Patent Publication Number: US-6658300-B2

Title: Telemetric reader/charger device for medical sensor

Description:
FIELD OF THE INVENTION 
     The present invention relates, in general, to telemetric medical devices. More particularly, the present invention relates to a novel telemetric medical system which is capable of various medical applications including the measurement of a parameter within a patient&#39;s body, particularly an organ. One such application of the present invention is as an implantable telemetric endocardial pressure system, its associated novel components and their novel methods of use. 
     BACKGROUND OF THE INVENTION 
     In general, the use of implantable medical sensors in a patient is known. One example for an implantable sensor is disclosed in U.S. Pat. No. 4,815,469 (Cohen et al.) incorporated herein by reference. The disclosure is directed to an implantable medical sensor which determines the oxygen content of blood. The sensor includes a miniaturized hybrid circuit that includes light-emitting diode means, phototransistor means, and a substrate to which the light-emitting diode means and phototransistor means are bonded in a desired circuit configuration. The hybrid circuit is hermetically sealed within a cylindrical body made from a material that is substantially transparent to light, such as glass. Feedthrough terminals provide means for making an electrical connection with the hybrid circuit. The light-emitting diode means is driven with a stair-stepped current pulse. The purpose of the sensor is to sense the reflective properties of body fluid, such as blood, for spectrophotometric analysis. In one embodiment, the sensor is embedded within a bilumen pacemaker lead and positioned near the distal electrode of the lead so that the sensor resides within the heart when the lead is implanted within a patient, thereby allowing the sensed oxygen content of the blood within the heart to be a physiological parameter that can be used to control the pacing interval of a rate-responsive pacemaker. 
     U.S. Pat. No. 5,353,800 (Pahndorf et al.) discloses an implantable pressure sensor lead having a hollow needle adapted to be screwed into a patient&#39;s heart. The pressure sensor is supplied electrical power through conductors in the sensor. 
     There are cases where permanent positioning of the sensor is needed. One such case, for example, is disclosed in U.S. Pat. No. 5,404,877 (Nolan et al.), which is incorporated herein by reference. A leadless implantable cardiac arrhythmia alarm is disclosed which continuously assesses a patient&#39;s heart function to discriminate between normal and abnormal heart functioning and, upon detecting an abnormal condition, generates a patient-warning signal. The alarm is capable of sensing impedance measurements of heart, respiratory and patient motion and, from these measurements, generating an alarm signal when the measurements indicate the occurrence of a cardiac arrhythmia. It is important to note that the sensor uses an antenna system having a coil inductor for generating an electromagnetic field into tissue for detecting changes in impedance which relate to a physiological phenomena. For example, the size of the inductor is preselected in order to match the dimensions of the organ or structure to be measured. 
     There are also several known implantable devices that employ telemetry for transmitting or receiving data from an external device. One such device is, for example, the system disclosed in U.S. Pat. No. 6,021,352 (Christopherson et al.). The device utilizes a pressure sensor as a transducer for sensing respiratory effort of the patient. Respiratory waveform information is received by an implantable pulse generator (IPG)/simulator from a transducer and inspiration synchronous simulation is provided by the IPG. 
     One other telemetric implantable device is disclosed in U.S. Pat. No. 5,999,857 (Weijand et al.). This reference discloses a telemetry system for use with implantable devices such as cardiac pacemakers and the like, for two-way telemetry between the implanted device and an external programmer. The system employs oscillators with encoding circuits for synchronous transmission of data symbols in which the symbols form the telemetry carrier. The system provides circuits for higher density data encoding of sinusoidal symbols, including combinations of BPSK, FSK, and ASK encoding. Embodiments of transmitters for both the implanted device and the external programmer, as well as modulator and demodulator circuits, are also disclosed. It is important to note that the implant device has its own power supply in the form of a battery for powering all of the circuitry and components of the implanted device. 
     It is also important to note, that to date, there has not been any telemetric medical system that is both a highly efficient system due to its components and their ease of use while providing extremely accurate information regarding a measured parameter in a patient&#39;s body. 
     SUMMARY OF THE INVENTION 
     The present invention is directed to a novel telemetric medical system for use with various medical applications such as monitoring medical conditions or measuring parameters within a patient&#39;s body for different types of organs, including tissue, as well as their function. 
     The present invention is a telemetric medical system comprising a telemetric medical sensor for implantation in a patient&#39;s body for measuring a parameter therein. The sensor comprises a housing, and a membrane at one end of the housing, wherein the membrane is deformable in response to the parameter. A microprocessor, which is in the form of a microchip, is positioned within the housing and operatively communicates with the membrane for transmitting a signal indicative of the parameter. 
     A signal reading and charging device is locatable outside of a patient&#39;s body and communicates with the sensor. The signal reading and charging device comprises a casing and a circuit within the casing. The circuit comprises a logic control unit and a processing unit operatively connected to the logic control unit. The logic control unit, through a deep detector, receives the transmitted signal from the sensor. The logic control unit also sends a powering signal to the sensor through a sine wave driver for remotely powering the sensor. The powering signal is a sinusoidal wave signal approximately 4-6 MHz. The processing unit includes an algorithm for converting the transmitted signal received from the sensor into a measured parameter. Additionally, the signal reading and charging device includes a power source operatively connected to the circuit and a power switch for activating and deactivating the device. 
     The signal reading and charging device also includes an antenna coil for sending the powering signal to the sensor and for receiving the transmitted digital signal from the sensor. The antenna coil has inductive coupling with the sensor. The signal reading and charging device also includes a display, which is an LCD screen, for displaying the measured parameter. 
     The microprocessor, which is in the form of a microchip, comprises an array of photoelectric cells which are arranged in staggered rows. The array also includes a reference photoelectric cell located at one end of the array. A light emitting diode (LED) transmits light at the photoelectric cells and the reference photoelectric cell. 
     The sensor further comprises a shutter connected to the membrane and moveable between the photoelectric cells and the LED in response to the deforming of the membrane. The sensor is arranged such that the reference photoelectric cell is not blocked by the shutter and remains exposed to the light emitted by the LED. 
     The microchip further comprises a plurality of comparators operatively connected to the photoelectric cells and a buffer operatively connected to the comparators for storing and transmitting the digital signal. The sensor further comprises an antenna, in the form of a coil, operatively connected to the microchip wherein the antenna is located at the exterior of the housing. Alternatively, the antenna is located within the housing of the sensor. Preferably, the antenna coil is made of wire comprising silver and platinum iridium. Additionally, the antenna has 20-25 turns. 
     The sensor according to the present invention further comprises a plurality of anchoring legs resiliently attached to the housing for anchoring the sensor into tissue. Additionally, the housing optionally includes a notch in an outer surface of the housing to facilitate deployment. The housing further optionally includes a circumferential groove at the notch to further facilitate deployment. 
     In another embodiment for the sensor, the housing further includes a tapered end and a piercing tip thereon. The tapered end further includes helical threads thereon for threading the sensor housing directly into tissue. An alternative embodiment includes a plurality of tissue barbs on the tapered end for anchoring the sensor housing directly into tissue. 
     The present invention also includes a method for telemetrically measuring a parameter in a patient&#39;s body comprising the steps of providing a telemetric medical sensor comprising a housing having a membrane at one end of the housing wherein the membrane is deformable in response to the parameter, and a microchip is positioned within the housing and operatively communicates with the membrane for transmitting a signal indicative of the parameter. The sensor is implanted at a site within the patient&#39;s body and the parameter is telemetrically measured from outside of the patient&#39;s body with a signal reading and charging device. The method also includes telemetrically powering the sensor from outside of the patient&#39;s body with the signal reading and charging device. The measured parameter is then displayed on the display of the signal reading and charging device. 
     The method according to the present invention also includes a method for telemetrically measuring a parameter in a patient&#39;s heart wherein the method comprises the steps of imaging the heart, through the use of transesophageal ultrasonic imaging, and identifying an implantation site in the heart. An opening is created in the tissue at the implantation site and a sensor comprising a housing, a membrane at one end of the housing wherein the membrane is deformable in response to the parameter, and a microchip positioned within the housing and operatively communicating with the membrane for transmitting a signal indicative of the parameter is provided. The sensor is placed within the opening and the parameter is telemetrically measured from outside of the patient&#39;s body based on the transmitted signal by the sensor. 
     The method also includes telemetrically powering the sensor from outside of the patient&#39;s body and displaying the measured parameter with a signal reading and charging device. Parameter measurements are made multiple times per second with the signal reading and charging device. 
     According to the present invention, the sensor is positioned within a chamber of the heart by using the septum as an implantation site, for instance, the fossa ovalis. Alternatively, the sensor is positionable at other anatomical sites within the heart and other organs and tissue. 
     One parameter that is measured with the system and method according to the present invention is hemodynamic blood pressure in a chamber of the heart. Accordingly, the method according to the present invention further includes taking between 10-20 parameter measurements per second. 
     Moreover, the method further includes creating the opening in the tissue with a needle. In one embodiment of the present invention, the sensor includes a plurality of anchoring legs on the sensor for anchoring the sensor to the tissue. Additionally, the sensor is coated with a nonthrombogenic agent in order to prevent thrombosis within the heart upon implantation of the sensor. 
     Another embodiment of the method according to the present invention includes a method for telemetrically measuring a parameter in a patient&#39;s heart wherein the method comprises the steps of imaging the heart with transesohageal ultrasonic imaging and identifying an implantation site in the heart. A sensor comprising a housing and a membrane at one end of the housing wherein the membrane is deformable in response to the parameter and a tapered distal end and piercing tip at the other end of the housing is provided. The sensor further comprises a microchip positioned within the housing and operatively communicating with the membrane for transmitting a signal indicative of the parameter. The sensor is implanted at the site with the piercing tip and the tapered distal end of the sensor. The parameter is telemetrically measured from outside of the patient&#39;s body based on the transmission signal by the sensor. Additionally, the sensor is telemetrically powered from outside of the patient&#39;s body. A signal reading and charging device is used outside of the patient&#39;s body to measure the parameter, power the sensor, and display the measured parameter. Accordingly, parameter measurements are made multiple times per second with the signal reading and charging device. 
     The sensor is positioned within a chamber of the heart and the implantation site is the septum, for instance, at the fossa ovalis. With the system and method according to the present invention, one parameter that is measured is hemodynamic blood pressure within a chamber of the heart. For instance, 10-20 parameter measurements are made per second for monitoring blood pressure in accordance with the present invention. 
     Alternatively, the sensor includes helical threads on the tapered distal end of the sensor and the sensor is anchored into the tissue at the site by threading the tapered distal end of sensor directly into the tissue. Alternatively, the sensor includes a plurality of tissue barbs on the tapered distal end of the sensor and the sensor is anchored into the tissue at the site with the tissue barbs. 
     The present invention will be more fully understood from the following detailed description of the preferred embodiments thereof, taken together with the drawings, in which: 
    
    
     BRIEF DESCRIPTION OF THE DRAWINGS 
     FIG. 1 is a schematic illustration of a telemetric implantable medical sensor according to the present invention; 
     FIG. 2 is a top view of the sensor of FIG. 1; 
     FIG. 3 is a schematic illustration of an alternative embodiment of the sensor of FIG. 1 having a tapered distal end with helical threads and tissue piercing tip for anchoring into tissue; 
     FIG. 4 is another alternative embodiment of the sensor of FIG. 1 having a tapered distal end with tissue piercing tip and a plurality of tissue piercing barbs thereon; 
     FIG. 5 is a partial perspective view of the sensor of FIG. 1 with some parts removed in order to reveal the internal components of the sensor; 
     FIG. 6A is schematic diagram illustrating a microprocessor circuit for the sensor according to the present invention; 
     FIG. 6B is a schematic diagram illustrating a logic circuit for the microprocessor circuit of FIG. 6A; 
     FIG. 7 is a schematic illustration depicting an array of photoelectric cells for the sensor according to the present invention; 
     FIG. 8 is a schematic illustration depicting the telemetric system according to the present invention including the sensor of FIG. 1 and a signal reading and charging device remotely located from and in communication with the sensor; 
     FIG. 9 is a schematic diagram illustrating a read/charge circuit for the signal reading and charging device of FIG. 8; 
     FIG. 10 is a schematic illustration of a patient&#39;s heart; and 
     FIG. 11 is a schematic illustration depicting the sensor fully deployed within a tissue aperture according to the present invention. 
    
    
     DESCRIPTION OF THE PREFERRED EMBODIMENTS 
     The present invention relates to a novel telemetric medical system  30 , as schematically illustrated in FIG. 8, as well as its novel components and methods of use useful for various medical applications, as explained and demonstrated herein. 
     One aspect of the system  30  of the present invention is to remotely sense and measure a characteristic or parameter (or number of various parameters including the magnitude of any parameter) within a patient&#39;s body, or within an organ or tissue of the patient&#39;s body, through the use of a novel implantable telemetric medical sensor  50 , which is completely wireless, and a novel signal reading and charging device  140  which operatively communicates with the sensor  50 . 
     Telemetric Sensor 
     As schematically illustrated in FIG. 1, the sensor  50  comprises a housing  52  made of a biocompatible material such as polysilicon or titanium. The housing  52  preferably has a cylindrical shape although any type of shape for the housing  52  is acceptable. The housing  52  has an approximate length ranging between 4-5 mm and an approximate diameter ranging from 2.5-3 mm in diameter. The housing  52  can also be smaller, e.g. 3 mm in length and a 1-2 mm outer diameter. The housing  52  includes cylindrical walls that are approximately 250 μm in thickness. A flexible membrane  56  made of a deformable material is fixed to one end of the housing  52 . A notch  58  and a circumferential groove  60  are provided on an exterior surface of the housing  52  for facilitating delivery and implantation of the sensor  50 . 
     The membrane  56  is made of a flexible or deformable material such as polysilicon rubber or polyurethane. The membrane  56  has an approximate thickness of 20 μm and has a diameter ranging from approximately 1.5-2 mm. The membrane  56  is normally biased outwardly from the housing  52  due to the interior pressure within the housing  52 . The membrane  56  is forced to bulge inwardly into the housing  52  whenever the pressure exterior of the housing  52  exceeds the internal pressure within the housing  52 . 
     Since the membrane  56  is deformable and normally biased outwardly from the housing  52 , the membrane  56  responds directly to the environment of the tissue or organ being monitored and/or measured for a particular characteristic or parameter. In response to even the slightest changes in these characteristics or parameters, the membrane  56  deforms inwardly toward the interior of the housing  52 . Accordingly, there is a direct relationship or correspondence between any change in measured characteristic or parameter and the amount or degree of deforming action or movement of the membrane  56 . 
     It is important to note that the membrane  56  has a relatively large area in dimension when compared to solid state membrane devices, such as piezoelectric sensors or fabricated memory chips utilizing membranes. Accordingly, the requirements from the electronics of the sensor  50  are less demanding. Additionally, the membrane  56  has a much larger deflection than that of the solid state membrane. 
     The sensor  50  also includes an antenna coil  68  which is operatively connected to the internal components of the sensor  50  by an antenna lead  70 . The antenna coil  68  is an inductance coil having a spiralled coil configuration. The material used for the antenna wire is approximately 90% silver content with a cladding of platinum iridium of approximately 10% content. The antenna coil  68  is preferably made of 20-25 turns of 30 μm thickness wire. The antenna outer diameter is 1.5-2.0 cm (FIG.  2 ). 
     Accordingly, due to these features, the antenna coil  68  possesses a very low parasitic capacitance. Additionally, the antenna coil  68 , due to its silver/platinum content wire has extremely high conductivity and is extremely flexible. 
     Although antenna  68  is described as being external of the housing  52 , it is well within the scope of the invention to include any type of suitable antenna, such as an antenna that is contained within the housing  52 . 
     The sensor  50  further includes anchoring legs  64  resiliently biased to the exterior of the housing  52 . The number of anchoring legs  64  can vary depending on the desired degree of anchoring and geography of the anatomy in which the sensor  50  is to be placed. The anchoring legs  64  are made from wire utilizing shape memory metal material, such as a nickel titanium alloy (NiTinol). The anchoring legs  64  have a concave configuration with a radius of curvature that curves into the tissue or organ in which the sensor  50  is to be anchored. Other appropriate configurations for the anchoring legs  64  are also contemplated herein. 
     If desireable, the sensor  50  is coated with a nonthrombogenic or anticoagulating agent such as Heparin prior to implantation in order to prevent thrombosis, clotting, etc. 
     FIG. 3 illustrates an alternative embodiment of the sensor  50  having a tapered end  54  on the housing  52 . The tapered end  54  has a tissue piercing tip  55  and helical threads  57  arranged on an outer surface of the tapered end  54  in order to facilitate the direct anchoring of the tapered end  54  of the housing  52  through direct threading into tissue. 
     FIG. 4 illustrates another alternative embodiment sensor  50  including a plurality tissue barbs  59  fixed to the tapered end  54  of the housing  52 . The barbs  59  have a tissue piercing tip curved outwardly away from the tissue piercing tip  55 . Accordingly, along with the tissue piercing tip  55 , the tissue barbs  59  grasp firmly into the tissue for firmly anchoring the housing  52  in the tissue. 
     As shown in FIG. 5, the interior of the housing  52  includes a microprocesser  90 , in the form of a microchip, fixed within one of the interior walls of the housing  52 . The lead  70  of the antenna coil  68  is operatively connected to the microprocessor  90 . Microprocessor  90  includes an array  92  of photoelectric cells  95  arranged in a patterned confirguration, e.g. eight staggered rows containing eight photoelectric cells  95  in each row. A reference photoelectric cell  97  is located at one end of the array  92  resulting in an array  92  having a total of sixty-five photoelectric cells such as illustrated in FIG.  7 . The photoelectric cell array  92  provides for 64 degrees of resolution. The pitch distance between each photocell  95  is approximately ¼ the size of a photocell  95 . Additionally, the reference photocell  97  has a dimension that is approximately the size of the pitch, e.g. ¼ the size of a photocell  95 , thus providing a resolution that is equal to a motion of ¼ of the photocell. 
     A light emitting diode (LED)  100  is operatively connected to the microprocessor  90  and is positioned above and spaced parallel and away from the photoelectric cell array  92 . A shutter  62  is connected to the inner surface of the membrane  56  and extends logitudinally from the membrane  56  within housing  52 . The shutter  62  has a substantially D-shaped configuration and logitudinally extends between the LED  100  and the photoelectric cell array  92 . The shutter  62  is made from an aluminum alloy and is positioned such that the planar surface of the shutter  62  directly faces the photoelectric cell array  92 . The shutter  62  is fixed to the deformable membrane  56  such that the shutter  62  moves in association with the membrane  56 . Accordingly, when the membrane  56  is deflected inwardly into the housing  52  (due to the monitored or measured tissue or organ parameter), the shutter  62  logitudinally extends over a number of photoelectric cells  95  in the array  92  in direct relation to the inward movement of the membrane  56  as it is being deformed. Likewise, when the membrane  56  is deflected outwardly from the housing  52 , the shutter  62  moves logitudinally outwardly from the end of the housing  52  along with the membrane  56 . Accordingly, the shutter  62  obscures or blocks a number of the photoelectric cells  95  in accordance with the degree of movement of the membrane  56 . Thus, when the shutter  62  is positioned over a specific number of photoelectric cells  95 , light from the LED  100  is prevented from reaching the photoelectric cells  95  and affects signal transmission from these cells  95 . This arrangement constitutes an analog-to-digital (A/D) conversion which is power effective since there is a simple counting of the number of photocells that are on or off as a measure of the shutter&#39;s motion. Hence, the analog-to-digital conversion. Accordingly, the microprocessor  90  operatively communicates with the membrane  56 . 
     The reference photoelectric cell  97  is never obscured or covered by the shutter  62  since it is located at the far end (end away from the membrane  56 ) of the array  92 . The shutter  62  and membrane  56  are calibrated such that even upon maximum deflection inwardly into the housing  52 , it results in the reference photoelectric cell  97  being permanently exposed to the LED  100  for use as a reference signal for the sensor  50 . Yet, the power dissipation of the photocell is very low. 
     As best shown in FIG. 6A, the microprocessor  90  is a circuit wherein the antenna coil  68  and a resonance capacitor  102  operate as a resonating oscillator for the sensor  50 . The antenna coil  68  receives transmitted RF signals sent by the signal reading and charging device  140  as illustrated in FIGS. 8 and 9. The RF signal received at the antenna coil  68  is a charging signal for powering the microprocessor  90 . Upon receiving the RF charging signal, the antenna coil  68  and capacitor  102  resonate and charge a charge capacitor  114  through diode  116 . Upon reaching a predetermined voltage threshold of approximately 1.2 V, the capacitor  114  powers the LED  100  and a logic circuit  91  through control unit  104 . Upon powering of the LED  100  by the charged capacitor  114 , the LED emits light to the photoelectric cell array  92  which is kept at negative voltage. 
     As illustrated in FIG. 6B, the photoelectric cell array  92  is designated P 1 , P 2 , . . . P 64  and P ref , respectively. Each photoelectric cell  95  (P 1 -P 64 ) are connected in parallel to a plurality of comparators  120  designated C 1 , C 2  . . . C 64 . The reference photoelectric cell  97  is operatively connected to each comparator  120  (C 1 -C 64 ) for providing a reference signal to each comparator  120  in comparison to the signal received from each respective photoelectric cell  95 . The logic circuit  91  is powered and controlled by the control unit  104  and a clock  106 . The control unit  104  is connected to each comparator  120 . 
     A buffer  126  having a plurality of buffer cells  129  (sixty-four total buffer cells corresponding to each comparator C 1 -C 64 ) is operatively connected to the comparators  120 . Each buffer cell  129  is a flip-flop, or memory cell, which receives a signal from its respective comparator C 1 -C 64  resulting in a binary number which is sixty-four digits long (a series of ones or zeros). All buffer cells  129  are filled in a single clock cycle and each buffer  129  has either “0” or “1” in it. After all sixty-four buffer cells  129  have been filled with its respective binary number, the digital signal representing all sixty-four bytes is sent to the signal reading and charging device  140  by the control unit  104 . After transmitting the digital signal, the control unit  104  is reset by the clock  106  awaiting further signal inputs from the signal reading and charging device  140 . Encryption of the binary number is provided by the signal reading and charging device  140  described in greater detail below. 
     Upon filling the sixty-fourth buffer cell, the digital signal is transmitted from the buffer  126  and activates switch  112  resulting in a transmission of the digital signal from the antenna coil  68  to the antenna coil  162  of the signal reading and charging device  140 . 
     One main aspect of the system  30  of the present invention is that the sensor  50  is both a wireless transponder and a low-powered device capable of fast update rate, despite its passive nature, due to the inherent analog-to-digital (A/D) conversion mechanism employed in the sensor  50 , e.g. the photoelectric cell array  92 , which directly converts the membrane  56  deflection into a digital signal, with no power consumption as would be required for a conventional electronic A/D converter. 
     Signal Reading and Charging Device 
     As illustrated in FIG. 8, the signal reading and charging device  140  according to the present invention is for use outside of a patient&#39;s body or at the exterior surface of the patient&#39;s body. The signal reading and charging device  140  includes a casing  145 , which is a housing, having a liquid crystal display (LCD) display screen  172  mounted in an opening in the housing  145 . The signal reading and charging device, also commonly referred to as a read/charge device, reader/charger or reader/charger device, is activated by a power switch or toggle  146  extending from the casing  145 . Antenna coil  162  operatively communicates with the antenna coil  68  of the sensor  50  by inductance coupling. 
     As shown in FIG. 9, once the logic circuit  91  transmits the digital signal from the sensor  50  through sensor antenna coil  68 , the coupling constant of the reader/charger antenna coil  162  is changed and is detected by a deep detector  168  operatively connected to the reader/charger antenna coil  162 . The deep detector  168  is sensitized to detect a change in the amplitude of the signal for as low as a 0.01% change in amplitude. 
     A read/charge logic control unit  154  is operatively connected to the deep detector  168  for determining the threshold for the deep detector  168 . The logic control unit  154  also includes a power source  151  for powering the components of the reader/charger device  140 . 
     The reader/charger circuit  150  further includes a processing unit  170  operatively connected to the logic control unit  154 . The processing unit  170  contains the algorithm for converting the digital signal received from the sensor  50  (FIG. 8) into a measured parameter for the medical parameter, condition or characteristic sensed at the implanted sensor  50 . Additionally, the processing unit  170  includes encryption code for encryption of the digital signal (sixty-four bit signal) using encryption algorithms such as exclusive-OR (XOR), RSA methods (RSA Security, Inc.), etc. 
     For example, where the parameter being measured is hemodynamic blood pressure, within an organ such as the chamber of a heart, once the processing unit  170  receives the digital signal, the processing unit  170 , through its algorithm, converts the digital signal (binary number) to a pressure value, using a look-up comparison table, or analytical expression representing the relation between the shutter  62  deflection in the sensor  50  versus the exterior sensor pressure at the membrane  56 , which is given below: 
     
       
           P =( KD   3   /A   2 ) X   2    
       
     
     where P is the pressure value, D is the thickness of the membrane, A is the membrane radius, X is the deflection from the equilibrium and K is a constant. 
     The LCD display  172  is operatively connected to the processing unit  170  for displaying the measured parameter (hemodynamic blood pressure in the example above) converted from the digital signal in real time. 
     By utilizing the signal reading and charging device  140  at the exterior of the patient&#39;s body, continuous parameter readings (for determining aspects of the parameter such as magnitude) are obtainable for both the mean and active or individual values of the sampled parameter. 
     When measuring characteristics of a body fluid such as blood, the signal reading and charging device  140  maintains an active reading volume around the sensor  50 , ranging anywhere from 5-25 cm, and preferably, an active reading volume ranging approximately 10-15 cm. Moreover, with the telemetric medical system  30 , through the sensor  50 , and the signal reading and charging device  140 , it is possible to sample multiple readings per second. Preferably, approximately 10-20 readings per second are possible with the present invention. 
     Other attributes associated with the present invention when utilized as a pressure monitor in a chamber of the heart include monitoring a pressure range of +/−30 mmHg; an accuracy (at 5 mSec. integration) of +/−1 mmHg with a repeatability (at 5 mSec. integration) of +/−1 mmHg. It is important to note that the pressure boundaries can be changed easily by changing the size and dimensions, such as width, of the membrane without any change to the electronics. This is important for allowing the present invention to be adapted for various applications while using the same design. 
     The control unit  154  is also operatively connected to a sine-wave driver  158  for generating a sinusoidal wave signal of approximately 4 to 6 MHz. The sinusoidal wave signal is generated by the sine-wave driver  158  through capacitor  160  to the reader/charger antenna coil  162  for transmission or sending to the antenna coil  68  of the sensor  50  in order to power or charge the sensor  50  as described above. 
     Medical Procedures 
     As mentioned above, the telemetric medical system  30  according to the present invention is useful for nearly any type of medical diagnostic procedure where it is desireable to implant the sensor  50  at a portion of the body, particularly tissue or organ of interest. The telemetric medical system  30  according to the present invention allows for remote monitoring and diagnosis of a condition of the tissue or organ by being able to rapidly sample various parameters or variables of any physical condition within the patient&#39;s body at the site of interest. Since the telemetric medical system  30  is wireless, these types of procedures are conducted in a completely non-invasive manner with minimal trauma to the patient. 
     One particular example for the telemetric medical system  30  according to the present invention, its components and their method of use, is in the field of congestive heart failure (CHF). CHF is defined as a condition in which a heart  400  (FIG. 10) fails to pump enough blood to the body&#39;s other organs. This can result from narrowed arteries that supply blood to the heart muscle (due to coronary artery disease), past heart attack, or myocardial infarction, with scar tissue that interferes with the heart muscle&#39;s normal work, high blood pressure, heart valve disease due to past rheumatic fever (in valves such as semilunar valve, tricuspid valve  417  or mitral valve  418 ) or other causes, primary disease of the heart muscle itself, called cardiomyopathy, defects in the heart present at birth such as congenital heart disease, infection of the heart valves and/or heart muscle itself (endocarditis and/or myocarditis). 
     The ailing heart  400  keeps functioning but not as efficiently as it should. People with CHF cannot exert themselves because they become short of breath and tired. As blood flowing out of the heart  400  slows, blood returning to the heart  400  through the veins backs up, causing congestion in the tissues. Often swelling (edema) results, most commonly in the legs and ankles, but possibly in other parts of the body as well. Sometimes fluid collects in the lungs and interferes with breathing, causing shortness of breath, especially when a person is lying down. Heart failure also affects the ability of the kidneys to dispose of sodium and water. The retained water increases the edema. 
     CHF is the most common heart disease in the United States and it is estimated that over 5 million patients suffer from it. One of the more predictive hemodynamic parameters being measured in patients with CHF is blood pressure in the left atrium  410 , e.g. left atrial (LA) pressure. To date, this parameter is measured by employing invasive right heart catheterization with a special balloon catheter such as the Swan-Gantz catheter. 
     Accordingly, in moderating for effects of CHF, it is desireable to measure the blood pressure in a particular chamber (either right atrium  415 , right ventricle  419 , left atrium  410  or left ventricle  420 ) in the heart  400  utilizing the telemetric medical system  30  according to the present invention. 
     Accordingly, in conducting one preferred method according the present invention, blood pressure can be directly monitored in the left atrium  410  of the heart  400 . Accordingly, it is desireable to implant the sensor  50  at fossa ovalis  407  within the septum  405 . 
     With respect to the specific anatomy of the septum  405 , in approximately 15% of the normal population, the fossa ovalis  407  has a pre-existing hole or opening that either remains open or patent and is normally covered by a small flap of tissue. In approximately 85% of the normal population, the fossa ovalis  407  is completely occluded, e.g. there is no hole in the septum  405 . 
     (1) Transcatheter Approach 
     In accordance with the method according to the present invention, a transcatheter approach has been found to be particularly useful for the patient population already having the pre-existing hole at the fossa ovalis  407 . Accordingly, in performing this method according to the present invention, first, a transesophageal ultrasonic probe (not shown) is inserted into the patient&#39;s mouth and placed in the esophagus. In most cases, the transesophageal ultrasonic probe is positioned approximately 30-35 cm from the mouth, i.e. in most cases positioned just above the patient&#39;s stomach. 
     Under transesophageal ultrasonic guidance, a wire (not shown) is inserted into the right atrium  415  through an appropriate vessel such as the inferior vena cava  408  wherein the wire is guided through the fossa ovalis  407  by gently lifting the tissue flap away from the patent opening at the fossa ovalis  407 . Once the wire is inserted through the fossa ovalis  407 , the wire is guided to one of the pulmonary veins  416  for placement of the distal end of the wire in order to properly position and anchor the wire in the opening of the pulmonary vein  416 . Accordingly, the pulmonary vein  416  has been proven to be a very reliable and steady anchoring point for the wire. 
     Once the wire is properly positioned in the fossa ovalis  407  and anchored in the pulmonary vein  416 , a catheter sheath (“over-the-wire” type —not shown) is guided over the wire through the right atrium  415  and the fossa ovalis  407  and positioned within the left atrium  410 , for instance, very close to the opening of the pulmonary vein  416 . 
     Once the catheter sheath has been properly positioned, the wire is removed from the patient&#39;s heart  400  and the sensor  50  is delivered through the catheter sheath by one of the many standard catheter-based delivery devices (not shown). Accordingly, the sensor  50  can be delivered to the fossa ovalis  407  by any of the typical catheter-based delivery devices normally associated with implantable pacemakers, electrodes, atrial septal defect (ASD) occlusion devices, etc. Accordingly, the sensor  50  is deliverable with typical delivery devices such as the Amplatzer® Delivery System, manufactured by AGA Medical Corporation of Golden Valley, Minn. 
     After placement of the catheter sheath, the sensor  50  is deployed from the catheter sheath within the fossa ovalis  407  as best illustrated in FIG.  11 . Upon deployment, the sensor  50  utilizes the anchoring legs  64  for anchoring the sensor  50  to the septum  405  and occluding the opening at the fossa ovalis  407 . 
     (2) Anterograde Approach 
     The sensor  50  is placed in the fossa ovalis  407  for those patients not having a pre-existing opening in the fossa ovalis  407  through means of an anterograde approach. Once again, a transesophageal ultrasonic probe is positioned in the patient&#39;s esophagus as described above. Under transesophageal ultrasonic imaging guidance, an opening is made in the septum  405  at the fossa ovalis  407  in order to place and accommodate the sensor  50 . Thus, the opening is made with a standard needle catheter (not shown) such as the BRK™ Series Transseptal Needle manufactured by St. Jude Medical, Inc. of St. Paul, Minn. Accordingly, under transesophageal ultrasonic guidance, the needle catheter is initially placed in the right atrium  415  and positioned at the fossa ovalis  407 . At this point, the tip of the needle of the needle catheter penetrates the fossa ovalis  407  and the catheter is inserted through the fossa ovalis  407  into the left atrium  410  through the newly created opening in the fossa ovalis  407  by the needle catheter. Once the opening in the fossa ovalis  407  is created, the sensor  50  is introduced with the delivery device, such as the delivery device described above, and placed in the fossa ovalis opening as shown in FIG.  11 . Upon deployment of the anchoring legs  64 , the opening in the fossa ovalis  407  is occluded around the sensor housing  52  and the sensor  50  fixed to the septum  405  in a secure fashion. 
     It is important to note that transesophageal ultrasonic imaging is utilized for both the transcatheter and the anterograde approach as described above in accordance with each method step of the present invention. Since either method according to the present invention can be utilized with the transesophageal ultrasonic guidance, other imaging modalities such as flouroscopy can be eliminated. As such, the methods according to the present invention can be conducted in an outpatient clinic or doctor&#39;s office as a bedside procedure. By eliminating the need for a flouroscope, the method according to the present invention also eliminates the need for conducting the procedure in a catheter lab which only adds additional time and cost to the procedure and additional time and inconvenience to the patient. 
     After the sensor  50  has been implanted in the patient&#39;s septum  405 , the patient is provided with standard treatment to prevent excessive coagulation or endothelialization. For instance, it is common practice to prescribe aspirin and/or an anticoagulant such as Heparin for a period of time such as six months. 
     With either of the methods described above, the sensor  50  is fixed to the septum  405  in order to provide real time pressure monitoring in the left atrium  410 . Since the sensor  50  is a wireless transponder and a battery low power receiver, the sensor  50  does not impede the natural function of the heart  400  and is truly minimally invasive. 
     By utilizing the signal reading and charging device  140  at the exterior of the patient&#39;s body, continuous pressure readings are obtainable for both the mean and pulsating values of pressure in the left atrium  410  provided by the sensor  50 . 
     With the telemetric system  30 , the signal reading and charging device  140  maintains an active reading volume around the sensor  50  ranging anywhere from 5-25 cm, and preferably, an active reading volume ranging approximately 10-15 cm. Moreover, with the sensor  50 , and the signal reading and charging device  140 , it is possible to sample multiple readings per second. Preferably, approximately 10-20 readings per second are possible with the present invention. 
     Other attributes associated with the present invention when utilized as a pressure monitor in a chamber of the heart include monitoring a pressure range of plus/minus 30 mmHg; and accuracy (at five Mmsec. integration) of plus/minus 1 mmHg and a repeatability (at 5 msec. integration) of plus/minus 1 mmHg. 
     Although preferred embodiments are described hereinabove with reference to a medical system, devices, components and methods of use, it will be understood that the principles of the present invention may be used in other types of objects as well. The preferred embodiments are cited by way of example, and the full scope of the invention is limited only by the claims.