An Integrated Metal Clad Leaky Waveguide Sensor for Detection of

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Anal. Chem. 2005, 77, 232-242

An Integrated Metal Clad Leaky Waveguide Sensor for Detection of Bacteria Mohammed Zourob,† Stephan Mohr,† Bernard J. Treves Brown,† Peter R. Fielden,† Martin B. McDonnell,‡ and Nicholas J. Goddard*,†

Department of Instrumentation and Analytical Science, The University of Manchester, P.O. Box 88, Manchester M60 1QD, U.K., and Dstl, Porton Down, Salisbury, Wiltshire SP4 0JQ, U.K.

An integrated optical metal clad leaky waveguide (MCLW) sensor device has been developed for the detection of bacteria. This is more sensitive than waveguide sensors currently in use. The MCLW device has been fabricated to extend the evanescent field to provide significant light intensity over the entire volume of the bacteria bound on the chip surface within this field. This in turn increases the interaction of the light with the entire volume of the bacteria. MCLW devices have been used for detecting refractive index changes, scattering, and fluorescence from bacterial spores captured on an immobilized antibody. The detection limit of Bacillus subtilis var. niger bacterial spores using refractive index detection was 8 × 104 spores/mL. The scattering intensity of the BG spores was found to be three times greater than the scattering intensity generated using surface plasmon resonance. The extended light propagation along the direction of flow for a few millimeters provides an effective interrogation approach to increase the area of detection to detect low concentrations down to 1 × 104 spores/mL. The sensor was then optimized by studying the key factors affecting sensor performance including changing the pH of the medium, type of antibody immobilization matrix, sensor surface regeneration approaches, and longevity of the sensor. Bioterrorism has become a great international concern following Bacillus anthracis-related incidents in the United States. The need for on-line and point-of-use analysis devices that can produce a rapid, accurate, sensitive, and cost-effective analysis, as well as specific detection and identification of biological agents, is the first step in responding to these threats. The current need for realtime detection of bacteria in many applications has fueled the development of sensors for the detection of pathogenic bacteria. Examples include monitoring the quality of indoor air and potable water, detecting spoilage in the food industries, controlling contamination in fossil and nuclear power plants and wastewater treatment plants, and in the military and defensive forces. Massive epidemics caused by pathogenic bacteria have been reported throughout human history. Yersinia pestis, the causative * To whom correspondence should be addressed. Tel: (+44) 161-2004895. Fax: 161-2004911. E-mail: [email protected]. † The University of Manchester. ‡ Dstl.

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agent of bubonic and pneumonic plague, has afflicted humans for many centuries. The fact that plague continues to be endemic in many parts of the world indicates the ever present danger of pathogenic bacteria. Rapid detection and identification of bacteria is essential if a biosensor is to provide the information needed to take effective countermeasures to enable protection from pathogenic microorganisms.1 It is estimated that in the United States alone 76 million cases of food-borne illnesses per year associated with bacterial contamination of meat, poultry, and eggs occur, leading to more than 5000 deaths. U.S. estimates have placed the economic impact of such food-borne illness as high as $ 14 billion/year.2,3 In the U.K., the public laboratory service has indicated that, in 2001, there were 85 468 food poisoning notifications, which represent a 600% increase from 1982. Another serious problem is hospital infection by opportunistic enterobacteria as a result of their high resistance to antibiotics. It was suggested by Whyte et al.4 that bacteriological standards of airborne bacteria at critical sites, such as in operating theaters and intensive care units, should not exceed 10 colony forming units (cfu) per cubic meter. They have also suggested an upper limit of 20 cfu at peripheral sites of all important hospital departments. Bacteria are also associated with biological fouling (biofouling) and microbially induced corrosion (MIC), which pose serious problems in industrial water-handling systems, i.e., corrosion in fluid conduits, mechanical parts, and other construction materials. The annual worldwide cost of combating biofouling and MIC has been estimated to be billions of U.S. dollars.5 Hitherto a large number of different technologies have been developed for detection of bacteria utilizing optical, electrochemical, biochemical, and physical properties, as well as the conventional detection methods such as dry weight measurement, viable counting, and turbidity measurements. Conventional microbiological methods for determining the cell counts of bacteria employ selective culture, biochemical, and serological characterization. (1) Koch, S.; Wolf, H.; Danapel, C.; Feller, K. Biosens. Bioelectron. 2000, 14, 779-784. (2) Buzby, J. C.; Roberts, T.; Lin, J.; McDonald, J. M. USDA, Economic Report 741, Washington, DC, 1996. (3) Mead, P. S.; Slutsker, L.; Dietz, V.; McCaige, L. F.; Bresse, J. S.; Shapiro, C.; Griffin, P. M.; Tauxe, R. V. Emerg. Infect. Dis. 1999, 5, 607-725. (4) Whyte, W.; Lidwell, O. M.; Lowbury, E. J.; Blowers, R. J. Hosp. Infect. 1983, 4, 133-139. (5) Zeikus, G., Johnson, E. A., Eds. Mixed Cultures in Biotechnology; McGrawHill: New York, 1991; pp 341-372. 10.1021/ac049627g CCC: $30.25

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Although these achieve sensitive and selective bacterial detection, they typically require days to weeks to yield a result. Enzymelinked immuosorbent assay (ELISA) methods have been used for the detection of bacteria down to 1 × 103 cells/mL but are quite slow as incubation and washing steps are required.6 Solid-phase ELISA has been used for detection of bacteria, which is based on using fused-silica capillaries as a support for immobilizing the antibodies; however, it provides only semiquantitative information.1 Dipstick tests were devised to shorten the analysis time. A dipstick immunoassay to detect Escherichia coli O157:H7 in ground beef using a sandwich assay took 16 h.7 Oxoid (Oxoid Limited, Hampshire, U.K.) has a qualitative rapid test for Listeria on the market that uses primary and secondary enrichment.8 An immunomagnetic concentration procedure can detect between 101 and 103 Salmonella cells/mL, but the bacteria require counting after separation from the beads.6 A flow injection analysis system was able to detect E. coli in less than 30 min with a detection limit of 5 × 107 cfu/mL.9 Similarly, the detection limit of E. coli by a glucose biosensor coupled to microdialysis sampling in a flow injection analysis system was 106 cfu/mL, with an analysis time of over 5 h.10 The combination of immunomagnetic separation and flow cytometry enabled the detection of 103 cells/mL E. coli O157: H7 within 1 h.11 However, the equipment is complex and expensive and is not portable. In another study, Perez et al.,12 using immunomagnetic separation with mediated flow injection analysis, achieved an amperometric detection of viable E. coli O157:H7 at concentrations down to 105 cfu/mL in 2 h. Salmonella was detected at a concentration of 104 cfu/mL using the commercially available BAX polymerase chain reaction (PCR) system, but the method required between 24 and 26 h to perform.13 Recently a portable PCR instrument was used for identification of bacteria in minutes.14,15 The AG-9600 AmpliSensor analyzer (an automated fluorescence-based system detection of PCR) was found to have a detection limit of less than 2 cfu/mL Salmonella typhimurium per PCR reaction. Partis and Newton16 reported that the ELISA method is more robust, as PCR is more susceptible to contamination, and requires a high level of laboratory skills to avoid carryover when used for different samples. Piezoelectric devices have been used for detection of E. coli concentration over the concentration range 1 × 106-5 × 108 cells/mL.17 A quartz crystal microbalance detected 1 × 106 cells/mL E. coli, Campylobacter, Schigella, Yersinia, and Salmonella.18 A surface acoustic wave (6) Hobson, N. S.; Tothill, I.; Turner, A. P. F. Biosens. Bioelectron. 1997, 61, 279-286. (7) Kim, M. S.; Doyle, M. P. Appl. Environ. Microbiol. 1992, 58, 1764-1767. (8) http://www.oxoid.com, 10-02-2003. (9) Bouvrette, P.; Loung, J. Int. J. Food Microbiol. 1995, 27, 129-137. (10) Palmisano, F.; De Santis, A.; Tantillo, G.; Volpicella, T.; Zamonin, P. G. Analyst 1997, 122, 1125-1128. (11) Seo, K. H.; Brackett, R. E.; Frank, J. F.; Hilliard, S. J. Food Prot. 1998, 61, 812-816. (12) Perez, F. G.; Mascini, M.; Tothill, I. E.; Manning, A. P. F. Anal. Chem. 1998, 70, 2380-2386. (13) Bailey, J. J. Food Prot. 1998, 61, 792-795. (14) Belgrader, P.; Benett, W.; Hadley, D.; Richards, J.; Statton, P.; Mariella, R., Jr.; Milanovich, F. Science 1999, 284, 484-449. (15) Higgins, J. A.; Nasarabadi, S.; Karns, J. S.; Shelton, D. R.; Cooper, M.; Gbakima, A.; Koopman, R. P. Biosens. Bioelectron. 2003, 18, 1115-1123 (16) Partis, L.; Newton, K. Australian Government Analytical Laboratories Puplic Interest Program Research and Development Report Series 98-6, 1998. (17) Muramatsu, H.; Watanabe, Y.; Hikuma, M.; Atka, T.; Kubo, I.; Tamara, E.; Karube, I. Anal. Lett. 1989, 22, 2155-2166. (18) Konig, B.; Gratzel, M. Anal. Lett. 1993, 26, 1567-1585.

device was used by different workers19-21 for the detection of S. typhymurium bacteria. The sensor had a low detection limit of a few hundred cells/mL and a response time of less than 100 s over the concentration range of 102-1010 cells/mL. These piezoelectric sensors are attractive because they can be used for direct, labelfree detection of bacteria; however, the disadvantages of such sensors are the relatively long bacteria-sensor incubation time required, the numerous washing and drying steps, and the problem of surface regeneration.22 Pyun et al.23 detected E. coli (K12 and J5) in natural water using an immunoaffinity layer on the acoustogravimetric flexural plate wave transducer. The detection limit of this biosensor was reported to be in the range of 3.0 × 105-6.2 × 107 bacterial cells/mL. The signals were amplified up to 5-fold by using secondary E. coli antibody coupled with microspheres. A flow injection amperometric immunofiltration assay system has been used to detect 50 cells/mL E. coli and Salmonella with an overall analysis time of 35 min involving using many steps and an antibody conjugated with horseradish peroxidase (HRP).24,25 Gau et al.26 reported a sensor for amperometric detection of E. coli based on the integration of a microelectromechanical system with self-assembled monolayers of streptavidin, which capture rRNA from E. coli, DNA hybridization and enzyme amplification techniques. The system was capable of detecting 103 cells/mL E. coli in 40 min. A Bead ARay Counter (BARC) biosensor was used by Edelstein et al.27 to detect and identify biological warfare agents using DNA hybridization, magnetic microbeads, and a giant magnetoresistive sensor. Optical biosensors have been the subject of intense interest over the past two decades. This is due to the numerous advantages provided by optical methods, such as they can be miniaturized, have multiplexing capabilities, and because optical biosensor technology combines rapid response times with high sensitivity for analyte evaluation. Hence, the use of such sensors in the realtime detection of bacteria appears promising.28 Fluorescent nucleic acid stain (SYTO 13) has been demonstrated for optical bacterial detection in aerosols and aqueous samples by Chuang et al.29 However, the sensor is nonspecific as it detects all bacterial species since all organisms contain nucleic acid. Nogami et al.30 reported a system that can detect all kinds of bacteria, including growing, dead, and resting cells of bacteria in ultrapure water. These workers used a method that works by (19) Pathirana, S.; Barbaree, J.; Chin, B.; Hartell, M.; Neely, W. Biosesns. Bioelectron. 2000, 15, 135-141. (20) Prusak-Sochaczewski, E.; Luong, J. Enzyme Microbiol. Technol. 1990, 12, 173-177. (21) Howe, E.; Harding, G. Biosesns. Bioelectron. 2000, 15, 641-649. (22) Wadkins, R. M.; Golden, J. P.; Pritsiolas, L. M.; Ligler, F. S. Biosens. Bioelectron. 1998, 13, 407-415. (23) Pyun, J.; Beutel, H.; Meyer, J.; Ruf, H. Biosens. Bioelectron. 1998, 13, 839845. (24) Abdel-Hamid, I.; Ivnitski, D.; Atanasov, P.; Wilkins, E. Anal. Chem. Acta 1999, 399, 99-108. (25) Abdel-Hamid, I.; Ivnitski, D.; Atanasov, P.; Wilkins, E. Biosens. Bioelectron. 1999, 14, 309-316. (26) Gau, J.-J.; Lan, E. H.; Dunn, B.; Ho, C.-M.; Woo, J. C. S. Biosens. Bioelectron. 2001, 16, 745-755. (27) Edelstein, R.; Tamanaha, C.; Sheehan, P.; Miller, M.; Baselt, D.; Whitman, L.; Colton, R. Biosens. Bioelectron. 2000, 14, 805-813. (28) Gill, Harrison, J.; Holwill, I.; Lowe, P. A.; Hoare, M. Protein Pept. Lett. 1996, 3, 199-206. (29) Chuang, H.; Patrick, M.; Tabacco, M. Anal. Chem. 2001, 73, 462-466. (30) Nogami, T.; Ohto, T.; Kawaguchi, O.; Zaitsu, Y.; Sasaki, S. Anal. Chem. 1998, 70, 5296-5301.

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trapping the bacterial cell in a filter and detecting the chemiluminescence emitted by the bacterial DNA, which is marked with an anti-DNA antibody conjugated with HRP. Diffraction grating intensity patterns from bacteria captured from solution to regular arrays of antibodies have been used for detection of concentrations equal to, or higher than, 106 cells/mL.31,32 A number of different types of optical evanescent waveguide biological sensors have been reported by employing surface plasmon resonance (SPR),33-35 resonant mirror (RM),36 and fiber optical techniques.37-40 MANTIS, a portable automated fiber-optic biosensor prototype with multianalyte capabilities, was introduced.41 After numerous improvements, the MANTIS was eventually renamed the RAPTOR. This instrument was used for the detection of Bacillus subtilis var. niger (BG) spores using fluoroimmunoassay in the field. The detection limit for this instrument was 105 cells/mL in 10 min.42 Rowe et al.43 utilized a standard sandwich immunoassay format with antigen-specific “capture” antibodies immobilized in a patterned array on the surface of a planar waveguide. The assay needed 14 min to detect 105 cfu/ mL BG. An evanescent waveguide Mark 1.5 instrument (IVD systems, Santa Barbara, CA) was used to detect BG spores and Erwinia herbicola (EH) (live bacterium) as mimics for biological warfare agents. The method used involves mixing the analyte with a fluorescent-labeled antibody that is brought in contact with the capture antibody array. The detection limit for BG and EH was 5 × 105 cfu/mL.44 A resonant mirror biosensor, an IAsys instrument (Affinity Sensors, Cambridge, U.K.), has been used to distinguish between bacterial strains on the basis of difference in cell surface proteins. Staphylococcus aureus (Cowan-1), which produce protein-A at the cell surface, were detected by binding to human IgG. The detection limit for S. aureus (Cowan-1) was quoted as 8 × 106-8 × 107 cells/mL, S. aureus strain (Wood-46), which does not express protein A, was not detected. The sensitivity of detection was increased by a 1000-fold when using a human IgGcolloidal gold complex (30 nm in diameter) in a sandwich assay format.36 A Biacore SPR biosensor has been reported for the real-time detection binding of E. coli O157:H7 with a specific antibody.33,34 (31) John, P.; Davis, R.; Cady, N.; Czajka, J.; Batt, C. A.; Craighead, H. Anal. Chem. 1998, 70, 1108-1111. (32) Morhard, F.; Pipper, J.; Dahint, R.; Grunze, M. Sens. Actuators, B 2000, 70, 232-242. (33) Medina, M.; Houten, L.; Cooke, P.; Tu, S. Biotechnol. Tech. 1997, 11, 173176. (34) Fratamico, P.; Strobaugh, T.; Medina, M.; Gehring, A. Biotechnol. Tech. 1998, 12, 571-576. (35) Perkins, E.; Squirrell, D. Biosens. Bioelectron. 2000, 14, 853-859. (36) Watts, H.; Lowe, C.; Pollard-Knight, D. Anal. Chem. 1994, 66, 2465-2470. (37) Ferreria, A. P.; Werneck, M. M.; Ribeiro, R. M. Biosens. Bioelectron. 2001, 16, 399-408. (38) Ferreira, A. P.; Werneck, M. M.; Ribeiro, R. M. Biotechnol. Tech. 1999, 13, 447-452. (39) DeMarco, D. R.; Saaski, E. W.; McCrae, D. A.; Lim, D. V. J. Food Protect. 1999, 62, 711-716. (40) DeMarco, D. R.; Lim, D. V. J. Rapid Methods Autom. Microbiol. 2001, 9, 241-257. (41) King, K. D.; Anderson, G. P.; Bullock, K. E.; Regina, M. J.; Saaski, E. W.; Ligler, F. S. Biosens. Bioelectron. 1999, 14, 163-170. (42) Anderson, G. P.; King, K. D.; Gaffney, K. L.; Johnson, L. H. Biosens. Bioelectron. 2000, 14, 771-777. (43) Rowe, C.; Tender, L.; Feldstein, M.; Golden, J.; Scruggs, S.; MacCraith, B.; Cras, J.; Ligler, F. Anal. Chem. 1999, 71, 3846-3852. (44) Sipe, D.; Schoonmaker, K.; Herron, J.; Mostert, M. Proc. SPIE 2000, 3913, 215-222.

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The sensitivity of SPR biosensing for bacteria is, however, less than satisfactory.35 Detection limits are no better than (5-7) × 107 cfu/mL for E. coli, and achieving even this level required the use of a secondary antibody to enhance the signals. Perkins and Squirrell35 employed an SPR sensor to detect bacterial spores by a more sensitive detection phenomenon using scattering in conjunction with refractive index (RI) detection, but the shortness of the evanescent field extension and the light propagation of SPR (2 µm using gold and 19 µm using silver at λ ) 630 nm) diminished its benefit in bacteria detection.45 Based on the extent of recent research work and development in optical waveguides for bacteria detection reported in the literature, it can be concluded that the most significant drawback in commonly used waveguides such as SPR and RM is the penetration depth of the evanescent field.35 For example, the size of a bacterial spore, ∼1 µm, places the bulk volume of the bound cell outside the region where the SPR and RM evanescent fields have a significant intensity. To increase the sensitivity, the penetration depth of the evanescent field should be increased to include all the bound bacterial volume on the sensor surface within the evanescent field. The second key factor is that generally RI detection is not an adequately sensitive technique for low concentrations of bacteria. To achieve useful limits of detection, RI should be used in conjunction with other techniques such as fluorescence and scattering. The third factor is the light propagation at the sensor surface should also increase to widen the detection region to avoid direct illumination of the particles by the light source. These limitations have been addressed in our previous paper,46 which described for the first time, to the best of our knowledge, extending the evanescent field and light propagation at the sensor surface by developing an optical metal clad leaky waveguide (MCLW) sensor for the detection of particles. In this paper, the integrated MCLW device with the grating coupler was tested for the detection of a nonpathogenic bacterial simulant BG spores for pathogenic bacteria such as B. anthracis as whole bacterium using RI, scattering and fluorescence detection. CHIP FABRICATION AND THE SENSING MECHANISM Whenever a light wave is prevented from propagating by an aperture or an interface, the conservation of momentum and energy of a wave cause the generation of an evanescent field. The evanescent wave propagates in parallel to the interface, and it decays exponentially from the interface, thus allowing interaction with the outside medium by evanescent field coupling when this field is exposed. This evanescent wave light illumination can produce information about the interaction at the sensor surface.47 It can be used in detection of refractive index changes or to excite scattering or fluorescence. The principle of this technique is that any material of unmatched refractive index located near the interface, i.e., in the evanescent field, will cause a change in the reflectivity angle which in turn is related to changes in the refractive index at the sensor surface, or it will fluoresce or scatter light at an intensity proportional to the intensity of the evanescent wave. To enhance the sensitivity, the extension of the evanescent (45) Homola, J.; Yee, S. S.; Gauglitz, G. Sens. Actuators, B 1999, 54, 3-15. (46) Zourob, M.; Mohr, S.; Fielden, P. R.; Goddard, N. J. Sens. Actuators, B 2003, 90, 296-307. (47) Rohrbach, A. Biophys. J. 2000, 78, 2641-2654.

Figure 1. Instrumental setup: (a) LED or laser; (b) collimating lens and polarizer; (c) cylindrical lens; (d) grating coupler; (e) linear CCD; (f) chip and flow cell; (g) imaging optics; (h) CCD chip.

field should be increased to include the entire particle within the evanescent field of the sensor. This has been achieved by developing a metal clad leaky waveguide. The description of the MCLW sensor structure is given elsewhere.46,48 It can be summarized as follows. The structure of the metal clad leaky waveguide sensor consists of a 1-mm-thick BK7 glass substrate (or normal microscope glass slides, n ) 1.51) coated with an 8.5nm-thick titanium layer (n ) 2.11-2.88i at 2 eV/620 nm) followed by a 300-nm-thick layer of vacuum-deposited silica (n ) 1.47), giving a penetration depth of ∼1 µm into the water layer and a propagation length along the waveguide for a few millimeters. The mode profile of the MCLW is shown in Figure S-1 in the Supporting Information. Small changes in key parameters such as the desired thickness of the silica waveguiding layer will vary the characteristics of the modal coupling. For example, a reduction in the silica layer thickness to below 300 nm prevents any guided modes. Similarly, increasing the thickness of the silica layer beyond 500 nm significantly reduces the overlap of the leaky mode into the aqueous sample layer and results in supporting both TE0 and TE1 modes within the system. Therefore, the thickness of the silica waveguiding layer was chosen to support a single TE0 moded leaky waveguide operating near cutoff. There are a number of advantages associated with the insertion of a thin metal layer between the substrate and waveguide. First, the metal layer increases the penetration depth of the evanescent field, increasing sensitivity to particle detection on the surface. Second, the increased reflectivity of the waveguide/metal interfaces decreases the leakage rate of the leaky mode, increasing the propagation distance. Third, off resonance (i.e., when the waveguide mode is not excited), almost all of the incident optical energy is deposited

in the metal in the form of heat. Therefore, at resonance, there is a sharp peak in the reflectivity of the waveguide. This peak shows up as illuminated line on the charge-coupled device (CCD) detector without using a polarizer; hence, this simplifies the instrument.48 Finally, the metal layer reduces background intensity by absorbing off-resonance incident light and any scattering or fluorescence generated in the substrate.

(48) Zourob, M.; Mohr, S.; Fielden, P. R.; Goddard, N. J. Sens. Actuators, B 2003, 94, 304-312.

(49) Hulme, J.; Mohr, S.; Goddard, N. J.; Fielden, P. R. Lab Chip 2002, 2 (4), 203-206.

EXPERIMENTAL SECTION Instrumentation. A schematic of the optical arrangement used is shown in Figure 1. For refractive index detection, the same optical arrangement that has been detailed elsewhere48 has been used. The LED used in the input side had a peak wavelength of 620 nm (RS Components, Corby, U.K.). The LED was polished flat and mounted in a (25-mm-diameter) tube that was placed at the focal plane length of the cylindrical lens. The mounting tube incorporated the collimating lens with a 40-mm focal length and cylindrical lens with a 75-mm focal length to focus the light into a line. An interference filter 620 ( 10 nm wavelength, 25 mm in diameter, and a 25-mm diameter polarizer was used. All optical components were obtained from Comar Instruments (Cambridge, U.K.) unless otherwise stated. A holographic patterned epoxy optical grating film (25 400 lines.in., Edmund Optics, York, U.K.) was fabricated onto the back of the 2.5 cm long × 2.5 cm wide MCLW chip, to couple the light, as published before.49 On the output side, a 5000 pixel linear CCD (Sony ILX506A, pixel pitch 7 µm) positioned at a distance of 17 cm away from the axis of rotation of the grating was used to monitor the resonance angle, data collection being via a 12-bit analog-to-digital converter with a high-speed parallel link to a computer with software written inhouse.

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For the production of scattered light and fluorescence excitation from bacterial spores, a 3-mW semiconductor laser (RS Components) with a 635-nm wavelength was used as a light source. A high-resolution digital camera (Pulnix TM-1001, Sunnyvale, CA) set to a 1-ms exposure time was used to observe the scattering and fluorescence. The camera consisted of a highresolution 1-in. monochrome progressive scanning 1024 (H) × 1024 (V) interline transfer CCD imager. The imaging lens was constructed in-house and consisted of two achromatic lenses with dimensions as follows: 40-mm diameter and 80-mm focal length (Edmund Optics Ltd.). With an f/number of ∼0.8, the magnification was 2:1, which covers an area of 4 × 4 mm on the sensor chip. When using the TM 1001 camera, the resolution is ∼4 µm/ pixel. For detection of scattering from the BG bacterial spores, a 635 ( 10 nm interference filter was placed in front of the CCD camera to suppress unwanted laser emission of different wavelengths. For fluorescence detection from BG bacterial spores, a 650 ( 10 nm interference filter was placed in front of the camera to allow the transmission of the emitted fluorescence light to the CCD camera while blocking scattered excitation light. An injectionmolded poly(methyl methacrylate) flow cell with inlet and outlet ports was used as described before.48 A peristaltic pump (Minipuls3, MP4, Gilson Inc.) was used to pump solutions at 50 µL min-1 through the flow cell. Materials. Glutaraldehyde, bovine serum albumin (BSA), ethanolamine hydrochloride, N-ethyl-N′-(dimethylaminopropyl)carbodiimide (EDC), succinic anhydride, bromoacetic acid, dextran, N,N-dimethylformamide, sodium phosphate buffer (pH 7.4), N-2-hydroxyethylpiperazine-N′-2-ethanesulfonic acid (HEBES buffered saline), guanidine, N-hydroxysuccinimide (NHS), and fetal calf serum (FCS) were purchased from Sigma (Gillingham, U.K.). 3-Aminopropyl)triethoxysilane (APTS), N-(2-aminoethyl)-(3-aminopropyl)trimethoxysilane (EDA), trimethoxysilylpropyldiethylenetriamine (DETA), dimethyl sulfoxide, ethanol (95%), sodium hydroxide, HCl, and 11-mercaptoundecanoic acid were purchased from Aldrich (Gillingham, U.K.). Anti-BG antibodies and BG spores were obtained from Dstl (Porton Down, U.K.). Biotin-LCNHS ester was purchased from Pierce Biotechnology, Inc. Rockford, IL). Cyanine (Cy5) labeling kit (FluoroLink Cy5 Reactive Dye 5-pack) was purchased from Amersham Life Sciences (Little Chalfont, Bucks, U.K.). NeutrAvidin and 0.5-mL capacity SlideA-Lyzer dialysis cassettes (product 66383) were purchased from Pierce (Cheshire, U.K.). The MCLW chip was fabricated by IMEC (Louvain, Belgium). Normal microscope glass slides (Dow Corning supplied by BDH, Poole, U.K.) were coated with an 8.5-nm-thick layer of titanium followed by a ∼300-nm-thick layer of silica. SPR chips were produced in the Electrical Engineering Department at UMIST using normal microscope glass slides from BDH and were coated with 4 nm of chromium and 48 nm of gold. Preparation and Silanization of Substrates. The 2.5-cm2 MCLW chips were cleaned with aqueous ethanol before use, thoroughly rinsed using deionized water, and subsequently dried at room temperature. Silanization was carried out by immersing the chips in a solution of 5% APTS in ethanol or EDA or in DETA in 1 mM acetic acid in deionized water for 30 min at room temperature. The chips were then thoroughly rinsed three times with ethanol and deionized water, dried under nitrogen, and then 236

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placed in an oven for 1 h at 110 °C. Optimum results were obtained when silanized chips were modified promptly with the chemistry supports. Modification of Silanized Chips. For modification of the amino surface by succinic anhydride, the protocol described by Janolino et al. was used.50 Succinic anhydride (0.5 g) was dissolved in 10 mL of phosphate buffer, while the pH was maintained at 6-7 by adding 3 M sodium hydroxide solution. The chips were immersed in this solution. After 1 h, the chips were rinsed with phosphate buffer and dried at room temperature. Standard EDC/ NHS coupling chemistry was used to attach the antibodies to the carboxyl groups on the succinic anhydride treated surface.51 For the glutaraldehyde monomer activation, the silanized chip was immersed in a 5% v/v glutaraldehyde solution for 2 h. The chip was subsequently thoroughly rinsed with deionized water. To activate the chip with polymerized glutaraldehyde, the glutaraldehyde was polymerized by taking 10 mL of 5% v/v glutaraldehyde and adding 1 mL of 0.1 M NaOH. The solution was left for 30 min to polymerize before the mixture was neutralized with 1 mL of 0.1 M HCl. The presence of polymerized glutaraldehyde was monitored by noting the increase of the absorption peak at 234 nm, whereas the peak due to the monomer at 280 nm decreased. The chip was then immersed in the polymerized glutaraldehyde for 2 h followed by thorough rinsing with deionized water. Then the anti-BG antibodies were coupled to the activated glutaraldehyde chip surface by immersing the MCLW chips for 4 h in 10 mM phosphate buffer (pH 7.4) containing 100 µg mL-1 anti-BG antibody.52 For dextran coating, the silanized chips were activated by immersion in 5% glutaraldehyde in water for 30 min. Following thorough rinsing with water, the chip was immersed overnight in a 20% w/w solution of dextran in 0.1 M NaOH. Excess dextran was rinsed off after 14 h, and the chip was treated with 1 M bromoacetic acid in 2 M NaOH for 6 h at room temperature. After a thorough rinse with deionized water, the dried chips were ready for antibody immobilization using EDC/NHS coupling chemistry.52 For the biotin-labeled antibody and NeutrAvidin (Pierce), the silanized chips were incubated in 10 mL of 50 mM bicarbonate buffer (pH 8.5) mixed with 100 µL of N,N-dimethylformamide containing 10 mg of biotin-LC-NHS ester for 2 h at 4 °C, to form the covalent bond between the biotin-LC-NHS and the surface amino group. The chip then washed with deionized water, incubated in 1.5 µM NeutrAvidin in 10 mM sodium phosphate/ 10 mM NaCl, at pH 7.4 for 2 h at room temperature, and then washed in sodium phosphate buffer (pH 7.4). Biotin-conjugated anti-BG antibody (100 µg/mL in 10 mM phosphate buffer/10 mM NaCl/0.05% Tween 20 solution) was incubated over the NeutrAvidin-coated chip for 4 h at 4 °C. This was followed by a thorough rinsing of the chip with 10 mM phosphate buffer/10 mM NaCl/ 0.05% Tween 20/1 mg/mL BSA solution for 30 min before a final rinse with phosphate buffer.42,53 The SPR chip was treated by soaking in 11-mercaptoundecanoic acid overnight at 40 °C and subsequently washed with ethanol followed by deionized water. The chip was then functionalized using EDC/NHS coupling chemistry to attach the antibod(50) Janolino, V. G.; Swaisgood, H. E. Biotechnol. Bioeng. 1982, 24, 1069-1080. (51) Buckle, P.; Davies, R.; Kinning, T.; Yeung, D.; Edwards, P.; Pollard-Knight, D.; Lowe, C. R. Biosens. Bioelectron. 1993, 8, 355-363. (52) http://www.affinity-sensors.com, 10-02-2003.

ies to the carboxyl groups on the 11-mercaptoundecanoic acidtreated surface. The antibodies were used at the same concentration as above: 100 µg mL-1 in 10 mM sodium acetate buffer at pH 4.5 for 4 h. Unreacted surface groups were blocked by 1 M ethanolamine hydrochloride (pH 8.5) followed by immersion of the chip in 5% v/v FCS.35 Fluorescent labeling tracer antibodies were labeled with Cy5 bisfunctional reactive dye (λex ) 635 nm, λem ) 650 nm, according to the method recommended by the manufacturer. First the antibodies were diluted to 1 mg/mL in 0.1 M carbonate/ bicarbonate buffer (pH 9.3) A 1-mL aliquot of the antibody solution was added to the dye vial, which was capped and shaken to mix thoroughly. The vial of dye was then incubated, in the dark, at 25 °C for 30 min, with additional mixing every 10 min. The reaction mixture was dialyzed overnight into 0.1 M carbonate/bicarbonate buffer with two changes to separate free dye. Molar ratios of dye to protein ranged from 1.5 to 2.5.42,53 Binding and Regeneration. After selecting the appropriate immobilization matrix, the optimal binding pH, antibody concentration, binding periods, and flow rate were determined. The optimum conditions determined were then used for further experiments. Following immobilization of the antibody onto the sensor chip surface, 1 × 106 BG spores/mL suspensions were used (after 3× centrifugal washing in 10 mM HEBES buffered saline to show the effects of the pH on the binding of the spores. The bacterial suspensions (in HBS, pH 4-7) were then pumped across the surface of the chip at a flow rate of 100 µL/min for 1 h to allow the bacteria to interact with the immobilized antibodies. The surface was subsequently washed with HBS to remove any unbound material. After selecting the optimum pH (5.0), other binding conditions were determined by observing the combined interaction after a binding time of 1 h. Flow rates from 30 to 200 µL/min were used. The antibody-bound sensor surface was regenerated using 6 M guanidine hydrochloride (pH 1) for 1 min. A control surface for the bacteria was generated by covalently immobilizing BSA or FCS in place of the anti-BG antibodies and using BG as an antigen. A second control surface was generated by covalently immobilizing anti-BG antibody and using E. coli as a control antigen. To determine the limit of detection, 10-fold serial dilutions (1 × 101-1 × 1010 spores/mL) of BG spores prepared in HBS were irrigated over the immobilized antibody surface. As both specific and control surfaces were exposed to the bacteria or buffer, images were recorded using the CCD camera and simultaneously collected on a personal computer for 1 h. Quantification of the signal intensity from the light scattering or fluorescence for a particular particle was calculated by summing all the pixels belonging to that particle whose value exceeded a preset threshold value. RESULTS AND DISCUSSION Detection of Bacterial Spores. Bacterial spores were detected using RI in conjunction with detection of the bacteria as particulate entities using sensitive scattering and fluorescence image processing. Dark-field scattering or fluorescence images were captured for the bacterial spores and then analyzed to count the number of particles present on the sensor surface using (53) Sapsford, K. E.; Charles, P. T.; J. Patterson, C. H.; Ligler, F. S. Anal. Chem. 2002, 74, 1061-1068.

Figure 2. RI sensorgram for 1 × 105, 3 × 105, and 7 × 105 spores/ mL bacteria using MCLW sensor. (Arrows indicate the times at which various solutions were added).

software written in-house using LabVIEW v 6.1 (National Instruments, Austin, TX). The system provided real-time, direct-view, high-resolution optical images for the detection of bacteria in natural aqueous environments with a significant reduction in analysis complexity by eliminating sample preparation stages such as: cell lysis, protein separation, freezing, dehydration, staining, shadowing, marking, or other such manipulations. Refractive Index Detection. RI is a favorable technique for biomolecular interactions, allowing real-time analysis of biospecific interactions without the need for biomolecular labeling techniques. Here, two different approaches were employed for bacteria detection using RI detection with the MCLW chip. The first approach used the anti-BG antibody immobilized via an EDC/ NHS-modified dextran surface. Unreacted groups were then blocked, and the sensor was exposed to different concentrations of BG spores (1 × 103-1 × 108 spores/mL). The total change in response position, after allowing for bulk refractive index shifts by subtraction of the small signal change produced by the buffer wash, can be related to the cell concentration applied. Figure 2 shows the MCLW sensor response for different concentrations (1 × 105, 3 × 105, and 7 × 105 spores/mL) and the regeneration step. The limit of detection was estimated to be ∼8 × 104 cells/ mL when 5 mL of BG spores suspension was pumped through for 20 min. This value has been taken as the concentration corresponding to a response of the intercept plus three times the standard deviation of this value. The second approach was the immobilization of 1 × 105 spores/mL onto the EDC/NHS-activated dextran prior to the binding of a specific antibody. It was noted that the pixel response resulting from the bacteria immobilization is similar to the pixel response observed for binding of the antibody. Suspensions of different bacterial spores (1 × 105, 3 × 105, and 7 × 105 spores/ mL) were immobilized onto the EDC/NHS-modified dextran surface followed by 100 µg/mL anti-BG antibody. It was noted that the increase in the RI from antibody addition was related to the concentration of the bacterial spores immobilized onto the sensor surface. This has been confirmed by fluorescence emission from the binding of Cy5-labeled anti-BG to the immobilized BG. This indicates that there is proportionality between the immobilized bacteria and the captured antibodies. It is clear from Analytical Chemistry, Vol. 77, No. 1, January 1, 2005

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the results that increasing the evanescent field extension to place the whole volume of the cells within the field improves the detection limit for bacteria monitoring. Figure S-2 (Supporting Information) shows a linear relationship between the refractive index that is represented by a shift in pixels associated with BG binding and the mean area covered by the bacterial spores in captured images using a CCD camera. The area used is the average area resulting from three analyses calculated as the ratio of area covered by bacterial spores to the whole area of the image, using a program written using LabVIEW. It can be calculated that the shift in pixel position due to a single BG binding event is 5.10 × 10-5 of a pixel, assuming a typical BG spore is a sphere 1 µm in diameter and the flow cell area can bind a maximum of 1.0 × 106 cells/mm2. From these data it is possible to estimate the resonance shift due to one BG cell binding as 5.01 × 10-5 pixels, which is similar to that calculated previously.54 Scattering and Fluorescence Detection. Evanescent wave light illumination is a favorable method of generating scattering, fluorescence excitation, and detection of objects near the surface of the sensor. The technique gives a dark background and so provides a significant improvement in spatial resolution for objects near surface structures, in comparison with other microscopy techniques.48 Scattering and fluorescence are very sensitive techniques, which are capable of detecting low concentrations of bacteria where refractive index sensors give a negligible response. This combination affords a higher sensitivity than might be achieved using conventional SPR.35 In initial experiments, anti-BG antibody was immobilized on the sensor surface and activated with glutaraldehyde monomer. BG spores were then subsequently let flow across the surface of the sensor chip. BG spores flowing across the surface of the chip appeared either as a diffuse area of moving light or as brighter, smaller points, moving more slowly over the area of light propagation. In the latter case, the BG spores appeared to be moving close to the surface and could occasionally be seen to come to an instantaneous stop, presumably having been captured by the immobilized antibody. No such attachment was observed with BG spores at control surfaces coated with BSA or FCS or with an E. coli strain on anti-BG antibody coated surfaces. Upon stopping the flow, the diffusely emitting particles appeared to settle onto the MCLW chip where they became brighter and more sharply defined. Capture at the surface was indicated by cessation of movement and the bacteria remaining in place when the flow was restarted. Figure 3a shows an intensity plot for the sensor surface before capturing bacterial cells (background). Panels b and c of Figure 3 show typical images for scattered and fluorescence light emission from captured BG bacterial spores. The sharp spikes represent individual particles on the surface. The plots were analyzed to count the number of particles present on the sensor surface. Generally, unlike previously reported systems,35 the background signal from the MCLW sensors is low enough to eliminate the need for a stored reference image that has to be subtracted from the image containing the particle signals. Figure 4 shows the increase in the number of bound bacterial spores with variation in the length of BG exposure time to the (54) Wyatt, P. Method Microbiol. 1973, 8, 183-263.

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Figure 3. Intensity plot of (a) background, (b) scattering of BG spores, and (c) fluorescence of Cy5-labeled anti-BG spores using the MCLW.

anti-BG antibody-coated surface. The figure was generated after feeding the captured image from the camera to MATLAB (v 6, MathWork, Inc., Natick, MA) to convert the image from a 2-D matrix into a 1-D histogram (x, y). Hence x represents the pixel intensity and y represents the frequency (number of particles having the same intensity). The specificity of the anti-BG immobilized surface to BG was tested and the results obtained are shown in Figure 5. The experiment demonstrated the effect of the presence and absence of BG spores on the scattered light intensity on a BSA-coated control surface. No significant surface capturing of the BG spores was found to occur, and no significant change in the intensity of

Figure 4. Scattering intensity as the number of bound bacterial spores varying with the length of exposure time of BG to the antibodycoated MCLW chip (background, HBS buffer only), capturing after exposure to 106 BG spores/mL for 15 min, 30 min, and 1 h).

Figure 6. Specificity test results for the anti-BG antibody-coated MCLW. Other bacteria are not bound nonspecifically. (a) Scattering intensity before exposure to spores; (b) scattering intensity after 60min exposure to 106 cells/mL E. coli; (c) scattering intensity after 60 min following exposure to 106 BG spores/mL. Table 1. Comparison Scattering and Fluorescence Intensity of BG Spores Using MCLW and SPR intensity (au)

Figure 5. Response of a control surface without immobilized antibody, showing background signal and the scattering intensity after exposure for 1 h to 106 BG spores/mL in HBS buffer on the MCLW chip blocked with BSA; showing that BG spores do not bind nonspecifically.

the scattered light was noted, so a relatively low level of nonspecific binding (less than 5%) of BG was found to occur on the surface. A further test of specificity was made by exposing the anti-BG antibody surfaces to E. coli, with a final check using BG to confirm the activity of the antibody. This is demonstrated in Figure 6, which shows no significant capturing occurred in the case of the nonspecific antigen (E. coli) to the anti-BG antibody, but capturing did occur when BG was passed across the surface. Comparison with SPR. Figure S-3 (Supporting Information) shows the intensity of evanescent field of SPR and MCLW sensors (y-axis) as a function of distance from the sensor surface (x-axis).

chip

scattering ( RSD (%)

fluorescence ( RSD (%)

SPR MCLW

73 ( 9 231 ( 11.5

45 ( 2 125 ( 3

From Figure S-3 (Supporting Information) it can be seen that the SPR evanescent field has a short penetration depth at the surface, while the MCLW evanescent field extends greatly into the sensing area. This results in a greater volume of large particles, such as bacteria, being illuminated with the MCLW evanescent field light than is the case with the SPR. The scattering and fluorescence intensities generated by BG spores using the SPR and MCLW chips are shown in Table 1. It is clear that the scattering and fluorescence intensity from the MCLW is approximately three times stronger than that observed with SPR. Optimization of Bacteria Detection. (1) Effect of pH on Bacteria Capture. The effect of pH on bacteria-antibody binding on the sensor surface was studied. The highest capturing was observed at pH 5.0 and was found to be 10 times greater than that observed at pH 7.0. A similar optimum pH of 5.0 has been reported using E. coli O157:H7-antibody.33 It has also been reported that the antibody has a higher activity at pH 5 than at pH 7 although the precise reason for this is unclear.55 In our case, it may be due to the fact that a number of bacteria have isoelectric points at pH e 4.56 It was documented by Howe and Harding21 that using a running buffer with a pH close to or below the isoelectric point reduces the repulsion between the bacteria and the negatively charged SiO2 surface. (55) Lin, J. N.; Andrade, J. D.; Chang, I. N. J. Immunol. Methods 1989, 125, 67-77. (56) van der Wal, A.; Minor, M.; Norde, W.; Zehnder, A. J. B.; Lyklema, J. Langmuir 1997, 13, 165-171.

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Figure 7. Effect of type of silanization on bacteria coverage and scattering intensity.

(2) Immobilization and Binding. The ability to immobilize active biomolecules on the transducer surface is a key feature of most biosensors. Previous studies have shown that when antibodies are immobilized by adsorption directly onto the dielectric surfaces of RM devices, they exhibit low antigen-binding activity indicative of denaturation.51 It is preferable that the biomolecule should be covalently immobilized on the surface. Not only does this greatly enhance the proportion of active antibody and packing density, but it also improves the sensitivity of the sensor to the complementary antigen.51 Immobilization of the antibody onto the chip surface was carried out using different activation chemistries involving 100 µg/mL antibody at a flow rate of 100 µL/min for 4 h. If a higher binding surface is needed, the immobilization flow rate can be reduced or the immobilization time increased. However, the conditions described in this study were sufficient to meet the research objectives. A comparative study of different silanization techniques used for preparing the MCLW chips before antibody immobilization was performed to find the method giving the best bacteria coverage. This was determined by comparing the results of the scattering surface images observed after 1 h of capturing BG bacterial spores, using the three different approaches. The results are shown in Figure 7. It can be concluded from Figure 7 that as the chain of silane reagent is increased the bacteria capturing is increased. A comparative study of several surface antibody immobilization protocols for preparation of bacteria immunosensors was performed to find the method that gave the highest scattering signal intensity and the greatest capturing density of bacterial spores on the surface. This was determined by comparing the averaged results of the scattering intensity from the captured bacteria spores at the sensor surface after 1 h for five runs, for different approaches. The results are shown in Figure 8. It can be concluded from Figure 8 that the conventional technique using linkage of dextran chains to the aminosilanecovered chip via glutaraldehyde produced a coating containing a low density of captured spores. This technique gave the lowest scattering intensity, due to the hydrogel occupying the most sensitive region of the evanescent field, immediately adjacent to the interface. This results in a decrease in the scattering intensity by ∼48% due to the exponential decay of the evanescent field. Despite being well suited to interactions involving lower molecular 240 Analytical Chemistry, Vol. 77, No. 1, January 1, 2005

Figure 8. Normalized spore coverage after 1-h incubation and scattering intensity using different supports.

weight analytes, the potential steric problems created by binding cellular analytes to a carboxymethyl-dextran-immobilized anti-BG layer made hydrogel coatings unsuitable for bacterial detection. Consequently, anti-BG was immobilized using the other four activation techniques for sensor preparation. These were three methods for activation of the aminosilanized surfaces: activation by glutaraldehyde (both monomer and polymerized) and succinic anhydride activated by EDC/NHS. The fourth technique was biotin-labeled antibody/NeutrAvidin. The amount of BG captured by the anti-BG, and the resultant scattering intensity, varied greatly depending upon the surface modification and immobilization chemistry used. A 2-fold increase in scattering intensity of the BG spores was detected with the other coating techniques compared to the dextran coating technique (Figure 8). NeutrAvidin/ biotin, monomer glutaraldehyde, and succinic anhydride activated by EDC/NHS all gave a high capturing ability and strong scattering intensity. The polymerized glutaraldehyde gave the lowest level of capture. This may be as a result of antibody inactivation due to cross-linking of the active sites of the antibody. This effect could occur due to excess chemical bonds between the antibody and the available double bounds and aldehyde groups in the polymer chain.55 It was also noted that the scattering intensity from bacterial spores captured on the polymerized glutaraldehyde-activated surface was lower than that associated with the monomer-activated surface, due to the longer polymer chain. The dextran surfaces (mw 70 000) gave captured spores results approximately midway between the NeutrAvidin and the polyglutaraldeyhde coating surface. The same concentration (4 × 107 cells/mL) was applied to each surface so cell concentrations cannot account for the differences in binding and illumination intensity at the surface. Such differences could be due to the following: differences in the amount of immobilized anti-BG, the relative accessibility of the immobilized anti-BG to the antigen at the BG cell surface, or the distance of the captured cell from the sensor surface once it has bound to the immobilized anti-BG, i.e., the position of the cell within the evanescent field. Repulsion between the cells and the waveguide silica surface, or repulsion and attraction of charged groups on the immobilized moieties, e.g., dextran, etc., on the cell surfaces is possible, but as nonspecific controls performed

Figure 9. Detection of different concentrations of BG bacterial spores ([). Control surface without antibody (2) capturing after exposure for 1 h to 106 BG spores/mL in HBS buffer on (MCLW blocked with BSA), showing that BG spores do not bind nonspecifically (9). Specificity test showing that the anti-BG antibody coated MCLW did not bind another bacteria nonspecifically after 60-min exposure to 106 cell/mL E. coli and after 60 min following exposure to 106 spores/mL. Error bars represent (1 standard deviation, n ) 4.

with E. coli were low in comparison to the specific ones, nonspecific interactions appear to be limited. (3) Regeneration. Another desirable property of biosensors is the possibility of surface regeneration and the capability of performing multiple measurements with a single device, for reproducibility as well as for economic and other practical reasons. There was found to be an apparent reduction, 2-20%, in the capture of bacteria in the second analysis when measuring duplicate samples. The decrease in antibody capture can be attributed to saturation of the binding sites by the previous analysis. It would appear that the regeneration step did not completely remove captured bacteria or, alternatively, that the analytical process may have changed the antibody structure. The former is indicated by the optical analysis of the regenerated immobilized surface, which showed binding of Cy5-labeled antiBG, which indicated that some of the binding sites were still occupied by antigen. Results from replicate analyses showed regeneration using 6 M guanidine hydrochloride to be more successful than regeneration using phosphoric acid. Phosphoric acid regeneration was also found to gradually inhibit the bacteria-antibody binding, showing a 45% decrease observed from the first replicate analysis. This result has also been observed by other workers.33 In the present study, the sensor chip surface could be used for at least 10 different analyses before any significant differences in the binding of the bacteria to the immobilized antibody were observed. (4) Dose Response. Figure 9 shows the normalized area occupied by captured BG spores, as determined from surface imaging, plotted against bacterial spores concentration (1 × 103-1 × 1010 cells/mL). It is clear that there is a dependent relationship between the spores captured by the immobilized antibody and the concentration of the bacterial spores in the tested solution. The dose response was found to be linear over five decades of bacterial concentration (regression coefficient r > 0.988, the

probability of the linear fit p < 0.001). Figure 9 shows the results from four replicates of eight assays monitoring the binding of BG to an immobilized nonspecific protein BSA. Figure 9 also shows the results from E. coli binding with anti-BG. No significant captured spores were demonstrated. It can also be seen from Figure 9 that the standard curve appears to plateau at cell concentrations above 8 × 107 cells/mL, corresponding to 50-60% of the sensor surface being occupied by cells. This plateau may be due to the cells already at the surface sterically hindering other spores interacting with the immobilized antibody. Negative charges on the waveguide silica surface and cell-to-cell repulsive forces may prevent close packing of cells. Indeed, as cell surfaces consist of many different charged groups, short and longer range interactive forces between cells would be expected.36 The sensitivity of the sensor, as determined by the slope of the linear portion of the dose response curve, was found to be 0.202 ( 0.017 per decade of BG concentration. This value was based on results obtained from 40 experimental measurements. The theoretical detection limit estimated from extrapolation of the dose response curve is 450 ( 47 cells/mL, but this could not be achieved experimentally. The interaction of BG with the antibody is shown to be specific since the sensor did not respond to the E. coli bacteria (Figure 9). (5) Longevity. Longevity experiments show that the sensors preserve ∼60% of the original sensitivity (fraction of original activity versus log number of test runs) over 10 runs and g15% over 20 runs when the sensor was stored at 4 °C. The observed stability of the biosensor compares well with that previously reported for E. coli O157:H7 in a SPR sensor.33 CONCLUSION The results show that it has been possible to develop a biosensor for bacterial detection more sensitive than currently available sensors. A MCLW device has been fabricated to increase the overlap of the evanescent field extension from the sensor surface with the bacteria bound at the sensor surface. Increasing the overlap of the field extension at the sensor surface and increasing the propagation of the light along the direction of flow for a few millimeters gives a better detection limit and stronger illumination for scattering and fluorescence detection of bacteria. The sensor was then optimized by studying the key factors affecting sensor performance including changing the pH of the medium, type of antibody immobilization matrix, sensor surface regeneration approaches, and longevity of the sensor. ACKNOWLEDGMENT We gratefully acknowledge financial support for M.Z. from the Overseas Research Studentship (ORS), the British Government, and UMIST. The authors also gratefully acknowledge funding from the British Ministry of Defence for this work. The authors thank IMEC and the Electrical Engineering Department at UMIST for providing the waveguide coatings. The authors also acknowledge Dr. Sara Baldock and Dr. Andrew Knight for their revision and corrections to the manuscript. Analytical Chemistry, Vol. 77, No. 1, January 1, 2005

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SUPPORTING INFORMATION AVAILABLE Figure S-1, the MCLW mode profile; Figure S-2, the % area occupied by BG plotted against the pixel shift; Figure S-3, comparison of the calculated evanescent field intensity of SPR, RM, and MCLW (y-axis) as a function of distance from the sensor

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surface (x-axis). This material is available free of charge via the Internet at http://pubs.acs.org. Received for review March 9, 2004. Accepted August 23, 2004. AC049627G