A. Field of the Invention
This invention relates generally to grating-based biochemical sensor devices and detection instruments for such devices. Grating-based sensors are typically used for optical detection of the adsorption of a biological material, such as DNA, protein, viruses or cells, small molecules, or chemicals, onto a surface of the device or within a volume of the device. The sensor of this invention has a grating structure that is constructed in a manner for use in two different applications: (a) label-free binding detection, and (b) fluorescence detection, for example wherein the sample is bound to a fluorophore or emits native fluorescence.
B. Description of Related Art
1. Label-Free Detection Sensors
Grating-based sensors represent a new class of optical devices that have been enabled by recent advances in semiconductor fabrication tools with the ability to accurately deposit and etch materials with precision less than 100 nm.
Several properties of photonic crystals make them ideal candidates for application as grating-type optical biosensors. First, the reflectance/transmittance behavior of a photonic crystal can be readily manipulated by the adsorption of biological material such as proteins, DNA, cells, virus particles, and bacteria on the crystal. Other types of biological entities which can be detected include small and smaller molecular weight molecules (i.e., substances of molecular weight<1000 Daltons (Da) and between 1000 Da to 10,000 Da), amino acids, nucleic acids, lipids, carbohydrates, nucleic acid polymers, viral particles, viral components and cellular components such as but not limited to vesicles, mitochondria, membranes, structural features, periplasm, or any extracts thereof. These types of materials have demonstrated the ability to alter the optical path length of light passing through them by virtue of their finite dielectric permittivity. Second, the reflected/transmitted spectra of photonic crystals can be extremely narrow, enabling high-resolution determination of shifts in their optical properties due to biochemical binding while using simple illumination and detection apparatus. Third, photonic crystal structures can be designed to highly localize electromagnetic field propagation, so that a single photonic crystal surface can be used to support, in parallel, the measurement of a large number of biochemical binding events without optical interference between neighboring regions within <3-5 microns. Finally, a wide range of materials and fabrication methods can be employed to build practical photonic crystal devices with high surface/volume ratios, and the capability for concentrating the electromagnetic field intensity in regions in contact with a biochemical test sample. The materials and fabrication methods can be selected to optimize high-volume manufacturing using plastic-based materials or high-sensitivity performance using semiconductor materials.
Representative examples of grating-type biosensors in the prior art are disclosed in Cunningham, B. T., P. Li, B. Lin, and J. Pepper, Colorimetric resonant reflection as a direct biochemical assay technique. Sensors and Actuators B, 2002. 81: p. 316-328; Cunningham, B. T., J. Qiu, P. Li, J. Pepper, and B. Hugh, A plastic colorimetric resonant optical biosensor for multiparallel detection of label-free biochemical interactions, Sensors and Actuators B, 2002. 85: p. 219-226; Haes, A. J. and R. P. V. Duyne, A Nanoscale Optical Biosensor: Sensitivity and Selectivity of an Approach Based on the Localized Surface Plasmon Resonance Spectroscopy of Triangular Silver Nanoparticles. Journal of the American Chemical Society, 2002. 124: p. 10596-10604.
The combined advantages of photonic crystal biosensors may not be exceeded by any other label-free biosensor technique. The development of highly sensitive, miniature, low cost, highly parallel biosensors and simple, miniature, and rugged readout instrumentation will enable biosensors to be applied in the fields of pharmaceutical discovery, diagnostic testing, environmental testing, and food safety in applications that have not been economically feasible in the past.
In order to adapt a photonic bandgap device to perform as a biosensor, some portion of the structure must be in contact with a test sample. Biomolecules, cells, proteins, or other substances are introduced to the portion of the photonic crystal and adsorbed where the locally confined electromagnetic field intensity is greatest. As a result, the resonant coupling of light into the crystal is modified, and the reflected/transmitted output (i.e., peak wavelength) is tuned, i.e., shifted. The amount of shift in the reflected output is related to the amount of substance present on the sensor. The sensors are used in conjunction with an illumination and detection instrument that directs light into the sensor and captures the reflected or transmitted light. The reflected or transmitted light is fed to a spectrometer that measures the shift in the peak wavelength.
The ability of photonic crystals to provide high quality factor (Q) resonant light coupling, high electromagnetic energy density, and tight optical confinement can also be exploited to produce highly sensitive biochemical sensors. Here, Q is a measure of the sharpness of the peak wavelength at the resonant frequency. Photonic crystal biosensors are designed to allow a test sample to penetrate the periodic lattice, and to tune the resonant optical coupling condition through modification of the surface dielectric constant of the crystal through the attachment of biomolecules or cells. Due to the high Q of the resonance, and the strong interaction of coupled electromagnetic fields with surface-bound materials, several of the highest sensitivity biosensor devices reported are derived from photonic crystals. See the Cunningham et al. papers cited previously. Such devices have demonstrated the capability for detecting molecules with molecular weights less than 200 Daltons (Da) with high signal-to-noise margins, and for detecting individual cells. Because resonantly-coupled light within a photonic crystal can be effectively spatially confined, a photonic crystal surface is capable of supporting large numbers of simultaneous biochemical assays in an array format, where neighboring regions within ˜10 μm of each other can be measured independently. See Li, P., B. Lin, J. Gerstenmaier, and B. T. Cunningham, A new method for label-free imaging of biomolecular interactions. Sensors and Actuators B, 2003.
There are many practical benefits for label-free biosensors based on photonic crystal structures. Direct detection of biochemical and cellular binding without the use of a fluorophore, radioligand or secondary reporter removes experimental uncertainty induced by the effect of the label on molecular conformation, blocking of active binding epitopes, steric hindrance, inaccessibility of the labeling site, or the inability to find an appropriate label that functions equivalently for all molecules in an experiment. Label-free detection methods greatly simplify the time and effort required for assay development, while removing experimental artifacts from quenching, shelf life, and background fluorescence. Compared to other label-free optical biosensors, photonic crystals are easily queried by simply illuminating at normal incidence with a broadband light source (such as a light bulb or LED) and measuring shifts in the reflected color. The simple excitation/readout scheme enables low cost, miniature, robust systems that are suitable for use in laboratory instruments as well as portable handheld systems for point-of-care medical diagnostics and environmental monitoring. Because the photonic crystal itself consumes no power, the devices are easily embedded within a variety of liquid or gas sampling systems, or deployed in the context of an optical network where a single illumination/detection base station can track the status of thousands of sensors within a building. While photonic crystal biosensors can be fabricated using a wide variety of materials and methods, high sensitivity structures have been demonstrated using plastic-based processes that can be performed on continuous sheets of film. Plastic-based designs and manufacturing methods will enable photonic crystal biosensors to be used in applications where low cost/assay is required, that have not been previously economically feasible for other optical biosensors.
The assignee of the present invention has developed a photonic crystal biosensor and associated detection instrument for label-free binding detection. The sensor and detection instrument are described in the patent literature; see U.S. patent application publications U.S. 2003/0027327; 2002/0127565, 2003/0059855 and 2003/0032039. Methods for detection of a shift in the resonant peak wavelength are taught in U.S. Patent application publication 2003/0077660. The biosensors described in these references include 1- and 2-dimensional periodic structured surfaces applied to a continuous sheet of plastic film or substrate. The crystal resonant wavelength is determined by measuring the peak reflectivity at normal incidence with a spectrometer to obtain a wavelength resolution of 0.5 picometer. The resulting mass detection sensitivity of <1 pg/mm2 (obtained without 3-dimensional hydrogel surface chemistry) has not been demonstrated by any other commercially available biosensor.
A fundamental advantage of the biosensor devices described in the above-referenced patent applications is the ability to mass-manufacture with plastic materials in continuous processes at a 1-2 feet/minute rate. Methods of mass production of the sensors are disclosed in U.S. Patent application publication 2003/0017581. As shown in FIG. 1, the periodic surface structure of a biosensor 10 is fabricated from a low refractive index material 12 that is overcoated with a thin film of higher refractive index material 14. The low refractive index material 12 is bonded to a base sheet of clear plastic material 16. The surface structure is replicated within a layer of cured epoxy 12 from a silicon-wafer “master” mold (i.e. a negative of the desired replicated structure) using a continuous-film process on a polyester substrate 16. The liquid epoxy 12 conforms to the shape of the master grating, and is subsequently cured by exposure to ultraviolet light. The cured epoxy 12 preferentially adheres to the sheet 16, and is peeled away from the silicon wafer. Sensor fabrication was completed by sputter deposition of 120 nm titanium oxide (TiO2) high index of refraction material 14 on the cured epoxy 12 grating surface. Following titanium oxide deposition, 3×5-inch microplate sections are cut from the sensor sheet, and attached to the bottoms of bottomless 96-well and 384-well microtiter plates with epoxy.
As shown in FIG. 2, the wells 20 defining the wells of the microtiter plate contain a liquid sample 22. The combination of the bottomless microplate and the biosensor structure 10 is collectively shown as biosensor apparatus 26. Using this approach, photonic crystal sensors are mass produced on a square-yardage basis at very low cost.
The detection instrument for the photonic crystal biosensor is simple, inexpensive, low power, and robust. A schematic diagram of the system is shown in FIG. 2. In order to detect the reflected resonance, a white light source illuminates a ˜1 mm diameter region of the sensor surface through a 100 micrometer diameter fiber optic 32 and a collimating lens 34 at nominally normal incidence through the bottom of the microplate. A detection fiber 36 is bundled with the illumination fiber 32 for gathering reflected light for analysis with a spectrometer 38. A series of 8 illumination/detection heads 40 are arranged in a linear fashion, so that reflection spectra are gathered from all 8 wells in a microplate column at once. See FIG. 3. The microplate+biosensor 10 sits upon a X-Y addressable motion stage (not shown in FIG. 2) so that each column of wells in the microplate can be addressed in sequence. The instrument measures all 96 wells in ˜15 seconds, limited by the rate of the motion stage. Further details on the construction of the system of FIGS. 2 and 3 are set forth in the published U.S. Patent Application 2003/0059855.
The descriptions and discussions below refer to the label-free technology described above as BIND technology. BIND is a trademark of the assignee SRU Biosystems, Inc.
2. Fluorescence Amplification Sensors
U.S. Pat. No. 6,707,561 describes a grating-based biosensing technology that is sometimes referred to in the art as Evanescent Resonance (ER) technology. This technology employs a sub-micron scale grating structure to amplify a luminescence signal (e.g., fluorescence, chemi-luminescence, electroluminescence, phosphorescence signal), following a binding event on the grating surface, where one of the bound molecules carries a fluorescent label. ER technology enhances the sensitivity of fluorophore based assays enabling binding detection at analyte concentrations significantly lower than non-amplified assays.
ER technology uses grating generated optical resonance to concentrate laser light on the grating surface where binding has taken place. In practice, a laser scanner sweeps the sensor at some angle of incidence (theta), typically from above the grating, while a detector detects fluoresced light (at longer optical wavelength) from the sensor surface. By design, ER grating optical properties result in nearly 100% reflection, also known as resonance, at a specific angle of incidence and laser wavelength (λ). Confinement of the laser light by and within the grating structure amplifies emission from fluorophores bound within range of the evanescent field (typically 1-2 um). Hence, at resonance, transmitted light intensity drops to near zero.
As noted above, the label-free biosensors described in the above-referenced patent applications employ a sub-micron scale grating structure but typically with a significantly different grating geometry and objective as compared to gratings intended for ER use. In practical use, label-free and ER technologies have different requirements for optical characteristics near resonance. The spectral width and location of the resonance phenomena describes the primary difference. Resonance width refers to the full width at half maximum, in wavelength measure, of a resonance feature plotted as reflectance (or transmittance) versus wavelength (also referred to as Q factor above). Resonance width can also refer to the width, in degrees, of a resonance feature plotted on a curve representing reflectance or transmittance as a function of theta, where theta is the angle of incident light.
Optimally, a label-free grating-based sensor produces as narrow a resonance peak as possible, to facilitate detection of small changes in peak position indicating low binding events. A label-free sensor also benefits from a high grating surface area in order to bind more material. In current practice, one achieves higher surface area by making the grating deeper (though other approaches exist). Current commercial embodiments of label-free sensors produce a resonance near 850 nm, thus BIND label-free detection instrumentation has been optimized to read this wavelength.
Conversely, practical ER grating sensor designs employ a relatively broad resonance to ensure that resonance occurs at the fixed wavelength laser light and often fixed angle of incidence in the presence of physical variables such as material accumulation on the grating or variation in sensor manufacture. Because field strength generally decreases with resonance width, practical ER sensor design calls for a balance in resonance width. By choosing an appropriate ER resonance width, one ensures consistent amplification across a range of assay, instrument and sensor variables while maintaining ER signal gain. A typical application uses a 633 nm wavelength to excite a popular fluorescent dye, known in the art as Cy5. Some ER scanning instrumentation permits adjustments to incident angle to “tune” the resonance towards maximum laser fluorophore coupling. This practice, however, may induce an unacceptable source of variation without proper controls.
Known ER designs also employ more shallow grating depths than optimal label-free designs. For example, the above-referenced '561 patent specifies the ratio of grating depth to “transparent layer” (i.e., high index coating layer) thickness of less than 1 and more preferably between 0.3 and 0.7. Optimal label-free designs employ gratings with a similarly defined ratio of greater than 1 and preferably greater than 1.5. Label-free designs typically define grating depth in terms of the grating line width or half period. For example, currently practiced commercial label-free sensors have a half period of 275 nm and a grating depth of approximately 275 nm, thus describing a 1:1 geometric ratio. This same sensor design employs a high index of refraction oxide coating on top of the grating with a thickness of approximately 90 nm. Thus, according to the definition in '561 patent, this sensor has a grating depth:oxide thickness ratio of approximately 3:1.
This disclosure reports grating-based sensor designs which are constructed in a manner such that it is optimized for both modes of detection (label-free and fluorescence amplification), in a single device. Such a grating dramatically increases the diversity of applications made possible by a single product.
All the previously cited art is fully incorporated by reference herein.
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