensor response was higher for bacteria in apple juice with an adjusted pH of 7.4 than for those in buffer or natural apple juice with a pH of 3.7. The LOD for L. monocytogenes was determined to be about 3 × 103 cfu/mL for detection in both buffer and apple juice with a pH of 7.4.234 Campylobacter jejuni was detected by Taylor et al.234 They used a custom-built multichannel SPR sensor with wavelength modulation and sandwich detection format. Biotinylated polyclonal antibodies against C. jejuni were immobilized via streptavidin attached to a mixed SAM of oligo(ethylene glycol) alkanethiolate and biotinylated alkanethiolate. Detection of heat-killed C. jejuni was performed for buffer solutions containing only C. jejuni as well as a mixture of C. jejuni and other bacteria and apple juice spiked with C. jejuni. The LOD for C. jejuni was 1 × 105 cfu/mL in buffer

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and about 5 × 104 cfu/mL in apple juice.234 Staphylococcus aureus was detected by means of an SPR biosensor by Subramanian et al.242 They used the commercial SPR sensor SR 7000 and detected S. aureus directly or in sandwich detection format. Alkane monothiol and dithiol dendritic tether-based SAMs were examined for subsequent attachment of S. aureus antibodies using amine coupling chemistry. The LOD was determined to be 107 cfu/mL for direct detection and 105 cfu/mL for sandwich format for both sensing surfaces.242 Balasubramanian et al. reported an SPRbased detection of S. aureus using lytic phage as a biorecognition element.243 In that work a commercial SPR sensor, Spreeta, was used as a detection platform, and lytic phage was immobilized on the sensor surface by direct physical adsorption. The LOD for S. aureus in buffer was found to be 104 cfu/mL.243 Recently, Taylor et al. demonstrated the simultaneous detection of the four above-mentioned bacteriasE. coli O157:H7, C. jejuni, S. typhimurium, and L. monocytogeness on a custom-built multichannel SPR sensor.234 All bacteria were heat-killed and ultrasonicated prior to detection. Simultaneous detection of individual bacteria in the mixtures showed good agreement with detections of individual bacteria in buffer. Detections of individual bacteria and mixtures were also performed in apple juice samples. LODs for all four cases were established at 104, 5 × 104, 5 × 104, and 104 cfu/mL for E. coli O157:H7, C. jejuni, S. typhimurium, and L. monocytogenes, respectively.234 Yersinia enterocolitica was detected by Choi’s group,244 using the commercial SPR sensor Multiskop and monoclonal Y. enterocolitica antibodies immobilized via protein G attached to an alkanethiolate SAM on the sensor surface. The LOD for Y. enterocolitica in buffer was determined to be 102 cfu/mL.244 Vibrio cholerae O1 was detected by means of an SPR biosensor by Choi’s group.245 They used the commercial SPR sensor Multiskop and monoclonal antibodies immobilized via protein G attached to an alkanethiolate SAM on the sensor surface. The LOD for V. cholera O1 in buffer was determined to be about 4 × 105 cfu/mL.245 Protozoan parasite Cryptosporidium parVum oocyst, was detected by an SPR biosensor by Kang et al.246 In that work, the authors used the commercial SPR sensor platform Biacore 2000 and direct detection approach. Immobilization of biotinylated monoclonal antibody against C. parVum oocyst was performed on a mixed alkanethiolate SAM with attached streptavidin. The biosensor was able to detect C. parVum oocyst in buffer directly at a concentration of 106 oocysts/ mL. The authors also explored an alternative detection format consisting of the immunoreaction step between the biotinylated antibody and oocysts followed by the binding step of antibody-oocysts complex on the streptavidin-incubated surface. In this format, the LOD for C. parVum oocyst in buffer was established at 102 oocysts/mL.246 Detection of a fungal pathogen, Fusarium culmorum, in wheat using an SPR sensor was demonstrated by Zezza et al.247 Detection of F. culmorum was based on extraction of DNA from a sample, amplification of specific DNA fragment of F. culmorum, and subsequent detection of the amplicon using the SPR method via hybridization with complementary sequence immobilized on the SPR sensor (Biacore X) surface. The detection limit for the F. culmorum amplicon was 0.25 ng/µL. In 30 ng of durum wheat DNA, the smallest

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detectable amount of specific F. culmorum DNA was 0.06 pg.247

5.2.2. Toxins Toxins implicated in food safety include mainly toxins produced by bacteria, fungi, and algae. Staphylococcal enterotoxin B (SEB) was detected by an SPR sensor by Nedelkov et al. in 2000.248,249 They used the commercial SPR sensor Biacore X and SEB antibody immobilized in a carboxymethyldextran layer on the sensor surface via amine coupling chemistry. SEB was detected directly in milk and mushroom samples at levels down to 1 ng/mL. The SPR detection was followed by identification of the bound toxin by matrix-assisted laser desorption/ ionization time-of-flight (MALDI-TOF) mass spectrometry.248,249 In 2002 Homola’s group reported detection of SEB using a fiber optic SPR sensor.250 A double layer of antibodies was physisorbed on the surface of the SPR sensor and cross-linked with glutaraldehyde. SEB was detected directly, and the limit of detection for SEB in buffer was established at 10 ng/mL.250 Naimushin et al. detected SEB using a prototype of an SPR sensor developed by Texas Instruments Co. and antibodies immobilized on the sensor surface via a gold binding peptide.251 SEB was detected in buffer and seawater using direct detection or sandwich format with one or two amplification antibodies. The LOD for direct detection was 0.2 nM (5.6 ng/mL) in buffer and 1 nM (28 ng/mL) in seawater. Using a one-step amplification, concentrations of 20 pM (0.6 ng/mL) and 50 pM (1.4 ng/mL) were detected in buffer and seawater, respectively. The use of a second amplification antibody was shown to improve the LOD in buffer to 100 fM (2.8 pg/mL).251 Homola et al. reported the detection of SEB in buffer and milk.91 In that work, a custom-built wavelength modulation SPR sensor was employed, and polyclonal SEB antibody was immobilized on a mixed SAM of alkanethiolate using amine coupling chemistry. Detection of SEB was performed directly or using sandwich detection format. The LOD for direct detection of SEB in buffer was 5 ng/mL. Using a secondary antibody the LOD was improved to 0.5 ng/mL for both buffer and milk.91 In 2003 Medina reported the detection of SEB using an inhibition detection format.252 Sample containing SEB was incubated with a known concentration of SEB antibody for 20-30 min, and then the mixture was analyzed by the SPR sensor. The LOD for SEB in milk was established at 0.3 ng/mL. The sensing surface was demonstrated to be regenerable for repeated use by 100 mM hydrochloric acid.252 Detection of SEB in another food matrix was demonstrated by Medina, who detected SEB in ham tissue.253 In that work the commercial SPR sensor Biacore 1000 was used, and polyclonal antibody was immobilized in a carboxymethyldextran layer on the sensor surface via amine coupling chemistry. The LOD for sandwich detection format was determined to be 2.5 ng/mL in both buffer and ham tissue extract.253 Medina also demonstrated detection of Staphylococcal enterotoxin A (SEA) in raw eggs using the commercial SPR sensor Biacore 1000 and competitive detection format.254 SEA was immobilized in a carboxymethyldextran layer on the sensor surface via amine coupling chemistry. Homogenized raw egg samples were clarified by centrifugation. Anti-SEA was added to the sample, allowing SEA to bind with anti-SEA. The bound complex was separated from the free antibody by centrifugation. The supernatant was injected

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over the SEA-coated surface. Using this approach, SEA was detected in whole egg at concentrations down to 1 ng/mL.254 Domoic acid (DA) was detected using an SPR biosensor by Lotierzo et al.255 They used the commercial SPR sensor Biacore 3000 and a molecularly imprinted polymer photografted on a gold chip as a biorecognition element. Detection was performed in a competitive binding format in which free DA competed with its conjugate with horseradish peroxidase. The sensor was demonstrated to be able to detect DA in buffer at a concentration as low as 5 ng/mL.255 In 2005 Yu et al. demonstrated detection of DA using a custom-built SPR and inhibition detection format.256 DA was immobilized on a mixed SAM of OEG-containing alkanethiolates using amine coupling chemistry. The effect of regeneration and storage on the performance of the SPR sensor was investigated. The LOD of DA in buffer was established at 0.1 ng/ mL.256 Traynor et al. demonstrated SPR-based detection of DA in extracts of shellfish species.257 They used the commercial SPR sensor Biacore Q, and immobilization of DA in a carboxymethyldextran layer on the sensor surface was performed by amine coupling chemistry. Detection of DA was performed in inhibition format. Detection limits for DA in mussels, oysters, and cockles were determined to be about 1, 4.9, and 7 µg/g, respectively.257 Stevens et al. detected domoic acid in clam extracts using a portable SPR biosensor employing Spreeta 2000 modules and inhibition detection format.258 Polyclonal DA antibodies were immobilized on the sensor surface via a gold binding peptide. Detection was performed in buffer and in diluted clam extracts. The sensor was able to detect DA at 3 ng/mL in both buffer and diluted clam extracts.258 Aflatoxin B1 was detected using the commercial SPR sensor Biacore 1000 by Daly et al.259 In that work, aflatoxin B1 was conjugated to BSA and immobilized on carboxymethylated dextran using amine coupling chemistry. Detection was performed using inhibition detection format. Detection of aflatoxin B1 in buffer was demonstrated at levels down to 3 ng/mL.259 Dunne et al. demonstrated an SPR sensor for aflatoxin B1 using scFvs as a biorecognition element.148 Detection was performed in the commercial SPR sensor Biacore 3000 using inhibition detection format. Aflatoxin B1 derivative was immobilized on the carboxymethylated dextran layer on the sensor surface using amine coupling chemistry. Regeneration protocol was developed enabling at least 75 detection/regeneration cycles. The LOD for aflatoxin B1 in buffer was 375 pg/mL for monomeric scFv and 190 pg/mL for dimeric scFv.148 Detection of deoxynivalenol in buffer and wheat was demonstrated by Tu¨do¨s et al.260 In that work, the commercial SPR system Biacore Q system was used and detection of deoxynivalenol was performed using the inhibition detection format. Deoxynivalenol conjugated to casein was immobilized on a carboxymethylated dextran layer on the sensor surface using amine coupling chemistry. The sensor was demonstrated to detect deoxynivalenol in buffer down to 2.5 ng/mL and showed good agreement with liquid chromatography-tandem mass spectrometry measurements on wheat samples.260

5.2.3. Veterinary Drugs Another important field in which SPR biosensor technology has been increasingly applied is testing for veterinary drug residues (e.g., antibiotics, β-agonists, and antiparasitic drugs) in food.261

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An SPR sensor for the detection of antibiotics (penicillins and cephalosporins) in milk was demonstrated by Cacciatore et al.262 Their approach was based on the noncompetitive inhibition of the binding of digoxigenin-labeled ampicillin (DIG-AMPI) to a soluble penicillin-binding protein 2× derivative (PBP 2×*) of Streptococcus pneumoniae by other β-lactam antibiotics. Subsequently, the DIG-AMPI/PBP 2×* complex was detected using the commercial SPR platform Biacore 3000 and digoxigenin antibody immobilized on the sensor chip. The LODs for the selected antibiotics in raw milk were established to be below 2 ng/mL for benzylpenicillin, ampicillin, and amoxicillin, 15 ng/mL for cloxacillin, 50 ng/mL for cephalexin, and 25 ng/mL for cefoperazone.262 Chloramphenicol and chloramphenicol glucuronide residues in various food matrices were detected by Ashwin et al.263 They used the commercial SPR platform Biacore Q and direct detection format. They detected chloramphenicol in extracts from honey, prawns, and dairy products and chloramphenicol glucuronide in extracts of porcine kidney at concentrations below 0.2 µg/kg.263 Detection of chloramphenicol and chloramphenicol glucuronide using an SPR sensor and inhibition assay was performed by Ferguson et al.264 They used the commercial SPR sensor Biacore Q and a chip with immobilized chloramphenicol derivative (Qflex Kit Chloramphenicol, Biacore). A known concentration of drug-specific antibody was mixed with the sample and injected over the surface of a sensor chip on which a chloramphenicol derivative was immobilized. Chloramphenicol and chloramphenicol glucuronide in extracts from food matrices were detected at levels down to 0.005 µg/kg (poultry), 0.02 µg/kg (honey), 0.04 µg/kg (prawn), and 0.04 µg/kg (milk).264 Dumont et al. demonstrated an SPR sensor for the detection of fenicol antibiotic residues in shrimps.265 They used the commercial SPR sensor Biacore Q and inhibition detection format. Analyte molecules were immobilized on carboxymethylated dextran using amine coupling chemistry. Chloramphenicol, florefenicol, florefenicol amine, and thiamphenicol were detected in extract from shrimp at levels down to 1, 0.2, 250, and 0.5 ng/mL, respectively.265 In 2007 Moeller et al. reported an SPR biosensor for the indirect detection of tetracycline in honey and milk.266 Their approach was based on the resistance mechanism against tetracycline in Gram-negative bacterias tetracyclines release Tet repressor protein (TetR) from the tet operator (tetO). Biotinylated single-strain DNA containing the sequence of the tetracycline operator tetO1 was immobilized on a streptavidin-coated sensor chip. The repressor protein TetR was attached to the chip-bound operator tetO. Injection of a solution containing tetracycline allowed tetracycline to bind to TetR. This resulted in the release of a conformationally changed protein, which was continuously flushed away from the sensor surface. The decrease in surface density was measured using the commercial SPR sensor Biacore 3000. The LOD was estimated to be 1 ng/mL for tetracycline in buffer and 15 ng/mL and 25 µg/kg for tetracycline in raw milk and honey, respectively.266 Detection of the antibiotic tylosin was demonstrated by Caldow et al.267 They used the commercially available SPR sensor Biacore Q and inhibition detection format. Immobilization of tylosin on a carboxymethyldextran layer on the sensor surface was performed by amine coupling chemistry. Tylosin was detected in extract from honey at levels down to 2.5 µg/ kg.267

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5.2.4. Vitamins Vitamin B2 (riboflavin) was detected by Caelen et al. using the commercial SPR sensor platform Biacore Q and inhibition detection format.268 A riboflavin derivative was immobilized on the carboxymethylated dextran using amine coupling chemistry. A known concentration of riboflavin binding protein was mixed with a sample, and the amount of unreacted protein was measured using the SPR sensor. Riboflavin was detected in milk-based products, and the LOD was established at 70 ng/mL.268 Haughey et al. reported an SPR sensor for the detection of vitamin B5 (pantothenic acid).269 They used the commercial SPR sensor Biacore Q and inhibition detection format. Vitamin B5 derivative was immobilized in carboxymethylated dextran using amine coupling chemistry. Detection of vitamin B5 was performed in extracts from various foods (e.g., infant formula, cereal, pet food, egg powder). The LOD was 4.4 ng/mL.269 An SPR sensor for the detection of vitamin B8 (biotin) and vitamin B9 (folic acid) was demonstrated by Indyk et al.270 They used the commercial SPR sensor Biacore Q, a biotin sensor chip, and inhibition detection format. In their experiments, they detected biotin and folic acid in infant formulas and milk powders at concentrations as low as 2 ng/mL.270 SPRbased detection of vitamin B12 (cobalamine) was demonstrated by Indyk et al.271 They used the commercial SPR sensor Biacore Q and inhibition detection format. Vitamin B12 was immobilized in carboxymethylated dextran layer on the sensor surface using amine coupling chemistry. The LOD for cobalamine in milk or infant formula or an extract from beef was determined to be 0.06 ng/mL.271

5.2.5. Hormones The steroid hormone progesterone in cow’s milk was detected by Gillis et al.272,273 They used the commercial SPR sensor platform Biacore 2000 and inhibition assay. Progesterone derivative was immobilized in the carboxymethylated dextran layer on the sensor surface using amine coupling chemistry. A known concentration of monoclonal antibody was incubated with sample (buffer or cow’s milk), and the amount of unreacted antibody was detected by the SPR sensor. In their earlier work in 2002, Gillis et al. established the LOD for progesterone in raw bovine milk at 3.6 ng/mL.272 Optimization of the assay reported in 2006 allowed Gillis et al. to detect progesterone in buffer and bovine milk down to 60 pg/mL and 0.6 ng/mL, respectively.273 Mitchell et al. used the commercial SPR sensor Biacore 2000 and inhibition detection format combined with gold nanoparticles and proteins to improve the sensitivity of the detection.212 Progesterone was immobilized to a dextran layer through covalent immobilization using an OEG linker attached to the progesterone. Detection formats employing gold particles conjugated with streptavidin and attached to biotinylated monoclonal antibody in both label prebinding and sequential binding formats were explored. Prelabeling format allowed detection of progesterone down to 143 pg/mL, and sequential binding formats yielded a LOD of 23.1 pg/mL. Secondary antibody labeling produced an 8-fold signal enhancement and a LOD of 20.1 pg/mL, whereas the use of secondary antibody conjugated with a gold nanoparticle improved the LOD to 8.6 pg/mL.212

5.2.6. Diagnostic Antibodies Detection of Mycoplasma hyopneumoniae antibody in pig serum using the SPR method was demonstrated by Kim et

SPR Sensors

al.274 They used the commercial Autolab Esprit SPR system and a recombinant 30 kDa fragment of P97 adhesin as an antigen. The performance of the SPR biosensor was compared with that of enzyme-linked immunosorbent assay (ELISA) using 70 pig serum samples. There was found to be a strong positive correlation between these two methods and, in terms of the LOD, the SPR sensor outperformed ELISA by a factor of 10.274 Classical swine fever virus (CSFV) antibody in pig serum was detected by Cho and Park.275 They used the commercial Autolab Esprit SPR system and recombinant gp55 protein as an antigen. They used the same immobilization method and methodology as in their previous study.274 One hundred and seventy pig serum samples were analyzed by the SPR sensor and ELISA. It was determined that the LOD of the SPR sensor was better by a factor of 10 than that of ELISA.275

5.2.7. Allergens Direct detection of peanut allergens by means of SPR sensor was demonstrated by Mohammed et al.276 They used a miniature commercial SPR sensor Spreeta and peanutspecific antibodies adsorbed on the sensor surface. The LOD for the peanut allergen in buffer was estimated to be 700 ng/mL.276 Food allergens were also detected by Malmheden Yman et al.277 They used the commercial Biacore Q SPR instrument and affinity-purified antibodies raised against egg white, protein conalbumin, sesame seed protein, peanut protein, hazelnut protein (corylin), and crab meat, which were immobilized on the carboxymethylated dextran by the amine coupling method. Detection of allergens was performed directly and by sandwich detection format. The second antibody in the sandwich assay was demonstrated to improve both the sensitivity and specificity of the detection. Peanut proteins in chocolate diluted only 10 times before the analysis were detected down to 1 µg/g. Conalbumin in pasta was detected at levels as low as 0.3 µg/g. Sesame seed protein was detected down to 0.125 µg/mL, corresponding to 12.5 µg/g in solid food (e.g., bread). Tropomyosin in pasta was detected at the level of 10 µg/g.277

5.2.8. Proteins Simultaneous detection of three caseins in milk using an SPR method was demonstrated by Dupont and MullerRenaud.278 They used the commercial SPR sensor platform Biacore 3000 and sandwich assay format employing two monoclonal antibodies directed against the N- and C-terminal extremities of each of the caseins, respectively. Antibody against C-terminal extremities of each of the caseins was immobilized in a separate sensing channel of the four-channel SPR sensor via carboxymethylated dextran matrix and amine coupling. Three major caseins (Rs1, β, and κ) were detected in milk samples. The LODs were estimated at 85 ng/mL for β-casein, 870 ng/mL for Rs1-casein, and 470 ng/mL for κ-casein.278 Indyk et al. detected proteins such as immunoglobulin G, folate-binding protein, lactoferrin, and lactoperoxidase in bovine milk using the commercial SPR biosensor Biacore Q.279 Respective monoclonal antibodies were immobilized on the dextran matrix using amine coupling chemistry. The detection was performed directly in milk samples diluted in buffer. The LODs were established at 16.8 ng/mL for immunoglobulin G, 0.7 ng/mL for folate-binding protein, 1.1 ng/mL for lactoferrin, and 75 ng/mL for lactoperoxidase.279

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5.2.9. Chemical Contaminants Detection of 4-nonylphenol in shellfish using an SPR biosensor was demonstrated by Samsonova et al.280 They used the commercial SPR platform Biacore Q and inhibition detection format. 9-(p-Hydroxyphenyl)nonanoic acid was immobilized on dextran matrix using amine coupling chemistry. Using monoclonal antibodies, a detection limit of 2 ng/mL in buffer was achieved. The detection was performed in <3 min including a 30 s regeneration step. The sensor was regenerated by 100 mM sodium hydroxide in 10% acetonitrile. In shellfish samples, 4-nonylphenol was detected at concentrations down 10 ng/g.280 A suspected carcinogen, insulin-like growth factor-1 (IGF-1), can occur in milk of cows treated by recombinant bovine somatotropin treatment. Guidi et al. detected IGF-1 using the commercial Biacore SPR sensor platform and inhibition detection format.281 Recombinant IGF-1 was immobilized to a carboxymethylated dextran matrix via amine coupling chemistry. Polyclonal antibody was incubated with a sample for 2 h, and then the sample was injected in the flow cell of the SPR sensor. On the basis of the reported data, the LOD for IGF-1 in buffer and milk can be estimated to be below 10 ng/mL.281

5.3. Medical Diagnostics Fast, sensitive, and specific detection of molecular biomarkers indicating normal biologic processes, pathogenic processes, or pharmacologic responses to a therapeutic intervention presents an important goal for modern bioanalytics.282,283 SPR biosensors have been demonstrated to hold promise for the detection of analytes related to medical diagnostics such as cancer markers, allergy markers, heart attack markers, antibodies, drugs, and hormones.

5.3.1. Cancer Markers The prostate-specific antigen (PSA) is a marker for prostate cancer.284 Detection of PSA in PBS buffer using the commercial SPR sensor Ibis II has been reported by Besselink et al.285 In their work, monoclonal antibodies against PSA were immobilized on the sensor surface via amine coupling chemistry. After incubation of the sensing surface with sample containing PSA, the sensor response was amplified with rabbit anti-PSA polyclonal antibodies followed with either biotinylated goat anti-rabbit IgG and streptavidin-coated latex microspheres or goat anti-rabbit IgG-coated colloidal gold. The detection format employing gold particle enhancement provided a LOD as low as 0.15 ng/mL. Huang et al. investigated the detection of PSA using direct and sandwich detection formats and the commercial SPR sensor Biacore 2000.286 PSA-receptor molecules consisting of a single-domain antigen-binding fragment were covalently immobilized on the sensor surface via a mixed alkanethiolate SAM. PSA concentrations as low as 10 ng/ mL were detected in buffer. Sandwich detection format involving a biotinylated secondary antibody and streptavidinmodified gold nanoparticles improved the LOD for PSA below 1 ng/mL. Recently, the determination of a complex of PSA with R1-antichymotrypsin (PSA-ACT) in both HBS buffer and human serum was demonstrated by Cao et al.201 using the commercial SPR sensor Biacore 2000. Mixed alkanethiolates were optimized to provide a stable surface for sequential attachment of biotin, streptavidin, and biotinylated antibodies against PSA-ACT. The PSA-ACT complex in HBS buffer and human serum was detected

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directly at concentrations as low as 20.7 and 47.5 ng/mL, respectively. The LOD for the PSA-ACT complex was improved by employing sandwich detection format and PSA polyclonal antibody to 10.2 and 18.1 ng/mL for detection in the HBS buffer and serum, respectively.201 The quantitation of a pancreatic cancer marker, carbohydrate antigen (CA 19-9) was performed by Chung et al. using a miniature commercial SPR sensor Spreeta.287 The antibody against CA 19-9 was immobilized on the sensor surface via a SAM of alkanethiolates. The sensor was shown to be able to detect CA 19-9 at a concentration of 410.9 U/mL directly and 66.7 U/mL using a sandwich assay.287 Protein vascular endothelial growth factor (VEGF), which plays a role in breast cancer, lung cancer, and colorectal cancer, was detected using an SPR imaging method and RNA aptamer microarray. The adsorption of proteins onto the RNA microarray was detected by the formation of a surface aptamer-protein-antibody complex. The sensor response was amplified using a localized precipitation reaction catalyzed by the enzyme horseradish peroxidase conjugated to the antibody. The sensor was demonstrated to be able to detect VEGF at a concentration of 1 pM.220 Yang et al.288 measured levels of interleukin-8 (IL-8) protein in the saliva of healthy individuals and patients with oropharyngeal squamous cell carcinoma using the commercial SPR sensor Biacore X. The sandwich detection format using two monoclonal antibodies recognizing different epitopes on the IL-8 was used. A monoclonal antibody against IL-8 was immobilized in the dextran layer via amine coupling chemistry. Saliva samples were first centrifuged to clarify the supernatants. The supernatants were then aspirated and separated from the cellular pellet. The detection limit for IL-8 was determined to be 2.5 pM (∼0.02 ng/mL) for detection in buffer and 184 pM (∼1.5 ng/mL) for detection in saliva samples. Carcinoembryonic antigen (CEA), a marker related to colorectal cancer, was detected by the SPR method of Tang et al.289 They used the commercial SPR sensor Autolab Springle and protein A adsorbed on the SPR sensor surface for subsequent attachment of carcinoembryonic antibody. The LOD of 0.5 ng/mL was achieved, and the sensor was demonstrated to be regenerable for repeated use.289 Tian et al. demonstrated an SPR biosensor for the detection of fibronectin, a glycoprotein implicated in carcinoma development.212 In that work, a research SPR instrument employing an acousto-optic tunable filter and wavelength modulation was used. Fibronectin antibody was immobilized on the self-assembled alkanethiolate monolayer using amine coupling chemistry. Fibronectin was detected directly, in sandwich detection format, and with additional amplification using colloidal gold. Regenerability of the sensor was demonstrated. The LODs for fibronectin in buffer were established at 2.5, 0.5, and 0.25 µg/mL for direct detection, sandwich detection, and colloidal Au-enhanced detection, respectively.212

5.3.2. Antibodies against Viral Pathogens SPR biosensors for the detection of hepatitis virus specific antibodies were reported by several research groups. In 2003 Rojo et al. described the SPR biosensor-based detection of antibodies against hepatitis G in serum.150 They used the commercial SPR sensor Biacore 1000 and synthetic peptide as a biorecognition element, which was immobilized in the dextran layer on the sensor surface via amine coupling

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chemistry. Threshold measurements on sera of chronic hepatitis C patients as well as control samples from healthy patients obtained with the sensor were consistent with those obtained by ELISA. Chung et al. used the commercial SPR sensor Spreeta for the detection of antibodies against human hepatitis B virus (hHBV). The hHBV antigen was immobilized on the SAM of alkanethiolates via amine coupling chemistry. Antibodies against hHBV were detected in 5% serum in PBS. The LOD for direct detection was established at 9.2 nM. Amplification methods based on sandwich detection and avidin-biotinylated antibodies were shown to yield amplification factors of 7 and 14, respectively. Using peroxidase-antiperoxidase complex, 17-fold amplification of the sensor response was obtained and the LOD was lowered to 0.64 nM.202 Wittekindt et al. demonstrated an SPR sensor for the detection of antibodies against herpes simplex virus type 1 and type 2 (HSV-1, HSV-2) in human sera.151 They used the commercial SPR sensor Biacore X and SPR chips coated with streptavidin on which two biotinylated peptides, used as biorecognition elements, were immobilized. Human serum samples (diluted 1:100 in HBS buffer) were tested using the SPR biosensor and immunoblotting (reference method). A good agreement between the SPR biosensor and immunoblotting was obtained (correlation of 83 and 86% for antibodies against HSV-1 and HSV-2, respectively).151 Direct detection of antibodies against Epstein-Barr virus (anti-EBNA) in 1% human serum was reported by Homola’s group.152 A short synthetic peptide was used as biorecognition element and was immobilized on the surface of a wavelengthmodulated SPR biosensor via hydrophobic and electrostatic interactions. The LOD for anti-EBNA was determined to be 0.2 ng/mL. A procedure for the regeneration of the sensor was developed and demonstrated to allow at least 10 repeated anti-EBNA detection experiments without a significant loss in sensor sensitivity.152 Regnault et al. reported the SPR biosensor-based detection of anti-protein S antibodies following Varicella-Zolter virus infection.290 In that work, protein S was immobilized in the layer of dextran via amine coupling chemistry and the experiments were performed using the commercial SPR sensor Biacore X. A high sensor response was observed to diluted plasma (1:5) of an infected patient, whereas samples from healthy patients generated a minimum response.290 Antibodies against human respiratory syncytial virus (RSV) in sera were detected by the commercial SPR biosensor Biacore 2000 by McGill et al.291 Monoclonal antibodies against the virus glycoproteins (F- and G-glycoproteins) were covalently attached to the dextran matrix via amine coupling chemistry and then used to immobilize the respective virus glycoproteins. Serum samples isolated from patients’ respiratory tracts were diluted in HBS buffer (1:10) and filtered. The SPR biosensor was demonstrated to be able to recognize the antigenic differences between the two different genotypes of the virus (G- and F-virus glycoproteins).291 Abad et al. demonstrated SPR-based detection of isotypespecific anti-adenoviral antibodies in patients dosed with an adenoviral-based gene therapy vector. The antibodies were detected using the commercial SPR instrument Biacore 3000 and intact virus immobilized in the thin layer of dextran on the surface of the sensor by amine coupling chemistry.292 Patient serum samples or ascites fluid samples were diluted 1:10 with HEPES prior to the analysis.

SPR Sensors

5.3.3. Drugs and Drug-Induced Antibodies Dillon et al. demonstrated the SPR biosensor-based detection of morphine-3-glucuronide (M3G), the main metabolite of heroin and morphine.293 They used the commercial SPR sensor Biacore 1000 and inhibition detection format. M3G-ovalbumin conjugate was immobilized on the dextran matrix via amine coupling chemistry. Two polyclonal antibodies were produced, purified, and tested for the detection of M3G. Regeneration protocols were developed for both polyclonal antibodies and allowed for approximately 60 cycles for the first antibody and 50 cycles for the second antibody. The LOD for M3G in buffer and in urine (diluted 1:250) was found to be <1 ng/mL for both antibodies.293 Detection of the oral anticoagulant warfarin by the SPR method was performed by Fitzpatrick and O’Kennedy.294 They used the commercial SPR sensor Biacore 3000 and inhibition detection format. 4′-Aminowarfarin or 4′-azowarfarin-BSA was immobilized on a dextran matrix via amine coupling chemistry. Detection of warfarin was performed in plasma ultrafiltrate (diluted 1:100). The sensor was demonstrated to detect warfarin at concentrations down to 4 ng/mL and to be regenerable for more than 70 detection cycles.294 Gobi et al. reported SPR sensor-based detection of insulin.203 In that work the commercial SPR sensor SPR670 and inhibition format were employed. Insulin was covalently bound to the activated monolayer of heterobifunctional OEG-dithiocarboxylic acid derivative. After 5 min of incubation of sample with a known concentration of monoclonal anti-insulin antibody, the mixture was injected in the SPR sensor and the concentration of the unreacted antibody was measured. A regeneration protocol was developed, and the chip was shown to be reusable for more than 25 detection cycles without an appreciable change in the sensor activity. The LOD for insulin in buffer was 1 ng/mL. The lowest detectable concentration of insulin in the serum samples spiked with insulin was 6 ng/mL.203 Rini et al. reported an SPR sensor for threshold measurement of antibodies against granulocyte macrophage colony stimulating factor (GM-CSF) used in therapies for various kinds of cancer.295 Rini et al. used the commercial SPR instrument Biacore 2000 and SPR chips with GM-CSF antigen immobilized on the carboxymethylated dextran via amine coupling chemistry. Antibodies against GM-CSF were induced in prostate cancer patients by repeated administration of GM-CSF, and their presence was measured in diluted sera (1:5). The SPR measurements showed the presence of GMCSF reactive antibodies for all prostate cancer patients treated with GM-CSF, which was in agreement with reference ELISA measurements.295 An SPR biosensor for the detection of antibody against insulin was demonstrated by Kure at al.296 Insulin antibodies can cause insulin resistance or hypoglycemia in diabetic patients treated with human insulin. In that work the commercial SPR biosensor Biacore 2000 was employed and purified human insulin, as a biorecognition element, was immobilized on the sensor surface via amine coupling chemistry. In calibration experiment in buffer, monoclonal antibodies against insulin were detected at concentrations as low as 0.6 µg/mL. Serum samples were pretreated to remove insulin and filtered before SPR measurements. Insulin antibodies were detected in eight diabetic patient serum samples and determined to fall in the range of 2.91-16.3

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µg/mL. No insulin antibodies were detected in the control group.296

5.3.4. Hormones A marker of pregnancyshuman chorionic gonadotropin hormone (hCG)shas been a frequent target of optical biosensor technologies. Jiang’s group reported SPR sensorbased detection of hCG exploiting a wavelength-modulated SPR sensor and DNA-directed antibody immobilization method.165 The immobilization consisted of non-covalent attachment of streptavidin to a biotinylated SAM of alkanethiolates followed with the binding of biotinylated oligonucleotides to available streptavidin binding sites. Antibodies chemically modified with oligonucleotides with a complementary sequence were finally attached to this surface via DNA hybridization. The detection limit for direct detection of hCG in buffer was determined to be 0.5 ng/ mL. Recently, detection of hCG in urine was performed using a sequential detection method developed by Chung et al.208 They used the commercial SPR sensor Spreeta and immobilized two molecular recognition elements (antibody against HCG and antibody against human albumin) in a single sensing channel of Spreeta sensor using SAM of alkanethiolates and amine coupling chemistry. Amplification polyclonal anti-hCG antibodies were used to increase the sensor response. In 10-fold diluted urine, the detection limit for hCG was established to be 46 mIU/mL.208 SPR-based detection of 17β-estradiol was demonstrated by Miyashita et al.297 They used the commercial SPR sensor Biacore X and inhibition detection format. Estradiol-BSA conjugate was immobilized on carboxymethylated dextran layer by amine coupling chemistry. The binding of unreacted antibody to 17β-estradiol conjugates at the surface of the sensor was measured. The 17β-estradiol was detected down to 0.47 nM (∼0.14 ng/mL).297 Teramura and Iwata demonstrated an SPR sensor for the detection of R-fetoprotein (AFP) in human plasma.298 They used a research SPR sensor system with angular modulation and a SAM of tri(ethylene glycol), and carboxyl groupterminated hexa(ethylene glycol) was employed for covalent attachment of monoclonal AFP-antibody. Detection was performed in sandwich detection format. The SPR signal shift was further enhanced by applying a polyclonal antibody against the second antibody. The polyclonal antibody against the second antibody was demonstrated to amplify the sensor response to AFP by a factor of 7. The LOD for AFP in blood plasma was estimated to be in nanograms per milliliter levels.298

5.3.5. Allergy Markers Measurement of immunoglobulin E (IgE) antibody levels plays an important role in the diagnostics of allergies. Imato’s group reported direct detection of IgE antibody using an SPR biosensor.199 In that work, anti-IgE antibody was immobilized on the surface of the commercial SPR sensor SPR 20 by physical adsorption. A sample containing IgE antibody was mixed with an anti-IgE(H) antibody solution to form an antiIgE(H) complex via the Ce2 domain of the IgE antibody. The solution was introduced in the SPR sensor and the immunocomplex of the IgE-anti-IgE(H) reacted with the anti-IgE(D) antibody immobilized on the sensor chip via the Ce3 domain of the IgE antibody. The LOD for the IgE antibody was about 10 ppb.199

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The same group reported SPR-based detection of histamine (β-imidazole ethylamine)sa protein involved in allergic reactions. Their SPR sensor was based on the commercial SPR-20 sensor and inhibition detection format. Histamine was immobilized on the sensor surface using a self-assembled alkanethiolate monolayer and amine coupling chemistry. A regeneration protocol was developed, and it was demonstrated that the sensor can be used for more than 10 detection/ regeneration cycles. The limit of detection was 3 ppb.299

Hwang et al. reported an SPR sensor for the detection of hepatitis B surface antigen (HBsAg), an early indicator of hepatitis B.307 In their work, they used the commercial SPR sensor Biacore 3000 and immobilized anti-HBsAg polyclonal antibody to the dextran layer on a sensor chip surface using amine coupling chemistry. HBsAg was detected at concentrations as low as ∼1 µg/mL.307

5.3.6. Heart Attack Markers

Analytes of environmental concern225,308 targeted by SPR biosensors include, in particular, pesticides, aromatic hydrocarbons, heavy metals, phenol, polychlorinated biphenyls, and dioxins.

Detection of a marker of cardiac muscle injury, troponin (cTn I), in serum was demonstrated by Wei et al.300 In that work, biotinylated antibodies against cTn I were immobilized on the avidin layer created using amine coupling chemistry on an activated SAM of alkanethiolates. cTn I was detected directly and in sandwich detection format. The LODs were determined to be 2.5 and 0.25 ng/mL for direct and sandwich detection format, respectively.300 Detection of cTn I was also demonstrated by Booksh’s group.301 They used a miniature fiber optic SPR sensor on which human anti-cardiac troponin I was immobilized via a dextran layer and amine coupling chemistry. The LOD for cTnI in buffer was established at 3 ng/mL.301

5.3.7. Other Molecular Biomarkers Detection of antibodies against glucose 6-phosphate isomerase (GPI) in synovial fluids of rheumatoid arthritis and osteoarthritis patients (diluted 1:100 in Hepes) using a Biacore 2000 is presented in the work of Kim et al.302 Recombinant human GPI proteins produced from E. coli were immobilized on the dextran sensor surface via amine coupling chemistry. The synovial fluid samples from rheumatoid arthritis patients showed a significantly higher level of binding to the recombinant GPI proteins than samples from osteoarthritis patients. SPR-based detection of anti-glutamic acid decarboxylase (GAD) antibodies for diagnosing type I diabetes mellitus was reported by Sim’s group.303,304 They used the commercial SPR sensor Biacore 2000 and biotinylated GAD coupled to streptavidin molecules anchored covalently on a mixed SAM of hydroxyl- and carboxyl-terminated alkanethiolates. Optimization of SAM composition was carried out. The SPR sensor employing an optimized SAM was shown to be able to detect antibody in HBS-EP buffer in sub-micromolar levels.304 Detection of c-reactive protein (CRP), a human blood serum marker for inflammatory processes, using SPR biosensor technology was demonstrated by Meyer.305 They used the commercial SPR sensor Plasmonic and sandwich detection format. Biotinylated monoclonal antiCRP antibody C6 was immobilized on the biotin-coated sensor surface via streptavidin. Buffer spiked with CRP was injected in the SPR sensor cuvette and then replaced by a solution containing the secondary antiCRP antibody C2. The sandwich assay was completed typically within 30-60 min. The LOD for CRP in buffer was determined to be 1 µg/mL.305 SPR-based detection of the cystatin C marker of the glomerular filtration rate (GFR), a critical measure of normal kidney function, was demonstrated by Corn’s group.306 They used an SPR imaging instrument with antibodies immobilized on the alkanethiolate-modified sensor surface using carbonyldiimidazole surface reaction. The sensor was shown to be able to detect cystatin C at 1 nM levels.306

5.4. Environmental Monitoring

5.4.1. Pesticides Following the demonstration of SPR biosensors for the detection of atrazine309 and simazine137,138 in the 1990s, various pesticides have been targeted by SPR sensor technology. Lechuga’s group reported a portable SPR sensor for the detection of atrazine in water using inhibition format.310 An atrazine derivative was immobilized on the alkanethiolate SAM formed on the gold-coated sensor surface. The sample was incubated with polyclonal antibodies, and then the mixture was analyzed by the SPR sensor. The LOD was established at 20 pg/mL. A measurement/regeneration cycle required about 25 min.310 An alternative approach to the detection of atrazine based on specifically expressed mRNA in Saccharomyces cereVisiae bacteria exposed to atrazine was reported by Lim et al.311 The cells were brought into contact with the analyzed sample and disrupted, and the amount of expressed P450 mRNA was measured by an SPR biosensor with complementary oligonucleotide probes immobilized on the sensor surface using streptavidin-biotin chemistry. The LOD (1 pg/mL) presents a substantial improvement compared to previous works.18,309 Lechuga’s group demonstrated SPR biosensors for the detection of organophosphate pesticide chlorpyrifos and carbaryl.312-314 They used inhibition detection format in which a pesticide derivative with BSA was covalently immobilized on a self-assembled alkanethiolate monolayer formed on the SPR sensor surface. Typical limits of detection were around 50 pg/mL for chlorpyrifos312,314 and 1 ng/mL for carbaryl.312,313 A protocol for regeneration of the sensing surface was developed, and the sensor was able to perform ∼200 detection cycles without degradation in performance.312 The detection cycle was completed in 20 min. The sensors were tested in ground, river, and drinking water samples without the observation of significant matrix effects. Lechuga’s group also demonstrated an SPR sensor for the detection of dichlorodiphenyltrichloroethane (DDT).315 They used a portable SPR sensor now commercially available from Sensia (Spain) and inhibition detection format. DDT derivative was immobilized on a self-assembled alkanethiolate monolayer on the SPR sensor surface. Two monoclonal antibodies, specific to DDT and specific to DDT and its metabolites, were used in the inhibition detection format. Regeneration of the sensor surface was developed, and 270 detection cycles were performed. The LOD in distilled water was established at 15 pg/mL for the DDT-specific assay and at 31 pg/mL for the DDT group-selective assay. Nearly the same performance was achieved when the sensor was used to analyze river water samples.315 Detection of three pesticidessDDT, carbaryl, and chlorpyrifossusing a single

SPR Sensors

channel of an SPR sensor and inhibition detection format was demonstrated by Lechuga’s group.209 DDT, carbaryl, and chlorpyrifos derivatives were attached to carboxylic terminal groups on a SAM of alkanethiolates. A sample was mixed with antibody against one target and injected in the flow cell of the sensor. After the sensor response had been established, a sample mixed with antibody against another target was injected with or without previous regeneration of the sensing surface. The LODs for this multianalyte detection approach were established at 18 pg/mL for DDT, 50 pg/mL for carbaryl, and 52 pg/mL for chlorpyrifos. These detection limits were comparable with those obtained using singleanalyte functionalizations.209 Miura’s group applied the SPR method to the detection of 2,4-dichlorophenoxyacetic acid (2,4-D).316,317 Initially, they used inhibition detection format and a conjugate of 2,4-D derivative and BSA (2,4-D-BSA) immobilized on the sensor by physisorption. The LOD was established at 0.5 ng/mL.316 In their later study, they used a SAM of alkanethiolates for covalent attachment of 2,4-D-BSA conjugate on the sensor surface. This functionalization approach led to an improvement in the LOD to 10 ppt.317 Regeneration of the sensor for up to 30 detection cycles was demonstrated using pepsin.317

5.4.2. 2,4,6-Trinitrotoluene (TNT) SPR biosensors for the detection of TNT, which is a prime constituent of most of landmines and also exhibits toxic, mutagenic, and carcinogenic effects, have been extensively researched by Miura’s group. They used inhibition detection format involving various conjugates and antibodies. The conjugates used in their experiments included 2,4,6-trinitrophenol-BSA (TNP-BSA).318,319 TNP-ovalbumin (TNPOVA),318 2,4,6-trinitrophenyl-keyhole limpet hemocyanin (TNP-KLH),320 and TNP-β-Alanine-ovalbumin (TNPβ-Ala-OVA).321 The antibodies (monoclonal and polyclonal) were homemade321 or commercial. Most of their sensors delivered a LOD below 10 pg/mL (10 ppt);318,320-322 the best LOD (1 ppt) was achieved using the immunoreaction between a homemade polyclonal anti-2,4,6-trinitrophenylkeyhole limpet hemocyanine antibody with a physically immobilized TNP-OVA.318 The sensor surface was regenerated by pepsin. The detection cycle was typically completed in <20 min. The inhibition-based SPR biosensors of TNT were found to exhibit low cross-sensitivity to other similar compounds such as 1,4-DNT, 1,3-DNB, 2A-4,6-DNT, and 4A-4,6-DNT.323 Larsson et al. studied the effect of composition of molecular coatings on the performance of the SPR biosensor for the detection of TNT.324 They used the commercial SPR platform Biacore2000 and inhibition detection format. Two types of thiols, OEG-alkylthiols terminated with a hydroxyl group and a TNT analogue (2,4-dinitrobenzene), were self-assembled on the surface of an SPR chip. The ratio of TNT analogues and hydroxyl-terminated OEGthiols was optimized to provide highly selective and sensitive biochips with minimum nonspecific binding. The LOD for TNT in buffer was demonstrated to be <10 pg/mL.324

5.4.3. Aromatic Hydrocarbons An SPR biosensor for the detection of 2-hydroxybiphenyl (HBP, a metabolite of BaP) was demonstrated by Miura’s group. The sensor was based on the inhibition immunoassay format. An antibody against HBP was mixed with a sample and, after incubation, the unreacted antibody was detected

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using an SPR biosensor functionalized with an HBP conjugate. HBP-BSA was immobilized on the surface of the SPR sensor by physical adsorption.325,326 The LOD using the commercial SPR biosensor SPR-20 was demonstrated to be 0.1 ppb (ng/mL). Most recently, this detection approach was combined with a portable SPR sensor, and the LOD of 0.1 ppb (ng/mL) was reproduced also on this sensor platform. A simple regeneration procedure using a pepsin solution allowed more than 30 measurement cycles without appreciable deterioration of sensor response. The measurement cycle was completed in 20 min. The same group also demonstrated detection of benzo[a]pyrene (BaP). They used the commercial SPR biosensor SPR-20 and inhibition assay. BAP molecules were attached to the sensor surface either by immobilizing BaP-BSA conjugate on the sensor via physical adsorption327,328 or by immobilizing BaP analogue on the sensor surface with a mixed SAM of alkanethiolates.329 The biomolecular coating incorporating a BaP analogue was determined to yield a more sensitive sensor with a LOD for BaP in buffer as low as 0.05 ppb.329 In 2005, Kawazumi et al. reported the simultaneous detection of benzo[a]pyrene and 2-hydroxybiphenyl using a compact, portable SPR instrument.52 They employed inhibition detection format and BSA-BaP and BSA-HBP conjugates immobilized on the surface of a dual-channel SPR sensor via physical adsorption. The sensor was shown to be able to detect BaP and HBP in buffer down to parts per billion levels.52

5.4.4. Heavy Metals An SPR sensor for direct detection of Cu2+ ions was reported by Ock et al.330 Their sensor employed a thin polymer layer containing squarylium dye (SQ), which changes its refractive index absorption properties when interacting with Cu2+ ions. Owing to anomalous dispersion accompanied with this absorption, a substantial refractive index change can be observed when SQ dye is exposed to Cu2+ ions. The sensor responded to Cu2+ in buffer at levels as low as 1 × 10-12 M.330 Another SPR sensor for the detection of heavy metals was demonstrated by Wu et al., who used a Biacore X instrument with rabbit metallothinein coupled to the dextran matrix on the sensor surface.286 Metallothinein is a protein that can be found in cells of many organisms and is known to bind to metals (especially cadmium and zinc). Model experiments in which metallothein was used as a receptor demonstrated the potential of this sensor to directly detect Cd, Zn, and Ni in buffer at concentrations down to 100 ng/mL. Forzani et al. demonstrated another approach to the direct detection of Cu2+ and Ni2+ ions. In their work they used a differential SPR sensor coated by properly selected peptides specifically binding metal ions. Detection limits for Cu2+ and Ni2+ in deionized water were 32 and 178 pM, respectively.153

5.4.5. Phenols An SPR biosensor for the detection of bisphenol A (BPA) was developed by Soh et al.332 They used the commercial SPR-20 sensor and inhibition detection format. The sensor surface was modified with a thiol monolayer on which BPA molecules were immobilized through BPA succinimidyl ester. Using a monoclonal antibody, detection of BPA in buffer at concentrations as low as 10 ng/mL was achieved.332 Detection time was approximately 30 min, and the sensor was demonstrated to be regenerable using 0.01 M hydrochloric acid. Another SPR biosensor for the detection of BPA

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Table 1. Overview of SPR Biosensors for Food Quality and Safety Analysis analyte

limit of detection

sensor system

detection matrix

Multiskop Multiskop custom-built

buffer buffer detergent-lysed bacteria heat-killed bacteria untreated bacteria BPV solution aqueous solution PBST solution heat-killed bacteria in buffer apple juice, pH 7.4 pasteurized milk milk; apple juice; ground beef

104 cells/mL 102 cells/mL 104 cfu/mL 105 cfu/mL 106 cfu/mL 8.7 × 106 cfu/mL 106 cells/mL 103 cfu/mL 1.4 × 104 cfu/mL 105 cfu/mL 25 cfu/mL 102-103 cfu/mL

Biacore 2000

heat-killed bacteria in buffer HBS-EP solution buffer buffer milk buffer apple juice, pH 7.4 pasteurized milk

106 cfu /mL 1.7 × 105 cfu/mL 102 cfu/mL 102 cfu/mL 105 cells/mL 4.4 × 104 cfu/mL 104 cfu/mL 23 cfu/mL

custom-built Biacore 3000 custom-built

heat-killed bacteria in buffer PBS solution apple juice, pH 7.4; buffer

custom-built

SR 7000

detection format

ref

Escherichia coli O157:H7

Spreeta Spreeta SR 7000 custom-built Biacore 2000 Spreeta Salmonella spp. S. enteritidis S. group A, B, D, E S. typhimurium S. paratyphi S. choleraesuis S. enterica Lysteria monocytogenes

custom-built Biacore 3000 Multiskop Multiskop Plasmonic custom-built

direct direct sandwich

228 229 230

direct direct sandwich sandwich

231 232 233 234

direct direct

235 236

direct sandwich direct direct sandwich sandwich

162 237 238 239 240 234

direct

235

107 cfu /mL 105 cells/mL 3 × 103 cfu/mL

direct competitive sandwich

162 241 234

heat-killed bacteria in buffer apple juice, pH 7.4

1 × 105 cfu/mL 5 × 104 cfu/mL

sandwich

234

PBST soluiton

105 cfu/mL 107 cfu/mL 104 cfu /mL

sandwich direct direct

242

104-5 × 104 cfu/mL

sandwich

234

102 cfu/mL

direct

244

4 × 105 cfu/mL

direct

245

106 oocyst/mL 102 oocyst/mL

direct

246 246

Campylobacter jejuni

Staphylococcus aureus

Spreeta buffer mixture of E. coli, C. jejuni, S. typhimurium, and L. monocytogenes cutom-built heat-killed bacteria in buffer; apple juice Yersinia enterocolitica Multiskop buffer Vibrio cholerae O1 Multiskop buffer Cryptosporidium parVum Biacore 2000 buffer

243

Fusarium culmorum

Staphylococcal enterotoxins SEB

Biacore X

wheat

0.25 ng/µL

direct, PCR amplicon

247

Biacore X fiber optic Spreeta

1 ng/mL 10 ng/mL 5.6 ng/mL 28 ng/mL 0.6 ng/mL 1.4 ng/mL 5 ng/mL 0.5 ng/mL 0.3 ng/mL 2.5 ng/mL 1 ng/mL

direct direct direct direct sandwich sandwich direct sandwich inhibition sandwich competitive

248, 249 250 251

Biacore 1000 Biacore 1000 Biacore 1000

milk buffer buffer seawater buffer seawater buffer; milk buffer; milk milk buffer; ham tissue extract raw eggs

Biacore 3000 custom-built Biacore Q Spreeta

buffer buffer shellfish species extract buffer; clam extracts

5 ng/mL 0.1 ng/mL 1-7 µg/g 3 ng/mL

competitve inhibition inhibiton inhibition

255 256 257 258

Biacore 1000 Biacore 3000

buffer buffer

3 ng/mL 190 pg/mL

inhibition inhibition

259 148

Biacore Q

buffer

2.5 ng/mL

inhibition

260

custom-built

SEA domoic acid

91 252 253 254

aflatoxin B1

deoxynivalenol

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Chemical Reviews, 2008, Vol. 108, No. 2 487

Table 1 (Continued) analyte antibiotics benzylpenicillin, ampicillin, amoxicillin Row70cloxacillin cephalexin cefoperazon chloramphenicol, chloramphenicol glucuronide

sensor system Biacore 3000

Biacore Q

tylosin vitamins B2 (riboflavin) B5 (pantothenic acid) B8 (biotin), B9 (folic acid) B12 (cobalamine) hormones progesterone diagnostic antibodies Mycoplasma hyopneumoniae classical swine fever virus allergens peanut allergens peanut proteins conalbumin seasame seed protein tropomyosin proteins β-casein Rs1-casein κ-casein IgG folate-binding protein lactoferrin lactoperoxidase chemical contaminants 4-nonylphenol IGF-1

raw milk

honey extract; prawn; dairy products; porcine kidney poultry honey prawn milk shrimps

limit of detection

detection format

ref

2 ng/mL

inhibition

262

15 ng/mL 50 ng/mL 25 ng/mL 0.2 µg/kg

direct

263

inhibition

264

inhibition

265

indirect

266

inhibition

267 268

Biacore Q

buffer raw milk honey honey extract

0.005 µg/kg 0.02 µg/kg 0.04 µg/kg 0.04 µg/kg 1 ng/mL 0.2 ng/mL 250 ng/mL 0.5 ng/mL 1 ng/mL 15 ng/mL 25 µg/kg 2.5 µg/kg

Biacore Q Biacore Q Biacore Q Biacore Q

milk-based products various foods milk powder; infant formulas milk; infant formulas

70 ng/mL 4.4 ng/mL 2 ng/mL 0.06 ng/mL

inhibition inhibition inhibition inhibition

269 270 271

Biacore 2000

60 pg/mL 0.6 ng/mL 8.6 pg/mL

inhibition

273

Biacore 2000

buffer milk buffer

imhibition, amplification

212

Autolab Esprit SPR Autolab Esprit SPR

pig serum pig serum

direct direct

274 275

Spreeta Biacore Q

buffer chocolate pasta bread pasta

700 ng/mL 1 µg/g 0.3 µg/g 12.5 µg/g 10 µg/g

direct direct sandwich

276 277

Biacore 3000

milk

sandwich

278

Biacore Q

milk

85 ng/mL 870 ng/mL 470 mg/mL 16.8 ng/mL 0.7 ng/mL 1.1 ng/mL 75 ng/mL

direct

279

Biacore Q

buffer shellfish buffer, milk

2 ng/mL 10 ng/g 10 ng/mL

inhibition

280

inhibition

281

Biacore Q

chloramphenicol florephenicol florephenicol amine thiamphenicol tetracycline

detection matrix

Biacore Q

Biacore 3000

Biacore

was demonstrated by Matsumoto’s group.334 Their sensor also used inhibition detection format and BPA-OVA conjugate immobilized on the sensor surface by physical adsorption. The sensor was shown to be able to detect BAP at 1 ng/mL (1 ppb) levels.334 Imato’s group developed an SPR biosensor for the detection of 2,4-dichlorophenol based on competitive detection format.333 They used the commercial SPR sensor SPR20 functionalized with monoclonal antibodies against 2,4dichlorophenol immobilized on the sensor surface via gold binding peptide and protein G. Detection was based on the competition between the analyte present in sample and added BSA-2,4-dichlorophenol conjugate. The LOD was established at 20 ng/mL.333

5.4.6. Polychlorinated Biphenyls Karube’s group demonstrated the detection of PCB 3,3′,4,4′,5-pentachlorobiphenyl using the commercial SPR sensor Biacore 2000 and competitive detection format.333 In

their work, the sample was mixed with a conjugate of PCBHRP and flowed across the sensor surface with polyclonal antibodies immobilized in the dextran matrix. The presence of PCB was detected as a decrease in the binding of PCBHRP conjugate. The LOD was determined to be 2.5 ng/mL. The sensor surface was regenerable using 0.1 M hydrochloric acid. The detection was completed in 15 min.335

5.4.7. Dioxins Karube’s group also demonstrated an SPR biosensor for the detection of 2,3,7,8-TCDD.335 They used the commercial SPR sensor Biacore 2000 and competitive detection format. Monoclonal antibody was immobilized in the dextran layer on the sensor chip by amine coupling chemistry. The sample was mixed with a of 2,3,7,8-TCDD-HRP conjugate and injected into the sensor. A LOD of 0.1 ng/mL was attained, and the sensor was shown to be regenerable using 0.1 M hydrochloric acid.335 Detection was completed in 15 min.

488 Chemical Reviews, 2008, Vol. 108, No. 2

Homola

Table 2. Overview of SPR Biosensors for Medical Diagnostic analyte cancer markers prostate-specific antigen (PSA)

sensor system Ibis II Biacore 2000

3% BSA in PBS buffer

PSA-ACT complex (R1-antichymotrypsin)

Biacore 2000

carbohydrate antigen (CA 19-9) vascular endothelial growth factor (VEGF) interleukin-8 (IL-8)

Spreeta

carcinoembryonic antigen (CEA) fibronectin

Autolab Springle

antibodies against viral pathogens hepatitis virus specific

anti-hepatits simplex virus anti-EBNA anti-protein S anti-RSV anti-adenoviral drugs and drug-induced antibodies morphine-3-glucuronide warfarin insulin anti-GM-CSF anti-insulin hormones hCG 17β-estradiol R-fetoprotein allergy markers IgE histamine heart attack markers troponin (cTn I) other molecular biomarkers Glc-6-P isomerase anti-GAD c-reactive protein cystatin C HBsAg

limit of detection

detection matrix

detection format

ref

HBS buffer HBS buffer human serum human serum buffer buffer buffer

0.15 ng/mL 10 ng/mL 1 ng/mL 20.7 ng/mL 10.2 ng/mL 47.5 ng/mL 18.1 ng/L 410.9 U/mL 66.7 U/mL 1 pM

sandwich direct sandwich direct sandwich direct sandwich direct sandwich sandwich

buffer human saliva buffer

2.5 pM (∼0.02 ng/mL) 184 pM (∼1.5 ng/mL) 0.5 ng/mL

sandwich

288

direct

289

custom-built

buffer buffer buffer

2.5 µg/mL 0.5 µg/mL 0.25 µg/mL

direct sandwich amplification

212

Biacore 1000 Spreeta

human serum 5% serum in PBS

150 202

Biacore X custom-built Biacore X Biacore 2000 Biacore 3000

serum in HBS (1:100) 1% human serum plasma 1:5 HBS Buffer ascites fluid in Hepes (1:10)

direct direct sandwich amplified direct direct direct direct direct

Biacore 1000 Biacore 3000 SPR-670

urine in buffer (1:250) plasma (1:100) buffer serum serum (1:5) buffer serum

1 ng/mL 4 ng/mL 1 ng/mL 6 ng/mL 0.6 µg/mL 2.91 µg/mL

inhibition inhibition inhibition inhibition direct direct direct

custom-built Spreeta Biacore X custom-built

buffer urine HBS-EP buffer plasma

0.5 ng/mL 46 mIU/mL 0.47 nM (∼0.14 ng/mL) ng/mL

direct sandwich inhibition sandwich

165 208 297 298

SPR 20 SPR 20

buffer buffer

10 ppb 3 ppb

direct inhibition

199 299

2.5 ng/mL 0.25 ng/mL 3 ng/mL

direct sandwich direct

300

custom fiber-optic

serum serum buffer

direct

302

direct sandwich direct direct

303, 304 305 306 307

cutom-built SPRI Biacore X

Biacore 2000 Biacore 2000

Biacore 2000 Biacore 2000 Plasmonic custom-built SPRI Biacore 3000

synovial fluids in Hepes (1:100) HBS-EP buffer buffer buffer buffer

5.5. Summary Over the past 5 years, more than 100 SPR biosensors for the detection of a variety of chemical and biological analytes were demonstrated. Most of these biosensors are based on prism coupling and angular or wavelength spectroscopy of surface plasmons. Commercial SPR systems have played an important role in the development of detection applications due to their increasing spread and the availability of special SPR platforms and kits dedicated to specific applications (e.g., Biacore Q for food analysis). Data collected in Tables 1-3 illustrate recent applications of SPR biosensors and achieved levels of performance. The performance figures should be compared with caution as performance of an SPR

9.2 nM 4.4 nM 0.64 nM 0.2 ng/mL

<µM 1 µg/mL 1 nM 1 µg/mL

285 286 201

287 220

151 152 290 291 292 293 294 203 295 296

301

biosensor is a result of a multitude of factors (performance of optical platform, characteristics of the employed biorecognition element, suitability and degree of optimization of the immobilization method, detection format, and methodology), and thus low performance of one part of the biosensor (e.g., optical platform) can be compensated for by high performance of another component (e.g., biorecognition elements). Clearly, analytes implicated in food safety have received the most attention (Table 1). Bacterial pathogens such as E. coli and Salmonella were the most frequently targeted analytes. Detection limits below 102 bacteria/mL were reported. A great deal of research has been devoted to the

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Chemical Reviews, 2008, Vol. 108, No. 2 489

Table 3. Overview of SPR Biosensors for Environmental Monitoring analyte pesticides atrazine chlorpyrifos carbaryl DDT DDT carbaryl chloropyrifos 2,4-dichlorophenoxyacetic acid

sensor system

detection matrix

Sensia

water

custom-built custom-built Sensia

water water water water

SPR-20 custom-built

limit of detection

detection format inhibition

buffer buffer

20 pg/mL 1 pg/mL 50 pg/mL 1 ng/mL 15 pg/mL 18 pg/mL 50 pg/mL 52 pg/mL 0.5 ng/mL 10 ppt

Biacore 2000 Biacore 2000

buffer buffer

SPR-20 SPR-20 SPR-20

ref

inhibition inhibition inhibition inhibition

310 311 312, 314 312, 313 315 209

inhibition inhibition

316 317

10 pg/mL 10 pg/mL

inhibition inhibition

318, 320-322 324

buffer buffer buffer

0.1 ng/mL 0.01 ng/mL 0.05 ng/mL

inhibition inhibition inhibition

325, 326 327, 328 329

custom-built Biacore X custom-built

buffer buffer deionized water

∼pM ∼µM 32 pM 178 pM

direct direct direct

330 331 153

SPR-20 Spreeta SPR-20 Biacore 2000 Biacore 2000

buffer buffer buffer buffer buffer

10 ng/mL 1 ng/mL 20 ng/mL 2.5 ng/mL 0.1 ng/mL

inhibition inhibition competitive competitive competitive

332 334 333 335 335

2,4,6-trinitrotoluene (TNT) aromatic hydrocarbons 2-hydroxybiphenyl benzo[a]pyrene heavy metals Cu2+ ions Cd, Zn, Ni Cu2+ Ni2+ phenols bisphenol A 2,4-dichlorphenol PCB 2,3,7,8-TCDD

development of SPR biosensors for other significant groups of analytes such as Staphylococcal enterotoxins (best demonstrated LODs < 1 ng/mL) and antibiotics (best LODs < 1-10 ng/mL depending on the substance). Several analytes have been detected also in complex food matrices. In the field of medical diagnostics (Table 2), the most attention has been paid to the development of SPR sensors for the detection of cancer markers (best LODs < 1-100 ng/mL) and antibodies (best LODs < 1-100 ng/mL). However, most of the detection experiments were performed in buffers rather than in clinical samples. The development of SPR biosensors for environmental monitoring (Table 3) has focused mainly on the detection of pesticides. The best LODs ranged from 1 to 100 pg/mL, depending on the analyte. Detection experiments were performed in buffers or real-world water samples.

6. Conclusions In the past 5 years, SPR biosensor technology has made substantial advances in terms of both sensor hardware and biospecific coatings. SPR biosensors have been applied for the detection of a variety of chemical and biological analytes. We envision that the performance of SPR biosensor technology will continue to evolve and that advanced SPR sensor platforms combined with novel biospecific surfaces with high resistance to the nonspecific binding will lead to robust SPR biosensors enabling rapid, sensitive, and specific detection of chemical and biological analytes in complex samples in the field. These biosensors will benefit numerous important sectors such as medical diagnostics, environmental monitoring, and food safety and security.

7. Abbreviations ATR, attenuated total reflection; CCD, charge-coupled device; DNA, deoxyribonucleic acid; ELISA , enzyme-linked

immunosorbent assay; LED, light-emitting diode; LOD, limit of detection; MALDI-TOF, matrix-assisted laser desorption/ ionization time-of-flight mass spectrometry; RIU, refractive index unit; RNA, ribonucleic acid; scFvs, single-chain antibody fragment; SAM, self-assembled monolayer; SP, surface plasmon; SPR, surface plasmon resonance; WDM, wavelength division multiplexing.

8. Acknowledgment I gratefully acknowledge the financial support of the Grant Agency of the Academy of Sciences of the Czech Republic (IAA400500507, KAN200670701).

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CR068107D

Chem. Rev. 2008, 108, 494−521

494

Nanostructured Plasmonic Sensors Matthew E. Stewart,† Christopher R. Anderton,† Lucas B. Thompson,† Joana Maria,‡ Stephen K. Gray,§ John A. Rogers,†,‡ and Ralph G. Nuzzo*,†,‡ Department of Chemistry, University of Illinois at Urbana−Champaign, Urbana, Illinois 61801, Department of Materials Science and Engineering, University of Illinois at Urbana−Champaign, Urbana, Illinois 61801, and Chemistry Division and Center for Nanoscale Materials, Argonne National Laboratory, Argonne, Illinois 60439 Received June 2, 2007

Contents 1. Introduction 2. Theoretical Considerations: Optical Properties of Metal Nanoparticles and Nanoholes 3. Synthesis and Fabrication of Plasmonic Nanostructures 3.1. Solution-Phase Syntheses 3.2. Top-Down Lithography 3.3. Unconventional Lithographic Techniques 3.3.1. Nanosphere Lithography 3.3.2. Colloidal Lithography 3.3.3. Soft Lithography 4. Applications of Plasmonic Nanostructures in Sensing and Chemical Imaging 4.1. Colorimetric Sensing Based on Particle−Particle Coupling 4.2. Nanoparticle-Enhanced Surface Plasmon Resonance 4.3. Exploiting Rayleigh Scattering for Sensing and Imaging 4.4. Label-Free Optical Detection Based on Changes in Refractive Index 4.4.1. Nanoparticle Dispersions 4.4.2. Surface-Immobilized Nanoparticles 4.4.3. Periodic Nanohole Arrays 4.4.4. Random Nanohole Arrays 4.5. Surface-Enhanced Spectroscopies 4.5.1. Surface-Enhanced Raman Scattering 4.6. Plasmonics for Detection Beyond the Diffraction Limit 4.6.1. Plasmon-Enhanced Fluorescence 5. Concluding Remarks 6. Acknowledgments 7. References

494 495 498 498 499 500 500 500 501 502 502 503 504 506 506 506 508 512 512 513 514 515 515 516 516

1. Introduction Surface plasmons (SPs) are coherent oscillations of conduction electrons on a metal surface excited by electromagnetic radiation at a metal-dielectric interface. The growing field of research on such light-metal interactions * To whom correspondence should be addressed. Phone: 217-244-0809. Fax: 217-244-2278. E-mail: [email protected]. † Department of Chemistry, University of Illinois at Urbana-Champaign. ‡ Department of Materials Science and Engineering, University of Illinois at Urbana-Champaign. § Argonne National Laboratory.

is known as ‘plasmonics’.1-3 This branch of research has attracted much attention due to its potential applications in miniaturized optical devices, sensors, and photonic circuits as well as in medical diagnostics and therapeutics.4-8 Plasmonics is also a highly active area due, in part, to recent advances in nanofabrication methodologies.9-12 These methodologies have led to the realization of metal nanostructures composed of nanoparticles (NPs),13 nanoholes,14 and other components15 with precisely controlled shapes, sizes, and/ or spacings.16,17 Such exquisite synthetic control in combination with advances in theory and the emergence of quantitative electromagnetic modeling tools has provided a better understanding of the optical properties of isolated and electromagnetically coupled nanostructures of various sizes and shapes. In addition to control over the geometry and optical properties of nanostructures,18,19 various strategies for modifying the surfaces of these materials make it possible to effect the selective binding and detection of specific targets for chemical and biological sensing.20,21 Detection schemes based on techniques that utilize plasmons experience enhancements that are commensurate with the magnitudes of the associated electric fields. These enhancements lead to new competencies for chemical sensing that are both useful and extraordinarily sensitiveswith detection levels in some cases reaching that of single molecules.22 As the field of plasmonic-based sensing grows, it is understood that the explicit control of nanostructured components will continue to provide techniques with unparalleled sensitivity, improved ease of fabrication, and thus enhanced utility outside of the laboratory. Two types of surface plasmon resonances (SPRs) are used in surface-based sensing: (i) propagating surface plasmon polaritons (SPPs) and (ii) nonpropagating localized SPRs (LSPRs) (The terms ‘propagating’ and ‘nonpropagating’ are used here to describe evanescently confined surface plasmons. It should be kept in mind that a ‘nonpropagating’ LSPR is intimately coupled to ordinary or nonevanescent propagating light since the LSPR is excited by and scatters such light.). SPPs can be excited on thin metal films using grating or prism couplers.23 These plasmons propagate tens to hundreds of micrometers along the metal surface with an associated electric field that decays exponentially from the surface (normal to the dielectric-metal interface).24 Changes in the refractive index above the metal shifts the plasmon resonance condition, which can be detected as intensity, wavelength, or angle shifts in sensing applications.25 SPR sensors that utilize propagating SPPs are covered in an article by Homola in this issue of Chemical ReViews. LSPRs are

10.1021/cr068126n CCC: $71.00 © 2008 American Chemical Society Published on Web 01/30/2008

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Matthew Stewart received his B.S. degree in Chemistry from Wittenberg University (Springfield, OH) in 2002, graduating summa cum laude with both departmental and university honors. In the summer of 2003 he worked as a flavor chemist at Givaudan, developing yeast blends, conducting accelerated aging experiments, and testing production lots for quality control. He joined Professor Ralph Nuzzo’s group at the University of Illinois at Urbana−Champaign in the fall of 2003 to pursue his Ph.D. degree in Analytical Chemistry. His research interests include plasmonics, microfabrication, nanostructured materials, biomolecular−surface interactions, and microanalytical systems.

Lucas B. Thompson was born on March 27, 1981, in Cleveland, OH. Lucas received his B.A. degree from The College of Wooster (Wooster, OH) in 2003. While at Wooster, he computationally modeled the reconstruction of adlayers on metal surfaces under the supervision of Professor Wingfield V. Glassey. In the fall of 2003 he joined Professor Ralph Nuzzo’s group at The University of Illinois at Urbana−Champaign to pursue his Ph.D. degree in Materials Chemistry. His research interests include collecting and analyzing cellular release with microfluidic devices, SPR as an analytical tool for selectively sensing biomolecules, and nanoparticle-enhanced polymer and hydrogel sensors.

Christopher Anderton was born in Colorado Springs, CO. He received his Bachelors of Science degree in Chemistry with highest distinction at the University of Colorado, Colorado Springs (UCCS). While at UCCS he worked under Professor David Weiss, researching flavonoid activity in many commercially available teas and green tea supplements. He also completed two NSF Research Experiences for Undergraduates under Professor Donald Dittmer at Syracuse University and under Professor Francisco Zaera at the University of California, Riverside. He is a third year analytical chemistry graduate student at the University of Illinois at Urbana−Champaign under Professor Ralph Nuzzo. His research interests include plasmonics, plasmonic enhancement from nanomanipulations, and microanalytical systems.

Joana Maria was born and raised in Lisbon, Portugal. She graduated with a degree in Engineering Physics from the New University of Lisbon and is currently pursuing her Ph.D. degree in Materials Science and Engineering under the guidance of Professor John A. Rogers at the University of Illinois at Urbana−Champaign. She does research in soft optical lithography and plasmonic sensors.

nonpropagating plasmon excitations that can be resonantly excited on metal NPs and around nanoholes or nanowells in thin metal films. The spectral position and magnitude of the LSPR depends on the size, shape, composition, and local dielectric environment.26,27 This latter property has been exploited for label-free optical sensing where adsorbateinduced refractive index changes near or on plasmonic nanostructures are used to monitor binding events in real time.28 Electromagnetic field enhancements also accompany these plasmonic resonances, which are used for performing surface-enhanced spectroscopies. This review focuses mostly, although not exclusively, on LSPRs and their use in chemical29 and biological30 sensing and surface-enhanced spectroscopies.31

2. Theoretical Considerations: Optical Properties of Metal Nanoparticles and Nanoholes Gold (Au) and silver (Ag) metal NPs are frequently studied because they can exhibit strong SPRs in the visible wavelength range.32 At these wavelengths, their optical properties are best described by a complex, wavelength-dependent dielectric constant

(λ) ) r(λ) + ii(λ)

(1)

where  ) m2 and m ) n + ik is the complex refractive index given as a function of the refractive index, n, and the absorption coefficient, k. Noble-metal NPs can support LSPRs33 when the incident photon frequency is resonant with the collective oscillation of the conduction electrons confined in the volume of the NPs (Figure 1).34 The simplest type of LSPR is a dipolar LSPR, which can be viewed as following in the limit of the particle’s diameter, d, being much smaller than the wavelength of the incident light, λ (d , λ). The

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Stephen Gray was born in Sherbrooke, Que´bec, Canada. He received his B.Sc. (Hons) degree in Chemistry from Ottawa’s Carleton University in 1977. Carrying out undergraduate thesis research work with Professor James Wright inspired him to pursue a career in theoretical chemistry. He went on to obtain his Ph.D. degree in Chemistry from the University of California at Berkeley in 1982 with Professor William Miller as his Ph.D. advisor. This was followed by post-doctoral work with Professor Mark Child at Oxford University and Professor Stuart Rice at The University of Chicago. He was an Assistant Professor of Chemistry at Northern Illinois University in DeKalb from 1986 to 1990 before joining the scientific staff at Argonne National Laboratory. His research interests include the quantum dynamics of chemical reactions and electrodynamics of metallic nanostructures.

Ralph G. Nuzzo is the William H. and Janet G. Lycan Professor of Chemistry at the University of Illinois at Urbana−Champaign, where he also holds an appointment as a Professor of Materials Science and Engineering. He received his AB degree with High Honors and Highest Distinction in Chemistry from Rutgers College in 1976, where he was also recognized as a Henry Rutgers Scholar, awarded the Merck Prize for undergraduate research, and elected to Phi Beta Kappa. He earned his Ph.D. degree in Organic Chemistry from the Massachusetts Institute of Technology in 1980. He accepted the position of Member of Technical Staff in Materials Research at Bell Laboratories in Murray Hill, NJ, in 1980, where he was named a Distinguished Member of the Staff in Research in 1987sa title held until he left to join the Illinois faculty in 1991. He is a fellow of the American Academy of Arts and Sciences, the World Innovation Foundation, and the American Vacuum Society. In 2006 he was recognized by the Wall Street Journal Innovators Award for Semiconductors and the Adamson Award of the American Chemical Society in 2003 for original discoveries leading to the development of self-assembled monolayers. He currently serves as a Senior Editor of Langmuir as well as a member of numerous advisory boards.

John A. Rogers is Founder Professor of Engineering at the University of Illinois at Urbana−Champaign with primary appointment in the Department of Materials Science and Engineering, where his research includes materials for large area and unusual format electronics. From 1995 to 1997, he was a Junior Fellow in the Harvard University Society of Fellows. He joined Bell Laboratories as a Member of the Technical Staff in the Condensed Matter Physics Research Department at the end of 1997 and served as Director of this department from the end of 2000 to the end of 2002. He has been on the faculty at Illinois since January 2003. He has published more than ∼180 papers and is co-inventor on ∼70 patents and patent applications, more than 40 of which are licensed or in active use. His research has been recognized with many awards including most recently the Xerox Distinguished Lectureship (2006), the Dorn Lectureship at Northwestern University (2007), the 2007 Daniel Drucker Eminent Faculty Award, the highest honor from the University of Illinois College of Engineering, and the 2007 Baekeland Award from the American Chemical Society for outstanding achievement by a chemist under the age of 40. Rogers was elected a Fellow of the American Physical Society in 2006. He serves or has recently served on several Editorial Boards, including those for Applied Physics Letters, Journal of Applied Physics, and Nano Letters. He is Associate Editor of IEEE Transactions on Nanotechnology and SPIE Journal of Microlithography, Microfabrication and Microsystems.

Figure 1. Schematic illustration of a localized surface plasmon of a metal sphere showing the displacement of the electron charge cloud relative to the nuclei. Reprinted with permission from ref 34. Copyright 2007 by Annual Reviews.

conduction electrons inside the particle will all move in phase upon plane-wave excitation. This leads to the buildup of polarization charges on the particle surface that will act as a

restoring force, allowing a resonance to occur at a specific frequency known as the particle dipole plasmon frequency.33,35 A resonantly enhanced field builds up inside the NP, which in the small particle limit is homogeneous throughout its volume, while a dipolar field is produced outside. This results in strong light scattering, the appearance of intense surface plasmon absorption bands, and the enhancement of the near-field in the immediate vicinity of the particle surface. The spectroscopic responses of larger metallic NPs are modified due to the excitation of higher order modes such as quadrupoles and retardation and skin depth effects.33,35 The bandwidth, peak height, and position of the absorption maximum depend on the particle material, size, and geometry (Figure 2) and the dielectric function of the surrounding environment.33,36,37 Classical Mie theory38 corresponds to the rigorous analytical solution of Maxwell’s equations for the optical properties of a spherical particle. It assumes that the particle and the surrounding medium are homogeneous.33,35 Solutions have

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Figure 2. (a) Dark-field microscopy image and corresponding SEM images and (b) light scattering spectra of Au nanocrystals of different shapes. Reprinted with permission from ref 54. Copyright 2003 American Institute of Physics.

subsequently been developed for other particle shapes, including more general spheroidal forms;32,39-41 such solutions are also often referred to as “Mie Theory”. These analyses are greatly simplified when the particles are much smaller than the wavelength of light since only the lowest (dipolar) order of the Mie theory scattering coefficients need to be retained. In the long wavelength, electrostatic dipole regime, the extinction E(λ)sthe sum of absorption and scattering cross sectionssof a spheroid metallic NP is given by the following equation33

E(λ) ∝

[

i

]

(r + χmed)2 + i2

(2)

where med is the dielectric constant of the surrounding medium, λ is the excitation wavelength, χ is a form factor that describes the NP’s aspect ratio (χ ) 2 for a sphere and increases directly with the aspect ratio of the NP), and r and i are the real and imaginary parts of the dielectric function of the metallic NP, respectively. For a spherical metal NP with |r| . i, eq 2 has a maximum when r(ω) ≈ -2med and this is the LSPR. The position of this resonance red shifts with an increase in magnitude of the dielectric constant of the medium surrounding the NP due to the buildup of polarization charges on the dielectric side of the interface, which is responsible for the weakening of the total restoring force. For nonspherical metallic NPs, the surface plasmons are unevenly distributed around them, manifesting in a shape dependence of the LSPR absorption spectra.42 The plasmon resonance of metallic nanorods, for example, splits into two peaks: (i) a strongly red-shifted long axis or longitudinal mode (L ) polarization parallel to the long axis) and (ii) a slightly blue-shifted transverse mode (T ) polarization perpendicular to the long axis). As the aspect ratio of a nanorod increases, the separation between the two plasmon bands becomes more pronounced (Figure 3).16,43-51 Triangular metallic NPs exhibit multiple plasmon resonances, a longitudinal (bulk) plasmon mode, and very large field enhancements at their sharp tips.52,53 Although a generalization of the quasistatic approach to metallic NPs of arbitrary shape has been suggested and demonstrated good results,54

numerical methods such as the T-matrix method,55 the discrete dipole approximation,56,57 or finite-difference timedomain simulations58,59 usually have to be used to calculate the optical properties (resonant frequencies; local field enhancement at the NP surface; absorption, extinction, and scattering efficiencies) for these more complex cases. For larger metallic NPs beyond the Rayleigh approximation (d > 30 nm), the dipolar resonance red shifts and suffers a substantial broadening. The red shift arises due to a reduction of the depolarization field caused by retardation effects60 where the conduction electrons do not all move in phase, as is true for smaller NPs, which leads to a reduced depolarization field at the particle center generated by the surrounding polarized matter. Radiative losses61 also start to contribute significantly to the plasmon damping, eventually dominating it totally for Au and Ag NPs with diameters d > 100 nm, and for this reason can impact the analytical sensitivity measurements. The latter effect causes significant broadening of the resonance peak. Scattering processes at the NP’s surface are thought to begin to contribute to the total damping for NPs smaller than the free-electron scattering length.33 The depolarization field and additional damping mechanisms for large and small particles can be seen as lowest order corrections to the quasistatic theory leading to decreases in the total enhancement of the excitation field. Mie theory only applies to noninteracting NPs well separated in the solid state or present at low concentration in solution. Recent advances in particle synthesis and fabrication techniques, however, have allowed the assembly of ordered arrays of interacting metallic NPs, leading to interesting new optical responses.3,62-68 The plasmon resonances of interacting particles are split and shifted, depending on the polarization of the incident light, relative to those of noninteracting NPs.65-67 In such cases, each NP with a diameter much smaller than the wavelength of the exciting light acts as an electric dipole. Two types of electromagnetic interactions prevail in this context, depending on the spacing d between adjacent NPs: (i) near-field coupling and (ii) farfield dipolar interactions.62 Far-field dipolar interactions with a d-1 dependence dominate at NP spacings on the order of the wavelength (λ) of the exciting light,69 while near-field dipolar interactions with a dependence of d-3 dominate for spacings much smaller than λ.63 These distance dependences have important consequences for sensing based on the LSPR peak shifts caused by changes in the electromagnetic interactions that occur upon aggregation (or dissociation) of NPs, a feature that has been broadly exploited for colorimetric sensing.70 LSPRs analogous to metal NP LSPRs can be excited around nanoscale holes in thin metal films. This is not surprising in view of Babinet’s principle, which relates the diffraction properties of particles to holes.71 For example, Prikulis et al.72 demonstrated remarkable correlations between the light scattering of holes in metal films and light scattering by disk-shaped NPs. Nanoholes tend to exhibit somewhat broader scattering features due to the fact that, in addition to the possibility of exciting LSPRs, the holes can serve as point sources for SPP waves in the thin film.73,74 As noted in the Introduction, propagating SPP waves form the basis of the SPR sensors reviewed by Homola in this issue. In the case of periodic hole arrays in thin metal films that are reviewed by us here (section 4.4.3), the periodic analog of SPPs and related diffractive phenomena such as

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Figure 3. Transmission electron micrographs (top), optical spectra (left), and photographs (right) of aqueous solutions of Au nanorods of various aspect ratios. The seed sample has an aspect ratio of 1. Samples a, b, c, d, and e have aspect ratios of 1.35 ( 0.32, 1.95 ( 0.34, 3.06 ( 0.28, 3.50 ( 0.29, and 4.42 ( 0.23, respectively. Scale bars: 500 nm for a and b, 100 nm for c-e. Reprinted with permission from ref 16. Copyright 2005 American Chemical Society.

Wood’s anomalies75-77 are intimately entangled with LSPRs and so must also be factored into our discussion. We will consider these influences and their consequences for sensing applications in the sections below.

3. Synthesis and Fabrication of Plasmonic Nanostructures Formation of metal nanostructures has been an active area of research due in part to their growing importance in diverse applications including photonics and optoelectronics,2,3,6,78-82 electronics,83-86 chemical and biosensing,14,87-92 and medical diagnostics and therapies.93-97 It is now well appreciated that the optical, electronic, and catalytic properties of metal nanostructures can be tuned very broadly by controlling their size, shape, and composition.25,35,98-101 This has resulted in a wealth of literature on synthetic methodologies for generating isotropic and anisotropic nanostructures with well-controlled sizes and shapes from a variety of materials.16,17,93,100-104 Nanostructures are typically formed using either so-called top-down or bottom-up approaches.10 Top-down techniques involve using various forms of conventional lithographic techniques to pattern nanostructures (e.g., onto planar substrates),65-67,105 whereas bottom-up methods exploit the interactions of atoms, molecules, or more complex mesoscale objects, in conjunction with the controlling influences of process kinetics, to “assemble” nanostructures either on substrates or in solution.72,106-111 The following sections will describe in more detail the use of these approaches for synthesizing representative metal nanostructures that support propagating and/or localized plasmons and exhibit interesting/useful optical properties.

3.1. Solution-Phase Syntheses Bottom-up solution-phase synthesis is a versatile approach to forming NPs that allows control over their size,112-117 shape,17,101 composition,98,118-120 and structure (e.g., solid or shell).95,121,122 This approach generally involves the reduction of metal salts in a solution containing an appropriate stabilizer to control the growth and suppress the aggregation of the NPs.123,124 The stabilizerscommonly ligands, surfactants, ions, organic acids, or polymerssadsorbs or coordinates to the surface of the NPs and inhibits aggregation by Coulombic

repulsion20,114,125 and/or steric hindrance.124,126,127 Reduction of the metal salt can be carried out electrochemically,15,128-132 photochemically,114,133-136 sonochemically,137-139 or using chemical reductants such as citrate,114,117,140 hydrides,141,142 alcohols,143,144 hydrogen,104 hydroxylamine,145,146 or hydrazine.147,148 The specific choice of reductant, stabilizer, temperature, and relative concentrations of the reagents can all affect the size and shape of the NPs. Recent work has investigated biosynthetic approaches149-153 and other environmentally friendly methods of synthesizing NPs.126,154,155 This topic has been reviewed in a recent issue of Chemical ReViews.156 Solution-phase synthesis tends to yield approximately spherical particles since the lowest surface energy shape is that of a sphere.4,100 (These particles have facets and should be more correctly called ‘quasi-spheres’,4 but will be referred to here as spherical particles for simplicity.) One of the most commonly used procedures for making spherical Au NPs is the citrate reduction of HAuCl4 in an aqueous solution, which was first reported by Turkevitch et al. in 1951.157 In this protocol the citrate acts as both a reducing agent and an electrostatic stabilizer, and the size of the NPs can be tuned by controlling the citrate to HAuCl4 ratio.113,158 Citrate can also be employed in the synthesis of Ag particles.114 Another broadly adopted solution-phase synthesis that yields highly stable, monodisperse thiol-protected particles with controllable sizes is the Brust-Schiffrin method.159 This approach uses a biphasic synthesis (an aqueous and organic phase) and tetraoctylammonium bromide as a phase-transfer agent to reduce the [AuCl4]-1 anion with NaBH4 in the presence of alkanethiols to yield thiol-stabilized particles that are one to several nanometers in size.159 The particle sizes can be controlled between 1.5 and 5.2 nm by adjusting the temperature, reduction rate, and thiol to Au ratio.13,112,156,160,161 The thiols on these so-called monolayer-protected cluster (MPCs) can be easily switched to introduce functionality,162,163 such as thiolated oligonucleotides164 or proteins,165 through simple place-exchange reactions.166-169 Other ligands such as amines and phosphines can also be used to stabilize and control the growth of NPs,170-174 and more recent singlephase adaptations of the Brust method have eliminated the need for phase-transfer agents.170,175-177

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Well-defined NPs can also be fabricated using seedmediated growth techniques.99,102,145,146,178-182 Nanorods are commonly prepared using this approach,100,102,178,179 although they can also be prepared electrochemically49,183-185 and photochemically.133,186 In seed-mediated syntheses, a stable growth solution is prepared containing a metal salt, a mild reducing agent (e.g., ascorbic acid), a surfactant molecule (e.g., cetyl-trimethylammonium bromide, CTAB), and possibly other additives.45,46,100,179,187 A NP ‘seed’, the nucleating agent, is added to the solution and the metal salt is reduced directly on the surface of the seed with minimal nucleation occurring in solution.178 The surfactant acts as an organic micellar template for anisotropic growth. The morphology, aspect ratios, and synthetic yields of the seeding approach are controlled by the composition and concentration of the surfactant, additives, seeds, metal salts, and reducing agent.45,46,100,179,188-191 Seeding growth can also be used to grow Au and Ag nanorods directly on surfaces,192-194 where the initial seed concentration affects the resulting nanorod aspect ratios and size distributions.195 As noted above, nanorods exhibit two distinct plasmon resonance modes: one associated with electron oscillations parallel to the longitudinal axis and the other with electron oscillations parallel to the transverse axis.16,43-51 The longitudinal plasmon mode can be tuned by adjusting the aspect ratio of the nanorod, making it a particularly useful structure for applications in photonics and biotechnology.16,43-51 Au nanoshells are another type of metal nanostructure that have highly tunable plasmon resonances.95,196 Halas and coworkers formed such core-shell structures by seeded metallization of colloidal silica spheres.197 The silica particles are functionalized with a monolayer of an amine-terminated silane (aminopropyltriethoxysilane (APTES)), which is subsequently used to bind small colloidal Au particles. After this initial seeding, more Au is deposited by an electroless plating procedure. The thickness of the final Au shell can be controlled by adjusting the initial Au seed coverage and the amount of reductant used in the plating process.197,198 The plasmon resonance frequency is tuned by adjusting the relative sizes of the core and shell dimensions (Figure 4).95,196 This synthetic approach has also been used to form a socalled nanorice structure,199 prolate spheroidal NPs consisting of a hematite core and a Au shell. The polyol process is a highly versatile synthetic procedure that can be used to form metal and alloy NPs with a variety of shapes, sizes, and optical properties.17,101,144 This synthesis uses a polyol such as ethylene glycol as both a solvent and reducing agent (at elevated temperatures) for a metal salt precursor. Xia and co-workers used the polyol synthesis to produce a variety of Ag nanostructures with well-defined shapes (including cubes, rods, wires, or spheres) by adjusting the relative amounts of the capping agent (poly(vinyl pyrrolidone)) and the precursor salt (AgNO3) in solution.17,101,144,200,201 Nanoplates and nanobelts are formed using a different capping agent (such as sodium citrate).17 Hollow Au or porous Ag/Au alloy nanostructures can also be produced using Ag NPs as a physical template in a wellknown galvanic replacement reaction.121,202 For example, Au nanocages have been generated by simply adding HAuCl4 to a suspension of Ag nanocubes.203 Adjusting the volume of the HAuCl4 solution added to the Ag nanocube suspension allows formation of a variety of different nanostructures such as hollow Au/Ag alloy nanoboxes and/or Au nanocages. The LSPR of these structures can be tuned from the visible to

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Figure 4. Visual demonstration of the tunability of metal nanoshells (top), and optical spectra of Au shell-silica core nanoshells (the labels indicate the corresponding Au shell thickness). Reprinted with permission from ref 95. www.tcrt.org.

near-infrared by varying the amount of HAuCl4 solution added to the Ag nanocube suspension.17,203 Surfactant aggregates such as micelles, reversed micelles, and microemulsions are also used to make restricted volume nanoreactors for the in-situ synthesis of NPs.204-209 The interested reader is referred to recent review articles on this synthetic approach204,205 as well as to a recent Chemical ReView article for more detailed descriptions of NP syntheses.13,156,210

3.2. Top-Down Lithography The size, shape, and interparticle spacing of surface-bound metallic nanostructures can be exquisitely controlled using scanning beam lithographies, such as electron beam lithography (EBL)67,105 and focused ion beam (FIB) lithography.211 This type of control is highly important for making reproducible substrates with tunable optical properties for conducting systematic studies of SERS212-214 and plasmonenhanced fluorescence.215 In EBL a tightly focused beam of electrons is scanned across a thin layer of resist (a radiationsensitive polymer), which makes it either more or less soluble in an organic developer solution. The patterned resist is then used as a sacrificial mask in subsequent etching or deposition processes to generate nanostructured metallic patterns with well-controlled geometries. EBL can be used to attain sub20 nm resolution using specialized resists such as hydrogen silsesquioxane (HSQ)216,217 or NaCl crystals218 or using more traditional organic resists such as poly(methylmethacrylate) (PMMA) in conjunction with ultrasonically assisted development.219 FIB is a related technique that uses a focused beam of ions (typically Ga+) to perform both additive and subtractive patterning by physical or chemically assisted processes. These include (1) FIB milling,211,220,221 (2) ion-assisted etching,222,223 and (3) FIB-induced deposition.224,225 FIB is capable of forming patterns with ∼10 nm resolution using either PMMA226 or inorganic resists.227 Various metallic nanostructures such as circular slits,221 nanoholes,228,229 slit

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Figure 5. Schematic illustration of hexagonally close pack colloidal crystal mask (left), and a representative AFM image of a triangular NP array (right). Reprinted with permission from ref 109. Copyright 2001 American Chemical Society.

gratings,230,231 and V-grooves6,232 have been fabricated by FIB for research in the growing field of plasmonics. While scanning beam lithographies are capable of precise control over the size, shape, and spacing of metallic nanostructures, more recent research has focused on unconventional lithographic techniques that are capable of patterning large areas in parallel at low cost.

3.3. Unconventional Lithographic Techniques 3.3.1. Nanosphere Lithography Nanosphere lithography (NSL) is an inexpensive and versatile hybrid bottom-up procedure, popularized by Van Duyne, for fabricating periodic arrays of metallic nanostructures on surfaces.109,233 This unconventional approach to nanofabrication is a variant of a technique originally named “natural lithography” where monolayers of nanospheres are used as deposition or etch masks.234,235 NSL extended natural lithography with the development of double-layer colloidal masks, which provide a degree of control over the size and shape of the resulting NPs.109 Single-layer NSL begins with the deposition of a singlelayer colloidal crystal mask of hexagonally close-packed latex or silica spheres on an appropriate substrate, which yields defect-free domains that are approximately 10-100 µm2 in size.109,236,237 The colloidal crystal contains triangular void spaces that are created between three neighboring particles, which make these crystals useful as masks in subsequent deposition or etching processes.91,109,238 In the case of additive lithography, a metal or other material is deposited through the mask at normal incidence by physical vapor deposition to produce a metallic film over nanosphere (MFON) structure.31,239,240 This approach has been used to produce Ag FON surfaces that are useful structures for applications based on surface-enhanced Raman spectroscopy (SERS).240-242 Alternatively, the nanosphere mask can be removed by sonicating in solvent to yield arrays of surface-bound triangular NPs with P6mm symmetry. These structures can be used for LSPR sensing and SERS (Figure 5).109,243-246 Nanostructured films composed of periodic spherical voids can also be formed using self-assembled polystyrene (PS) colloidal crystal masks. These ‘nanovoid arrays’ support delocalized Bragg and localized Mie plasmons247,248 and are formed by electrochemical deposition of Au through a singlelayer self-assembled colloidal template. The thickness of the metal is controlled by monitoring the current passed through the plating solution and the plating time.18,249 This allows

Figure 6. SEM images of nanovoid arrays with void diameters of 600 nm at three normalized thicknesses t-)0.2, 0.5, and 0.9 (left), and schematic illustrations of the surface at each thickness (right). Reprinted with permission from ref 247. Copyright 2006 The American Physical Society.

control over the geometry and thus the spectroscopic/ plasmonic properties of the nanostructured films.18,250 After completing the metal deposition step, the PS spheres are dissolved in tetrahydrofuran to yield a ‘nanovoid array’ (Figure 6). These structures have been used with some success as substrates for SERS.249-252 Self-assembled colloidal crystals also can be used as an etch mask to form arrays of submicrometer triangular pits in an underlying substrate.238 Anchored arrays of triangular metal NPs110 and metal films over nanowells253 can be formed using this subtractive processing technique. The anchored NPs are formed by depositing metal at normal incidence by physical vapor deposition on the colloidal mask and underlying etched substrate. The spheres are subsequently removed by sonication in an appropriate solvent, leaving behind an array of substrate-embedded triangular NPs.110 Alternatively, metal deposition can be performed after the spheres are removed from the etched substrate. This results in structured metallic surfaces that support a single, narrow plasmon resonance that exhibits sensitivity to external changes in refractive index.253 The attractive features of NSL include its low cost, versatility, and ability to produce well-ordered sub-100 nm array structures. The geometry and shapesand thus the optical propertiessof the particles can be tuned by moving the sample during metal deposition,254 annealing the latex colloidal mask prior to metal deposition,255 as well as changing the thickness of the deposited metal, size of the colloidal spheres, number of colloidal layers, or angle of the metal deposition.109,256,257 The size and shape of the NPs can also be changed in a controllable manner using electrochemistry to oxidize the substrate-bound particles.258 The surfacebound NPs formed by NSL can be released from the substrate into solution by adding surfactant and sonicating to generate both isolated particles and dimers.111

3.3.2. Colloidal Lithography Colloidal lithography72 is a versatile technique that can be used to form random arrays of nanoholes,259 nanodisks,260

Nanostructured Plasmonic Sensors

Figure 7. Schematic depiction of nanoring fabrication. (a) (1) Polystyrene colloidal particles are deposited by electrostatic selfassembly onto the substrate in a dispersed layer. (2) A 20-40 nm thick Au film is evaporated onto the particle-coated substrate at normal incidence. (3) Argon-ion beam etching is used to remove the Au film. During the etching, secondary sputtering creates a Au shell around the base of the polystyrene particles. (4) The remainder of the polystyrene particles are removed by UV-ozone treatment, resulting in free-standing Au nanorings on the substrate. SEM images of Au nanorings (150 nm diameter) made using a 40 nm thick sacrificial Au layer at normal incidence (b) and 80° tilt (c). Reprinted with permission from ref 383. Copyright 2007 American Chemical Society.

and nanorings.108 This unconventional nanofabrication technique involves adsorbing polystyrene (PS) particles onto a substrate via electrostatic self-assembly. The distance between the self-assembled spheres is governed by the particleparticle repulsion, which can be controlled by adjusting the electrolyte concentration of the colloidal solution.107 The randomly adsorbed particles are then used as a mask for subsequent evaporation and/or etching processes wherein the size of the sphere dictates the size of the resulting nanostructures. The main difference between colloidal lithography and NSL is that the colloidal spheres do not form an hcp monolayer on the substrate in colloidal lithography. Random nanoholes in Au films can be formed by evaporating metal on top of an assembled colloidal mask followed by lift off of the PS particles by tape stripping or boiling the sample in ethanol.259 Nanorings are formed on substrates in a similar process with the exception that prior to particle removal an Ar+-ion beam is used to etch the Au film, during which secondary sputtering creates Au shells around the bottom of the PS particles.108 The remains of the particles are then removed by an UV-ozone treatment, which leaves free-standing Au rings (Figure 7). Nanodisks can be formed in two ways.260,261 In one approach, nanodisks are formed by assembling a colloidal mask on top of a Au film followed by etching of the Au and removal of the colloidal mask.260 In the second approach, a colloidal mask is self-assembled on a PMMA film followed by deposition of a thin Au film. The Au-capped PS particles are then removed by tape stripping, leaving behind holes in the Au film. The exposed PMMA is then etched from the holes, and metal is deposited through the holes onto the substrate.

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Figure 8. Fabrication and structural characterization of large-area hole arrays. Preparation of free-standing films of subwavelength hole arrays (a). SEM image of a portion of a free-standing 100 nm Au film perforated with 250 nm holes (b). Scale bar: 500 nm. Optical micrograph of a free-standing film placed on a glass substrate (c). SEM images of representative areas of the film illustrating the uniformity of the nanohole array (d and e). The holes are spaced 1.6 × 2.4 µm. Scale bar: 2 µm. Reprinted with permission from ref 270. Copyright 2005 American Chemical Society.

The metal-coated PMMA film is removed by lift off, leaving a random array of metal disks on the substrate.261

3.3.3. Soft Lithography Soft lithography refers to a set of microfabrication techniques that use a structured elastomer as a stamp, conformable photomask, or mold to pattern a material of interest.262-264 The most commonly used elastomer, Sylgard 184 poly(dimethylsiloxane) (PDMS), has a low modulus, which limits the utility of this elastomer for patterning in the nanometer regime.265,266 The development of siloxane-based elastomers with larger moduli, such as the socalled hard PDMS (h-PDMS)266-268 and UV-curable PDMS (hV-PDMS),269 has extended the patterning ability of soft lithography to the nanometer regime. Composite stamps consisting of a thin layer of structured h-PDMS supported by a thicker planar layer of compliant PDMS are typically used for patterning at the nanoscale due to the mechanical instabilities and difficulty of achieving conformal contact with nonplanar surfaces using h-PDMS stamps. In one embodiment of soft lithography for fabricating plasmonic structures, Odom et al. used high-resolution composite PDMS stamps as conformable phase masks to generate large-area, free-standing 2D nanohole arrays in Au.270,271 A schematic illustration of the fabrication process is shown in Figure 8. In this process, an array of posts of positive photoresist with diameters of ∼250 nm is patterned first on a Si(100) wafer by phase-shifting photolithography using a conformable composite PDMS photomask.266,271,272

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thane film with a composite h-PDMS/PDMS stamp presenting surface relief features in the geometry of a square array of cylindrical posts. The composite PDMS stamp is pressed into the liquid polyurethane film and cured by UV light passed through the stamp. The stamp is then removed, and a thin layer of Au (∼50 nm) is uniformly deposited by e-beam evaporation onto the raised and recessed regions of the structured polyurethane film. This creates a Au film with an array of nanoscale holes (top surface) that is physically separated from a second level of Au disks at the bottom of the wells (Figure 9).

4. Applications of Plasmonic Nanostructures in Sensing and Chemical Imaging 4.1. Colorimetric Sensing Based on Particle−Particle Coupling

Figure 9. Plasmonic crystal fabrication process: (a) imprint, (b) cure, (c) remove stamp, and (d) Au deposition (top). SEM of a crystal and a high-magnification SEM (inset) that shows the upper and lower levels of Au (bottom). Reprinted with permission from refs 8 (top) and 77 (bottom). Copyright 2005 OSA (top) and Copyright 2006 The National Academy of Sciences of the USA (bottom).

A thin layer of chromium (Cr) is then deposited by electronbeam (e-beam) evaporation followed by removal of the photoresist posts. This yields an array of holes patterned in a thin film of Cr, which acts as an etch mask and as a sacrificial release layer to generate the free-standing nanostructured Au films. The exposed Si is then anisotropically etched using a KOH/isopropyl alcohol (IPA) solution to produce pyramidal-shaped voids directly below the Cr nanoholes. A layer of Au can then be deposited by e-beam evaporation to form a 2D nanohole array Au film on the top surface and pyramidal mesoscale Au particles in the lower pyramidal voids. The top Au nanohole array can be released by selectively etching the Cr film to form a free-standing structure (Figure 8), and the Au pyramids can be released by etching the Si with KOH/IPA. The material and chemical functionality of these nanostructures are easily controlled by depositing other metals or a combination of metals in a layerby-layer mode using e-beam evaporation.270 The periodic grating structure of the nanohole array allows the direct coupling of light to SPPs on these films.270,273 This fabrication protocol has been used to produce pyramidal metallic particles274 and particle arrays.275 Imprint lithography is another unconventional lithographic technique holding broad utility for patterning materials at the nanoscale over large areas.8,77,276-278 This rapidly emerging technology is used to replicate features on a hard or soft stamp in a thermoplastic or UV-curable polymeric material by embossing or molding (Figure 9). A metal film can be deposited on the resulting polymeric replicas to produce plasmonic structures useful for chemical and biosensing applications.8,77,279,280 In a recent example, large-area spatially coherent arrays of quasi-3D plasmonic crystals were formed by soft nanoimprint lithography and used for multispectral sensing and imaging of molecular binding events. The arrays were fabricated by embossing a thin UV-curable polyure-

Colorimetric detection is perhaps one of the most powerful and simple nanosensing methods available. In an exemplary model of this approach, Mirkin et al.70 reported an assay using oligonucleotide-functionalized Au NPs that exhibit strong red shifts upon aggregation in the presence of a complementary nucleotide (Figure 10).281 The color change in this case results from particle-particle plasmonic coupling as well as aggregate scattering281 and provides a “litmus test” method for determining nucleic acid targets. The optical properties of these assays are due to the aforementioned resonantly exited LSPRs of the NPs. The enhanced electronic fields are confined within a small area around the NPs (typically on the order of the particle radius) and decay approximately exponentially thereafter.65,66 As the distance between the NPs decreases, near-field coupling begins to dominate, leading to a strong enhancement of the localized electric field within the interparticle spacing producing pronounced red shifts of the LSPR frequency.65-67 Most reports of colorimetric assays exploit the LSPRs that develop on spherical Au NPs, but the method is also amenable to nonspherical particles21,282 as well as particles of other noble metals.283,284 Some benefits to using elongated particles include their inherent higher sensitivity to changes occurring in the local dielectric environment285 and the capability they afford for multiplexed detection schemes.286 Methods exploiting nucleotide interactions have been the most reported of these particle-particle coupling systems. This field of research has advanced rapidly, and biological analyses using oligonucleotide-modified Au NPs have been developed that achieve limits of detection (LOD) for a variety of analytes in the low-picomolar to mid-femtomolar range.281 Han et al. used oligonucleotide-functionalized Au NPs to determine the relative binding strengths of a variety of duplex and triplex DNA-binding molecules.287,288 Such data can offer insights into the activity of an array of possible anticancer drugs due to the correlations that can be made between ligand binding strength and biological activity. Kanaras and coworkers289 demonstrated the use of DNA-Au NP interactions for the determination of the enzymatic cleavage of DNA, in which a recognition site for the restriction endonulease EcoRI was designed to monitor the enzymatic cleavage activity, and consequent disassembly of the closely coupled Au NPs, via blue to red color shifts. Peptide nucleic acid (PNA) and aptamer-modified Au NPs have also been used to control assembly rates and aggregate sizes of Au NPs for sensing applications.87,290-293 The dissolution of aggregated NPs can also be used to monitor binding events.

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Figure 10. In the presence of complementary target DNA, oligonucleotide-functionalized Au NPs will aggregate (left), resulting in a change in the color of the solution from red to blue (right). Reprinted with permission from ref 281. Copyright 2005 American Chemical Society.

Figure 11. Schematic representation of colorimetric detection of adenosine. Absorbance spectra of the adenosine sensor before (blue) and 10 s after (red) addition of adenosine. Reprinted with permission from ref 87. Copyright 2006 Wiley-VCH Verlag GmbH & Co. KGaA, Weinheim.

In a specific example, Liu and Lu87 employed aptamerfunctionalized Au NP aggregates specific to adenosine and cocaine that become unstable when the target was bound, leading to a color shift from purple to red (Figure 11).87 Functionalized Au NPs that do not exploit nucleotide interactions also have been reported that provide a viable route to colorimetric detection.284,294-303 In a prototypical example, the ubiquitous biotin-avidin system linked to Au NPs was utilized to create a colorimetric assay for kinase inhibitors that could be applied to microplates amenable to parallel screening.301 Rapid detection of cholera toxin has been demonstrated using lactose-stabilized Au NPs, in which the cholera toxin binds to a lactose derivative inducing NP aggregation, which is detected visually by a red to purple color shift.299 In an expansion of the current methods for colorimetric detection, cation-specific functionalized Au NPs have been developed.88,282,297,304,305 The detection of Pb2+ is an area that has received much attention due to its physiological impacts. Demonstrating the use of DNAzymes in Au NP assembly for the detection of Pb2+, Liu and Lu304 optimized the DNA configuration for NP alignment, the NP size, Pb2+-specific DNAzyme concentration, and salt concentration, leading to a colorimetric detection method for Pb2+ that can be performed in the field in less than 10 min. In contrast to the previously mentioned method of sensing lead, Lin et al.297 used a crown ether specific to Pb2+ for colorimetric detection of Pb2+ using Au NPs ligated with a mixed monolayer of crown ether thiols and carboxylic acids.306 Recently, Lee and co-workers88 demonstrated a detection method for the mercuric ion (Hg2+) analogous to Liu and Lu,305 which takes advantage of the Hg2+ affinity toward T-T mismatches in DNA. Aside from the cation determination methods mentioned here, additional methods have been reported for determining other relevant charged molecules.296,302,303 All of these methods employ the plasmonic coupling inherent in aggregated NP systems and demonstrate that

detection of desirable biological compounds and other relevant molecules can be accomplished without the use of complex instrumentation. These methods are interesting in that they are sensitive at relevant time scales and that the NPs are easily functionalized to provide chemical and biological selectivity. The optical properties of NPs also have been exploited to enhance the sensitivity of conventional SPR systems as described below.

4.2. Nanoparticle-Enhanced Surface Plasmon Resonance As mentioned above, SPR spectroscopy is a surfacesensitive technique that can be used to detect refractive index changes that occur within the evanescent field of propagating SPPs excited at metal-dielectric interfaces.307 A change in refractive index shifts the plasmon resonance condition, which can be detected as intensity, wavelength, or angle shifts to provide quantitative information about the binding event.307-313 Small refractive index changes caused by the binding of low molecular weight analytes or small quantities of a larger analyte can challenge the sensitivity/detection limits of SPR spectroscopy.314-319 The sensitivity/detection limits can be improved by coupling the molecular recognition of analyte at the surface of the metal with another event that leads to larger changes in the SPR signal.314-319 This signal enhancement can be achieved using competitive inhibition or sandwich assays320-323 and enzymatic amplifaction.312,313,324,325 The SPR signal also can be enhanced by labeling the target analyte with dielectric or plasmonic NPs. These labels increase the refractive index at the metal surface and can electromagnetically couple to the flat metal film in the case of plasmonic NPs, leading to larger SPR shifts.314-319,326 Figure 12a shows a possible architecture for performing plasmonic NP-enhanced SPR immunoassays.314 In this approach a surface-immobilized antibody (e.g., anti-human

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Figure 12. Schematic of plasmonic NP-enhanced SPR immunoassay (a). In situ SPR curves of an evaporated Au film modified with a-h-IgG(γ) (solid line) followed by sequential exposure to a 0.045 mg/mL solution of h-IgG (‚‚‚) and a 8.5 mg/mL solution of a-h IgG(Fc) (- -) (b). A film modified with a-h-IgG(γ) (solid line) followed by sequential exposure to a 0.045 mg/mL solution of h-IgG (‚‚‚) and a solution of a-h-IgG(Fc)-10-nm Au colloid conjugate (- -) (c). Reprinted with permission for ref 314. Copyright 1998 American Chemical Society.

immunoglobulin G (a-h-IgG)) is sequentially exposed to antigen (e.g., h-IgG) and a solution of secondary antibodyconjugated Au NPs (e.g., a-h-IgG/Au NPs). The presence of the Au NP leads to an enhancement of the measured SPR shift as shown in Figure 12b and c. Figure 12b shows that the shift of the SPR peak upon binding the secondary antibody alone is small compared to the shift observed when the secondary antibody is conjugated to a Au NP.314 The particle size, composition, and surface coverage as well as the substrate metal and the distance between the substrate and NP can affect the observed signal enhancement.314,315,327-329 This approach has been used not only for performing immunoassays314,330 but also for detecting DNA hybridization,319,331 protein conformational changes,332 small molecule binding interactions,317,318 and single nucleotide polymorphisms.316 Recently, Fang et al. reported the detection of attomoles of microRNA using NP-enhanced SPR imaging.333 In addition to flat film SPR, NPs can also be used to enhance the peak shifts obtained from transmission LSPR sensing using Au island films.334-336 This platform has been used with responsive polymer brushes, surface-immobilized singlestranded DNA, and molecularly imprinted polymers to detect pH changes,334 DNA hybridization,336 and cholesterol,335 respectively.

4.3. Exploiting Rayleigh Scattering for Sensing and Imaging Biosensing based on the optical scattering properties of plasmonic NPs is regarded as a potentially more powerful, yet remains a less widely exploited, modality of NP sensing than the extinction-based colorimetric assays described above.337,338 A single 80 nm Au NP, for example, exhibits a light scattering power equivalent to the emission of ∼106 fluorescein molecules.339,340 Unlike molecular fluorophores, however, plasmonic NPs do not undergo photobleaching or blinking. These are enabling distinctions that facilitate long-term single-particle measurements and tracking.285,341-343 These optical properties have been exploited in a variety of applications where NPs are used for sensing285,344 and as labels for immunoassays and DNA

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Figure 13. Colorimetric detection of nucleic acids using scattered light (a). Step 1: DNA-Au NP probes (A and B) are hybridized to a DNA target in solution. Step 2: The samples are spotted onto a glass slide, which is illuminated with white light in the plane of the slide. The evanescent induced scatter from the Au NPs is visually observed. Individual 40 to 50 nm diameter Au probes scatter green light, whereas complexed probes scatter yellow to orange light because of a plasmon band red shift (b). Reprinted with permission from ref 352. Copyright 2004 Nature Publishing Group.

microarrays,345-348 imaging contrast agents,95,349 and molecular rulers.341,350,351 As noted in section 4.1, the changes in extinction (absorption and scattering) that occur due to NP aggregation can be used to detect the presence of a specific target DNA or protein.70,287,288,301 Monitoring only the changes in the scattering properties of NPs upon target-induced aggregation can also be used to detect target DNA.344 This method of sensing was demonstrated by immobilizing two different oligonucleotide fragments (A′ and B′, both 15 bases long), separately, on 13 nm Au NPs to create two sets of probe particles that aggregate in the presence of a complementary 30 base long target DNA (AB). Changes in the scattered light intensity upon target-induced aggregation were measured using a commercially available spectrofluorimeter. This approach was used to detect target DNA at picomolar concentrations and allowed the detection of single-nucleotide polymorphisms (SNPs) without the need for temperature control. Even lower concentrations of target DNA were detected by Muller et al.352 using 50 nm Au particles in conjunction with a light-scattering-based spot test. In this work, a target sequence was added to a solution containing oligonucleotidemodified Au NPs to induce hybridization and aggregation of the probes. An aliquot of the solution was then spotted on a glass slide (a planar waveguide) that was illuminated using a planar fiber optic illuminator (Figure 13a). The evanescently coupled light was scattered from the particles at the surface of the waveguide and imaged with a complementary metal-oxide-semiconductor (CMOS) camera. Au NPs not exposed to target DNA scattered green light (control), whereas Au NPs aggregated in the presence of the target DNA scattered orange light due to particle-particle coupling (Figure 13b). This approach allowed the detection of femtomolar concentrations of target DNA without the need for PCR or signal amplification.352 The high sensitivity and simple readout make this approach highly promising for use in point-of-care molecular diagnostics.

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Figure 14. Schematic of DNA hybridization to microarrays and detection using Ag-amplified Au NP probes. Reprinted with permission from ref 346. Copyright Elsevier B.V.

Wavelength-ratiometric plasmon light scattering is another sensing method that measures changes in scattered light to detect target-induced aggregation or dissociation of NPs in solution.353,354 In this approach, the ratio of the scattered light intensity at two wavelengths is used to quantify a target analyte, such as glucose.353 Other ‘ratiometric’ methods of biosensing include measuring the ratio of scattered light intensity355 or polarization356 at two different angles as a function of target-induced aggregation or dissociation of plasmonic NPs. The benefits of ratiometric approaches are that the measurements become independent of source and/ or detector fluctuations and NP concentration, and in this way improve the analytical stability of protocols relative to single-wavelength measurements. Fluorescent tagging is the most common method of labeling targets for optical detection; however, the demand for greater sensitivity and simplicity (by removing the need and thus cost and complexity of PCR oligonucleotide target amplification) has led to research on and development of alternative labels, such as plasmonic NPs.345-348 As an example, the large scattering cross section of plasmonic NPs has also led to their use as alternative labels in microarraybased technologies.345-348 Microarrays are made by immobilizing biologically relevant moieties, such as DNA or proteins, as discrete spots, typically 10-500 µm in size, on a substrate.345 These spatially arranged capture agents are then used to bind targets through specific interactions such as hybridization (nucleic acids) or ligand-receptor binding (proteins) to allow multiplexed detection of multiple analytes in a complex solution. Plasmonic NP labels have been shown to be very promising for improving the sensitivity of microarray-based analyses.345-348 Muller et al.347 recently used 15 nm oligonucleotide-modified Au NPs to label captured target DNA on a microarrayed glass slide using a three-component sandwich assay (Figure 14). Once the Au NPs were immobilized, a Ag amplification step was used to increase the target signal by electroless reduction of Ag ions to metal at the surface of the Au NPs.346 The glass slide was then used as a waveguide, and evanescently coupled light scattered from the NPs was imaged. The Ag-amplified Au NPs provided an approximately 1000-fold increase in sensitivity compared to the Cy3 fluorescent labels commonly used in microarray analyses.346 This protocol was used to detect femtomolar concentrations of target sequences in human genomic DNA samples without prior PCR amplification.347 A similar procedure was used to perform multiplexed SNP genotyping in total human genomic DNA without prior

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complexity reduction or target amplification, which is a major advancement toward point-of-care diagnostic medical applications.348 It is interesting to note that these light scattering measurements were performed using commercially available systems357 that can achieve attomolar and zeptomolar sensitivity when used in conjunction with immobilized transcyclopentane-modified peptide nucleic acid capture strands358 or bio-bar-code techniques,359 respectively. The large scattering cross section339 and ability to tailor the scattered wavelength of Au NPs,95 in conjunction with their biocompatibility198 and the availability of well-characterized surface conjugation chemistries,360 also makes them attractive candidates for use as contrast agents for imaging applications.95,349 Such optical measurements provide a promising route to noninvasive, high-resolution diagnostic imaging of tissues and cells with high sensitivity and chemical specificity.95,350,361 For example, cancer cells can be labeled by conjugating plasmonic NPs to antibodies that target a protein that is overexpressed by cancerous cells. AntiHER293,94 and anti-EGFR362-364 are commonly used antibodies that target epidermal growth factor receptors (EGFRs),96,364,365 which are transmembrane glycoproteins that are overexpressed in many types of cancers such as cervical, bladder, breast, lung, and oral cancers.365 The cells or tissues are incubated with the antibody-conjugated plasmonic NPs, and the labeled cells are then examined using an appropriate form of optical imaging (e.g., dark field microscopy, twophoton luminescence, etc.).97,362-364 Both spherical Au NPs362-364 and nanorods96,97 have been conjugated with antibodies for use in imaging a variety of cancer cells. Recently, Sokolov et al. used multifunctional Au NPs that incorporated both cytosolic delivery and targeting moieties for real-time intracellular imaging of biomarkers in live cells.366 The authors used water-soluble 20 nm Au NPs formed by citrate reduction that were functionalized with (i) TAT-HA2 peptides, (ii) anti-actin antibodies, and (iii) 5000 molecular weight PEG-SH (poly(ethylene glycol)thiol)). The first component serves two purposes: (a) the TAT protein transduction domain induces endocytic uptake of the functionalized Au NPs into the cell, and (b) the HA2 protein destabilizes the endosomal lipid membrane, which causes release of the NPs into the cytosol.367 The second component allows the NPs to bind to actin, and the third component improves the biocompatibility of the multifunctional particles. NIH3T3 fibroblasts were labeled with these probes and imaged using dark field reflectance microscopy. Actin labeling by these intriguing plasmonic probes was observed as an increase in red scattering due to dipole-dipole coupling between the NPs, and live cell imaging was used to track actin rearrangement. While the scattering properties of NPs are useful for labeling and optically detecting cancer cells, their large optical absorption cross section can also be exploited for photothermal therapeutic treatments.93,94,96 The treatment begins with the labeling of cancerous cells with molecular probe-conjugated plasmonic NPs. The labeled cells are then irradiated with light, which is absorbed by the NPs and converted to heat. The local heating causes irreversible damage and kills the cancerous cells. Research has focused on developing nanostructures with absorption peaks in the near-infrared (NIR) due to the transmissivity of blood and tissue at these wavelengths.198,368 Au nanoshells,94 nanocages,93 and nanorods96 with NIR absorption peaks have all been used to demonstrate this mode of photothermal treat-

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ment, although spherical Au NPs with absorption bands in the visible region of the electromagnetic spectrum have also been used.369

4.4. Label-Free Optical Detection Based on Changes in Refractive Index Plasmons are exceptionally sensitive reporters for chemical phenomena that influence the refractive index of the local environment of a probe. SPR spectroscopy307,370,371 and imaging24,28,313,333,372,373 are well-known label-free optical detection methods that use this property to monitor surface binding events in real time (as discussed in a separate contribution in this issue). In these techniques, changes in the dielectric environment at the surface of a flat or periodically structured noble-metal film shift the observed SPR resonance, which can be measured using angular interrogation, wavelength interrogation, or intensity measurements. This well-established technique enables determination of kinetic and thermodynamic data for a wide variety of molecular binding events, especially those involving biomolecular targets.371,374,375 Progress is currently being made to develop nanostructured plasmonic materials for performing similar analyses. Changes in the dielectric environment of nanostructured metals, such as NPs, results in measurable shifts of the LSPR peak position and/or magnitude that can be used to perform label-free chemical or biosensing in real time.342,343 A variety of noble-metal nanostructures such as NPs in solution26,286 or immobilized on surfaces,5,29,376 nanoholes and nanohole arrays,8,14,77,90,377,378 and nanoisland films379-381 have been used in this way. Substrate-bound nanostructures offer several attractive features as platforms for chemical sensing. These include (i) the shape, size, composition, and spacing of the NPs can be readily controlled to provide tunable peak positions and widths65,77,105,109,256 and (ii) the NPs are free of the capping agents or stabilizers used in solution-phase NP synthesis, making their surfaces readily accessible for functionalization with specific receptors or ligands.77,92,382,383 In the following discussion we will provide some examples of chemical and biosensing based on refractive index changes near the surface of nanostructured metals in solution and on substrates, where the LSPR is monitored using extinction or scattering spectroscopy. The refractive index sensitivity of each sensing platform will also be given where validated data is available. This sensitivity is commonly defined in terms of the change in an experimentally measurable parameter (typically peak position or magnitude) per ‘refractive index unit’ (RIU), which corresponds to a change of 1 in the refractive index. These measurements are usually performed by taking spectra of a plasmonic nanostructure in solutions of increasing refractive index while monitoring peak position or intensity changes.

4.4.1. Nanoparticle Dispersions Au nanostructures are known to exhibit strong plasmonic bands that are dependent on their shape, size, and surrounding media.26,286,384 Ghosh et al.26 studied the effects of changing solvents and ligands on the LSPR of Au NPs dispersed in solution. It was found that the surface plasmon absorption maximum of the Au NPs varied between 520 and 550 nm, depending on the refractive index and chemical nature of the surrounding solvent. The authors found that the LSPR peak red shifted linearly with the refractive index of the solvent when using solvents that do not possess active

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functional groups that could complex with the surface of the Au NPs. A nonlinear relationship between the LSPR peak position and refractive index was found, however, when using solvents with nonbonding electrons capable of complexing to the surface of the Au NPs. Interestingly, these authors found that the LSPR peak position blue shifted ∼3 nm for every one carbon atom when the NPs were dispersed in alcohols with varying linear carbon chain lengths.26 This rather atypical trend reverses in the presence of more strongly coordinating ligands. In such cases, the magnitude of the red shift caused by stabilizing ligands, such as alkyl amines or thiols, increases when the headgroup of the ligand interacts more strongly with the surface of the Au NPs. In a recent study,286 a dispersion of gold nanorods (GNRs) with different aspect ratios was used to perform a multiplexed bioanalytical sensing measurement in solution. This work exploited the fact that small changes in the aspect ratio of GNRs lead to drastic changes in their optical properties (the longitudinal plasmon mode red shifts with increasing aspect ratio, as described in section 2). A series of GNRs with aspect ratios (length/width) of 2.1, 4.5, and 6.5sannotated as GNR 1, 2, and 3, respectivelyswere functionalized with a different recognition molecule (Figure 15a) and dispersed into a single solution.286 The solution yielded a composite absorption spectrum with three pronounced longitudinal LSPR maxima, each one corresponding to one of the three types of GNR (Figure 15b). The longitudinal peaks, in order of decreasing energy, correspond to GNR 1 > GNR 2 > GNR 3. Targets that were complementary to the different recognition molecules on the GNRs were then added to the solution to initiate binding events that caused a selective red shift of one or more of the longitudinal peaks. For example, when target 1 (complementary to the receptors on GNR 1) was added to the solution, a red shift in the longitudinal peak corresponding to GNR 1 was observed with smaller shifts in the peaks corresponding to GNR 2 and 3 (Figure 15b). When targets 1 and 2 (complementary to the receptors on GNR 1 and 2, respectively) were added to the solution, a shift in the peaks corresponding to GNR 1 and 2 was seen, while only a small shift in the LSPR maximum associated with GNR 3 occurred (Figure 15c). All three peaks red shifted when their corresponding targets were added to the solution (Figure 15d). To our knowledge, this report is the first to take full advantage of the multiplexing potential offered by GNRs and demonstrates the potentially enabling qualities offered by the ability to tailor the optical properties of metallic nanostructures.

4.4.2. Surface-Immobilized Nanoparticles NPs synthesized in solution can also be immobilized on surfaces for potential ‘on-chip’ sensing applications and thereby engender more useful multiplexing capabilities.5,376 Thin Cr or titanium (Ti) films are conventionally applied to substrates such as glass to promote adhesion of Au or Ag to the substrate. These metal layers are known to attenuate and broaden plasmon resonance bands246,385,386 and lead to markedly adverse effects on conventional SPR and LSPR sensing with noble metals.385,387 For this reason, it is beneficial to use organic adhesion layers of amine- or mercapto-terminated silanes to immobilize Au or Ag NPs to oxide-bearing substrates such as silica (Figure 16).5,376,385,388,389 The LSPR of these Ag or Au colloidal monolayers can be measured with a commercially available UV-Vis spectrometer in a simple collinear transmission configuration,5,376,390 and binding events can be monitored

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Figure 15. Schematic of a GNR molecular probe (a). Multiplexing detection of various targets using GNR molecular probes: one target (b), two targets (c), and three targets (c). Reprinted with permission from ref 286. Copyright 2007 American Chemical Society.

Figure 16. Schematic representation of the applied biosensing principle based on Au or Ag NPs deposited on a quartz substrate. From left to right: light source, quartz substrate, mercaptosilane adhesion layer, Au or Ag NPs, self-assembled monolayer of functional thiols, antibodies and antigens. The resulting absorbance spectra increase upon binding of analytes to the NPs is shown on the right-hand side of the figure. Reprinted with permission from ref 385. Copyright 2003 American Chemical Society.

in real time by integrating the functionalized substrate with a flow cell.5,376 Changes in refractive index occurring near the surface of the immobilized NPs (caused, for example, by the adsorption of a protein from a dilute solution) shifts the LSPR peak position and magnitude.5,376,385,391-393 These devices exhibit peak position sensitivities of ∼167, ∼76, and ∼252 nm/RIU for spherical Ag391 and Au NPs5 and Au nanorods394 respectively, and changes in extinction at an off-peak wavelength of ∼0.4-1.2 and ∼0.8 per RIU for spherical Au NPs376 and Au nanorods, respectively.394 This sensitivity can be enhanced using a multireflection attenuated total reflection setup.389 Surface-immobilized spherical NPs have recently been used to detect the binding of antibodies to BSA and human serum albumin (HSA)392 and for the selective detection of phosphopeptides on titania-coated Au NPs in complex samples at nanomolar concentrations.393 The at-

tractive features of this approach to sensing are the relatively easy and low-cost fabrication process and the simple optical setup. This approach has also been used to immobilize NPs on optical fibers,388,395 which have shown the ability to detect streptavidin (using a model biotin-avidin assay) and staphylococcal enterotoxin B (using a model antibody-antigen immunoassay) at picomolar concentrations.29,395 Surfaceimmobilized core shell NPs, such as spherical silica-Au (core-shell, Figure 4) and rice-shaped hematite-Au (core-shell) NPs, have been shown to exhibit bulk refractive index sensitivities of ∼555396 (dipole resonance) and ∼800 nm/RIU,199 respectively. Nanostructures for refractive index sensing have also been formed directly on substrates using NSL,7,337,382,397 colloidal lithography,259,383,398,399 soft nanoimprint lithography,77 or metal thin film evaporation.379,400,401 Arrays of nanostructures offer the advantage of tunability of the wavelength response.

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For example, arrays of triangular Ag NPs formed on substrates by NSL (Figure 5) exhibit an LSPR extinction band that can be tuned from the near-UV to the mid-IR by simply changing the size and shape of the NPs.257 The peak position λmax is sensitive to the local dielectric environment and red shifts linearly with increasing solvent refractive index with a sensitivity of ∼200 nm/RIU.397 The distance dependence of the LSPR resonance of these triangular NPs has been systematically studied using SAMs,402 layer-by-layer self-assembly,403 and atomic layer deposition.404 These studies provided several important insights into applicationrelevant parameters including (i) the sensing volume can be systematically tuned by controlling the composition, size, and shape of the NPs, (ii) the sensor response is linear with respect to analyte coverage/quantity when the binding occurs at short distances from the surface of the NPs, and (iii) the sensor response varies in a complex, nonlinear fashion when binding occurs at large distances from the surface of the NPs.403 The refractive index sensitivity of triangular Ag NP arrays has been exploited in sensing applications where peak shifts were used to monitor binding events at the surface of the NPs.7,92,243,382,402,403 The measured peak shifts were used in conjunction with a simple mathematical formalism developed for flat film SPR spectroscopy to quantitatively determine target analyte concentrations.405 This allowed the quantitative detection of Concanavalin A,382 streptavidin,7 and antibiotin243 as well as the thermodynamic evaluation of their binding constants. In a recent and notable example, triangular NP arrays were used to detect biomarkers for Alzheimer’s disease in both synthetic and human patient samples.92 In this work, synthetic amyloid-β-derived diffusible ligands (ADDL) were detected at femtomolar concentrations using a sandwich assay.92 The sensitivity of these types of measurements is enhanced if the molecular resonance of the analyte overlaps with the intrinsic LSPR of the NPs or if the analyte is labeled with a marker that has a resonance that overlaps with the intrinsic LSPR of the NPs.406,407 One of the advantages of using plasmonic nanostructures for sensing is their relatively small footprintsone that is more amenable to miniaturization than flat film SPR detection. For example, LSPR sensing has been demonstrated at the single-NP level using spherical (Au343 and Ag408), triangular (Ag),285,408,409 disk-like (Au),410 and cubic (Ag) NPs.342 Scattering-based spectroscopies must be used to characterize the optical properties of single NPs103,285,343,408-412 since the absorbance of individual NPs is close to the shot noisegoverned limit of detection.285,337 As an example, McFarland et al. have shown that a LSPR peak shift of ∼40 nm occurs upon binding of ∼100 zeptomoles of 1-hexadecanethiol to a single triangular Ag NP (as measured by resonant Rayleigh scattering spectroscopy). This high sensitivity and small transducer size suggests that single NPs could be useful for the analysis of precious or limited-volume samples. Rubinstein developed the so-called transmission LSPR (T-LSPR) spectroscopy method.379 In this protocol one measures the changes that occur in the extinction band of the LSPRs of discontinuous Au or Ag films upon analyte binding, which is monitored in transmission mode using a standard spectrophotometer. Discontinuous and random island films for this plasmonic measurement are prepared by direct evaporation of an ultrathin (e10 nm nominal thickness) layer of the desired metal onto a transparent substrate like quartz, mica, or polystyrene. Random Au island

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films, for example, display a SP extinction peak at 550800 nmsone whose shape, intensity, and position depend on the island morphology, which in turn is determined by the evaporation conditions and postdeposition treatment. Rubinstein and co-workers demonstrated the potential of this method by measuring changes in the position and intensity of the SP extinction band that result from the binding of various molecules to Au islands.379,400,401,413 A linear relationship was shown to exist between the surface coverage of the adsorbing molecules, ones bound either directly to the Au or through a receptor layer, and the plasmon intensity or wavelength changes.379,401,413 In an interesting modification of typical RI-based measurement methods, Au island films were modified with a biotinylated monolayer and used to monitor the binding of avidin based on changes in plasmon intensity (rather than the more commonly used wavelength shift) since the binding event caused only a small change in the peak position. T-LSPR spectroscopy was shown to be widely applicable with a sensitivity (under optimized conditions)380 that is comparable to that of conventional forms of SPR sensing.379 The instability of T-LSPR sensors, however, is a source of concern. Changes in the optical properties of metal island films due to morphological changes occurring upon immersion in organic solvents and aqueous solutions has been noted to appreciably introduce uncertainties into NP-based sensing measurements. To obtain metal island films with stable and reproducible optical properties, new design schemes for stabilizing the structures of the evaporated film have been devised. The most useful of these reported to date consists of depositing an ultrathin silica layer (1.5 nm thick) on the metal island film by a sol-gel procedure.381 Au nanorings formed directly on substrates by colloidal lithography (Figure 7) have recently been used for chemical and real-time biosensing.383 These plasmonic nanostructures exhibit a bulk refractive index sensitivity of ∼880 nm/RIU and show a peak shift of ∼5.2 nm per CH2 unit when SAMs of varying alkanethiol chain lengths were formed on the nanorings by chemisorption. This peak shift per CH2 unit (short-range refractive index sensitivity) for the Au nanorings is greater than the peak shifts reported for triangular Ag NPs (∼3.1 nm per CH2 unit) and nanoprisms (∼4.4 nm per CH2 unit).402,409 The greater sensitivity of Ag plasmonic nanostructures compared to Au plasmonic nanostructures suggests that greater sensitivities could be achieved by forming Ag nanorings.403 Real-time, label-free biosensing was demonstrated using these structures by monitoring the optical changes that occur due to the nonspecific binding of biotinBSA to the Ag nanorings followed by the specific binding of NeutrAvidin (NA).

4.4.3. Periodic Nanohole Arrays The enhanced transmission of light through periodic arrays of subwavelength holes in metal films has generated considerable interest since it was first reported by Ebbesen et al. in 1998.75 Nanohole arrays are typically formed by the serial process of the FIB milling of holes in a thin film of Au supported on a transparent substrate. While Ebbesen et al. attributed their enhanced transmission to surface plasmons,75 several authors have pointed out alternatives ranging from waveguide modes to even more novel surface waves.414-416 In actual fact, and depending on the specific details of the plasmonic structures involved, a

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variety of mechanisms can be operative. For this reason, it is necessary to carefully analyze each situation to determine what sorts of resonance and diffractive effects are possible. In most cases involving thin Au films with periodic nanoholes, however, surface plasmons do play a key role.76,273 The features in transmission spectra of nanohole arrays are thought to arise from a combination of LSPRs, Bloch wave SPPs (BW-SPPs), and Wood’s anomalies.76 An approximate relation for the allowed wavelengths of BW-SPPs excited by normal incidence illumination on a square array of subwavelength holes in a metal film is given by76,77,417,418

λ)

P

xnx2 + ny2

Re

x

(λ)med

(λ) + med

(3)

where (λ) is the wavelength-dependent relative dielectric constant of the metal, med is the relative dielectric constant of an adjacent medium, P is the nanohole lattice spacing, and nx and ny are integer scattering or diffraction orders from the array. Another important equation in analyzing the features of transmission spectra is that for the positions of Wood’s anomalies (WAs)75,76

λ)

Pxmed

xnx2 + ny2

(4)

which is written here for the case of normal incidence for simplicity. The WAs in this case may be thought of as light moving parallel to the metal surface, i.e., light that is diffracted 90° relative to the normal incident direction. Note that in one classification scheme of WAs this particular WA is often referred to, more precisely, as a Rayleigh anomaly.419 Equations 3 and 4 predict a variety of discrete wavelength positions for the BW-SPPs and WAs. Originally a correlation between transmission minima and WAs and transmission maxima and BW-SPPs was noted;386,417 however, this may not always be the case. At a naı¨ve level, for example, one might associate both the BW-SPP and WA positions with transmission minima, not maxima, because they represent wavelengths where incident light is channeled into light moving completely counter to the direction of the transmitted light. Indeed, careful numerical studies made in the absence of experimental imperfections and uncertainties have shown a definite correlation between transmission minima and eqs 3 and 4.76 It remains a fact, however, that nonradiating BW-SPPs and WAs are an idealization. One can regard the BW-SPPs or WAs as being analogous to zero-order quantum mechanical bound states embedded in a zero-order continuum. In the present case a zero-order continuum state would be light that is predicted to be transmitted by the hole/ film structure disregarding the possibility of zero-order bound states. The full problem is described by a coupling of the zero-order bound and continuum states. A Fano resonance line shape (generally an asymmetric line shape with a distinct minimum and maximum) can often describe each BW-SPP or WA feature very well.76,419,420 It often turns out that the same order BW-SPP and WA have wavelengths that are close to one another, which can complicate this picture. Coupled with the fact that the WA features can be narrow and difficult to resolve experimentally, it sometimes follows that only a minimum/maximum resonance feature associated with a particular BW-SPP is seen, along with possibly a dip or kink in the region where the WA would be ex-

pected.76,421 Finally, we note that the features discussed here can occur alongside or be superimposed on generally broader LSPR resonance features. With all these factors taken into account and playing varying roles depending on the specific details of the experiment, it is perhaps not surprising that there has been some confusion in the literature regarding the mechanisms of enhanced transmission. The involvement of LSPRs, BW-SPPs, and WAs in the optical response of nanohole arrays suggests that the features in the transmission spectra should be sensitive to changes in the local refractive index, making these nanostructures useful for sensing applications. Indeed, the spectral features of 2D Au nanohole arrays have been shown to depend on the external dielectric environment,14,422,423 which has been used to perform label-free detection of SAM formation on Au followed by subsequent adsorption of bovine serum albumin (BSA).14 The sensitivity of these structures to changes in the external dielectric environmentsmeasured as peak shift per RIUsis ∼333-400 nm/RIU.14,377 The advantages of plasmonic nanostructures over flat film SPR systems are their relatively small footprints (i.e., small patterned areas on the order of micrometers14,377 or single holes259 or NPs285,342) and simple optical setups (i.e., normal incidence transmission or reflection).14,342,377 These advantages were recently exploited by integrating a microfluidic device with a nanohole array for detecting refractive index changes and surface binding events ‘on-chip’.377 Plasmonic quasi-3D periodic nanohole arrays have been formed on surfaces using a type of soft nanoimprint lithography (see section 3.3.3 and Figure 9).8,77 These multilayered structures consist of arrays of nanoscale holes in Au films with a second level of Au disks at the bottom of the embossed wells. The normal incidence transmission spectrum of these structures is complex (Figure 17a, blue spectrum) and shows high transmission despite the fact that no ‘line of sight’ exists through the sample (i.e., unlike 2D nanohole arrays, the cylindrical holes or perforations are ‘capped’ with Au disks in a quasi-3D geometry). Rigorous 3D finite-difference time domain (FDTD) simulations were performed to assist in the interpretation of the experimental results (Figure 17a). (See also refs 76 and 424 for related FDTD studies of hole arrays.) The calculated electromagnetic field distributions in and around the metal nanostructures are shown in Figure 17b and c. The two largest features in the spectra (Figure 17a) are labeled B and C. Peak B is associated with excitations of LSPRs on the rims of the nanoholes in the upper Au film near the air/Au interface, while peak C involves overlapping Wood’s anomaly and BW-SPPs excitations on the Au disk/polymer side of the device. The intensity associated with peak C also extends vertically up to the hole opening, showing a strong coupling between disk and hole, a feature that is absent in random and ordered 2D nanohole arrays. Quantitative modeling of the experimental spectra required consideration of fine structural details in this region of strong coupling. Good agreement between the experimental and theoretical spectra required the addition of small (20-30 nm), isolated grains of Au on the sidewalls of the nanoholes, just above the edges of the Au disks at the bottom of the nanowells (Figure 17a, red spectrum). The transmission properties of the quasi-3D nanohole array are sensitive to the nature of the adjacent dielectric medium at the crystal surface (Figure 18a). The bulk refractive index sensitivity of these devices was determined by passing

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difference spectra, as referenced to the spectrum at time t ) 0, illustrate changes in transmission due to both peak shifts and intensity changes throughout the wavelength range as solutions of increasing PEG concentration were injected into the flow cell. As shown in Figure 18b, the transmission can increase or decrease depending on the measurement wavelength. The particular crystal geometry used in this study (hole diameter and periodicity of ∼480 and 780 nm, respectively) exhibited the largest multiwavelength response in the near-infrared region (900-1250 nm) as shown in Figure 18c. The response over all wavelengths, including both positive and negative transmission changes, was calculated using the following equation

R ) ∫|∆(%T(λ))dλ

Figure 17. Correlation of transmission spectral features with hole/ disk plasmonic excitations. (a) Normal incidence transmission spectrum of a quasi-3D plasmonic crystal (blue), and rigorous electrodynamics modeling of the spectrum for an ideal crystal (green) and one that includes subtle isolated nanoscale grains of Au near the edges of the Au disks (red). (b) Computed electromagnetic field distribution associated with the resonance at 883 nm (labeled B in a). The intensity is concentrated at the edges of the nanoholes in the upper level of the crystal. (c) Field distribution associated with the resonance at 1138 nm (labeled C in a), showing strong coupling between the upper and lower levels of the crystal. Reprinted with permission from ref 77. Copyright 2006 The National Academy of Sciences of the USA.

solutions of increasing concentration (0-7.6 wt %) of polyethylene glycol (PEG) through a fluid flow cell containing a plasmonic crystal. Changes in both peak positions and intensities were observed over a wide spectral range as the refractive index of the PEG solution was increased. The most sensitive peak at ∼1023 nm linearly red shifted with a sensitivity of ∼700-800 nm/RIU and linearly increased in intensity with a sensitivity of ∼2.5-3.5 Abs/RIU as the refractive index of the solution was increased (determined using several devices).77 SPR sensing is typically performed by following the response of an individual peak or wavelength to binding events at the surface of the sensor. This method of analysis, however, does not fully capture the sensitivity of the quasi3D plasmonic sensor since it does not take advantage of all the peak shifts and intensity changes occurring at multiple plasmonic resonances over the spectral range created by the coherent couplings of the LSPRs, BW-SPPs, and Wood’s anomaly responses exhibited by this type of system. This wide spectral response was exploited using a type of full, multispectral analysis as described below. Multispectral analysis of a PEG calibration of a quasi-3D plasmonic crystal is shown in Figure 18a. A series of

(5)

This total or integrated response, R, has units of ∆%T‚nm (Figure 18d) and changes linearly with the refractive index of the PEG solution with a sensitivity of ∼22,000 ∆%T‚ nm/RIU (inset of Figure 18d). This multispectral analysis also improves the signal-to-noise ratio by a factor of 3-10 times that of a single-wavelength response.77 The linear response of the integrated metric was used to perform quantitative sensing and imaging of binding events. The biotin-avidin ligand-receptor conjugate was used to illustrate the utility of these devices for performing quantitative sensing (Figure 19). The device was first exposed to a solution of biotinylated BSA (bBSA), which led to an increase and plateau of the integrated response of the sensor (Figure 19b) upon formation of a bBSA monolayer. This layer rendered the surface of the sensor inert to further nonspecific adsorption, which was demonstrated by the lack of response after rinsing the bBSA monolayer with buffer and exposing it to a solution of nonfunctionalized BSA. Subsequent exposure to avidin, however, led to a response due to a specific binding interaction between the avidin and the initial bBSA monolayer. The surface-immobilized avidin was then used to complete the assay by binding a layer of bBSA to the remaining free biotin binding sites on the avidin layer (inset Figure 19b). This resulted in a response that was smaller than that observed for the initial bBSA adsorption step, an observation that follows the patterns of layerdependent mass coverage generated in assays of this sort.77 The integrated response can be converted to an effective protein thickness using a mathematical formalism.405 Although this model was developed for quantifying binding events measured using flat film SPP-based SPR sensors (i.e., it assumes a uniform plasmon evanescent field), it provides approximate protein coverages that agree with literature values.425-427 One notable advantage of soft nanoimprint lithography is the ability to pattern over large areas in parallel with high spatial uniformity and low defect densities, which facilitates large-area imaging for multiplexed microarray-based analyses. The high-quality, large-area patterning capability along with the capacities for quantitative biosensing was combined to perform quantitative imaging. To this end, five lines of nonspecifically adsorbed fibrinogen were patterned on the surface of a plasmonic crystal using a microfluidic device (Figure 20a). Figure 20b shows changes in transmission measured relative to an interchannel region on the crystal that did not come in contact with the protein. The spectral image shows five stripes with the expected geometries, each corresponding to a line of nonspecifically adsorbed fibrino-

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Figure 18. Optical response of a plasmonic crystal to sequential injections of increasing concentrations of aqueous PEG solutions. The color contour plot of the change in transmission (T) as a function of wavelength and time (with the corresponding injection sequence overlaid on the plot) (a), change in T as a function of time during the injection sequence, evaluated at several wavelengths (b), absolute value of the change in T as a function of wavelength evaluated at different times (c), and integrated multispectral plasmonic response as a function of time (d). (Inset) A linear correlation to the change in refractive index. Reprinted with permission from ref 77. Copyright 2006 The National Academy of Sciences of the USA.

Figure 19. Plasmonic crystal used in a biotin-avidin assay. The color contour plot of the change in T as a function of wavelength and time (a). The overlaid injection sequence corresponds to PBS (1), bBSA (2), BSA (3), and avidin (4). The integrated multispectral plasmonic response and corresponding effective thickness of the biotin-avidin-biotin assay (schematically illustrated in the upper inset) (b). The noise limited refractive index resolution of the crystals corresponds to submonolayer coverages (lower inset). Reprinted with permission from ref 77. Copyright 2006 The National Academy of Sciences of the USA.

gen. Analysis of the step edges (Figure 20c, inset) shows a width of ∼20 µm, which is only slightly larger than the ∼17 µm resolution limit of the imaging optics. The ∼3 µm of additional width in the plasmonic image can be associated with the propagation lengths of plasmons on these structures.428 The integrated spatial response of the spectral image could be converted to an effective protein thickness of ∼7 nm using the PEG calibration and the mathematical model described above.77,405 This thickness is consistent with the molecular dimensions of fibrinogen.429,430 The micrometer-

Figure 20. Spatial imaging of fibrinogen nonspecifically adsorbed to the surface of a crystal. A schematic illustrating the use of a multichannel PDMS microfluidic network to pattern the surface of a crystal (shown here with the multicolored appearance that characterizes these crystals) (a), spectroscopic difference image of fibrinogen lines patterned on a crystal (b), and spatially resolved integrated response and corresponding effective thickness illustrating binding events in the geometry of the microfluidic channels (c). (Inset) A measured step edge between a fibrinogen line and bare area of the crystal (blue symbols) and a fitted step edge with a Gaussian width of ∼20 µm. Reprinted with permission from ref 77. Copyright 2006 The National Academy of Sciences of the USA.

scale imaging resolution and the large area defect-free (and thus uniform response) aspects of the crystals suggest a promising platform for performing parallel diagnostic bioassays. This was further demonstrated by performing quantitative 2D imaging on these quasi-3D devices.279

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4.4.4. Random Nanohole Arrays Random nanohole arrays formed by colloidal lithography have been used in enabling ways for the sensing of chemical and biological targets.90,259,398,399 Randomly arranged holes do not exhibit the long-range higher-order diffractive coupling effects that are observed in the highly ordered, periodic nanohole arrays discussed above.72,75,90,422 They instead exhibit LSPRs that are concentrated at the edges of the holes with decay lengths on the order of 10-20 nm, which have been used for sensing binding events occurring near or in the holes.90,259,398,399 The refractive index sensitivity of these random nanohole arrays depends on the metal thickness398 and the detection modality. As noted above, refractive index sensitivities are often reported as a change in a given measurable parameter per RIU. Sensitivities of ∼100-270 nm/RIU,72,90,398,399 0.23-1 Abs/RIU,398,399 and ∼70 nm/ RIU259 have been reported for measurement of changes in the extinction peak position, extinction peak magnitude, and scattering peak position, respectively. In one notable example of biosensing carried out with such devices, Dahlin et al. used random nanohole arrays to detect membrane-mediated recognition of bioanalytes using surface-immobilized phospholipid bilayers (SPBs). These SPBs were supported on the silica surface within the interior of the nanoholes (Figure 21). This was accomplished by first exposing the nanohole array to bBSA, which preferentially adsorbs to the Au rather than the SiO2,431 making it inert to lipid vesicle fusion. The SPBs were then formed on the SiO2 substrate at the bottom of the nanoholes by performing vesicle adsorption in the presence of Ca2+, which promotes the vesicle fusion process.432 The top of Figure 21a shows schematically the steps leading to the adsorption of the protein NA within the nanoholes containing the biotin-modified lipids (left). In the absence of biotin-modified lipids the NA binds only to the bBSA on the Au surface, whereas in the presence of biotinmodified lipids the NA binds to both the SPB and the Au surface. The bottom panel of Figure 21 shows the measured changes in extinction as a function of time at ∼725 nm for the two cases, with the signal being more than a factor of 3 greater when NA binds to lipids inside the holes. This demonstrates that the holes are highly sensitive regions for detecting protein-binding events. The top of Figure 21b shows schematics illustrating the binding of cholera toxin to SPBs containing ganglioside GM1 glycolipids (left) and hybridization of 15-base single-stranded DNA (ssDNA) to SPBs presenting a complementary strand (right). The bottom of Figure 21b shows the measured changes in extinction as a function of time at ∼725 nm upon binding of the cholera toxin and complementary and noncomplementary ssDNA. These examples demonstrate the potential for using random nanohole arrays as a platform for real-time plasmonic-based label-free sensing. The figure of merit (FOM) for this system appears to be quite good in that analytically discriminable signals could be obtained from zeptomole quantities of protein. Related nanohole arrays also have been used to detect cancer biomarkers where the detected signal (peak shift) was estimated to arise from the binding of picograms of an antigen (a specific tumor biomarker, cancer antigen 19-9) to its corresponding surface-immobilized antibody.90

4.5. Surface-Enhanced Spectroscopies The optical techniques discussed thus far are amenable to chemical functionalizations that engender capacities for molecule-specific sensing, ones that the methods alone lack.

Figure 21. Detection of lipid-membrane-mediated binding events using SPB in random arrays of gold nanohole. (a) Temporal variation in extinction measured at the longer wavelength inflection point (725 nm) of the LSPR peak upon addition of ∼0.3 µM NA. In both cases, Au is modified with biotin-BSA. In one case, biotinmodified SPB patches cover the SiO2 regions (NAAu+SiO2), while in the other (top right illustration in a), unmodified SPB patches cover the SiO2 regions (NAAu) (top left illustration in a). (b) Variation in extinction measured upon addition of ∼0.5µM cholera toxin (CT) to GM1-modified (5 wt %) SPBs (purple curve, illustration of SPB-modified nanohole shown in upper left of b). (Inset) Magnification of changes in extinction versus time upon addition of a 15 base long noncomplementary (0.2 µM, blue) and a fully complementary strand (0.2 µM, red) to SPB patches modified with a DNA construct carrying two cholesterol moieties at its one end and a 15-base-long single strand available for hybridization at the other. Reprinted with permission from ref 399. Copyright 2005 American Chemical Society.

It is useful then to consider cases where other properties besides explicit forms of chemical recognition can be used to discriminate a molecularly specific event. Techniques such as Raman scattering and fluorescence have the capability of providing molecule-specific data, and perhaps most interesting in the context of this review, surface plasmons are able to markedly enhance the sensitivity of these techniques. The utility of Raman spectroscopy, for example, stems from its ability to probe molecular vibrations, but the method suffers from relatively weak signal intensity. In most cases the number of inelastically scattered photons (which directly corresponds with the signal level) is extremely small, corresponding to scattering cross sections of 10-30-10-25 cm2.22 As a result, even though Raman scattering is a type of spectroscopy that can give very specific molecular information, it lacks the intrinsic sensitivity required of a viable technique for high-throughput detection. Plasmonic architectures provide one general method for enhancing Raman signals to levels that can enable many analytical applications, ones ranging from DNA sequencing to forensics.337,433,434 The sections below examine several of the more interesting prospects emerging from current research.

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4.5.1. Surface-Enhanced Raman Scattering The first discovery of surface-enhanced Raman scattering (SERS) showed that it was possible for a roughened noblemetal surface to enhance the intensity and sensitivity of Raman spectroscopy.435,436 Using the SERS effect in a quantitative form is intricately tied to the ability to fabricate nanostructured surfaces with well-defined morphologies. New methods to create nanostructured materials are constantly evolving, and research centered around SERS has followed that trendswork that also has increased the sensitivity to levels required for advanced sensing applications. Signal enhancements of up to 1014 over normal Raman scattering have been observed in special cases,437 a value sufficient to study analytes even at the single-molecule level. The electromagnetic theory of SERS437-440 suggests that the enhancement effect is a result of the creation of surface plasmons that can transfer energy to the bound molecule through the associated electric field. We direct the interested reader to a number of reviews that have already been published on this subject.437-440 The ability to enhance the Raman signal via SERS is fundamentally linked to the precise details of the structured nature of the surface as well as the choice of the metal itself, which includes but is not limited to Ag, Au, and Cu.438 These metals are used because they have the appropriate values of both the real and imaginary parts of the dielectric constant, which allows surface plasmons to propagate at wavelengths of interest for spectroscopic applications. As the discussions above have revealed, the plasmonic excitations that a system can support depend very sensitively on the metal structure, a factor that in turn drastically affects its overall ability to enhance the Raman scattering cross section. The advances made in nanomaterial synthesis and NP fabrication have greatly empowered competencies in this latter regard. Such forms of modern SERS substrates span a vast geometrical range encompassing colloids, templated colloidal crystal films, electrochemically roughened electrodes, deposited metal island films, nanohole arrays, and lithographically defined thin films. The ability to use these different structures as SERS substrates allows a variety of experimental setups including solution- and surface arraybased sensing.435,441-446 Solution-based SERS has been shown to provide a viable option for the determination of binding constants and monitoring DNA/RNA mononucleotide recognition.447 For the latter, detection of mononucleotides at a submicromolar level (3 ppb) was achieved for 2′-deoxyadenosine 5′-monophosphate (dAMP) in the presence of MgSO4, an aggregating agent (Figure 22).447 The data indicates that it is essential for the analyte to be bound within close proximity to the NPs. Solution-based sensing with NPs has also been applied to the field of cellular biology.448 In a notable example taken from recent research, it was demonstrated that it is possible to introduce 60 nm Au NPs into a single osteosarcoma cell and use their optical responses to map the distribution of the cellular constituents.449 It is possible to further enhance the SERS signal of an analyte by altering the physical properties of the NPs.450-452 For example, using Au octahedrons instead of Au nanospheres allows for a more than 3× greater enhancement in the spectrum of 2-napthalenethiol.450 Substrate-bound nanostructures can also be used to enhance Raman signals.441 Nanosphere lithography, for example, has been used to fabricate two intriguing forms of surface-immobilized structures for SERS: (i) triangular NP

Figure 22. Effect of aggregating agents to the SERS signal of dAMP: 1000 ppm dAMP mixed with Ag colloid (a); mixture after aggregation with 0.1 M MgCl2 (b); 0.1 ppm dAMP mixed with MgSO4-aggregated (0.1 M) Ag colloid (c). Spectra d, e, and f used the same conditions as c except 0.03, 0.01, and 0.003 ppm dAMP. The inset shows the calibration plot of the dAMP (4 s accumulation times). Reprinted with permission for ref 447. Copyright 2006 American Chemical Society.

Figure 23. Ambient contact-mode atomic force microscope image of 200 nm Ag over 542 nm diameter polystyrene spheres. Array of spheres (10 µm × 10 µm) and image (600 nm × 600 nm) of one sphere showing substructure roughness (a and b, respectively). AgFON electrode SER spectrum of 50 mM pyridine in 0.1 M KCl at -0.7 V vs Ag/AgCl (c). Reprinted with permission from ref 453. Copyright 2002 American Chemical Society.

arrays (Figure 5) and (ii) metal film over nanosphere (MFON) surfaces (Figure 23). Triangular NP arrays have been shown to be quite sensitive with SERS enhancement factors of ∼108.337 Analytes detected in this way include Alzheimer’s precursors,93 glucose,433 and Concanavalin

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A.382 The highly corrugated MFON surfaces have been used to detect a key component of an analogue to anthrax stimulant at a concentration lower than the infectious dosage.242,445 The MFON substrates are also useful as electrodes with an improvement in stability and reproducibility over other types of SERS active electrodes.453 The latter are critical parameters since most SERS active substrates are not particularly stable. It is notable, therefore, that recent work has shown that a sub-1 nm thick alumina film deposited over a MFON surface can extend the stability of the sensor up to seven times, corresponding to a shelf life of at least nine months.242 Electrochemistry can be used to roughen an electrode surface for SERS applications, but this method generally yields substrates with irreproducible enhancement factors.453 An alternative method that can be used to form SERS-active electrode structures is to use a template to create nanowires. As an example, Au nanopillar films can be formed by physical/electrodeposition of Au in an anodized aluminum oxide (AAO) template followed by dissolution of the AAO template.454 These nanopillar substrates are highly reproducible (to within 20% of the SERS intensity profile) and capable of detecting thionine as a model analyte at concentrations near 10-8 M.454 Nanopillars of Ag can also be fabricated in the form of aligned pillars of controllable length, giving SERS enhancement factors on the order of 108.455 Nanodots,443 nanohole arrays,444 and even nanostructured multicore optical fibers456 have been fabricated for use in SERS applicationsswork that evidence a broadening research landscape in response to requirements for SERS substrates that are both stable and ultrasensitive. 4.5.1.1. Single-Molecule Surface-Enhanced Raman Scattering. Highly specific single-molecule detection is a challenging goal for chemical sensing.22,457,458 It is interesting, therefore, to consider how the marked enhancements afforded by SERS compare to this most demanding analytical FOM. Kniepp and co-workers elucidated the fact that singlemolecule SERS (SMSERS) detection levels can be achieved both in solution and on surfaces, the latter using patterned metal substrates. Each protocol requires the use of a proximal probe in which incident light is restricted to encompass a probe volume of a few femto- to picoliters. This makes it possible to examine limited volume samples with analyte concentrations as low as 10-12-10-14 M. Solution-based SMSERS, for example, has been used to detect pseudoisocyanine (PIC) in an aqueous solution at a concentration of 10-14 M, corresponding to an enhancement factor on the order of 1014.22 Ag colloids also have been used to detect yeast cytochrome c at near single-molecule levels.459 Patterned Ag surfaces have been shown to give enhancements that make it possible to detect single molecules of enkephalin, in this case by monitoring the ring-breathing mode of phenylalanine at 1000 cm-1.22 The small sample probe volume has other (potentially beneficial) impacts as diffusion brings single molecules into and then out of the probed volume at very low analyte concentrations due to Brownian motion. This leads to a variation in the Raman signal that can be correlated to the number of molecules in the probed volume via models based on Poisson statistics.434

4.6. Plasmonics for Detection Beyond the Diffraction Limit One limitation of SMSERS (or for that matter any optical technique) is that the spatial resolution of the probed area is

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dictated by the spot size of the incident light. Optical detection methods are constrained by the wave nature of light (which determines the diffraction limit); however, it is particularly interesting that plasmonic nanostructures make it possible to probe systems with a spatial resolution that is well beyond the diffraction limit of light. Two methods that use the enhanced local electric fields associated with plasmons for detection at length scales beyond the diffraction limit are apertureless near-field optical microscopy (aNSOM) and tip-enhanced Raman spectroscopy (TERS). Both of these methods have enabled the scientific community to optically probe surfaces with resolutions previously only afforded by techniques such as electron microscopy and surface probe microscopy (SPM). Near field optical microscopy (NSOM), with a typical aperture, is a technique that affords capabilities of sub-100 nm resolution; however, it is important to note that the transmission of light through the fiber tip is typically quite low. One way to improve the resolution and limited transmission is to move to a fully metallized tip that has no aperture (a-NSOM). The apertureless tip is then capable of supporting surface plasmons that lead to enhancement factors of the electric field of 10-1000 times, and the ability to sharpen the tip affords resolution limits on the order of 10 nm.460 For example, tips coated with 15-25 nm of Ag were used to detect Alexa 532 dye molecules with a resolution of 15 nm.461 The ability to garner optical images from nearfield microscopy is a useful technique for imaging nanostructures with sub-100 nm resolution but lacks the specific chemical information afforded by techniques such as TERS. The basis of the SERS effect is tied to the nanoscaled structures present on a metal surface, and it is interesting to note that the SERS effect also can be observed using a metallized SPM tip. The surface of the probe tip must be coated with a metal (e.g., Ag, Au, Cu) that exhibits strong SERS enhancements. In this case the Raman scattering is constrained to within a few nanometers of the tip.462 The enhancement factors realized to date (102-104) are still moderately low and considerably less than the predicted maximum values (1010).462 One particularly interesting aspect of the advances coming from this research is the ability to combine the spatial resolution of SPM with the molecule-specific detection afforded by SERS. An example of this is illustrated by the data shown in Figure 24. In this case the nanoscale roughness, achieved with a 2 nm abrupt step on the surface of a metal-coated sharp TERS tip, can lead to an increase in Raman scattering cross section over a standard proximal probe by over an order of magnitude for benzenethiol adsorbed on Au.463 The tips of silicon nanowires have also been modified with Au droplets to improve detection limits.464 The ability to create tips of varying geometrical parameters allows the spatial resolution of the technique to be optimized. These methods also can be used to characterize features of crystalline materials such as GaN with high spatial resolution and enhancement factors as large as 104.465 The ability to break the diffraction limit while gaining chemically specific information constitutes an area of particular opportunity for progress in research, one whose progress will be dictated directly by capabilities to create better tips and to optimize plasmonic excitation and correlated collection optics.462

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are attracting growing attention in research.468,472,476 In one specific example, Au nanohole arrays exhibited an enhancement of up to 2 times over flat metal films when the fluorophores were beyond the inherent quenching distance.468 Surface-immobilized Ag NP arrays have also been used as a medium to enhance the fluorescence of common dyes with leading work directed to determine how the geometry of the NP films affects the enhancement factor.215 It was found that the observed enhancements (of some 10-20 times) were dependent on the width and height of the particles with fluorescein and Cy3 exhibiting a maximum enhancement for particle sizes in the range of 85-95 nm.215 We believe that the ability to enhance fluorescence through the use of plasmons is an area of research that is not yet mature and will continue to grow in importance and impact as the method is optimized. Figure 24. TERS mapping on a rough Au surface. An STM image of the sample is shown in a. TERS data was collected at the positions indicated by the arrows. The cross section of the topography image is shown in b, and the TERS collection sites are labeled with crosses. (c) Corresponding TERS sequence. The numbers denote the sites where the spectra were collected. Reprinted with permission from ref 463. Copyright 2007 American Chemical Society.

4.6.1. Plasmon-Enhanced Fluorescence Plasmon-enhanced fluorescence (PEF) is a relatively new technique that can be used to increase sensitivity and detection limits. This is a capability that involves integrating the measurement within a device form factor that can exploit the enhanced electrical fields associated with surface plasmons. Fluorescence microscopy and spectroscopy are perhaps the most widely adopted methods used to study complex biomolecular systems.457,466-468 The use of surface plasmon modes to enhance fluorescence has recently attracted considerable interest as a way to extend the already considerable FOM of this spectroscopic measurement.215,356,469,470 Fluorophore quenching occurs in close proximity to metallic surfaces (typically < 50 Å);471 however, at longer distances ranging from 5 to 200 nm an interesting coupling between the fluorophore and the enhanced local field near the metal structure occurs, resulting in both increased absorption cross section and radiative decay rates.470 The increase of the radiative decay rate is central to a number of other phenomena observed in PEF including increased photostability, enhanced wavelength-tunable emission, and the ability to probe fluorescence resonance energy transfer (FRET) at larger distances than previously possible.470 These enhancements have opened new opportunities to improve detection and imaging schemes. A number of different metallic structures such as nanohole arrays, NPs, flat thin films, and deposited nanostructured films have been used for PEF.215,472-474 Au NPs, for example, have been used to sense mouse IgG at concentrations as low as 7 fM using a sandwich assay format coupled to an optical fiber.469 Sensing of metal ions such as Cu2+ has also been accomplished using fluorophore-labeled Au NPs with a detection limit of 1 µM.475 This particular sensing scheme has advantages over other methods because it shows increased quantum yields even in the presence of Cu2+, which generally quenches fluorescence. In the same vein, Au NPs have also been used to detect the presence of Hg2+ in pond water at concentrations as low as 2.0 ppb.473 Nanohole arrays can also be used for PEF detection, and such applications

5. Concluding Remarks The rich literature summarized in this review develops a compelling story about both the health of and the prospect for technological impacts following from research in the interdisciplinary field of plasmonics. Whether in the form of refractive-index-based detection schemes for analytes of extremely low concentrations or as optics for chemically sensitive imaging, the reports appearing from many laboratories at this time demonstrate an accelerating pace of underlying progress, one driven by rapidly improving capacities for nanoscale materials synthesis and methods of fabrication. An interesting question to ask at this point is one related to context. Specifically, will the most important impact going forward with respect to chemical analysis come from the development of systems that embed “new physics” (such as those emerging within the vastly interesting topic of metamaterials)477-480 or from compelling applications of systems that exploit what is currently known? Our hope is that the answer comes from both. There is no doubt that the rapid progress being made in both the fabrication and theory will greatly enhance progress in the field to render the enabling science in a form that begins to become a true form of predictive design for function. Essentially all applications in sensing will benefit from this advent. Still though, what might happen in the context of “new physics” and the opportunities it might engender? An interesting (and nascent) model here is provided by an area only touched on briefly in this reviewsuse of plasmonics as devices for subwavelength “imaging”. It is well appreciated that the spatial resolution in optical microscopy is limited to λ/2, according to the Rayleigh criterion. This limit is caused by the loss of the evanescent field intensity in the far field, which carries high spatial frequency information. Complex methods can overcome this limit and have been devised in such notable forms as nearfield scanning optical microscopy.460 The complexity of the methods used there carry with them limitations that have tended to limit its widespread use for biological imaging, especially cellular imaging. Metal nanohole arrays have been shown to inherently give anomalous high transmission of light,75 and this, in conjunction with “a perfect lens” based on negative index metamaterials,481 is theorized to make possible subwavelength resolution microscopy down to limits as small as 25 nm. Recent reports have demonstrated482-485 that a surface plasmon’s near-field dimensions are smaller than free-space radiation at the same frequency. Srituravanich et al. examined

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this effect by exposing photoresist through a nanohole array, obtaining features with sizes of 90 nm spaced by 170 nm.485 Fields with dimensions as small as 25 nm were obtained using slightly more complex plasmonic structures.486 It is therefore most intriguing that wide field optical microscopy with subwavelength resolution, based on the concepts of plasmon-coupled transmission and optical near fields, has been proposed along with methods to implement it.487-489 More recently, two different types of magnifying metamaterials-based superlenses that can be integrated into a conventional far-field optical microscope were reported.490,491 Liu et al. demonstrated a magnifying optical hyperlens consisting of a curved periodic stack of Ag (35 nm) and Al2O3 (35 nm) deposited on a half-cylindrical cavity in quartz.491 This setup was able to image a pair of 35 nm lines spaced 150 nm apart. In a later report, Smolyaninov et al. demonstrated a superlens design based on a multilayer photonic metamaterial consisting of alternating layers of positive and negative refractive index.490 With this design a resolution of 70 nm was reported. These are advances that serve to illustrate the inspirational opportunities for progress coming from work in this field.

6. Acknowledgments We gratefully acknowledge the support of aspects of our work by the National Science Foundation (CHE 04-02420) and the Department of Energy (DEFG02-91ER45439). S.K.G. was supported by the Office of Basic Energy Sciences, Division of Chemical Sciences, Geosciences, and Biosciences, U.S. Department of Energy (DE-AC0206CH11357).

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CR068126N

Chem. Rev. 2008, 108, 522−542

522

Microcantilevers: Sensing Chemical Interactions via Mechanical Motion Karen M. Goeders,† Jonathan S. Colton,‡ and Lawrence A. Bottomley*,† School of Chemistry & Biochemistry and George W. Woodruff School of Mechanical Engineering, Georgia Institute of Technology, Atlanta, Georgia 30332 Received September 21, 2007

Contents 1. Introduction 2. Theory of Operation 2.1. Static Deflection 2.2. Dynamic Response 2.2.1. Mass−Spring−Dashpot System. 2.2.2. Quality Factor 2.2.3. Plane Strain 2.2.4. Bending Mode Frequency Response 2.2.5. Effect of Damping Due to Viscous Fluids 2.2.6. Effect of Air Damping on Quality Factor 2.3. Lateral and Torsional Mode Frequency Responses 2.4. Temperature Effects 2.4.1. Effect on Material Properties 2.4.2. Effect on Geometry 3. Detection Schemes 3.1. Optical Lever 3.2. Interferometer 3.3. Piezoresistive 3.4. Capacitive 4. Design, Materials, and Fabrication 4.1. Design Considerations 4.2. Fabrication of Silicon-based Cantilevers 4.2.1. Film Deposition 4.2.2. Photolithography 4.2.3. Etching 4.2.4. Doping 4.3. Fabrication of Polymeric Cantilevers 5. Chemical Selectivity 6. Chemical Applications 6.1. Volatile Organics 6.2. Chemical Warfare Agents 6.3. Explosives 6.4. Toxic Metal Ions 7. Biological Applications 7.1. Cells 7.2. Viruses 7.3. Antigen−Antibody Interactions 7.4. DNA Hybridization 7.5. Enzymes 8. Recommendations for Future Work * To whom correspondence should be addressed. [email protected]. † School of Chemistry & Biochemistry. ‡ George W. Woodruff School of Mechanical Engineering.

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8.1. Guidelines for Reporting Sensor Performance 8.2. Experimental Design Considerations 8.3. Fruitful Areas for Further Research 8.3.1. More Selective Coatings 8.3.2. Increased Sensitivity and Faster Response 9. References

538 538 539 539 539 539

1. Introduction Micromechanical devices comprise emerging sensor platforms with straightforward sensing mechanisms. Molecular adsorption onto the sensing element, typically a cantilever, shifts its resonance frequency and changes its surface forces (surface stress). Adsorption onto the sensing element composed of two chemically different surfaces produces a differential stress between the two surfaces and induces bending. The analyte that induces the mechanical response may be physi- or chemisorbed onto the cantilever in a reversible or irreversible process. Devices that respond to chemical stimuli in this manner are more commonly referred to as microcantilever sensors.1-6 A compelling feature of microcantilever sensors is that they can be operated in air, vacuum, or liquid. The rapid growth in microcantilever-based sensor technology parallels advancements in micromachining methodologies and is in response to the need for more sensitive and selective detection of airborne and waterborne toxic and pathogenic substances. The purpose of this review is to critically examine the current state of theory, modes of detection, design considerations, and innovative applications of this sensing platform. Each will be addressed separately in the following sections. At the conclusion of this review, we will identify areas that warrant further investigation and suggest guidelines for reporting the performance of microcantilever sensors to facilitate comparison of microcantilevers with other sensing platforms.

2. Theory of Operation Microcantilever sensors rely on their deflection to indicate sensing. This section describes the theory for the mechanical response of microcantilevers in the bending, lateral, and torsional modes when used as sensors. This discussion is divided into the two modes of microcantilever deflection, static and dynamic, that are used in sensing applications. The means for detecting deflection are discussed in a separate section. The static mode of deflection occurs when an adsorbed species causes differential surface stresses on the oppo-

10.1021/cr0681041 CCC: $71.00 © 2008 American Chemical Society Published on Web 01/30/2008

Sensing Chemical Interactions via Mechanical Motion

Karen Meloy Goeders is pursuing her doctorate in analytical chemistry from the Georgia Institute of Technology. She came to the School of Chemistry & Biochemistry after completing her B.S. degree in chemistry from Louisiana State University. Her dissertation research focuses on developing microcantilever array technology for enzymatic assays.

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Dr. Lawrence A. Bottomley is a professor of chemistry at the Georgia Institute of Technology. He obtained his B.S. in chemistry at California State University, Fullerton, and his Ph.D. in analytical chemistry at the University of Houston. His current research interests include the biological and nanotechnological applications of scanning probe microscopy, electroanalytical chemistry, and microcantilever array sensing.

Figure 1. Microcantilever geometry and nomenclature.

Dr. Jonathan Colton is a professor of mechanical engineering at the Georgia Institute of Technology. He obtained his S.B., S.M., and Ph.D. in Mechanical Engineering at the Massachusetts Institute of Technology. He serves as the director of the Center for Polymer Processing. He is a fellow of the American Society of Mechanical Engineers and of the Society of Plastics Engineers and is a registered professional engineer in the state of Georgia. His research interests include polymer and polymer composites processing, bioMEMS sensors, biomedical devices, and dielectric materials.

site surfaces of the microcantilever. The equations that describe the static deflection of microcantilevers will be presented, and the response of the deflection to surface stress will be discussed. In the dynamic mode of detection, the frequency of vibration of the beam changes as species are adsorbed onto the microcantilever. The equations describing the vibration of the beam in air will be presented. The damping effects of measurement in a viscous gas and liquid will be described next. Thermal effects also will be discussed. A microcantilever can be modeled as a cantilever beam (thickness, t; width, w; and length, L), which is built in (fixed) at one of its ends (see Figure 1).7 Note that, in Figure 1, z denotes the deflection in the thickness direction along the beam length and time [i.e., z(x,T)] and does not indicate the origin of the coordinate system. The discussion first will be limited to pure bending of a beam; lateral and torsional (twisting) motions will be discussed at the end of this section. Figure 2 shows the bending (a), lateral (b), and torsional (c) deflections of a built-in beam.7

Figure 2. Schematic of the first bending (a), lateral (b), and torsional (c) modes of a resonating cantilever. The heavy lines denote the undeformed cantilever; the shaded regions denote the deformed cantilever.

2.1. Static Deflection Static deflection is used to determine the amount of material adsorbed onto a microcantilever. The more material that is adsorbed, the more the microcantilever will deflect.8-10 Deflection results from two mechanisms: added mass and surface stress from adsorbed species.11-13 However, the surface stress may not necessarily correlate with the amount

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Figure 3. End-loaded beam. Figure 5. Fixed beam subjected to surface stresses.

t is the thickness of the coating, and d is the thickness of the beam. The radius of curvature, R, of a microcantilever under the influence of surface stresses on its top and bottom surfaces is typically reported as eq 518 Figure 4. Beam loaded by a uniform load (wo/unit length).

of material adsorbed. The deflection of the free end of the beam depends on the type of loading to which the beam is subjected. If a concentrated load, F, is applied to the free end of a rectangular beam (see Figure 3), then the deflection of the free end of the beam, δ, is given by eq 114

δ)

FL3 3EI

(1)

where E is the Young’s elastic modulus of the beam material, L is the length of beam, and I is the second moment of the beam’s cross-sectional area. I is a function of the beam width, w, and thickness, t, and is equal to wt3/12. The resistance to bending deformation (i.e., stiffness) in bending of a rectangular beam (k) is given by eq 2.15

k)

3EI L3

(2)

An example of this type of deflection is a tipped microcantilever used in atomic force microscopy to measure surface geometry. Equation 2 is commonly used to calculate the force constant of imaging probes. However, since the dynamic etches used to create these probes modify the shape and dimensions of both the tip and the beam, use of this equation to compute k results in significant error. Poggi et al.16 have presented an improved method for determining beam stiffness that takes into account the actual geometry of the cantilever. If the beam is uniformly loaded along its length by a load per unit length, wo, (see Figure 4), the deflection of its free end is given by eq 3.14

δ)

woL4 8EI

(3)

An example of this type of deflection is when a species is adsorbed uniformly to a cantilever’s surfaces. For example, if one assumes that the species of interest absorbs only on one surface of a microcantilever, a surface stress results on that side. The difference in the stresses on the top and bottom surfaces of the cantilever generates a deflection that is independent of that due to the adsorbed mass. Stoney’s equation17 has been used to relate the difference in surface stresses on each surface of a beam to its deflection (see Figure 5)

Pt 1 )6 2 r Ed

(4)

where r is the radius of curvature of the beam, P is the surface stress due to a coating on one surface of the beam,

1-ν 1 ) 6 2 (∆σ1 - ∆σ2) R Et

(5)

where V is the Poisson’s ratio of the material; it is included to reflect the plane strain condition of the microcantilever (see discussion of the Searle parameter below). ∆σ1 and ∆σ2 are the stresses that act on the top and bottom surfaces of the beam, and t is the thickness of the beam. The deflection of such a beam can be calculated using the geometric relation R-1 ) 2∆z/L2, where ∆z is the beam’s end deflection. Only a difference in absorption between the top and bottom surfaces will cause deflection; equal absorption to both top and bottom surfaces will counteract each other, resulting in no deflection. One will note that the units on the two sides of eq 5 do not match (left-hand side (LHS) ) distance-1 vs right-hand side (RHS) ) distance-2); this is due to the fact that the surface stresses in Stoney’s equation are reported on the basis of a per unit thickness of the layer that causes the deflection in the beam. As is convention in the literature, the term stress is used and its units are N/m. In practice, j ftf, where σ j f is the average normal stress acting on ∆σs ) σ the cross-sectional area residing in a plane that is normal to the neutral axis of the beam-coating composite and tf is the coating thickness. Hence, ∆σs is visualized as the normal force per unit width acting on a normal section of the coating. Stoney’s equation (eqs 4 and 5) is an exact solution for plate bending that is unrestrained at all edges and assumes no interaction between adsorbed species. Hence, a more accurate equation is necessary. For a cantilever with length L (x), width w (2y), and thickness t (z) (see Figure 1), Sader19 has developed a more complete solution for the deflection of a point on a fixed cantilever beam under the influence of a surface stress, wcant(X, Y), as shown in eq 620

{

wcant(X, Y) ) ΩL2 X2 + 2νX[τ1-1 + τ2-2]

() w

-

L

[12-1 + 2υ(τ1-2 + τ2-2 + τ1-1τ2-1) 2 w2 ∑di(12-1 + 2ντi -2) × exp(-τiXLw-1)] L + i)1

()

2

Y2[1 - ∑di exp(-τiXLw-1)] i)1

}

(6)

where X ) xL-1, Y ) yL-1, di ) τ3-i(τ3-i - τi)-1, Ω ) ∆σst/ [4D(1 + υ)], where ∆σs is the surface stress and D ) Et3{12(1 - υ2)}-1 is the cantilever bending rigidity. The τi are defined by eq 7.

τi ) 2x3[5(1 - ν + (-1)ix10(1 - ν)(2 - 3ν))] (7)

Sensing Chemical Interactions via Mechanical Motion

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Figure 6. Mass-spring-dashpot system.

For the case where the deflection at the end of the tip is desired, x ) L and y ) w/2. Klein provides an expression for the error in using Stoney’s equation for a multilayer laminate.21 Stoney’s and Sader’s equations can be used to relate the deflections of microcantilevers to the surface stresses resulting from adsorbed species. Typically, the surface stress is calculated following measurement of deflection. It is difficult to determine a priori the surface stresses due to molecular attachment. In the following section on dynamic methods, we will develop expressions for the effect of surface stress on the dynamic response of a microcantilever.

2.2. Dynamic Response In this section, the dynamic response of a simple massspring-dashpot system first is described to introduce the reader to the appropriate concepts. Then, the quality factor, a well-used metric of microcantilever performance, is discussed. Next, the vibration of a fixed microcantilever in air and shifts in resonant frequency resulting from added mass layers and changes in surface stress are presented. The damping effects of a liquid on vibrational frequency then are presented, which correlates to the submersion of a microcantilever in viscous gas and liquid media. Finally, the effect of temperature is discussed.

2.2.1. Mass−Spring−Dashpot System A mass-spring-dashpot system, such as that in Figure 6, has the equation of motion shown in eq 821

m

d2x dx + c + kx ) F(t) 2 dt dt

(8)

where m is the mass, c is the dashpot constant, k is the spring constant, and x is the displacement of the mass. F(T) represents the generalized force that is exciting the system. The natural frequency of vibration (ωnat) (i.e., with no excitation force, F(t) ) 0) of the undamped beam (i.e., with no dashpot, c ) 0) is given by eq 9.

ωnat ) xk/m

(9)

When damping is present (i.e., one includes the effect of the dashpot), one can use the natural frequency to rewrite the equation of motion with no excitation force (F(t) ) 0) as eq 10. 2

dx dx + 2ζωnat + ωnat2x ) 0 2 dt dt

(10)

The parameter ζ is known as the system’s damping and is defined for this system in eq 11.

ζ)

c c ≡ 2ωnatm 2xkm

(11)

Figure 7. Generic frequency response curves.

Similarly, a freely vibrating microcantilever beam (i.e., due to thermal excitation) will resonate at its natural frequency (eq 9). Thus, as material absorbs onto the beam forming a coating, the microcantilever’s vibrational frequency will decrease and, depending upon the thickness of the adsorbed layer, its spring constant may change.

2.2.2. Quality Factor The quality factor of a microcantilever characterizes the shape of its frequency response curve (e.g., a plot of the displacement amplitude versus frequency) near a resonance mode.22 Accordingly, each resonance mode has its own quality factor. Mathematically, the ith mode quality factor, Qi, is defined as the ratio of the resonance frequency of the ith mode, fi, to the full width of the resonance peak evaluated at the half-maximum (FWHM ) full width half-maximum) of the peak. The quality factor indicates the narrowness of a resonant peak. Figure 7 shows generic frequency response curves and their quality factors and fi values.7 The definition and value of the quality factor for a lightly damped onedegree-of-freedom system, such as an AFM microcantilever, is given by eq 1221

Qi )

fi ωi 1 ) ) ∆ω 2ζ FWHM

(12)

where ωi ) 2πfi and ∆ω ) 2πFWMH. The quality factor depends on the cantilever geometry and the fluid in which the cantilever is immersed.23-30 Increased damping effects lead to a lower Q value. A higher Q value is desired because it lowers the minimum detectable resonance shift (i.e., it increases the frequency resolution). For a quality factor of 10, the minimum detectable resonance frequency shift is roughly 25 Hz, whereas a quality factor of 100 allows for a frequency resolution below 10 Hz.

2.2.3. Plane Strain The discussions in this paper assume that microcantilevers are in a plane strain situation.18,31 The Searle parameter32 is defined as βb ) w2κb/t in a bending mode (Figure 2a) and βl ) t2κl/w in a lateral mode (Figure 2b), where κb and κl are the principal curvatures in the bending and lateral modes, respectively, and dictate the deformational situation a beam is undergoing. A Searle parameter value < 1 indicates a plane stress situation, whereas a Searle parameter value > 100 indicates a plane strain situation.33 From the Euler-Bernoulli beam theory,34 the expression for the maximum curvature

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of a beam subjected to a transverse deflection at its end (causing a bending or lateral deformation) of δ is 3δ/L2. This implies that the maximum Searle parameter value for a bending mode is βbmax ) 3δw2/L2t and for a lateral mode is βlmax ) 3δt2/L2w. Assuming an end deflection at thermal resonance of 100 nm for both the bending and lateral modes (the lateral mode will have a smaller deflection)35 and using the L, w, and t values of 500, 100, and 0.8 µm, typical of silicon microcantilevers,7,18,31 the maximum Searle parameters for the bending and lateral modes are βbmax ) 0.015 and βlmax ) 7.7 × 10-9. These are ,1, indicating a plane stress situation; hence, equations derived in this article are valid for plane strain situations.

2.2.4. Bending Mode Frequency Response An unloaded beam freely vibrating in a bending mode (see Figure 2a) in a vacuum will have a number of resonance frequenciessfrequencies at which it will naturally vibrate under thermally induced excitation. The following discussion is taken from McFarland and co-workers.31 The general expression for the ith mode resonance of the beam, fi, is given by eq 13

fi )

( )x

1 Ri 2π L

2

EI Fbwt

(13)

Equation 19 shows the effects of the mass (+m) and the thickness (+k)

fi+m,+k )

( )x

t Ri 4π L

2

Iads )

EI 3Fb

(15)

Ri2 2πx3

x

k Mb

(16)

where Mb is the mass of the microcantilever beam. If a layer of molecules or other species is added to (coats) a beam, the mass of the composite beam will increase by ∆M; hence, the new resonance frequencies (+m) due to the mass increase will be given by eq 17. The following discussion is taken from McFarland20 and can be used to determine the sensitivity of a microcantilever’s response to added mass.

fi+m )

Ri2 2πx3

x

k Mb + ∆M

(17)

Such an added layer also will increase the second moment of the microcantilever’s cross-sectional area, I, making the beam stiffer. Equation 18 shows this effect (+k).

fi+k )

Ri2 2πx3

x

3EadsIads k + 3 Mb LM b

(18)

b

[( )

]

wtads3 tads + wtads - tb - ycm 12 2

2

(20)

The centroid of the cross section including the adsorbed layer is given by eq 21

Ebtb2 + Eads(2tadstb + tads2) ycm ) 2Eadstads + 2Ebtb

(21)

where tb is the beam thickness and tads is the thickness of the adsorbed layer. The effect of surface stress (+∆σ) effect is given by eqs 22 and 23 +∆σ

)

fi

(Ri∆σ)2 2πx3

[

Ri∆σ ) Ri 1 +

x

k Mb

]

2σL3 π2EI3

(22)

1/4

(23)

where σ is the surface stress. Equations 22 and 23 can be rearranged to allow calculation of surface stress based on the resonant frequency for a microcantilever of rectangular cross section (L, w, and t) as shown in eq 24.7,31

[( ) ]

fi+∆σ 2 π2Ewt3 -1 fi 24L3

σ)

Equation 13 can be modified further by inserting eq 2, resulting in eq 16

fi )

2πx3

(14)

The solutions to eq 14 can be found in the report by Han and co-workers.36 For a rectangular beam with I ) wt3/12, eq 15 can be used to determine fi

fi )

3EadsIads k + 3 (19) Mb + ∆M L (M + ∆M)

where the adsorbed layer (denoted by subscript “ads”) has elastic modulus Eads and second moment Iads are given by eq 20 via the parallel axis theorem.

where Fb is the cantilever material density and Ri is obtained numerically from the frequency relation (eq 14).

cosh Ri cos Ri + 1 ) 0

x

Ri2

(24)

One can combine eqs 19, 22, and 23 to arrive at the most general case of a change in frequency due to adsorbed mass (+m), to increased stiffness due to a change in thickness (+k), and to surface stress (+∆σ); see eq 25.18,20,37

fi+m,+k,+∆σ )

x

(Ri∆σ)2 2πx3

3EadsIads k + 3 (25) Mb + ∆M L (M + ∆M) b

2.2.5. Effect of Damping Due to Viscous Fluids For most cases, air is assumed not to affect the operation of the cantilever; hence, the equations derived above are valid for microcantilevers operating in air. If the microcantilever is used in a liquid or gas that does influence its operation, then damping effects will influence the response of the microcantilever.38 The effect of damping can be estimated by eq 26, which relates f vi , the ith mode frequency of a beam with density Fb oscillating in a vacuum, to f Di , the frequency when oscillating in, and hence damped by, a fluid of density F for a Reynolds number . 1.39 This is an inertiaresistance dominated situation, so the resistance is roughly proportional to the acceleration of the cantilever.40

Sensing Chemical Interactions via Mechanical Motion

f Di f vi

(

) 1+

)

πFw 4Fbt

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-1/2

(26)

It should be noted that, while the damping effects in air have only a minute effect on the resonance frequency, the effect on the quality factor can be quite dramatic, with Q jumping from the order of 10-100 in air to the order of 1 000-10 000 in vacuum. For flows where the Reynolds number is ,1 (i.e., damping mainly due to the viscosity of the fluid surrounding the beam), an expression (eq 27) analogous to eq 26 can be derived23,39

f Di f

v i

)

x

1-

1 4Qi2

(27)

where Qi is the ith frequency quality factor for the damped microcantilever (i.e., submerged in a viscous gas or fluid).

2.2.6. Effect of Air Damping on Quality Factor The effect of air damping on the quality factor of microcantilevers is discussed in more detail by Newell.41 He presents a number of expressions for the quality factor for a microcantilever damped by air. The first case is where the pressure is so low that air damping is negligible. Here, the quality factor is independent of pressure and must be determined empirically. The second case is where air damping is the dominant mechanism but the air molecules are so far apart that they do not interact with each other. In this case, the quality factor is given by eq 28

[ ] ( )

R0T M0 π 3/2 Ftfi Qi ) 2 1 P

()

1/2

(28)

where F is the density of the microcantilever, t is the thickness of the microcantilever, fi is the resonant frequency of the microcantilever, R0 is the universal gas constant, M0 is the molar mass of the air, T is the temperature, and P is the air pressure. The final case is where the air molecules do interact with each other, and here one assumes that the air acts as a viscous fluid. Since viscosity will be independent of pressure, the quality factor also is independent of pressure. If one uses Stokes’ law for damping, eq 29 results

[

]( )

w(EF)1/2 t Qi ) 24 µ L

2

(29)

where w is the width of the microcantilever, t is its thickness, L is its length, and µ is the viscosity of air. Thus, the value of the quality factor is strongly influenced by the media that surrounds the microcantilever.

2.3. Lateral and Torsional Mode Frequency Responses Until now, only pure bending of a microcantilever beam has been discussed. Lateral (Figure 2b) and torsional (Figure 2c) motions can be modeled in similar manners to those presented above. The results presented here are for undamped microcantilevers. The frequency for a freely vibrating beam deflecting in the jth lateral mode is described by eqs 30 and 317

f lj )

( )x

l w Rj 4π L

2

EI 3Fb

(30)

cosh Rlj cos Rlj + 1 ) 0

(31)

where the superscript l signifies the lateral mode of deformation. In a similar manner, the sth torsional resonant modes for a freely vibrating beam can be modeled as eq 327

f ts )

x

2s - 1 4L

Gξ FbIp

(32)

where the superscript t denotes the torsional mode of deformation, G is the shear modulus (G ) E/(2(1 + υ)), Ip is the polar moment of the cross section defined by eq 3342 for a rectangular cross section beam, and ξ is defined by eq 34. For the approximate solution of eq 34, the reader is referred to the paper by McFarland et al.7

Ip )

(

1 (tw3 + wt3) 12

(33)

)

nπw 1 w 192 ∞ 1 tanh ξ ) t4 ∑ 3 t 2t π5 n)1n5

(34)

These equations for lateral and torsional modes of deflection can be used to derive situation-specific resonant frequencies in an analogous manner for the bending mode equations shown above, e.g., for added mass, added thickness, and fluid viscosity effects.

2.4. Temperature Effects Thermomechanical noise (vibration due to thermal agitation) is a consequence of a microcantilever being in thermal equilibrium with its environment. This discussion, taken from Newell41 and Yasumura et al.,28 is for an undamped microcantilever; one utilizes the material discussed above to include damping. Energy dissipation in a microcantilever causes the stored mechanical energy to be converted into heat. The interaction of a microcantilever with the many microscopic degrees of freedom in its environment will subject the microcantilever to constant random excitation. The relationship between energy dissipation and random thermal excitation is embodied in the “fluctuation-dissipation theorem” of statistical mechanics. The net result is that, the lower the mechanical Q of the system is, the larger is the noise force. The mean square vibration amplitude associated with a mode of oscillation at temperature T can be determined from the equipartition theorem as shown in eq 35

1 1 k T ) k〈z2〉 2 B 2

(35)

where kB is Boltzmann’s constant, k is the cantilever stiffness, and z is the microcantilever’s deflection. If one assumes that the noise spectrum is white (i.e., frequency independent), then the spectral density SF ) 4kkbT/ω0Q and the force noise (F) in a bandwidth (B) is given by eq 36

Fmin )

x

4kkBTB ω0Q

(36)

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where ω0 is the cantilever resonance frequency and is equal to 2πf0. For a simple rectangular cantilever, the minimum detectable force can be expressed by eq 37

Fmin )

( ) wt2 LQ

1/2

(kBTB)1/2(EF)1/2

(37)

where w is the microcantilever’s width, t is its thickness, L is its length, E is its Young’s modulus, and F is its density. Similarly, the mean square root deflection is given by eq 38.

zrms )

() ( )() kT kB

1/2

)

2kT wE

1/2

L3 t3

1/2

(38)

Equations 37 and 38 can be used to design the sensitivity of microcantilevers, but one can see that very high quality factors are necessary for ultrasensitive devices.

2.4.1. Effect on Material Properties Many of the material properties of microcantilevers depend on temperature. For example, as the temperature increases, the elastic modulus decreases. The temperature dependence of the elastic modulus of silicon in the high-temperature limit has been modeled semiempirically by eq 3943

( )

E(T) ) E0 - BT exp -

T0 T

(39)

where E0 is the Young’s modulus at 0 K. The constants B > 0 and To > 0 are temperature independent. For aluminum oxide, E0 is reported as approximately 4.6 × 1012 dyn/cm2, B as 4.41 × 108 dyn/cm2, and T0 as 373 K.43 The frequency shift of silicon microcantilevers and variations in the Q-factor over a range of temperatures has been studied.44

2.4.2. Effect on Geometry Temperature also affects the geometry of a microcantilever, with an increase in temperature generally being related to an increase in dimensions through a parameter termed the coefficient of thermal expansion (CTE). For example, the thermal expansion of silicon is on the order of 3.2 parts per million (ppm) per °C and those of a polymer are on the order of 50-100 ppm/°C. As one can see from the discussion in this section, it is critical that calibration and operation of the microcantilever be performed at the same temperature and that the temperature is controlled within very tight tolerances. In the absence of temperature control, differential measurements utilizing pairs of coated and uncoated microcantilevers must be performed.

3. Detection Schemes Vertical, lateral, or torsional movement of a cantilever changes its position. This movement ranges from several angstroms to a micrometer or more, depending upon the dimensionality of the cantilever and the magnitude of surface stress. In this section, methods for measurement of cantilever deflection are examined.

3.1. Optical Lever The optical lever method, illustrated schematically in Figure 8, is the most widely utilized method of quantifying

Figure 8. Schematic of the experimental setup with liquid cell, optical readout of cantilever deflections, and sample liquid exchange system: VCSEL ) vertical cavity surface emitting lasers, PSD ) position sensitive detector. Reproduced with permission from Arntz, Y. et al. Nanotechnology 2003, 14, 86. Copyright 2003 Institute of Physics Publishing.

static and dynamic cantilever deflections. The method involves reflection of a beam of light off the cantilever onto a segmented photodiode or a position-sensitive detector (PSD). Light emitting diodes (LED) and laser diodes are the sources typically used to generate the beam of light. Photodiodes, divided into two or four segments, transduce the light energy striking each segment into an electrical signal that can be compared, amplified, and displayed. Motion of the cantilever changes the position of the reflected light beam on the photodiode and, consequently, the level of light energy incident on each segment. Quad-type photodiodes can, in principle, measure all modes of deflection (bending, lateral motion, and twisting) simultaneously. Typically, the reflected beam is centered on photodiode so that each segment has the same level of illumination at the beginning of each experiment. Then, as the cantilever bends, the laser spot changes location on the photodiode array. By comparing the outputs of the segments, the location of the centroid of the reflected laser spot, and, hence, the deflection of the microcantilever, can be determined. Segmented photodiodes are employed extensively in atomic force microscopes. PSDs are monolithic PIN (positive intrinsic negative) photodiodes with uniform resistance in one or two dimensions. Incident light on the photosensitive region of the PSD generates two photocurrents, each inversely proportional to the distance of the spot from the end of the region. The difference in photocurrents is converted to a voltage, amplified, converted, and displayed. PSDs possess high position resolution and fast response speed and require simple operating circuits. Establishing the relationship between output signal of either segmented photodiodes or PSDs with the magnitude of deflection requires careful calibration. Optical lever detection is currently the most sensitive method for measuring deflection; vertical deflections as small as a few angstroms can be reliably measured with this technique. An intrinsic limitation of this technique is that the laser diode, positioning system, and detector must be external to the air or fluid stream passing by the cantilever. Their dimensions are large in comparison to the microcantilever. Also, this technique is ineffective when the sample passing over the cantilever absorbs or scatters light, e.g., smoky air streams,45 and fluids with suspended particles.46 The optical lever technique is well-suited for detection of cantilever arrays. A number of formats have been published; two have been commercialized. One approach is to have multiple beam sources and detectors, one pair for each cantilever in the array. While this approach enables simultaneous measurement of all cantilevers in the array, the

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integrated source, detector, and signal-processing system is quite complex to design and expensive to manufacture. Sequential reflection of a light beam off each cantilever in the array onto a single detector significantly reduces the complexity of the system and dramatically lowers its cost. One way to achieve this is to scan a single laser source across all of the beams in a cantilever array.47,48 Another way is to sequentially illuminate each element in an array of LEDs or vertical cavity surface emitting lasers.11,49-56 A third approach is to illuminate all cantilevers in the array with a single collimated beam, reflecting the light onto the image plane of a charge-coupled device camera.57,58 Deflection of each cantilever in the array is computed from changes in reflection spot location in images acquired over time.

3.2. Interferometer Interferometric detection of cantilever deflection is based on constructive and destructive interferences that occur when a collimated beam of light reflects off two surfaces displaced from one another.59 In the majority of applications of this technique, cantilevers containing a deformable diffraction grating consisting of a reference and movable set of interdigitated fingers were used. These can be intrinsic to a given cantilever or between cantilever pairs. Chemisorption onto the movable set displaces them relative to the reference fingers and alters the intensity of the diffracted orders is altered. The order intensity is measured with a photodiode array. This technique is capable of measuring very small deflections (as small as 0.01 Å)60 but has a very limited dynamic range. As with the optical lever technique, the interferometric detection technique is ineffective when the sample stream absorbs or scatters the incident or reflected beams. Interferometric detection is being used for hightemperature vibration sensors,61 while Gimzewski and coworkers62 used strobed interferometric microscopy to study the different resonance modes of cantilevers in arrays.

3.3. Piezoresistive The electrical conductivity of a piezoresistive material changes when stress is applied to it. Thus, when a piezoresistive element is integrated onto the cantilever during fabrication, cantilever bending is proportional to the change in resistance. The change in resistance is measured with a Wheatstone bridge, often located at the base of the cantilever.63,64 Piezoresistive elements fabricated onto or into cantilevers comprise either semiconductor or metallic strain gauges. A semiconductor strain gauge is smaller and lower in cost than a metallic foil resistance sensor described below. While the higher unit resistance and sensitivity of semiconductor sensors are definite advantages, their greater sensitivity to temperature variations and tendency to drift are disadvantages in comparison to metallic foil sensors. Another disadvantage of semiconductor strain gauges is that the resistance-to-strain relationship is nonlinear, varying 10-20% from a straightline equation, although this limitation can be overcome through software compensation. Metallic foil strain gauges measure the change in resistance of a metal as it is stretched. By appropriate calibration, the relation between the strain and the change in resistance can be determined and used to determine the strain in the substrate. The gauge factor (GF) of a material is used to characterize its strain sensitivity and is defined by eq 40

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GF )

∆R/R ∆L/L

(40)

where ∆R is the change in resistance, R is the initial resistance, ∆L is the change in length, and L is the initial length. The numerator is also known as the strain. The relation between strain and resistance change is linear. The application of thin, narrow gold traces to a microcantilever similarly can be used to measure its deflection. Piezoresitive detection is the second most common technique used for measuring cantilever deflection, even though its sensitivity is less than that of the optical lever.55 It is applicable to cantilever arrays of almost any size. The read-out electronics can be integrated onto the chip containing the cantilever array. This technique is unaffected by lightabsorbing or scattering components in the analyte stream. Because current is flowing through the cantilevers while measurements are being made, local heating can occur. It can be manipulated by changing the amount of current flowing through the resistive layer.60 Other drawbacks to this technique are thermal, electronic, and conductance fluctuation noise, thermal drifts, nonlinearity in piezoresponse, and poor sensitivity.45

3.4. Capacitive In this detection mode, the cantilever acts as one of the parallel plates of a capacitor. As the cantilever deflects, the distance between the two plates changes and this changes the capacitance of the system. The advantage of capacitive detection is in the simplicity of the associated electronics.65 This technique is not one of the more common ones used because of a number of limitations. To accurately record cantilever deflection, the dielectric material between the conductive plates must be constant throughout the experiment. The presence of analyte within the gap often changes its effective dielectric constant. Additionally, if the parallel plates are brought in too close proximity, they may stick together, which terminates the collection of useful data until they become separated. This phenomenon is frequently encountered when solvent vapor is passed over the cantilever and the solvent condenses onto the surfaces. Also, although the capacitive cantilevers can be integrated onto a microchip,66 scaling down the size of the capacitive cantilever will lower its overall sensitivity because the capacitance of a capacitor is directly proportional to its surface area. For gas sensing, Amirola and co-workers67,68 used capacitive detection of gaseous molecules and found the limit of detection (for their specific cantilever set up) to be 50 ppm for toluene and 10 ppm for octane. Verd and co-workers report sensitivity on the order of 10-8 g for their specific capacitive cantilever system.69

4. Design, Materials, and Fabrication Fabrication of cantilevers is an attractive option for groups with the appropriate resources, facilities, and time available. Creating cantilevers in-house allows greater flexibility in the design of the cantilever to enhance its suitability for the intended application. Only recently have commercial sources for cantilever arrays become available (see, for example, the following websites: http://www.concentris.ch; http://www. micromotive.de/Octosensis_e.php; and http://www.cantion. com). This section examines the interplay between the shape of the cantilever, the material from which it is made, and

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the fabrication methods required to achieve cantilevers with desired mechanical properties.

4.1. Design Considerations The shape of the cantilever often depends upon the detection technique. For example, square pads on the end of cantilever beams are used with capacitive detection systems to increase sensitivity because the measured capacitance is proportional to the surface area of the parallel plates.70 Piezoresistive cantilevers are often U-shaped, with components of the Wheatstone bridge circuit manufactured at their base points. For optical detection schemes, rectangular, paddle, and T-shaped cantilevers are quite common. However, as pointed out by Mertens and co-workers48,71 and previously in the Theory of Operation section, the actual cantilever deflection does not agree exactly with Stoney’s equation, especially when the shape of the beam is different from that which Stoney used in deriving his equation. The implicit assumption behind the equation is that surface stress will cause a uniform curvature of the beam. Because the cantilever beam is clamped on one end, the surface-stress induced curvature is not uniform, and this can cause the beam to twist.48 To reduce cantilever torsion caused by the additional stress at the clamped end, Plaza et al. used T-shaped microcantilever arrays.72 The “T” allows the major part of the beam to be mechanically decoupled from the twist-inducing stress at the clamped end.

4.2. Fabrication of Silicon-based Cantilevers The microfabrication process for silicon-based (i.e., silicon, silicon nitride, and silicon dioxide) cantilevers comprises four main techniques that, when used in combination, yield multiple cantilever chips with the desired shape and mechanical properties. These techniques are film deposition, photolithography, etching, and doping,73 the same as those commonly used in fabricating integrated circuits. The intent of this section is to provide an overview of each technique. For in-depth information, the reader is referred to the review by Hierlemann et al.73 and Madou’s text.74 The fabrication process typically begins with a polished monocrystalline wafer of silicon or silicon-on-insulator (SOI).75,76 SOI wafers are composed of a thick bottom later of single-crystal silicon, a middle silicon oxide layer, and a top layer of single-crystal silicon or silicon nitride.77 SOI wafers are useful because the buried oxide layer acts as an etch stop during the fabrication process. The thin top layer of single-crystal silicon (or silicon nitride) is commonly used as the material of the actual cantilever, so it is important that the defects in this layer are minimized.75-77

4.2.1. Film Deposition Deposition of thin films onto the wafer is carried out by spin-coating, either chemical (CVD) or physical (PVD) vapor deposition, and electroplating. Spin-coating is useful for the formation of polymer thin films, most commonly photoresist, whose utility is described in the next section. The wafer to be coated with the polymer is placed onto a vacuum chuck, which holds it in place. An aliquot of photoresist is dropped onto the wafer, and then the wafer is rotated at thousands of rotations per minute to distribute the polymer evenly over the wafer. Generally, spin-coated polymer thin films have thicknesses of 1-2 µm.74

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CVD is used to deposit silicon oxide and silicon nitride layers that can be used as insulation, masks, and etchstops.73,74 In CVD, gaseous reactants are introduced into the vacuum chamber containing a heated wafer substrate. A thin layer is deposited onto the heated substrate via thermally induced reaction. Depending on the material deposited, CVD films can range in thickness from 20-1500 nm.73 Metals are generally deposited using PVD, i.e., sputtering and evaporation.73,74 In a PVD process, a thin film accumulates on the substrate from a heated reservoir of material in a linear alignment. Metal also can be deposited through electroplating. Metals are useful as reflective surfaces, electrode material, electronic interconnects, thermistors, and chemically reactive binding sites (because alkanethiols covalently bind to gold).

4.2.2. Photolithography Photolithography is the process used to transfer a pattern onto the wafer. First, a thin film of a UV-active polymer in a volatile solvent (i.e., photoresist) is placed on the wafer by spin-coating. Excess solvent is evaporated by heating the wafer in an oven. Next, a glass plate with transparent and opaque regions (mask) that contains the desired pattern is placed close to the wafer; then the mask and wafer are exposed to UV light. Depending upon the tone of photoresist used, UV light exposure initiates chemical bonding between adjacent polymer strands (cross-linking) or chemical bond cleavage along a strand. The reaction is completed as the wafer is placed in the oven for the postbake. Placement of the exposed photoresist wafer into a developer solution dissolves away the uncross-linked polymer and products of the chemical bond cleavage reaction. Etching is the final step that transfers the pattern from the photoresist onto the wafer. The remaining photoresist protects the underlying wafer from the etchant. After the wafer has been etched, the remaining photoresist is removed.

4.2.3. Etching Etching is a process used to remove parts of a thin film or the wafer. There are many different etching reagents; both liquid and dry etchants are available. The specific chemicals used for etching are chosen so that they preferentially etch one type of material over another. This way, thin film layers of various materials can prevent certain features of the wafer from being etched and transfer the desired pattern onto the wafer.

4.2.4. Doping Doping refers to the process of introducing specific impurities into the silicon lattice to alter the electrical conductivity of the silicon. Ion implantation and thermal diffusion are two methods of doping. The type of doping describes whether the dopant contains more or less valence electrons than silicon. For example, incorporation of boron or gallium into the silicon lattice results in p-type doping. These elements have one less electron than silicon; at their location in the lattice, a “hole” is momentarily created. Similarly, incorporation of phosphorus or arsenic into the lattice results in N-type doping. These elements have one more valence electron than silicon; at their location in the lattice, an unpaired electron resides. The unpaired electron and hole are charge carriers and can move about the lattice. Thus, the resistivity of the silicon wafer is determined by

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the dopant type and concentration. Doping is commonly employed in piezoresistive cantilevers. The actual fabrication sequence depends upon the intended use and detection scheme. Silicon-based cantilevers used in optical detection schemes require fewer fabrication steps than those used in piezoresistive, piezoelectric, or capacitive detection schemes. For optical detection schemes, sequences of film deposition, photolithography, and etching are used. A reflective coating often is evaporated onto the surface of the finished cantilever beam to enhance the reflectivity of the beam. Silicon cantilevers used in piezoresistive, piezoelectric, and capacitive detection schemes require more fabrication steps because the detection mechanism is integrated onto the cantilever or the chip holding the cantilevers. Piezoresistive cantilevers require doping in specific areas to create the resistors of a Wheatstone bridge. Detailed information concerning the fabrication sequence and process optimization is readily available in the literature for piezoresistive cantilevers,78-82 piezoelectric cantilevers,83-90 and capacitive cantilevers.68-70,91,92 The number of steps can be reduced by use of silicon-on-insulator (SOI) wafers as the starting substrate.76,77,93 SOI wafers are composed of a thick bottom later of single-crystal silicon, a middle silicon oxide layer, and a top layer of single-crystal silicon. These wafers are commercially available in a variety of layer thicknesses and dopant levels. SOI wafers are useful because the buried oxide layer acts as an etch-stop during the fabrication process. The thin top layer of single-crystal silicon is commonly used as the material of the actual cantilever, so it is important that the defects in this layer are kept at a minimum.

4.3. Fabrication of Polymeric Cantilevers Microcantilevers fabricated from polymers inherently possess readily tailorable mechanical and chemical properties. To alter the stiffness of silicon-based cantilevers, their geometry must be changed or a rigid coating must be applied to the surface. In contrast, the stiffness of polymeric cantilevers requires only a change in material. In this way, microcantilevers with the same geometry but different properties can be produced. This reduces manufacturing costs and simplifies the apparatus required for detection. The materials used for polymer-based cantilevers span a wide range of thermosets, thermoplastics, and polymeric composites. Examples of polymer composites include silver nanoparticles and SU-8,94 carbon nanotubes, poly(m-phenylenevinylene-co-2,5-dioctoxy-p-phenylenevinylene),95 and many other combinations. Polymeric microcantilevers can be fabricated in a variety of ways; the method used is determined by the type of polymer to be used. For example, microcantilevers have been fabricated out of SU-8, a photopolymerizable epoxy-acrylate polymer. The process for fabricating SU-8 cantilevers is quite similar to that used for silicon-based cantilevers. A thin film of SU-8 is deposited onto a wafer by spin-coating. Photolithography then is used to define the regions that will comprise both the cantilever and the chip to which it is attached. The unwanted material is removed and the polymer cantilevers are released from the substrate by immersion in appropriate solvent. SU-8 cantilevers have been made into arrays for optical lever96-98 and piezoresistive99,100 detection schemes. Calleja et al. compared the deflection of silicon nitride cantilevers to SU-8 cantilevers with similar dimensions.96 When a surface stress change of 1 mN/m was applied to the

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Figure 9. Photograph of polymer microcantilevers produced by injection molding.

SU-8 cantilever, a deflection of 11 nm was observed using optical lever detection. When the same stress was applied to the silicon nitride cantilever, a deflection of only 1.2 nm was measured. On the basis of a minimum detectable deflection of 0.5 nm, Calleja et al. concluded that SU-8 cantilevers could be used to detect surface stress changes as small as 60 µN/m. Injection molding, an economical, mass production technique, has also been used to fabricate microcantilevers out of thermoplastic polymers.15,18,37,101 In this process, a molten polymer is forced under pressure into a steel cavity (mold); the shape of the cavity defines the dimensions of both the base and the cantilever(s), as shown in Figure 9. Microcantilevers with thicknesses down to 2 µm and lengths of up to 500 µm have been produced. Because of the small size of microcantilevers, the mold must be heated to the temperature of the molten polymer to ensure mold filling. Any thermoplastic polymer material can be formed into microcantilevers with injection molding; examples include polystyrene, polypropylene, liquid crystal polymer, polymethylmethacrylate, and nanoclay-filled nylon. Cantilevers with tips also have been molded in this manner.101 Injection-molded microcantilevers have been shown to be of equal caliber to commercial silicon microcantilevers. McFarland and coworkers15,18,37,101 detail the fabrication of injection-molded microcantilevers. Despite their advantages over silicon-based cantilever arrays, polymeric cantilever arrays are not commercially available.

5. Chemical Selectivity To achieve selectivity in response, one or more surfaces of the cantilever must be modified to promote binding of desired analytes to the surface and inhibit interfering substances from doing so. A variety of approaches have been used to impart selectivity to microcantilever sensors. The efficacy of a specific approach depends, to a large extent, on the complexity of the sample matrix in which the sensor is used and the chemical reversibility of analyte binding to the cantilever coating. For detection of a gaseous analyte in an air stream, metallic or ceramic films with a high affinity for the analyte are useful. When all surfaces of the cantilever are coated with a selective thin metallic or ceramic film, then the concentration of analyte in the air stream is proportional to the change in frequency. When only one side of the cantilever is laden with the selective thin film, then the concentration of the analyte in the air stream is proportional to the extent of deflection. Thin metallic or ceramic films are applied to the

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desired surfaces of the cantilever using the film-deposition technique described above. To prevent delamination, an intermediate adhesion-promoting layer is often employed. As an example, if a thin film of gold is to be applied to one side of a cantilever by evaporation, then a thin underlayer of titanium or chromium is used to promote adhesion of the gold film onto the silicon cantilever. Mercury vapor in air chemisorbs onto gold films with high affinity.5,102-104 Known interfering substances include water vapor and volatile organic compounds, which bind to either the gold or mercury surface (e.g., thiols and nitriles). Thus, high fidelity detection of mercury vapor in air is possible with a single gold-coated cantilever sensor only when the air stream has been dried and scrubbed of these interfering substances. The chemical selectivity of metallic and ceramic surfaces is significantly diminished in fluid. To enhance selectivity, several research groups have self-assembled monolayer films onto one or more faces of the cantilever. Reactive terminally substituted thiols, silanes, and siloxanes are commonly used to impart specific chemical functionality to the surface. The choice of reactive group depends upon the cantilever surface composition; the choice of terminal group depends upon the specific chemical interaction desired to attract the analyte to the cantilever surface. While this approach is effective in creating densely packed films on the surface, the analyte binding capacity is limited. Construction of multilayer films, through self-assembly, dip-coating, or spin-coating, is one means of increasing the capacity of the chemically selective film. Examples of this approach in the literature include the use of trialkoxysilanes,105-107 cyclodextrins,108-110 hydrogels,111-118 and polymers119,120 as chemically selective coatings. While the capacity of the film increases with increasing film thickness, as the film thickness increases, the following occur: • The rate of analyte transport into and out of the film diminishes, thereby slowing the temporal response of the sensor. • The added mass changes the effective spring constant of the cantileversthe degree of change depends upon the uniformity of coverage of the film on the cantilever. • The viscoelastic response of the film impacts the temporal response of the cantilever and its Q factor. • The number of compounds that partition into the coating increases, thereby reducing the chemical selectivity of the film. Thus, there is a clear tradeoff between film capacity and both detection specificity and temporal response. The impact of this tradeoff is minimized through the use of arrays in which each cantilever in the array has a different coating. A variety of coatings are available; the identification of the most appropriate coatings for detection of specific analytes in either gas or fluid streams has been aided by the application of chemometrics to this field.42,49,121-125 Perhaps the most promising area for development and application of cantilever sensing technology is biology. Deliberate attachment of biological molecules to a cantilever surface opens the possibility of highly selective interactions between the capture molecule and its binding partner. The large number of highly selective binding pairs in biology suggests that, through judicious selection of the coating, cantilever sensor systems can be designed to detect single analytes in complex media with high fidelity. The challenges in creating chemically selective biofilms lie in the following: controlling the spatial distribution and orientation of

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Figure 10. Immersion of a cantilever array into an array of glass microcapillaries filled with food coloring for demonstration purposes. Reproduced with permission from Bietsch et al. Nanotechnology 2004, 15, 873. Copyright 2004 Institute of Physics Publishing.

capturing agent; minimizing the nonspecific binding by components in the sample matrix to the cantilever; maximizing the chemical sensitivity and dynamic range in sensor response; and extending the time before the capture agent denatures, thereby eliminating selectivity in cantilever response. Various approaches for immobilizing capturing agents on cantilever surfaces have been published. Most involve covalent attachment of the biomolecule directly to the cantilever surface or through a hetero-bifunctional linker molecule (e.g., alkanethiol or -siloxane). The use of a linker facilitates uniform distribution of biomolecule on the surface and minimizes denaturing caused by interactions with the cantilever surface. Blocking agents (e.g., polyethylene glycol and bovine serum albumin) typically are employed to reduce nonspecific binding. Further details regarding biomolecular coatings on cantilevers are provided below in the Biologicial Applications section. Coating individual cantilevers within arrays can be challenging. One way is to insert the desired cantilever into a capillary filled with reagent using a micromanipulator.51 The capillary must have an internal diameter larger than the width of the beam, and the wall of the capillary must be thin enough to fit between the cantilever beams in the array. The capillary is held in place for an allotted amount of time required for functionalization and then retracted. When several different cantilevers within the array require functionalization, this approach becomes time-consuming and tedious. Three approaches have proven useful for functionalizing multiple cantilevers: capillary arrays, inkjet printing, and contact printing. All cantilevers within the array can be simultaneously inserted into an array of capillaries (or smallvolume reaction wells) using an appropriately designed micromanipulator (see Figure 10).126,127 All sides of the cantilever are wetted using this approach. Thus, if only one side of the cantilever is to be modified, then the reaction chemistry of the fluid within the capillary must be designed to react only with the desired region (e.g., photochemically initiated reaction). Inkjet printing is also useful for coating individual cantilevers within an array.128,129 Commercial micro-inkjet printing systems are available from several manufacturers (e.g., Cantisens and Microdrop Technologies). Micro-inkjet printing affords efficient and controlled functionalization of only one side of the cantilever (see Figure

Sensing Chemical Interactions via Mechanical Motion

Figure 11. Inkjet printing of individual droplets onto a cantilever array: (a) schematic and (b) image from a video camera. A positioning system allows accurate placement of single droplets onto selected cantilevers. Reproduced with permission from Bietsch et al. Nanotechnology 2004, 15, 873. Copyright 2004 Institute of Physics Publishing.

11), and it is faster than the capillary array. Coatings also can be applied by contact-printing methods using dip-pen lithography130-134 or specially designed stamps.135,136 Commercial dip-pen lithography systems are available (e.g., BioForce, NanoInk, and Nanonics Imaging).

6. Chemical Applications Numerous applications of chemomechanical sensors in environmental monitoring, medical diagnostics, and chemical detection in air and flowing liquids have been published. Several reviews have been published over the past decade.55,60,137-144 This section highlights recent and innovative applications. Because of editorial restrictions on the number of citations, the work cited herein illustrates only some of the applications currently being explored. Our selections are, without doubt, subjective.

6.1. Volatile Organics In a series of papers emanating from IBM Zurich, the University of Basel, and the Paul-Scherrer-Institute, the efficacy of cantilever arrays for detection of specific analytes in complex gaseous mixtures has been demonstrated. Lang and co-workers11,50,145 showed that the diffusion of various alcohols into polymethylmethacrylate coating induces resonance frequency shifts and differential bending of cantilevers. Baller et al.49 coated each cantilever in the eight-cantilever array with a specific polymer layer to transduce a physical process or a chemical reaction into a nanomechanical response. Chemisorption of the analyte induced polymer swelling; the kinetics of the swelling process was related to the vapor pressure and the solubility characteristics of the analyte in the polymers. The array format enabled the use of some cantilevers as reference sensors. Baller et al. distinguished different mixtures of alcohols using principal component analysis (PCA) with mixtures that had been previously characterized. They could not determine the mixing ratio of individual analytes directly from the cluster positions of the mixture’s constituents in the PCA plot because desorption kinetics of analyte mixtures do not depend on the mixing ratio in a predictable way. A year later, the same group demonstrated the simultaneous detection of

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deflection and resonant frequency shifting of cantilevers within the array to the same analytes.42 Using artificial neural network analysis of the deflection and resonant frequency shift data, they demonstrated the utility of polymer-coated cantilevers for both qualitative and quantitative analysis of gaseous mixtures with well-defined composition. In related work, Betts and co-workers119 evaluated two polymeric chromatographic stationary phases as cantilever coatings for select vapor phase analytes. Fagan et al.107 evaluated sol-gels as cantilever coatings for nonpolar vapor phase analytes. Maute and colleagues146-149 used polydimethylsiloxane, polyetherurethane, zinc phthalocyanine, and ethyl cellulose as coatings for the detection of volatile organic compounds in the gas phase. Differential chemisorption of analytes into each polymer film and principal component analysis of the response of each cantilever to the analyte provided a means of qualitative and quantitative determination. Improved performance was found when using higher resonant modes for detection. Headrick et al.109 used focused ion milling of the cantilever surface to create submicron channels across the width of one side of the cantilever. Responses of the nanostructured, coated cantilevers to a series of volatile organic compounds were compared to smooth, coated cantilevers. The results showed that roughened cantilevers were more sensitive, i.e., exhibited an increase in differential stress to the analytes investigated. Lange et al.150 compared the performance of cantilever arrays to thickness shear mode resonators and to surface acoustic wave devices to detect volatile organic compounds in vapor phase. From parallel analyses performed by these transducers on a mixture of n-octane and toluene, it was shown that the limit of detection achieved with cantilever sensors is comparable to that of other acoustic wave-based gas sensors. To enable improved quantification of analyte mixtures, Kurzawski et al.151 evaluated the performance of a singlechip, three-transducer, complementary metal oxide semiconductor gas sensor microsystem. This system comprised a mass-sensitive cantilever, a thermoelectric calorimetric sensor, and an interdigitated capacitive sensor. Each sensing element was coated with various partially selective polymers and then was exposed to different volatile organic compounds. The sensitivities of the three different polymercoated transducers to defined sets of gaseous analytes were determined. These workers have demonstrated that each transducer responds to fundamentally different molecular properties. Thus, the response of each transducer to an analyte provides orthogonal data from which analytes present in the mixture can be quantified, using appropriate signalprocessing and pattern-recognition techniques. Fadel et al.26,152 investigated the analysis of gaseous mixtures using piezoresistive cantilevers of millimeter dimensions. They showed that the choice of the cantilever dimensions and the polymer thickness for gas detection requires compromises concerning sensitivity, response time, quality factor, and resonant frequency. Their comparison between millimeter-size and micrometer-size cantilevers shows the importance of noise in the design of an integrated sensor. The Ziegler group8,153 showed that electrostatic or magnetic actuation of the cantilever results in the enhancement of the quality factor by over 3 orders of magnitude for commercial cantilevers. With actuation, cantilever sensors possess a 1000-

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fold higher mass sensitivity compared to quartz crystal microbalances.

6.2. Chemical Warfare Agents Cantilever technology has been applied to detect chemical warfare agents.154-159 Most published reports have centered on detection of the nerve agent simulant dimethylmethylphosphonate in their studies. Limits of detection were in the 0.5-20 ppb range, depending upon the coating used and the detection mode.160-164 In only a few instances has the selectivity of detection been investigated.165 Assessment of the practical utility of this sensing technology for chemical warfare agent detection awaits a systematic study of the selectivity of coatings to common components found in the atmospheres of cities and on the battlefield.

6.3. Explosives Two approaches have been explored in developing cantilever sensors for detection of explosives: deflagration of particles placed on the cantilever155,166-168 and chemisorption of vapor into thin coatings on cantilevers. The low volatility of most explosives limits the utility of the latter approach and sensitivities obtained to date are less than those of competitive technologies.169-171

6.4. Toxic Metal Ions The concentration of a variety of metal ions in solution has been determined using cantilever technology. For metal ions that chemisorb (or amalgamate) with the metallic coatings used to increase reflectivity for optical lever measurements of cantilever detection, quantization of metal ion concentration is straightforward (e.g., detection of Hg2+ with gold-coated cantilevers).172 For other metals, cantilever deflection can be induced through ion exchange of the analyte onto thin film coatings. For example, the concentrations of Cr2+, Ca2+, Cs+, and CrO42- can be determined via ionexchange with ω-modified alkanethiol monolayers selfassembled onto gold-coated cantilevers.123,173-175 In some instances, very low limits of detection are obtained (e.g., Cs+ and CrO42-), whereas in others (e.g., Ca2+), the limit of detection, dynamic range, and selectivity of the method are not competitive with ion-selective electrode technology.173,175 Hydrogel coatings can also be used in quantifying metal ion concentrations.116,117 Monolayers composed of alkanethiols modified with crown ethers have proven to be an effective way to improve selectivity and sensitivity for specific ion detection.175

7. Biological Applications 7.1. Cells In 2001, Ilic and co-workers176 first reported the detection of Escherichia coli cells using a cantilever array. Selective binding to the cantilever was achieved with the anti-E. coli antibodies immobilized onto the cantilever. Resonant frequency shifts correlated with the number of cells bound to the surface. The sensitivity of the method was sufficient to detect the binding of a single cell. These findings were confirmed by Zhang and Ji.177 Campbell and Mutharasan178 extended the work of Ilic and co-workers, showing that a composite self-excited cantilever made of a PZT film and glass of a few millimeters in length and coated with anti-E.

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coli antibodies can be used to detect E. coli O157:H7 in fluid. Sensitivity achieved was in the order of tens of nanograms, with a limit of detection of only 700 bacterial cells/mL.179 Gfeller et al.180,181 demonstrated that an oscillating cantilever can be used as a sensor for active bacterial growth. Their approach was elegant in its simplicity. E. coli cells were deposited onto cantilevers coated with a thin nutritive agarose layer and kept in a humid environment. Within minutes, the cells started to grow and assimilate water, protein, salts, and carbohydrates from the nutritive layer. To regain equilibrium with the humid environment, the nutritive layer absorbed water; the resultant mass increase produced a commensurate shift in resonant frequency. When they compared the observed frequency shifts due to additional mass loading onto the cantilevers with a conventional bacterial growth curve, all characteristic bacterial growing phases were observed. By incorporation or omission of antibiotics in the cantilever coating, they demonstrated the utility of their approach for rapidly assessing antibiotic resistance. This new application of cantilever array technology offers numerous advantages over conventional bacterial detection methods including rapid real-time detection, labelfree and small analyte volume, and high sensitivity. Ramos and co-workers182 recently showed that the response of oscillating cantilevers to bacteria adsorption depends on the added mass, the site of immobilization of the cell on a cantilever, and the stiffness of the bacterial cells. They predicted that detection sensitivities can be increased by an order of magnitude or more by monitoring higher vibrational modes or scaling down cantilever size. However, the mechanical properties of adsorbed molecules became increasingly important as the size of the resonator was decreased. Taken collectively, these reports portend of the use of microcantilever-based sensors for detection of pathogenic bacteria in medical diagnostics and monitoring of our food supply.

7.2. Viruses Ilic et al.183 first reported on the use of cantilever arrays to detect immunospecific binding of viruses, captured from liquid. Baculovirus particles bound selectively to an AcV1 antibody monolayer immobilized onto the cantilever surface. The resonant frequency shift resulting from the adsorbed mass of the virus particles distinguished solutions of virus concentrations varying between 105 and 107 pfu/mL. Single virus particle detection was achieved using specially designed cantilevers. Similar findings were reported by Gupta and coworkers184 using vaccinia virus, a member of the PoxViridae family and the virus that forms the basis of the smallpox vaccine.185 Ji and co-workers159 showed the utility of antibody-antigen binding interactions for detection of biowarfare agents ricin and tularemia. Dhayal and co-workers130 demonstrated the utility of peptide-functionalized silicon cantilever arrays for detection of whole B. subtilis spores (a nonpathogenic B. anthracis simulant) in liquids. Real-time detection was achieved by monitoring stress changes in the cantilever due to spore binding. Estimates for the induced stress per binding event were obtained. They also observed a higher sensitivity to resonant frequency shifts by monitoring with the fifth mode of vibration. There results suggest that real-time detection of multiple pathogenic organisms can be realized using peptide-funtionalized microcantilever arrays. Campbell and Mutharasan186,187 investigated the detection of pathogen Bacillus anthracis spores in liquid under both

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stagnant and flow conditions. They reported the detection of B. anthracis spores at a very low concentration (300 spores/mL) using piezoelectrically excited millimeter-sized cantilever sensors coated with antibody specific to B. anthracis. High selectivity was demonstrated by detecting B. anthracis spores in the presence of another Bacillus spore (Bacillus thuringiensis) at ratios up to 1/500. More complicated spore mixtures also have been examined.188 In these, the presence of non-antigenic Bacillus species reduced the binding kinetics of B. anthracis spores but did not alter the steady-state response of the sensor. Nugaeva and co-workers189 explored the use of cantilever arrays for selective immobilization and rapid detection of fungal spores. Cantilever arrays were exposed to either the mycelial form Aspergillus niger or the unicellular yeast form Saccharomyces cereVisiae, as models to explore their utility for growth detection of eukaryotic organisms using cantilever arrays. These workers exploited the specific biomolecular interactions of surface-grafted proteins (concanavalin A, fibronectin, or immunoglobulin G) with the molecular structures on the fungal cell surface to achieve selective immobilization of the spores. They found that these proteins have different affinities and efficiencies to bind the spores. Maximum spore immobilization, germination, and mycelium growth were observed on the immunoglobulin G functionalized cantilever surfaces. They also found that spore immobilization and germination of the mycelial fungus A. niger and yeast S. cereVisiae led to shifts in resonance frequency within a few hours, in contrast to conventional techniques that require several days. Measured frequency shifts were proportional to the mass of single fungal spores, and this biosensor could detect the target fungi in a range of 103-106 CFU/mL. This work exemplifies an important application of cantilever array technology in medical and agricultural diagnostics and food- and water-quality monitoring.

7.3. Antigen−Antibody Interactions Raiteri and colleagues60 reviewed the working principles behind cantilever-based sensors based on antigen-antibody interactions. The reader is referred to this review for a critical analysis of the literature in this area up through the year 2000. Several reports have appeared over the last 7 years that utilize antigen-antibody binding for selectivity. For example, Arntz et al.51 presented continuous label-free detection of two cardiac biomarker proteins, creatine kinase and myoglobin, using anti-creatine kinase and anti-myoglobin antibodies covalently anchored to a cantilever array. Binding of the antigen to the anchored antibodies generated sufficient surface stress to enable detection via cantilever deflection. Both myoglobin and creatine kinase could be detected independently using cantilevers functionalized with the corresponding antibodies, in unspecific protein background. These workers showed the utility of reference cantilevers to eliminate thermal drift, undesired chemical reactions (i.e., nonspecific binding), and turbulence from injection of liquids into the cell. They achieved a sensitivity detection of myoglobin below 20 mg mL-1. Grogan et al.190 investigated the activity, stability, lifetime, and reusability of monoclonal antibodies to myoglobin covalently immobilized onto cantilever surfaces. Sucrose was shown to be an effective stabilizing agent for the immobilized antibody layer; with it, the immobilized antibody was found to have a stable active lifetime for up to 7 weeks.

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Alvarez and co-workers191 reported the use of a synthetic hapten conjugated with bovine serum albumin as a bioselective layer for cantilever-based detection of the pesticide dichlorodiphenyltrichloroethane (DDT). Exposure to a solution of a specific monoclonal antibody to the DDT hapten derivative results in deflection of the cantilever. Specific detection is achieved by performing competitive assays in which the cantilever is exposed to a mixture of the monoclonal antibody and DDT. Backmann and colleagues192 showed that single-chain Fv (scFv) antibody fragments can be used as receptors to detect antigens by the static deflection of cantilevers. The authors reported that the performance of the microcantilever-based immunosensor was comparable with surface plasmon resonance. By simultaneously tracking deflection of sensing and reference cantilevers, the differential deflection signal revealed specific antigen binding and was proportional to the concentration of antigen in solution. Dutta et al.193 reported the first demonstration of chiral discrimination using microcantilever sensors. Stable, reusable protein bioaffinity phases based on unique enantioselective antibodies were created by covalently linking monoclonal anti-D- and anti-L-R-amino acid antibodies to nanostructured cantilever surfaces. The temporal response of the cantilever (∆deflection/∆time) was linearly proportional to the concentration of chiral amino acid and allowed quantization of enantiomeric purity up to an enantiomeric excess of 99.8%. Hwang and co-workers194,195 have fabricated a selfactuating and self-sensing piezoelectric cantilever for labelfree detection of a prostate-specific antigen. Cantilevers were coated with parylene-c, deposited by chemical vapor deposition, to electrically insulate the oscillator circuitry for use in fluids.196,197 Specificity in detection of PSA was achieved through its binding to a PSA antibody that was immobilized via host-guest interactions with a proprietary calixcrown self-assembled monolayer. The resonance frequency shift of the cantilever was proportional to antigen concentration. This strategy also was used for detection of C-reactive protein.198,199 Kang and co-workers200 reported the assay of myoglobin concentration using PZT cantilevers coated with biotinylated myoglobin antibodies immobilized onto the surface through streptavidin conjugation. Most of the published works in this area focus on demonstrating that specific antigen-antibody pair interactions lead either to mass increases that can be sensed by shifts in cantilever resonance or to changes in surface stress that produce measurable cantilever deflections. Little attention is directed to performing cantilever-based immunoassays in a clinically relevant setting. A noteworthy exception is the work of Wu and colleagues.201 These authors report the detection of two forms of prostate-specific antigen (PSA) over a wide range of concentrations from 0.2 ng/mL to 60 mg/mL in a background of human serum albumin (HSA) and human plasminogen (HP) at 1 mg/mL. Prostate-specific antigen is a particularly useful marker for early detection of prostate cancer and in patient monitoring for disease progression. In serum, this biomarker exists in two forms: uncomplexed and complexed with the serum protease inhibitor R1antichymotrypsin. Early diagnosis of prostate cancer requires an accurate measure of both the total concentration and the ratio of the complexed to uncomplexed forms of the antigen in serum. In addition, the clinically useful range spans from 0.01 to >10 ng/mL. The dose-response curve they obtained

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Figure 12. Steady-state cantilever deflections as a function of uncomplexed (fPSA) and complexed (cPSA) prostate specific antigen concentrations for three different cantilever geometries. Reprinted with permission from Wu et al. Nat. Biotechnol. 2001, 19, 856. Copyright 2001 Nature Publishing Group.

with cantilevers of differing length is shown in Figure 12. This curve was obtained under static conditions that included thermal regulation. In contrast to the conventional enzyme-linked immunosorbent assay for this antigen, the cantilever-based assay required no labels and was performed in a single reaction without additional reagents. A logical extension of this work involves the use of an array of microcantilevers to perform multiple assays. High fidelity, clinically relevant detection of this biomarker for prostate cancer would be anticipated via the coating of individual cantilevers within the array with antibodies selective for different epitopes on this antigen. A number of such antibodies are now commercially available.

7.4. DNA Hybridization Fritz and his colleagues125 pioneered the use of cantilevers for detection of nucleic acid hybridization. Deflection of each cantilever in the array was measured using the optical beam deflection technique. 5′-Thio-modified synthetic DNA oligonucleotides with different base sequences were covalently immobilized on the gold-coated side of the cantilevers the array. When solutions containing the complementary oligo were injected into the liquid cell, hybridization resulted in a change in surface stress between the functionalized gold and the nonfunctionalized Si surface, bending the cantilever. This is shown schematically in Figure 13. This work stimulated interest in exploiting the sensitivity of chemomechanical detection of DNA hybridization. A crucial test for any DNA hybridization sensor is its ability to discern mismatches. Fritz and colleagues125 observed a small but measurable difference in surface stress between a pair of complementary oligos and a pair with a single base mismatch between two DNA sequences that can be detected. Hansen and co-workers202 further evaluated the capability of cantilever sensors for detecting single base mismatches. They found that the direction of cantilever bending, whether tensile or compressive, depended up the number and location of mismatch sites along the strand pairs. Wu et al.203 showed that the magnitude of cantilever deflection during hybridization depends upon the ionic strength of the matrix. McKendry et al.127 systematically examined the impact of single strand extensions on cantilever deflection during DNA hybridization. In all cases, compressive surface stress results from hybridization, regardless of whether the complementary

Figure 13. Schematic illustration of the hybridization assay. Each cantilever is functionalized on one side with a different oligo (red or blue). (A) The differential signal is set to zero. (B) After injection of the first complementary DNA strand (green), hybridization occurs on the cantilever laden with the matching sequence (red), increasing the differential signal. (C) Injection of the second complementary DNA oligo (yellow) causes the cantilever functionalized with the second oligo (blue) to bend. Reprinted with permission from Fritz et al. Science 2000, 288, 316. Copyright 2000 American Association for the Advancement of Science.

oligomers have the same number of nucleotides. They found that hybridization between two complementary 12-mers generated an average compressive surface stress of 2.7 mN/m. Hagan et al.204 presented an explanation of cantilever deflections resulting from adsorption and subsequent hybridization of DNA molecules. Using an empirical model, they predicted deflections upon hybridization that are consistent with experimental results. They asserted that hydration forces, not conformational entropy or electrostatics, are the dominant contributors to deflections arising from DNA hybridization. They showed that predicted deflections before and after hybridization strongly depend on surface coverage as well as the degree of disorder on the surface. The latter point was experimentally verified by Alvarez and co-workers.205

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In a follow-up report from the Majumdar group, Stachowiak et al.206 provided experimental evidence that the surface stresses resulting from hybridization depend on oligo length, grafting density, and hybridization efficiency. At low ionic strength, the osmotic pressure of counterions dominates the intermolecular forces, while at higher ionic strength, the grafting density is independent of the ionic strength and hydration interactions dominate. They also showed that, regardless of the length and grafting density of the singlestranded probe DNA, surface stress was related exponentially to the density of hybridized DNA. The same group207,208 observed surface stress changes in response to thermal dehybridization, or melting, of double-stranded DNA oligonucleotides that were grafted on one side of a microcantilever beam. Changes in surface stress occur when one complementary DNA strand melts and diffuses away from the other, resulting in alterations in the degree of hydration and electrostatic interactions between the remaining neighboring surface-grafted DNA molecules. They were able to distinguish changes in the melting temperature of dsDNA as a function of salt concentration and oligomer length. Recent effort has focused on improving sensor performance. Several groups have evaluated piezoresistive detection of DNA hybridization as an alternative to optical methods.128,209,210 While piezoresistive detection is less sensitive than the optical lever method, piezoresistive methods are sufficiently sensitive to detect hybridization and single base mismatches. Improvement in sensitivity is anticipated with continued optimization of piezoresistive cantilevers. Others have focused on fabricating cantilevers from polymers in hopes of lowering the limit of detection through reduction in cantilever spring constant without significantly changing the active area.96,211 Su et al.212 used gold nanoparticle modified oligos to improve the mass sensitivity of resonant frequency-based microcantilever detection of nucleic acid hybridization. Their method is capable of detecting DNA concentrations as low as 0.05 nM. As clearly pointed out by Alvarez and coworkers,205 detection of nucleic acid hybridization requires reference cantilevers sensitized with noncomplementary DNA to decouple the molecular recognition signal from nonspecific binding events and matrix effects. This highlights the need for an improved understanding of the mechanisms responsible for surface stress due to the biomolecular interactions. Such knowledge is crucial for the development of immobilization procedures in which the geometry of the receptor molecules is addressed to generate high interaction forces between neighboring molecules during molecular recognition.

7.5. Enzymes Subramanian et al.213 reported on the first microcantileverbased enzymatic assay. The enzyme glucose oxidase was immobilized onto a gold-coated silicon cantilever with glutaraldehyde following coating of the gold surface with poly-L-lysine. Quantifiable deflection of the cantilever was observed in the presence of analyte. Their analysis of the heat of the enzymatic reaction and the thermal sensitivity of the cantilever suggested that cantilever deflection is not simply a result of reaction-generated heat but appears to result from surface-induced stresses. They offered two hypotheses to account for the surface stress: entropic effects due to the continual binding of glucose at the active site of the enzyme and changes in the local chemical environment that result from glucose conversion to gluconic acid and peroxide.

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Figure 14. Schematic illustration of protein binding and enzymatic assays with cantilevers.

Yan and co-workers sought to clarify the source of cantilever deflection when immobilized glucose oxidase is exposed to glucose.214,215 Their immobilization strategy differed from that of Subramanian et al. in that the enzyme was electrostatically immobilized within an alternately charged polyelectrolyte multilayer structure that comprised poly(sodium 4-styrenesulfonate) and polyethyleneimine. The multilayer approach provided improved performance. They proposed that cantilever bending results from both a conformational change of the enzyme in the presence of glucose and from protonation of the polymer multilayer structure as a consequence of the enzyme-catalyzed oxidation of glucose to gluconic acid. Pei et al.216 further characterized the performance of a glucose oxidase-based cantilever sensor. They cross-linked the enzyme to bovine serum albumen chemisorbed onto the surface. They concluded that the deflection response of the cantilever cannot be due to the heat of the enzymatic reaction and attributed the deflection mechanism to changes in the local chemical environment of the coating layer. They noted that the poor reproducibility of results for this enzyme-based glucose sensor is likely due to the corrosive nature of peroxide produced by the enzymatic reaction. Clearly, the mechanism behind the surface-induced stress observed for this type of glucose sensor remains unknown. Bottomley and co-workers reported the use of microcantilevers as sensors of enzymatic function.217 Exposure of cantilevers coated with enzyme substrates to enzymes capable of changing substrate mass, conformation, and charge results in measurable deflection of the cantilever as shown in Figure 14. Enzyme inhibitors also can be identified using this approach. Stevenson and colleagues218 monitored the restriction and ligation of cantilevers coated with DNA. An oligo containing the Hind III restriction site was immobilized on the cantilever and then digested with that enzyme; strand scission produced cantilever bending and left behind a shortened oligo with a single-stranded sticky end. Exposure of a second oligo with a compatible end to the DNA on the cantilever in the presence of ligase resulted in the extension of the immobilized oligo and commensurate cantilever deflection in the opposite direction. The authors point out that, since most DNA restriction and ligation enzymes require dithiothreitol to retain their activity, immobilization of the oligo through thiol linkages must be avoided. Otherwise, displacement of the thiolated DNA from the gold surface by dithiothreitol will produce cantilever deflection and complicate detection of the restriction and ligation events. Liu and colleagues219 presented a new approach to track enzyme action with cantilevers. Their technique relies on

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the detection of bead detachment from the cantilever due to the enzyme cleavage of the linker tethering the bead to the cantilever. To illustrate this principle, they used the enzymatic action of Botulinum neurotoxin type B on its substratum, the synaptic protein synaptobrevin 2. Nickel-agarose beads were functionalized with recombinant synaptobrevin 2 conjugated to six consecutive His residues at its C terminus. To suspend the bead off the cantilever tip, they used proteinprotein interaction (synaptobrevin 2 with another synaptic protein, syntaxin 1A). In the presence of zinc ion, the neurotoxin cleaves synaptobrevin 2, leading to the detachment of the bead from the tip. Since the mass of the bead is many times larger than that of the immobilized protein, its detachment chemically amplifies the mass loss of the protein fragment. The bead detachment technique is general and can also be used for any cleavage reaction. For example, Weizmann et al.220 utilized the endonuclease scission of magnetic beads functionalized with sequence-specific DNAs to detect single base mismatch specificity of the endonucleases. Magnetic beads were used to reduce thermal motion and amplify the mechanical motion of the cantilever to enzymatic action. In a subsequent report, they extended this approach to the development of enzyme-based AND or OR logic gates.221 The bead detachment technique is not limited to cleavage reactions; it is also suitable for displacement reactions, such as in receptor-ligand pairs, where the introduction of one chemical leads to the displacement of another.

8. Recommendations for Future Work 8.1. Guidelines for Reporting Sensor Performance To facilitate comparison with other sensing platforms, we suggest the following figures of merit be included in all future publications regarding the performance of microcantilever sensors: detector sensitiVity, limit of detection, dynamic range, and sensitiVity of the analysis. Within the microcantilever community, the term sensitiVity is used to describe several different parameters. Some workers use this term to describe the minimum concentration of analyte that can be detected. Others use it to describe the performance characteristics of the sensing technique used to measure shifts in cantilever resonance frequency or changes in cantilever deflection. A third group uses this term to describe the slope of the calibration curve. Multiple usages of the same word can lead to confusion and misinterpretation on the part of the reader. We suggest more explicit terminology be used in reporting results. The term “limit of detection” should be used to describe the minimum concentration of analyte that can be reliably detected. Convention within the analytical chemical community is that this is the concentration that gives a sensor response signal equal to three times the background noise level. The performance of the sensor to changing concentrations of analyte should be reported and distinguished from the sensitivity of the device used to measure cantilever movement. We suggest that the term “detector sensitiVity”, the measured change in signal per unit value of the sensor response, be employed in characterizing device performance. For optical level and piezoresistive detection of cantilever deflection, the detector sensitivities would have units of V/nm and (∆R/R)/nm, respectively. This term is a function of the properties of the cantilever, the deflection measurement technique, and the signal amplification techniques employed.

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This figure of merit would be of benefit to those trying to compare cantilever designs, materials, and detection techniques. We suggest that the performance of the sensor to changing concentrations of analyte be labeled “sensitiVity of the analysis”. This parameter is determined from the slope of the linear region of the dose-response curve. As acquisition of this parameter requires exposure of the coated cantilever to varying concentrations of analyte and measurement of the system’s response, a measure of the chemical reversibility of analyte binding to the coating is readily obtainable and also should be reported. Similarly, in the course of determining the chemical sensitivity, the experimenters should determine and report the “dynamic range” of the sensor for the specific analyte under study. While this issue may seem obvious to the majority of readers, the omission of this information in many of the published papers has made it difficult for us, during the course of writing this review, to evaluate the scientific contribution of many papers and include them in the context of important, new applications of microcantilever sensor technology.

8.2. Experimental Design Considerations In the first few years following the invention of this sensor technology,1,3-6,222,223 experimentation with a single cantilever was commonplace. Since then, it has become well-established that cantilevers respond to small changes in temperature, viscosity, and ionic strength of the medium in which they are immersed, as well as to the flow dynamics of the cell that houses the cantilever chip. Thus, the utility of single cantilever experimentation in fluid streams is, at best, questionable. In many instances, interpretation of results of present-day research involving single-cantilever experimentation are based largely on the assumption of fixed conditions between sequential experiments. Often, experiments expressly designed to test the validity of the assumption are unreported. Microcantilever arrays are the preferred format. They enable control experiments to be performed simultaneously with analyses and provide more reliable control of empirical factors such as thermal drift, changes in viscosity, and solution flow dynamics. They also provide a straightforward means and correct for nonspecific adsorption and nonspecific chemical reactions that may occur on or within the chemically selective coating. In addition, multiple targets can be detected simultaneously, leading to high-throughput measurements and producing distinct recognition patterns from complex mixtures.139 We recommend that all future work with microcantilever sensor technology be performed using cantilever arrays. The field of microcantilever sensors has matured to the point where reports of new applications of this technology should include performance testing under relevant conditions with measures of the fidelity and selectivity of detection. For example, if a new chemically selective cantilever coating is developed that provides a means for detecting a volatile analyte in air, then the report should include the performance of this coating when exposed to a variety of air samples (e.g., compressed air, laboratory air, auto exhaust, etc.). Similarly, reports of new biological applications should include results of tests carried out in the fluids where the analyte is typically found (e.g., sputum, serum, urine, and cell lysate). While there is some value to disseminating results of analyses carried out with pristine solutions, reports of successful detection of specific analytes in complex mixtures signifi-

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cantly advances the field and provides strong impetus for increased commercial participation in the development of this sensor technology.

8.3. Fruitful Areas for Further Research To become competitive with existing commercial sensing technologies (e.g., quartz crystal microbalance, surface plasmon resonance, and surface acoustic wave), microcantilever sensors must provide faster, cheaper, more sensitive, rugged, and reliable analyses. In addition, the microcantilever sensing system must be easy to operate and field deployable. On the basis of these benchmarks and the present state-ofthe-art, there is a need for more research in the following areas.

8.3.1. More Selective Coatings The quartz crystal microbalance and surface acoustic wave device are two commercialized sensing technologies that rely on changes in mass for detection. To compete with these technologies, analyses based on shifts in cantilever resonance either should be performed on short, stiff cantilevers with resonance frequencies in the MHz range, or by tracking shifts in one of the higher resonance modes of conventional cantilevers. The latter is preferred as the small dimensions of short, stiff cantilevers reduce the capacity of the sensor and, thus, the dynamic range of detection. We suggest increased effort in the development of new, highly selective coatings that give rise to large changes in surface stress upon analyte binding. It seems likely that these coatings will utilize highly specific biomolecular interactions. Also needed are novel packaging approaches to increase the shelf life of existing biomolecular coatings.

8.3.2. Increased Sensitivity and Faster Response To compete favorably with benchmark sensing technologies, the speed and sensitivity of analysis with microcantilever sensors must be improved. Shortening the temporal response of cantilevers to analyte passing over the selective layer requires additional insight into the analyte-binding mechanism(s) as well as a dramatic reduction in the volume of the compartment in which the cantilever array is housed. The latter will require careful attention to the mass transport of analyte to the sensor surface and modeling of the fluid dynamics of analyte flow through the compartment and about the cantilevers. Thus, incorporation of arrays in microfluidics cartridges would seemingly be one way to shorten the response time. Another approach is being pursued by the Manalis group at MIT.224-226 They have achieved significantly enhanced sensitivity and very low limits of detection for fluidborne analytes using specially designed cantilevers that have integrated microfluidic channels within them (see Figure 15). The analyses are performed using optical lever detection with the cantilever under vacuum and sample flowing through the interior of the cantilever. This approach eliminates both the damping normally encountered when the cantilever is immersed in fluid and light scattering or absorption by the fluid sample, which negatively impacts optical lever detection. Selectivity in detection is achieved by precoating the walls of the microfluidics channel.225 This very recent advance suggests that the present shortcomings which impede many applications of microcantilever technology will be removed through innovations in the design of cantilevers,

Figure 15. Schematic illustration of cantilevers with integrated microchannels developed by Manalis and co-workers. (a) A suspended microchannel translates mass changes into changes in resonance frequency. Fluid continuously flows through the channel and delivers biomolecules, cells, or synthetic particles. (b) While bound and unbound molecules both increase the mass of the channel, species that bind to the channel wall accumulate inside the device, and, as a result, their number can greatly exceed the number of free molecules in solution. This enables specific detection by way of immobilized receptors. Reprinted with permission from Burg et al. Nature 2007, 446, 1066. Copyright 2007 Nature Publishing Group.

detection devices, and sample delivery systems; intelligent design of coating-layer composition and high-throughput methods for their application; and incorporation of chemometric methods of analysis for processing data acquired with cantilever arrays. With more advances such as these, microcantilever technology will enable rapid detection of harmful agents that may be present in the air we breathe and the fluids we ingest.

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CR0681041

Chem. Rev. 2008, 108, 543−562

543

Radionuclide Sensors for Environmental Monitoring: From Flow Injection Solid-Phase Absorptiometry to Equilibration-Based Preconcentrating Minicolumn Sensors with Radiometric Detection Jay W. Grate,*,† Oleg B. Egorov,†,§ Matthew J. O’Hara,† and Timothy A. DeVol‡ Pacific Northwest National Laboratory, P.O. Box 999, Richland Washington 99352, and Environmental Engineering and Earth Sciences, Clemson University, 342 Computer Court, Anderson, SC 29625−6510 Received March 7, 2007

Contents 1. Introduction 2. Background 2.1. Flow Injection and Sequential Injection Analysis 2.2. Preconcentrating Minicolumn Sensors 2.3. Bead Injection and Renewable Surface Sensing 2.4. Automated Radiochemical Separation and Analysis 3. Radionuclide Sensors 3.1. Challenges of Radionuclide Sensing in Water 3.2. Minicolumn Sensors Based on Extractive Scintillating Resins 3.3. Composite Bed Scintillating Minicolumn Sensors 3.4. Sensor Regeneration or Renewal 3.5. Equilibration-Based Sensing 3.6. Chromatographic Theory for Equilibration-Based Sensing 3.7. Engineered Radiometric Preconcentrating Minicolumn Sensors for Groundwater Measurements 3.8. Planar Dual-Functionality Radionuclide Sensors 3.9. Planar Radionuclide Sensors Based on Diodes 3.10. Fiber-Based Sensor 3.11. Whole-Column Chromatographic Sensor 3.12. Dual-Functionality Sensor for Tritiated Water in Air 4. Discussion 5. Acknowledgment 6. References

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1. Introduction The development of in situ sensors for ultratrace detection applications in process control and environmental monitoring * To whom correspondence should be addressed. E-mail: [email protected]. Telephone: 509-376-4242. Fax: 509-376-5106. † Pacific Northwest National Laboratory. ‡ Clemson University. § Present address: Isoray Medical, Inc., 350 Hills St., Suite 106, Richland, Washington 99354.

remains a significant challenge. Such sensors must meet difficult detection limit requirements while selectively detecting the analyte of interest in complex or otherwise challenging sample matrixes. Nowhere are these requirements more daunting than in the field of radionuclide sensing for R- and β-emitting radionuclides in water. The detection limit requirements can be extremely low. Nevertheless, a promising approach to radionuclide sensing based on preconcentrating minicolumn sensors has been developed. In addition, a method of operating such sensors, which we call equilibration-based sensing, has been developed that provides substantial preconcentration and a signal that is proportional to analyte concentration, while eliminating the need for reagents to regenerate the sorbent medium following each measurement. While this equilibration-based sensing method was developed for radionuclide sensing, it can be applied to nonradioactive species as well, given a suitable on-column detection system. By replacing costly sampling and laboratory analysis procedures, in situ sensors could have a significant impact on monitoring and long-term stewardship applications. The preconcentrating minicolumn sensor relies on a solid phase, typically a packed bed of particles or beads, to collect and concentrate the analyte species of interest within a detector. A portion, or ideally all, of the solid phase is within the detected volume. Typically the solid phase is packed in a small column with fluid flow parallel to the column axis. This can be a straight column or a column that has been coiled to fit within the detection system. Some sensors, however, are prepared with the solid phase in a disk or plate geometry with radial flow from the center to the periphery. Optical or luminescent methods predominate for the detection of analyte species, or their reaction products, captured on the solid phase. The radionuclide sensors described in this review are primarily preconcentrating minicolumn sensors that rely on the detection of scintillation photons from a dualfunctionality column. The column contains selectively sorbent functionality and scintillating properties in the same material, or in materials that are in close proximity to one another. The preconcentrating minicolumn sensor is shown schematically in Figure 1a, along with the radionuclide sensor concept in Figure 1b, where the scintillation photons are detected with a pair of photomultiplier tubes. The fluidic format is an efficient means of collecting analyte species from much larger sample volumes and for concentrating them on the solid phase for on-column detection. This precon-

10.1021/cr068115u CCC: $71.00 © 2008 American Chemical Society Published on Web 01/03/2008

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Grate et al. rations, and equilibration sensing. His research interests include radiochemical analysis, selective radionuclide sensing, nuclear waste process monitoring, medical radioisotope production, and laboratory automation.

Jay W. Grate is a Laboratory Fellow at the Pacific Northwest National Laboratory (PNNL) and an Affiliated Professor with the Chemistry Department of the University of Washington. He received a B.A. in Chemistry at Rollins College and received his Ph.D. in Chemistry from the University of California, San Diego. After postdoctoral research at the University of California Irvine, he joined the Naval Research Laboratory in 1984, moving to PNNL in 1992. He spent a sabbatical at the Scripps Research Institute prior to joining PNNL. Dr. Grate’s research has focused on chemically interactive polymers and nanomaterials, chemical vapor sensors, radiochemical separations and sensing, and bioanalytical fluidics for biothreat detection. His work integrates aspects of the chemical sciences, material sciences, and measurement sciences into new microanalytical principles, methods, and systems. His work in the radioanalytical field has entailed the development of new radiochemical analysis methods using sequential injection separations, fully automated radiochemical process monitors, and sensors for radionuclides in water. He has published over 100 papers in peer-reviewed journals and numerous book chapters, and he is author or coauthor on over a dozen patents, several of which have been licensed. He received an R&D 100 Award in 2004 and an American Chemical Society Regional Industrial Innovation Award in 2007.

Prior to joining IsoRay Medical Inc. in 2005 as a Director of Research and Development, Dr. Egorov has worked as a senior research scientist at the Pacific Northwest National Laboratory. He received his Ph.D. in Analytical Chemistry from the University of Washington in 1998. In addition to his Ph.D., Dr. Egorov has an M.S. degree specializing in Radiochemistry from Moscow State University in Moscow, Russia, and an M.S. in Environmental Analytical Chemistry from the University of Washington. His research at PNNL has specialized in microfluidic systems and their application toward automation of radionuclide separations and analysis, where he has authored or co-authored several key publications, including invited review articles and book chapters. Dr. Egorov pioneered the application of flow-injection techniques for automating radiochemical analyses of nuclear wastes and process monitoring, renewable surface sensing and sepa-

Matthew J. O’Hara is a scientist at the Pacific Northwest National Laboratory. He received B.A. degrees in Chemistry and in Geology from the University of Montana in 1996, and a Masters degree in Business Administration from Washington State University in 2004. He has been involved in scientific research in the fields of radioanalytical chemistry and laboratory automation for over 10 years. The primary focus of his research has been the selective preconcentration, separation, and detection of actinides and radioactive fission products from various matrices using automated fluid handling and detection systems. The instrumentation he has developed has targeted specific Rand β-emitting radionuclides with the objective of activity quantification ranging from ultralow activities in environmental waters to high-level activity in nuclear waste samples. His research has resulted in the development of medical isotope separation systems, sensors for groundwater monitoring, and prototype process monitors for Hanford’s nuclear waste treatment plant.

Timothy A. DeVol is a Professor of Environmental Engineering and Earth Science at Clemson University. Dr. DeVol earned a B.S. in engineering physics from The Ohio State University, and a M.S. and Ph.D. in nuclear engineering from the University of Michigan. Dr. DeVol has been teaching and conducting research at Clemson University since 1993. He teaches courses in radiation and health physics, ionizing radiation detection, and radioactive waste management. His major research interests are in the detection and measurement of ionizing radioactivity in the environment, environmental radiochemistry, and statistical analysis of monitoring data. Dr. DeVol has published over 40 papers in peer-reviewed journals and has made over 130 scientific presentations.

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Throughout the narrative, we will focus on sensors for Tc as the prototypical examples for illustrating the detection principles. These are the most mature radionuclide sensors to date, and 99Tc is an important radionuclide to detect in environmental monitoring. It is generated from the thermal fission of 235U with a high production yield of 6% and is a significant radioactive contaminant at U.S. Department of Energy sites associated with nuclear weapons production. It has a long radioactive half-life of 2.13 × 105 years, and it is highly mobile in the environment in the Tc(VII) oxidation state as the pertechnetate oxyanion, TcO4-. Hence, this contaminant will persist in the environment, and it must be monitored as it is unlikely to stay in one place. As a portion of the background material, we will cite selected material on preconcentrating minicolumn sensors using transduction mechanisms other than radioactivity to detect analytes ranging from metal ions to organic pharmaceuticals and nutrients. This work has largely occurred from 1985 to the present. These types of fluidic sensors were developed within the fields of flow injection and sequential injection analysis; therefore, these topics will be introduced in the background material. In addition, the method of “bead injection” has been developed where the sorbent material is delivered to the detection flow cell for each measurement and then released. Bead injection represents a renewable surface preconcentrating sensor. Before turning our attention entirely to radionuclide sensors for water monitoring, we will also provide some background on automated radiochemical analysis. Finally, this review is not concerned with assays that use radionuclides as labels, such as the scintillation proximity assay (SPA).19-21 This method is used for studies of binding interactions of biologically relevant compounds, using scintillating microspheres and radiolabeled molecules of high specific activity. It is designed to discriminate between bound and unbound molecules. Although this assay combines chemical selectivity with scintillation, its purpose is not focused on environmental radiochemical analysis applications. 99

Figure 1. Schematic diagrams for (a) preconcentrating minicolumn sensors and (b) radionuclide sensors based on dual-functionality materials in preconcentrating minicolumn sensors. The column may also be a coil within the detection zone.

centration increases sensitivity and reduces detection limits. In addition, concentrating the analyte in a smaller volume can simplify and reduce the size of the detection method used. These features are all desirable for environmental sensors, where analytes are typically present at very low concentrations and the sensor should ideally be suitable for at site or in situ deployment. In typical use, the fluidic system containing the sensor processes a sample aliquot of defined volume, and the analytical signal is taken within a defined time range of the process. In this regard, the sensor may be regarded as part of an assay system that determines the quantity of the analyte in that particular sample volume. However, as we shall illustrate with the equilibration-based radionuclide sensor to be described below, such a sensor can also function as a true sensor whose signal goes up and down with the ambient concentration. As long as the sensor is allowed to reach equilibrium, the signal is not dependent on the volume of the sample processed through the flow cell; it depends on the concentration of the analyte and its interaction with the solid phase. The aim of this review is to cover radionuclide sensors for R- and β-emitting radionuclides that combine some form of selective sorption with a radiometric detection method andsas a primary aimsto comprehensively review preconcentrating minicolumn sensors for radionuclide detection. This work that has largely occurred from 1995 to the present.1-18 As a secondary aim, we will cover radionuclide sensors that combine sorption and scintillation in formats other than minicolumn sensors. We are particularly concerned with the detection of R- and β-emitting radionuclides in liquids, which presents particular challenges as we shall describe below. We will not cover systems to detect γ rays or the radionuclides that emit them, since γ rays can readily pass through condensed media to radiation detectors and the γ-ray energy spectrum provides considerable selectivity. Nonetheless, preconcentrating sensor methodologies to be described in this paper can also offer advantages for lowlevel sensing of γ-ray emitters where preconcentration is required.

2. Background 2.1. Flow Injection and Sequential Injection Analysis A large number of sensors fitting our definition of preconcentrating minicolumn sensors were developed as detectors for flow injection analysis systems.22-24 Flow injection consists of a fluidic analysis approach where sample and reagent solutions are driven in a continuous forward flow paradigm through a progression of mixing, reaction, and/or separation steps to a flow-through detector. The prototypical assay was colorimetric with an optical absorbance detector. Figure 2a provides a schematic diagram of a flow injection system. The original fluid drive was typically one or more peristaltic pumps coupled with at least one valve for sample injection. By incorporating an appropriate solid phase within the detector, creating an optosensor, the analyte or reaction products could be captured, preconcentrated, and focused within the detection zone for greater sensitivity. Sequential injection arose as a subsequent approach to flow-based analysis; a schematic diagram is shown in Figure 2b.25-28 Sequential injection relies on programmed bidirectional flow. The preferred fluid drive is a digital syringe pump coupled to a multiposition valve with a holding coil

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Figure 2. Schematic diagrams of prototypical (a) flow injection analysis, (b) sequential injection analysis, and (c) sequential injection separation systems.

in between. In a typical analysis, samples and reagents are pulled as zones into the holding coil using reverse flow, stepping the multiposition valve from one position to another for each solution. These zones are then propelled forward though the multiposition valve to the analysis system consisting of a reactor or separator and detector. In a simple colorimetric assay, the zones intermix by dispersion, generating reaction products for detection. Sequential injection systems are fully automated with a computer providing precise control of volumes, flow rates, and timing. The detector in some examples has been a preconcentrating minicolumn optosensor. Because sequential injection systems provide such a versatile system for fluid handling, and they scale well for handling milliliter size to microliter size samples, they have been used in a great variety of analytical approaches beyond the simple example of mixing, reaction, and detection just given. Sequential injection separations (Figure 2c) and bead injection represent two prominent examples, which will be described in more detail below. The use of solid phases in sequential injection systems has recently been reviewed.29

2.2. Preconcentrating Minicolumn Sensors Early work on preconcentrating minicolumn sensors in flow-based analysis was reviewed in 1993.30 Another review on such sensors appeared in 2004.31 Our own surveys indicate there are over 100 papers on such sensors. Generally, these sensors are columns containing sorbent solid phases and fit the general idea shown in Figure 1a, with all or part of the column in the detection zone. In the fluidic systems shown in Figure 2, a and b, the sensor serves as the detector. Many such sensors are spectrophotometric, measuring absorbance of the packed bed in the visible or UV wavelengths. Accordingly, they have been described using terms such as optosensors, optosensing, solid-phase absorptiometry, and ion-exchanger phase absorptiometry. The flow cell may be a modified cuvette fitting in a conventional spectrophotometer, or it may be a flow cell configured with fiber optics. Luminescent methods, such as fluorescence, phosphorescence, and chemiluminescence, and luminescent methods using energy transfer processes have also been widely employed. Even photoacoustic detection has been adapted to flow cells containing sorptive solid phases.32 The solid phases are typically ion exchangers, ligandloaded complexing resins, or hydrophobic phases such as C18-modified silica. Molecularly imprinted polymers have also been used. Ligand-loaded complexing resins have been reviewed.33 The detected analyte may be the “native” ionic

or molecular species, a reaction product or complex formed upstream of the flow cell, or such products formed upon interaction with the solid-phase material. The range of species detected using these methods are extremely broad, including metal ions (including the rare earths), complexes, inorganic anions such as iodide and phosphate, pharmaceuticals and metabolites, nutrients and other food components, oxygen, and aromatic hydrocarbons and phenolics. Typically, the analyte or analyte reaction product from a certain volume of sample is captured on the solid phase, measured, and then released using a suitable reagent after the completion of the measurement. In some cases the interaction is weak enough that the species migrates down the column and is detected as it traverses the optical path. For example, Yoshimura used an ion-exchange resin in an optosensor flow cell to capture copper ions from a 0.17-mL injected sample volume.34 Using a 0.014 M nitric acid concentration, where the distribution ratio was D ) 62000 [mol of copper sorbed/kg of resin]/[mol of copper/L of solution], the copper ions were completely retained and were detected with a spectrophotometer at 800 nm. The sensor was regenerated by perfusion with 2 M nitric acid solution. On the other hand, if the copper ions were captured from 0.28 M nitric acid, where the distribution coefficient was D ) 340, the ions were eluted in a “fairly short time” in additional carrier solution. Recently, an alternative methodology for operation of a preconcentrating minicolumn sensor with optical absorption detection has been described.18 Hexavalent chromium ions were accumulated in an equilibration-based sensing approach, where the entire bed of the anion exchange column sensor was equilibrated with the analyte in the sample by flowing an excess of sample through the column. Once the column was fully equilibrated, the entering and exiting chromium ion concentrations were the same. At trace concentrations (e.g., on the linear portion of a sorption isotherm), the amount retained on the column at equilibration is proportional to the sample concentration. Because it is dynamic equilibrium, pumping a sufficient volume of a blank solution through the column will eventually elute the analyte. This approach can be regarded as a true sensor whose response can go up and down with analyte concentration, rather than representing an assay on a specific volume of sample solution. Furthermore, reagents are not required to regenerate the sensor. The concept of equilibration-based preconcentrating minicolumn sensors will be described in more detail below for sensing the radionuclide 99Tc (as pertechnetate, 99TcO4-).

Radionuclide Sensors for Environmental Monitoring

2.3. Bead Injection and Renewable Surface Sensing As an alternative to methods that elute the analyte from the solid phase, as just described, a methodology has been developed to automatically provide a fresh solid phase for each measurement. Renewable surface sensing using solid phases has also been dubbed “bead injection”.28,35-40 In this approach, the solid phase is again within the observed region of a detector, and most detection methods are optical or microscopic. However, fluidic procedures and specialized flow cells have been developed so that the solid phase is delivered as a liquid suspension to the flow cell, captured within the flow cell, perfused with the sample for interaction and measurement, and then removed or released from the flow cell, all under computer control. In this way, a fresh solid phase with a new surface can be provided for each sample measurement, hence the phrase “renewable surface sensing”. These approaches can be used in both separation and sensing for a variety of species; however, they have been particularly useful in bioanalytical measurements where sensitive biochemical interfaces are involved. Nevertheless, renewable surface separations41,42 and sensing3 have both been described in the field of radiochemistry as well. A variety of flow cells have been designed for implementing renewable surface techniques, including a “jet ring cell” with a moveable tubing end in contact with a transparent plate,35,38,40 a machined flow cell with a moveable solid rod intersecting the flow channel,35,38,43 and a rotating rod design where the angled end of a solid rod intersects an angular flow path in one position but allows beads to pass when rotated 180°.44,45 Methods that can direct the flow and beads toward one channel with a frit or another channel without a frit have also been developed.41,42,46 These will not be described in detail here. Some have been reviewed together in connection with nucleic acid-based analyses.46

2.4. Automated Radiochemical Separation and Analysis Radiochemical analysis is concerned with the determination of radionuclides from a variety of sample matrixes. If the radionuclide of interest cannot be determined nondestructively by detection of a γ-ray emission and identification from the γ-ray energy spectrum, then chemical separations are normally a necessary and critical aspect of the analysis. The radionuclides of interest must be separated from the sample matrix and concentrated for determination by either radiometric or mass spectrometric techniques. The classical methods for performing such separations, including precipitation, solvent extraction, and manual ion exchange, are tedious and time-consuming. Significant advances in separation materials for columnbased separations have simplified radiochemical analysis. At the same time, fluidic and in some cases robotic methods have been developed to automate column-based radiochemical separation and detection. The fluidic automation methods have been based on flow injection and sequential injection methods as described briefly above. In particular, the coupling of sequential injection fluidics to small separation columns lead to “sequential injection separations”, as shown in Figure 2c. In this approach, the sequential injection fluidics provide fluid handling to deliver samples, reagents, and eluants; the column provides selective separations based on the sorbent material in the column; and the analytes that are separated from the matrix and eluted from the column are

Chemical Reviews, 2008, Vol. 108, No. 2 547 Scheme 1

detected downstream. Typically, the separation is based on using a separation chemistry where the radionuclide(s) are retained under the sample load conditions, the sample matrix and unretained radionuclides are removed in a wash step, and then the retained radionuclide(s) are released in one or more steps using a change in solution conditions that greatly reduces the affinity of the radionuclide for the separation material. These separations can rapidly separate individual radionuclides or groups and tolerate significant sample loading. The separation materials for automated radiochemical separations can be conventional ion-exchange resins or more recent extraction chromatography or solid-phase extraction materials. These same separation materials can be used in the development of radionuclide sensors to be described below. Extraction chromatographic materials47 for radionuclide separations, using selective or semiselective extractants impregnated on macroreticular polymer supports, have been developed by Horwitz and co-workers and commercialized by Eichrom Technnologies, Inc.48-52 The uptake properties and chemical selectivities of these materials are well characterized in the literature, and hence the selectivities of sensors can be rationally designed and understood. The chemical structures of the extractants in some of these resins are shown in Scheme 1. A variety of solid-phase extraction materials dubbed “SuperLig” have been developed using “molecular recognition” ligands on solid supports and commercialized by IBC Advanced Technologies (American Fork, Utah).53-56 These ligands are covalently bound to various polymeric or silica gel supports. Automated sequential injection separation can be illustrated with data for 99Tc analysis, the same radionuclide we will use to illustrate the principles of radionuclide sensors. A system similar to that in Figure 2c was set up with a separation column containing the extraction chromatographic resin known as TEVA-resin.9,42,57 TEVA-resin is a macroreticular polymer support impregnated with Aliquat 336, a liquid anion exchanger consisting of a mixture of long-chain quaternary ammonium ions. This resin is known to retain 99 Tc(VII) as pertechnetate under neutral to weakly acidic conditions and to release it in strongly acidic conditions.50 As shown in Figure 3, the majority of the radionuclides in a nuclear waste sample pass through the column in the wash step following sample injection, resulting in a large

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Figure 3. Detector traces from a sequential injection separation system set up to isolate 99Tc as pertechnetate from a nuclear waste sample, using TEVA-resin as the separation material. Figure reprinted with permission from reference 57. Copyright 1998 American Chemical Society.

transient peak as detected by a flow-through scintillation detector. Pertechnetate is retained until a strongly acid eluant solution releases it, resulting in a small 99Tc peak seen in the inset. Flow injection and sequential injection separations using extraction chromatographic separations have been developed for a variety of radionuclides.3,41,57-65 Automated radiochemical separation methods have been reviewed.9,42

3. Radionuclide Sensors 3.1. Challenges of Radionuclide Sensing in Water Prior to the development of preconcentrating minicolumn sensors for radiochemical sensing, there had been very little development of radiochemical sensors suitable for rapid and selective quantification of β- and R-emitting radionuclides in water or process streams. Thus, although there were many radioactivity detectors, there were not any selective radiochemical sensors. This state of affairs is evident, for example, in a review of “Emerging Technologies for Detecting and Measuring Contaminants in the Vadose Zone” in the Handbook of Vadose Zone Characterization and Monitoring, published in 1994.66 This review contained a section on “Radiochemical Sensors”, yet it was notably lacking in any examples of such sensors. Instead, it discussed various ways of detecting and analyzing for radionuclides and heavy metals, including general radioactivity detection techniques, inductively coupled plasma mass spectrometry (ICPMS), inductively coupled plasma atomic emission spectroscopy (ICPAES), neutron activation analysis (NAA), and X-ray fluorescence (XRF) spectrometry. Radioactivity detection and instrumental analysis techniques such as these are not radiochemical sensors and have significant limitations for field analysis. The article correctly noted that radioactivity detection “usually requires some form of sample preparation to concentrate the radionuclides prior to counting to achieve a reasonable degree of sensitivity”. The article further stated that “much of current analytical work is still done in fixed chemical laboratories using conventional radiochemical analysis” and that “conventional detection techniques... are confined to fixed or mobile laboratories”. Thus, the conventional analytical methods for R- and β-emitting radionuclides in water consist of concentration, separation, and source

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preparation methods prior to either radioactivity counting or mass spectrometry, activities that are largely performed in centralized laboratories. A radionuclide sensor for water monitoring must succeed at achieving results similar to those of multistep laboratory procedures, all in a compact sensor package that operates automatically. The detection and quantification of R- and β-emitting radionuclides in water present a number of basic challenges: the required detection limits are typically extremely low, the particles emitted have short penetration ranges in condensed media, and the R/β decay events in condensed media provide limited spectroscopic information for distinguishing one radionuclide from another (i.e., for selectivity). Detection limit requirements determined by regulations such as drinking water standards or maximum contaminant levels,3 typically defined in radioactivity units, translate into chemical detection limits that are well below parts per billion (ppb) levels. For example, the 33 Bq/L (900 pCi/L) drinking water standard for 99Tc67-69 translates to 0.05 µg/L , which is the same as 0.05 ppb. Required mass detection limits for other radionuclides such as 90Sr, 129I, and various transuranic actinides are from 1 to 6 orders of magnitude lower. Consequently, chemical detection with a sensor is simply not feasible; radiometric detection methods are required for measurement at and below the standards-based requirements. In addition, chemical detection alone does not distinguish between stable isotopes (that may be natural) and radioactive isotopes of concern. Uranium is an exception to this conclusion, since detection at the required tens of ppb is feasible with chemical sensing approaches, such as stripping voltammetry.70-72 Taking radiometric detection as a given, the properties of R and β emissions in water must then be considered. In contrast to γ rays, which are characterized by relatively long mean-free paths through solid and liquid media, β and especially R particles are characterized by short ranges and rapid energy dispersion in condensed media. For example, the ranges in water for β particles emitted by 90Y (Emax ) 2282 keV), 90Sr (Emax ) 546 keV), and 99Tc (Emax ) 294 keV) are 1.1 cm, 1.8 mm, and 750 µm, respectively. The range of a 5.5 MeV R particle emitted by 241Am is only 47 µm in water. Furthermore, the energies of R and β particles detected in liquids do not provide well-resolved energy spectra that can be used for selective radionuclide identification. β particles are emitted with broad energy spectra, the β spectra of different radionuclides are not well separated, and the particles lose energy as they travel through water or other condensed media. Although R particles are emitted with characteristic energies, detection by scintillation does not provide adequate energy resolution for selective individual detection of R emitters, and again, the particles lose energy as they travel through water. High-resolution R spectroscopy requires preparation of very thin counting sources placed in a close proximity to a solid-state diode detector, typically in vacuum. Even then, radiochemical separations are required to overcome R energy peak overlap problems and interferences from the sample matrix. The challenges listed above lead to a number of requirements for radiometric sensors for R and β emitters in water. (1) Due to the short radiation travel distances, the species of interest must be spatially localized within a detector volume of well-defined geometry in close proximity to the transducing medium. Localization can be achieved by sorbing the species in a material that is in close proximity to the transducing medium. Typically this transducing medium is a scintillating material, although semiconductor diodes may

Radionuclide Sensors for Environmental Monitoring

also be used. (2) Due to insufficient energy information for discrimination, the method for localizing the analyte must also be selective for particular species and separate them from potentially interfering radionuclides. (3) Due to the challenging detection limit requirements, the species must be collected from a large sample volume and preconcentrated. The preconcentrating minicolumn sensor configuration meets these requirements. The flow-based sensor comprising the sensing material in combination with the radioactivity detection method captures the analyte from the matrix, does so according to its selectivity, and can achieve very low detection limits. This approach achieves the same preconcentration and separation results that would conventionally be obtained as the result of a multistep, manual procedure. In terms of the automated radiochemical methods described above, the preconcentrating minicolumn sensor combines the separation column and the radioactivity detection, as shown in series in Figure 2c, into a single functional unit as shown in Figure 1b.

3.2. Minicolumn Sensors Based on Extractive Scintillating Resins Using a preconcentrating minicolumn sensor for radionuclide detection via scintillation requires that the column have dual functionality. It must provide chemical selectivity for capture and separation of the radionuclide of interest, and it must scintillate. Given the short ranges of the R or β particles, these functions are most readily achieved by (1) creating a column packing where the packing medium has dualfunctionality or (2) combining and intimately mixing scintillating media with selectively sorbent media in one column. The latter approach, which we call a composite bed column, will be described in the next section. The creation of polymeric beads with both ion-exchange and scintillating properties was reported over 40 years ago;73 however, there was practically no follow-up related to radiochemical analysis. The collection of radionuclides from the sample and subsequent counting were performed manually in separate steps; consequently, this did not yet represent a sensor. In 1994, the oxidation of a scintillating plastic to create ion-exchange sites was described.74 The purpose of this work was to create a tool for studying ion exchange, where the β particles from 45Ca (simulating Ca in hard water) would result in a signal when they were absorbed. The creation of dual-functionality materials for radiochemical analysis and sensing was pursued in the late 1990s by independent teams at Clemson University and Pacific Northwest National Laboratory. An investigation by DeVol et al. into the adsorption of uranium ions onto CaF2:Eu scintillator particles (Eu-doped CaF2) represented a combination of sorption and scintillation in a flow cell.1 Detection efficiencies of 60% were reported, and scintillation pulse height spectra were measured. Detection efficiency, Ed, is the ratio of observed counts to the number of decay events that occur within the detector. Subsequently, these authors described scintillating glass beads coated with organic extractants for the detection of radionuclides.2 These investigators also reported impregnation of polymer beads with extractants and scintillators for radionuclide sensing.75 Similarly, Egorov and co-workers described impregnation of both extractant and scintillating fluors in polymer beads, and characterized the analytical performance of the material in minicolumn sensors for 99Tc.3 Both groups now have a number of publications, sometimes jointly, on dual-functionality, preconcentrating minicolumn sensors for radionuclides.2-18

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The separation properties of extractive scintillating resins were modeled after the extraction chromatographic resins used for radiochemical separations. Scintillating properties could be obtained by co-impregnating the resins with fluor molecules. For example, the fluor 2,5-diphenyloxazole, PPO, has been used as a primary fluor which captures energy deposited in the polymeric material and subsequently emits light. It has been combined with 1,4-bis(2-methylstyryl)benzene, bis-MSB,3 or with 1,4-bis(4-methyl-5-phenyloxazol2-yl)benzene, DM-POPOP,4 as a secondary fluor to shift the emitted wavelength. Alternatively, resins have also been prepared with 2-(1-napthyl)-5-phenyloxazole, R-NPO, as the primary fluor, without a secondary fluor. The chemical extractant is chosen according to the radionuclide to be retained and detected. Structures of the fluors used in developing extractive scintillating resins are shown in Scheme 2. Egorov et al. described a sensor for technetium based on co-impregnating macroreticular acrylic polymer beads with Aliquat 336, PPO, and bis-MSB.3 The Aliquat 336 is a liquid anion exchanger, as noted above, for pertechnetate separations. Its use for an automated SI separation of 99Tc was shown in Figure 3. Characterization results for the dualfunctionality material are shown in Figure 4. Analyte retention, shown as the capacity factor, k′,76 in the upper plot, is very high in low acid to neutral conditions, which is favorable for uptake from groundwater. At higher acid concentrations, pertechnetate is released; hence, acidic solutions can be used to regenerate the sensor. These results are consistent with the known uptake characteristics of Aliquat 336. The lower plot in Figure 4 shows the instrumental pulse height spectra of 99Tc obtained using the selective sensor material (trace A) in a static liquid scintillation spectrometer. This result is compared with the 99Tc spectrum in liquid scintillation cocktail (trace B). The luminosity of the sensor material is lower than that of the liquid scintillator, but the detection efficiency remains sufficiently high (56%) for practical analytical applications. This sensor material was packed into a minicolumn flow cell which was placed between the photomultiplier tubes of a Packard Radiomatic 515A flow-through scintillation detector. In this configuration the detection efficiency Ed (observed counts divided by the radioactive events from 99Tc quantitatively captured on the column) was 45%. The sensor flow cell was configured as part of a computer controlled sequential injection fluidic system for sample and reagent delivery. Detector traces illustrating selective 99Tc sensing are shown in Figure 5. Analyte capture and measurable scintillation light output are observed upon injection of an aliquot of 99Tc standard (duplicate traces shown as A) in dilute acid. The signal persists as the sensor column is washed with dilute

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Figure 4. Analyte uptake and scintillation properties of a dualfunctionality sensor material for 99Tc. (Top) Plot of the sensor material capacity factor as a function of nitric acid concentration. (Bottom) Pulse height spectra for the dual-functionality sensor (A), and liquid scintillation spectrometer (B). Figures reprinted with permission from reference 3. Copyright 1999 American Chemical Society.

acid. By contrast, radioactive species that do not have a high affinity for the sensor material appear as only transient peak signals, and are promptly removed from the system using a small volume of wash solution. This is illustrated with 137 Cs in Figure 5 (trace C). In the presence of interferences, the light output integrated over time after the wash step provides a quantitative measure of the 99Tc in the sample. If interferences are not a problem, the analyte can be quantified from either the slope of the uptake signal or from the steady-state signal as shown in the calibration traces in the middle plot of Figure 5. Each standard was measured using a freshly packed column followed by release of the packing after the measurement, as in a renewable column sensor. The lower plot in Figure 5 illustrates detection of 99Tc in actual groundwater from the Hanford nuclear site. The water was acidified to pH 2 and analyzed with and without a spike. The detection limit for this sensor was estimated to be 6.2 Bq/L (167 pCi/L) or 0.0098 ppb, based on a 50-mL sample size and a 30-min signal accumulation. This detection limit is well below the 33 Bq/L (900 pCi/L) drinking water standard for 99Tc.67-69,77 These results illustrate the extremely low detection limits that can be achieved with a preconcentrating minicolumn radiometric sensor. Because the sensor material exhibits high binding affinity toward pertechnetate, large sample volumes can be preconcentrated using a small sensor column. DeVol et al. investigated a wide range of dual-functionality column packing materials for on-line and off-line radionuclide measurements.4 Subsequent studies examined these materials in greater detail and introduced additional materials.5-7,12-14,17 These studies included a number of extrac-

Figure 5. (Top) Sensor response to 99Tc(VII) analyte and a potentially interfering species (137Cs) unretained by the sensor material. Flow rate 1 mL min-1, injected sample volume 0.1 mL. Following the injection the sensor bed is washed with 10 mL of 0.02 M nitric acid. (Middle) Calibration traces for 99Tc sensing. (a) Sample load step (10 mL), (b) sensor wash step (10 mL), and (c) ejection of the sensing material from the column. (Bottom) Detector traces from the analysis of acidified Hanford groundwater (GW 1) sample. Flow rate used was 2 mL min-1; (a) sample load step (50 mL); (b) sensor wash step using 5 mL of 0.05 M HNO3; and (c) 30-min stopped-flow counting interval. Time zero corresponds to the beginning of the sample load step. Figures reprinted with permission from reference 3. Copyright 1999 American Chemical Society.

tive scintillating materials, extractant-coated glass scintillator particles, and a heterogeneous mixture of plastic scintillator beads with extraction chromatographic resin. The extractive scintillator materials were prepared by impregnating acrylic or styrene-based polymer beads with PPO and DM-POPOP, followed by impregnation with the extractants of interest. Extractive scintillating materials developed for radionuclide sensors are summarized in Table 1.

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Table 1. Scintillating Extractive Resins for Flow-Cell Minicolumn Sensors fluor or scintillator PPO, bis-MSB PPO, DM-POPOP R-NPO PPO, 9,10 diphenylanthracene PPO, DM-POPOP PPO, bis-MSB PPO, DM-POPOP PPO, DM-POPOP PPO, DM-POPOP PPO, DM-POPOP R-NPO PPO, 9,10 diphenylanthracene GS-20 scintillating glass GS-20 scintillating glass

extractant Aliquat 336 Aliquat 336 Aliquat 336 MnO2 “ABEC” Me-PEG-2000 HDEHP crown ether DtBuCH18C6 crown ether DtBuCH18C6 CMPO/TBP CMPO/TBP H2DEH[MDP] H2DEH[MDP] H2DEH[MDP] crown ether DtBuCH18C6

Materials for 99Tc detection were developed based on two extractants, Aliquat 336 and a monomethylated-polyethylene glycol (“ABEC”). The latter is known to bind pertechnetate from certain high ionic strength solutions which is useful in the analysis of nuclear waste or waste-processing streams.78 Resins with bifunctional organophosphorus extractant octyl(phenyl)-N,N-diisobutylcarbamoylmethylphosphine oxide (CMPO) in tri-n-butyl phosphate (TBP) were investigated for actinide retention and sensing. These were modeled after Eichrom TRU-resin.48,79 In later work,13 another material for actinide detection was developed based on bis(2-ethylhexyl)methane-diphosphonic acid, which is abbreviated as H2DEH[MDP] and also known as Dipex. This material is modeled after Eichrom Actinide Resin.80 For 90Sr sensing, the crown ether4,4′(5′)-bis(tert-butylcyclohexano)-18-crown-6(DtBuCH18C6) was used. This material was modeled after Eichrom Sr-resin, which contains the crown ether in 1-octanol solution impregnated in a polymeric resin.49,81-85 This material selectively binds strontium from nitric acid solutions as Sr(NO3)2(DtBuCH18C6). In separate work, 2-ethylhexylphosphoric acid, HDEHP, was used in a prototypical 90Sr sensor with uptake from 0.001 M HCl.3 These various extractive, scintillating resins were evaluated for the efficiency with which radionuclides were captured, the efficiency with which the radionuclides were recovered from the column by elution, and the detection efficiencies.4 Values from 30 to 100% were found for the detection efficiencies in the initial study,4 which are all suitable for development of sensors. A number of these were further investigated and tested against groundwater, synthetic groundwater samples, or nuclear waste samples.5-7,12-14,17 For example, further studies of 99Tc sensing were carried out using Aliquat 336 as the sorbent.5 The fluors, PPO and bis-MSB, were diffused into the macroporous acrylic-based (Amberchrom CG-71t2) polymer beads in a separate step prior to impregnation with the extractant. This resin in minicolumn format was used to analyze six contaminated groundwater samples from the Hanford site. The samples were acidified to pH 2 prior to analysis, and quantification was carried out using standard addition methods. Each measurement involved analysis of the following solutions: (1) reagent blank (0.01 M HCl or 0.02 M HNO3, depending on acid used for sample acidification), (2) acidified groundwater sample, and (3) spiked acidified groundwater sample. All solution delivery steps were performed at 2-mL min-1 flow rate, delivering 50-mL sample aliquots, and washing the flow cell with 5 mL of 0.02 M HNO3. Then the flow was stopped, and the count rate was determined over a 30min counting interval. Analysis results obtained using the flow-cell sensor and standard radiochemical methods were in excellent agreement. The minimum detectable concentra-

support macroreticular acrylic polymer macroporous styrenic polymer macroporous styrenic polymer polyvinyltoluene macroporous styrenic polymer macroreticular acrylic polymer macroporous acrylic polymer macroporous styrenic polymer macroporous styrenic polymer macroporous acrylic polymer macroporous styrenic polymer polyvinyltoluene GS-20 scintillating glass GS-20 scintillating glass

analyte 99Tc 99Tc 99Tc

U 99Tc 90Sr 90Sr 90Sr actinides actinides actinides actinides actinides 90Sr

refs 3,5 4,5 14 14 4,5 3 6 4,6 4,7, 12,17 4 13 13 2,4 2,4,6

Figure 6. 99Tc sensing in Paducah groundwater, acidified to 0.1 M nitric acid, using a preconcentrating minicolumn sensor in portable detection instrumentation. Volumetric flow rate of ∼1 mL min-1 results in a measured slope of 8.26 cps/L. Results are replotted from data published in reference 87.

tion (MDC) of the flow-cell procedure was calculated with the Currie equation,86 using average background levels and analyte loading and detection efficiencies. For the 50-mL groundwater samples and 30-min counting interval, the MDC was reported to be 6 Bq/L. An additional pertechnetate-selective resin was prepared by co-immobilization of the R-NPO fluor with Aliquat 336 extractant within an inert macroporous polystyrene resin. This resin was used in combination with portable and transportable instruments to demonstrate these sensors as potential field screening tools.14,87,88 A Hidex Triathler field portable scintillation counter was modified with a sensing flow cell containing a small coil of Teflon tubing containing the resin. The detection efficiency measured for this sensor instrument was ∼30%. Figure 6 illustrates the Triathler instrument response obtained during loading of a 400-mL quantity of acidified 99Tc-contaminated groundwater from the Paducah site. These data were collected by recording the count rate of the detector in 100-s intervals while the groundwater was continually pumped through the flow cell. A 99Tc activity of 26.8 Bq/L was determined from the slope of the count rate, in reasonable agreement with an independent radiochemical measurement of 22.0 Bq/L. In addition to the experiments with the Triathler, a minicolumn, flow-cell detection system was designed around an Eberline E-600 survey meter and a modified photomultiplier tube (PMT) housing. The advantage of this configuration is portability, but the disadvantage is the low detection efficiency of about ∼2%. Preconcentrating minicolumn sensors for radiostrontium were prepared using a crown ether chemistry (DtBuCH18C6) to concentrate the radioactive ions of interest.6 The polymeric resins (both acrylic and styrenic resins were investigated) were impregnated with PPO, DM-POPOP, and DtBuCH18C6

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Figure 7. Repeated loading and regeneration of a Sr-selective extractive scintillating resin. 89Sr was loaded in 4 M nitric acid and released in water.

in 1-octanol. Packed flow cells were monitored, and the signal was quantified using a commercially available scintillation detection system, the IN/US Beta-Ram model 1. Strontium ions were captured from 4 M nitric acid solutions with efficiencies of 99-100% and released in water, which is consistent with the known extraction chromatographic separation chemistry of DtBuCH18C6. Regeneration capability was demonstrated through multiple loading/elution cycles with 89Sr and 90Sr. Repeated capture and release of 89Sr are shown in Figure 7. There was only a slight degradation in the detection efficiency (59.7 ( 2.97% for 89Sr) over time. The detection efficiency for 90Sr was lower than that for 89 Sr because 90Sr has a weaker β energy. Detection of 89Sr was demonstrated at concentration as low as 120 Bq/L. Lower minimum detectable concentrations would require (1) that the detection efficiency be higher, (2) that the detection system background count rate be lower, (3) the use of longer count times, and/or (4) replacement of the extractant with one which has a higher distribution coefficient so that analyte can be collected from a larger sample volume without breakthrough. For uptake of 90Sr from weakly acidic to neutral solutions, a scintillating extractant resin was created based on HDEHP extractant. In this case, both 90Sr and its daughter product 90 Y were extracted from 0.001 M HCl.3 However, they could be individually determined by selective elution of the 90Sr using 0.2 M HCl, leaving the 90Y on the column. Finally, 90 Y was eluted with 4 M HCl to regenerate the column. These steps are shown in Figure 8. Assuming 100% efficiencies of the capture and elution procedures, the individual absolute detection efficiencies, Ed, for 90Sr (Emax ) 546 keV) and 90Y (Emax ) 2282 keV) were determined to be 46 and 99% respectively. Thus, co-retention and selective elution steps can be used to quantify individual radionuclide species. The method of selective uptake of a group of radionuclides with selective elution steps has also been demonstrated in actinide sensing. The extractive scintillating resin containing fluors PPO and DM-POPOP was impregnated with CMPO in TBP.7 The resin was packed in an FEP Teflon tubing flowcell coil and placed into a dual photomultiplier tube coincidence detection system to obtain pulse height spectra and time-series data. Loading and elution experiments were conducted with 241Am (9.8 Bq), 239Pu (7.4 Bq), and 233U (10.2 Bq) as illustrated in Figure 9, where on-line sensor detection results are shown in plot A (top) along with the eluted radionuclides in plot B, as determined by fraction collection and liquid scintillation counting. The data shown in Figure 9, a and b, are complementary and were obtained essentially simultaneously using the same column. The intervals when

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Figure 8. Detector traces for 90Sr/90Y separation experiments using HDEHP-based scintillating extraction resin. The sample was 100 µL of 1.57 × 104 dpm 90Sr/90Y standard in 0.001 M HCl. Flow rate was 1 mL min-1; (a) sample injection and sensor wash step using 6 mL of 0.001 M HCl; (b) 90Sr elution step using 6 mL of 0.2 M HCl; (c) 90Y elution using 6 mL of 4 M HCl. Figure reprinted with permission from reference 3. Copyright 1999 American Chemical Society.

Figure 9. Detection of actinides 241Am, 239Pu, and 233U in the mixed actinide solution with selective elution in a dual functionality sensor (Plot A), and the chromatogram of the eluents as determined by fraction collection and counting (Plot B). Results are replotted from data published in reference 7.

solutions were pumped through the column are indicated by the lettered boxes along the x-axis between the plots, where interval “a” is the sample load in 2 M nitric acid. The captured actinides produce a counting plateau corresponding to the total actinide count rate of 24.88 cps for a 10-min counting time. Trivalent actinides were eluted with 4 M hydrochloric acid (e.g., 241Am during interval “b”). A solution 0.02 M TiCl3 in 4 M hydrochloric acid was applied during interval “c” to reduce the plutonium and elute it from the resin. After a column rinse with 2 M hydrochloric acid during interval “d”, uranium-233 was eluted with 0.1 M ammonium bioxalate during interval “e”. The count rate difference divided by the detection efficiencies for 241Am, 239Pu, and 233 U (96.5%, 77.5% and 96.6%, respectively) constitute the

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Table 2. Composite Bed Minicolumn Sensors scintillator BC-400 BC-400 GS-20 BC-400 yttrium silicate (YSO) BC-400 BC-400 BC-400

selective chemistry Aliquat 336 crown ether DtBuCH18C6 crown ether DtBuCH18C6 H2DEH[MDP] H2DEH[MDP] anion exchange, strongly basic anion exchange, weakly basic SuperLig 620 solid-phase extraction material

measured activity and were within 10% of the expected values. One advantage of this system is that quantification can be accomplished at low activities because long count times can be obtained by extending the time between eluents. This sensor resin was also applied to a digested high-level waste sludge and high-activity drain tank samples where the agreement between the on-line and off-line analyses was within 35%. The extractive scintillating resin and detection system just described was subsequently applied to low-level uranium concentration determination in acidified groundwater.12 Using Dipex extractant, a scintillating extraction resin containing R-NPO fluor was demonstrated for monitoring natural uranium in groundwater.13 This resin was packed into a flow cell designed for a modified Hidex Triathler. The average detection efficiencies were 51.7 ( 2.6% and 65.8 ( 10.1% for natural uranium and 241Am, respectively. The resin was stable for sample load volumes of up to 1000 mL, resulting in rapid real-time quantification of natural uranium in groundwater down to 30 µg/L, which is sufficient to meet the standard established in the U.S. Safe Drinking Water Act. In summary, extraction chromatographic materials containing organic scintillator fluors were developed and demonstrated in preconcentrating minicolumn sensors for uptake and detection of a variety of radionuclides, including 99 Tc, 90Sr, actinides, and natural uranium. The uptake characteristics are similar to those of the extraction chromatographic materials without the fluors, and detection efficiencies were good. Accurate detection in Hanford groundwater samples was demonstrated for 99Tc, and it could be detected to below drinking water standard limits. Detection of uranium in groundwater to drinking water standard levels was demonstrated. and samples up to at least one liter volume could be preconcentrated. Nevertheless, extractive scintillating materials have some potential drawbacks related to their stabilities under repeated or long-term use. The impregnated extractants can leach out. Under some conditions, chemiluminescence signals from the scintillating fluors are observed, which can interfere with the radiochemical measurement. In addition, the scintillating fluors are quenched or damaged and may lose their scintillating properties in the presence of some of the acid reagents used. It was noted, for example, that attempts to elute pertechnetate from resins that were co-impregnated with Aliquat 336 and the PPO/bis-MSB fluor combination, using 4 M nitric acid, resulted in a significant chemiluminescence signal and loss of scintillating properties in subsequent sensing experiments.3 Fluor impregnation methods based on diffusion or synthesis yield sensors with greater stability for sequential sensing experiments, but long-term stability remains a concern. The best combination of fluor and support for the extractive scintillating resin of the variations tested is the

solid phase TEVA resin Sr-resin Sr-resin Actinide resin Actinide resin AGMP1 AG 4-X4 silica gel-based solid-phase extraction material

analyte 99

Tc Sr 90 Sr actinides actinides 99 Tc 99 Tc 90 Sr 90

refs 4,5 6 6 13 13 8-11 16,18 10,18

R-NPO fluor impregnated into the macroporous polystyrene resin. For the combination of fluor and support to respond like a scintillator there needs to be good energy transfer from the support, which is the bulk of the scintillator, to the fluor. This π-electron energy transfer is more efficient with the polystyrene support. The procedure for producing the scintillating resin results in some fluor diffusing into the polystyrene and some just being retained within the pores of the resin. In the latter case, the fluor can interact with reagents and eluants, and may be leached out. Of the fluors evaluated, R-NPO resulted in little to no leaching from the polystyrene resin.

3.3. Composite Bed Scintillating Minicolumn Sensors Composite bed columns, consisting of a heterogeneous mixture of scintillating particles and chemically selective particles, represent an alternative to the extractive scintillating resins described above. Composite bed column materials are listed in Table 2. In most cases, the scintillating component consists of Bicron BC-400 beads, which are poly(vinyltoluene) scintillating plastic beads. These nonporous scintillating beads were found to have high chemical stability in sample and regeneration solutions. A further advantage of the composite bed sensing approach is that it facilitates the use of existing extraction chromatographic, ion-exchange, or solid-phase extraction materials for the chemically selective sorbent component of the bed. Given an intimate mixture of the sorbent particles and the scintillating particles, the R or β emission from a radionuclide captured on a sorbent particle has a reasonable probability of colliding with a neighboring scintillating bead, provided the travel distance of the radiation in the condensed phases is large enough in comparison to the distance across column pore spaces (in between particles) and particle diameters. Higher ratios of scintillating particles to sorbent particles increase this collision probability, increasing the detection efficiency, at the cost of reduced sorbent material for radionuclide capture in the composite bed. This approach is particularly well suited to detection of radionuclides that emit β particles, which have longer ranges in condensed media than typical R particles. For example, the β particle from 99 Tc has an estimated maximum range in water of 750 µm, which is greater than typical sorbent particle sizes in the 20200 µm size range. A number of composite bed columns were demonstrated for mixtures of extraction chromatographic resins and BC400 beads (100-200 µm), including those based on TEVA resin for 99Tc sensing,4,5 Sr-resin for 90Sr sensing,6 and Actinide resin for actinide measurements.13 Composite bed approaches were also developed using conventional anionexchange materials for 99Tc sensing.8-11,16,18

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For example, a composite bed flow cell was prepared by mixing equal masses of TEVA extraction chromatographic resin (100-150 µm) and BC-400 plastic scintillating beads (100-200 µm).5 The mixture was packed into a flow cell, and responses to 99Tc were quantified using an IN/US BetaRam model 1. The detection efficiency of 99Tc ranged from 7.5% to 16.4% during the subsequent performance tests. The uptake of 99Tc was quantitative. In addition to characterizing responses to 99Tc, tests to evaluate potential inference by 137 Cs, 90Sr, and 239Pu were investigated. Typically, the column was first exposed to a 99Tc standard, followed by an injection of a potential radioactive interference. A statistically significant increase in the sensor count rate after loading and washing the interference, while the 99Tc remained captured, was defined as an interference. For example, after loading the sensor with 5 mL of 24 Bq/mL 99Tc standard, resulting in a measurable steady-state signal, 1 mL of a 7200 Bq/mL 137 Cs standard was loaded and subsequently washed with additional carrier solution (2 M HCl for standards and wash). A transient peak was detectible while the 137Cs progressed through the flow cell, but no interference was detectable after washing. A trace amount of 137Cs was detectable in the eluant when the 99Tc was released in 8 M nitric acid. Interference trials with 90Sr and 239Pu at 54 and 240 Bq/mL, respectively, resulted in no detectable interference nor were these radionuclides detected in the final 99Tc eluant. At a higher activity of 24,000 Bq/mL, 90Sr resulted in a slight interference. Composite bed columns using conventional ion-exchange resins were demonstrated using AG MP-1, a strongly basic anion-exchange resin that has very high uptake affinity and good selectivity toward Tc(VII) ions in basic to weakly acidic media. The weight/volume distribution coefficient (Kd) for 99 Tc(VII) on AG MP-1 was 2.5 × 105 mL/g for 99Tc(VII) in unacidified Hanford groundwater. As a result, small volumes of the sorbent material can be used to preconcentrate analyte from large volumes of groundwater. A composite sensor bed was prepared by mixing 200-400 mesh AG MP-1 material (particle size about 40-70 µm) with BC400 plastic scintillator beads (particle size 100-250 µm) at a 30:1 weight ratio of scintillator to sorbent.8,10 The total bed volume was just 50 µL. The absolute detection efficiency of the composite sensor column was 34%, and analyte loading efficiency was 97%. This sensor was demonstrated to be effective in capturing 99Tc(VII) in unacidified Hanford groundwater; up to 60 mL of the groundwater sample could be preconcentrated without analyte breakthrough using this very small sensor column.9,11 The 99Tc(VII) selective composite bed sensor can be regenerated using a small volume of 2 M nitric acid solution, resulting in rapid elution of the retained analyte without loss of the scintillation properties.8,10 Composite bed sensor columns have also been prepared using the weakly basic anion-exchange resin AG 4-X4.16,18 The uptake affinity of this material for pertechnetate in weakly acidic to weakly basic conditions is substantially lower than that of the AG MP-1. Nevertheless, it has sufficient uptake to collect and preconcentrate pertechnetate, and it offers other practical advantages. The weakly basic anion exchanger is less prone to irreversible uptake of soil organic matter, such as humic acids, and pertechnetate can be readily eluted with sodium carbonate solutions, an environmentally benign reagent that provides an alternative to the use of strong acids for regeneration.16 It should be noted that high ratios of scintillator to sorbent (as was used in the AG MP-1 studies) are not necessary to

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Figure 10. Detection efficiency for 99Tc as a function of the dry scintillator to sorbent ratio for a composite bed sensor column. Scintillator is BC-400 (100-250 µm); sorbent is AG 4-X4 (100200 mesh).

obtain satisfactory detection efficiencies. This high ratio was motivated in part by the extremely high affinity of the strongly basic anion exchanger AG MP-1; dilution with a high ratio of scintillator reduces the overall retention volume of the column for such a high-affinity sorbent. The effect of ratio on detection efficiency was examined in more detail for weakly basic anion exchanger AG 4-X4. Various mixtures of 100-200 mesh (75-150 µm) AG 4-X4 and 100-250 µm beads of BC-400 were prepared. Detection efficiencies from 20 to 60% were found for scintillator to sorbent ratios of 1.5 to 14 (dry weight to dry weight) as shown in Figure 10. Sensing results using composite bed sensors based on AG 4-X4 will be described in more detail in connection with equilibration-based sensing below. BC-400 plastic scintillator beads have been mixed with solid-phase extraction material SuperLig 620 for the development of 90Sr sensors.18 Using a 1:1 weight ratio, a detection efficiency of 63% was reported. It was anticipated and observed that a higher detection efficiency could be obtained for 90Sr than 99Tc because the former has a higher-energy β emission.

3.4. Sensor Regeneration or Renewal Much like the preconcentrating minicolumn sensors developed in flow injection analysis, the radionuclide sensors described above are based on quantitative capture of the analyte from a certain volume of sample. Typically, the sensor is regenerated by passing a solution over the sensor such that the distribution coefficient between the stationary phase and mobile phase is significantly decreased and the analyte is released from the column and washed away. Then the sensor is ready for another sample. The chemistry used for radionuclide elution and sensor regeneration is typically taken directly from the elution conditions developed for extraction chromatographic purification of radionuclides in radiochemical analysis. Examples of analyte elution and sensor regeneration are shown in Figures 7 and 8 for 90Sr sensing. As described above, pertechnetate that is retained on anion-exchange materials at neutral to low pH can be released using more acidic solutions. However, an alternative approach based on the use of sodium carbonate was demonstrated for pertechnetate release from weakly basic anion-exchange materials.16 This environmentally benign reagent was proposed to be compatible with groundwater sensing applications. The alternative to elution or regeneration reagents is to replace the entire column packing using renewable, surfacesensing techniques as described above. This approach was proposed in one of the early examples of radionuclide sensing.3 No liquid reagent solutions would be required for

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elution in this approach, but it would require a supply of sensing material to be delivered in suspension to the flow cell. One advantage of this approach is that some of the typical concerns associated with reusing sensor materials or layers, such as long-term stability, reversibility, degradation, fouling, and potential analyte carryover from sample to sample, are also alleviated.

3.5. Equilibration-Based Sensing As just noted, the preconcentrating minicolumn sensors for flow injection and for radionuclides relied on quantitative capture of analyte and subsequent regeneration with reagents. While this approach provided new radionuclide sensors for water monitoring and succeeded at meeting stringent detection limit requirements, the use of reagents to regenerate the sensor column for each and every measurement is a potential drawback for in situ monitoring applications. In addition, as discussed above, this approach represents an assay on a sample volume, as opposed to a sensor that responds to changes in sample concentration. To address these issues, a new modality for the preconcentrating minicolumn sensor was developed that we call “equilibration-based” sensing.8-11,16,18 The equilibration-based approach sets out to deliberately achieve full breakthrough conditions where the analyte concentration exiting the column is the same as the analyte concentration entering the column. Under these conditions, the sensing material in the column has equilibrated with the analyte concentration in solution. At low concentrations typical of trace detection applications, the linear portion of the sorption isotherm applies, and the amount captured on the column material is proportional to the analyte concentration. The sorbent material is not “saturated”; it is equilibrated. The amount of analyte collected has not reached the total capacity of the sorbent (as would be the case at higher concentrations corresponding to the plateau of the sorption isotherm). The key features of the equilibration-based approach are: (1) a steady-state response once the sorbent phase is equilibrated, (2) a response that varies with the analyte activity or concentration, and (3) reversibility of the response because it is based on dynamic equilibrium. When a sample containing a different analyte concentration or activity is pumped through the column, the phase will re-equilibrate, and the signal will go up or down accordingly. If this sample is blank, the signal will go down as if the column were regenerated with a reagent. Thus, in principle at least, no consumable reagents are required in this equilibration-based approach. The responses of a 99Tc sensor using this approach are shown in Figure 11. All three key features just mentioned are apparent in this figure. Steady-state responses are obtained while pumping 225 mL samples through the column, and a final blank shows the reversibility. The sensor response levels and the calibration curve in the inset illustrate how the signal varies with sample activity. The activity of the lowest level standard is equivalent to the drinking water standard for 99Tc. The column in this case had internal dimensions of 4 mm i.d. × 29 mm length for a bed volume ) 0.364 mL. The experimentally determined retention volume, Vr , was 81 mL. Retention volume is equivalent to the sample volume containing the same quantity of analyte as the fully equilibrated sorbent phase in the column. Comparison of the

Figure 11. Responses of a composite bed (1:4 w/w ratio AG 4-X4: BC-400) preconcentrating minicolumn sensor to samples at increasing activities of 99Tc standards (in 0.01 M nitric acid) in the pertechnetate form, followed by a final blank sample. The data were obtained by delivering 225-mL aliquots of the following solutions at 2-mL min-1 syringe pump flow rate: (1) blank sample, (2) 0.033 Bq/mL; (3) 0.13 Bq/mL; (4) 0.34 Bq/mL; (5) 0.65 Bq/mL; (6) blank sample. Figure reprinted with permission from reference 18. Copyright 2006 American Chemical Society.

Figure 12. Detector trace showing responses of the 90Sr sensor column (1:1 w/w ratio SuperLig 620:BC-400) to 90Sr standards in Hanford groundwater acidified to pH ∼2.1 with nitric acid: (1) 1.02 Bq/mL; (2) 0.33 Bq/mL; (3) 0.10 Bq/mL; (4) blank sample. Figure reprinted with permission from reference 18. Copyright 2006 American Chemical Society.

retention volume with the column bed volume illustrates the high degree of preconcentration achieved with this sensor (a factor greater than 200). The experimentally determined column theoretical plates, N, was 12, and the detection efficiency, Ed, was 38%. Whereas the 99Tc column sensor was based on anionexchange chemistry, a sensor for 90Sr was developed using a solid-phase extraction sorbent, SuperLig 620, which consists of a silica support with covalently bound crown ether ligands. A composite bed sensor was created with a 1:1 ratio of the SuperLig sorbent to BC-400 scintillating plastic beads. The sorbent is capable of 90Sr uptake from neutral or acidic solutions. However, because the daughter product 90Y is also slightly retained at neutral conditions and unretained under acidic conditions, experiments to illustrate equilibration-based sensing were carried out in samples acidified to pH 2 with nitric acid. The sensor responses to sequential 450-mL sample volumes of acidified groundwater containing 90Sr are shown in Figure 12. The detection efficiency Ed of this composite bed sensor was 63%, and the column theoretical plates were just N ) 3.4. These results illustrate how a different sorbent chemistry can be used to tailor this sensor concept to different radionuclide analytes, and that even with a column of rather low theoretical plates, a satisfactory sensor can be created.

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Figure 13. Schematic illustration of column chromatography concepts. A discrete injection and subsequent Gaussian peak illustrates conventional chromatography in the upper pair of plots. A step function input with a subsequent sigmoidal breakthrough profile is shown in the lower pair of plots illustrating frontal chromatography.

3.6. Chromatographic Theory for Equilibration-Based Sensing The characteristics of the preconcentrating minicolumn sensor in equilibration-based sensing mode, and the attainment of the steady-state response, can be understood in terms of concepts from chromatography. In conventional chromatography with a discrete injection, the analytes are sorbed to the stationary phase in a dynamic process and migrate down the column until they elute with a typically Gaussian peak. This discrete input and the peak-shaped output are shown schematically in Figure 13. The peak shape results from a normal distribution of the velocities with which individual molecules traverse the column length, all starting at essentially the same time. Individual molecules or ions have different net velocities due to random factors including the path through the packed bed, diffusion in various directions while in the mobile phase, and variations in the amount of time each molecule or ion spends being immobilized in the stationary phase. In frontal chromatography, and in the operation of the preconcentrating minicolumn sensor in equilibration-based mode, the input is a step change in analyte concentration, as shown in the lower plots of Figure 13. As analytes migrate down the column, the step input is transformed into a sigmoid-shaped concentration profile in the column effluent.89 In this mode, the analyte molecules or ions still traverse the column with a distribution of individual velocities, but they do not all start at the same time. The sigmoid shaped effluent concentration profile is typically represented by an integral of the Gaussian distribution function.89,90 The inflection point of the sigmoidal breakthrough profile corresponds to the maximum in the Gaussian distribution. In practical terms, this inflection point also corresponds to the retention volume, Vr. These characteristics are shown in detail for the 99Tc sensor in Figure 14.18 The 99Tc activity in the input is first stepped from 0 to 1 Bq/mL for a 250-mL sample volume and then stepped back down from 1 to 0 Bq/mL. This sample volume is clearly more than that which is necessary to equilibrate the column in each step. The sensor response shown in the upper trace indicates the amount of 99Tc captured on the column. Initially the 99Tc is captured quantitatively, resulting

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Figure 14. (A) Trace showing sensor equilibration with 1 Bq/mL 99Tc solution followed by a reequilibration to a blank solution in a matrix of 0.01 M nitric acid. Sensor cell (dimensions 4 mm i.d. × 29 mm height) is composed of a 1:4 dry w/w ratio AG 4-X4/BC400. (B) Fractions collected immediately downstream of the sensor cell show the breakthrough profile of 99Tc from the sensor column. Both the Gaussian model (black lines in each plot) and low plate model (not shown) provide a good fit to the observed data. Figure reprinted with permission from ref 18. Copyright 2006 American Chemical Society.

in a linear increase in the signal. As 99Tc begins to break through, the response begins to level out. The lower plot shows the sigmoid-shaped output profile, i.e., the effluent concentration profile as a function of time or volume, as determined by analyzing collected fractions on the sensor output. The amount accumulating on the column, leading to the sensor response, is the total amount that has been delivered to the column minus the total amount that has broken through and exited the column. When the output concentration equals the input concentration, the amount accumulated on the column stops increasing. After the 250-mL sample load, the input was stepped from the 1 Bq/mL 99Tc activity back down to zero. The sensor signal begins to drop and the effluent concentration also begins to drop, until the retained 99Tc is completely removed from the column. Frontal chromatography theory in equation form has been adapted to derive theoretical equations that express the effluent concentration as a function of volume or time (i.e., the breakthrough curve), and similarly the sensor response as a function of the solution volume or time. The derivation and equations are given in detail in ref 18. The total amount that has exited the column is the integral of the effluent concentration profile, as shown in the first half of Figure 14b, which is itself the integral of a normal distribution (an integral of an integral). The amount that has accumulated on the sensor column, as a function of volume or time, is the amount delivered to the column minus the amount that has exited. While the plot in Figure 14 illustrates step changes from zero to one and one down to zero, the equations can be derived for step changes from any arbitrary concentration to another, and converted to radiometric count rates on the sensor. In equation form, the function f(V) is used to represent the normalized analyte concentration profile in the column

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effluent, i.e., the breakthrough curve as shown schematically in the lower right plot in Figure 13 and the first half of the lower plot in Figure 14. For an input step with concentration changing from initial concentration C0 ) 0 to subsequent concentration C1 (C1 > 0), the effluent concentration as a function of volume, Cef(V), can be expressed as

Cef(V) ) C1f(V)

(1)

where f(V) goes from 0 to 1.89 In conventional linear frontal chromatography with a sufficient number of theoretical plates, the shape of function f(V) is usually assumed to be represented by an integral of a Gaussian distribution:89-91

1 1 f(V) ) + erf 2 2

(( ) x ) V -1 Vr

N 2

(2)

where erf(x) is the error function defined as erf(x) ≡ 2/xπ ∫x0 e-t2dt , and N is the number of theoretical plates. The parameter Vr is the analyte retention volume, as usual. The amount of the analyte that has exited the sensor column after delivering sample volume V, Mel(V), is given by the integral of the effluent concentration,

Mel(V) ) C1 ∫0 f(V)dV V

(3)

Then the amount of analyte, Ms,C1(V), present on the sensor column as a function of the sample volume can be calculated as the amount delivered C1V minus the amount that has exited the column:

Ms,C1(V) ) C1V - C1 ∫0 f(V)dV ) C1V - C1F(V) V

(4)

where the integral of the normalized breakthrough profile is expressed by the simplified notation:

F(V) ) ∫0 f(V)dV V

(5)

For an arbitrary step input with prior and subsequent analyte concentrations equal to C0 and C1 the amount of analyte present on the sensor column as a function of the sample volume with analyte concentration C1 can be expressed as:

Ms,C0,C1(V) ) C0Vr + V(C1 - C0) - (C1 - C0)F(V) (6) where, V is the volume of the sample solution with concentration C1. In radiometric detection, the number of radioactive decay events per second is proportional to the number of analyte atoms, while the fraction of the total decay events being detected is expressed as the absolute detection efficiency, Ed. Therefore, the radiometric sensor response can be expressed by using the following equations:

Rc/s,eq ) EdAaVr Rc/s,C0,C1 ) Ed[A0Vr + V(A1 - A0) - (A1 - A0)F(V)]

(7) (8)

Equation 7 expresses the radiometric count rate, Req, in counts/second (c/s) of a sensor column that is fully equilibrated with a sample containing analyte activity, Aa , in Bq/mL. Equation 8 gives the sensor count rate, Rc/s,C0,C1, as

a function of sample volume for an arbitrary activity step with initial and subsequent activities A0 and A1, respectively. Recalling that F(V) is an integral of f(V) (see eq 5), which depends on the column parameters Vr and N (see eq 2), it follows that the sensor response as a function of volume depends on the column chromatographic parameters, i.e., the retention volume Vr and number of theoretical plates N, and the detection efficiency Ed. In turn, the retention volume Vr depends on the volume of the stationary phase, Vs, and the analyte partition coefficient, K ) Cs/Ca, where Cs and Ca are the concentrations of the analyte in the sorbent and aqueous phases at equilibrium, respectively. It is the dependence on the partition coefficient that accounts for the capture of the analyte from the aqueous phase due to interactions between the analyte and sorbent, while variations in the partition coefficient among analytes and potential interferences provide the sorbent selectivity. The sensor responses in Figure 14 for each step change in input concentration were fit to the models using a nonlinear least-squares optimization with the detection efficiency, Ed , retention volume, Vr , and column theoretical plates, N, as regression parameters. A Gaussian function was used to model the normal distribution. These fits are shown as solid lines in Figure 14, where fits were determined for the step change from zero to the sample concentration, and separate fits were determined for the step change from sample concentration to zero. The experimental data can be fit extremely well, and one obtains important sensor and column chromatographic parameters. The fit parameters were in excellent agreement with independent experimental determinations of these parameters. Fits to the model were also used to create the solid lines in Figure 11 for the 99Tc sensor. Strictly speaking, the Gaussian distribution assumption is valid only for columns with relatively high numbers of theoretical plates.90 The sensor columns do not have high numbers of theoretical plates. However, for plate numbers greater than 5 or 10, the Gaussian model provides satisfactory fits and extracts parameters that are in agreement with other measurements. For lower numbers of theoretical plates, the Gaussian model still provides reasonable looking fits to the sensor response, but the fit parameters are unreasonable. Lo¨vkvist et al. reviewed several breakthrough profile equations and proposed an alternative function that can be used to describe the breakthrough profile in low plate number frontal chromatographic systems.90,91 This function was also integrated into the theory for the equilibration-based column sensors and dubbed the Low Plate Model.18 When applied to experimental data from columns with plate numbers of about 3, it was found to give good fits and parameters. It is not surprising that the composite bed minicolumn sensors generally have low numbers of theoretical plates, given that they have short lengths and the bed itself contains the sorbent diluted by the plastic scintillator beads. In addition, the sorbent materials have relatively big particle sizes, and linear flow velocities through the column are high. Nevertheless, such columns do make effective sensors and can be modeled. In general, extremely low numbers of theoretical plates (e.g., less than two) are undesirable because the volume in excess of the retention volume that is necessary to equilibrate the column becomes rather large, i.e., the sigmoid-shaped elution profile becomes extended horizontally. Thus, the

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Figure 15. Sensor flow cell and PMTs of the first-generation prototype 99Tc(VII) preconcentrating minicolumn sensor for equilibration-based sensing.

design of the sensor requires a balance between theoretical plates and the volumes required to equilibrate it, and these must be balanced with the required detection limits. The fact that sensors can be characterized in terms of chromatographic parameters and modeled theoretically can help to guide decisions about sensor design.

3.7. Engineered Radiometric Preconcentrating Minicolumn Sensors for Groundwater Measurements The preconcentrating minicolumn sensor for radiometric detection of 99Tc has been engineered into a sensor probe in a rod-shaped form suitable for monitoring in groundwater via well bore holes. Images of the sensor flow cell with two PMTs is shown in Figure 15, while images of first- and second-generation prototypes are shown in Figure 16. The first-generation prototype, which incorporates an anticoincidence shield around the sensing flow cell to reduce background counts, has been described in some detail.16 Using AG 4-X4 anion exchanger and BC-400, the performance of this sensor in equilibration-based sensing mode was investigated with chemically unmodified Hanford groundwater that was spiked with known amounts of 99Tc(VII). Sample volumes of 150 mL, pumped at 3 mL min-1, were used to equilibrate the column, followed by a 30-min counting interval. A linear calibration curve was obtained. Actual contaminated Hanford groundwater samples were also analyzed with results that were in agreement with independent laboratory analysis results, using the method of standard addition for calibration. It was shown in this study that a hydroxyapatite prefilter could be used to extend the life of the sensor by reducing fouling by colored impurities (likely soil organic matter) in the groundwater. No change in sensor response was evident after pumping a volume equivalent to 36 samples of 150 mL each. For a radionuclide sensor in equilibration-based sensing mode, the time required to obtain a measurement depends on the amount of sorbent material, the flow rate, the volume of sample required to equilibrate the sorbent phase, and the time to devote to counting once the sensor has equilibrated. Although one may spend minutes to hours equilibrating the column and counting the captured radionuclides, fast response time is typically not required for subsurface monitoring applications and long-term environmental stewardship. Changes occur slowly in the subsurface, and the interval between taking data points is long. The Hanford Site’s groundwater-monitoring program typically requires the sampling and analysis of monitoring wells on a quarterly basis. A deployed sensor could be equilibrated and the signal determined and then sealed until the next measurement. At a later date, a new sample could be introduced and

Figure 16. (Left) First-generation package for the engineered 99Tc(VII) sensor. Hands at the bottom where the water intake is located provide a scale. (Right) Second generation prototype, with a water intake, filtration, and pumping components at the bottom in clear plastic, linked to the sensor module and electronics in black with a universal joint to provide some flexibility. The diameters of the two prototypes are the same, the second is longer overall.

equilibrated to get another measurement. A reversible equilibration-based sensor requiring no reagents could potentially work in the field for tens of measurements or years of use if it is sufficiently selective within the sample matrix and it does not become fouled with organic matter or bacteria.

3.8. Planar Dual-Functionality Radionuclide Sensors The column format for preconcentrating radionuclide sensors is convenient but not required. DeVol et al. have described a radial flow format for a disc-shaped composite bed with scintillation detection.15 The flow-cell was based on a planar fountain cell design,92 shown in Figure 17, which introduces solution flow from the center of the resin bed to the periphery. The sensor was interfaced with a single photomultiplier tube (PMT). This design initially concentrates most of the retained activity in the center, where the PMT has the highest sensitivity. For 99Tc capture and detection, a number of sorptive scintillating media were investigated, including (1) an extractive scintillator (dual-functionality beads) combining a porous polystyrene resin with the extractant Aliquat-336 and fluor R-NPO, (2) a composite bed of plastic scintillator beads (BC-400) and Tc-selective TEVA resin, and (3) a composite bed of inorganic scintillator particles (CaF2:Eu) with either TEVA resin or strongly basic anion-exchange resin (Dowex1 × 8-400(Cl-)). These sensors were operated in a quantitative capture mode (as opposed to the equilibration-based modality just discussed), and the capture efficiencies with these materials in this flow-cell

Radionuclide Sensors for Environmental Monitoring

Figure 17. Planar flow cell (fountain cell) coupled to a PMT detector for a dual functionality preconcentrating radiometric sensor. The upper image with a gray background shows a side crosssectional view, while the lower image represents a top crosssectional view.

configuration were all above 98%. The detection efficiencies ranged from 10% for the extractive scintillator resin to 50% for the CaF2:Eu/Dowex. These results show that this alternative geometry provides both effective capture and reasonable detection efficiencies, even using just one photomultiplier tube. Planar configurations offer flexibility in the geometry of the selective materials and transducing materials, as well as creating opportunities for placing permeable membranes between the sample and the separation/transduction materials. Such a membrane could allow diffusion of ions to the resin bead while stopping suspended particles, thus preventing sensor fouling.

3.9. Planar Radionuclide Sensors Based on Diodes Semiconductor diode detectors represent an alternative to scintillation for radiometric detection. The principles of R and β radiation detection using silicon semiconductor diodes have been described in detail previously.93,94 The diode device necessarily leads to planar sensor formats. For sensors in liquids, the requirement to capture the R- and β-emitting radionuclides close to the transducing medium remains. This can be achieved by placing a selective thin film on the diode surface and using a flow cell like the fountain cell in Figure 17. Diode detectors offer a number of potential benefits. First, they offer superior energy resolution compared to scintillation detectors. For R particles the resolution can be as good as 20 keV FWHM, as opposed to several hundred keV for scintillation detectors. Thus, sensors based on diode detectors have the potential to offer R energy spectra of radionuclides, which could be useful for distinguishing actinides that are co-retained in a semiselective film. Second, diodes for R detection offer much lower background noise levels than scintillation approaches for R detection. Third, diodes offer the potential to discriminate R particle detection from β and γ radiation. Diode detectors with a selective layer to capture R emitters have been described recently.10,95 Polymeric layers containing

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extractants were applied to thicknesses from 0.25 to 5 µm. Thin polymeric films containing supported extractants have been extensively used in preparation of ion-selective electrodes. Typically, high-molecular weight PVC is the matrix, but other materials such as polyacrylates have been utilized.96,97 Extractant-loaded PVC membranes for radioactive cesium and strontium have also been described.98 Similar layers can be used to create radionuclide sensors. Using a surface passivated ion-implanted silicon diode, actinide sensors were demonstrated using bis(2-ethylhexyl)phosphoric acid (HDEHP) as the extractant and plasticizer in a PVC film on the surface.10,95 This extractant effectively binds actinides from aqueous solutions of low acidity, and the capture of 241Am on the diode from dilute acid was demonstrated. The analyte remained on the diode surface during a wash step, but the diode sensor can be regenerated by eluting retained 241Am with 4 M nitric acid. The absolute analytical efficiency was ∼30%. Using layers of 0.25 µm thickness, characteristic R energy peaks were obtained for 241 Am and 233U with peak widths of 35 keV FWHM. In an alternative format, DeVol et al. used a planar silicon diode in combination with ion-exchange beads for the detection of 99Tc.99 The sample was mixed with a small volume of these beads, which were then allowed to settle on the diode surface at the bottom of the flow cell. This process captured 99Tc from solution and brought it into close proximity to the diode. The experimental results indicated that this approach could detect to levels below the drinking water standards. In these experiments, the passivated ionimplanted planar silicon semiconductor detector was spray coated with a layer of Teflon AF to make the surface resistant to contact with aqueous solutions.

3.10. Fiber-Based Sensor Scintillating fibers have a significant history in radiation detection. One example of their use as a radiation detector in environmental monitoring is the direct detection of 90Sr and 238U in soil using a “blanket” of scintillating fibers, a method developed by Schilk and co-workers and marketed by Beta-Scint Company.100 This radiation detector responds to high-energy β particles from 90Y and 234mPa as an indication of 90Sr and 238U contamination in soil. Scintillating fibers were combined with a selective sorbent resin to create dual-functionality sensors for 137Cs in aqueous samples.101 The resin contained phenol-formaldehyde oligomers grafted on a polystyrene backbone with diphosphonate ligands along the chain. Particles of this resin (3-5 µm diameter) were bonded to Bicron BCF-12 scintillating fiber with an epoxy resin layer approximately one micrometer thick. In typical experiments, the fiber was allowed to equilibrate with a volume of sample. Radiometric measurements were made separately, after removing the fiber from the sample. Detection of 137Cs in alkaline simulated Hanford tank waste samples was demonstrated with selectivity over a number of other metals.

3.11. Whole-Column Chromatographic Sensor Whole-column radiometric detection was demonstrated in the chromatographic retention and separation of positronemitting analytes.102 Link and Synovec created a column flow cell out of BC-400 scintillating plastic materials and packed it with Dionex C14 media. Thus, the device as a whole was a dual-functionality sensor although the packing itself served

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only as the sorbent functionality. The planar column geometry was just 0.5 mm thick, while positrons from the analyte were emitted with a distribution of energies with an energy maximum of 0.96 MeV, and a maximum range of 3.5 mm. Using two photomultiplier tubes, a detection efficiency of 96% could be obtained. A flow-through scintillation detector without a packing was placed downstream from the packed column in order to compare on-column radiometric detection with downstream radiometric detection. Injected analytes on this system were concentrated by the sorbent medium and then released using a gradient elution, resulting in chromatographic migration down the column. Counts were recorded for the entire elution period. A plateau in the detector count rate was observed while the entire quantity of injected analyte was within the column sensor. By contrast, the downstream flow-through detector produced typical transient peaks of much lower area. The wholecolumn chromatographic sensor improved detection limits by a factor 50 compared to the downstream flow-through detector. Moreover, upon injection of samples with two or three components, a chromatographic separation was observed, enabling detection and quantification of multiple components in a single injected sample.

3.12. Dual-Functionality Sensor for Tritiated Water in Air While all the selectively sorbent sensors described above were directed toward the detection of radionuclides in water, the combination of sorptive and scintillating properties in a material has also been used to develop a novel sensor for tritiated water in air.103 Tritium has a low-energy β emission. The sensor material in this case is a Eu-doped zeolite of molecular sieve type 13X, where the zeolite structure provides the sorptive properties for water. The zeolite is also intrinsically a scintillator whose scintillation efficiency is increased by the Eu doping. A disk of the zeolite material was combined with a PMT and an air flow system for tritiated water monitoring. Because tritiated hydrogen gas is not adsorbed like water, the detector can distinguish between the oxidized (tritiated water) and elemental (tritiated hydrogen gas) forms.

4. Discussion The key concept of the preconcentrating minicolumn sensor is that analytes can be efficiently collected from a large sample volume into a small detection volume on a solid-phase sorbent in a flow-cell. The macroscopic quantity of sorbent has a capacity for a significant accumulation of analyte. By filling such a flow cell with a radionuclideselective material and a method for transducing radiation events into detectable light pulses, this approach has been demonstrated to provide very sensitive radionuclide sensors for water monitoring. These sensors potentially have applications in monitoring radionuclides in nuclear fuel reprocessing or nuclear waste processing plants, or for radionuclides in environmental matrixes such as groundwater. Using quantitative capture of radionuclide analytes with subsequent regeneration of the columns, these sensors were similar to the optosensors originally developed as detectors for flow injection analysis systems. These sensors assayed the quantity of analyte in a particular volume of solution, or the rate of analyte uptake as a function of sampled volume. With the development of equilibration-based sensing approach, the preconcentrating minicolumn sensors for

Grate et al.

Figure 18. Optical detector response to the hexavalent chromium column sensor in response to chromium standards in Hanford groundwater matrix; data was obtained by pumping at ∼1.5-mL min-1 net flow rate: (1) 45-mL blank sample, (2) 60 mL of 7.1 ppb Cr(VI); (3) 60 mL of 10.6 ppb Cr(VI); (4) 60 mL of 14.2 ppb Cr(VI); (5) 100-mL blank sample. Figure reprinted with permission from ref 18. Copyright 2006 American Chemical Society.

radionuclides became true sensors whose signals rise and fall with the ambient analyte concentration, albeit with a time constant determined in part by the time necessary to equilibrate the column. In this approach, the signal magnitude is no longer dependent on the volume of the sample once the minimum equilibration volume has been delivered. Although developed primarily for radionuclide sensors, the method can also be applied to other aqueous analytes where an on-column detection method of sufficient sensitivity is available. For example, hexavalent chromium can be captured in an anion exchange-based preconcentrating minicolumn sensor and observed with spectrophotometric detection.18 (This was mentioned briefly in the Background section on Preconcentrating minicolumn sensors.) This sensor can be operated in the equilibration sensing modality as shown in Figure 18. The three key features of this sensing modality are seen, with steady-state responses, reversibility, and a signal that varies with sample concentration. Equilibration-based sensing is well-known for sorbent thin films on microsensor device surfaces, but has not been widely practiced for minicolumn-type sensors. Thin films minimize the equilibration time by creating a short diffusion distance within the sorbent material, but the amount of analyte that can be captured is limited by the small amount of sorbent material. A thin film sensor, for example, may have a film mass in the scale of a few micrograms or less, while a preconcentrating minicolumn sensor may have tens of milligrams of sorbent. Thus, a thin film may have a very low capacity compared to a packed column. In addition, a thin film sensor creates a potentially long diffusion distance from the solution to the sorbent film for very low analyte concentrations. The preconcentrating minicolumn sensor overcomes long diffusion distances from solution to sorbent by actively pumping the sample through the column so that all portions of the solution pass within close proximity to the sorbent. The equilibration-based approach maximizes the extent of preconcentration for a given column material and geometry, and the performance can be modeled using chomatographic concepts. This approach can also alleviate the need for column regeneration solutions. All these features are advantageous for environmental sensing and monitoring applications. The additional time required to reach equilibration of

Radionuclide Sensors for Environmental Monitoring

a macroscopic amount of sorbent material is not a significant disadvantage in groundwater monitoring because changes in the subsurface occur slowly, and monitoring intervals are long. Thus, one obtains high sensitivity and operational simplicity in return for a response time penalty that is not important. A reversible equilibration-based sensor requiring no reagents could work in the field for tens of measurements or years of use if all other practical issues with regard to selectivity and stability in the particular application scenario are addressed. To date, the preconcentrating minicolumn sensor represents the best available sensing approach for meeting the daunting detection limit requirements for R- or β-emitting radionuclides in groundwater.

5. Acknowledgment We gratefully acknowledge sustained funding from U.S. DOE Office of Science Environmental Management Science Program and the Environmental Remediation Science Program. Funding for the development of the engineered sensor probe was provided by the DOE Environmental Management Advanced Monitoring Systems Initiative. We thank Dr. John Hartman at the Pacific Northwest National Laboratory (PNNL) for leading efforts in the development of this prototype. J.W.G. acknowledges the William R. Wiley Environmental Molecular Sciences Laboratory, a U.S. DOE scientific user facility operated for the DOE by PNNL. The Pacific Northwest National Laboratory is a multiprogram national laboratory operated for the U.S. Department of Energy by Battelle Memorial Institute.

6. References (1) DeVol, T. A.; Keillor, M. E.; Burggraf, L. W. IEEE Trans. Nucl. Sci. 1996, 43, 1310-1315. (2) DeVol, T. A.; Roane, J. E.; Harvey, J. T. Scintillating Extraction Chromatographic Resin for Quantification of Aqueous Radioactivity. In IEEE Nucl. Sci. Symp. Conf. Rec.; IEEE: New York, NY, 1997; pp 415-419. (3) Egorov, O. B.; Fiskum, S. K.; O’Hara, M. J.; Grate, J. W. Anal. Chem. 1999, 71, 5420-5429. (4) DeVol, T. A.; Roane, J. E.; Williamson, J. M.; Duffey, J. M.; Harvey, J. T. Radioact. Radiochem. 2000, 11, 34-46. (5) DeVol, T. A.; Egorov, O. B.; Roane, J. E.; Paulenova, A.; Grate, J. W. J. Radioanal. Nucl. Chem. 2001, 249, 181-189. (6) DeVol, T. A.; Duffey, J. M.; Paulenova, A. J. Radioanal. Nucl. Chem. 2001, 249, 295-301. (7) Roane, J. E.; DeVol, T. A. Anal. Chem. 2002, 74, 5629-5634. (8) Egorov, O.; O’Hara, M. J.; Grate, J. W. Spectrum 2002, Reno, NV. (9) Grate, J. W.; Egorov, O. B. Automated Radiochemical Separation, Analysis and Sensing. In Handbook of RadioactiVity Analysis, 2nd ed.; L’Annunziata, M. F., Ed.; Academic Press: Boston, 2003; pp 1129-1164. (10) Egorov, O. B.; O’Hara, M. J.; Addleman, R. S.; Grate, J. W. ACS Symp. Ser. 2004, 868, 246-270. (11) Grate, J. W.; Egorov, O. B.; O’Hara, M. J. ACS Symp. Ser. 2004, 904, 322-341. (12) Roane, J. E.; DeVol, T. A. J. Radioanal. Nucl. Chem 2005, 263, 51-57. (13) Hughes, L.; DeVol, T. A. Nucl. Instrum. Methods Phys. Res., Sect. A 2003, 505, 435-438. (14) Ayaz, B.; DeVol, T. A. Nucl. Instrum. Methods Phys. Res., Sect. A 2003, 505, 458-461. (15) Hughes, L. D.; DeVol, T. A. Anal. Chem. 2006, 78, 2254-2261. (16) Egorov, O. B.; O’Hara, M. J.; Grate, J. W.; Knopf, M.; Anderson, G.; Hartman, J. J. Radioanal. Nucl. Chem. 2005, 264, 495-500. (17) Fjeld, R. A.; Roane, J. E.; Leyba, J. D.; Paulenova, A.; DeVol, T. A. ACS Symp. Ser. 2004, 868, 105-119. (18) Egorov, O. B.; O’Hara, M. J.; Grate, J. W. Anal. Chem. 2006, 78, 5480-5490. (19) Bosworth, N.; Towers, P. Nature 1989, 341, 167-168. (20) Cook, N. D. Drug DiscoVery Today 1996, 1, 287-294.

Chemical Reviews, 2008, Vol. 108, No. 2 561 (21) L’Annunziata, M. F. In Handbook of RadioactiVity Analysis; L’Annunziata, M. F., Ed.; Academic Press: San Diego, 1998; pp 556-565. (22) Valcarcel, M.; Luque de Castro, M. D. Flow-Through (Bio)Chemical Sensors; Elsevier: Amsterdam, 1994. (23) Ruzicka, J.; Hansen, E. H. Flow Injection Analysis, 2nd ed.; WileyInterscience: New York, 1988; Vol. 62, p 498. (24) Fang, Z. Flow Injection Separation and Preconcentration; VCH: Weinheim, 1993. (25) Ruzicka, J.; Marshall, B. D. Anal. Chim. Acta 1990, 237, 329. (26) Ruzicka, J. Anal. Chim. Acta 1992, 261, 3-10. (27) Ivaska, A.; Ruzicka, J. Analyst 1993, 118, 885-889. (28) Ruzicka, J. Collect. Czech. Chem. Commun. 2005, 70, 1737-1755. (29) Miro, M.; Hansen, E. H. Trends Anal. Chem. 2006, 25, 267-281. (30) Yoshimura, K.; Matsuoka, S. Lab. Rob. Autom. 1993, 5, 231-244. (31) Miro, M.; Frenzel, W. Trends Anal. Chem. 2004, 23, 11-20. (32) Yoshimura, K.; Yamada, S. Talanta 1992, 39, 1019-1024. (33) Torre, M.; Marina, M. L. Crit. ReV. Anal. Chem. 1994, 24, 327361. (34) Yoshimura, K. Anal. Chem. 1987, 59, 2922-2924. (35) Ruzicka, J.; Scampavia, L. Anal. Chem. 1999, 71, 257A-263A. (36) Ruzicka, J. Anal. Chim. Acta 1995, 308, 14-19. (37) Egorov, O.; Ruzicka, J. Analyst 1995, 120, 1959-1962. (38) Ruzicka, J. Analyst 1994, 119, 1925-1934. (39) Pollema, C. H.; Ruzicka, J. Anal. Chem. 1994, 66, 1825-1831. (40) Ruzicka, J.; Pollema, C. H.; Scudder, K. M. Anal. Chem. 1993, 65, 3566-3570. (41) Egorov, O.; O’Hara, M. J.; Grate, J. W.; Ruzicka, J. Anal. Chem. 1999, 71, 345-352. (42) Grate, J. W.; Egorov, O. B. Anal. Chem. 1998, 70, 779A-788A. (43) Dockendorff, B.; Holman, D. A.; Christian, G. D.; Ruzicka, J. Anal. Commun. 1998, 35, 357-359. (44) Bruckner-Lea, C. J.; Stottlemyre, M. S.; Holman, D. A.; Grate, J. W.; Brockman, F. J.; Chandler, D. P. Anal. Chem. 2000, 72, 41354141. (45) Grate, J. W.; Bruckner-Lea, C. J.; Jarrell, A. E.; Chandler, D. P. Anal. Chim. Acta 2003, 478, 85-98. (46) Chandler, D. P.; Brockman, F. J.; Holman, D. A.; Grate, J. W.; Bruckner-Lea, C. J. Trends Anal. Chem. 2000, 19, 314-321. (47) Cortina, J. L.; Warshawsky, A. Ion Exch. SolVent Extr. 1997, 13, 195-293. (48) Horwitz, E. P.; Chiarizia, R.; Dietz, M. L.; Diamond, H.; Nelson, D. M. Anal. Chim. Acta 1993, 281, 361-372. (49) Horwitz, E. P.; Chiarizia, R.; Dietz, M. L. SolVent Extr. Ion Exch. 1992, 10, 313-336. (50) Horwitz, E. P.; Dietz, M. L.; Chiarizia, R.; Diamond, H.; Maxwell, S. L.; Nelson, M. R. Anal. Chim. Acta 1995, 310, 63-78. (51) Dietz, M. L.; Horwitz, E. P. LC-GC 1993, 11, 424-426, 428, 430, 434, 436. (52) Maxwell, S. L. Radioact. Radiochem. 1997, 8, 36-44. (53) Izatt, R. M.; Bradshaw, J. S.; Bruening, R. L. Pure Appl. Chem. 1996, 68, 1237-1241. (54) Izatt, R. M. J. Inclusion Phenom. Mol. Recognit. Chem. 1997, 29, 197-220. (55) Izatt, R. M.; Bradshaw, J. S.; Bruening, R. L. Pure Appl. Chem. 1995, 68, 1237-1241. (56) Izatt, R. M.; Bradshaw, J. S.; Bruening, R. L.; Bruening, M. L. Am. Lab 1994, 26, 28C. (57) Egorov, O. B.; O’Hara, M. J.; Ruzicka, J.; Grate, J. W. Anal. Chem. 1998, 70, 977-984. (58) Egorov, O. B.; O’Hara, M. J.; Farmer, O. T., III; Grate, J. W. Analyst 2001, 126, 1594-1601. (59) Grate, J. W.; Fadeff, S. K.; Egorov, O. Analyst 1999, 124, 203210. (60) Grate, J. W.; Egorov, O. B.; Fiskum, S. K. Analyst 1999, 124, 11431150. (61) Grate, J. W.; Egorov, O. Anal. Chem. 1998, 70, 3920-3929. (62) Egorov, O.; Grate, J. W.; Ruzicka, J. J. Radioanal. Nucl. Chem. 1998, 234, 231-235. (63) Grate, J. W.; Strebin, R. S.; Janata, J.; Egorov, O.; Ruzicka, J. Anal. Chem. 1996, 68, 333-340. (64) Egorov, O. B.; O’Hara, M. J.; Grate, J. W. Anal. Chem. 2004, 76, 3869-3877. (65) Egorov, O.; O’Hara, M. J.; Grate, J. W. J. Radioanal. Nucl. Chem. 2005, 263, 629-633. (66) Koglin, E. N.; Poziomek, E. J.; Kram, M. L. In Handbook of Vadose Zone Characterization & Monitoring; Wilson, L. G., Everett, L. G., Cullen, S. J., Eds.; Lewis Publishers: Ann Arbor, 1994. (67) Hartman, M. J., Morasch, L. F., Webber, W. D., Eds. Hanford Site Groundwater Monitoring for Fiscal Year 2005, Pacific Northwest National Laboratory, PNNL-14670, 2006.

562 Chemical Reviews, 2008, Vol. 108, No. 2 (68) Hartman, M. J.; Dresel, P. E.; editors, Hanford Site Groundwater Monitoring for Fiscal Year 1997, Pacific Northwest National Laboratory, PNNL-11793 UC-402, 403, 702, 1998. (69) The drinking water standard for a β emitter in water as established by the EPA is 4 mrem/year. Using the dose conversion factor from National Bureau of Standards Handbook 69 (U.S. Department of Commerce, as amended August 1963) and the other parameters established by the EPA, one can calculate an equivalent concentration of 33 Bq/L (900 pCi/L) assuming 99Tc is the sole β emitter. (70) Wang, J.; Lu, J.; Wang, J.; Luo, D.; Tian, B. Anal. Chim. Acta 1997, 354, 275-281. (71) Wang, J. Stripping Analysis: Principles, Instruments, and Applications; VCH Publishers, Inc.: New York, 1985. (72) Olsen, K. B.; Wang, J.; Setiadji, R.; Lu, J. EnViron. Sci. Technol. 1994, 28, 2074-2079. (73) Heimbuch, A. M.; Gee, H. Y.; DeHaan, A. J.; Leventhall, L. Radioisotope Sample Measurement Techniques in Medicine and Biology; International Atomic Energy Agency Symposium, Vienna, May 24-28, 1965. (74) Li, M.; Schlenoff, J. B. Anal. Chem. 1994, 66, 824-829. (75) DeVol, T. A.; Roane, J. E.; Williamson, J. M.; Duffey, J. M.; Harvey, J. T. 44th Annual Conference on Bioassay, Analytical, and Environmental Radiochemistry, Alburquerque, NM, Novermber 15-20, 1998. (76) Capacity factors, k′, were calculated as k′) A × Dw × (Vs/Vm ) after equilibrating known quantities of sensor beads with solutions of known volume and 99-Tc activity . The ratio Vs/Vm is the stationary/ mobile phase volume ratio. Weight distribution ratios, Dw, were calculated according to the formula Dw ) ((A0 - As)/W)/(As/V), where A0 is the activity of the blank solution after equilibration, As is the activity of the test solution after equilibration with beads, W is the weight of the beads (g), and V is the volume of the equilibrated solution (mL). Finally, the coefficient A is the conversion factor from Dw to volume distribution ratio). Further details are in the cited work. (77) “National Primary Drinking Water Regulations; Radionuclides; Final Rule,“ Federal Register, Vol. 65, No. 236, December 7, 2000, pp 76708-76753. (78) Rogers, R. D.; Bond, A. H.; Griffin, S. T.; Horwitz, E. P. SolVent Extr. Ion Exch. 1996, 14, 919-946. (79) Horwitz, E. P.; Dietz, M. L.; Diamond, H.; LaRosa, J. J.; Fairman, W. D. Anal. Chim. Acta 1990, 238, 263-271. (80) Horwitz, E. P.; Chiarizia, R.; Dietz, M. L. React. Funct. Polym. 1997, 33, 25-36. (81) Dietz, M. L.; Horwitz, E. P.; Nelson, S. M.; Wahlgren, M. Health Phys. 1991, 61, 871-877.

Grate et al. (82) Horwitz, E. P.; Dietz, M. L.; Fisher, D. E. Anal. Chem. 1991, 63, 522-525. (83) Horwitz, E. P.; Dietz, M. L.; Chiarizia, R. J. Radioanal. Nucl. Chem. Articles 1992, 161, 575-583. (84) Vajda, N.; Ghods-Esphahani, A.; Cooper, E.; Danesi, P. R. J. Radional. Nucl. Chem. 1992, 162, 307-323. (85) Jeter, H. W.; Brob, B. Radioact. Radiochem. 1994, 5, 8-16. (86) Currie, L. A. Anal. Chem. 1968, 40, 586-591. (87) DeVol, T. A.; Roane, J. E.; Leyba, J. D. In LSC 2001 AdVances in Liquid Scintillation Spectrometry; Mobius, S., Noakes, J. E., Schonhofer, F., Eds.; University of Arizona: Tuscon, 2002; pp 415-424. (88) DeVol, T. A.; Roane, J. E.; Hughes, L.; Ayaz, B. Spectrum 2002: 9th Biennial International Conference on Nuclear and Hazardous Waste Management, Reno, NV, 2002. (89) Reilley, C.; Hildebrand, G. P.; Ashley, J. W. Anal. Chem. 1962, 34, 1198-1213. (90) Lovkvist, P.; Jonsson, J. A. Anal. Chem. 1987, 59, 818-821. (91) Lovkvist, P.; Jonsson, J. A. J. Chromatogr. 1986, 356, 1-8. (92) Scudder, K. M.; Pollema, C. H.; Ruzicka, J. Anal. Chem. 1992, 64, 2657-2660. (93) Fettweis, P. F.; Verplancke, J.; Venkataraman, R.; Young, B. M.; Schwenn, H. In Handbook of RadioactiVity Analysis, 2nd ed.; L’Annunziata, M. F., Ed.; Academic Press: Boston, 2003; pp 239346. (94) Knoll, G. F. Radiation Detection and Measurement; John Wiley & Sons: New York, 1989. (95) Addleman, R. S.; O’Hara, M. J.; Grate, J. W.; Egorov, O. B. J. Radioanal. Nucl. Chem. 2005, 263, 291-294. (96) Edmonds, T. E., Ed. Chemical Sensors; Chapman and Hall: New York, 1988. (97) Janata, J.; Bezegh, A. Anal. Chem. 1988, 60, 62R. (98) Rais, J.; Mason, C. V.; Abney, K. D. Sep. Sci. Technol. 1997, 32, 951-969. (99) Hughes, L. D.; DeVol, T. A. J. Radional. Nucl. Chem. 2006, 267, 287-295. (100) Schilk, A. J.; Knopf, M. A.; Thompson, R. C.; Hubbard, C. W.; Abel, K. H.; Edwards, D. R.; Abraham, J. R. Nucl. Instrum. Methods Phys. Res., Sect. A 1994, 353, 477-481. (101) Headrick, J.; Sepaniak, M.; Alexandratos, S. D.; Datskos, P. Anal. Chem. 2000, 72, 1994-2000. (102) Link, J. M.; Synovec, R. E. Anal. Chem. 1999, 71, 2700-2707. (103) Campi, F.; Edwards, A. H.; Ossiri, A.; Pacenti, P.; Terrani, S. Health Phys. 1998, 75, 179-182.

CR068115U

Chem. Rev. 2008, 108, 563−613

563

Higher-Order Chemical Sensing Andreas Hierlemann* ETH Zu¨rich, Laboratory of Physical Electronics, Wolfgang-Pauli-Strasse 16, 8093 Zu¨rich, Switzerland

Ricardo Gutierrez-Osuna Texas A&M University, Department of Computer Science, College Station, Texas 77843-3112 Received April 3, 2007

Contents 1. Introduction 1.1. Integrated versus Discrete Sensor Arrays 1.1.1. Materials and Fabrication Processes 1.1.2. Performance 1.1.3. Auxiliary Sensors/Smart Features 1.1.4. Connectivity 1.1.5. Sensor Response Time 1.1.6. Package 1.1.7. Summary 1.2. Outline 2. Arrays and Systems Comprising Identical Transducers 2.1. Parameter Variations 2.2. Dynamic Methods and Transient Signals 2.3. Extracting Information in Different Domains 3. Arrays and Systems Comprising Different Transducers 3.1. Metal- and Metal-Oxide-Based Gas Sensors 3.2. Polymer-Based Gas Sensors 3.3. Gas Sensor Arrays Relying on Different Transducer and Sensitive-Material Types 3.4. Liquid-Phase Chemo- and Biosensors 3.5. Cell-Based Biosensors 4. Operational Considerations for Higher-Order Devices 4.1. Setup and Manifold Considerations 4.2. Multitransducer Operation Example 5. Sensor-Based Microanalytical Systems 6. Are More Sensors Better? 6.1. Characteristics of High-Dimensional Vector Spaces 6.2. Orthogonality versus Independence 6.3. Cross-sensitivity and Diversity 6.4. Multiple Roles of Redundancy 7. Data Preprocessing 7.1. Baseline Correction 7.2. Scaling 7.2.1. Global Techniques 7.2.2. Local Techniques 7.2.3. Nonlinear Transforms

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* To whom correspondence should be addressed. Voice: +41 44 633 3494. Fax: +41 44 633 1054. E-mail: [email protected].

8. Drift Compensation 8.1. Univariate Drift Compensation 8.2. Multivariate Drift Compensation 9. Feature Extraction from Sensor Dynamics 9.1. Transient Analysis 9.1.1. Oversampling Procedures 9.1.2. Ad hoc Transient Parameters 9.1.3. Model-Based Parameters 9.1.4. Comparative Studies 9.2. Temperature-Modulation Analysis 10. Multivariate Calibration 10.1. Multiway Analysis 10.2. Dynamical Models 11. Array Optimization 11.1. Sensor Selection 11.2. Feature Selection 11.3. Optimization of Excitation Profiles 12. Conclusion and Outlook 13. References

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1. Introduction Desired properties of a chemical sensor include high sensitivity, a large dynamic range, high selectivity or specificity to a target analyte, related low cross-sensitivity to interferents, perfect reversibility of the physicochemical detection or sensing process (short sensor recovery and response times), and long-term stability of the sensor and sensing material.1-9 Unfortunately, a sensor exhibiting all these properties is a largely unrealizable ideal. Sensor sensitivity, selectivity, speed of response, and reversibility are determined by the thermodynamics and kinetics of sensor material/analyte interactions. In particular, high sensitivity and specificity on the one hand and perfect reversibility on the other hand impose contradictory constraints on the sensor design: high sensitivity and selectivity are typically associated with strong interactions, whereas perfect reversibility requires weak interactions. Consequently, it is necessary to compromise, and, in most cases, sensors showing partial selectivity to only some of the detected species are used to ensure reversibility. The output of an individual sensor consists of, e.g., a certain current value measured at a fixed potential or a resistance value of a certain material in response to a chemical stimulus.1-9 This means that, usually, one feature per sensor is monitored at a time, preferably

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Andreas Hierlemann received his diploma in chemistry in 1992 and the Ph.D. degree in physical chemistry in 1996 from the University of Tu¨bingen, Germany. After working as a Postdoc at Texas A&M University, College Station, TX (1997), and Sandia National Laboratories, Albuquerque, NM (1998), he is currently associate professor at the Physical Electronics Laboratory at ETH Zurich in Switzerland. The focus of his research activities is on CMOS-based microsensors and interfacing CMOS electronics with electrogenic cells.

Ricardo Gutierrez-Osuna received a B.S. degree in Electrical Engineering from the Polytechnic University of Madrid, Spain, in 1992, and M.S. and Ph.D. degrees in Computer Engineering from North Carolina State University in 1995 and 1998, respectively. From 1998 to 2002, he served on the faculty at Wright State University. He is currently an associate professor of Computer Engineering at Texas A&M University. Dr. GutierrezOsuna is a recipient of the National Science Foundation Career Award for his research on machine olfaction with chemical sensors arrays. His research interests include pattern recognition, neuromorphic computation, chemical sensor arrays, and machine perception.

during an equilibrium-type or steady-state-type situation, in which a certain analyte concentration can unequivocally be correlated to the resulting sensor response. Individual sensors and the analysis of the respective individual signals or features, however, show limited selectivity performance in most practical applications, as mentioned above. Therefore, arrays of several sensors are commonly used, which effectively extends the “feature space”. A feature space is an abstract space in which each sample (e.g., sensor measurement value) is represented as a point in n-dimensional space, whose dimension is determined by the number of features evaluated. Features are the individual measurable heuristic properties of the phenomena being observed, in our case, e.g., sensor measurement values. The acquired information is then processed using pattern recognition and multicomponent analysis tools.10-12 Increasing the measurement or feature space dimensionality is an attractive approach to obtain a substantial improvement in analytical capabili-

Hierlemann et al.

ties,13,14 provided that the additional dimensions carry complementary information and provided that suitable dataanalysis techniques are used, i.e., techniques to handle small numbers of samples. The notion of “order” has been frequently used in instrumental analysis to categorize the dimensionality of the feature space and, consequently, the richness of the information obtained by a certain device or analytical instrument15,16 and can, within certain limitations, also be applied to sensor or sensor array configurations.17 A zero-order device or sensor would be a single, moreor-less specific sensor. This sensor would be aimed at the detection of a certain target compound. However, it would not be possible to perform any reliable analyte quantification in the presence of interferents. Even worse, there would not be any evidence to let the operator know that the measurement had been influenced by interferents. A first-order device is then a sensor array, the sensors of which differ in one domain, i.e., an array of identical transducers such as chemoresistors featuring different sensitive materials or layers or an array of chemoresistors featuring the same sensitive material but operated at different temperaturessprovided that temperature has a distinct effect on the sensor selectivity. This is the case for, e.g., hightemperature metal-oxide-based sensors.18-22 A prerequisite for the successful use of such first-order devices is the establishment of a calibration model that includes the signals of the species of interest and of all possible interferents. Then, multicomponent analysis and outlier detection will be possible, but any unexpected interferent will invariably upset the respective predictions. In other words, the calibration of a first-order array allows interferences to be detected but not to be compensated for. This problem can be addressed by using second-order or higher-order sensor devices. These devices utilize analyte characteristics in at least two domains, which should be ideally orthogonal, or, inasmuch as possible, independent from each other (see section 6.2 for a definition of these terms). Under certain conditions, the calibration of secondorder instruments allows the target analytes to be quantified even in the presence of unknown interferents; this property is commonly referred to as the “second-order advantage”.23 Additional potential benefits of second-order devices include the ability to perform calibration with a single mixture sample and recover the response profiles of the individual target analytes.24 Higher-order sensor devices can, according to a paper published by Go¨pel,25 be generated by making use of various features to be exploited in chemical sensing. Though the estimated overall number of features may be overly optimistic (Figure 1), in particular with regard to “independent features”, it may be interesting to briefly summarize the findings of this paper. The number of sensitive materials to convert a chemical into, e.g., electrical information is estimated to be on the order of 108, the number of suitable transducers (e.g., chemoresistor, microbalance, optical fiber, etc.) is estimated to be ∼101, the number of transducer geometries (e.g., electrode distance, shape, etc.) is estimated to be on the order of 102, and the features that can be added through variation of modulation of external (e.g., gas switching strategies, use of filters, catalysts, etc.) and internal parameters (e.g., light frequency, operation temperature, bias voltage, etc.) is estimated to be on the order of 102. The distinctive way to modulate these parameters (e.g., stepwise, sinusoidal, ramp, etc.) is assumed to account for an additional

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Figure 1. “Hyperspace” of chemical sensor features with more than 1021 independent features according to Go¨pel. Redrawn with persmission from ref 25. Copyright 1998 Elsevier.

102 features, and, finally, multiple modulations that can be performed simultaneously (e.g., simultaneous variation of bias voltage, temperature, and gas flow) are assumed to account for another 106 features. While the above considerations are of hypothetical nature, and the number of viable parameters and realizable variations will be massively lower, the exercise shows that a large variety of parameters is available that can be used to identify or quantify a specific analyte or complex mixture. Higher-order (i.e., “hyphenated”) instrumentation, such as the combination of gas chromatography (GC) and mass spectrometry (MS), is vastly used in modern laboratory analytical chemistry. Although the data-analysis methods applied to sensor and analytic-instrument data are mostly the same, there is no direct analogy between the preprocessing of a sample to separate it into multiple, less complex samples that are then characterized by an analytical method and the attempts to enhance the selectivity of a set of sensors by modulation of physical parameters. Yet, it is striking that higher-order methods are still uncommon in the field of chemical sensors. This holds particularly true, since the higher-order advantage may be important due to unpredictable changes in the sample matrix composition. This situation is arguably a consequence of the requirements for sensors or sensor systems, which include low costs, small physical size, and ease of use. The progress in micro- and nanotechnology, microelectronics, and in data-processing speed and capability will help to address many, if not all, of these issues:26-30 rather complex and versatile microsensor and microanalysis systems operable directly through standard interfaces from a laptop or palmtop by means of standard software are already available, as will be also demonstrated in this article.

1.1. Integrated versus Discrete Sensor Arrays In this context, we want to address the advantages and disadvantages of integrated or even monolithic multisensor arrays or systems versus the use of sets of discrete sensors and electronics, particularly, since both types of sensor arrays or systems will appear in the following sections. There is a number of aspects that have to be taken into account, which will be briefly discussed here.31

1.1.1. Materials and Fabrication Processes For monolithic designs or integrated systems, the selection of materials is restricted to a few, e.g., CMOS-technologyrelated materials and CMOS-process-compatible materials, as well as to a set of specific fabrication steps32 (CMOS technology, complementary-metal-oxide-semiconductor technology, is an industrial standard fabrication technology for integrated circuits on silicon microchips). High-temperature steps (e.g., >400 °C) can be detrimental to metallizations (metal oxidation, diffusion) and may alter semiconductor

characteristics. For hybrid or discrete devices, any material, or the optimum sensor material, can be used, and a wealth of fabrication techniques is available.

1.1.2. Performance Microsensors frequently also generate “microsignals”, and perform pronouncedly better in monolithic designs owing to the fact that the signals can be conditioned at the site of generation (filters, amplifiers, etc.), e.g., by means of onchip electronics, so that the influence of parasitic and crosstalk effects can be reduced.33,34 On-chip analog-to-digital conversion is another feature that helps to generate a stable sensor output that can be easily transferred to off-chip units. For hybrid or discrete microsensors, it is sometimes very difficult to read out rather minute analog sensor signals.

1.1.3. Auxiliary Sensors/Smart Features Temperature or flow sensors can be monolithically cointegrated with the chemical sensors on the same chip. Calibration, control, and signal processing functions as well as self-test features can be realized on-chip. For hybrid designs, additional devices and off-chip components are required.

1.1.4. Connectivity The number of electrical connections prominently contributes to the overall system costs (failure probability and packaging costs). The monolithic implementation of, e.g., an array of gas sensors (see also later in this article) with multiplexer structures and interface units requires only a few connections.33,34 A hybrid/discrete approach will require many more connections, since each sensor has to be individually addressed and since there are no interface units available on the sensor side.

1.1.5. Sensor Response Time The response time of, e.g., a gas sensor array is, in most cases, determined by the volume of the measurement chamber and the flow rate (other relevant processes include also analyte diffusion or dissociation). Using the monolithic or integrated approach and a suitable packaging technique, such as flip-chip packaging, the volume of the measurement chamber can be kept very small as a consequence of a smallsize, flat and planar sensor or sensor array. Therefore, parameter modulation, such as flow variations or dynamic protocols, can be easily realized. For hybrid or discrete arrays, the volume is dependent on sensor geometries and array arrangements.

1.1.6. Package To package monolithic designs, microelectronics-derived packaging techniques can be modified and adapted, such as

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flip-chip technology or simple epoxy-based packaging methods. Hybrid implementations require complex packages to reduce sensor interference (e.g., high-frequency acousticwave sensors), to minimize electric crosstalk, and to optimize the critical connections. This further complicates the already difficult task of chemical sensor packaging.

1.1.7. Summary The main disadvantages of integrated or monolithic arrays or systems include the restriction in materials and the limited choice of fabrication processes and steps. On the other hand, integrated systems offer unprecedented advantages over hybrid or discrete arrays, especially with regard to signal quality, device performance, increased functionality, and available packaging solutions.35 These advantages, in our opinion, clearly outweigh the drawbacks and limitations.31 In the case of well-established physical sensors, such as acceleration and pressure sensors, a trend toward integrated monolithic solutions can be identified for large production volumes and severe cost restrictions.

1.2. Outline The topic of this review is “higher-order devices”, i.e., neither single sensors nor homogeneous transducer arrays featuring only different coatings will be treated any further. The latter constitute, according to the text in the introduction above, a first-order system, since analyte exposure generates a one-dimensional data vector (row or column). In some publications, an array of sensors with different coatings is referred to as a zero-order array,9 with an array being itself a first-order device. We here prefer to use a categorization according to the data output format (0th order, single value; first order, vector; second order, matrix; third order, tensor; etc.). We decided to apply a rather broad scope in this review in order to give the reader a comprehensive overview on strategies to increase the information that can be extracted from sensor systems or arrays. In section 2, devices consisting of arrays of identical transducers (with different coatings) will be described, for which an extension to a higher order has been realized by adding additional dimensions such as time (sensor dynamics and transients) or by varying internal and external operation parameters such as temperature modulation or the use of a catalyst for analyte decomposition. In this section, it will also be shown how different information, e.g., in the physical and chemical domains, can be extracted from an array of identical transducers. Section 3 focuses on different monolithic and discrete sensor arrays making use of more than one transduction principle. In section 4, we will briefly describe practical measurement and setup considerations for using multitransducer sensor arrays and for recording transients or applying parameter variations. Section 5 includes sensor-based complex microanalytical systems, consisting of preconcentration, separation, and detection stages. In section 6, we will discuss the relations between dimensionality, information, cross-selectivity, and redundancy, concepts that are important when dealing with higher-order sensor systems. Section 7 presents a review of two important data preprocessing procedures for chemosensors: baseline correction and scaling. Section 8 will be dedicated to methods for drift compensation. In section 9, we will review computational methods to extract information from transient and temperature-modulated responses of chemosensors. Section 10 is dedicated to multivariate

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calibration for higher-order sensing, with an emphasis on multiway and dynamical models. Section 11 will include methods to optimize arrays for specific applications, including sensor selection, feature selection, and optimization of temperature programs. This article will be concluded by a short summary and outlook. Before embarking upon the subject of multisensor arrays, there is a last note on terminology. The term “electronic nose” has been very popular for more than a decade to describe multisensor and multitransducer arrays.6,36,37 We believe, however, that this term can be very misleading for several reasons. Following a definition of Gardner and Bartlett, an “electronic nose” is “an instrument comprising an array of electronic chemical sensors with partial specificity and an appropriate pattern recognition system, capable of recognizing simple or complex odors”.6,36,37 The majority of such “electronic noses” may be capable of differentiating between analytes or analyte mixtures from the headspace of different foods or beverages, but, in most cases, the sensor response patterns cannot be directly correlated with human olfactory perception. Another more important point concerns the general applicability implied by the term “electronic nose”. Most of the sensor systems perform well in certain key applications, but there are few systems, if any, that exhibit the enormously broad applicability spectrum, at once including the sensitivity and discriminating power of a human or animal nose. In addition, successful sensor systems have to be designed and optimized with the key application in mind to guide the selection of coatings, transduction mechanisms, etc. As yet, however, there is no universally applicable system that invariably provides satisfactory performance under all circumstances. Therefore, we will use this term sparingly and only in quotation marks. Similar considerations apply to “electronic tongues”, ion-sensitive or lipid-film based sensors in liquid phase. An article on “electronic noses” by Weimar38 is included in this issue.

2. Arrays and Systems Comprising Identical Transducers In this section, we will treat arrays of identical transducers with, e.g., different coatings, for which an extension to a higher order has been realized by adding additional dimensions in the feature space. These additional dimensions may include time (sensor dynamics and transients), or the variation of internal and external sensor operation parameters (temperature variation, use of a catalyst for analyte decomposition). It will also be shown how information in the physical and chemical domains can be extracted from microarrays and microsystems. A short glance at differential or ratiometric methods concludes this section.

2.1. Parameter Variations Following Go¨pel and Weimar,25,39 the parameters that can be varied during sensor operation include internal parameters, such as sensor temperature, electrode bias voltage, or measurement frequencies, and external parameters, such as the use of filters or catalysts to change the gas composition. In looking at the literature, it is evident that one type of internal variation, the sensor operation temperature variation, is very popular in particular for conductometric metal-oxidebased sensors.40 This development has been fueled by the appearance of microhotplates with low thermal mass,41 which allow for millisecond-scale temperature variations, so that

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Figure 2. Scanning-electron micrograph of microhotplates. The suspended plates exhibit a polysilicon heater, an aluminum plane for homogeneous heat distribution, and electrodes for measuring the resistance of a semiconductor metal oxide. Reprinted with permission from ref 51. Copyright 1998 American Chemical Society.

the temperature variations are faster or on the same time scale as the chemical processes occurring during gas/metaloxide interaction. This enables effective gas discrimination via the use of analyte-specific dedicated temperature profiles. More details on temperature-modulation strategies can be found in section 9.2 of this article, as well as in the articles by Benkstein and Semancik42 and Nakamoto43 in this issue. A variety of hotplate structures has been used including membranes,18,41,44-47 spiderlike structures48-51 (see Figure 2), and bridgelike structures.52 The most recent developments in temperature-variable microhotplates include standalone CMOS-based microsystems featuring temperature-control loops, transistor heaters, digital circuitry, and standard interfaces, which allow for the application of any arbitrary temperature profile to three differently coated hotplates via standard software and a USB interface.53,54 Besides microhotplate-based systems, there have been also static approaches to temperature variation using 38 metaloxide sensor elements (the array features a noble metal doping gradient) on a 4 × 8 mm2 bulk silicon substrate equipped with 4 meander heaters, which create temperature gradients between 3 and 7 °C per mm in the array area, producing a temperature difference of 50 °C over the array.55-57 This system consumes up to 6 W at operation temperatures between 300 and 400 °C, which is 3-4 times the power consumption of microhotplate-based systems per detection spot or sensor. The gas reactions at the metal-oxide surfaces and, hence, the sensor selectivity or sensitivity patterns, are highly depending on the operation temperature:1,58-66 carbon monoxide is usually detected at lower operation temperatures using a tin-oxide-based sensitive layer, whereas higher temperatures are used for monitoring, e.g., methane. The variation of the operation temperature of a single sensor or a small set of sensors can lead to a degree of selectivity that would otherwise require arrays of fixed-temperature sensors and, thus, effectively extends the feature space of single sensors or small arrays. With regard to static, fixed-temperature sensors, the fast temperature variation of microhotplates, which generates a large set of “virtual” sensors, is clearly preferable due to the almost infinite number of possible and target-analytespecific temperature-variation profiles (sinusoidal, ramp, rectangular); the arbitrarily selectable temperature interval,

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within which the variations can be realized; and the massively lower instrumental complexity and overall power consumption. A temperature-modulation example is shown in Figure 3,47 which shows a sinusoidal modulation of the operation temperature of a single tin-oxide sensor between 200 and 400 °C that produces characteristic frequency-dependent resistance features. Resistance changes of the micromachined sensor upon exposure to CO, NO2, and a mixture of CO and NO2 in synthetic air at 50% relative humidity are displayed. By applying this temperature-modulation profile and by using fast Fourier transformation techniques for feature extraction and data evaluation, a single sensor could be used to qualitatively and quantitatively analyze a binary mixture.47 The authors ascribe the possibility to differentiate between the two gases, CO and NO2, to the different reaction kinetics of the two gases at the sensor surface and, in particular, to the presence of oxygen species at the surface at low temperature as a consequence of the fast temperature modulation. These surface oxygen species would not exist on the surface under equilibrium conditions at the lower temperatures in the cycling range (200 °C).47 Additionally, an example for external parameter variation will be given, which includes the use of a catalyst located upstream of the sensor array in the analyte gas inlet.67-69 The noble metal catalyst is heated to different temperatures and decomposes (oxidizes) the incoming analyte molecules or promotes chemical reactions of those. The resulting reaction products are then detected by an array of, e.g., electrochemical sensors.67-69 By varying the catalyst temperature, the sensor responses can be modified, and operation regimes can be optimized for the detection of specific target compounds. The catalysts included, e.g., rhodium or platinum filaments.67-69 A test analyte pattern for 8 different electrochemical sensors (4 CO sensors, 2 hydrogen sulfide sensors, 1 sulfur dioxide sensor, and 1 nitrogen monoxide sensor) and 7 different catalyst temperature steps is shown in Figure 4.67 The sensor-response patterns vary according to temperature and sensor type upon exposure to the 19 analytes. Again, the use of the catalyst generates virtual sensors and efficiently extends the feature space.

2.2. Dynamic Methods and Transient Signals The sensor-signal evolution over time can be used to extend the feature space of a sensor array. The information content that can be extracted from a transient signal f(t) is considerably higher than that from a steady-state signal f; whereas the steady-state signal is given by a single number f, the transient signal f(t) provides a series of measurement values at discrete time intervals t. More detailed information on transient analysis can be found in section 9 of this article and also in the article by Nakamoto43 in this review issue. An example of transient signals is displayed in Figure 5, which shows how the creation of exposure steps and transients of varying length can help to discriminate between methanol and ethanol using a polymer-coated capacitive device.70 It is noteworthy that the recording of transient signals stringently requires a dedicated gas manifold70,71 (permanent gas flows, crossover valves, small dead volumes between valves and measurement chamber, and small-volume chamber; for more information, see section 4.1 of this article), so that the recorded dynamics represent the real sensor dynamics and not those of the manifold or of the gas exchange in the measurement chamber. The time required for a full exchange

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Figure 3. Sinusoidal modulation of the operation temperature of a tin oxide sensor between 200 and 400 °C (bottom) leads to characteristic frequency-dependent resistance features (upper part). Changes of the resistance (R sensor) of the micromachined sensor upon exposure to 50 ppm CO, 1 ppm NO2, and a mixture of 50 ppm CO and 1 ppm NO2 in synthetic air (50% relative humidity). Reprinted with permission from ref 47. Copyright 1997 Elsevier.

Figure 4. Analyte test data of a sensor array consisting of 8 electrochemical sensors detecting the analyte gas reaction products at 7 different catalyst temperatures (30, 100, 200, 500, 600, 750, 900 °C) of an upstream Pt filament. The analytes included a set of 10 alcohols, 2 ketones, ammonia, an amine, 2 sulfides, and 3 aldehydes, all of which are characteristic for fish freshness. Reprinted with permission from ref 67. Copyright 1994 Elsevier.

Figure 5. Sensor signals for a series of concentration steps of decreasing lengths from 160 down to 1 s. The capacitor was coated with a 4-µm-thick layer of poly(epichlorohydrin). The envelope of the response profile is highlighted in gray. It is analyte-specific and depends on the analyte absorption and desorption times in the respective polymer. Reprinted with permission from ref 70. Copyright 2006 American Chemical Society.

of the measurement chamber volume is often underestimated. A recording of the setup and manifold dynamics using sensors with very fast response times, e.g., sensors coated with very thin sensitive layers,70 is, therefore, recommendable.

A wealth of parameters can be extracted from a sensorsignal-versus-time representation and can be used as input to multicomponent-analysis or pattern-recognition algorithms. These parameters can include simple parameters like pulse heights, derivatives, and integrals calculated directly from the response curves or coefficients estimated from different models of the transient response like polynomial functions, exponential functions, or autoregressive models that have been fitted to the response curves.72 For most types of sensors (metal oxides and polymer-based sensors), the respective response times to reach equilibrium state are on the order of tens of seconds. In general, there are two different mechanisms determining the transient sensor response upon a sharp analyte concentration increase or decrease:73 (a) diffusion within the sensitive layer, whereat the diffusion processes in the measurement chamber should be significantly faster, and (b) surface or bulk reaction kinetics of the sensitive material. A nonlinear diffusionreaction model for thick-film metal-oxide sensors has been proposed by Gardner,73 and similar models have been used by other authors.74-77 The sensors included, in most cases,

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Taguchi-type thick-film sensors, and it was found that the response time is predominantly diffusion-limited (porous thick-film layer) and not reaction-limited.73 In many publications, polymer-based sensors are used, for which no reactions occur, so that bulk dissolution processes, i.e., analyte molecule diffusion into and out of the sensitivematerial matrix, determine the transient characteristics.71,78-80 Mass-sensitive devices, such as thickness-shear-mode resonators, have been used by several authors to identify, e.g., the aromas of alcohols,80 a variety of organic volatiles,71,79 or wine aroma compounds.78 The polymeric layers are usually <1 µm thick. The temporal or transient characteristics of sensor responses upon different analytes can also be induced by applying modulation techniques similar to the ones described in the preceding section on parameter variation: the exposure interval of the sensors to the analytes can be varied by actuating valves and by switching between analyte-loaded and pure carrier gas as displayed in Figure 5.70 The sensor signals in Figure 5 are given in Hertz, since on-chip electronics convert the minute capacitive signals into the frequency domain.81 A poly(epichlorohydrin)-coated (4 µm thickness) capacitive sensor has been used in these experiments, since measurements can be made very rapidly with this transducer (no extended gate time needed as for resonant sensors), and since the dynamic sensor signal neither relies on secondary effects like analyte-induced conductivity changes (conducting polymers) nor is influenced by accompanying effects such as analyte-induced film plasticization (acoustic-wave-based devices). Methanol and ethanol exposure steps of varying duration (from 160 s exposure duration down to 1 s exposure duration) were applied to the sensor. By applying long exposure intervals, all analytes reach absorption equilibrium and maximum signal amplitude, whereas for short intervals, this holds true only for fastdiffusing analytes. The sensor responses to methanol shown in Figure 5 reach saturation and sorption equilibria, even for relatively short exposure duration. For ethanol, which is a larger molecule with a smaller diffusion constant, the sensor signals do not reach equilibrium for medium or short exposure durations; as a result, the amplitude of the ethanol response begins to decrease much earlier in comparison to methanol. In conclusion, variations in the exposure interval can be used to facilitate the discrimination of analytes that belong to even the same homologous series.70 Additionally, modulation techniques to produce transients and to reveal the temporal sensor signal characteristics can be combined with any other parameter modulation in the preceding section, such as temperature modulations or the use of a catalyst.

2.3. Extracting Information in Different Domains In this last subsection on arrays of identical transducers, we will detail examples on how to extract, e.g., physical sensor data such as temperature or magnitude of flow from chemical sensors in addition to the chemical information they provide. A very simple multipurpose sensor/actuator structure offering three sensor operation modes (temperature, conductivity, and amperometric measurements) and two actuator operation modes (local heating and pH gradient control) was proposed by Langereis et al.82 and is displayed in Figure 6. The temperature can be measured along two different resistive paths between pads A and B or between C and D. By short-circuiting A and B as well as C and D, an

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Figure 6. Sketch of the multipurpose sensor structure: temperature between pads A and B or C and D, conductivity between pads A, B and C, D, and amperometric working electrode with all pads (A-D) short-circuited against an additional reference electrode. Reprinted with permission from ref 82. Copyright 1998 Elsevier.

interdigitated structure for conductivity measurements results, as has been shown for various potassium nitrate concentrations. By short-circuiting all four pads (A-D), the resulting electrode can be used for amperometric measurements against an additional reference electrode in a two-electrode setup, as has been demonstrated for different hydrogen peroxide concentrations. The heating can be performed by applying a voltage between A and B, or C and D, and by using one of the meanders or both as resistive heaters. One meander can be used as a temperature sensor. Finally, by supplying a current to all four pads (A-D) against an additional counter electrode, the aqueous environment can be electrolyzed and the local pH can be either increased or decreased (production of protons or hydroxide ions), as has been demonstrated by a coulometric titration of an aqueous acetic acid solution. In another approach, multifunctional modules have been realized on the basis of an array of ion-sensitive field-effect transistors (ISFETs).83-86 A schematic is shown in Figure 7.86 The setup is arranged around a flow-through cell hosting a so-called “hybrid module” (2 ISFETs, 1 Pt-wire counter electrode, and 1 gold generator electrode), and an Ag/AgCl reference electrode, all connected to external measurement equipment. The ISFET sensors are either pH-ISFETs with sensitive Ta2O5 films (55-58 mV/pH-unit) or enzymemodified ISFETs (penicillinase adsorbed on Ta2O5). The temperature is measured by a differential measurement of the two ISFETs operated at different working points. A flowvelocity measurement has been realized by using the generator electrode to electrochemically generate ions (H+ ions, electrolysis) and by measuring the ion concentration downstream upon their arrival at one of the ISFETs (pH-ISFET). By placing the ion generator electrode between two ISFETs, the flow direction and flow velocity can be determined. If the solution is not pumped through the setup, the diffusion of generated ions away from the generator electrode can be measured, and diffusion coefficients can be determined. An extended version including two of the “hybrid chips” in series has been detailed by the same authors, and, in the same paper, the use of an ISFET as a liquid-level sensor has been described.85 The multiparameter detection systems as described above have been used to detect, besides the physical parameters (temperature, flow, and diffusion), potassium ion concentrations (limit of detection (LOD) ) 5 µM) via valinomycincontaining poly(vinyl chloride) (PVC) membranes on the Ta2O5 gates85 and pH changes, since also the penicillin sensor (LOD ) 5 µM) detected the concentration of H+ ions resulting from an enzymatic penicillin hydrolysis.83-86 Examples for other enzyme-based ISFETs are given by the same authors.87 An open question that remains is, whether or not the concept of using chemosensors with their well-known drift- and stability problems for measuring physical param-

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Figure 7. Schematic of a multiparameter detection system with a hybrid module in a flow-through setup: PCB, printed circuit board; GE, generator electrode; CE, counter electrode; ISFET, ion-sensitive field-effect transistor; PenFET, penecillinase-modified ISFET; RE, reference electrode; PC, personal computer. Redrawn with permission from ref 86. Copyright 2003 Elsevier.

eters (temperature, flow, and liquid level) will prove itself in the long run, in particular, since rather simple fully integrated temperature/flow units are commercially available. As will be seen later in this review, the co-integration of temperature sensors has been realized with the aim to enhance the reliability of the chemical sensor signals.81,88

3. Arrays and Systems Comprising Different Transducers In this section, different monolithic and discrete sensor arrays making use of more than one transduction principle will be detailed. The information gained from the different transducers should be as orthogonal or complementary as possible (see discussion in section 6.2), i.e., different analyteinduced changes in the properties of the coating materials, such as resistivity and work function changes, or different properties of the analyte molecules themselves, such as dielectric coefficients and mass, should be exploited. The application of different transduction principles for monitoring changes in the same or in highly correlated physical properties upon analyte dosage will not provide significantly more information than applying only one transducer. In this section, several multisensor arrays will be presented, which include discrete transducers and fully integrated complex microsystems. We have categorized the different systems with regard to the thermodynamic phase (gas and liquid) they are operating in and then subcategorized them according to the types of sensitive materials.

3.1. Metal- and Metal-Oxide-Based Gas Sensors In the case of metal-oxide-based sensors, several approaches have been made to extract more than only resistance/conductance or impedance values. A rather obvious possibility is to monitor gas-reaction-induced temperature changes on the heated stage of the metal-oxide-coated sensor. The reaction of, e.g., CO, methane, or alcohols at heated metal-oxide surfaces featuring catalytic metals such as Pd or Pt leads to changes in the heat budget, which either increase or decrease the temperature of the heated structure.89-91 To explain the occurring temperature effects, all processes involved in the gas interaction process and contributing to heat budget changes have to be considered: adsorption, dissociation, surface reaction, and desorption of the products. The particular thermal gas signature is dependent on these different contributions. CO was found to provide negative

Figure 8. Simultaneously recorded sensor temperature and resistance traces upon dosage of CO at concentrations between 7 and 200 ppm to a Pt-doped tin oxide sensor at 50% relative humidity. Reprinted with permission from ref 91. Copyright 1999 Elsevier.

calorimetric signals (temperature decrease) upon surface reaction with Pt- or Pd-doped tin oxide, though the oxidation reaction and formation of CO2 is generally exothermic.54,90,91 Exemplary resistance changes (resistance decreases) and simultaneously recorded temperature changes (temperature decreases) are displayed in Figure 8.91 The thermal signature can also be recorded for temperature-controlled microhotplate devices by monitoring the changes in the heating power (or in the source-gate voltage in the case of using a heating transistor54) for maintaining a preset temperature. The recording of work function data and catalytic activity in addition to the metal-oxide resistance has been reported on by several authors.92-94 A setup schematic for such measurements is shown in Figure 9 for the example of tin oxide.94 It includes a two-electrode resistance arrangement (Taguchi-type sensor), a Kelvin probe for the work function measurements (Kelvin probe relies on the displacement of one of the two surfaces in a periodic oscillation so that a sinusoidal current is produced, which is proportional to the

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Figure 9. Setup schematic for measurements of conductance, change in work function, and catalytic activity of tin-oxide-based sensors. The catalytic activity is assessed by measuring the concentration of the formed CO2 via spectroscopy or electrochemical sensors. Reprinted with permission from ref 94. Copyright 1990 Elsevier.

work function difference between the two surfaces), and spectroscopic methods or electrochemical sensors to, e.g., detect the concentration of CO2 formed through the catalytic reaction of CO at the tin-oxide-sensor surface. Since the setup and experimental efforts are quite substantial, this combination of methods has been used prevailingly to reveal details of surface reactions and sensor mechanisms for, e.g., CO.93 Conductivity or impedance measurements provide information on band bending effects and/or changes in the relative position of the Fermi level; the work function measurements a