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  • 1. 203© 2013 David G. Wild. Published by Elsevier Ltd. All rights reserved. http://dx.doi.org/10.1016/B978-0-08-097037-0.00014-2 A principal impetus for developing biosensor systems has been the need to produce a simple, very rapid, sensitive, and easy-to-use analytical system that does not need trained specialists to produce results. An aim for many applications is to develop a small, portable unit into which the test sample can be applied directly, without prepro- cessing. The result should be obtained within seconds, and the answer should not be subject to interference or modi- fication by the test-sample matrix, the user, or the envi- ronmental factors. Sensors that use immunological detection methods are among the most advanced type of biosensor technologies due to the ubiquity of traditional immunoassay techniques and because high detection spec- ificity is readily obtained with receptor–ligand interaction chemistry. Traditional “rapid” tests for blood glucose, fertility hor- mones, and drugs of abuse are well proven and cost-effec- tive. Immunosensors have become more important in recent years as a result of progress in point-of-care (POC) testing, most of which are “rapid” immunoassay technolo- gies, and due to an increased focus on monitoring environ- mental biohazards. POC technology made notable strides with the introduction of emergency room and doctor’s office rapid assays for cardiac stress proteins (troponins, B-type natriuretic peptide, CK-MB, and D-dimer) and the in-clinic tests for infectious agents, such as influenza A and B (Gubala et al., 2012). Among the notable emerging POC diagnostics being investigated is the WOUNDCHEK™ assay for protease status (http://www.systagenix.com/), used in monitoring inflammation in wound healing to aid in the treatment of chronic wounds such as venous leg, pressure, and diabetic ulcers. The diagnostics company Alere (formerly Biosite, Inc.) is investigating a POC test for plasma neutrophil gelatinase-associated lipocalin as an early predictive biomarker for acute kidney injury, a complication of cardiopulmonary bypass surgery (http://www.alere.com/ EN_US/products/alere-triage-ngal-test-poc-test/index. jsp). Alere is also conducting clinical trials to assess the accuracy and precision of the quantitative determination of C-reactive protein in whole blood, serum, and plasma from individuals being assessed for risk of cardiovascular disease, including a POC finger-stick test modality (http:// www.alere.com/EN_US/products/alere-cholestech-ldx- system-poc-test/index.jsp). Continuous, in vivo sensing for diagnostic monitoring and drug delivery is probably the most technically demand- ing application for immunosensors, though the field is developing with some limited commercial successes. A notable commercially produced, subcutaneous in vivo sensor is the Medtronic MiniMed Guardian Continuous Glucose Monitoring (CGM) System (http://www.medtronicdiabe- tes.com/treatmentoptions/continuousglucosemonitor- ingto). This system consists of a subcutaneous sensor and an external monitor; in clinical studies, it improved patient’s glycemic control and lowered their hemoglobin A1C values. Medtronic has combined this sensor with an insulin pump to better manage blood glucose in diabetic patients. Aside from diabetes care, this technology is being evaluated in clinical trials for its utility in other indications, such as use in diagnostic monitoring of patients with severe trauma and or thermal injury. A similar in vivo sensor, called the SEVEN® Plus, is marketed by Dexcom®, Inc. Abbott Dia- betes Care received approval for a subcutaneous CGM sen- sor device called FreeStyle Navigator in 2008, but in 2011 discontinued sale of the device in the United States, pre- sumably due to lack of market penetration. A promising but commercially unsuccessful noninvasive CGM system was the Glucowatch, a wrist watch-like device that uses iontophoresis to extract sample though the skin and standard amperometric enzyme detection to mea- sure glucose concentration (Diabetes Research Children Network, 2005). Originally marketed by Cygnus and later acquired by Johnson and Johnson, this device received Food and Drug Administration approval in 2002 but limi- tations in its stability, accuracy, and longevity resulted in it being discontinued in 2007. Detection of chemical and biological warfare agents is an important driver of sensor development, and many pro- totype detectors in this application utilize receptor–ligand or enzyme–reporter formats. For example, the company Tetracore produces a number of conventional test strip- based enzyme-linked immunosorbent assays for bacterial pathogens, such as anthrax, Yersinia pestis, and Francisella tularensis, and biotoxins, such as ricin and botulinum toxin (http://www.tetracore.com/elisa-kits/index.html). Northrop Grumman and Luminex are developing a highly multiplexed optical array biosensor for real-time biohazard detection. This sensor can detect a variety of chemical toxins and infectious agents in a complex back- ground such as human serum. Applications for immuno- sensors related to biohazard detection also include pollution monitoring, food safety, and industrial process control (Van Dorst et al., 2010). Overview A biosensor uses a biological system to measure a substance and differentiate this from other substances in a test sam- ple. It is a measurement device that is comprised of three components: a biological receptor of appropriate speci- ficity for the analyte (or test material to be measured); a transducer to convert the recognition event into a suitable physical signal; and a detection system, including analysis and processing, that is usually electrical. The physical sig- nal can, for example, be acoustic, electromagnetic, or mechanical. To bring together these three components for Immunological Biosensors Jason Reed (jreed@cnsi.ucla.edu) James K. Gimzewski (gimzewski@cnsi.ucla.edu) C H A P T E R 2.11
  • 2. 204 The Immunoassay Handbook the development of a biosensor, therefore, requires an integrated, multidisciplinary team of biologists, chemists, physicists, engineers, and computer experts. This blend of skilled personnel is not found in every establishment; so biosensor development has resulted primarily from inter- institutional collaborative projects or within industry. Generally, in clinical chemistry analyses, a lower thresh- old of detection at around micromolar concentrations is sat- isfactory for most analytes (Gubala et al., 2012). The demands for hormone measurements in the endocrinology clinic pushed the threshold for the detection to around nanomolar concentrations and then below this to the picomolar range; these measurements were performed almost exclusively with biological binding assays, often immunoassays. For some more recent applications, low detection limits are needed, for example, with DNA fragments or pesticide residues. The ability to detect a small amount of material in a test sample is determined by the signal-to-noise ratio. Jackson and Ekins (1983) have shown that the detection limit of an immunoassay system is directly proportional to the relative error in the signal and inversely proportional to the equilib- rium constant of the binding reagent. So, a decrease in the measurement error and a high affinity constant (a low-valued equilibrium constant) would contribute to a low detection limit for the analysis (see PRINCIPLES OF COMPETITIVE AND IMMUNOMETRIC ASSAYS (INCLUDING ELISA)). Most immuno- logical methods of analysis are subject to more-or-less severe interference from other components in the complex sample matrix. Washing or other additional manipulations are often used to minimize interference effects. The design features of a biosensor are little different from those of any modern laboratory instrumentation system: G instruments should be small, self-contained, cheap, and robust, capable of interfacing with existing central lab- oratory systems; G the user interface should be simple, for use by unskilled operators; G no volumetric measurement of the specimen, e.g., by pipette, should be necessary; G the test specimen alone should be added, with no fur- ther reagent addition being required; G results should be unaffected by the test specimen matrix, e.g., water, whole blood, serum, urine, or plasma; G the time between presentation of the specimen and final result should be rapid and ideally less than 5min; G built-in standardization and controls are required; G an easily understood record of the results should be available; G the detection limit should be appropriate to the analyte and in the sub-picomolar range for the most sensitive systems; G a wide analytical range is required if the same biosensor is to be adapted to many different analytes: for a generic analytical system, a capability for immunochemistry, clinical chemistry, enzymology, DNA probe measure- ments, and a variety of other applications is desirable; G the potential for simultaneous measurement of multi- ple analytes should be considered; G there should be a good correlation of results with already-established test methods; and G biosensor consumables must be cheap to manufacture in bulk and readily available. Amperometric Sensors Enzymes provide an attractive method for signal amplifica- tion. The continual turnover of a substrate generates a cas- cade signal that is large and can, therefore, be measured quite easily. Amperometric biosensors generally use reduction– oxidation (redox) enzyme systems. In the simplest case, a redox enzyme is immobilized by some convenient proce- dure at an electrode surface. The electrode is held at a fixed potential, adjusted so that electrons arising from an oxidized substrate are transferred to the electrode (or vice versa for a reduction reaction), and this regenerates the active form of a cofactor for another redox cycle by the enzyme. The speci- ficity of the reaction is determined by the enzyme. Because the rate of enzymatic reaction at a fixed temperature and pH is directly proportional to the substrate concentration, the current produced at the electrode is proportional to the rate of modification of the substrate by the enzyme (see Fig. 1). The rate of an enzymatic reaction is also dependent on ions, cofactors, and competitive or noncompetitive inhibi- tors (or activators) present in the test sample, and any redox compound present, such as oxygen, ascorbate, thi- ols, or certain drugs, can interfere with the reaction. To circumvent these interference effects, chemical elec- tron acceptors are used as mediators. For example, the ferrocene–ferrocenium redox system has been used in the mediation of electron transfer from glucose oxidase to graphite electrodes. FIGURE 1 An amperometric biosensor arrangement. A mediator is used to transfer electrons from an electrode to an enzyme-catalyzed redox reaction.
  • 3. 205CHAPTER 2.11 Immunological Biosensors Immunoassay methods are normally required for mea- surements of analyte concentrations in the sub-millimolar range. Various methods of detection of antibody–antigen interactions using enzyme-labeled reagents have been tried, coupling the enzymatic redox reaction to an amperometric detection system (Majumdar, 2002). However, the efficiency of coupling of the biological electron-generating steps to the electrode is not fully characterized. Additionally, the effects of interfering substances are greater when the substrate con- centration falls below millimolar. These and other technical difficulties, such as attaching the reagents to the membrane or the electrode, have hindered the development of systems based on this principle for immunosensing. Piezoelectric Mass Sensors Mass detection sensors are among the most widely used microanalytical technologies. These methods rely, in gen- eral, on measuring the changes in vibrational resonant fre- quency of piezoelectric quartz oscillators that result from changes in mass on the oscillator’s surface (see Fig. 2). Common configurations are the quartz crystal micro- balance (QCM) and the surface acoustic wave (SAW) devices. The QCM device consists of a quartz crystal disk driven by electrodes on either face. The mass of analytes that bind to the sensor is measured as a change in the crys- tal’s resonant frequency. This type of sensor is also known as a thickness-shear mode (TSM) device. The mass determination in a TSM sensor is given in terms of the Sauerbrey relation: In the SAW sensor, an acoustic wave is created by applying an alternating voltage to a metallized, interdigitated elec- trode plated onto one end of the thin piezoelectric planar substrate of the device. The acoustic wave acts on a sym- metric transducer at the opposite end of the substrate, where the energy is converted back into an electrical sig- nal. The SAW device can be altered for particular applica- tions by varying the interdigital spacing of the transducers, the distance between the transducers, and the thickness of the substrate. SAW devices respond to changes in the surface wave amplitude or, more commonly, in the velocity of the acoustic wave when it interacts with molecules bound to the surface of the substrate. The thin layer of bound mate- rial alters the elasticity, density, viscosity, and conductivity of the SAW substrate. In some early work, it is apparent that the exquisite temperature sensitivity of these devices was not taken into account. There are many modes of acoustic wave that can propagate on the SAW devices, which can be used variously for sensing applications of gases, liquids, or deposited solids. Alternative acoustic modes induce particle motion on both surfaces of the substrate; these devices are sometimes referred to as acoustic wave-guide devices. The modes can be grouped into families and, by suitable selection of the type of mode, high sensitivity and minimization of non- specific interferences can be achieved. As experience has been gained, it has become apparent that, for applications such as immunosensors, SAW devices should operate in FIGURE 2 Three different forms of piezoelectric biosensors. (Redrawn from Ward and Buttry, 1990).
  • 4. 206 The Immunoassay Handbook shear horizontal (SH) or surface transverse wave (STW) modes (Collings & Caruso, 1997). The so-called Love-wave device, sometimes referred to as a surface-skimming bulk wave, is obtained when a layer of the appropriate thickness and acoustic properties is deposited over a conventional SH device. The energy of the bulk wave is concentrated in the guiding layer because the shear acoustic velocity in this layer is lower than in the substrate, leading to the alternative name of the surface-guided SH wave. The sensitivity to mass load- ing is increased by focusing the energy in this layer, which is dependent on the layer thickness and its acoustic properties. The latter devices have the advantage of avoiding radia- tion losses in liquids, yet have much better sensitivity to mass loading by concentrating the energy near the surface. The SH wave may be guided along device surfaces by grat- ings as well as over-layers, as either SH waves or STWs. The improved sensitivity compared with simple SH devices, without much greater complexity in fabrication, holds much promise for their use in liquid-phase biosens- ing. This is because the viscous loading contribution of the solvent, in its attenuation of the shear wave in other modes, shifts the resonance of the transducer with consequent alteration in sensitivity. Using conventional antibody–antigen interactions on thin substrates, detection limits of nanomolar down to picomolar concentrations have been claimed with “ideal” test samples, though this is perhaps unlikely to be achieved with typical clinical specimens. Manufacture of the devices to produce cheap, disposable units with uniform character- istics may be relatively easy, although provision of multi- analyte analysis on a single unit would provide a challenge. Microcantilever Sensors The microcantilever is a particularly versatile class of sen- sor that is unique in its combination of simplicity, sensitiv- ity, and potential for miniaturization. Microcantilevers can sense and quantitate biological proteins, nucleic acids and a variety of organic and inorganic chemical species. The principal mechanism of action for sensing is a nanoscopic deflection caused by receptor- and ligand-induced stress on one face of the microcantilever. The deflection signal can be recorded with an optical lever, an interferometer, or a piezoresistive element. When operated in “active mode,” cantilever sensors are induced to oscillate at their resonant frequency and function very much like the mass detection sensors described above. Gimzewski and colleagues produced the first cantilever array immunosensor, which could distinguish species-specific binding of protein A to rabbit immunoglobulin G. The same sensor design could also detect single base-pair mismatches in DNA oligonucleotide hybridization experiments (Fritz et al., 2000). Another iteration of this device, a “nanomechan- ical nose,” used an array of nonspecific polymer probes to distinguish hydrogen, primary alcohols, natural flavors, and water vapor in air (Lang et al., 1999). Microcantilevers have been successfully used to detect prostate-specific antigen, heavy metal ions, neurotoxins, and glucose (Majumdar, 2002). The latter application utilized cantilever-bound glucose oxidase enzyme as a detector–reporter element. For miniaturization, fully integrated cantilever sensor systems have been fabricated using a complementary metal-oxide- semiconductor (CMOS) process to incorporate signal- conditioning electronics, on-chip actuation and self-sensing to eliminate the need for external optics (Boisen et al., 2000; Hierlemann and Baltes, 2003). Live-Cell Nanomechanical Assays An emerging class of biosensors utilizes live cells them- selves as the sensing element. This approach has the advantage of being able to integrate biological and chemi- cal sensing into a true functional readout: the cell’s physi- ologic state (Biran et al., 2003; Peters et al., 2010; Wozei et al., 2006). This concept involves a focus on one class of cell-based sensing that relies on a cell’s nanomechanical properties, such as stiffness, time-dependant deformabil- ity, or mass, to infer cell state in response to various stim- uli. A key advantage of this approach is that it does not require labeling with fluorescent proteins or optically active dyes, and thus both cultured cell lines and material derived from whole organisms can be used. It has been shown that cancerous cells can be distin- guished from normal cells present in a biological aspirate taken from patents with mesothelioma by measuring the cells’ elastic modulus (cancerous cells are softer) with atomic force microscopy (AFM) (Cross et al., 2007). This work has been extended to the detection of urinary and ovarian cancers (Rao et al., 2009; Sharma et al., 2011). Similarly, cell elasticity can be used as a real-time mea- surement of response to therapeutics such as cisplatin and green tea extract (Sharma et al., 2011; Cross et al., 2011). While AFM has the drawback of being a single-point contact technique, and thus has a relatively low throughput, cellular nanomechanical measurements can be parallelized using arrays of nanoparticles actuated magnetically and tracked using interference microscopy (Reed et al., 2008b). This method has been used to characterize the effect of cytoskeleton-altering drugs in populations of fibroblast cells in culture and the effect of neurotoxin in whole zebra fish embryos (Reed et al., 2009; Reed et al., 2008a). Rapid nanoscale measurement of single cell mass has become possible using quantitative phase-imaging tech- niques and microresonators (Reed et al., 2011; Popescu et al., 2008; Godin et al., 2007). This type of measurement can be used to characterize the growth of single cells over the course of minutes and to record the instantaneous response of cells to external chemical and physical stimuli. The technique has been used to rapidly detect chemother- apeutic response in multiple myelomas and in breast can- cer cells (Reed et al., 2011; Chun, 2012). Conclusions The amperometric sensors have been intensively investi- gated over the past 35 years since Clark and Lyons described
  • 5. 207CHAPTER 2.11 Immunological Biosensors the first glucose sensor (1962). It is notable that the most advanced commercial glucose sensors still use electrochem- ical detection methods, while nanomechanical sensors, such as microcantilevers, are likely candidates as the general- purpose bioanalytical sensors of the future (Arlett et al., 2011). They are easy to produce and have the potential for use over a wide range of analyte concentrations. They also approach the low detection limits required for immuno- sensing, while the reproducibility of results is similar to that achieved by conventional immunoassays. These devices rely on relatively well understood physical principles, and the associated instrumentation is quite simple to assemble. A major development in the biosensor field is directed at designing a reaction cell in which both the kinetics of ana- lytical reaction and flow characteristics at the sensor sur- face are optimized for ease of manufacture and use. The key to reproducible manufacture and consequent ease of use of these biosensor systems is the uniformity of physical and chemical properties of the biophysical interface. Instrumentation developments focus on miniaturization, improved signal processing, and analyzing multiple chemi- cal species simultaneously on the same reaction-cell sensor surface. The chemistry of reagent deposition, assay format, and other aspects related to the use of the reagents are, in general, not very different from those encountered with immunoassay systems. Attention has to be given to use the assay formats in which the signal change is large enough to measure. The assay system should be a homogeneous one, with no separation or measurement step required; so the design of the assay format becomes in some respects easier than for a conventional immunoassay. References Arlett, J.L., Myers, E.B. and Roukes, M.L. Comparative advantages of mechanical biosensors. Nat. Nanotechnol. 6, 203–215 (2011). 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