Biosensom & Bioelectronics 6 (1991) 461-415
Current developments in optical biochemical sensors* R. Narayanaswamy Department
of Instrumentation
and Analytical
Science, UMIST, PO Box 88, Manchester
(Received 4 May 1990; revised version received 15 September
M60 lQD, UK
1990; accepted 17 September
1990)
Abstract: By combining modern libre optics and opto-electronic instrumentation with chemical and biochemical reagent systems, it has become possible to fabricate optical biosensors. The current state of the art in this development is reviewed in this paper. Many developments describe selective and sensitive methods for sensing bioanalytes and it is likely that such a development will
continue to be a very active area of analytical research. However, these biosensing devices can be regarded as successful only if their practicality and reliability can be demonstrated. Keywords: biosensors, biochemical transducers, tibre optics, reflectance, fluorescence, immobilised reagents.
INTRODUCTION
*Paper presented at Biosensors 90, Singapore, 2-4 May W90.
optically using a ratiometric method whereby a part of the conducted light that is not affected by the measurement variable can be used for correcting any optical variations in the measurement system. The present state of optical fibre sensor technology has provided devices that have remote measurement capabilities and can be miniaturised. The problems that may be encountered with optical tibre biochemical transducers originate from ambient light interference, limited dynamic ranges, stability of reagent systems and limited response times. Ambient light interference may be eliminated by appropriate modulation of optical signals. Problems due to reagent stability, which may determine the lifetime of an optical transducer, may be compensated to some extent by employing multi-wavelength detection or by having a facility to renew the reagent phase. Rates of reaction and of diffusion determine the time reaching the steady state in the chemical/ biochemical transduction system, Current optical fibre devices utilise membranes to
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Optical tibre biochemical sensors are becoming of increasing interest because they possess several advantages for measurements in biochemistry and clinical chemistry. Optical methods are among the oldest and best established techniques for sensing chemical and biochemical analytes. The development of inexpensive, high-quality optical tibres for use in the communications industry has provided the essential component for this new technology. The several advantages offered by this sensing technology over conventional devices include electrical safety, immunity to electromagnetic interference, rugged construction in some applications and possible low cost. Optical fibre sensors do not require an external reference probe, unlike potentiometric sensors. The referencing can be carried out
Biosensors & Bioelectronics 6 ( 1991) 467-415
R Narayanaswamy
envelop the reagent phases, which offer barriers to mass transfer of analytes and thus contribute to the response time. The membrane can also prevent diffusion of the reagent away from the sensor into the measuring medium. If membranes are eliminated, the response time could be improved but the reagent phase could be exposed to possible poisoning by substances other than the analyte in the measurement medium. A further problem may be encountered with these optical sensors because many chemical and biochemical reactions for adaptation in the development of chemical transduction systems may be irreversible. Although such systems would possess high sensitivity for measurements, they can be used only as disposable or alarm-type devices. Optical tibre chemical/biochemical sensors may be classified as extrinsic or intrinsic devices, depending on the characteristics of the sensing element. While extrinsic sensors involve a sensing element that is external to the optical tibre itself, intrinsic sensors utilise the optical fibre as the sensing element. Extrinsic optical tibre sensors are the most common of these two types utilised for chemical and biochemical analyses. This paper reviews the current developments in the sensing of chemical and biochemical species through optical fibres. To understand how these sensors operate, it is important to understand the basic principles of light transmission in optical fibres. A number of books dealing with tibre optics (Lacy, 1982; Pedrotti & Pedrotti, 1987) are recommended to the reader for information on the basic physics of light transmission. Many reviews on optical fibre chemical sensors (Seitz, 1988; Ashworth & 1989) briefly explain the Narayanaswamy, essential fundamentals of tibre optics as applied in these sensors.
PRINCIPLES The basic concept of a chemical/biochemical sensor based on optical tibres is simple and can be diagrammatically represented as shown in Fig. 1. Light from a suitable source is transmitted into the tibre and directed to a region where it interacts with the measurand system or with a chemical/biochemical transducer element. This interaction results in a modulation of the optical intensity and the modulated light, which is now
Light
Photodetector
source
fibre
Feed
Mearurand Fig 1. Basic concept of an optical jibre chemical/ biochemical sensor (jkorn Narayanaswamy & &villa III (1988)).
encoded with chemical information, is collected by the same or another optical fibre and directed to a detection system. The interactions of light with the atoms or molecules comprising the transducer involve an exchange of energy and may lead to absorption, transmission, emission, scattering or reflection of light. The quantised nature of this energy transfer provides information about the composition of the system and has been exploited in spectroscopic methods of chemical and biochemical analysis. The background details of the spectroscopic techniques, that can be used in conjunction with optical tibres for chemical/biochemical sensing, and the principles ofchemical transduction are presented elsewhere (Seitz, 1988; Narayanaswamy & Sevilla, 1988a).
SENSOR DESIGN Two types of optical tibre sensors for chemical/ biochemical species have been developed. One type, which can be described as spectroscopic or plain-fibre sensors, do not contain any transducer and utilise the characteristic spectral property of the analyte species itself. In these sensors, the optical fibre functions only as a light guide, conveying light from the source to the sampling area, and from the sample to the light detection system. The light transmitted through the optical tibre interacts with the analyte species being sensed. The second type of sensor utilises a chemical or a biochemical transducer which normally contains immobilised chemical or biochemical reagents. The analyte species interacts with the transducer, whose optical
Biosensors & Bioelectronics 6 (1991) 467-475
properties are monitored through the optical tibre. This type of sensor is the more common of the two, since a great number of substances of analytical importance that need to be measured or sensed are colourless or non-luminescent. Furthermore, the incorporation of the chemical/ biochemical transducer imparts a great specificity to these sensing devices. A variety of configurations have been employed in the sensing region, and the design of the sensors can be grouped into: (i) a transmittive type where the interaction between the light and transducer occurs along the length of the optical fibre, and (ii) a non-transmittive type where the interactions take place at a distal end of the optical fibre system. These are dealt with clearly in early reviews (Narayanaswamy and Sevilla, 1988~; Seitz, 1988; Ashworth & Narayanaswamy, 1989), and are not presented here. However, representative examples in biochemical sensing are presented below together with a discussion of the trends in this new and important field. As far as biochemical sensing is concerned, optical sensors can be broadly categorised as follows:
(1) those
(2j
which respond selectively and perhaps reversibly to the chemical species in biological samples, where the transducer contains immobilised chemical reagent, e.g. pH sensor for blood measurements, and those which contain a biologically active material for the measurement of chemical species in any type of sample, where the biological/ transducer contains a biochemical reagent, e.g. enzyme-based sensor for glucose measurements.
The applications covered here are treated in the broadest sense as explained above.
APPLICATIONS Early use of optical fibres for sensing chemical species were chiefly as light carriers in conjunction with optical techniques utilising specific wavelengths of light transmitted through tibres. One such application of optical tibres in biosensing is the measurement of blood oxygen saturation (i.e. oximetry) (Polanyi & Hehir, 1962). Here, a plain-fibre sensor is used without a chemical transducer and the device is based on
Current developments in optical biochemical sensors
the fact that fully oxygenated (oxyhaemoglobin) and fully reduced haemoglobin have different absorption and reflectance spectra. Measurements made at a wavelength of about 630 nm using a reference wavelength of about 805 nm, are related to blood oxygen saturation. Several other optical biochemical sensors developed to date involve the use of immobilised chemical/biochemical reagents and these are briefly reviewed in the forthcoming sections.
pH, pCOz AND pOz SENSORS Optical fibre sensing applications have attracted by far the most attention in the development of sensors for continuous in vivo measurements of pH, pCOz and ~01. The advantages imparted to the measurements by the optical approach are particularly significant and there is a substantial potential market for these sensors for monitoring the status of critical care patients. Fibre optic sensors are considered safer for in vivo measurements than electrical devices. A few commercial systems for monitoring all three chemical species in blood are currently being investigated. For pH measurements, several indicators have been studied for monitoring in the physiological range. These include phenol red (Petersen et al., 1980) and the trisodium salt of 8-hydroxy-1,3,6pyrene sulphonic acid (HPTS) (Zhujun & Seitz, 1984~; Offenbacher et al., 1986). Phenol red is covalently immobilised and used in the sensor based on the ratio of reflected intensity at about 560 nm, where the base form of indicator absorbs, to reflected intensity at about 600 nm where neither form of the indicator absorbs. The chemical transducer has the configuration where it is placed at the tip with two separate 0.15 mm diameter optical fibres used to conduct light to and from the indicator. Because the wavelengths of measurement are in the visible region of the electromagnetic spectrum, plastic tibre could be used in the sensor. Another pH sensor based on reflectance measurements (Kirkbright et al., 1984) utilises bromothymol blue as the indicator, immobilised by adsorption on a styrenedivinyl benzene copolymer and retained at the tip of an optical tibre using a membrane. The schematic diagram of the probe is shown in Fig. 2. The probe utilises a plastic optical tibre (bundle or single) of nominal diameter 1 mm and the device is 469
Biosensors t Bioelectronics 6 (1991) 467-415
R. Narayanaswamy Reagent
Supporting
polymer
Incident light
Reflected light Fhcapsulating
Optical
Fibre
membrane Fig. 2. Cross-sectional diagram of an opticaljbre
approximately 2 mm in overall diameter. Changes in pH in the vicinity of the sensitive tip cause a variation in the attenuation of specific reflectance bands; at a wavelength of 590 nm in this case. Carbon dioxide sensors essentially consist of a pH sensor in contact with a reservoir of bicarbonate solution which is isolated from the sample by a COz-permeable membrane. At equilibrium, the pH sensed in the internal solution depends on the concentration of carbonic acid in the internal bicarbonate solution, which in turn is proportional to the partial pressure of CO* in the sample. Both phenol red and HPTS have been incorporated in COz sensors (Zhujun & Seitz, 19846; Gehrich etal., 1986) and also the fluorescence of 4methylumbelliferone (Luebbers & Opitz, 1983) has been studied. The major problem in developing a CO2 sensor is engineering the device in order to maximise the response in terms of the signal measured and the response time of the device. An optical sensor for 02 based on reflectance measurements of immobilised haemoglobin (Zhujun & Seitz, 1986) has been studied, and involves the following equilibrium: Hb+O
z-
Hb02
(Hb = haemoglobin)
This is a true chemical equilibrium requiring a mass transfer of oxygen to form HbOz. This may lead to some errors in measurement. Furthermore, it has been reported that immobilised haemoglobin degrades with time, which imposes serious limitations in the useful lifetime of this oxygen sensor. Most other optical oxygen sensors 470
reflectance pH sensor.
have been based on fluorescence quenching of an immobilised indicator. A sensor based on perylene dibutyrate adsorbed on Amberlite XAD-4 (a copolymer of styrene and divinylbenzene) has been studied for its suitability for in vivo oxygen measurements (Petersen et al., 1984). The quantitative relationship of the fluorescence of the immobilised reagent to the partial pressure of oxygen (~02) is given by the Stem-Volmer equation
IO -- 1 = kp02 I
where IO and I are the fluorescence intensities measured in the absence and presence of oxygen respectively. An oxygen-permeable porous the separates polypropylene membrane immobilised indicator from the sample and separate optical tibres conduct light to and from the indicator. This sensor has a diameter of 600 ,um and can be used to measure physiological pOz in the range O-150 Torr (O-20 kPa) with a precision of 1 Torr. The geometry of this sensor is considered suitable for in vivo blood gas measurements. ALKALI METAL ION SENSORS Chromogenic ionophores that change optical properties upon binding alkali metal ions have been developed and can serve as direct indicators in optical sensors. A chromogenic crown ether (Fig. 3) has been synthesised and shown to be potassium-ion selective when immobilised (Alder ec al., 1987). This sensor was sensitive to K+
Biosensors & Bioekctronics 6 (1991) 467-475
o/
c 0
c
4
Current developments in optical biochemical sensors
‘0
HO
10
B
Fig. 3. A chromogenic
2N
NO2
crown ether Cfrom Alder et al. (I 987)).
ions in aqueous solution in the concentration range 0~001-0~1 M. However, this reagent in the sensor was found to have substantially higher sensitivity towards Ca *+ than that towards K+, which might make it a suitable transducer for use in a calcium sensor. The binding coefficients for the immobilised reagent with Na+, K+, Ca*+ and Mg*+ are 17.7, 113, 938 and 62 respectively (Ashworth etal., 1988) which gives a selectivity ratio for Ca*+ over Na+, K+ and M$+ of 53,8 and 15 respectively. The response of the sensor to calcium (O-50 mM) in the presence and absence of other metal ions is shown in Fig. 4. Though this sensor, in the present form, may not be useful as an in vivo sensor, the study demonstrates that it may be feasible to determine alkali and alkaline earth metal ions optically at physiological concentrations by means of an immobilised chromogenic crown ether and optical fibres. An alternative approach based on ion pairing has been used to develop a reversible indicator system for sodium (Zhujun et al.. 1986). The transducer consists of an ionophore immobilised on a solid substrate, an anionic fluorophor, and a cationic polyelectrolyte that quenches fluorescence from bound fluorophor. In the absence of alkali metal ion, a combination of electrostatic and hydrophobic interactions causes the anionic fluorophor to bind to the cationic polyelectrolyte. Added alkali metal ion complexes with ionophore forming a hydrophobic cation which competitively ion pairs with the anionic fluorophor, drawing it away from the cationic polyelectrolyte and rendering it fluorescent. It was shown that the response to sodium was linear at low metal ion concentrations, selective, as established by the selectivity of the ionophore, and reversible. The sensitivity and the range of linear response depends on the concentrations of the fluorophor. Typically, it takes 1-3 min to reach 90% of steady state response.
25
CZ’
50
CONCENTRATION
ImM
Fig 4 Response versus Ca+’ ion concentration at pH 7.2, in the absence (0) of other metal ions, and in the presence (a) of Na+ (137 mM), K+ (4.3 mM) and Mg’+ (5 m.y) @om Ashworth et al. (1988)).
GLUCOSE There is also considerable biomedical interest in the development of a glucose sensor (Schultz etal., 1982; Narayanaswamy & Sevilla. 19886). An optical sensor for measurement of glucose concentrations has been developed by Schultz and co-workers (Schultz eral., 1982) which is based on the concept of affinity binding and measurement of fluorescence. This sensor is an implantable optical tibre glucose sensor that offers several advantages over more conventional electrochemical sensors. The principle of detection in this type of optical sensor involves the competitive binding of a particular metabolite (in this case, glucose) and a fluorescein-labelled 471
R. Narayanaswamy
Biosensors & Bioelectronics 6
Hollow dialysis membrane
Ianobilised Concanavalin
(1991) 467-475
A
Optical
fibret
/
Excitation
light Emitted
light
Giucose
Iluorescein-labelled dextran
Fig. 5. Schematic diagram of a competitive binding fluorescence sensor for glucose Cfrom Schultz et
analogue with receptor sites specific to the metabolite and the labelled ligand. The glucose sensor consists of a dialysis tube, on the inside of which Concanavalin A (Con A) is immobilised and fluorescein-labelled dextran bound to this natural substance. Con A has a high affinity to glucose and as a result of increasing glucose concentration (say, from blood), the fluorescent-labelled dextran will be liberated onto the optical path of the sensor. The intensity of fluorescence measured will be proportional to glucose concentration. A schematic diagram of this sensor is presented in Fig. 5. This sensor is a reversible one and has been shown to be stable for several days. It has been reported that the sensor has a linear response in the range of 50-400 mg glucose per 100 ml with a response time of 5-7 min. Such an optical sensor for blood glucose measurements could be a very important part of an insulin dispenser for diabetics. By substitution of other specitic immunoreagents, it should be possible to develop optical tibre sensors based on affinity binding, with good specificity for a variety of materials such as antibiotics and drugs. The enzyme-based glucose sensor (Narayanaswamy & Sevilla, 19886) utilised the enzyme glucose dehydrogenase immobilised on nylon, and the transducer also contains the coenzyme NAD. The enzyme-catalysed reduction of glucose in the presence ofNAD produces NADH, 412
al. (1982)).
whose fluorescence was monitored at a wavelength of 460 nm using an excitation wavelength of 340 nm. The probe response was reproducible and displayed good linearity in the concentration range of 1.1-11.0 mM glucose, with a limit 01 detection of 0.6 mM glucose.
HALOTHANE A useful optical sensor based on the use of the effect of fluorescence quenching is the measurement of concentration of the widely used inhalation narcotic halothane (Woltbeis et al., 1985). This sensor consists of a highly halothanesensitive indicator decacyclene, which is a polynuclear aromatic hydrocarbon, exposed to the sample. Interference by molecular oxygen is taken into account by a second sensor made highly sensitive towards oxygen by covering the sensor tip with a polytetrafluoroethylene (PTFE) layer. While the first sensor responds to both oxygen and halothane, the second responds to oxygen only. The two-sensor combination allows the determination of halothane or oxygen or both with a precision of + 5%. It is reported that other gases present in the inhalation gases or blood including CO, N20 or fluorans, do not interfere with the sensor system. The response time of 15-20 s for halothane and lo-15 s for oxygen for
Biosensors& Bioelectronics 6 (1991) 461-415
90% of the final values is considered short enough to allow gas analysis in the breathing circuit. Potential applications of this sensor include the continuous monitoring of halothane in blood with tibre optic catheters during operations and of anaesthetic gases in the breathing circuit.
ANTIGENS In principle, it is possible to design competitive binding sensors using immobilised antibodies as selective reagents and basing detection on the displacement of a labelled antigen by the analyte. Because antibodies are available for a whole host of antigens, antibody-based optical sensors have been the subject of considerable interest (Andrade et al., 1985). Practical reversible antibody-based sensors remain speculation only, owing to strong binding of antigen to antibody and vice versa. This causes difficulty in designing reversible sensors because the kinetics of antigen displacement are slow. One way to solve the response time problem is to use antibodies with relatively weak affinities for the antigen of interest. A weaker binding affinity could imply faster dissociation. Although this has yet to be demonstrated, weaker binding is likely to be accompanied by a loss of selectivity. Nevertheless, regenerate devices based on internal reflection for detecting and monitoring antibody-antigen reactions have been demonstrated (Sutherland etal., 1984). The antibody is immobilised on the surface of an optical libre core. The reaction of immobilised antibody with antigen in solution is detected by the evanescent wave of radiation propagating through the fibre. Because of the short penetration depth of the evanescent wave, substances are selectively excited when bound on the surface of the tibre core. Some of the resulting optical signal propagates through the libre which conducts it to a detector. After a measurement, the sensor is reactivated by changing the pH to reduce the binding constant. By this technique, it has been possible to measure immobilised-antibody binding of methotrexate by direct optical absorbance, and IgG by a two-site immunofluorimetric assay. The detection limit of immunoassay for methotrexate was about 2.7 X 10T7 M, and for IgG was 3 X IO-* M. With both immunoassays, the optical signal was monitored kinetically and was complete within 15 min. Such principles are
Current developmentsin optical biochemical sensors
useful in the study of optical immunoassay internal reflection techniques.
by
ENZYME ACTIVITY AND SUBSTRATE ANALYSIS Enzyme activity can be measured through an optical libre using immobilised substrate as the reagent. The feasibility of this approach has been demonstrated for esterases using the trisodium salt of 8-acetyl-1,3,6-pyrenetrisulphonic acid immobilised on an ion-exchange membrane as the reagent (Wolfbeis, 1986). The enzyme hydrolyses the non-fluorescent acetyl ester to the fluorescent phenol. The rate of increase in fluorescence, therefore, serves as a measure of enzyme activity. This reagent system has some attractive features in that the acetate, formed as the natural product of enzyme-catalysed hydrolysis, effectively regenerates the reagent by migration to the surface and forming the acetyl ester, after a measurement has been made. This allows a single reagent to be used for several measurements. Immobilised enzymes can be used as reagents for the determination of substrates that are converted into optically detectable products. This concept based on enzyme transduction of substrate has been demonstrated for the determination ofp-nitrophenyl phosphate using immobilised alkaline phosphatase to catalyse hydrolysis of the substrate to the coloured p-nitrophenoxide products (Arnold, 1985). The measured signal involves a steady state where the rate of product formation is balanced by the rate at which the product diffuses away from the fibre optic surface. In principle, this concept is widely applicable. For example, it can be used for enzymes that catalyse the reduction of NAD to the fluorescent product NADH. Other enzyme substrates that can be assayed this way include lactate, cholesterol, penicillin etc. An optical biosensor based on the use of intrinsic fluorescence of the enzyme has been studied, and demonstrated for the determination of lactate (Trettnak & Woltbeis, 1989). Here, the changes of intrinsic fluorescence of the enzyme lactate monooxygenase were measured during its interaction with lactate and the tibre optic device was fully reversible in the presence of molecular oxygen. The enzyme acts as both the recognising and transducing element. 473
R Narayanaswamy
ALBUMIN
An optical albumin sensor has been developed using immobilised bromocresol green as the indicator which could be reactivated (Goldfinch & Lowe, 1980). Albumin binds the base form of the indicator, effectively decreasing its pK,. Thus, the presence of albumin can be detected by the increase in the absorption of the base form of the dye, provided the actual pH of the medium is maintained constant at an appropriate value. In this work, the dye is covalently immobilised on a cellophane membrane. Adsorption of serum albumin to the membrane at pH 3.8 causes a characteristic yellow to blue-green colour change of the dye in the membrane, which is monitored. The response of the sensor is reported to be reproducible and linear over the albumin concentration range 5-35 mg ml-‘. This concept has been extended to the analysis of enzyme substrates, including penicillin G, urea and o-glucose, by co-immobilising the acid-base enzyme indicator and the appropriate (Goldfinch & Lowe, 1984).
CONCLUSIONS Optical fibres have induced a renaissance of optical methods of chemical analysis. By combining optical fibres with associated instrumentation, optical tibre biochemical sensors can be developed. The basic instrumentation associated with optical fibre transducers is simple (Narayanaswamy & Sevilla, 1988a). Along with the optical tibre, there is a light source, a photodetector, monochromators or filters and optical couplers. Nearly all these components are well advanced in their state of the art. Hence, progress in the field of optical biochemical sensors will be dependent on the development of appropriate reagent phases for the transducer. Ideally, optical tibre sensing devices should be characterised by high sensitivity, selectivity and reliability. Furthermore, the ability to perform measurements in situ (e.g. in the body), in a sitespecific fashion and on a continuous basis, would be important in the application ofthese sensors to practical analysis problems. The concept of multi-sensors for the simultaneous detection of several species has begun to attract considerable for in vivo Optical fibre sensors interest. continuous measurements of pH, oxygen and 474
Biosensom & Bioelectronics 6 ( 1991) 467-475
carbon dioxide have been bundled into a single probe, for use in medical diagnostics (Gehrich et al., 1986). Most of the applications described in this paper involve ‘proof-of-concept’ studies with the sensors, without actually demonstrating the practicality or reliability of the devices. Realistically, many of these sensors may never reach the point of practical development. Nevertheless, when one considers the abundance of chemical and/or biochemical reagents that have been successfully studied in the development of selective, and in many cases sensitive, biosensors, the motivation for developing optical fibre biosensors is well-established and it is likely that such development will continue to be a very active area of analytical research. This development also requires expertise from several areas, including indicator/reagent synthesis, polymer chemistry and biochemistry, analytical spectroscopy, libre optics and optoelectronics, because optical tibre sensor devices require an interdisciplinary approach to enhance their development. REFERENCES Alder, J. F.. Ashworth. D. C., Narayanaswamy. R, Moss, R. E. & Sutherland, I. 0. (1987). An optical potassium sensor. Analyst, 112, 1191-2. Andrade, J. D.. Vanwagenen, R. A., Gregonis, D. E.. Newby, K. & Linn, J. N. (1985). Remote fiber-optic biosensors based on evanescent-excited fluoroimmunoassay: concept and progress. IEEE Trans. Electron Devices, ED-32, 1175-9. Arnold, M. A. (1985). Enzyme-based fiber optic sensor. Anal. Chem.. 57, 565-6.
Ashworth, D. C., Huang, H. P. & Narayanaswamy, R (1988). An optical calcium ion sensor. Anal. Chim. Acta, 213, 251-7.
Ashworth. D. C. & Narayanaswamy, R. (1989). Transducer mechanisms for optical biosensors. Part 2: Transducer design, Comp. Meth. Prog. Biomed, 30, 21-31.
Gehrich, J. L., Lubbers, D. W., Opitz, N., Hansmann, D. R., Miller, W. W., Tusa, J. K. & Yafuso, M. (1986). Optical fluorescence and its application to an intravascular blood gas monitoring system. IEEE Trans. Biomed. Eng.. BME-33, 117-32. Goldfinch, M. J. & Lowe, C. R (1980). A solid-phase optoelectronic sensor for serum albumin. Anal. Biochem.. 109,216-21.
Goldfinch. M. J. & Lowe, C. R. (1984). Solid-phase optoelectronic sensors for biochemical analysis, Anal. B&hem..
138,430-6.
Biosensors & Bioelectronics 6 (1991) 467-415 Kirkbright, G. F., Narayanaswamy, R. & Welti, N. A. (1984). Fibre-optic pH probe based on the use of an immobilised calorimetric indicator, Analyst, 109, 1025-8. Lacy, E. A. (1982). Fibre Optics, Prentice-Hall, New York. Luebbers, D. W. & Opitz, N. (1983). Optical fluorescence sensors for continuous measurements of chemical concentrations in biological systems. Sew. Actuators, 4, 641-54. Narayanaswamy, R. & Sevilla III, F. (1988a). Optical fibre sensors for chemical species. J. Phys. E: Sci. Instrum., 21, 10-17. Narayanaswamy, R. & Sevilla III, F. (19886). An optical tibre probe for the determination of glucose based on immobilised glucose dehydrogenase. Analyt. Len.. 21, 1165-75. Offenbacher, H., Wolfbeis, 0. S. & Fuerlinger, E. (1986). Fluorescence optical sensors for continuous determination of near-neutral pH values. Sens. Actuators, 9, 13-84. Pedrotti. F. L. & Pedrotti, L. S. (1987). Introduction to Optics, Prentice-Hall, New York. Peterson, J. I., Fitzgerald, R. V. & Buckhold. D. K. (1984). A fiber-optic pOz sensor for physiological use. Anal. Chem., 56,62-l. Peterson, J. I., Goldstein, S. R.. Fitzgerald, R. V. & Buckhold. D. K (1980). Fiber-optic pH probe for physiological use. Anal. Chem.. 52, 864-9. Polanyi, M. L. & Hehir, R. M. (1962). In vivo oximeter with fast dynamic response. Rev. Sci. Instrum., 33, 1050-4. Schultz, J. S., Mansouri, S. & Goldstein, I. J. (1982).
Current developments in optical biochemical sensors Affinity sensor: a new technique for developing implantable sensors for glucose and other metabolites. Diabetes Care, 5, 245-53. Seitz, W. R. (1988). Chemical sensors based on immobilised indicators and fiber optics. CRC Crit. Rev. Anal. Chem., 19, 135-73. Sutherland, R. M., Daehne. C., Place, J. F. & Ringrose, A. S. (1984). Optical detection of antibody-antigen reactions at a glass-liquid interface, Clin. Chem., 30, 1533-8. Trettnak, W. & Wolfbeis, 0. S. (1989). A fully reversible fiber optic lactate biosensor based on the intrinsic fluorescence of lactate monooxygenase. Fres. Z. Anal. Chem., 334,428-30. Wolfbeis, 0. S. (1986). Fiber-optic probe for kinetic determination of enzyme activities. Anal. Chem., 58, 2874-6. Wolfbeis, 0. S., Posch, H. E. & Kroneis, H. (1985). Fiber optical fluorosensor for determination of halothane and/or oxygen. Anal. Chem., 57, 2556-61. Zhujun, Z. 8c Seitz, W. R. (1984a). A fluorescence sensor for quantifying pH in the range from 6.5 to 8.5. Anal. Chim. Acta. 160, 47-55. Zhujun, Z. & Seitz, W. R. (19848). A carbondioxide sensor based on fluorescence. Anal. Chim. Acta, 160, 305-9. Zhujun, Z. & Seitz, W. R. (1986). An optical sensor for oxygen based on immobilised haemoglobin. Anal. Chem., 58, 220-2. Zhujun. Z., Mullin, J. L. & Seitz. W. R (1986). Optical sensor for sodium based on ion-pair extraction and fluorescence. Anal. Chim. Acta, 184, 251-8.
475
Current developments in optical biochemical sensors* R. Narayanaswamy Department
of Instrumentation
and Analytical
Science, UMIST, PO Box 88, Manchester
(Received 4 May 1990; revised version received 15 September
M60 lQD, UK
1990; accepted 17 September
1990)
Abstract: By combining modern libre optics and opto-electronic instrumentation with chemical and biochemical reagent systems, it has become possible to fabricate optical biosensors. The current state of the art in this development is reviewed in this paper. Many developments describe selective and sensitive methods for sensing bioanalytes and it is likely that such a development will
continue to be a very active area of analytical research. However, these biosensing devices can be regarded as successful only if their practicality and reliability can be demonstrated. Keywords: biosensors, biochemical transducers, tibre optics, reflectance, fluorescence, immobilised reagents.
INTRODUCTION
*Paper presented at Biosensors 90, Singapore, 2-4 May W90.
optically using a ratiometric method whereby a part of the conducted light that is not affected by the measurement variable can be used for correcting any optical variations in the measurement system. The present state of optical fibre sensor technology has provided devices that have remote measurement capabilities and can be miniaturised. The problems that may be encountered with optical tibre biochemical transducers originate from ambient light interference, limited dynamic ranges, stability of reagent systems and limited response times. Ambient light interference may be eliminated by appropriate modulation of optical signals. Problems due to reagent stability, which may determine the lifetime of an optical transducer, may be compensated to some extent by employing multi-wavelength detection or by having a facility to renew the reagent phase. Rates of reaction and of diffusion determine the time reaching the steady state in the chemical/ biochemical transduction system, Current optical fibre devices utilise membranes to
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Optical tibre biochemical sensors are becoming of increasing interest because they possess several advantages for measurements in biochemistry and clinical chemistry. Optical methods are among the oldest and best established techniques for sensing chemical and biochemical analytes. The development of inexpensive, high-quality optical tibres for use in the communications industry has provided the essential component for this new technology. The several advantages offered by this sensing technology over conventional devices include electrical safety, immunity to electromagnetic interference, rugged construction in some applications and possible low cost. Optical fibre sensors do not require an external reference probe, unlike potentiometric sensors. The referencing can be carried out
Biosensors & Bioelectronics 6 ( 1991) 467-415
R Narayanaswamy
envelop the reagent phases, which offer barriers to mass transfer of analytes and thus contribute to the response time. The membrane can also prevent diffusion of the reagent away from the sensor into the measuring medium. If membranes are eliminated, the response time could be improved but the reagent phase could be exposed to possible poisoning by substances other than the analyte in the measurement medium. A further problem may be encountered with these optical sensors because many chemical and biochemical reactions for adaptation in the development of chemical transduction systems may be irreversible. Although such systems would possess high sensitivity for measurements, they can be used only as disposable or alarm-type devices. Optical tibre chemical/biochemical sensors may be classified as extrinsic or intrinsic devices, depending on the characteristics of the sensing element. While extrinsic sensors involve a sensing element that is external to the optical tibre itself, intrinsic sensors utilise the optical fibre as the sensing element. Extrinsic optical tibre sensors are the most common of these two types utilised for chemical and biochemical analyses. This paper reviews the current developments in the sensing of chemical and biochemical species through optical fibres. To understand how these sensors operate, it is important to understand the basic principles of light transmission in optical fibres. A number of books dealing with tibre optics (Lacy, 1982; Pedrotti & Pedrotti, 1987) are recommended to the reader for information on the basic physics of light transmission. Many reviews on optical fibre chemical sensors (Seitz, 1988; Ashworth & 1989) briefly explain the Narayanaswamy, essential fundamentals of tibre optics as applied in these sensors.
PRINCIPLES The basic concept of a chemical/biochemical sensor based on optical tibres is simple and can be diagrammatically represented as shown in Fig. 1. Light from a suitable source is transmitted into the tibre and directed to a region where it interacts with the measurand system or with a chemical/biochemical transducer element. This interaction results in a modulation of the optical intensity and the modulated light, which is now
Light
Photodetector
source
fibre
Feed
Mearurand Fig 1. Basic concept of an optical jibre chemical/ biochemical sensor (jkorn Narayanaswamy & &villa III (1988)).
encoded with chemical information, is collected by the same or another optical fibre and directed to a detection system. The interactions of light with the atoms or molecules comprising the transducer involve an exchange of energy and may lead to absorption, transmission, emission, scattering or reflection of light. The quantised nature of this energy transfer provides information about the composition of the system and has been exploited in spectroscopic methods of chemical and biochemical analysis. The background details of the spectroscopic techniques, that can be used in conjunction with optical tibres for chemical/biochemical sensing, and the principles ofchemical transduction are presented elsewhere (Seitz, 1988; Narayanaswamy & Sevilla, 1988a).
SENSOR DESIGN Two types of optical tibre sensors for chemical/ biochemical species have been developed. One type, which can be described as spectroscopic or plain-fibre sensors, do not contain any transducer and utilise the characteristic spectral property of the analyte species itself. In these sensors, the optical fibre functions only as a light guide, conveying light from the source to the sampling area, and from the sample to the light detection system. The light transmitted through the optical tibre interacts with the analyte species being sensed. The second type of sensor utilises a chemical or a biochemical transducer which normally contains immobilised chemical or biochemical reagents. The analyte species interacts with the transducer, whose optical
Biosensors & Bioelectronics 6 (1991) 467-475
properties are monitored through the optical tibre. This type of sensor is the more common of the two, since a great number of substances of analytical importance that need to be measured or sensed are colourless or non-luminescent. Furthermore, the incorporation of the chemical/ biochemical transducer imparts a great specificity to these sensing devices. A variety of configurations have been employed in the sensing region, and the design of the sensors can be grouped into: (i) a transmittive type where the interaction between the light and transducer occurs along the length of the optical fibre, and (ii) a non-transmittive type where the interactions take place at a distal end of the optical fibre system. These are dealt with clearly in early reviews (Narayanaswamy and Sevilla, 1988~; Seitz, 1988; Ashworth & Narayanaswamy, 1989), and are not presented here. However, representative examples in biochemical sensing are presented below together with a discussion of the trends in this new and important field. As far as biochemical sensing is concerned, optical sensors can be broadly categorised as follows:
(1) those
(2j
which respond selectively and perhaps reversibly to the chemical species in biological samples, where the transducer contains immobilised chemical reagent, e.g. pH sensor for blood measurements, and those which contain a biologically active material for the measurement of chemical species in any type of sample, where the biological/ transducer contains a biochemical reagent, e.g. enzyme-based sensor for glucose measurements.
The applications covered here are treated in the broadest sense as explained above.
APPLICATIONS Early use of optical fibres for sensing chemical species were chiefly as light carriers in conjunction with optical techniques utilising specific wavelengths of light transmitted through tibres. One such application of optical tibres in biosensing is the measurement of blood oxygen saturation (i.e. oximetry) (Polanyi & Hehir, 1962). Here, a plain-fibre sensor is used without a chemical transducer and the device is based on
Current developments in optical biochemical sensors
the fact that fully oxygenated (oxyhaemoglobin) and fully reduced haemoglobin have different absorption and reflectance spectra. Measurements made at a wavelength of about 630 nm using a reference wavelength of about 805 nm, are related to blood oxygen saturation. Several other optical biochemical sensors developed to date involve the use of immobilised chemical/biochemical reagents and these are briefly reviewed in the forthcoming sections.
pH, pCOz AND pOz SENSORS Optical fibre sensing applications have attracted by far the most attention in the development of sensors for continuous in vivo measurements of pH, pCOz and ~01. The advantages imparted to the measurements by the optical approach are particularly significant and there is a substantial potential market for these sensors for monitoring the status of critical care patients. Fibre optic sensors are considered safer for in vivo measurements than electrical devices. A few commercial systems for monitoring all three chemical species in blood are currently being investigated. For pH measurements, several indicators have been studied for monitoring in the physiological range. These include phenol red (Petersen et al., 1980) and the trisodium salt of 8-hydroxy-1,3,6pyrene sulphonic acid (HPTS) (Zhujun & Seitz, 1984~; Offenbacher et al., 1986). Phenol red is covalently immobilised and used in the sensor based on the ratio of reflected intensity at about 560 nm, where the base form of indicator absorbs, to reflected intensity at about 600 nm where neither form of the indicator absorbs. The chemical transducer has the configuration where it is placed at the tip with two separate 0.15 mm diameter optical fibres used to conduct light to and from the indicator. Because the wavelengths of measurement are in the visible region of the electromagnetic spectrum, plastic tibre could be used in the sensor. Another pH sensor based on reflectance measurements (Kirkbright et al., 1984) utilises bromothymol blue as the indicator, immobilised by adsorption on a styrenedivinyl benzene copolymer and retained at the tip of an optical tibre using a membrane. The schematic diagram of the probe is shown in Fig. 2. The probe utilises a plastic optical tibre (bundle or single) of nominal diameter 1 mm and the device is 469
Biosensors t Bioelectronics 6 (1991) 467-415
R. Narayanaswamy Reagent
Supporting
polymer
Incident light
Reflected light Fhcapsulating
Optical
Fibre
membrane Fig. 2. Cross-sectional diagram of an opticaljbre
approximately 2 mm in overall diameter. Changes in pH in the vicinity of the sensitive tip cause a variation in the attenuation of specific reflectance bands; at a wavelength of 590 nm in this case. Carbon dioxide sensors essentially consist of a pH sensor in contact with a reservoir of bicarbonate solution which is isolated from the sample by a COz-permeable membrane. At equilibrium, the pH sensed in the internal solution depends on the concentration of carbonic acid in the internal bicarbonate solution, which in turn is proportional to the partial pressure of CO* in the sample. Both phenol red and HPTS have been incorporated in COz sensors (Zhujun & Seitz, 19846; Gehrich etal., 1986) and also the fluorescence of 4methylumbelliferone (Luebbers & Opitz, 1983) has been studied. The major problem in developing a CO2 sensor is engineering the device in order to maximise the response in terms of the signal measured and the response time of the device. An optical sensor for 02 based on reflectance measurements of immobilised haemoglobin (Zhujun & Seitz, 1986) has been studied, and involves the following equilibrium: Hb+O
z-
Hb02
(Hb = haemoglobin)
This is a true chemical equilibrium requiring a mass transfer of oxygen to form HbOz. This may lead to some errors in measurement. Furthermore, it has been reported that immobilised haemoglobin degrades with time, which imposes serious limitations in the useful lifetime of this oxygen sensor. Most other optical oxygen sensors 470
reflectance pH sensor.
have been based on fluorescence quenching of an immobilised indicator. A sensor based on perylene dibutyrate adsorbed on Amberlite XAD-4 (a copolymer of styrene and divinylbenzene) has been studied for its suitability for in vivo oxygen measurements (Petersen et al., 1984). The quantitative relationship of the fluorescence of the immobilised reagent to the partial pressure of oxygen (~02) is given by the Stem-Volmer equation
IO -- 1 = kp02 I
where IO and I are the fluorescence intensities measured in the absence and presence of oxygen respectively. An oxygen-permeable porous the separates polypropylene membrane immobilised indicator from the sample and separate optical tibres conduct light to and from the indicator. This sensor has a diameter of 600 ,um and can be used to measure physiological pOz in the range O-150 Torr (O-20 kPa) with a precision of 1 Torr. The geometry of this sensor is considered suitable for in vivo blood gas measurements. ALKALI METAL ION SENSORS Chromogenic ionophores that change optical properties upon binding alkali metal ions have been developed and can serve as direct indicators in optical sensors. A chromogenic crown ether (Fig. 3) has been synthesised and shown to be potassium-ion selective when immobilised (Alder ec al., 1987). This sensor was sensitive to K+
Biosensors & Bioekctronics 6 (1991) 467-475
o/
c 0
c
4
Current developments in optical biochemical sensors
‘0
HO
10
B
Fig. 3. A chromogenic
2N
NO2
crown ether Cfrom Alder et al. (I 987)).
ions in aqueous solution in the concentration range 0~001-0~1 M. However, this reagent in the sensor was found to have substantially higher sensitivity towards Ca *+ than that towards K+, which might make it a suitable transducer for use in a calcium sensor. The binding coefficients for the immobilised reagent with Na+, K+, Ca*+ and Mg*+ are 17.7, 113, 938 and 62 respectively (Ashworth etal., 1988) which gives a selectivity ratio for Ca*+ over Na+, K+ and M$+ of 53,8 and 15 respectively. The response of the sensor to calcium (O-50 mM) in the presence and absence of other metal ions is shown in Fig. 4. Though this sensor, in the present form, may not be useful as an in vivo sensor, the study demonstrates that it may be feasible to determine alkali and alkaline earth metal ions optically at physiological concentrations by means of an immobilised chromogenic crown ether and optical fibres. An alternative approach based on ion pairing has been used to develop a reversible indicator system for sodium (Zhujun et al.. 1986). The transducer consists of an ionophore immobilised on a solid substrate, an anionic fluorophor, and a cationic polyelectrolyte that quenches fluorescence from bound fluorophor. In the absence of alkali metal ion, a combination of electrostatic and hydrophobic interactions causes the anionic fluorophor to bind to the cationic polyelectrolyte. Added alkali metal ion complexes with ionophore forming a hydrophobic cation which competitively ion pairs with the anionic fluorophor, drawing it away from the cationic polyelectrolyte and rendering it fluorescent. It was shown that the response to sodium was linear at low metal ion concentrations, selective, as established by the selectivity of the ionophore, and reversible. The sensitivity and the range of linear response depends on the concentrations of the fluorophor. Typically, it takes 1-3 min to reach 90% of steady state response.
25
CZ’
50
CONCENTRATION
ImM
Fig 4 Response versus Ca+’ ion concentration at pH 7.2, in the absence (0) of other metal ions, and in the presence (a) of Na+ (137 mM), K+ (4.3 mM) and Mg’+ (5 m.y) @om Ashworth et al. (1988)).
GLUCOSE There is also considerable biomedical interest in the development of a glucose sensor (Schultz etal., 1982; Narayanaswamy & Sevilla. 19886). An optical sensor for measurement of glucose concentrations has been developed by Schultz and co-workers (Schultz eral., 1982) which is based on the concept of affinity binding and measurement of fluorescence. This sensor is an implantable optical tibre glucose sensor that offers several advantages over more conventional electrochemical sensors. The principle of detection in this type of optical sensor involves the competitive binding of a particular metabolite (in this case, glucose) and a fluorescein-labelled 471
R. Narayanaswamy
Biosensors & Bioelectronics 6
Hollow dialysis membrane
Ianobilised Concanavalin
(1991) 467-475
A
Optical
fibret
/
Excitation
light Emitted
light
Giucose
Iluorescein-labelled dextran
Fig. 5. Schematic diagram of a competitive binding fluorescence sensor for glucose Cfrom Schultz et
analogue with receptor sites specific to the metabolite and the labelled ligand. The glucose sensor consists of a dialysis tube, on the inside of which Concanavalin A (Con A) is immobilised and fluorescein-labelled dextran bound to this natural substance. Con A has a high affinity to glucose and as a result of increasing glucose concentration (say, from blood), the fluorescent-labelled dextran will be liberated onto the optical path of the sensor. The intensity of fluorescence measured will be proportional to glucose concentration. A schematic diagram of this sensor is presented in Fig. 5. This sensor is a reversible one and has been shown to be stable for several days. It has been reported that the sensor has a linear response in the range of 50-400 mg glucose per 100 ml with a response time of 5-7 min. Such an optical sensor for blood glucose measurements could be a very important part of an insulin dispenser for diabetics. By substitution of other specitic immunoreagents, it should be possible to develop optical tibre sensors based on affinity binding, with good specificity for a variety of materials such as antibiotics and drugs. The enzyme-based glucose sensor (Narayanaswamy & Sevilla, 19886) utilised the enzyme glucose dehydrogenase immobilised on nylon, and the transducer also contains the coenzyme NAD. The enzyme-catalysed reduction of glucose in the presence ofNAD produces NADH, 412
al. (1982)).
whose fluorescence was monitored at a wavelength of 460 nm using an excitation wavelength of 340 nm. The probe response was reproducible and displayed good linearity in the concentration range of 1.1-11.0 mM glucose, with a limit 01 detection of 0.6 mM glucose.
HALOTHANE A useful optical sensor based on the use of the effect of fluorescence quenching is the measurement of concentration of the widely used inhalation narcotic halothane (Woltbeis et al., 1985). This sensor consists of a highly halothanesensitive indicator decacyclene, which is a polynuclear aromatic hydrocarbon, exposed to the sample. Interference by molecular oxygen is taken into account by a second sensor made highly sensitive towards oxygen by covering the sensor tip with a polytetrafluoroethylene (PTFE) layer. While the first sensor responds to both oxygen and halothane, the second responds to oxygen only. The two-sensor combination allows the determination of halothane or oxygen or both with a precision of + 5%. It is reported that other gases present in the inhalation gases or blood including CO, N20 or fluorans, do not interfere with the sensor system. The response time of 15-20 s for halothane and lo-15 s for oxygen for
Biosensors& Bioelectronics 6 (1991) 461-415
90% of the final values is considered short enough to allow gas analysis in the breathing circuit. Potential applications of this sensor include the continuous monitoring of halothane in blood with tibre optic catheters during operations and of anaesthetic gases in the breathing circuit.
ANTIGENS In principle, it is possible to design competitive binding sensors using immobilised antibodies as selective reagents and basing detection on the displacement of a labelled antigen by the analyte. Because antibodies are available for a whole host of antigens, antibody-based optical sensors have been the subject of considerable interest (Andrade et al., 1985). Practical reversible antibody-based sensors remain speculation only, owing to strong binding of antigen to antibody and vice versa. This causes difficulty in designing reversible sensors because the kinetics of antigen displacement are slow. One way to solve the response time problem is to use antibodies with relatively weak affinities for the antigen of interest. A weaker binding affinity could imply faster dissociation. Although this has yet to be demonstrated, weaker binding is likely to be accompanied by a loss of selectivity. Nevertheless, regenerate devices based on internal reflection for detecting and monitoring antibody-antigen reactions have been demonstrated (Sutherland etal., 1984). The antibody is immobilised on the surface of an optical libre core. The reaction of immobilised antibody with antigen in solution is detected by the evanescent wave of radiation propagating through the fibre. Because of the short penetration depth of the evanescent wave, substances are selectively excited when bound on the surface of the tibre core. Some of the resulting optical signal propagates through the libre which conducts it to a detector. After a measurement, the sensor is reactivated by changing the pH to reduce the binding constant. By this technique, it has been possible to measure immobilised-antibody binding of methotrexate by direct optical absorbance, and IgG by a two-site immunofluorimetric assay. The detection limit of immunoassay for methotrexate was about 2.7 X 10T7 M, and for IgG was 3 X IO-* M. With both immunoassays, the optical signal was monitored kinetically and was complete within 15 min. Such principles are
Current developmentsin optical biochemical sensors
useful in the study of optical immunoassay internal reflection techniques.
by
ENZYME ACTIVITY AND SUBSTRATE ANALYSIS Enzyme activity can be measured through an optical libre using immobilised substrate as the reagent. The feasibility of this approach has been demonstrated for esterases using the trisodium salt of 8-acetyl-1,3,6-pyrenetrisulphonic acid immobilised on an ion-exchange membrane as the reagent (Wolfbeis, 1986). The enzyme hydrolyses the non-fluorescent acetyl ester to the fluorescent phenol. The rate of increase in fluorescence, therefore, serves as a measure of enzyme activity. This reagent system has some attractive features in that the acetate, formed as the natural product of enzyme-catalysed hydrolysis, effectively regenerates the reagent by migration to the surface and forming the acetyl ester, after a measurement has been made. This allows a single reagent to be used for several measurements. Immobilised enzymes can be used as reagents for the determination of substrates that are converted into optically detectable products. This concept based on enzyme transduction of substrate has been demonstrated for the determination ofp-nitrophenyl phosphate using immobilised alkaline phosphatase to catalyse hydrolysis of the substrate to the coloured p-nitrophenoxide products (Arnold, 1985). The measured signal involves a steady state where the rate of product formation is balanced by the rate at which the product diffuses away from the fibre optic surface. In principle, this concept is widely applicable. For example, it can be used for enzymes that catalyse the reduction of NAD to the fluorescent product NADH. Other enzyme substrates that can be assayed this way include lactate, cholesterol, penicillin etc. An optical biosensor based on the use of intrinsic fluorescence of the enzyme has been studied, and demonstrated for the determination of lactate (Trettnak & Woltbeis, 1989). Here, the changes of intrinsic fluorescence of the enzyme lactate monooxygenase were measured during its interaction with lactate and the tibre optic device was fully reversible in the presence of molecular oxygen. The enzyme acts as both the recognising and transducing element. 473
R Narayanaswamy
ALBUMIN
An optical albumin sensor has been developed using immobilised bromocresol green as the indicator which could be reactivated (Goldfinch & Lowe, 1980). Albumin binds the base form of the indicator, effectively decreasing its pK,. Thus, the presence of albumin can be detected by the increase in the absorption of the base form of the dye, provided the actual pH of the medium is maintained constant at an appropriate value. In this work, the dye is covalently immobilised on a cellophane membrane. Adsorption of serum albumin to the membrane at pH 3.8 causes a characteristic yellow to blue-green colour change of the dye in the membrane, which is monitored. The response of the sensor is reported to be reproducible and linear over the albumin concentration range 5-35 mg ml-‘. This concept has been extended to the analysis of enzyme substrates, including penicillin G, urea and o-glucose, by co-immobilising the acid-base enzyme indicator and the appropriate (Goldfinch & Lowe, 1984).
CONCLUSIONS Optical fibres have induced a renaissance of optical methods of chemical analysis. By combining optical fibres with associated instrumentation, optical tibre biochemical sensors can be developed. The basic instrumentation associated with optical fibre transducers is simple (Narayanaswamy & Sevilla, 1988a). Along with the optical tibre, there is a light source, a photodetector, monochromators or filters and optical couplers. Nearly all these components are well advanced in their state of the art. Hence, progress in the field of optical biochemical sensors will be dependent on the development of appropriate reagent phases for the transducer. Ideally, optical tibre sensing devices should be characterised by high sensitivity, selectivity and reliability. Furthermore, the ability to perform measurements in situ (e.g. in the body), in a sitespecific fashion and on a continuous basis, would be important in the application ofthese sensors to practical analysis problems. The concept of multi-sensors for the simultaneous detection of several species has begun to attract considerable for in vivo Optical fibre sensors interest. continuous measurements of pH, oxygen and 474
Biosensom & Bioelectronics 6 ( 1991) 467-475
carbon dioxide have been bundled into a single probe, for use in medical diagnostics (Gehrich et al., 1986). Most of the applications described in this paper involve ‘proof-of-concept’ studies with the sensors, without actually demonstrating the practicality or reliability of the devices. Realistically, many of these sensors may never reach the point of practical development. Nevertheless, when one considers the abundance of chemical and/or biochemical reagents that have been successfully studied in the development of selective, and in many cases sensitive, biosensors, the motivation for developing optical fibre biosensors is well-established and it is likely that such development will continue to be a very active area of analytical research. This development also requires expertise from several areas, including indicator/reagent synthesis, polymer chemistry and biochemistry, analytical spectroscopy, libre optics and optoelectronics, because optical tibre sensor devices require an interdisciplinary approach to enhance their development. REFERENCES Alder, J. F.. Ashworth. D. C., Narayanaswamy. R, Moss, R. E. & Sutherland, I. 0. (1987). An optical potassium sensor. Analyst, 112, 1191-2. Andrade, J. D.. Vanwagenen, R. A., Gregonis, D. E.. Newby, K. & Linn, J. N. (1985). Remote fiber-optic biosensors based on evanescent-excited fluoroimmunoassay: concept and progress. IEEE Trans. Electron Devices, ED-32, 1175-9. Arnold, M. A. (1985). Enzyme-based fiber optic sensor. Anal. Chem.. 57, 565-6.
Ashworth, D. C., Huang, H. P. & Narayanaswamy, R (1988). An optical calcium ion sensor. Anal. Chim. Acta, 213, 251-7.
Ashworth. D. C. & Narayanaswamy, R. (1989). Transducer mechanisms for optical biosensors. Part 2: Transducer design, Comp. Meth. Prog. Biomed, 30, 21-31.
Gehrich, J. L., Lubbers, D. W., Opitz, N., Hansmann, D. R., Miller, W. W., Tusa, J. K. & Yafuso, M. (1986). Optical fluorescence and its application to an intravascular blood gas monitoring system. IEEE Trans. Biomed. Eng.. BME-33, 117-32. Goldfinch, M. J. & Lowe, C. R (1980). A solid-phase optoelectronic sensor for serum albumin. Anal. Biochem.. 109,216-21.
Goldfinch. M. J. & Lowe, C. R. (1984). Solid-phase optoelectronic sensors for biochemical analysis, Anal. B&hem..
138,430-6.
Biosensors & Bioelectronics 6 (1991) 467-415 Kirkbright, G. F., Narayanaswamy, R. & Welti, N. A. (1984). Fibre-optic pH probe based on the use of an immobilised calorimetric indicator, Analyst, 109, 1025-8. Lacy, E. A. (1982). Fibre Optics, Prentice-Hall, New York. Luebbers, D. W. & Opitz, N. (1983). Optical fluorescence sensors for continuous measurements of chemical concentrations in biological systems. Sew. Actuators, 4, 641-54. Narayanaswamy, R. & Sevilla III, F. (1988a). Optical fibre sensors for chemical species. J. Phys. E: Sci. Instrum., 21, 10-17. Narayanaswamy, R. & Sevilla III, F. (19886). An optical tibre probe for the determination of glucose based on immobilised glucose dehydrogenase. Analyt. Len.. 21, 1165-75. Offenbacher, H., Wolfbeis, 0. S. & Fuerlinger, E. (1986). Fluorescence optical sensors for continuous determination of near-neutral pH values. Sens. Actuators, 9, 13-84. Pedrotti. F. L. & Pedrotti, L. S. (1987). Introduction to Optics, Prentice-Hall, New York. Peterson, J. I., Fitzgerald, R. V. & Buckhold. D. K. (1984). A fiber-optic pOz sensor for physiological use. Anal. Chem., 56,62-l. Peterson, J. I., Goldstein, S. R.. Fitzgerald, R. V. & Buckhold. D. K (1980). Fiber-optic pH probe for physiological use. Anal. Chem.. 52, 864-9. Polanyi, M. L. & Hehir, R. M. (1962). In vivo oximeter with fast dynamic response. Rev. Sci. Instrum., 33, 1050-4. Schultz, J. S., Mansouri, S. & Goldstein, I. J. (1982).
Current developments in optical biochemical sensors Affinity sensor: a new technique for developing implantable sensors for glucose and other metabolites. Diabetes Care, 5, 245-53. Seitz, W. R. (1988). Chemical sensors based on immobilised indicators and fiber optics. CRC Crit. Rev. Anal. Chem., 19, 135-73. Sutherland, R. M., Daehne. C., Place, J. F. & Ringrose, A. S. (1984). Optical detection of antibody-antigen reactions at a glass-liquid interface, Clin. Chem., 30, 1533-8. Trettnak, W. & Wolfbeis, 0. S. (1989). A fully reversible fiber optic lactate biosensor based on the intrinsic fluorescence of lactate monooxygenase. Fres. Z. Anal. Chem., 334,428-30. Wolfbeis, 0. S. (1986). Fiber-optic probe for kinetic determination of enzyme activities. Anal. Chem., 58, 2874-6. Wolfbeis, 0. S., Posch, H. E. & Kroneis, H. (1985). Fiber optical fluorosensor for determination of halothane and/or oxygen. Anal. Chem., 57, 2556-61. Zhujun, Z. 8c Seitz, W. R. (1984a). A fluorescence sensor for quantifying pH in the range from 6.5 to 8.5. Anal. Chim. Acta. 160, 47-55. Zhujun, Z. & Seitz, W. R. (19848). A carbondioxide sensor based on fluorescence. Anal. Chim. Acta, 160, 305-9. Zhujun, Z. & Seitz, W. R. (1986). An optical sensor for oxygen based on immobilised haemoglobin. Anal. Chem., 58, 220-2. Zhujun. Z., Mullin, J. L. & Seitz. W. R (1986). Optical sensor for sodium based on ion-pair extraction and fluorescence. Anal. Chim. Acta, 184, 251-8.
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