2.8 Interference and Diffraction-Based Biosensors
2.8.4 Diffraction-Based Sensors
Biosensors have been developed based on optical diffraction from periodic grating structures. Cells are detected from the diffraction pattern caused by selective binding with antibody arrays immobilised on the substrate [10]. The adsorption of cells to the surface changes the intensity and distribution of the diffraction pattern produced by illuminating the slide with a laser beam. This variation was used to quantify the the amount of cells adsorbed from solution, to the substrate.
Figure 2.40: Diffraction patterns produced by laser illumination of biological gratings were imaged on a scattering screen using a CCD camera. Surface patterning (inset) was created using microcon- tact printing of Cell Tak followed by adsorption of anti-E.Coli as a recognition element for E.Coli
in solution (Figure based on [10]).
Microcontact printing was used to pattern a gold-coated silicon wafer with alkanethiol pat- terns. A PDMS stamp was inked with ‘Cell Tak’ (containing adhesive proteins from the sea shell Mytilus edulis) and then placed in contact with the gold coated silicon substrate.
The Cell Tak micropatterns were used to non-specifically bind receptor proteins or antibod- ies at well-defined adsorption sites to detect specific proteins in solution. Thus, a periodic biochemically active surface (a biological grating) was created (see Figure 2.40). The diffrac- tion pattern produced is sensitive to the period of the scattering elements, their shape and height. The micropatterns were illuminated with a 632.8 nmHeNe laser. Surface reactions (site specific cell adsorption) affect the shape of the scattering objects, affecting the diffrac- tion pattern in the Fraunhofer regime. The diffraction pattern image was captured on a CCD camera.
The substrate was patterned with anti-E.Coli IgG as a receptor recognition element for Es- cherichia Coli bacteria. Observing the formation and intensity of diffraction orders enabled study of the effect of solution concentration and time dependence on protein density. The anti-body was immobilised in an array of lines, with 10 µm line width and 20 µm grating period. As binding occurred a diffraction pattern built up with diffraction orders appearing after 10−20 minutes incubation time.
Diffraction has also been used previously for detecting bacterial Escherichia Coli by pro- ducing antibody diffraction gratings [158]. In this experiment oxidised silicon substrates were patterned by microcontact printing of antibody cells in a diffraction grating structure. The grating pattern used featured 10 µm wide lines of antibodies separated by 30 µm and diffraction was measured using a 632.8 nm HeNe laser. Gratings were characterised (prior to the attachment of target cells) for reflected (m = 0) and first order (m = 1) diffraction intensities. After attachment of target cells, diffraction intensities were measured again and the increase in intensities were correlated to the number of captured cells per unit area.
Studies have been conducted that use both the phenomena of diffraction and surface plasmon resonance for label-free detection of biomolecules. One such study coupled light in to the surface plasmon modes of a gold film via a prism and featured a grating structure on the reverse side of the gold [159, 160]. The grating structure was created by microcontact printing a biotin-thiol grating (100 µm period, 42 µm or 6 µm line width) directly on to an evaporated gold film (50 nm thick [161]) and a spacer thiol was used to passivate the unpatterned region. The contrast between the biotin thiol and spacer thiol was insufficient to generate detectable diffraction. Upon binding of anti-biotin antibody to the biotin lines sufficient optical contrast was generated to diffract the surface plasmon field and create a detectable diffraction pattern. The diffracted intensity was found to rise quadratically with the increase in antibody binding (at initial stages before saturation of binding sites was approached). A schematic for the experimental setup used is shown in Figure 2.41.
2.8 Interference and Diffraction-Based Biosensors 54
Figure 2.41: Schematic of the Surface Plasmon Diffraction Sensor based on a Kretschmann config- uration (Figure based on [11]).
This methodology was also applied to the label-free detection of DNA oligonucleotides [11]. The grating pattern was created in the same way as above, with thiol-biotin grating lines on a gold film and thiol spacer molecules covering regions in between. Streptavidin was then flowed over the grating and associated to the biotin regions. Probe DNA consisting of a biotinylated oligonucleotide sequence was then added to associate to the immobilised streptavidin. Upon addition of target DNA sequences, changes in diffraction intensities measured were correlated to hybridisation efficiencies. The lowest detectable coverage was calculated to be 1.1×1011 molecules percm2 with a maximum coverage of 1.2×1012. This
type of sensor is interesting, but to develop a more accurate sensor, the use of target DNA labels such as gold nanoparticles would introduce a much larger refractive index change. A further improvement would be to have the DNA probe layer closer to the gold surface where the evanescent field has a higher intensity and would offer lower detection limits.
Neuschafer et al developed ‘Evanescent Resonator’ chips and found that the fluorescence yield was enhanced up to 100 times over normal microarray chips [162]. By patterning the substrate for the hybridisation assay with sub-micron periodic structures in a thin dielectric layer, interference creates a resonant angle at which transmitted light destructively interferes, and virtually all light is reflected as it interferes constructively (see Figure 2.42). The concentration of light in specific directions means that the observed fluorescence signal is stronger at certain viewing orientations. These chips utilise diffraction effects resulting from the corrugated structure; a period of 360nm, step height of 20nm and width of 150 nm.
Figure 2.42: Evanescent resonator chips have been used to enhance fluorescent emission [162].
A fibre-optic biochemical sensor has been developed by Tang et al featuring a gold colloid modified fibre grating [163]. The resonant wavelength of the long period fibre grating created in the fibre core is sensitive to the refractive index of the cladding. In the grating region of the fibre, the cladding was removed and the core coated with≈8nmdiameter gold nanopar- ticles, to act as a sensing region. The bulk refractive index of the gold colloids is sensitive to biomolecular binding at the surface of the gold colloids and affects the transmission loss along the fibre. When there is a maximal refractive index contrast between the core and gold colloid cladding (such as in air), there is a minimum in the transmitted light intensity. As the refractive index surrounding the grating increases, the transmission loss reduces and the resonant wavelength at which the minimum transmission occurs is blue-shifted to shorter wavelengths. When the gold colloids were coated with dinitrophenyl (DNP), the detection limit of anti-DNP was found to be 9.5×10−10 M.
Sarov et al developed a microfluidic device using a reflecting diffraction grating to sense changes in the refractive index of the fluid flowing under the grating [164]. A gold reflecting diffraction grating (10µmperiod) was patterned on top of a microfluidic channel window of Si3N4. The transparent region in between the gold grating lines enables light to interact with
the fluid flowing through the channel, changing the amplitude and phase of light reflected from this region. The reflection properties from the gold regions remain constant and so a change in the optical properties of the fluid changes the optical contribution from the transparent region of the grating, changing the first order diffraction efficiencyIm=1/Im=0 of
the grating. A linear relationship was found between the change in diffraction efficiency and the change in refractive index of a sucrose solution as the sucrose content of the water was increased. The accuracy of diffraction efficiency change measurement of 5×10−3 correlated
to a change in refractive index of the fluid of ∆n= 1.7×10−3, a concentration of 1.2 % by
2.9 Conclusions 56