Abstract:
The present invention relates to an implantable microfabricated sensor device and system for measuring a physiologic parameter of interest within a patient. The implantable device is micro electromechanical system (MEMS) device and includes a substrate having an integrated inductor and at least one sensor formed thereon. A plurality of conductive paths electrically connect the integrated inductor with the sensor. Cooperatively, the integrated inductor, sensor and conductive paths defining an LC tank resonator.

Description:
CROSS REFERENCE TO RELATED-APPLICATION 
     This application claims priority to prior U.S. provisional application No. 60/263,327 (filed Jan. 22, 2001) and U.S. provisional application No. 60/278,634 (filed Mar. 26, 2001). 
    
    
     BACKGROUND OF THE INVENTION 
     Field of the Invention 
     The present invention generally relates to the field of MEMS (micro-electromechanical systems) sensors and more specifically to a wireless MEMS capacitive sensor for implantation into the body of a patient to measure one or more physiologic parameters. 
     A number of different biologic parameters are strong candidates for continuous monitoring. These parameters include, but are not limited to blood pressure, blood flow, intracranial pressure, intraocular pressure, glucose levels, etc. Wired sensors, if used have certain inherent limitations because of the passage of wires (or other communication “tethers”) through the cutaneous layer. Some limitations include the risks of physical injury and infection to the patient. Another risk is damage to the device if the wires (the communication link) experience excessive pulling forces and separate from the device itself. Wireless sensors are therefore highly desirable for biologic applications. 
     A number of proposed schemes for wireless communication rely on magnetic coupling between an inductor coil associated with the implanted device and a separate, external “readout” coil. For example, one method of wireless communication (well-known to those knowledgeable in the art) is that of the LC (inductor-capacitor) tank resonator. In such a device, a series-parallel connection of a capacitor and inductor has a specific resonant frequency, expressed as 1/√{square root over (LC)}, which can be detected from the impedance of the circuit. If one element of the inductor-capacitor pair varies with some physical parameter (e.g. pressure), while the other element remains at a known value, the physical parameter may be determined from the resonant frequency. For example, if the capacitance corresponds to a capacitive pressure sensor, the capacitance may be back-calculated from the resonant frequency and the sensed pressure may then be deduced from the capacitance by means of a calibrated pressure-capacitance transfer function. 
     The impedance of an LC tank resonator may be measured directly or it may also be determined indirectly from the impedance of a separate readout coil that is magnetically coupled to the internal coil. The latter case is most useful for biologic applications since the sensing device may be subcutaneously implanted, while the readout coil may be located external to the patient, but in a location that allows magnetic coupling between the implanted sensing device and readout coil. It is possible for the readout coil (or coils) to simultaneously excite the resonator of the implanted device and sense the reflected back impedance. Consequently, this architecture has the substantial advantage of requiring no internal power source, which greatly improves its prospects for long-term implantation (e.g. decades to a human lifetime). 
     Such devices have been proposed in various forms for many applications. Chubbuck (U.S. Pat. No. 4,026,276), Bullara (U.S. Pat. No. 4,127,110), and Dunphy (U.S. Pat. No. 3,958,558) disclose various devices initially intended for hydrocephalus applications (but also amenable to others) that use LC resonant circuits. The &#39;276, &#39;110, and &#39;558 patents, although feasible, do not take advantage of recent advances in silicon (or similar) microfabrication technologies. Kensey (U.S. Pat. No. 6,015,386) discloses an implantable device for measuring blood pressure in a vessel of the wrist. This device must be “assembled” around the vessel being monitored such that it fully encompasses the vessel, which may not be feasible in many cases. In another application, Frenkel (U.S. Pat. No. 5,005,577) describes an implantable lens for monitoring intraocular pressure. Such a device would be advantageous for monitoring elevated eye pressures (as is usually the case for glaucoma patients); however, the requirement that the eye&#39;s crystalline lens be replaced will likely limit the general acceptance of this device. 
     In addition to the aforementioned applications that specify LC resonant circuits, other applications would also benefit greatly from such wireless sensing. Han, et al. (U.S. Pat. No. 6,268,161) describe a wireless implantable glucose (or other chemical) sensor that employs a pressure sensor as an intermediate transducer (in conjunction with a hydrogel) from the chemical into the electrical domain. 
     The treatment of cardiovascular diseases such as Chronic Heart Failure (CHF) can be greatly improved through continuous and/or intermittent monitoring of various pressures and/or flows in the heart and associated vasculature. Porat (U.S. Pat. No. 6,277,078), Eigler (U.S. Pat. No. 6,328,699), and Carney (U.S. Pat. No. 5,368,040) each teach different modes of monitoring heart performance using wireless implantable sensors. In every case, however, what is described is a general scheme of monitoring the heart. The existence of a method to construct a sensor with sufficient size, long-term fidelity, stability, telemetry range, and biocompatibility is noticeably absent in each case, being instead simply assumed. Eigler, et al., come closest to describing a specific device structure although they disregard the baseline and sensitivity drift issues that must be addressed in a long-term implant. Applications for wireless sensors located in a stent (e.g., U.S. Pat. No. 6,053,873 by Govari) have also been taught, although little acknowledgement is made of the difficulty in fabricating a pressure sensor with telemetry means sufficiently small to incorporate into a stent. 
     Closed-loop drug delivery systems, such as that of Feingold (U.S. Pat. No. 4,871,351) have likewise been taught. As with others, Feingold overlooks the difficulty in fabricating sensors that meet the performance requirements needed for long-term implantation. 
     In nearly all of the aforementioned cases, the disclosed devices require a complex electromechanical assembly with many dissimilar materials, which will result in significant temperature- and aging-induced drift over time. Such assemblies may also be too large for many desirable applications, including intraocular pressure monitoring and/or pediatric applications. Finally, complex assembly processes will make such devices prohibitively expensive to manufacture for widespread use. 
     As an alternative to conventionally fabricated devices, microfabricated sensors have also been proposed. One such device is taught by Darrow (U.S. Pat. No. 6,201,980). Others are reported in the literature (see, e.g. Park, et al., Jpn. J. Appl. Phys., 37 (1998), pp. 7124-7128; Puers, et al., J. Micromech. Microeng. 10 (2000), pp. 124-129; Harpster et al., Proc. 14 th  IEEE Int&#39;l. Conf. Microelectromech. Sys. (2001), pp. 553-557). 
     Past efforts to develop wireless sensors have separately located the sensor and inductor and have been limited to implant-readout separation distances of 1-2 cm at most, rendering them impractical for implantation much deeper than immediately below the cutaneous layer. This eliminates from consideration wireless sensing applications, such as heart ventricle pressure monitoring or intracranial pressure monitoring, that inherently require separation distances in the range of 5-10 cm. In the present state-of-the-art, several factors have contributed to this limitation on the separation distance including 1) signal attenuation due to intervening tissue, 2) suboptimal design for magnetic coupling efficiency; and 3) high internal energy losses in the implanted device. 
     In view of the above and other limitations on the prior art, it is apparent that there exists a need for an improved wireless MEMS sensor system capable of overcoming the limitations of the prior art and optimized for signal fidelity, transmission distance and manufacturability. It is therefore an object of the present invention is to provide a wireless MEMS sensor system in which the sensing device is adapted for implantation within the body of patient. 
     A further object of this invention is to provide a wireless MEMS sensor system in which the separation distance between the sensing device and the readout device is greater than 2 cm, thereby allowing for deeper implantation of the sensing device within the body of a patient. 
     Still another object of the present invention is to provide a wireless MEMS sensor system in which the sensing device utilizes an integrated inductor, an inductor microfabricated with the sensor itself. 
     It is also an object of this invention to provide a wireless MEMS sensor system in which the sensing device is batteryless. 
     A further object of the present invention is to provide a wireless MEMS sensor system. 
     BRIEF SUMMARY OF THE INVENTION 
     In overcoming the limitations of the prior art and achieving the above objects, the present invention provides for a wireless MEMS sensor for implantation into the body of a patient and which permits implantation at depths greater than 2 cm while still readily allowing for reading of the signals from the implanted portion by an external readout device. 
     In achieving the above, the present invention provides a MEMS sensor system having an implantable unit and a non-implantable unit. The implantable unit is microfabricated utilizing common microfabricating techniques to provide a monolithic device, a device where all components are located on the same chip. The implanted device includes a substrate on which is formed a capacitive sensor. The fixed electrode of the capacitive sensor may formed on the substrate itself, while the moveable electrode of the capacitive sensor is formed as part of a highly doped silicon layer on top of the substrate. Being highly doped, the silicon layer itself operates as the conductive path for the moveable electrode. A separate conductive path is provided on the substrate for the fixed electrode. 
     In addition to the capacitive sensor, the implanted sensing device includes an integrally formed inductor. The integral inductor includes a magnetic core having at least one plate and a coil defining a plurality of turns about the core. One end of the coil is coupled to the conductive lead connected with the fixed electrode while the other end of the coil is electrically coupled to the highly doped silicon layer, thereby utilizing the silicon layer as the conductive path to the moveable electrode. 
     In order to optimize the operation of the inductor and to permit greater implantation depths, a novel construction is additionally provided for the magnetic core. In general, the optimized magnetic core utilizes a pair of plates formed on opposing sides of the substrate and interconnected by a post extending through the substrate. The windings of the coil, in this instance, are provided about the post. 
     The external readout device of the present system also includes a coil and various suitable associated components, as well known in the field, to enable a determination of the pressure or other physiologic parameter being sensed by the implanted sensing device. The external readout device may similarly be utilized to power the implanted sensing device and as such the implanted sensing device is wireless. 
     Integrally formed on the implanted device and microfabricated therewith, may be additionally be active circuitry for use in conjunction with capacitive sensor. Locating this circuitry as near as possible to the capacitive sensor minimizes noise and other factors which could lead to a degradation in the received signal and the sensed measured physiologic parameter. As such, the active circuitry may be integrally microfabricated in the highly doped silicon layer mentioned above. 
     Further object and advantages of the present invention will become apparent to those skilled in the art from a review of the drawings in connection with the following description and dependent claims. 
    
    
     
       BRIEF DESCRIPTION OF THE DRAWINGS 
         FIG. 1  is a schematic illustration of a wireless MEMS sensor system according the principles of the present invention; 
         FIG. 2  is a graphical illustration of impedance magnitude and phase angle near resonance, as sensed through a readout coil; 
         FIG. 3  is a cross-sectional representation of a sensing device embodying the principles of the present invention. 
         FIGS. 4A and 4B  are schematic illustrations of the magnetic field distribution with  FIG. 4A  illustrating the magnetic field distribution of prior art devices and with  FIG. 4B  illustrating the magnetic field distribution for a sensing device having a magnetic core embodying the principles of the present invention; 
         FIG. 5  is an enlarged cross-sectional view of the diaphragm portion of  FIG. 3  operating in what is herein referred to as a “proximity” mode; 
         FIG. 6  is a cross-sectional view similar to that seen in  FIG. 5  illustrating, however, the diaphragm operating in what is herein referred to as a “touch” mode; 
         FIG. 7  is a capacitance versus pressure curve in the proximity and touch modes of operation; 
         FIG. 8  is a top plane view of a second embodiment of the main electrode in the capacitive sensor portion of the implanted sensing device according to the principles of the present invention; 
         FIG. 9  is a diagrammatic illustration of one scheme for providing electrically isolated paths for the connections and electrodes of the capacitive sensor portion; 
         FIG. 10  is a diagrammatic illustration of another scheme for electrically isolating the conductive paths for the connections and contacts of the capacitive sensor portion; 
         FIG. 11  is a cross-sectional view, generally similar to that seen in  FIG. 3 , further incorporating active circuitry into the sensing device; 
         FIG. 12  is a block diagram illustrating one possible circuit implementation of the active circuitry when incorporated into the sensing device of the present wireless MEMS sensing system; 
         FIG. 13  illustrates one method of mounting, within the body of a patient, a sensing device embodying the principles of the presents invention; 
         FIG. 14  illustrates a second embodiment by which a sensing device embodying the principles of the present invention may be secured to tissues within the body of a patient 
         FIGS. 15 and 16  are diagrammatic illustrations of different embodiments for locating a sensing device according to the principles of the present invention, within a vessel in the body of a patient; 
         FIG. 17  illustrates a sensing device, according to the principles of the present invention, encapsulated in a material yielding a pellet-like profile for implantation into the tissues in the body of a patient; 
         FIG. 18  illustrates a sensing device according to the principles of the present invention being located within the electrode tip of an implantable stimulation lead, such as that used for cardiac pacing; 
         FIG. 19  illustrates a plurality of sensing devices according to the present invention located within a catheter and utilized to calculate various physiologic parameters within a vessel within the body of a patient; 
         FIG. 20  is a schematic illustration of multiple sensors being used to measure performance of a component in the body or a device mounted within the body of a patient; 
         FIG. 21  illustrates a sensing device according to the principles of the present invention being utilized to measure pressure externally through a vessel wall; 
         FIG. 22  illustrates a portion of a further embodiment of the present invention in which the pressure sensing features of the sensing device have been augmented over or replaced with a structure allowing a parameter other than pressure to be sensed; 
         FIG. 23  is schematic perspective view, with portions enlarged, illustrating an alternative embodiment for sensing according to the principles of the present invention; and 
         FIG. 24  is an embodiment generally similar to that seen in  FIG. 23  for sensing according to the principles of the present invention. 
     
    
    
     DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS 
     In order to provide for battery-less, wireless physiologic parameter sensing over significant distances greater than 2 cm (e.g. 10 cm), the present invention provides a wireless MEMS sensing system, generally designated at  10  and seen schematically in FIG.  1 . The system  10  includes a microfabricated implantable sensing device  12 , optimized for coupling with an external readout device  14 . The sensing device  12  is provided with an integrated inductor  16  that is conductive to the integration of transducers and/or other components necessary to construct the wireless sensing system  10 . As an example, the preferred embodiment integrates a capacitive pressure sensor  18  into a common substrate  20  with the integrated inductor  16 . A second inductor  24 , in the readout device  14 , couples magnetically  26  with the integrated inductor  16  of the sensing device  12 . 
     The readout device  14  is constructed according to techniques well known in the industry and in the sensing field in general. As such, the readout device  14  is not illustrated or described in great detail. It is noted, however, that the readout device  14  may be included, in addition to its inductor  24 , signal conditioning, control and analysis circuitry and software, display and other hardware and may be a stand alone unit or may be connected to a personal computer (PC) or other computer controlled device. 
     The magnetic coupling  26  seen in  FIG. 1  allows the impedance of the LC tank circuit  22  to be sensed by the readout device  14 . The typical impedance magnitude  28  and phase angle  30  near resonance  32 , as sensed through the readout coil  14 , is seen in FIG.  2 . Real-time measurement and analysis of this impedance and changes therein allows the sensed pressure to be determined as previously mentioned. 
     Referring now to  FIG. 3 , a cross section of a preferred embodiment of the sensing device  12  is illustrated therein. The sensing device  12  includes a main substrate  34  (preferably 7740 Pyrex glass) formed and located within recessed regions of the substrate  34  are those structures forming the integrated inductor  16 . The integrated inductor  16  is seen to include a magnetic core  33  defined by a top plate  36 , a bottom plate  38  and a post  40  connecting the top plate  36  to the bottom plate  38  and being continuous through the substrate  34 . The plates  36  and  38  and the post  40  are preferably constructed of the same material, a ferromagnetic material and are monolithic. The integrated conductor  16  additionally includes a coil  42 , preferably composed of copper or other high-conductivity material, successive turns of which surround the post  40  of the magnetic core  33 . 
     In  FIG. 3 , the coil  42  is seen as being recessed into the top plate  36 . The coil  42  may additionally be planar or layered and preferably wraps as tightly as possible about the post  40 . If the material of the coil  42  has a high electrical resistance relative to the material of the core  33 , (as in a copper coil and NiZn ferrite core system) the core  33 , and specifically the top plate  36  may be directly deposited on top of the coil  42  without need for a intermediate insulating layer. If the electrical resistance of the coil material relative to the coil material is not high, an intermediate insulating layer must be included between the successive turns of the coil  42  and the core  33 . 
     Top and bottom cap layers  44  and  46  respectively, are provided over upper and lower faces  48  and  50  of the substrate  20  and over the top and bottom plates  36  and  38  of the magnetic core  33 . To accommodate any portions of the magnetic core  33  that extend significantly above or below the upper and lower faces  48  and  50  of the substrate  20 , the cap layers  44  and  46  may be provided with recesses  52  and  54 , respectively. Preferably, the cap layers  44  and  46  are of monocrystalline silicon. Other preferred materials include polycrystalline silicon, epitaxially deposited silicon, ceramics, glass, plastics, or other materials that can be bonded to lower substrate and/or are suitable for fabrication of the sensor diaphragm. In lieu of a monolithic cap layer, several sub-pieces may be fabricated at separate process steps, together forming a complete cap layer after processing is finished. 
     The coupling effectiveness of the integrated inductor  16  is a function of the magnetic flux enclosed by the windings of the coil  42 ; therefore the coupling is greatest if the structure of the integrated inductor  16  maximizes the flux encompassed by all of the winding loops.  FIG. 4A  shows schematically the magnetic field distribution  56  in a known inductor structure having a single core layer  58  and associated windings  60 . Schematically shown in  FIG. 4   b  is the magnetic field distribution  62  for an inductor structure  16 ′ having upper and lower plates  36 ′ and  38 ′, connected by a post  40 ′ about which windings of a coil  42 ′ are located, as generally seen in the present invention. The design of the present invention optimizes the inductor geometry for maximum field coupling. Placing the plates  36  and  38  on opposite sides of the substrate  20 , as in  FIG. 3 , increases the plate-to-plate spacing. The increased plate spacing creates a localized path of least resistance for the free-space magnetic field of an external readout coil, causing the magnetic field to preferentially pass through the post  40  of the integrated inductor&#39;s magnetic core  33 . This increases device effectiveness since the coupling efficiency between the sensor and a readout unit increases with the total magnetic flux encompassed by the windings of the inductor. A greater coupling efficiency increases the maximum separation distance between the sensor and a readout unit. 
     The materials used to form the integrated inductor  16  should be chosen and/or processed to maximize the above mentioned effect and to minimize drift in the inductance value across time, temperature, package stress, and other potentially uncontrolled parameters. A high-permeability material such as NiZn ferrite is used to maximize this effect on the magnetic field and to minimize drift. Other preferred materials include nickel, ferrite, permalloy, or similar ferrite composites. 
     To the right of the integrated inductor  16  seen in  FIG. 3  is the capacitive pressure sensor  18 . The capacitive pressure sensor  18  may be constructed in many forms commonly know to those familiar with the art. In the illustrated embodiment, the upper cap layer  44  is formed to define a diaphragm  64 . The diaphragm  64  constitutes and may also be referred to as the moveable electrode of the pressure sensor  18 . The fixed electrode  66  of the pressure sensor  18  is defined by a conductive layer formed on the upper face  48  of the substrate  20 , in a position immediately below the moveable electrode or diaphragm  64 . If desired, a conductive layer may additionally be located on the underside of the moveable electrode  64 . To prevent shorting between the upper electrode  64  (as defined by either the diaphragm itself or the diaphragm and the conductive layer  68 ) and the lower electrode  66 , one or both of the electrodes  64  and  66  may be provided with a thin dielectric layer (preferably less than 1000 Å) deposited thereon. 
     To improve performance of the capacitive pressure sensor  18 , as seen in  FIG. 8 , one or more secondary electrodes designated at  70  may be located about the fixed electrode  66  near the projected edge of the diaphragm  64  where pressure induced deflection of the diaphragm  64  is minimal. The secondary electrodes  70  experience all of the capacitance-effecting phenomena seen by the main electrode  66 , with the exception of any pressure-induced phenomena. The secondary electrodes  70 , as such, operate as reference electrodes and by subtracting the secondary electrodes&#39; capacitive measurement from the capacitive measurement of the main electrode  66 , most or all non-pressure-induced capacitance changes (signal drift) may be filtered out. Examples as sources of signal drift, that may be filtered out by this method, include thermally induced physical changes and parasitics resulting from an environment with changing dielectric constant, such as insertion of the sensor into tissue. In a preferred embodiment, the secondary (or reference) electrodes  70  would require an additional coil, similar to construction of the previously mentioned coil  42  to form a separate LC tank circuit. It is noted, that both coils may, however, share the same core post  40 . 
     Under normal operation, pressure applied to the exterior or top surface of the capacitive pressure sensor  18  causes the diaphragm  64  (or at least the center portions thereof) to deflect downward toward the fixed electrode  66 . Because of the change in distance between the fixed electrode  66  and the moveable electrode  64 , a corresponding change will occur in the capacitance between the two electrodes. The applied pressure is therefore translated into a capacitance. With this in mind, it is seen that the capacitance pressure sensor  18  may be operated in either of two modes. 
     A first mode, hereinafter referred to as the “proximity” mode, is generally seen in FIG.  5 . In this mode of operation, the starting gap between the fixed electrode  66  and the moveable electrode  64 , as well as the material and physical parameters for the diaphragm  64  itself, are chosen such that the fixed electrode  66  and the moveable electrode  64 will be spaced apart from one another over the entire operating pressure range of the sensor  18 . For the standard equation of parallel plate capacitance, C=∈A/d, the plate separation d will vary with the applied pressure, while the plate area A and the permittivity ∈ remain constant. 
     In the touch mode of operation, generally seen in  FIG. 6 , the geometry (e.g., initial gap spacing between the fixed electrode  66  and the moveable electrode  64 ) as well as the material and physical parameters of the diaphragm  64  itself, are chosen such that the fixed electrode  66  and the moveable electrode  64  will progressively touch each other over the operating pressure range of the sensor  18 . Accordingly, the area  72  of the fixed electrode  66  and the moveable electrode  64  in contact with each other will vary with the applied pressure. In the touch mode of operation, the dominant capacitance is the capacitance of the regions of the fixed electrode  66  and the moveable electrode  64  in contact with one another (if the dielectric coating  74  is thin compared to the total gap thickness, thereby yielding a relatively small effective plate separation distance d). In the capacitance equation mentioned above, plate separation d and permittivity ∈ will remain constant (at approximately that of the dielectric thickness) while the plate contact area A varies with the applied pressure. 
     In the graph of  FIG. 7 , capacitance-pressure relationship in the proximity and touch modes, respectively designated at  76  and  78 , are seen. From a practical standpoint, the operational mode may be chosen based upon sensitivity, linearity, and dynamic range requirements. The touch mode typically yields higher sensitivity with a more linear output, but involves mechanical contact between surfaces and therefore requires a careful choice of the materials to avoid wear induced changes in performance of the pressure sensor  18 . 
     To permit the innermost turn of the coil  42  to be electrically connected to the moveable electrode  66 , a post  80  (formed integral with the substrate  20 ) extends upward through the top plate  36  and a conductive trace  82  runs up the side of the post  80 . The trace  82  begins at the innermost turn of the coil  42  and proceeds to a point where the trace  82  makes electrical contact with the upper cap layer  44 . Preferably of monocrystalline silicon and highly doped to be conductive, the upper cap layer  44  serves as the electrical connection between the trace  82  and moveable electrode  64 . If the upper cap layer  44  is not conductive, an additional conductive trace along the upper cap layer  44  to the moveable electrode  64  will be utilized. The outermost turn of the coil  42  is connected by an electrical trace  84 . Where the upper cap layer  44  is conductive, a dielectric layer  86  insulates the trace  84  from the upper cap layer  44 . Alternatively, a p-n junction structure (as further described below) could be used. 
     It is noted that the inner and outer turns of the coil  42  may be alternatively connected respectively to the fixed electrode  66  and the moveable electrode  64 , thereby reversing the polarity of the LC tank circuit  22  if desired. Additionally, the particular paths between the coil  42  and the electrodes  66  and  64  may also be varied (e.g., such that both are included on the substrate  20 ) as best suited by the fabrication process. In all cases, the resistance of the electrical path through the traces  82 ,  84  and the upper cap layer  44  (if used) should be minimized. 
     The upper and lower cap layers  44  and  46  are bonded to the substrate  20  preferably via a hermetic sealing process. Alternatively, a post-bond coating of the entire sensing device  12  may be used to establish hermeticity. In either situation, steps are taken to minimize the residual gas pressure within the sensing device  12  after a hermetic seal is established. Once the initial hermetic seal is achieved, gas may be trapped in the interior of the sensing device  12  due to continued outgassing of the interior surfaces and/or the bonded regions. Gas pressure of the residual gas will increase within the interior chamber  90  of the pressure sensor  18  as the diaphragm  64  deflects during normal operation. This residual gas may effect the overall sensitivity of the pressure sensor  18  by effectively increasing the spring constant of the diaphragm  64 . Additionally, the residual gas will expand and/or contract with changes in the temperature of the sensing device  12  itself, causing signal drift. 
     To compensate for the various negative effects of any residual gas, the pressure sensor  18  is provided with a displacement cavity  88 . This displacement cavity  88  is generally seen in FIG.  3  and is in communication either directly or through a small connecting channel with the interior chamber  90  of the pressure sensor  18 , defined between the diaphragm  64  and the fixed electrode  66 . The displacement cavity  88  is sized such that the total internal sensor volume, the combined volume of the displacement cavity  88  and the interior chamber  90 , varies minimally with deflection of the diaphragm  64  over its operational range of displacement. By minimizing the overall change in volume with deflection of the diaphragm  64 , the effect of the residual gasses are minimized and substantially eliminated. In the preferred embodiment, the volume of the displacement cavity  88  is approximately ten times greater than the volume of the chamber  90 . To further reduce temperature induced drift and to increase the sensitivity of the device  12 , lower pressures within the internal volume  90  should be used. 
     In addition to the preferred embodiment, other configurations for the sensing device  12  are possible. Depending on the relative sizes of the diaphragm  64  and coil  42 , the diaphragm  64  may be located within, above, or below the turns of the coil  42 , as well as off to one end or side of the device  12  as seen in FIG.  3 . The post  40  and/or one of the plates  36  or  38  of the magnetic core  33 , may be omitted to simplify fabricating. However, this would be to the detriment of performance. Alternate lead transfer schemes may be used instead of the disclosed traces  82  and  84  that connect the coil  42  to the sensor  18 . More or fewer wafer layers may be used to adapt manufacturing processing to available technologies. For example, the entire magnetic core  33  could be formed on the top side of the substrate  20 , thereby eliminating the need for lower cap layer  46 . Multiple coil layers could also be implemented to increase the coil turn count. Finally, the overall shape of the device  10  may be square, round, oval, or another shape. 
     To isolate the internal volume of the pressure sensor  18  from the internal volume of the integrated inductor  16 , a hermetic lead transfer can be provided as a substitute for the dielectric layer  86 . A hermetic lead transfer would eliminate outgassing from the inductor coil  42  and magnetic core  33  as a source of drift for the pressure sensor  18 , thereby improving long-term stability. The hermetic lead transfer may be accomplished by any of several means that provide a sealed and electrically isolated conductive path. One example, of a mechanism for achieving a sealed and electrically isolated conductive path is through the use of a p-n junction structure  92  in the sensor  18 ′. This is illustrated in FIG.  9 . The p-n junction structure  92  (with p-material forming the diaphragm) forms an electrically isolated path in a silicon layer and provides for electrical contact between a fixed electrode  66 ′ and a lead trace  94  but not from the fixed electrode  66 ′ to the diaphragm  66 ′. 
     In another alternative construction, a separate polysilicon layer  96  forms a conductive path to a fixed electrode  66 ″. The conductive layer  96  is insulated, by a separate insulating layer  98 , from the doped silicon rim  100  of the sensor  18 ″. 
     An alternative embodiment of the present sensing device, designated as  12 ″, includes active circuitry for immediate processing of the data including logging, error correction, encoding, analysis, multiplexing of multiple sensor inputs, etc. Since the sensing device  12 ″ of this embodiment, seen in  FIG. 11 , includes numerous structures which are the same or identical to the structures seen in the embodiment illustrated in  FIG. 3 , like structures are accordingly provided with like designations and are not repetitively discussed. Reference should therefore be accordingly made to the preceding sections of this description where those structures are discussed in connection with FIG.  3 . 
     The block diagram of  FIG. 12  illustrates one possible circuit implementation for the active circuitry  102  seen in FIG.  11 . In the illustrated configuration, the integrated inductor  16  serves as an antenna for RF telemetry with the external readout device  14 . Using RF modulation schemes well know to those skilled in the art, the RF magnetic field  26  transmitted from the device  14  provides both data communication and necessary power to the circuitry  102 . The received energy across inductor  16  is rectified and stored temporarily in an onboard capacitor or power supply designated at block  104 . The input decoder  103  may receive digital data pertaining to short or long term memory or real time clock signals, and may transfer this information to the control logic  107 . The front end conditioning circuitry  109  converts an analog sensor signal into a form that is encoded and amplified by the output driver  105 . The integrated inductor  16  then serves to transmit the RF signal back to the external readout device  14 , where the information can be processed, stored, or displayed. The many variations for circuit implementations of the rectifier of  104 , modulation and coding schemes encompassing blocks  103  and  105 , analog circuitry  109  and needed control logic  103  will be appreciated. 
     A key issue for sensing physiologic parameters in medical applications is that the sensor must be biocompatible. Biocompatibility involves two issues: the effect of the sensor on the body (toxicity), and the effect of the body on the sensor (corrosion rate). While the fabrication of the substrate  20  of Pyrex glass, as described in connection with  FIG. 3 , would be advantageous since Pyrex is highly corrosion resistant, additional measures must be taken to include the corrosion resistance of the silicon and other components of the sensing device  12 . One method of improving those structures of the sensing device  12  formed of silicon, such as the upper and lower cap layers  44  and  46 , is to fabricate those structures of heavily boron-doped silicon. Heavily boron-doped silicon is believed to be largely corrosion resistant and/or harmless to tissues in biologic environments. 
     Another method by which corrosion resistance of the implanted device  12  may be improved is through coating of the device  12  with titanium, iridium, Parylene (a biocompatible polymer), or various other common and/or proprietary thick and thin films. Such a coated device provides two levels of corrosion resistance: and underlying stable surface and a separate, stable coating (which may also be selectively bioactive or bioinert). Provided with these two levels of corrosion resistance, even if the outer coating contains pinholes, cracks, or other discontinuities, the device  12  retains a level of protection. 
     A number of different, and at times application-specific, schemes can be envisioned for long-term use of the sensing device  12  of the present invention. In general, it is necessary to anchor the device  12  so that migration of the device  12  does not occur within the patient. A dislodged device  12  may migrate away from the physiologic parameter intended to be sensed, thereby rendering the device  12  useless for its intended purpose and requiring implantation of another device  12 . A variety of such anchoring schemes is discussed below. 
     Referring now to  FIG. 14 , a screw (or stud)  104  is attached to the lower cap layer  46  of the sensing device  12 . Preferably, the screw  104  is attached to the lower cap layer  46  with biocompatible epoxy or a similar method. The screw  104  is then embedded into tissue  106  of the patient and the device  12  retained in place. Preferred materials for the screw  104  include stainless steel and titanium. 
     Another scheme for securing the sensing device  12  within a patient is seen in FIG.  14 . As seen therein, the sensing device  12  has secured to the lower cap layer  46  a sheet of mesh  108 . The mesh  108  becomes encapsulated by tissue of the patient over time, thus anchoring the sensing device  12 . Sutures  110  may be used to hold the sensing device  12  in place until encapsulation occurs. Preferred materials for the mesh  108  include loosely woven, biocompatible cloth and the mesh  108  may range in size from 1 to 20 mm. 
     An endoluminal attachment scheme is illustrated within FIG.  15 . In this application, sensing device  12  is attached to stent-like spring cage  112 . As such, the sensing device  12  may be non-surgically injected into a blood vessel  114  or other body cavity containing fluid flow. After ejection from the insertion apparatus (not shown), the spring cage  112  expands and lodges the sensing device  12  at the sensing location, while allowing blood (or other fluid) to continue flowing past the sensing device  12 . To expand outward, the spring cage  12  is formed so that the arms  115  thereof are resiliently biased outward. Preferred materials for the arms  115  include stainless steel or titanium. The arms  115  may also be in wire or other forms. 
     Another endoluminal attachment scheme is shown in FIG.  16 . In this embodiment, the sensing device  12  is anchored in place within vessel  114  by a set of radially outwardly expandable spring arms  116 . The spring arms  116  may be provided with depth-limited anchoring tips  118  on their ends to further secure the sensing device  12 . The arms  116  may be in wire, ribbon or other form and are biased outwardly to cause engagement of the anchoring tips  118  with the wall of the vessel  114 . Preferred materials for the arms  116  and for the anchoring tips  118  include stainless steel or titanium. 
     In  FIG. 17 , the sensing device  12  is encapsulated in a biocompatible material such as poly(methyl methacrylate), yielding a pellet-like profile designated at  120 . A recess  122  formed in the pellet  120  allows access to the movable element  64 . In addition to providing an alternate form factor that may be less mechanically irritating to tissue  124  both during and after implantation, such an embodiment may better allow the sensing device  12  to be incorporated into the body of a medical device, such as an extrusion, injection-molded part, soft rubber, or other material, that otherwise would poorly anchor to a rectangular or other geometrically shaped sensing device  12 . Obviously, encapsulation could be used to give the sensing device other profiles or form factors as well. 
     From the above, it can be seen that many applications exist for the system  10  of the present invention. Some illustrative examples of such applications are described hereafter. 
     One application of the described technology, depicted in  FIG. 18 , locates the sensing device  12  in an electrode tip  126  of an implantable stimulation lead  128 , such as a stimulation lead used for cardiac pacing. In such an arrangement, the sensing device  12  could be used with the read-out device  14  for monitoring arterial, atrial, ventricular, and/or other blood pressures. 
     In the application seen in  FIG. 19 , three sensing devices  12  are being used to calculate a diameter  130  of a flow path  132  defined by walls  134 . In addition to the diameter  130 , mass and/or volumetric blood or other fluid flow rates through the flow path  132  may be calculated. The sensing devices  12  are located in a variable diameter catheter  136  or similar geometric construction conductive to taking such measurements. Computational fluid dynamics (CFD) models and calculations utilizing the distances between the sensing devices  12  (L 1  and L 2 ) and pressure changes ΔP 1  and ΔP 2  therebetween, can be used to derive the desired parameters from suitably precise pressure data. 
     Cardiac monitoring applications can particularly benefit from the present system  10  in its various embodiments. One possibility is to locate the sensing devices  12  (either by means of a multiple-sensor catheter or individually placed sensor devices  12  (or placed as a tethered pair)) at appropriate locations around a natural or artificial heart valve or other biologic valve, to monitor the pressure on either side of, and/or the flow through, the valve. The same setup may also be used to monitor pressure along a vascular stent  137 , as shown in FIG.  20 . Sensing devices  12  may be placed at one or more locations  138 - 142  along the length of the stent. 
     Referring now to  FIG. 21 , a sensing device  12  is located such that pressure is measured externally through a vessel wall  144 , such as the wall of a blood vessel. The sensing device  12  is placed in intimate contact with the wall  144  through use of a variety of means, including adhesive clips  146  (of a biocompatible material), tissue growth or other methods. The sensing device  12  is oriented so that the moveable element  64  is adjacent the vessel wall  144  and measures pressure transduced through the vessel wall  144 . A calibration factor in active circuitry may be used to adjust the measured value to an actual value so as to account for the effects of sensing the pressure through the vessel wall  144 . 
     As an alternative to the foregoing embodiments, the pressure sensor  18  of the sensing device  12  may be augmented and/or replaced with a structure or sensor  18 ′ that allows a parameter other than pressure to be sensed. For clarity, in  FIG. 22  only the sensor  18 ′ portion of the sensing device  12  is shown, the nonillustrated elements being as previously discussed. In the sensor  18 ′, a chemical-sensitive substance  148  is placed in a confinement cavity  149  and contact with and exterior surface of sensor diaphragm  150 . Osmotic expansion of the substance  148 , in response to the concentration of a target chemical, generates a pressure on the diaphragm  150  and allowing the concentration of the chemical to be monitored. For convenience, only the substrate  20  is illustrated, the fixed electrode and associated structures be omitted. This sensor  18 ′ may optionally include cap structure  152  to restrict the expansion of the chemical sensitive substance  148  to the center of the diaphragm  150  to maximize deflection of the diaphragm  150 . A micromachined mesh, grid, or semipermeable membrane  154 , also optional and either integral to the cap or attached separately thereto, may be included to prevent the chemical sensitive substance  148  from escaping (or bulging out of) the confinement cavity  149 , and/or to prevent foreign materials from entering the cavity  149 . The mesh  154  could also exist on the molecular level, being formed of a material such as a cross-linked polymer. 
     In another alternative parameter sensing embodiment, a material with high thermal coefficient of expansion is placed between moveable and fixed electrodes in a structure otherwise constructed similar to a capacitive sensor structure, thereby forming a temperature sensor. 
       FIG. 23  illustrates an alternative capacitive sensor  156  on the substrate  20 , additional structures are omitted for clarity. In this sensor  156 , the capacitance changes due to a varying dielectric constant within the capacitive gap defined between electrodes  158  and  160 . The gap is filled with sensing substance  162  chosen such that its dielectric constant changes in response to the particular physiologic stimulus being evaluated.  FIG. 24  depicts an alternate implementation of the above embodiment, with the electrodes  158 ′ and  160 ′ and the sensing substance  162  being stacked vertically on the substrate  20 , as opposed to the lateral orientation in FIG.  23 . 
     The pressure, temperature or other data sensing technology, in its various forms, may be incorporated into an open or closed-loop therapeutic system for the treatment of medical conditions which require or benefit from regular, subcutaneous monitoring of pressures or other parameters. The system may be used, for example, to control the administration of drugs. One particular application of this would be to control hyper- or hypotension. In the preferred embodiment, pressure data from the sensor, alone or in conjunction with other real-time or preexisting data, is used to adjust drug or other therapy for hypo- or hypertensive patient. Therapy is provided by means of a control module worn by, or implanted within, the patient (similar to e.g., an insulin pump for diabetics). The module may alert the user to take action, directly administer a drug intravenously, and/or initiate other invasive or non-invasive responses. Furthermore, relevant information (including, but not limited to, measure physiologic parameters, treatment regimens, data histories, drug reservoir levels) can further be transmitted from the control module to other locations via cellular phone, wireless infrared communication protocols or other communication methods and mechanisms. 
     Other applications of the implantable wireless sensing device of this invention include, without limitation, the following: 1) Monitoring congestive heart failure patients such as left ventricle pressure monitoring, left atrium pressure monitoring and pulmonary artery pressure monitoring; 2) other hemodynamics parameters including blood pressure, blood flow velocity, blood flow volume and blood temperature; 3) diabetic applications including glucose level monitoring; 4) urinary applications such as bladder pressure and urinary tract pressure measuring; and 5) other blood parameters including O 2  saturation, pH, CO 2  saturation, temperature, bicarbonate, glucose, creatine, hematocirt, potassium, sodium, chloride; and 6) cardiac parameters including (previously discussed) valve pressure gradients and stent pressure gradients. 
     In addition to single sensor, an array of different sensors may be fabricated or assembled on one sensing device to enhance artifact removal and/or selectivity/differentiation between signals. A discussion of such a construction best details this construction. Local pressure or pH variations can add spurious signals to a pressure- or pH-based glucose sensor. To compensate for these spurious signals, adjacent pH or pressure reference sensors may be implemented to measure these environmental parameters. External sensors may also be used to compensate for factors such as atmospheric pressure. A combination of sensor arrays, fuzzy logic, look-up tables, and/or other signal-processing technologies could all be used to effect such compensation. 
     The foregoing disclosure is the best mode devised by the inventor for practicing the invention. It is apparent, however, that several variations in accordance with the present invention may be conceivable to one of ordinary skill in the relevant art. Inasmuch as the foregoing disclosure is intended to enable such person to practice the instant invention, it should not be construed to be limited thereby, but should be construed to include such aforementioned variations, and should be limited only by the spirit and scope of the following claims.