Abstract:
An external device, charge, system and method for an implantable medical device having therapeutic componentry, a secondary coil operatively coupled to the therapeutic componentry and an internal telemetry coil. A primary coil is capable of inductively energizing the secondary coil when externally placed in proximity of the secondary coil. An external telemetry coil is capable of communicating with the internal telemetry coil. Driver circuitry is selectively operatively coupled to the primary coil and to the external telemetry coil. The driver circuitry is switchable between (1) driving the primary coil for inductively energizing the secondary coil and (2) driving the external telemetry coil for communicating with the internal telemetry coil.

Description:
FIELD OF THE INVENTION 
     This invention relates to implantable medical devices and, in particular, to energy transfer devices, systems and methods for implantable medical devices. 
     BACKGROUND OF THE INVENTION 
     Implantable medical devices for producing a therapeutic result in a patient are well known. Examples of such implantable medical devices include implantable drug infusion pumps, implantable neurostimulators, implantable cardioverters, implantable cardiac pacemakers, implantable defibrillators and cochlear implants. Of course, it is recognized that other implantable medical devices are envisioned which utilize energy delivered or transferred from an external device. 
     A common element in all of these implantable medical devices is the need for electrical power in the implanted medical device. The implanted medical device requires electrical power to perform its therapeutic function whether it be driving an electrical infusion pump, providing an electrical neurostimulation pulse or providing an electrical cardiac stimulation pulse. This electrical power is derived from a power source. 
     Typically, a power source for an implantable medical device can take one of two forms. The first form utilizes an external power source that transcutaneously delivers energy via wires or radio frequency energy. Having electrical wires which perforate the skin is disadvantageous due, in part, to the risk of infection. Further, continuously coupling patients to an external power for therapy is, at least, a large inconvenience. The second form utilizes single cell batteries as the source of energy of the implantable medical device. This can be effective for low power applications, such as pacing devices. However, such single cell batteries usually do not supply the lasting power required to perform new therapies in newer implantable medical devices. In some cases, such as an implantable artificial heart, a single cell battery might last the patient only a few hours. In other, less extreme cases, a single cell unit might expel all or nearly all of its energy in less than a year. This is not desirable due to the need to explant and re-implant the implantable medical device or a portion of the device. One solution is for electrical power to be transcutaneously transferred through the use of inductive coupling. Such electrical power or energy can optionally be stored in a rechargeable battery. In this form, an internal power source, such as a battery, can be used for direct electrical power to the implanted medical device. When the battery has expended, or nearly expended, its capacity, the battery can be recharged transcutaneously, via inductive coupling from an external power source temporarily positioned on the surface of the skin. 
     Several systems and methods have been used for transcutaneously inductively recharging a rechargeable used in an implantable medical device. 
     U.S. Pat. No. 5,411,537, Munshi et al, Rechargeable Biomedical Battery Powered Devices With Recharging and Control System Therefor, (Intermedics, Inc.) discloses a hermetically-sealed automatic implantable cardioverter-defibrillator (AICD) or any other bioimplantable device which may be operated on a single rechargeable cell, or a dual power source system, the rechargeable complement being recharged by magnetic induction. Included in the implantable devices are lithium rechargeable chemistries designed to sense the state-of-charge or discharge of the battery; a battery charge controller specifically designed to recharge a lithium battery rapidly to less than 100% full charge, and preferably 90%, more preferably 80%, of full rated charge capacity; and charging means for multi-step charging. The batteries are based on lithium chemistries specially designed to yield higher currents than conventional primary lithium chemistries and to permit long-term performance despite sub-capacity recharging. 
     U.S. Pat. No. 5,690,693, Wang et al, Transcutaneous Energy Transmission Circuit For Implantable Medical Device, (Sulzer Intermedics Inc.) discloses a transcutaneous energy transmission device for charging rechargeable batteries in an implanted medical device. A current with a sinusoidal waveform is applied to a resonant circuit comprising a primary coil and a capacitor. Current is induced in a secondary coil attached to the implanted medical device. Two solid-state switches are used to generate the sinusoidal waveform by alternately switching on and off input voltage to the resonant circuit. The sinusoidal waveform reduces eddy current effects in the implanted device which detrimentally increases the temperature of the implanted device. The batteries are charged using a charging protocol that reduces charging current as the charge level in the battery increases. The controller is constructed as a pulse with modulation device with a variable duty cycle to control the current level applied to the primary coil. An alignment indicator is also provided to insure proper and alignment between the energy transmission device and the implanted medical device. 
     U.S. Pat. No. 5,733,313, Barreras, Sr., FR Coupled Implantable Medical Device With Rechargeable Back-Up Power Source, (Exonix Corporation) discloses an implantable, electrically operated medical device system having an implanted radio frequency (RF) receiving unit (receiver) incorporating a back-up rechargeable power supply and an implanted, electrically operated device, and an external RF transmitting unit (transmitter). RF energy is transmitted by the transmitter and is coupled into the receiver which is used to power the implanted medical device and/or recharge the back-up power supply. The back-up power supply within the receiver has enough capacity to be able to, by itself, power the implanted device coupled to the receiver for at least 24 hours during continual delivery of medical therapy. The receiver is surgically implanted within the patient and the transmitter is worn externally by the patient. The transmitter can be powered by either a rechargeable or non-rechargeable battery. In a first mode of operation, the transmitter will supply power, via RF coupled energy, to operate the receiver and simultaneously recharge the back-up power supply. In a second mode of operation, the receiver can, automatically or upon external command from the transmitter, acquire its supply of power exclusively from the back-up power supply. Yet, in a third mode of operation, the receiver can, automatically or upon command from the transmitter, alternatively acquire it supply of power from either, FR energy coupled into the receiver or the internal back-up power supply. 
     U.S. Pat. No. 6,308,101, Faltys et al, Fully Implantable Cochlear Implant System, (Advanced Bionics Corporation) discloses a fully implantable cochlear implant system and method including an implantable cochlear stimulator unit that is connected to an implantable speech processor unit. Both the speech processor unit and the cochlear stimulator unit are in separate, hermetically-sealed, cases. The cochlear stimulator unit has a coil permanently connected thereto through which magnetic or inductive coupling may occur with a similar coil located externally during recharging, programming, or externally-controlled modes of operation. The cochlear stimulator unit further has a cochlear electrode array permanently connected thereto via a first multi-conductor cable. The cochlear stimulator unit also has a second multi-conductor cable attached thereto, which second cable contains no more than five conductors. The second cable is detachably connected to the speech processor unit. The speech processor unit includes an implantable subcutaneous microphone as an integral part thereof, and further includes speech processing circuitry and a replenishable power source, e.g., a rechargeable battery. 
     U.S. Pat. No. 6,324,430, Zarinetchi et al, Magnetic Shield For Primary Coil of Transcutaneous Energy Transfer Device, (Abiomed, Inc.) discloses a transcutaneous energy transfer device which has a magnetic shield covering the primary winding of the device to reduce sensitivity of the device to conducting objects in the vicinity of the coils and to increase the percentage of magnetic field generated by the primary coil which reaches the secondary coil. The shield is preferably larger than the primary coil in all dimensions and is either formed of a high permeability flexible material, for example a low loss magnetic material and a flexible polymer matrix, with perforations formed in the material sufficient to permit ventilation of the patient&#39;s skin situated under the shield, or the shield may be formed of segments of a very high permeability material connected by a flexible, porous mesh material. 
     U.S. Pat. No. 6,516,227, Meadows et al, Rechargeable Spinal Cord Stimulator System, (Advanced Bionics Corporation) discloses a spinal cord stimulation system providing multiple stimulation channels, each capable of producing up to 10 milliamperes of current into a one kilohm load. The system further includes a replenishable power supply, e.g., a rechargeable battery that requires only an occasional recharge, and offers a life of at least 10 years at typical settings. The replenishable power source may be replenished using non-invasive means. The system monitors the state of charge of the internal power source and controls the charging process by monitoring the amount of energy used by the system, and hence the state of the charge of power source. A suitable bidirectional telemetry link allows the system to inform the patient or clinician regarding the status of the system, including the state of the charge, and makes requests to initiate an external charge process. 
     U.S. Pat. No. 6,505,077, Kast et al, Implantable Medical Device With External Recharging Coil Electrical Connection, (Medtronic, Inc.) discloses a rechargeable implantable medical device with an improved external recharging coil electrical connection resistant to corrosion. The electrical connection couples the external recharging coil to a recharge feedthrough. The rechargeable implantable medical device can be a medical device such as a neuro stimulator, drug delivery pump, pacemaker, defibrillator, diagnostic recorder, cochlear implant, and the like. The implantable medical device has a housing, electronics carried in the housing configured to perform a medical therapy, a rechargeable power source, and a recharging coil. 
     European Patent Application 1,048,324, Schallhorn, Medical Li+ Rechargeable Powered Implantable Stimulator, (Medtronic, Inc.) discloses an implantable stimulator having a rechargeable lithium ion power source and delivers electrical stimulation pulses, in a controlled manner, to a targeted site within a patient. The lithium ion power source can supply sufficient power to the implantable stimulator on an exclusive basis over at least about 4 days. The power source includes a high value, small size lithium ion storage unit having a power rating of at least 50 milliamp hours. The implantable stimulator also has an inductor adapted to gather EMF power transmissions. The implantable stimulator can be replenished with electrical power by an electrical power replenisher, external to the implantable stimulator, to replenish the lithium ion power source up to its maximum rated voltage by generating the EMF power transmission near the inductor. 
     PCT Patent Application No. WO 01/83029 A1, Torgerson et al, Battery Recharge Management For an Implantable Medical Device, (Medtronic, Inc.) discloses an implantable medical device having an implantable power source such as a rechargeable lithium ion battery. The implantable medical device includes a recharge module that regulates the recharging process of the implantable power source using closed-loop feedback control. The recharging module includes a recharge regulator, a recharge measurement device monitoring at least one recharge parameter, and a recharge regulation control unit for regulating the recharge energy delivered to the power source in response to the recharge measurement device. The recharge module adjusts the energy provided to the power source to ensure that the power source is being recharged under safe levels. 
     PCT Patent Application No. WO 01/97908 A2, Jimenez et al, An Implantable Medical Device With Recharging Coil Magnetic Shield, (Medtronic, Inc.) discloses a rechargeable implantable medical device with a magnetic shield placed on the distal side of a secondary recharging coil to improve recharging efficiency. The rechargeable implantable medical device can be wide variety of medical devices such as neurostimulators, drug delivery pumps, pacemakers, defibrillators, diagnostic recorders, and cochlear implants the implantable medical device has a secondary recharging coil carried over a magnetic shield and coupled to electronics and a rechargeable power source carried inside the housing electronics are configured to perform a medical therapy. Additionally a method of for enhancing electromagnetic coupling during recharging of an implantable medical device is disclosed, and a method for reducing temperature rise during recharging of an implantable medical device is disclosed. 
     Transcutaneous energy transfer through the use of inductive coupling involves the placement of two coils positioned in close proximity to each other on opposite sides of the cutaneous boundary. The internal coil, or secondary coil, is part of or otherwise electrically associated with the implanted medical device. The external coil, or primary coil, is associated with the external power source or external charger, or recharger. The primary coil is driven with an alternating current. A current is induced in the secondary coil through inductive coupling. This current can then be used to power the implanted medical device or to charge, or recharge, an internal power source, or a combination of the two. 
     For implanted medical devices, the efficiency at which energy is transcutaneously transferred is crucial. First, the inductive coupling, while inductively inducing a current in the secondary coil, also has a tendency to heat surrounding components and tissue. The amount of heating of surrounding tissue, if excessive, can be deleterious. Since heating of surrounding tissue is limited, so also is the amount of energy transfer which can be accomplished per unit time. The higher the efficiency of energy transfer, the more energy can be transferred while at the same time limiting the heating of surrounding components and tissue. Second, it is desirable to limit the amount of time required to achieve a desired charge, or recharge, of an internal power source. While charging, or recharging, is occurring the patient necessarily has an external encumbrance attached to their body. This attachment may impair the patient&#39;s mobility and limit the patient&#39;s comfort. The higher the efficiency of the energy transfer system, the faster the desired charging, or recharging, can be accomplished limiting the inconvenience to the patient. Third, amount of charging, or recharging, can be limited by the amount of time required for charging, or recharging. Since the patient is typically inconvenienced during such charging, or recharging, there is a practical limit on the amount of time during which charging, or recharging, should occur. Hence, the size of the internal power source can be effectively limited by the amount of energy which can be transferred within the amount of charging time. The higher the efficiency of the energy transfer system, the greater amount of energy which can be transferred and, hence, the greater the practical size of the internal power source. This allows the use of implantable medical devices having higher power use requirements and providing greater therapeutic advantage to the patient and/or extends the time between charging effectively increasing patient comfort. 
     Prior art implantable medical devices, external power sources, systems and methods have not always provided efficiency of operational componentry. 
     BRIEF SUMMARY OF THE INVENTION 
     In one embodiment, the present invention provides, an external device for an implantable medical device having therapeutic componentry, a secondary coil operatively coupled to the therapeutic componentry and an internal telemetry coil. A primary coil is capable of inductively energizing the secondary coil when externally placed in proximity of the secondary coil. An external telemetry coil is capable of communicating with the internal telemetry coil. Driver circuitry is selectively operatively coupled to the primary coil and to the external telemetry coil. The driver circuitry is switchable between (1) driving the primary coil for inductively energizing the secondary coil and (2) driving the external telemetry coil for communicating with the internal telemetry coil. 
     In another embodiment, the present invention provides, a charger for an implantable medical device having a rechargeable power source, therapeutic componentry operatively coupled to the rechargeable power source, a secondary coil operatively coupled to the therapeutic componentry and an internal telemetry coil. A primary coil is capable of inductively energizing the secondary coil for charging the implantable medical device when externally placed in proximity of the secondary coil. An external telemetry coil is capable of communicating with the internal telemetry coil. Driver circuitry is selectively operatively coupled to the primary coil and to the external telemetry coil. The driver circuitry is switchable between (1) driving the primary coil for inductively energizing the secondary coil and (2) driving the external telemetry coil for communicating with the internal telemetry coil. 
     In another embodiment, the present invention provides a system for transcutaneous energy transfer. An implantable medical device has therapeutic componentry for producing a therapeutic output, a secondary coil operatively coupled to therapeutic componentry and an internal telemetry coil operatively coupled to the therapeutic componentry. An external power source has a primary coil capable of inductively energizing the secondary coil when externally placed in proximity of the secondary coil, an external telemetry coil capable of communicating with the internal telemetry coil and driver circuitry selectively operatively coupled to the primary coil and to the external telemetry coil. The driver circuitry is switchable between (1) driving the primary coil for inductively energizing the secondary coil and (2) driving the external telemetry coil for communicating with the internal telemetry coil. 
     In another embodiment, the present invention provides a method of transcutaneous energy transfer to a medical device implanted in a patient having therapeutic componentry, a secondary coil operatively coupled to the therapeutic componentry and an internal telemetry coil. A primary coil is positioned externally of the patient in proximity of the secondary coil. The secondary coil is energized with the primary coil via driver circuitry. An external telemetry coil is positioned externally of the patient in proximity of the internal telemetry coil. Communication is established between the external telemetry coil and the internal telemetry coil via the driver circuitry. The driver circuitry is switched between (1) driving the primary coil for inductively energizing the secondary coil and (2) driving the external telemetry coil for communicating with the internal telemetry coil. 
     In a preferred embodiment, the driver circuitry operates (1) at a relatively low frequency when switched for driving the primary coil for inductively energizing the secondary coil and (2) at a relatively high frequency when switched for driving the external telemetry coil for communicating with the internal telemetry coil. 
     In a preferred embodiment, the implantable medical device has a rechargeable power source operatively coupled to the therapeutic componentry and wherein the secondary coil charges the rechargeable power source when driven by the primary coil. 
    
    
     
       BRIEF DESCRIPTION OF THE DRAWINGS 
         FIG. 1  illustrates an implantable medical device implanted in a patient; 
         FIG. 2  is a block diagram of an implantable medical device; 
         FIG. 3  is a detailed block diagram of an implantable medical device implanted sub-cutaneously and an associated external charging device in accordance with an embodiment of the present invention; 
         FIG. 4  is a perspective view of an internal antenna associated with an implantable medical device; 
         FIG. 5  is a side view of the internal antenna of  FIG. 4 ; 
         FIG. 6  is an exploded perspective view an external antenna and associated bracket in accordance with an embodiment of the present invention; 
         FIG. 7  is a top view of an external antenna in accordance with an embodiment of the present invention; 
         FIG. 8  is a perspective view of an external antenna and bracket combination in accordance with an embodiment of the present invention; 
         FIG. 9  is a cross-sectional side view of an implantable medical device implanted sub-cutaneously and an associated bracket for use with an external antenna; 
         FIG. 10  is a cut-away top view of view a primary coil and associated magnetic core in accordance with an embodiment of the present invention; 
         FIG. 11  is a cross-sectional view of the primary coil and associated magnetic core of  FIG. 10  taken through section line B-B; 
         FIG. 12  is an exploded view a portion of an external antenna constructed in accordance with an embodiment of the present invention showing the magnetic core and a core cup assembly; 
         FIG. 13  is block diagram of an external charging unit and an associated inductively coupled cradle for recharging the external charging unit; 
         FIG. 14  is a detailed block diagram of the external charging unit of  FIG. 13 ; 
         FIG. 15  is a flow chart illustrating a charging process in accordance with an embodiment of the present invention; and 
         FIG. 16  is a schematic diagram of a dual range temperature sensor. 
     
    
    
     DETAILED DESCRIPTION OF THE INVENTION 
       FIG. 1  shows implantable medical device  16 , for example, a drug pump, implanted in patient  18 . The implantable medical device  16  is typically implanted by a surgeon in a sterile surgical procedure performed under local, regional, or general anesthesia. Before implanting the medical device  16 , a catheter  22  is typically implanted with the distal end position at a desired therapeutic delivery site  23  and the proximal end tunneled under the skin to the location where the medical device  16  is to be implanted. Implantable medical device  16  is generally implanted subcutaneously at depths, depending upon application and device  16 , of from 1 centimeter (0.4 inches) to 2.5 centimeters (1 inch) where there is sufficient tissue to support the implanted system. Once medical device  16  is implanted into the patient  18 , the incision can be sutured closed and medical device  16  can begin operation. 
     Implantable medical device  16  operates to infuse a therapeutic substance into patient  18 . Implantable medical device  16  can be used for a wide variety of therapies such as pain, spasticity, cancer, and many other medical conditions. 
     The therapeutic substance contained in implantable medical device  16  is a substance intended to have a therapeutic effect such as pharmaceutical compositions, genetic materials, biologics, and other substances. Pharmaceutical compositions are chemical formulations intended to have a therapeutic effect such as intrathecal antispasmodics, pain medications, chemotherapeutic agents, and the like. Pharmaceutical compositions are often configured to function in an implanted environment with characteristics such as stability at body temperature to retain therapeutic qualities, concentration to reduce the frequency of replenishment, and the like. Genetic materials are substances intended to have a direct or indirect genetic therapeutic effect such as genetic vectors, genetic regulator elements, genetic structural elements, DNA, and the like. Biologics are substances that are living matter or derived from living matter intended to have a therapeutic effect such as stem cells, platelets, hormones, biologically produced chemicals, and the like. Other substances may or may not be intended to have a therapeutic effect and are not easily classified such as saline solution, fluoroscopy agents, disease diagnostic agents and the like. Unless otherwise noted in the following paragraphs, a drug is synonymous with any therapeutic, diagnostic, or other substance that is delivered by the implantable infusion device. 
     Implantable medical device  16  can be any of a number of medical devices such as an implantable therapeutic substance delivery device, implantable drug pump, cardiac pacemaker, cardioverter or defibrillator, as examples. 
     In  FIG. 2 , implantable medical device  16  has a rechargeable power source  24 , such as a Lithium ion battery, powering electronics  26  and therapy module  28  in a conventional manner. Therapy module  28  is coupled to patient  18  through one or more therapy connections  30 , also conventionally. Rechargeable power source  24 , electronics  26  and therapy module  28  are contained in hermetically sealed housing  32 . Secondary charging coil  34  is attached to the exterior of housing  32 . Secondary charging coil  34  is operatively coupled through electronics  26  to rechargeable power source  24 . In an alternative embodiment, secondary charging coil  34  could be contained in housing  32  or could be contained in a separate housing umbilically connected to electronics  26 . Electronics  26  help provide control of the charging rate of rechargeable power source  24  in a conventional manner. Magnetic shield  36  is positioned between secondary charging coil  34  and housing  32  in order to protect rechargeable power source  24 , electronics  26  and therapy module  28  from electromagnetic energy when secondary charging coil  34  is utilized to charge rechargeable power source  24 . 
     Rechargeable power source  24  can be any of a variety power sources including a chemically based battery or a capacitor. In a preferred embodiment, rechargeable power source is a well known lithium ion battery. 
       FIG. 3  illustrates an alternative embodiment of implantable medical device  16  situated under cutaneous boundary  38 . Implantable medical device  16  is similar to the embodiment illustrated in  FIG. 2 . However, charging regulation module  42  is shown separate from electronics  26  controlling therapy module  28 . Again, charging regulation and therapy control is conventional. Implantable medical device  16  also has internal telemetry coil  44  configured in conventional manner to communicate through external telemetry coil  46  to an external programming device (not shown), charging unit  50  or other device in a conventional manner in order to both program and control implantable medical device and to externally obtain information from implantable medical device  16  once implantable medical device has been implanted. Internal telemetry coil  44 , rectangular in shape with dimensions of 1.85 inches (4.7 centimeters) by 1.89 inches (4.8 centimeters) constructed from 150 turns of 43 AWG wire, is sized to be larger than the diameter of secondary charging coil  34 . Secondary coil  34  is constructed with 182 turns of 30 AWG wire with an inside diameter of 0.72 inches (1.83 centimeters) and an outside diameter of 1.43 inches (3.63 centimeters) with a height of 0.075 inches (0.19 centimeters). Magnetic shield  36  is positioned between secondary charging coil  34  and housing  32  and sized to cover the footprint of secondary charging coil  34 . 
     Internal telemetry coil  44 , having a larger diameter than secondary coil  34 , is not completely covered by magnetic shield  36  allowing implantable medical device  16  to communicate with the external programming device with internal telemetry coil  44  in spite of the presence of magnetic shield  36 . 
     Rechargeable power source  24  can be charged while implantable medical device  16  is in place in a patient through the use of external charging device  48 . In a preferred embodiment, external charging device  48  consists of charging unit  50  and external antenna  52 . Charging unit  50  contains the electronics necessary to drive primary coil  54  with an oscillating current in order to induce current in secondary coil  34  when primary coil  54  is placed in the proximity of secondary coil  34 . Charging unit  50  is operatively coupled to primary coil by cable  56 . In an alternative embodiment, charging unit  50  and antenna  52  may be combined into a single unit. Antenna  52  may also optionally contain external telemetry coil  46  which may be operatively coupled to charging unit  50  if it is desired to communicate to or from implantable medical device  16  with external charging device  48 . Alternatively, antenna  52  may optionally contain external telemetry coil  46  which can be operatively coupled to an external programming device, either individually or together with external charging unit  48 . 
     As will be explained in more detail below, repositionable magnetic core  58  can help to focus electromagnetic energy from primary coil  46  to more closely be aligned with secondary coil  34 . Also as will be explained in more detail below, energy absorptive material  60  can help to absorb heat build-up in external antenna  52  which will also help allow for a lower temperature in implantable medical device  16  and/or help lower recharge times. Also as will be explained in more detail below, thermally conductive material  62  is positioned covering at least a portion of the surface of external antenna  52  which contacts cutaneous boundary  38  of patient  18 . 
     In a preferred embodiment of internal antenna  68  as shown in  FIG. 4  and  FIG. 5 , secondary coil  34  and magnetic shield  36  are separate from but adjacent to housing  32  encompassing the remainder of implantable medical device  16 . Internal antenna  68  is contained in a separate housing  74  which is attachable to housing  32  so that implantable medical device  16  can be implanted by a medical professional as essentially one unit. Secondary coil  34  is electrically attached to charging regulation module  42  through leads  82 . 
     In order to achieve efficient inductive coupling between primary coil  54  of external antenna  52  and secondary coil  34 , it is desirable to place primary coil  54  of external antenna  52  as close to secondary coil  34  as possible. Typically, external antenna  52  is placed directly on cutaneous boundary  38  and, since the location of implantable medical device  16  is fixed, the distance across cutaneous boundary  38  between primary coil  54  and secondary coil  34  is minimized as long as external antenna  52  is kept adjacent cutaneous boundary  38 . 
     In a preferred embodiment, external antenna  52  is attachable to patient  18  with bracket  84  when charging rechargeable power source  24 .  FIG. 6  is an exploded illustration of a preferred embodiment of external antenna  52  attachable to bracket  84 . Primary coil  54  is contained in bobbin assembly  86  which sits in bottom housing  88 . Primary coil is connectable to cable  56 . The bottom of external antenna  52  is formed from a thermally conductive material  90 . Rotating core cup assembly  92  is held in place by top housing  94 . Rotating core cup assembly  92  is rotatable is allowed to rotate within external antenna  52 . Detents  96  engage detent spring  98  to position rotatable core cup assembly  92  in one of a plurality of detent positions. External antenna may be secured together, for example, with screws (not shown) holding top housing  94  and thermally conductive material  90  together. 
     Bracket  84  is adapted to be attached to the body of patient  18  with a belt (not shown) attachable to bracket  84  with belt loops  102 . Ears  104  are adapted to mate with tabs  106  in top housing  94  and pivotally secure external antenna  52  in bracket  84  when charging is to be accomplished. Bracket  84  has an opening  108  allowing thermally conductive material  90  of external antenna  52  to contact the skin of patient  18  when external antenna  52  is pivotally secured in bracket  84 . 
     As bracket  84  is attached to patient  18  with a belt via belt loops  102 , the skin surface of patient  18  is typically not completely flat. For example, if implantable medical device  16  is implantable in the body torso of patient  18 , then the belt attached via belt loops  102  will typically pass around the torso of patient  18 . Since the torso of patient  18 , and especially the torso of patient  18  near the location of implantable medical device  16 , bracket  84  may not sit completely flat on patient  18 . This may be especially true as patient  18  moves and the torso flexes during such movement. It is preferred that bracket  84  be conformal and flexible in order to conform to the shape of the body of patient  18 . However, it is also preferred that bracket  84  be rigid enough so that opening  108  in bracket  84  maintain it shape in order to properly receive external antenna  52 . Bracket  84  is preferably constructed of PCABS. To maintain the proper position of bracket  84  with the skin of patient  18 , the surface of bracket  84  closest to patient  18  contains material  109  constructed from a high durometer, e.g., 40 Shore A, or “sticky” material such as a material known under the tradename of “Versaflex” manufactured by GLS Corp. of McHenry, Ill. This will help external antenna to sit more closely to the skin surface of patient  18  and remain there during movements of patient  18  throughout the charge or recharge cycle. In addition, external antenna  52  is allowed to pivot by way of ears  104  on tabs  106 . Bracket  84  is configured to allow thermally conductive material  90  to extend through opening  108  and contact the skin surface of patient  18 . Allowed pivoting of external antenna  52  and, hence, thermally conductive material  90 , permits thermally conductive surface to sit more closely to the skin surface of patient  18 . 
       FIG. 7  is a partially cut away top view of external antenna  52  is assembled form and attached to cable  56 . Rotatable core cup assembly  92  is shown located inside of primary coil  54  and positionable in selected rotated positions via detents  96  and detent spring  98 . In  FIG. 7 , rotatable core cup assembly is positioned between with detent spring  98  between detents  96  illustrating that while multiple detent positions are available, rotatable core cup assembly can be positioned between detent positions and, indeed, at any rotated position. 
     In  FIG. 8 , the assembly of external antenna  52  with bracket  84  is shown connected to cable  56 . It is preferred that bracket  84  be affixed to patient  18  through belt loops  102  and then, after bracket  84  has been affixed to patient  18 , external antenna  52  be attached to bracket  84 . Affixing bracket  84  to patient  18  first allows for bracket  84  to be used to laterally position external antenna close to the position of implantable medical device  16 . 
     Typical prior art positioning systems rely on the external antenna for lateral positioning. The external antenna is moved around on the body of the patient  18  until the best lateral position is found. When the best lateral position is found, the external antenna is removed from the body and the bottom of the external antenna (the portion of the external antenna) contacting the patient&#39;s body) is made to be resistant to lateral movement. As an example, one way is to remove a protective liner exposing a sticky surface allowing the external antenna to be relatively fixed in location. However, the very act of lifting the external antenna in order to remove the protective liner and replacing the external antenna on the body of the patient  18  causes crucial positioning information to be lost. There is no guarantee, and in fact it is not likely, that the external antenna will be replaced in the exact same position as the position previously found to be best. 
     In contrast, bracket  84  of the present invention can be used to roughly find the optimum position for external antenna  52 . This can be done relatively easily due to opening  108  in bracket  84 . Implantable medical device  16 , when implanted, usually leaves an area of the body of patient  18  which is not quite as flat as it was before implantation. That is, implantable medical device  16  usually leaves an area of the skin of patient  18  which bulges somewhat to accommodate the bulk of implantable medical device  16 . It is relatively easy for patient, medical professional or other person, to place bracket  84  in the general area of implantable medical device  16  and move bracket  84  around until the bulge caused by implantable medical device  16  is most closely centered in opening  108 . As bracket  84  is moved laterally, opening  108  tends to naturally center on the bulge created by implantable medical device  16 . Once positioned in this manner, bracket  84  can be secured to the body of patient  18  with belt (not shown) attached via belt loops  102 . Securing and/or tightening, by pulling the belt tight or snapping a buckle, for example, can be without removing bracket  84  from the body of patient  16 . Thus, bracket  84  can be relatively easily positioned over the general location of implantable medical device  16  and secured in that position without be removed from the body of patient  18 . 
       FIG. 9  is cross-sectional view of implantable medical device  16  implanted in patient  18  approximately one centimeter under cutaneous boundary  38  creating bulging area  110 , an area of the body of patient  18  in which the skin of patient  18  is caused to bulge slightly due to the implantation of implantable medical device  16 . Bulging area  110  is an aid to locating the position of external antenna  52  relative to secondary coil  34 . Bracket  84  can be positioned roughly in the area where implantable medical device  16  is implanted. Opening  108  in bracket  84  can aid is establishing the location of implantable medical device. Bracket  84  can be roughly centered over bulging area  110 . After external antenna  52  is coupled to bracket  84 , then primary coil  54  can be generally centered on implantable medical device  16 . 
     However, secondary coil  34  may not be centered with respect to implantable medical device  16 . This can occur due to a variety of reasons such as the need for operatively coupling secondary coil  34  to charging regulation module  42 . Connections to make this operative coupling may require physical space on one side of internal antenna  68  which may cause secondary coil  34  not to be centered on implantable medical device  16 . It is also possible that the attachment of internal antenna  68  to housing  32  can cause secondary coil  34  not to be centered on implantable medical device  16 . Regardless of the cause, if secondary coil  34  is not centered on implantable medical device  16 , then centering bracket  84  on bulging area  110  may not optimally position primary coil  54  with respect to secondary coil  34 . Any offset in the position of primary coil  54  and secondary coil  34  may not result in the most efficient energy transfer from external antenna  52  to implantable medical device  16 . 
     A magnetic core  58  is positioned within primary coil  54  in order to focus energy generated by primary coil  54 . Magnetic core  58  attracts the magnetic flux lines generated by primary coil  54 . The position of magnetic core  58  within primary coil  54  the lateral location of the largest amount of the flux lines generated by primary coil  54 .  FIGS. 10 and 11  show cut-away top and cross-sectional views of magnetic core  58  used with primary coil  54 . Magnetic core  58  is moveable within primary coil  54 . Lower portion  122  of magnetic core  58  can be rotated to a plurality of positions within primary coil  58  by rotating core cup assembly  92  (see  FIG. 12 ). In a preferred embodiment, the travel path of magnetic core  58  can be locked in a plurality of discrete positions. In a preferred embodiment, magnetic core  58  is locked in four (4) different positions by detents  96  and detent spring  98  (see  FIG. 6 ). Magnetic core  58  has an upper planar portion  120  and a smaller lower portion  122 . 
     As magnetic core  58  is repositioned within primary coil  54 , the focus of magnetic flux generated by primary coil  54  is also repositioned. As noted above, external antenna  52  is generally aligned with implanted medical device  16  using palpatory sensation. Moveable magnetic core  58  can then be used to provide a “fine” adjustment to the lateral positioning of external antenna  52  with respect to secondary coil  34 . After bracket  84  has been secured to patient  18 , external antenna  52  is attached to bracket  84 . Magnetic core  58  is then moved until the best lateral alignment with secondary coil  34 . 
     Magnetic core  58  is shown positioned within external antenna  52  of  FIG. 12 . Core cup assembly  92  holds magnetic core  58  within the assembly of external antenna  52 . Lower portion  122  (not visible in  FIG. 12 ) of magnetic core  58  fits into recess  124  of core cup assembly  92  while upper portion  120  of magnetic core  58  rests upon ledge  126  of core cup assembly  92 . Preferably, magnetic core  58  is a ferrite core. Still more preferably, magnetic core  58  is constructed from MN60LL high performance, low loss ferrite manufactured by Ceramic Magnetics, Inc., Fairfield, N.J. Magnetic core  58  has an initial permeability of 6,500 and a maximum permeability of 10,500 (typical) with a volume resistivity of 500 ohm-centimeters. 
     One surface, preferably the top, of magnetic core  58  is lined with an adhesive coated foam  127  and contained in core cup assembly  92 . Magnetic core  58  has a tendency to be brittle. Containing magnetic core  58  is core cup assembly assures that even if magnetic core  58  has one or more fractures, magnetic core  58  will still be properly positioned and continue to function. Foam  127  also helps to hold magnetic core  58  together and minimize gaps between fractured segments of magnetic core  58 . Further, foam  127  adds mechanical stability to magnetic core  58  helping to cushion magnetic core  58  against mechanical impacts, such as from dropping external antenna  52  against a hard surface, and helps to prevents audible rattles which may otherwise develop from a fractured magnetic core  58 . 
     As shown in  FIG. 13 , external charging device  48  can be powered either directly from internal (to charging unit  50 ) batteries  160  or indirectly from desktop charging device  162 . Desktop charging device is connectable via power cord  164  to a source of AC power, such as a standard readily available wall outlet. Desktop charging device  162  can be configured as a cradle which can receive charging unit  50 . Other forms of connection from desktop charging device  162  to a power source, such as by a dedicated line cable can also be utilized. Desktop charging device  162  can charge and/or recharge batteries  160  in charging unit  50 , preferably by inductive coupling using coil  167  positioned in desktop charging device  162  and coil  168  positioned within charging unit  50 . Once charged and/or recharged, batteries  160  can provide the power through internal circuitry  168  and cable  56  to external antenna  52 . Since charging unit  50  is not, in a preferred embodiment, coupled directly to the line voltage source of AC power, charging unit  50  may be used with external antenna  52  to transfer power and/or charge implanted medical device  16  while desktop charging device  162  is coupled to a line voltage source of AC power. The inductive coupling using coil  167  and coil  168  break the possibility of a direct connection between the line voltage source of AC power and external antenna  52 . Batteries  160  also allow charging unit  50  and, hence, external charging device  48 , to be used in transferring power and/or charging of implanted medical device  16  while completely disconnected from either a line voltage source of AC power or desktop charging device  162 . This, at least in part, allows patient  18  to be ambulatory while transferring power and/or charging implanted medical device  16 . 
       FIG. 14  is a block diagram of external charging device  48  controlled by microprocessor  212 . Transmit block  214  consists of an H-bridge circuit powered from 12 volt power supply  216 . Transmit block  214  drives primary coil  54  in external antenna  52 . H-bridge control signals and timing are provided conventionally by microprocessor  212 . H-bridge circuit in transmit block  214  is used to drive both primary coil  54 , used for power transfer and/or charging, and telemetry antenna  218 . Drive selection is done by electronically controllable switch  220 . During power transfer and/or charging, H-bridge circuit is driven at 9 kiloHertz. During telemetry, H-bridge circuit is driven at 175 kiloHertz. 
     Receive block  222  is used only during telemetry, enabled by switch  224 , to receive uplink signals from implanted medical device  16 . Twelve volt power supply  216  is a switching regulator supplying power to transmit block  214  during power transfer and/or charging as well as telemetry downlink. Nominal input voltage to 12 volt power supply  216  is either 7.5 volts from lithium ion batteries  226  or 10 volts from desktop charging device  162  ( FIG. 13 ). 
     Current measure block  226  measures current to 12 volt power supply  216 . Current measured by current measure block  226  is used in the calculation of power in along with the voltage of batteries  160 . As noted above, power in is used in the calculation of efficiency of power transfer and/or charging efficiency to determine, in part, the best location of external antenna  52  and/or rotating core cup assembly  92 . 
     Rotating core cup assembly  92  is rotated in external antenna  52  for better lateral alignment of primary coil  54  and secondary coil  34 . A feedback mechanism is used to determine the best rotation of core cup assembly  92 . External charging device  48  can determine whether the current position of rotating core cup assembly  92  is optimally aligned for energy transfer and/or charging. External charging device  48  measures the power out of external charging device  48  divided by the power into external charging device  48 . This calculation is a measure of the efficiency of external charging device  48 . The power out is gauged by the power induced in implantable medical device  16  and is determined by multiplying the voltage of rechargeable power source  24  by the charging current in implantable medical device  16 . These values are obtained by telemetry from implanted medical device  16 . The power in is gauged by the power generated by charging unit  50  and is determined by multiplying the voltage of the internal voltage of charging unit  50 , e.g., the voltage of a battery or batteries internal to charging unit  50 , by the current driving external antenna  52 . 
     The ratio of power out divided by power in can be scaled displayed to patient  18 , or a medical professional or other person adjusting rotatable core cup assembly  92  or positioning external antenna  52 . For example, the available efficiency can be divided into separate ranges and displayed as a bar or as a series of lights. The separate ranges can be linearly divided or can be logarithmic, for example. 
     Using efficiency as a measure of effective coupling and, hence, as a measure of proper location of external antenna  52  and rotatable core cup assembly  92  works not only at high charging or power transfer levels but also at reduced charging levels, as for example, when charging at reduced levels toward the end or beginning of a charging cycle. 
     If, after patient  18  or other person has moved rotatable core cup assembly  92  through all of the range of positions on external antenna  52  and can not achieve an acceptable efficiency level, patient  18  or other person can remove external antenna  52  from bracket  84 , realign bracket  84  with bulging area  110 , reattach external antenna  52  to bracket  84  and restart the alignment and coupling efficiency process. 
       FIG. 15  is a flow chart illustrating an exemplary charging process using external antenna  52 . The process starts [block  126 ] and a charging session begins [block  128 ] with a test [block  130 ]. The charging system performs start-up checks [block  132 ]. If the start-up checks are not performed successfully, the actions taken in Table 1 are performed. 
     
       
         
               
               
               
             
           
               
                   
                 TABLE 1 
               
               
                   
                   
               
               
                   
                 Check 
                 Screen/Message 
               
               
                   
                   
               
             
             
               
                   
                 System Errors: e.g., stuck key 
                 System Error 
               
               
                   
                 External Charger Battery Status 
                 Recharge Complete 
               
               
                   
                   
                 Battery Low 
               
               
                   
                   
                 Recharge External Charger 
               
               
                   
                 External Charger Connected 
                 Recharge in Process Icon 
               
               
                   
                 to External Antenna 
               
               
                   
                 Antenna Disconnect 
                 Connect Antenna 
               
               
                   
                   
               
             
          
         
       
     
     If the start-up checks are successful, telemetry with implantable medical device  16  is checked [block  134 ]. If telemetry is successful, the error messages indicated in Table 2 are generated. 
     
       
         
               
               
               
             
           
               
                   
                 TABLE 2 
               
               
                   
                   
               
               
                   
                 Failure 
                 Screen/Message 
               
               
                   
                   
               
             
             
               
                   
                 Poor Communication 
                 Reposition Antenna 
               
               
                   
                 External Charger Error Code Response 
                 Call Manufacturer 
               
               
                   
                 Communication Error 
                 Communication Error 
               
               
                   
                 External Charger Fault 
                 Call Manufacturer 
               
               
                   
                 Antenna Disconnect 
                 Connect Antenna 
               
               
                   
                 Antenna Failure 
                 Antenna Failure Icon 
               
               
                   
                   
               
             
          
         
       
     
     If telemetry checks are successful, external charging device  48  is able to monitor [block  136 ] charging status. Monitoring charging status can includes providing feedback to an operator to help determine the best rotational position of core cup assembly  92 . 
     Charge events are checked [block  138 ]. If no charge events are noted, the actions indicated in Table 3 are executed. 
     
       
         
               
               
             
           
               
                 TABLE 3 
               
               
                   
               
               
                 Event 
                 Screen/Message 
               
               
                   
               
             
             
               
                 Telemetry Failure 
                 (See Messages From Table 2) 
               
               
                 Implantable Medical Device Battery Low 
                 Device Battery Low 
               
               
                 External Charger Battery Low 
                 Charger Battery Low 
               
               
                 External Charger Battery Depleted 
                 Recharge Charger 
               
               
                 External Charger Recharge Complete 
                 External Charger 
               
               
                   
                 Recharge Complete 
               
               
                 Implantable Medical Device Will 
                 Recharge Device 
               
               
                 Not Provide Therapeutic Result 
               
               
                 Until Recharged: Therapy 
               
               
                 Unavailable/Sleep Mode 
               
               
                 Antenna Disconnect 
                 Connect Antenna 
               
               
                   
               
             
          
         
       
     
     If a charge event occurs, then the process checks to determine if charging is complete [block  140 ]. Once charging is complete, the process terminates [block  142 ]. 
     As energy is transferred from primary coil  54  of external antenna  52  to secondary coil  34  of implantable medical device  16 , heat may also be generated in implantable medical device  16  in surrounding tissue of patient  18 . Such heat build-up in tissue of patient  18 , beyond certain limits, is undesirable and should be limited as acceptable values. Generally, it is preferable to limit the temperature of external antenna  52  to not more than forty-one degrees Centigrade (41° C.) and to limit the temperature of implanted medical device  16  and the skin of patient  18  to thirty-nine degrees Centigrade (39° C.). In order to ensure that implantable medical device  16  is less than the upper limit of thirty-nine degrees Centigrade (39° C.), it is preferred that the actual temperature of external antenna  52  be less than thirty-nine degrees Centigrade (39° C.). In general, the temperature of external antenna  52  should be maintained to be less than or equal to the desired maximum temperature of implanted medical device  16 . While the temperature limits discussed above are preferred under current conditions and regulations, it is recognized and understood that conditions and regulations may change or be different in different circumstances. Accordingly, the actual temperatures and temperature limits may change. In a preferred embodiment, such temperature limits are under software control in charging unit  50  so that any such temperatures or temperature limits can be modified to fit the then current circumstances. 
     Magnetic shield  36  serves to at least partially protect the portion of implantable medical device  16  contained within titanium housing  32  from the effects of energy transfer from external charging device  48  produced through inductive coupling from primary coil  54 . Magnetic shield  36  is constructed of Metglas magnetic alloy 2714A (cobalt-based) manufactured by Honeywell International, Conway, S.C. Magnetic shield  36  is positioned between secondary coil  34  and housing  32  of implantable medical device  16  with secondary coil  34  facing cutaneous boundary  38 . Magnetic shield does not interfere with the operation of secondary coil  34  because magnetic shield  36  is positioned away from primary coil  54 . Also, magnetic shield does not interfere with telemetry between implantable medical device  16  and an external programmer because magnetic shield  36  is smaller than internal telemetry coil  44 . That is, internal telemetry coil  44  lies outside of magnetic shield  36 . 
     However, the material of magnetic shield  36  substantially limits the electromagnetic energy induced by primary coil  54  from penetrating beyond magnetic shield. Electromagnetic waves induced by primary coil  54  that reach titanium housing  32  will tend to be absorbed by titanium housing  54  and its components and will tend to cause the temperature of titanium housing  54  to rise. As the temperature of titanium housing  54  rises, such temperature increase will be disadvantageously transferred to the surrounding tissue of patient  18 . However, any electromagnetic waves which are prevented from reaching titanium housing  32  will not cause such a temperature rise. 
     Thermally conductive material  62  of external antenna  52  is positioned to contact the skin of patient  18  when external antenna  52  is placed for energy transfer, or charging, of implanted medical device  16 . Thermally conductive material  62  tends to spread any heat generated at the skin surface and spread any such heat over a larger area. Thermally conductive material  62  tends to make the temperature of the skin surface more uniform than would otherwise be the case. Uniformity of temperature will tend to limit the maximum temperature of any particular spot on the skin surface. The skin itself is a pretty good conductor of heat and initially spreading any heat generated over a larger area of the skin will further assist the skin in dissipating any heat build-up on to surrounding tissue and further limit the maximum temperature of any particular location on the surface of the skin. 
     Thermally conductive material  62  is molded into the surface of external antenna  52  which will contact the skin surface of patient  18  when external antenna  52  provides energy transfer to implanted medical device  16 . Since thermally conductive material  62  should pass electromagnetic energy from primary coil  54 , thermally conductive material  62  should be constructed from a non-magnetic material. It is desirable that thermally conductive material  62  have a thermal conductivity of approximately 5.62 BTU inch/hour feet 2  degrees Fahrenheit (0.81 W/meters degrees Kelvin). In a preferred embodiment, thermally conductive material is constructed from a proprietary composite of approximately forty percent (40%) graphite, seven percent (7%) glass in RTP 199×103410 A polypropylene, manufactured by RTP Company, Winona, Minn. It is also preferable that thermally conductive material not be electrically conductive in order to reduce eddy currents. In a preferred embodiment, thermally conductive material has a volume resistivity of approximately 10 3  ohm-centimeters and a surface resistivity of 10 5  ohms per square. 
     Energy absorptive material  62  is placed in and/or around primary coil  54  of external antenna  52  in order to absorb some of the energy generated by primary coil  54 . In a preferred embodiment, energy absorptive material  62  fills in otherwise empty space of rotating core cup assembly  92 . Heat generated by energy produced by primary coil  54  which is not effectively inductively coupled to secondary coil  34  will tend to cause a temperature rise in other components of external antenna  52  and, possibly, the skin of patient  18 . At least a portion of this temperature rise can be blocked through the use of energy absorptive material  62 . Energy absorptive material  62  is chosen to absorb heat build-up in surrounding components and tend to limit further temperature increases. Preferably, energy absorptive material  62  is selected to be material which undergoes a state change at temperatures which are likely to be encountered as the temperature of surrounding components rises during energy transfer, e.g., charging, using external antenna  52 . 
     If it is a goal to limit the temperature of the skin of patient  18  to thirty-nine degrees Centigrade (39° C.), it is desirable to use of energy absorptive material  62  which has a state change at or near the temperature limit. In this example, the use of an energy absorptive material  62  having a state change in temperature area just below thirty-nine degrees Centigrade (39° C.), preferably in the range of thirty-five degrees Centigrade (35° C.) to thirty-eight degrees Centigrade (38° C.), can help limit the rise in the temperature of the skin of patient  18  to no more than the desired limit, in this example, thirty-nine degrees (39° C.). 
     As the temperature of surrounding components of external antenna  52  rise to a temperature which is just below the temperature at which energy absorptive material  62  changes state, at least a portion of further heat energy generated by primary coil  54  and surrounding components of external antenna  52  will go toward providing the energy necessary for energy absorptive material  62  to change state. As energy absorptive material  62  is in the process of changing state, its temperature is not increasing. Therefore, during the state change of energy absorptive material  62 , energy absorptive material  62  is serving to at least partially limit a further rise in the temperature of components of external antenna  52 . As the state change temperature of energy absorptive material has been preferably selected to be near or just below the temperature limit of the skin of patient  18 , energy absorptive material  62  will tend to limit the temperature components of external antenna  52  from reaching the temperature limit and, hence, will also tend to limit the temperature of the skin of patient  18  from reaching the maximum desired temperature limit. 
     In a preferred embodiment, energy absorptive material  62  is constructed from wax and, in particular, a wax which has change of state temperature of approximately the maximum temperature at which external antenna  52  is desired to reach, such as thirty-eight (38) or thirty-nine (39) degrees Centigrade. Thus, it is preferred that the wax material of which energy absorptive material is constructed melt at that temperature. 
     Inductive coupling between primary coil  54  of external antenna  52  and secondary coil of implantable medical device  16  is accomplished at a drive, or carrier, frequency, f carrier , in the range of from eight (8) to twelve (12) kiloHertz. In a preferred embodiment, the carrier frequency f carrier , of external antenna  54  is approximately nine (9) kiloHertz unloaded. 
     However, the inductive coupling between primary coil  54  of external antenna  52  and secondary coil  34  of implantable medical device is dependent upon the mutual inductance between the devices. The mutual inductance depends upon a number of variables. Primary coil  54  is preferably made from a coil of wire that has an inductance L and a series or parallel tuned capacitance C. The values of both inductance L and capacitance C are fixed. For instance, if the desired drive frequency, f carrier , of the energy transfer system was to be 1 megaHertz and external antenna  52  had an independence of one microHenry, capacitance would be added so that the resonant frequency of the energy transfer system would equal that of the drive frequency, f carrier . The total capacitance added can be found using the equation f resonate  equals one divided by two times pi (π) times the square root of L times C where L is the inductance of the energy transfer system. In this example, the value of capacitance C required to tune external antenna  52  to resonate at the carrier frequency of 1 megaHertz is calculated as approximately 25 nanofarads. 
     However, when the electrical properties of external antenna  52  change, either by the reflected environment or due to a physical distortion or change in the composition of the external antenna  52 , the inductance, L, may be altered. The inductance, L, can be altered because it is made up of two separate parts. The first part is the self-inductance, L self , of external antenna  52  at f carrier . The second part is the mutual inductance, L mutual , which is a measure of the change in current driving external antenna  52  and the magnetic effect, or “loading”, which the environment has on external antenna  52 . When the electrical characteristics of the environment of external antenna  52  change, L self  remains constant while L mutual  varies. The effect of a change in the overall inductance, whether that change is from L self  or from L mutual , is a change in the resonant frequency, f resonate . Since C was chosen in order to have the resonant frequency, f resonate , match the drive frequency, f carrier , in order to increase the efficiency of energy transfer from primary coil  54  of external antenna  52  to secondary coil  34 , a change in either or can result in the resonant frequency, f resonate , being mismatched with the drive frequency, f carrier . The result can be a less than optimum efficiency of energy transfer to implantable medical device  16 . 
     As the drive frequency, f carrier, varies with respect to the resonant frequency, f resonate , apparent impedance of the energy transfer system, as seen by primary coil  54 , will vary. The apparent impedance will be at a minimum when the drive frequency, f carrier , exactly matches the resonant frequency, f resonate . Any mismatch of the drive frequency, f carrier , from the resonant frequency, will cause the impedance to increase. Maximum efficiency occurs when the drive frequency, f carrier , matches the resonant frequency, f resonate . 
     As the impedance of the energy transfer system varies, so does the current driving primary coil  54 . As the impedance of the energy transfer system increases, the current driving primary coil  54  will decreases since the voltage being applied to primary coil  54  remains relatively constant. Similarly, the current driving primary coil  54  will increase as the impedance of the energy transfer system decreases. It can be seen then that point of maximum current driving primary coil  54  will be at a maximum when the impedance of the energy transfer system is at a minimum, when the resonant frequency, f resonate , matches the drive frequency, f carrier , and when maximum efficiency occurs. 
     The impedance of the energy transfer system can be monitored since the current driving primary coil  54  varies as a function of drive frequency, f carrier . The drive frequency can be varied and the current driving primary coil can be measured to determine the point at which the impedance of the energy transfer system is at a minimum, the resonant frequency, f resonate , matches the drive frequency, f carrier , and when maximum efficiency occurs. 
     In a preferred embodiment, instead of holding the drive frequency, f carrier , constant for a nominal resonant frequency, f resonate , the drive frequency, f carrier , is varied until the current driving primary coil  54  is at a maximum. This is not only the point at which the impedance of the energy transfer system is at a minimum but also the point at which maximum efficiency occurs. 
     Maximum efficiency is not as important in systems, such as telemetry systems, which are utilized in a static environment or for relatively short periods of time. In a static environment, the resonant frequency, f resonate , may be relatively invariable. Further, efficiency in not terribly important when energy or information transfer occurs over a relatively short period of time. 
     However, transcutaneous energy transfer systems can be utilized over extended periods of time, either to power the implanted medical device  16  over an extended period of time or to charge a replenishable power supply within implanted medical device  16 . Depending upon capacity of the replenishable power supply and the efficiency of energy transfer, charging unit  50  can be utilized for hours and typically can be used as patient  18  rests or over night as patient  18  sleeps. Further, over the extended period of time in which charging unit  50  is utilized, external antenna  52  is affixed to the body of patient  18 . As patient  18  attempts to continue a normal routine, such as by making normal movement or by sleeping, during energy transfer, it is difficult to maintain external antenna  52  in a completely fixed position relative to secondary coil  34 . Movement of external antenna  52  with respect to secondary coil  34  can result in a change in mutual inductance, L mutual , a change in impedance and a change in the resonant frequency, f resonate . Further, any change in spatial positioning of the energy transfer system with any external conductive object, any change in the characteristics of external antenna  52 , such as by fractures in magnetic core  58 , for example, a change in the charge level of rechargeable power source  24  of implantable medical device  16  or a change in the power level of charging unit  50 , all can result in a change of mutual inductance, L mutual . 
     In a preferred embodiment, drive frequency, f carrier , is varied, not only initially during the commencement of energy transfer, e.g., charging, but also during energy transfer by varying the drive frequency, f carrier , in order to match the drive frequency, with the resonant frequency, f resonate , and, hence, maintaining a relatively high efficiency of energy transfer. As an example, drive frequency, f carrier , can be constantly updated to seek resonant frequency, f resonate , or drive frequency, f carrier , can be periodically updated, perhaps every few minutes or every hour as desired. Such relatively high efficiency in energy transfer will reduce the amount of time charging unit  50  will need to be operated, for a given amount of energy transfer, e.g., a given amount of battery charge. A reduced energy transfer, or charging, time can result in a decrease in the amount of heating of implanted medical device  16  and surrounding tissue of patient  18 . 
     In a preferred embodiment, external charging device  48  incorporates temperature sensor  87  in external antenna  52  and control circuitry in charging unit  50  which can ensure that external antenna  52  does not exceed acceptable temperatures, generally a maximum of thirty-eight degrees Centigrade (38° C.). Temperature sensor  87  in external antenna  52  can be used to determine the temperature of external antenna  52 . Temperature sensor  87  can be positioned in close proximity to thermally conductive material  62  in order to obtain reasonably accurate information on the temperature of the external surface of external antenna  52  contacting patient  18 . Preferably, temperature sensor  87  is affixed to thermally conductive material  62  with a thermally conductive adhesive. Thermally conductive material  62  smoothes out any temperatures differences which otherwise might occur on the surface of external antenna  52  contacting patient  18 . Positioning temperature sensor  87  in the proximity or touching thermally conductive material  62  enables an accurate measurement of the contact temperature. 
     Control circuitry using the output from temperature sensor  87  can then limit the energy transfer process in order to limit the temperature which external antenna  52  imparts to patient  18 . As temperature sensor  87  approaches or reaches preset limits, control circuitry can take appropriate action such as limiting the amount of energy transferred, e.g., by limiting the current driving primary coil  54 , or limiting the time during which energy is transferred, e.g., by curtailing energy transfer or by switching energy transfer on and off to provide an energy transfer duty cycle of less than one hundred percent. 
     When the temperature sensed by the temperature sensor is well below preset temperature limits, it may be acceptable to report the temperature with relatively less precision. As an example, if the temperature sensed by temperature sensor  87  is more than two degrees Centigrade (2° C.) away from a preset limit of thirty-eight degrees Centigrade (38° C.), it may be acceptable to know the temperature with an accuracy of three degrees Centigrade (3° C.). 
     However, when the temperature of external antenna  52  approaches to within two degrees Centigrade (2° C.), it may be desirable to know the temperature with a much greater accuracy, for example, an accuracy of within one tenth of one degree Centigrade (0.1° C.). 
     It is generally difficult, however, to produce a temperature which has a high degree of accuracy over a very broad temperature range. While a temperature sensor can easily be produced to provide a resolution within one-tenth of one degree Centigrade (0.1° C.) over a relatively narrow range temperatures, it can be difficult to produce a temperature sensor providing such a resolution over a broad range of temperatures. 
     In a preferred embodiment, a dual range temperature sensor is utilized. This temperature sensor has a first, broad, less accurate range of measurement from thirty-one degrees Centigrade (31° C.) to forty degrees Centigrade (40° C.) having an accuracy within three degrees Centigrade (3° C.). Further, this temperature sensor has a second, narrow, more accurate range of measurement over four degrees Centigrade (4° C.), from thirty-six degrees Centigrade (36° C.) to forty degrees Centigrade (40° C.), having an accuracy within one-tenth of one degree Centigrade (0.1° C.). 
       FIG. 16  illustrates a preferred embodiment of a dual range temperature sensor utilizing temperature sensor  87 . Temperature sensor  87 , located in external antenna  52 , is coupled to amplifier  170  which has been pre-calibrated to operate only in the range of from thirty-six degrees Centigrade (36° C.) to forty degrees Centigrade (40° C.). Components of amplifier  170  have an accuracy reflecting a temperature within one-tenth of one degree Centigrade (0.1° C.). The analog output of amplifier  170  is sent to analog-to-digital converter  172  producing a digital output  173  having an accuracy of one-tenth of one degree Centigrade (0.1° C.). The analog output of amplifier  170  is also sent to comparator  174  which compares the analog output against a known reference voltage  176  which is set to at a predetermined level to produce a positive output  178  when temperature sensor  87  reflects a temperature of thirty-eight degrees Centigrade (38° C.), the maximum temperature permitted for external antenna  52 . Control logic in charging unit  50  can then take appropriate action to limit further temperature increases such as by ceasing or limiting further energy transfer and/or charging. Temperature sensor  87  is also coupled to amplifier  182 . Components of amplifier  182  have an accuracy reflecting a temperature within three degrees Centigrade (3° C.), much less accuracy than amplifier  170 , but amplifier  182  can operate over the much larger temperature range of thirty-one degrees Centigrade (31° C.) to forty-five degrees Centigrade (45° C.). The output of amplifier  182  is sent to analog-to-digital converter  184  producing a digital output  186  having an accuracy of three degrees Centigrade (3° C.). 
     Some or all of the various features of implantable medical device  16  and charging unit  50  described enable a system for transcutaneous energy transfer having a relatively high efficiency of energy transfer, especially in situations involving some latitude of maladjustment of external antenna  52  with secondary coil  34 . High efficiency of energy transfer can enable a rechargeable power source  24  of implantable medical device  16  to be charged, or recharged, within a shorter period of time than would otherwise be possible. Alternatively or in addition, high efficiency of energy transfer can enable transcutaneous energy transfer to occur at higher rate than would otherwise be possible since more of the energy generated by charging unit  50  is actually converted to charging rechargeable power source  24  instead of generating heat in implanted medical device  16  and/or surrounding tissue of patient  18 . Alternatively or in addition, high efficiency of energy transfer can result in lower temperatures being imparted to implanted medical device  16  and/or surrounding tissue of patient  18 . Alternatively or in addition, high efficiency of energy transfer can enable a greater degree of maladjustment of external antenna  52  with secondary coil  34  effectively resulting in patient  18  being able to be more ambulatory. 
     Thus, embodiments of the external power source for an implantable medical device having an adjustable magnetic core and system and method related thereto are disclosed. One skilled in the art will appreciate that the present invention can be practiced with embodiments other than those disclosed. The disclosed embodiments are presented for purposes of illustration and not limitation, and the present invention is limited only by the claims that follow.