Abstract:
An integrated optical chip device for molecular diagnostics comprising a tunable laser cavity sensor chip using heterodyned detection at the juncture of a sensor laser and a reference laser, and including a microfluid chip to which the sensor chip is flip-chip bonded to form a sample chamber that includes exposed evanescent field material of the tunable laser cavity to which fluid to be diagnosed is directed.

Description:
CROSS-REFERENCE TO RELATED APPLICATION 
     This application claims the benefit of U.S. Provisional Patent Application Ser. No. 60/221,624, filed Jul. 28, 2000. 
    
    
     STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH OR DEVELOPMENT 
     This invention was made with Government Support under Grant No. DAAD19-00-1-0400, awarded by the Department of the Army, and Grant No. N00014-96-1-G014, awarded by the Office of Naval Research. The Government has certain rights in this invention. 
    
    
     FIELD OF THE INVENTION 
     This invention relates to the detection of a molecular species using heterodyned laser light. 
     BACKGROUND OF THE INVENTION 
     There has been long recognized a desire to automate the analysis of a wide variety of substances including chemical and biochemical materials, contaminants, biological warfare agents, and generally any substance, the presence and/or amount of which is desired to be determined. In recent years, on-chip systems have been developed for molecular diagnostics, e.g., for the detection of antigens by combination with antibodies or the analysis of nucleic acids via hybridization. The systems require the mixing of conjugate antibodies or the use of fluorescent antibodies or hybridizing fluorescent molecules during preparation, and, while being miniaturized nevertheless still require macroscopic techniques such as external light sources, external electro-optical detectors, and electronic instrumentation, all of which significantly limit the size and flexibility of such on-chip devices. Particularly as would be applied to military operations there is a need for fully integrated, field portable, and sensitive chip technology which can work reliably in demanding situations. Simply scaling down existing technologies, such as fluorescent measurement schemes, to the chip scale does not provide effective solutions. Moreover, any new technology must minimize meticulous sample preparation and handling steps, which limits the robustness of current technologies. 
     There has also been a growing need to develop microscale devices that can manipulate and transport relatively small volumes of fluids. These devices have applications in many areas of engineering, including propulsion and powered generation of micro-satellites, micro-air vehicles, inkjet printer heads, and bioanalytical instruments. See for example “PIV measurements of a microchannel flow” by C. D. Meinhart et al.,  Experiments in Fluids  (1999) 414-419, the disclosure of which is incorporated herein by reference. When dealing with minute quantities of contaminants, for example, methods of separating or isolating the molecules to be diagnosed become important. Electrophoretic systems have been developed which aid in such techniques. Such systems separate molecules by their unique directed motions in an electric field. 
     In recent years, lasers have been put to use in molecular diagnostics. Robert Frankel et al. U.S. Pat. No. 5,637,458 (the disclosure of which is incorporated herein by reference) describes a system for biomolecular separation and detection of a molecular species that uses a solid state laser detector formed with a sample channel. The presence of a molecular species is indicated by a frequency shift in the laser&#39;s output which is detected by optical heterodyning the laser&#39;s output with the output of a reference laser. The interior of the sample channel can, optionally, be coated with a ligand for binding a molecular species of interest. The system involves rather complex preprocessing of the sample by electro-osmotic separation in channels that are lithographically formed in a two dimensional planar substrate and/or by a nanostructural molecular sieve formed of spaced apart posts defining narrow channels. Although an attempt at integrated system is provided by U.S. Pat. No. 5,637,458, it does not entirely provide a fully integrated optical chip device. 
     Also recently, highly coherent semiconductors, lasers and laser arrays have been developed primarily for telecommunications applications. See for example, C. E. Zah et al., IEEE Photon. Technol. Lett. Vol. 8 pp. 864-866, July 1996. In addition, widely tunable semiconductor lasers have been developed, in particular, sampled-grating distributed Bagg reflector (SGDBR) lasers. See, for example “Tunable Sampled-Grading DBR Lasers with Integrated Wavelength Monitors,” by B. Mason et al.,  IEEE Photonics Technology Letters , Vol. 10, No. 8 August 1998; 1085-1087 and “Ridge Waveguide Sampled Grating DBR Lasers with 22-nm Quasi-Continuous Tuning Range,” by B. Mason et al.,  IEEE Photonics Technology Letters , Vol. 10, No. 9 September 1998, 1211-1213. These widely tunable lasers are based on the use of two-multi-element mirrors as described in Coldren, U.S. Pat. No. 4,896,325. The foregoing also includes a Y-branch splitter with a detector in each branch for wavelength determination: Disclosures of the foregoing three publications and Coldren, U.S. Pat. No. 4,896,325 are incorporated herein by reference. 
     SUMMARY OF THE INVENTION 
     The present invention provides a fully integrated optical sensor for on-chip analysis of immunoassays and molecular diagnostics. The present invention measures minute changes in the index of refraction (−10 −7 ), within one micron of a microchannel surface, which can be the result of a specific heterogeneous chemical reaction or an antigen-antibody binding event. 
     The present invention does not require mixing of conjugate antibodies or fluorescent molecules during sample preparation as used in related art devices and techniques. Further, the present invention does not require external devices such as external light sources, fluorescent filters, or external recording optics. Unlike fluorescence imaging, which is a macroscopic technique that is applied to bio-chips, the present invention operates at the microscopic scale. The system has sensitivities that can detect single molecules, is fully integratable into the chip, and avoids mixing steps during sample preparation. 
     In particular, an integrated optical chip device usable for molecular diagnostics in what we term a tunable laser cavity sensor (TLCS) is flip chip bonded to a microfluidic chip. The TLCS is formed from a reference laser and a sensor laser, each comprising a waveguide having a gain section, a partially transmissive mirror section, and a coherent light beam output section, one or both of the waveguides having a phase control section. The light beam output sections of the reference and sensor lasers are joined to enable the coherent light from these sections to interfere, providing a heterodyned frequency. The sensor laser has a thinned waveguide region exposing evanescent field material to form a cavity and which detects the presence of a molecule by a heterodyned frequency shift. 
    
    
     
       BRIEF DESCRIPTION OF THE DRAWINGS 
         FIG. 1  is a schematic view of an assembled biosensor/analyzer showing the optical sensor flip-chip bonded to the biofluidic chip; 
         FIG. 2  is a top plan schematic view of a heterodyned tunable reference and sensor lasers with an intracavity sensor region; 
         FIG. 3  is a schematic, cross-sectional vertical view of a microfluidic chip of this invention and the optical sensor flip-chip bonded thereto; 
         FIG. 4  is a bottom perspective view showing the tunable laser cavity sensor with control electrodes for gain, phase, and mirror currents; 
         FIG. 5  shows an exploded perspective view of the assembled biosensor/analyzer, similar to  FIG. 1 , but showing how the tunable laser cavity sensor is flip-chip bonded to the microfluidic chip; 
         FIG. 6  is a cross-section of a vertical schematic view of the assembled tunable laser cavity chip and microfluidic chip showing electrical and gasket connection and the interaction region thereof; 
         FIG. 7  is a top plan schematic view of a one-dimensional tunable laser cavity sensor array composed of multiple heterodyne tunable lasers with intracavity interaction regions; 
         FIG. 8  is a schematic plan view of the tunable laser cavity sensor of  FIG. 4 ; 
         FIG. 9  is a cross-sectional, schematic view of a ridge waveguide usable in the present invention; 
         FIG. 10  is a cross sectional perspective view of reference and sensor ridge waveguides; 
         FIG. 11  is a cross sectional schematic view of a buried rib waveguide usable in the present invention; 
         FIG. 12  is a cross sectional perspective schematic view of reference and sensor buried-rib waveguides; 
         FIG. 13  is a schematic plan view of the tunable laser cavity sensor of  FIG. 4 ; and 
         FIG. 14  is a schematic plan view of the tunable laser cavity sensor similar to that of  FIG. 13 , but with left and right side sampled-grating mirrors. 
     
    
    
     DETAILED DESCRIPTION OF THE INVENTION 
     Referring to  FIG. 1  an integrated optical chip device  10  in accordance with this invention is formed by flip-chip bonding an InP-based laser/detector chip  12  to a Si-biofluidic chip  14 . The InP optical sensor chip measures slight frequency shifts due to evanescent wave interactions with fluidic medium in a laser cavity, as will be described in more detail below, and can be referred to as a tunable laser cavity sensor (TLCS). 
     Referring additionally to  FIG. 3 , the biofluidic chip  14  contains a microfabricated flow cell with an opening  16  adjacent a cavity in the sensor chip  12  as will be described below in more detail. Microchannel  18  feeds fluid to the opening  16  and an outlet microchannel  20  removes effluent from the opening  16 . The opening  16  and the sensor cavity (not shown in  FIGS. 1 and 3 ) serve as a sample chamber, a diffusion-dominated region, where analyte can diffuse to a wall of the sensor cavity. An adsorbent, such as a capture antibody for immunoassays, or a ligand for a chemically reactive species, is provided, e.g., by deposition, adjacent the sensor cavity. When a particular reaction occurs on the surface, or an antigen binds to an antibody on the surface, a change in index of refraction will occur adjacent the surface and this changes the lasing frequency of the tunable laser cavity sensor, which is detected by a heterodyne detector, again as will be described in more detail below. 
     The fluidic channels  18  and  20  can be formed by deep reactive ion etching (DRIE) into a 300 micron thick Si wafer. DRIE provides an excellent means for machining high aspect ratio channels with good tolerances. Access ports  19  and  21  respectively, for the inlet and outlet channels  18  and  20  are etched into the bottom of the Si substrate. The inlet and outlet channels  18  and  20  are etched through the entire depth of the wafer. The opening  16  which connects to the sensor laser cavity is formed by a nominally 100×100 micron channel etched between the inlet and outlet channels  18  and  20  on the top surface of the chip  14 . A glass cover slip  23  seals the access ports  19  and  20  and is provided with corresponding openings  25  and  27 . 
     In the embodiment shown in  FIG. 3 , a pressure gradient, such as a syringe pump can be used to propel fluid through the device. See, for example, C. D. Meinhart et al., supra. 
     The TLCS optical sensor element is shown schematically in FIG.  2 . Two distributed-bragg reflector (DBR) tunable lasers  22  and  24  are integrated with a Y-branch coupler  26  and a photodetector  28 . One of the DBR tunable lasers  22  is a reference laser, the other  24  being a sensor laser. The photodetector  28  provides heterodyne detection of small changes in amplitude or frequency of the sensor laser  24  relative to the reference laser  22 . As is known, the frequencies of the reference and sensor lasers can be set, as indicated at  30  and  31  by adjustment of the control sections, more particularly by adjustment of the respective gain  32 ,  34  and phase  36 ,  38  sections of the waveguides. Each waveguide has a partially transmissive grating mirror section  40  and  42  and a coherent light beam output section  44  and  46  which are joined at the mixer detector section  28 . 
     The interactive region  29  of the sensor waveguide is formed between the gain and phase control sections, respectively  34  and  38 , and the sampled grating mirror section  42 . However, the particular order of the components between the mirrors is not critical and other configurations are equally useable. Thus, all permutations of the locations of the gain section  34 , phase control section  38  and interactive region  29  can be used. For example, the order from the cleaved facet  24  ( FIG. 4 ) can be phase control section  42 , gain section  38  and interactive region  29 , etc. Also, while a phase control section is shown on both the reference laser  22  and sensor laser  24 , it is sufficient to have it on only one of the lasers in order to tune one to the other. As indicated, the left ends of the lasers  22  and  24  are formed by cleaved facets. Both the left end facet mirrors and the right side grating mirrors can be sampled-grating mirrors to provide for wider tunability of the lasers output wavelength, in which case, the opposed sampled-grating mirrors would preferably have different sampling periods. Using lasers with different sampled grating periods is described in the aforementioned Coldren, U.S. Pat. No. 4,896,325. 
     As shown, the frequency output of the sensor waveguide differs by ±Δλ from the frequency of their reference waveguide. By adjusting the tuning electrodes as shown in  FIG. 2 , one can enhance the measurement resolution by tuning to possible molecular bond resonances, e.g. in the 1550 nm wavelength range. Researchers at the University of California in Santa Barbara have pioneered DBR lasers with extended tuning ranges—so called sampled grating-DBR lasers. The lasing wavelengths of these lasers can be tuned up to 100 nm, enabling the measurement of the index of the perturbing species versus wavelength over a relatively wide range to better identify their chemical nature. 
       FIGS. 13 and 14  show schematic plan views of TLCSs using either a simple DBR partially transmissive mirror or two SGDBRs, respectively. The TLCS of  FIG. 13  is that of  FIG. 4  shown in plan view, with corresponding lead lines. In the TLCS of  FIG. 14 , the SGDBR configuration replaces the simple grating on the right side as well as the opposite laser facet mirror with sampled grating mirrors, respectively  57  and  59 , for extended tuning range. 
     Referring to  FIG. 4 , the TLCS is shown in more detail. The tunable cavity sensor is fabricated by integrating a tunable DBR sensor laser  22  with a reference laser  24  and combining them into a heterodyning detector  28  to accurately monitor changes in the modal index for loss due to adsorbates or interactions at the surface of a thinned interaction region  48  on the sensor laser  22 . The InP chip  12  is formed with reference and sensor lasers  22  and  24 , as will be described in more detail hereinafter, each of which carries gain control electrodes, respectively,  50 ,  52  and phase control electrodes, respectively,  54 ,  56  spaced from mirror control electrodes, respectively,  58 ,  60  overlying a partially transmissive grating mirror  43 . As described with respect to  FIG. 4 , the reference and sensor coherent light beam output sections  62  and  64  join to deliver interfering light beams at the detector  28 , sensed at a detector electrode  66  thereon. Although a “Y-branch” waveguide combiner element  62  and  64  is shown, another type of waveguide combiner such as a “Multimode-interference” element, may also be employed as is well known to those skilled in the art. The cladding of the sensor laser waveguide  24  is thinned to form the sensor cavity  48  to expose the evanescent fields of the lasing mode, and provide an interaction region. 
     As in Frankel et al., U.S. Pat. No. 5,637,458, the surface of the cavity  48  can be coated with various ligands, such as capture antibodies, various binding molecules, or reactive molecules. After flip-chip bonding to the Si microfluidic chip, as described hereinafter, the thin waveguide region  48  then forms one side of an interaction chamber in which analytes can diffuse to the treated surface. When a particular reaction occurs on the surface, or an antigen binds to an antibody adsorbate on the surface, a change in index of refraction, Δn s , will occur at the region just above the surface. Since a portion of the laser mode, Γ xy , fills this transverse region, the modal index is changed by an amount, Γ xy  Δn. Also, the interaction region extends along the axis of the laser to fill an axial fraction Γ z , of the cavity, so that the net fill-factor for region in which the perturbation takes place is Γ xy  Γ z . 
     Since the lasing wavelength changes in direct proportion to the net weighted change in index (and frequency as the direct negative), the relative change in laser output wavelength, λ, (or frequency, f) is given by: 
         Δλ   λ     =         Γ   xy     ⁢     Γ   z     ⁢       Δ   ⁢           ⁢     n   s         n   _         =     -       Δ   ⁢           ⁢   f     f             
 
     For a typical sensing configuration, Δn s ,=0.01, and Γ xy  Γ z ,=0.01, and assuming the average index of the laser cavity is n=3.3, then Δλ=0.05 nm, or Δf=−6 GHz@λ=1550 nm. Now, if this deviation were to be measured in the optical domain, a quarter-meter or larger spectrometer would be necessary to obtain sufficient resolution to see the effect, which would be very difficult at the chip level. However, with an integrated heterodyne detector, the shifted optical frequency can be down converted to the VHF radio frequency range where simple frequency counters can be used to measure the difference frequency with 1 Hz accuracy. Using heterodyne detection with two semiconductor lasers, a 6 GHz frequency shift can be measured with an accuracy of about 10 MHz, because this is the approximate linewidth of such lasers. 
     Put another way, again assuming the index shift in the small perturbation region, ns=0.1, the net fill-factor of this region relative to the volume of the guided mode can be as small as □xy□z,=(10 MHz)(3.3)/(0.1)(193 THz)=1.7×10 6 . Then, for example, if the transverse over lap, □xy is only 0.1% (very conservative estimate of the evanescent field), the axial □z can be as small as 0.17%. Therefore, with a net laser cavity length of 500 μm, single submicron particles can be detected. 
       FIG. 5  depicts flip-chip bonding of the InP TLCS  10  to the Si-biofluidic chip  14 . In this embodiment, the biofluidic chip  14  is formed with integrated inlet and outlet channels, respectively,  68  and  70  leading to and from a sample cavity  72  having the gain and phase control circuitry  11  and heterodyne detection circuitry  13  integrated therewith, connecting to the InP chip components via the conductive lines, respectively,  15  and  17 , as previously shown in FIG.  1 . 
     In this fully-integrated design, the channels are sufficiently small so that capillary forces can be used to fill them or alternatively, an onboard pump could be used to propel the fluid. 
     Details of connection and operation of the integrated optical chip of the present invention are shown in FIG.  6 . The microfluidic chip  14  is shown with the direction of microchannel flow out of the plane. The chip  14  carries electrical contacts  74  and  76 , respectively, for the gain and phase control of the TLCS. The sample cavity  72  of the biofluidic chip  14 , (the thinned sensor cavity  48  of the TLCS) is interconnected by a gasket  78  to form a sample chamber  80  defining an interaction region. The exposed evanescent field material of the sensor chamber  48  is provided with an adsorbate layer  82 . The laser guided mode is illustrated at  84  showing propagation of the laser beam along the waveguide to and from the sensor mirror section  42  adjacent the sampled grating mirror  43 . 
     As shown, the InP optical sensor chip measures slight frequency shifts due to evanescent wave interactions with the fluid medium in the sample chamber  80 , which serves as a diffusion-dominated region where analytes can diffuse to the adsorbate layer  82 . The adsorbate layer, which can be referred to as an interaction layer, can be formed as a capture antibody for immunoassays for a ligand for some chemically reactive species. When a particular reaction occurs on its surface, or an antigen binds to an antibody on the surface, a change in index of refraction will occur adjacent the surface, and this changes the lasing frequency. The inclusion of an “interaction region within the cavity  48  of the sensor laser provides for a change in the modal index of refraction (gain or loss) within this region due to the surface absorption or chemical interaction, which overlap the evanescent fields of the laser mode. 
     The relative frequency change, Δf/f, of the laser is just equal to the relative modal index change times a fill factor, ΓΔn/n, and this frequency change, Δf, can be measured very accurately in the radio frequency (RF) range after down conversion by mixing with the unperturbed laser in the heterodyne detector, to measure changes in modal index of refraction inside the sensor laser cavity  48  with a resolution estimated at about Δf/f=10 MHz/200 THz˜10 −7 . 
     Antibody immobilization strategies utilizable with this invention can exhibit high sensitivity and high selectivity. For example, using antibodies immobilized to polystyrene and using waveguide illumination of fluorescence, it has previously been demonstrated that cTnI (troponinI) can be detected down to 1 pm. Sensitivity has been reported in the literature down to the fM range using thin-film silicon oxynitride waveguides approximately 1 micron thick; see Plowman et al., 1996. In a similar fashion, DNA has been detected down to the 50 fM level using evanescent planar waveguides with covalently attached capture oligonucleotides probes within twelve minutes; see Bucach et al. 1999. 
     In many situations it may be desired to detect more than one kind of molecular species or more than one kind of interaction. This may be possible by sweeping the wavelengths of the reference and sensor lasers by applying suitable currents to the control electrodes and observing characteristic resonances in the index measurement vs. λ. The use of a widely-tunable laser such as a sampled-grating DBR will facilitate this option. 
     Another approach to detect a multiplicity of species is to use a one-dimensional TLCS array on the same chip, as illustrated in  FIG. 7. A  plurality of TLCSs which can be a dozen or more, but of which only three TLCSs  86 ,  88  and  90  are shown. The TLCSs form an array interconnected by an elongate sample chamber  94 . The sample chamber can be contained on a Si biofluidic chip with separate sample cavities aligned with each sensor laser cavity, and/or a single gasket can surround a single sample cavity that runs across all of the TLCSs, forming a succession of sample chambers with successive interactive regions  98 ,  100  and  102 , whereby fluid flows serially from the first interactive region  98  to the last interactive region  102 , as shown by the arrow  104 . 
     Depending upon the binding chemistry deposited on the sensor cavity, each sensor cavity could measure a different constituent of the flow, such as pH, temperature, antigen, etc. A single fluidic flow cell doses each interaction region TLCS. The practical number of TLCS array elements and thus sensed properties, is mainly limited by the desired to finite chip size. The active elements, including the two DBR lasers are spaced, e.g., by about 500 μm so as to allow space for flip-chip contacts and to avoid cross talk. Thus, the device is applicable to the analysis of a broad range of chemical and biological assays. For example, one could test for such biological warfare agents as Botulinium Toxin, Ebala and Anthrax, by using Ovalbumin, MSZ and Bacillus Globigil to simulate the invasion by such warfare agents into a human bloodstream. Again, spectral index information can also supplement the index information at each element if the wavelengths are varied across some range. 
     In a further embodiment of the invention, illustrated in  FIG. 8 , a series of electrodes for dielectrophoresis (DEP) can be fabricated in the microchannel sample cavity  72  and, with the sensor cavity  48 , forms the sample chamber  80  (all with reference to the components of FIG.  6 ). Many biological particles (such as cells and large macromolecules) exhibit both positive and negative diaelectrophoretic, constants, depending upon the frequency of applied electric field. See Jones, 1995. By changing the frequency and intensity of the electric field, dielectrophoresis can be used to induce biological particles toward or away from the DEP electrodes and the sensor. The force due to DEP is proportional to particle volume, and therefore will be more effective for large particles. 
     When a biological particle exhibits a positive dielectrophoretic constant, the particles can be induced toward the DEP electrodes and will be less likely to deposit on the laser sensor cavity area.  FIG. 8 , shows the application of force to a fluid, at  106 , carrying particles  108 . When sensing analytes with low particle concentrations, the frequency of the electric field can be adjusted to increase the concentration of particles near the sensor area, as shown at  110 , making the measurements more sensitive. Considering the situation where one is continuously monitoring particle concentrations in a flowing fluid by observing the concentration of particles attached to the sensor wall, and knowing a prior the equilibrium constant for the reaction at the wall, the particle concentration can be measured most accurately over a limited range, depending on the optimum measurement concentration at the sensor. This range can be extended using DEP and the TLCS for feedback control. This system can be calibrated by applying known particle concentrations, varying DEP frequency and amplitude, and monitoring measurements from the TLCS. 
     DEP has been used to increase particle concentrations, separate particles, and capture particles with relatively low voltages compared to electrophoresis. Miles et al., (1999) used DEP to manipulate DNA,  Bacillus globigii  spores and  Erwinia herbicola  bacteria. They demonstrated the feasibility of capturing DNA molecules using DEP, with a relatively simple microfluidic device. While Washizu et al. (1994, 1995), used DEP to stretch and position DNA molecules and biopolymers. DEP coupled with field-flow-fractionation has been used successfully to separate polystyrene beads. Wang et al (1998), and to separate human breast cancer cells from normal blood cells, Yang et al. (1999). The technique of this invention therefore builds upon established technology in the field of optical immunosensors. These sensors use optical detection techniques to determine the presence and concentration of antigens by monitoring antigen/antibody binding reactions to capture antibodies that are immobilized to a wall, Rabbany et al. (1994). 
     The dielectrophoresis force of a lossless dielectric sphere is given by Jones (1995) as 
               F   DEP     =     2   ⁢     πɛ   1     ⁢     R   3     ⁢   K   ⁢     V   _     ⁢     E   o   2               (   I   )             
 
where ε 1  and ε 2  are the permittivity of the fluid medium and the lossles dielectric sphere, R is the radius of the sphere, E o  is the applied electric field. The dielectric constant K can be written using the Clausius-Mossotti function (Jones, 1995) 
       K   =         ɛ   2     -     ɛ   1           ɛ   2     +     ɛ   2             
 
Equation (I) indicates the DEP force is proportional and parallel to the gradient of the electric field squared, and proportional to the cube of the sphere radius. The DEP force is present only for spatially varying electric fields and works in either AC or DC fields. If the permittivity of a particle is greater than its surrounding medium, then K&gt;0 and the particle is said to have a positive dielectrophoretic constant and is attracted in increasing electric fields.
 
     Bahaj and Bailey (1979) state that for geometrically similar electrodes, the DEP force scales as 
               F   DEP     ≈       V   2       L   e   3               (   3   )             
 
where V is the magnitude of the applied voltage and L e  is the effective length of the electrodes. Therefore, smaller geometries will increase the sensitivity of a particle to the dielectrophoretic effect (Jones, 1995). In addition, for a constant DEP force decreasing the geometric length scale, allows for a reduction in the applied voltage.
 
     In the case of conductive losses, the DEP constant K can be a function of the applied voltage frequency. Therefore, the magnitude and direction of the DEP force can be manipulated by varying the voltage frequency. The biological particles that exhibit K&lt;0 can be passively levitated using DEP so that they will be less likely to deposit on channel walls. When high-sensitivity detection is desired, the electric field can be adjusted (i.e. in magnitude and frequency) so that the concentration of particles near the laser sensor interface is increased, making the molecular detection more sensitive. 
     In fabricating the TLCS chip, known InP growth and fabrication procedures and DBR laser fabrication characterization procedures can be used. Existing 3-D beam propagation modeling (BPM) software can be utilized to provide inclusion of lateral and transverse variations in straight guides, such as in the interaction region, as well as the actual variations in bends, such as in the Y-branches offset regions for gain and detector circuitry, as shown in  FIG. 5 , will be used. 
     Referring to  FIGS. 9 and 10 , after a first growth, the lower band gap gain/detector layers are removed in the passive sections and the grating lines are etched into the underlying passive guide in the grating mirror section.  FIG. 9 , a transverse cross section of a ridge waveguide is shown. The InGaAsP waveguide  112  is formed on an n-InP buffer and substrate  114 . A p-InP ridge waveguide  116  is formed on the InGaAsP waveguide (regrowth) to provide the top cladding and contact layers, the latter formed by InGaAs. Sampled grating lasers can be made with the same procedure. See for example Mason et al. (1998). 
     Referring to  FIG. 10 , to form the sensor cavity  48  containing the interaction region, the cladding over the optical waveguide is thinned to expose the vertical evanescent optical field. This results in a much smaller ridge height over the center of the guide but some lateral ridge structure must remain to provide lateral waveguiding. The resultant TLCS with its reference waveguide  116  and sensor waveguide  118  are thus formed. Inert polymer  120  is left at the corners of the ridge guides  116  and  118  to eliminate interactions with the fluid, which is especially important for the reference laser which is not to be affected by the fluid. 
     Referring to  FIGS. 11 and 12 , in another embodiment of the invention, the waveguides can be buried-rib waveguides formed by etching away all the layers outside of the desired optical channel. As shown in  FIG. 11 , the n-InP substrate  122  carries a waveguide  124  and adjacent quantum well  128  in a p-InP layer contained in an implanted region  126  under a SiNx layer  130 , an InGaAs contact layer  132  and Ti/Pt/Au contact layer  134  providing electrical contact. 
     As shown in  FIG. 12 , for the buried-rib embodiment, thinning results in a uniform lateral surface  136 , obtained by removing the passive waveguide layer beneath the surface. The result is a TLCS  138  containing reference and sensor waveguides  140  and  142  with the sensor cavity  144  defining the interactive region of the TLCS. 
     Referring again to  FIG. 6 , to form the adsorbent layer  82 , one can coat the InP laser cavity with a thin film of silicone or other hydrophobic polymer material. For example, antitroponin I can be deposited onto the thin film of silicone. Pluronics, block co-polymers, can be used as an intermediate in binding antibody to a surface. In one embodiment, the coated surfaces passivated or blocked with a sugar/protein mixture to both stabilize the deposited antibody and to cover portions of the InP surface where antibodies are not present. Ideally, the application process and drying process are optimized to the thinnest layer possible to make the surface immediately active, to minimize non-specific binding and to stabilize the antibody activity. The passivating can be sprayed onto the antibody-coated surface in a fine mist until the surface is wetted. The wetted surface can then be washed thoroughly with a buffer solution to remove excess protein and sugar, and antibody that has been loosened in the passivating process. A final layer of passivating material can then be applied to maximize the stability of the active antibody. 
     The captured chemistry can be deposited on the small 3 μm×500 μm interaction region of the laser cavity sensor. When an array of multiple laser cavity sensors are used in a single microfluidic channel, adjacent laser cavities, which are positioned approximately 500 μm apart, are each coated with a separate reference chemistry. 
     The detector signal from the heterodyne-mixed laser cavity sensor will contain a beat frequency, which will correspond to the amount of bound target analyte. The relationship between the beat-frequency versus time occurred and the target species concentration can be characterized. One way of handling the beat -frequency versus time relationship is to measure the time evolution of the beat frequency. One can then correlate the curve to a known concentrate of analyte, and a known flow condition. 
     While the invention has been described in terms of specific embodiments, various modifications can be made without departing from the scope of the invention. 
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