Abstract:
A biosensor embodying the invention includes a sensing volume having an array of pores sized for immobilizing a first biological entity tending to bind to a second biological entity in such a manner as to change an index of refraction of the sensing volume. The biosensor further includes a ring interferometer, one volumetric section of the ring interferometer being the sensing volume, a laser for supplying light to the ring interferometer, and a photodetector for receiving light from the interferometer.

Description:
ORIGIN OF THE INVENTION 
     The invention described herein was made in the performance of work under NASA contract, and is subject to the provisions of Public Law 96-517 (35 U.S.C. § 202) in which the Contractor has elected not to retain title. 
    
    
     BACKGROUND OF THE INVENTION 
     A biosensor can be used to detect the presence of a specific antigen. These typically involve a sample of antibodies specific to the antigen of interest. There is a large class of biosensors that use immobilized antibodies on a surface as the sensing agent. The antibodies on the surface are identical and bind to a specific antigen, so that the sensor is specific to that particular antigen. 
     The antibody-antigen binding event must be detected. The traditional detection scheme in antibody sensors is fluorescence. In a typical sandwich assay, the immobilized antibody binds to the antigen; the system is then exposed to a fluorophore conjugated to the antibody, which then binds to the antigen. This tags the bound system upon exposure to light of a suitable wavelength. 
     The binding event also can be detected by the change in refractive index of the surface that occurs whenever antigens become bound to the antibodies on the surface. Such a biosensor is disclosed in U.S. Pat. No. 5,663,790, in which the bound antibodies are on a surface overlying an optical ring resonator. The change in refractive index shifts the resonant wavelength of the optical ring resonator. By sweeping the light frequency while observing the light intensity in the ring, the shift in resonant wavelength is observed, indicating a shift in refractive index and the corresponding event of the binding of the antigens to the antibodies. 
     One disadvantage of such a sensor is that the light source must have a variable wavelength that can be swept across a range. A related disadvantage is that the binding event can only be inferred after the light source wavelength has been swept across the range and the optical ring resonator response compared across the range. Another disadvantage is that the coupling between the change in refractive index in the sample and the detected optical output is limited because the sample is adjacent to and not within the optical ring interferometer. Finally, there appears to be no way of enhancing sensitivity of the sensor. The problem is that a very dilute antigen sample may not contain a sufficient population of antigens to bind to more than a small fraction of the bound antibodies, so that the change in refractive index may be so slight that the sensor cannot detect it. 
     What is needed is an optical sensor having an enhanced sensitivity capable of detecting extremely small changes in refractive index so as to be capable of sensing and measuring extremely weak or dilute antigen samples. Moreover, what is needed is a sensor that does not require expensive optical features such as a variable wavelength light source, and which does not require sweeping the light source wavelength across a range to make a measurement. 
     SUMMARY OF THE INVENTION 
     A biosensor embodying the invention includes a sensing volume having an array of pores sized for immobilizing a first biological entity tending to bind to a second biological entity in such a manner as to change an index of refraction of the sensing volume. The biosensor further includes a ring interferometer, one volumetric section of the ring interferometer being the sensing volume, a laser for supplying light to the ring interferometer, and a photodetector for receiving light from the interferometer. 
     The array of pores can be nanometer-sized pores. The array of pores may be Sol-Gel. The first biological entity may be an antibody and the second biological entity may be an antigen that binds to the antibody. Alternatively, the first biological entity may be a printed polymer and the second biological entity may be an antigen that binds to the printed polymer. Or, the first biological entity may be a first DNA or RNA strand and the second biological entity may be a second DNA or RNA strand that is complementary to the first strand. 
    
    
     
       BRIEF DESCRIPTION OF THE DRAWINGS 
         FIG. 1  is a schematic diagram of a first embodiment of the invention. 
         FIG. 2  is a partial cut-away cross-sectional view qualitatively depicting a Sol-Gel structure constituting the sensing volume in the embodiment of  FIG. 1 . 
         FIG. 3  is a view corresponding to  FIG. 2  in which antibodies are held in the nanometer-size pores of the Sol-Gel structure. 
         FIG. 4  is a view corresponding to  FIG. 3  in which some of the antibodies have antigens bound to them, as symbolized by the shaded pores. 
         FIG. 5  is a graph illustrating the change in referactive index of the sensing volume in the embodiment of  FIG. 1  as a function of analyte density (density of antibody-antigen bound pairs in the Sol-Gel structure of  FIG. 4 ). 
         FIG. 6  is a graph illustrating the light output intensity of the ring interferometer of the embodiment of  FIG. 1  as a function of the phase shift through the sensing volume. 
         FIG. 7  is a graph illustrating the light output intensity of the ring interferometer as a function of analyte density. 
         FIG. 8  is a perspective view of a second embodiment of the invention. 
         FIG. 9  is a perspective view of a third embodiment of the invention. 
         FIG. 10  is a perspective view of a fourth embodiment of the invention. 
     
    
    
     DETAILED DESCRIPTION OF THE INVENTION 
     The problems described in the foregoing background discussion are solved by a biosensor in which the bound antibodies are immobilized in a nano-porous structure, such as Sol-Gel, that occupies a sensing volume within one arm or section of an optical ring interferometer. All the light traveling through that arm or section must pass through the bound antibody sensing volume, i.e., the nano-porous matrix, and thereby becomes a phase-shifted light beam due to the different index of refraction of the sensing volume. Another portion of the ring interferometer contains light that has not passed through the sensing volume, and interferes with the phase-shifted light beam. The degree of the interference is determined by the index of refraction of the sensing volume, which changes as antigens become bound to the immobilized antibodies in the sensing volume. The sensor is highly responsive in part because of the complete immersion of one portion of the ring interferometer in the Sol-Gel or nano-porous matrix. The sensor&#39;s sensitivity is greatly enhanced because the light makes multiple passes through the ring interferometer. The intensity of the interference pattern increases as the light circulates around the ring interferometer, thereby enhancing sensitivity of the sensor. 
     One advantage is that a variable wavelength light source is not required, and no frequency sweeping or comparison is required, so that the sensor is simple and its response is immediate. 
     Referring to  FIG. 1 , an optical ring interferometer  100  includes four mirrors  102 ,  104 ,  106 ,  108  establishing a closed path consisting of optical paths  110 ,  112 ,  114 ,  116  of the ring interferometer  100 . A laser  118  produces a light beam  120  that follows the optical paths  110 – 116  by reflection on the mirrors  102 – 108 . The light beam  120  enters the optical path  110  through the backside of the mirror  102 . The mirror  104  may be a beam splitter which transmits a portion of the light traveling along the optical path  110  to a detector  122  and reflects the remaining portion to the next optical path  112 . A section of the optical path  114  is immersed in a sensing volume  124  consisting of a 3-dimensional array or matrix of nano-meter sized pores that are sized to attract and immobilize antibodies of a selected type. All of the light in the optical path  114  passes through the sensing volume  124 . The array of nanometer-sized pores constituting the sensing volume  124  is best formed as a Sol-Gel structure of the type disclosed in Yamanaka et al.,  J. Sol - Gel Sci. Technol.  7, 117 (1996). 
       FIG. 2  is a qualitative illustration of a Sol-Gel structure  200  constituting the sensing volume  124 , before its pores are filled with any material. In  FIG. 2 , the structure consists of a 3-dimensional base  210  having many 3-dimensional pores  220  formed within it. Each pore  220  has an average pore size P. The sol-gel precursor is a liquid that can be poured into the structure, the evaporation of the liquid leaving the sol-gel structure. Reference is made to a periodical dedicated to this technology, the  Journal of Sol - gel Science Technology . The pore size P is chosen to correspond to the size of the antibody to be immobilized in each pore. This size ranges between a few nanometers and several hundred nanometers, depending upon the size of the particular antibody. The chosen antibody is one that binds to the antigen targeted by the sensor. In accordance with current practice, the base  210  can be formed of a material such as a silicon-based compound which is essentially inert with respect to biological materials such as antibodies and antigens, for example. The pore size P of the Sol-Gel is selected to attract and bind individual antibodies of a particular type within the pores, so that P is about the size of or somewhat larger than the selected antibody type, in accordance with conventional practice.  FIG. 3  illustrates the Sol-Gel structure of  FIG. 2  in which antibodies of a selected type have been immobilized in each pore  220 . The selected antibody type determines the type of antigen targeted by the sensor.  FIG. 4  illustrates the Sol-Gel structure of  FIG. 2  in which the Sol-Gel structure of  FIG. 3  has been immersed in or wetted with a fluid or analyte possibly containing the targeted antigen. 
     Depending upon the concentration of the targeted antigen in the analyte or fluid being tested, a certain fraction of the antibodies in the pores  220  receive and bind to antigens, as indicated by the dotted pores  220   a . The remaining pores  220   b  contain immobilized antibodies but without antigens bound to them. Thus,  FIG. 4  illustrates a case in which the Sol-Gel structure has been introduced to a relatively weak concentration of antigens so that only a portion of the immobilized antibodies in the Sol-Gel structure  200  receive and bind to antigens. The proportion of antigen-antibody bound pairs in the Sol-Gel structure  200  (i.e., their volume density) is a function of the concentration of antigens in the liquid or analyte introduced to the Sol-Gel structure  200 , and other factors. An analyte having a very high concentration of the targeted antigen may saturate the Sol-Gel structure  200 , so that all or nearly all of the immobilized antibodies may become bound to antigens. An analyte having a very weak concentration of the targeted antigen may cause antibody-antigen binding by only a small proportion of the immobilized antibodies. 
     The analyte may be introduced to the Sol-Gel structure  200  by flowing a liquid form of the analyte over the Sol-Gel structure  200 . The analyte may either be the liquid itself or may be contained in the liquid. Unbound analyte is washed out of the Sol-Gel structure  200 , leaving only the immobilized antibodies in the Sol-Gel structure  200  and any antigens from the analyte that became bound to the immobilized antibodies. 
     Binding of antigens to the immobilized antibodies in the Sol-Gel structure  200  changes the refractive index of the Sol-Gel structure  200  that constitutes the sensing volume  124 . The density of antigen-antibody bound pairs in the Sol-Gel structure  200  determines the resulting change in refractive index. The magnitude of the change in refractive index resulting from the introduction of the analyte to the Sol-Gel structure is a measure of the concentration of the antigens in the analyte. The sensor therefore provides two pieces of information: (a) a detectable change in refractive index indicates the presence of the targeted antigen type in the analyte, while (b) the magnitude of the change in refractive index is indicative of the concentration of the targeted antigen in the analyte. 
     Referring to  FIG. 1 , an “unshifted” light beam  120  from the laser  118  travels through the mirror  102  to the optical path  110  and is transmitted by the beam splitter mirror  104  to the detector  122 . A phase-shifted light beam consists of the portion of the light from the optical path  110  that is reflected by the beam splitter mirror  104  to the path  112 , so that it travels through the optical paths  114 ,  116  and  110  (in that order). This light beam is phase shifted by passing through the sensing volume  124  occupying one section of the optical path  114 . Thus, the light path  110  has a light beam traveling toward the beam splitter mirror  104  consisting of both the unshifted light beam and the phase shifted light beam. A portion of all the light in the optical path  110  is transmitted by the beam splitter mirror  104  to the detector  122 , so that both the phase shifted and unshifted light beams impinge on the detector  122 . The phase shift imposed by the sensing volume  124  causes interference between the shifted and unshifted light beams, which affects the intensity of the light sensed by the detector  122 . The length of the optical path  110  as well as the total length of the interferometer ring  110 ,  112 ,  114 ,  116  are preferably integral numbers of wavelengths of the laser  118 . 
     Referring to graph of  FIG. 5 , the horizontal axis corresponds to the concentration, ρ, of the targeted antigen in the analyte, while the vertical axis corresponds to the change in refractive index, Δn (left hand vertical axis) and, equivalently, to the phase shift, ΔΦ (right hand vertical axis) of the light traveling in the optical path  114  through the sensing volume  124  containing the Sol-Gel structure  200 .  FIG. 5  qualitatively depicts the behavior in which stronger concentrations of the antigen in the analyte produce proportionately larger changes in refractive index and, hence, larger phase shifts. Such phase shifts change the interference between the shifted and unshifted light beams at the photodetector  122 , producing corresponding changes in the light intensity at the photodetector  122 . Referring to the graph of  FIG. 6 , the horizontal axis corresponds to the phase shift ΔΦ, while the vertical axis corresponds to the light intensity, I, sensed by the photodetector  122 . As the phase shift between the “shifted” and “unshifted” light beams approaches 180 degrees, the light intensity at the photodetector approaches a minimum, as indicated by the graph of  FIG. 6 . The sensor may be constructed so that the phase shift is near zero in the absence of bound antibody-antigen pairs in the sensing volume  124  and approaches 180 degrees as the concentration of bound antibody-antigen pairs approaches saturation (100%). Two methods exist for modifying the index of refraction of the ring resonator in order to set its operating point. One method uses a polymer section of the waveguide which can have its index of refraction “written” by UV light, allowing device tuning during manufacture. Another method uses a small resistive heater on a portion of the waveguide to thermally tune the ring resonance point. In such implementations, the light intensity at the detector  122  decreases as the bound antibody-antigen concentration increases. This effect may be quantified by obtaining different samples of the photodetector output obtained using different concentrations of targeted antigen in different analytes. The results are illustrated in  FIG. 7 , in which the targeted antigen concentration ρ in the analyte is represented by the horizontal axis while the light intensity, I, corresponds to the vertical axis. In general, an increase in targeted antigen concentration in the analyte (over successive samples) causes a decrease in light intensity at the photodetector  122 . However, other arrangements of the sensor may be made by the skilled worker that may reverse the behavior illustrated in  FIG. 7 . In either case, the change in measured intensity I at the photodetector may be employed as a measure of targeted antigen concentration p in the analyte. 
     Synchronous detection may be employed in the sensor of  FIG. 1 . For this purpose, a synchronous detection control circuit  140  may control the pulse width and repetition rate of the laser  118  while enabling the detector  122  in synchronism with the pulses of the laser  118 . Alternatively, rather than pulsing the laser  118 , the synchronous detection control circuit  140  can dither one of the mirrors (e.g., the mirror  104 ) to pulse the light propagation in the ring interferometer. The synchronous detection control circuit  140  would enable the detector  122  in synchronism with the dithering of the mirror so that the detector  122  is enabled with each pulse of light. 
       FIG. 8  illustrates an integrated implementation of the sensor of  FIG. 1  implemented on a substrate  800 . The substrate is formed of an optically transparent material such as glass or plastic or a semiconductor material or a ceramic. The optical paths  110 ,  112 ,  114 ,  116  of  FIG. 1  are implemented in  FIG. 8  as a closed waveguide  801  having connected sections  810 ,  812 ,  814 ,  816 . The waveguide  801  is formed as a mesa or rib structure by etching the surface of the substrate  800  or by other methods well-known in the art. The cross-sectional height and width of the waveguide  801  may be selected by the skilled worker to support a single optical mode at the frequency of the laser  818 . Mirrors  802 ,  804 ,  806 ,  808  are placed at each corner between successive waveguide sections  810 ,  812 ,  814 ,  816 . A laser  818  is formed on the substrate  800  and feeds light into the waveguide  801  through the back of the mirror  802 . A photodetector  822  is formed on the substrate  800  and receives light transmitted through the mirror  804 . The mirror  804  performs the same function as the mirror  104  of  FIG. 1 , in that it transmits a portion of the light in the waveguide section  810  to the photodetector  822  while reflecting the remaining portion to the next waveguide section  812 . A sensing volume  824  occupies a section of the waveguide  814  and consists of Sol-Gel structure  200  of  FIGS. 2 ,  3  or  4 . One advantage of the integrated sensor of  FIG. 8  is that it can be highly compact, especially if the laser  818  is a semiconductor diode laser and the photodetector  822  is a semiconductor photodiode. 
     While the embodiments of  FIGS. 1 and 8  employ a four-sided optical path as the ring interferometer, the number of sides can be any number ranging from three up to any practical number. 
       FIG. 9  illustrates a modification of the sensor of  FIG. 8  in which the three waveguides sections  812 ,  814 ,  816  and the mirrors  806 ,  808  are replaced by a semicircular waveguide  910 , a section of which is occupied by the sensing volume  824 . 
       FIG. 10  illustrates an embodiment in which the ring interferometer is implemented using a full circular waveguide  1010  formed as a mesa structure on a substrate  1020 . Light is supplied to the circular waveguide  1010  by evanescent coupling from a linear waveguide  1030  adjacent or tangent to the circular waveguide  1010  at a contact point  1035 . The term “evanescent coupling” is defined in: Hunsperger, Robert M.,  Integrated Optics: Theory and Technology,  1991, pp. 110–113, Springer-Verlag, New York. As employed in this specification, the term “evanescent coupling” is the same as “optical tunneling” referred to in the foregoing publication. The linear waveguide  1030  has a laser  1018  at one end and a photodetector  1022  at the other end. The distance “a” between the laser  1018  and the contact point  1035 , the distance “b” between the contact point  1035  and the photodetector  1022  and the pathlength of the circular waveguide are all integral multiples of the wavelength of the laser  1018 . A sampling volume  1024  containing the Sol-Gel structure  200  of  FIGS. 2 ,  3  or  4  occupies a section of the circular waveguide  1010 . Light is coupled from the laser  1018  to the circular waveguide  1010  by evanescent coupling, and light is coupled from the circular waveguide to the photodetector by evanescent coupling between the circular and linear waveguides  1010 ,  1030  at the contact point  1035 . Thus, the photodetector  1022  receives a first light beam directly from the laser  1018  as well as a second (“shifted”) light beam that has passed through the sampling volume  1024 . The second light beam is phase shifted by an amount depending upon the number of antigen-antibody bound pairs in the sensing volume. Thus, the light intensity sensed by the photodetector  1022  is a function of the interference between the two light beams in the same manner as described above with reference to the sensors of  FIGS. 1 and 8 . 
     While the sensor has been described with reference to implementations in which an antibody for the targeted antigen is immobilized in the Sol-Gel structure, artificial antibodies, i.e., printed polymers, could be substituted for the antibodies immobilized in the Sol-Gel. As another alternative, the sensor could be employed to sense or measure other biological binding events, such as, for example, binding between a pair of complementary DNA strands. In this case, the Sol-Gel porous structure would be sized to immobilize single DNA strands. An analyte possibly containing the complementary DNA strand would be introduced to the Sol-Gel. Complementary DNA strands would bind to the immobilized DNA strands, thereby changing the refractive index of the Sol-Gel. Thus, in general, the sensor detects binding between complementary biological entities, one of which is immobilized in the Sol-Gel and one of which is the targeted entity to be detected. 
     While the invention has been described in detail by specific reference to preferred embodiments, it is understood that variations and modifications thereof may be made without departing from the true spirit and scope of the invention.