US20050075696A1

Movatterモバイル変換

Info

Publication number: US20050075696A1
Application number: US10/836,318
Authority: US
Inventors: John Forsberg; William Phillips; Andrew Schmeling; David Olson
Original assignee: Medtronic Inc
Current assignee: Medtronic Inc
Priority date: 2003-10-02
Filing date: 2004-04-30
Publication date: 2005-04-07
Also published as: US11318250B2; WO2005039698A1; US9463324B2; ATE355870T1; US8005547B2; DE602004005236T2; US20070167997A1; US20150321015A1; US20140213978A1; US20180036477A1; US20100076524A1; EP1682221A1; US9108063B2; ES2282905T3; US8165678B2; US10369275B2; US8725262B2; AU2004283661A1; US20190358395A1; US20110022125A1

Abstract

External power source, charger, system and method for an implantable medical device having therapeutic componentry and a secondary coil operatively coupled to the therapeutic componentry. A primary charging coil is capable of transcutaneously inductively energizing the secondary coil when externally placed in proximity of the secondary coil. Drive circuitry is operatively coupled to the primary charging coil for exciting the primary charging coil. A rechargeable power source is operatively coupled to the drive circuitry. A secondary recharging coil is operatively coupled to the rechargeable power source. A primary recharging coil is adapted to be coupled to a source of AC power which when placed in proximity of the secondary recharging coil can inductively energize the secondary recharging coil in order to charge the rechargeable power source.

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'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's mobility and limit the patient'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.

BRIEF SUMMARY OF THE INVENTION

An external power source for transcutaneously transferring energy to an implantable medical device must obtain power from a source. The external source of power is typically either from batteries carried within the external power source or directly from readily available AC power, for example, from a standard wall socket. Obtaining power from batteries carried within the external power source is convenient and may allow the patient to be at least somewhat mobile, the amount of power available is limited by the size and weight of the batteries before the batteries are exhausted. However, obtaining power directly from standard readily available AC power not only limits the patient's mobility but also poses a potentially serious health risk. Placing a primary inductive charging coil directly on the patient's skin while the external power source is connected directly to standard AC power risks shock to the patient in the event of a malfunction of the external power source.

In one embodiment, the present invention provides an external power source for an implantable medical device having therapeutic componentry and a secondary coil operatively coupled to the therapeutic componentry. A primary charging coil is capable of transcutaneously inductively energizing the secondary coil when externally placed in proximity of the secondary coil. Drive circuitry is operatively coupled to the primary charging coil for exciting the primary charging coil. A rechargeable power source is operatively coupled to the drive circuitry. A secondary recharging coil is operatively coupled to the rechargeable power source. A primary recharging coil is adapted to be coupled to a source of AC power which when placed in proximity of the secondary recharging coil can inductively energize the secondary recharging coil in order to charge the rechargeable power source.

Utilizing a second inductive charging system, the second one for charging batteries in the external power supply, allows the patient to be mobile when the batteries are sufficient to transfer energy to the implantable medical device and allows the patient to be safely coupled to a source of standard AC power without fear of shock or electrocution if the batteries need to be refreshed.

In another embodiment, the present invention provides a charger for an implantable medical device having an internal rechargeable power source, therapeutic componentry operatively coupled to the internal rechargeable power source. A secondary coil is operatively coupled to the therapeutic componentry and an internal telemetry coil. A primary charging coil is capable of inductively energizing the secondary coil for charging the internal rechargeable power source when externally placed in proximity of the secondary coil. Drive circuitry is operatively coupled to the primary charging coil for exciting the primary charging coil. A rechargeable power source is operatively coupled to the drive circuitry. A secondary recharging coil is operatively coupled to the rechargeable power source. A primary recharging coil is adapted to be coupled to a source of AC power which when placed in proximity of the secondary recharging coil can inductively energize the secondary recharging coil in order to charge the rechargeable power source.

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 and a secondary coil operatively coupled to the therapeutic componentry. An external power source has a primary charging coil capable of inductively energizing the secondary coil when externally placed in proximity of the secondary coil. Drive circuitry is operatively coupled to the primary charging coil for exciting the primary charging coil. A rechargeable power source is operatively coupled to the drive circuitry. A secondary recharging coil is operatively coupled to the rechargeable power source. A primary recharging coil is adapted to be coupled to a source of AC power which when placed in proximity of the secondary recharging coil can inductively energize the secondary recharging coil in order to charge the rechargeable power source.

In a preferred embodiment, said primary recharging coil is located in a cradle capable of receiving the secondary recharging coil for charging the rechargeable power source.

In a preferred embodiment, the implantable medical device has an internal rechargeable power source operatively coupled to the therapeutic componentry and wherein the secondary coil charges the internal rechargeable power source when driven by the primary 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 is operatively coupled to the therapeutic componentry. A primary charging coil is positioned externally of the patient in proximity of the secondary coil. The secondary is inductively energized with the primary charging coil via driver circuitry powered from a rechargeable power source. A secondary recharging coil which is operatively coupled to the rechargeable power source is positioned in proximity of a primary recharging coil adapted to be coupled to a source of AC power in order to charge the rechargeable power source.

In a preferred embodiment, the positioning step comprises setting the secondary recharging coil in a cradle containing the primary recharging 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 ofFIG. 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 ofFIG. 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 ofFIG. 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 implantablemedical device16, for example, a drug pump, implanted inpatient18. The implantablemedical device16 is typically implanted by a surgeon in a sterile surgical procedure performed under local, regional, or general anesthesia. Before implanting themedical device16, acatheter22 is typically implanted with the distal end position at a desiredtherapeutic delivery site23 and the proximal end tunneled under the skin to the location where themedical device16 is to be implanted. Implantablemedical device16 is generally implanted subcutaneously at depths, depending upon application anddevice16, of from 1 centimeter (0.4 inches) to 2.5 centimeters (1 inch) where there is sufficient tissue to support the implanted system. Oncemedical device16 is implanted into thepatient18, the incision can be sutured closed andmedical device16 can begin operation.

Implantablemedical device16 operates to infuse a therapeutic substance intopatient18. Implantablemedical device16 can be used for a wide variety of therapies such as pain, spasticity, cancer, and many other medical conditions.

The therapeutic substance contained in implantablemedical device16 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.

Implantablemedical device16 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.

InFIG. 2, implantablemedical device16 has arechargeable power source24, such as a Lithium ion battery, poweringelectronics26 andtherapy module28 in a conventional manner.Therapy module28 is coupled topatient18 through one ormore therapy connections30, also conventionally.Rechargeable power source24,electronics26 andtherapy module28 are contained in hermetically sealedhousing32.Secondary charging coil34 is attached to the exterior ofhousing32.Secondary charging coil34 is operatively coupled throughelectronics26 torechargeable power source24. In an alternative embodiment,secondary charging coil34 could be contained inhousing32 or could be contained in a separate housing umbilically connected toelectronics26.Electronics26 help provide control of the charging rate ofrechargeable power source24 in a conventional manner.Magnetic shield36 is positioned between secondary chargingcoil34 andhousing32 in order to protectrechargeable power source24,electronics26 andtherapy module28 from electromagnetic energy whensecondary charging coil34 is utilized to chargerechargeable power source24.

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 implantablemedical device16 situated undercutaneous boundary38. Implantablemedical device16 is similar to the embodiment illustrated inFIG. 2. However, chargingregulation module42 is shown separate fromelectronics26

controlling therapy module

28. Again, charging regulation and therapy control is conventional. Implantablemedical device16 also hasinternal telemetry coil44 configured in conventional manner to communicate throughexternal telemetry coil46 to an external programming device (not shown), chargingunit50 or other device in a conventional manner in order to both program and control implantable medical device and to externally obtain information from implantablemedical device16 once implantable medical device has been implanted.Internal telemetry coil44, 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 ofsecondary charging coil34.Secondary coil34 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 shield36 is positioned between secondary chargingcoil34 andhousing32 and sized to cover the footprint ofsecondary charging coil34.

Internal telemetry coil

44, having a larger diameter thansecondary coil34, is not completely covered bymagnetic shield36 allowing implantablemedical device16 to communicate with the external programming device withinternal telemetry coil44 in spite of the presence ofmagnetic shield36.

Rechargeable power source

24 can be charged while implantablemedical device16 is in place in a patient through the use ofexternal charging device48. In a preferred embodiment,external charging device48 consists of chargingunit50 andexternal antenna52. Chargingunit50 contains the electronics necessary to driveprimary coil54 with an oscillating current in order to induce current insecondary coil34 whenprimary coil54 is placed in the proximity ofsecondary coil34. Chargingunit50 is operatively coupled to primary coil bycable56. In an alternative embodiment, chargingunit50 andantenna52 may be combined into a single unit.Antenna52 may also optionally containexternal telemetry coil46 which may be operatively coupled to chargingunit50 if it is desired to communicate to or from implantablemedical device16 withexternal charging device48. Alternatively,antenna52 may optionally containexternal telemetry coil46 which can be operatively coupled to an external programming device, either individually or together withexternal charging unit48.

As will be explained in more detail below, repositionablemagnetic core58 can help to focus electromagnetic energy fromprimary coil46 to more closely be aligned withsecondary coil34. Also as will be explained in more detail below,energy absorptive material60 can help to absorb heat build-up inexternal antenna52 which will also help allow for a lower temperature in implantablemedical device16 and/or help lower recharge times. Also as will be explained in more detail below, thermallyconductive material62 is positioned covering at least a portion of the surface ofexternal antenna52 which contactscutaneous boundary38 ofpatient18.

In a preferred embodiment ofinternal antenna68 as shown inFIG. 4 andFIG. 5,secondary coil34 andmagnetic shield36 are separate from but adjacent tohousing32 encompassing the remainder of implantablemedical device16.Internal antenna68 is contained in aseparate housing74 which is attachable tohousing32 so that implantablemedical device16 can be implanted by a medical professional as essentially one unit.Secondary coil34 is electrically attached to chargingregulation module42 through leads82.

In order to achieve efficient inductive coupling betweenprimary coil54 ofexternal antenna52 andsecondary coil34, it is desirable to placeprimary coil54 ofexternal antenna52 as close tosecondary coil34 as possible. Typically,external antenna52 is placed directly oncutaneous boundary38 and, since the location of implantablemedical device16 is fixed, the distance acrosscutaneous boundary38 betweenprimary coil54 andsecondary coil34 is minimized as long asexternal antenna52 is kept adjacentcutaneous boundary38.

In a preferred embodiment,external antenna52 is attachable to patient18 withbracket84 when chargingrechargeable power source24.FIG. 6 is an exploded illustration of a preferred embodiment ofexternal antenna52 attachable tobracket84.Primary coil54 is contained inbobbin assembly86 which sits inbottom housing88. Primary coil is connectable tocable56. The bottom ofexternal antenna52 is formed from a thermallyconductive material90. Rotatingcore cup assembly92 is held in place bytop housing94. Rotatingcore cup assembly92 is rotatable is allowed to rotate withinexternal antenna52.Detents96 engagedetent spring98 to position rotatablecore cup assembly92 in one of a plurality of detent positions. External antenna may be secured together, for example, with screws (not shown) holdingtop housing94 and thermallyconductive material90 together.

Bracket

84 is adapted to be attached to the body ofpatient18 with a belt (not shown) attachable tobracket84 withbelt loops102.Ears104 are adapted to mate with tabs106 intop housing94 and pivotally secureexternal antenna52 inbracket84 when charging is to be accomplished.Bracket84 has anopening108 allowing thermallyconductive material90 ofexternal antenna52 to contact the skin ofpatient18 whenexternal antenna52 is pivotally secured inbracket84.

Asbracket84 is attached topatient18 with a belt viabelt loops102, the skin surface ofpatient18 is typically not completely flat. For example, if implantablemedical device16 is implantable in the body torso ofpatient18, then the belt attached viabelt loops102 will typically pass around the torso ofpatient18. Since the torso ofpatient18, and especially the torso ofpatient18 near the location of implantablemedical device16,bracket84 may not sit completely flat onpatient18. This may be especially true aspatient18 moves and the torso flexes during such movement. It is preferred thatbracket84 be conformal and flexible in order to conform to the shape of the body ofpatient18. However, it is also preferred thatbracket84 be rigid enough so that opening108 inbracket84 maintain it shape in order to properly receiveexternal antenna52.Bracket84 is preferably constructed of PCABS. To maintain the proper position ofbracket84 with the skin ofpatient18, the surface ofbracket84 closest to patient18 contains material109 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 ofpatient18 and remain there during movements ofpatient18 throughout the charge or recharge cycle. In addition,external antenna52 is allowed to pivot by way ofears104 on tabs106.Bracket84 is configured to allow thermallyconductive material90 to extend throughopening108 and contact the skin surface ofpatient18. Allowed pivoting ofexternal antenna52 and, hence, thermallyconductive material90, permits thermally conductive surface to sit more closely to the skin surface ofpatient18.

FIG. 7 is a partially cut away top view ofexternal antenna52 is assembled form and attached tocable56. Rotatablecore cup assembly92 is shown located inside ofprimary coil54 and positionable in selected rotated positions viadetents96 anddetent spring98. InFIG. 7, rotatable core cup assembly is positioned between withdetent spring98 betweendetents96 illustrating that while multiple detent positions are available, rotatable core cup assembly can be positioned between detent positions and, indeed, at any rotated position.

InFIG. 8, the assembly ofexternal antenna52 withbracket84 is shown connected tocable56. It is preferred thatbracket84 be affixed topatient18 throughbelt loops102 and then, afterbracket84 has been affixed topatient18,external antenna52 be attached tobracket84. Affixingbracket84 topatient18 first allows forbracket84 to be used to laterally position external antenna close to the position of implantablemedical device16.

In contrast,bracket84 of the present invention can be used to roughly find the optimum position forexternal antenna52. This can be done relatively easily due to opening108 inbracket84. Implantablemedical device16, when implanted, usually leaves an area of the body ofpatient18 which is not quite as flat as it was before implantation. That is, implantablemedical device16 usually leaves an area of the skin ofpatient18 which bulges somewhat to accommodate the bulk of implantablemedical device16. It is relatively easy for patient, medical professional or other person, to placebracket84 in the general area of implantablemedical device16 and movebracket84 around until the bulge caused by implantablemedical device16 is most closely centered inopening108. Asbracket84 is moved laterally, opening108 tends to naturally center on the bulge created by implantablemedical device16. Once positioned in this manner,bracket84 can be secured to the body ofpatient18 with belt (not shown) attached viabelt loops102. Securing and/or tightening, by pulling the belt tight or snapping a buckle, for example, can be without removingbracket84 from the body ofpatient16. Thus,bracket84 can be relatively easily positioned over the general location of implantablemedical device16 and secured in that position without be removed from the body ofpatient18.

FIG. 9 is cross-sectional view of implantablemedical device16 implanted inpatient18 approximately one centimeter undercutaneous boundary38 creating bulgingarea110, an area of the body ofpatient18 in which the skin ofpatient18 is caused to bulge slightly due to the implantation of implantablemedical device16. Bulgingarea110 is an aid to locating the position ofexternal antenna52 relative tosecondary coil34.Bracket84 can be positioned roughly in the area where implantablemedical device16 is implanted. Opening108 inbracket84 can aid is establishing the location of implantable medical device.Bracket84 can be roughly centered over bulgingarea110. Afterexternal antenna52 is coupled tobracket84, thenprimary coil54 can be generally centered on implantablemedical device16.

However,secondary coil34 may not be centered with respect to implantablemedical device16. This can occur due to a variety of reasons such as the need for operatively couplingsecondary coil34 to chargingregulation module42. Connections to make this operative coupling may require physical space on one side ofinternal antenna68 which may causesecondary coil34 not to be centered on implantablemedical device16. It is also possible that the attachment ofinternal antenna68 tohousing32 can causesecondary coil34 not to be centered on implantablemedical device16. Regardless of the cause, ifsecondary coil34 is not centered on implantablemedical device16, then centeringbracket84 on bulgingarea110 may not optimally positionprimary coil54 with respect tosecondary coil34. Any offset in the position ofprimary coil54 andsecondary coil34 may not result in the most efficient energy transfer fromexternal antenna52 to implantablemedical device16.

Amagnetic core58 is positioned withinprimary coil54 in order to focus energy generated byprimary coil54.Magnetic core58 attracts the magnetic flux lines generated byprimary coil54. The position ofmagnetic core58 withinprimary coil54 the lateral location of the largest amount of the flux lines generated byprimary coil54.FIGS. 10 and 11 show cut-away top and cross-sectional views ofmagnetic core58 used withprimary coil54.Magnetic core58 is moveable withinprimary coil54.Lower portion122 ofmagnetic core58 can be rotated to a plurality of positions withinprimary coil58 by rotating core cup assembly92 (seeFIG. 12). In a preferred embodiment, the travel path ofmagnetic core58 can be locked in a plurality of discrete positions. In a preferred embodiment,magnetic core58 is locked in four (4) different positions bydetents96 and detent spring98 (seeFIG. 6).Magnetic core58 has an upperplanar portion120 and a smallerlower portion122.

Asmagnetic core58 is repositioned withinprimary coil54, the focus of magnetic flux generated byprimary coil54 is also repositioned. As noted above,external antenna52 is generally aligned with implantedmedical device16 using palpatory sensation. Moveablemagnetic core58 can then be used to provide a “fine” adjustment to the lateral positioning ofexternal antenna52 with respect tosecondary coil34. Afterbracket84 has been secured topatient18,external antenna52 is attached tobracket84.Magnetic core58 is then moved until the best lateral alignment withsecondary coil34.

Magnetic core

58 is shown positioned withinexternal antenna52 ofFIG. 12.Core cup assembly92 holdsmagnetic core58 within the assembly ofexternal antenna52. Lower portion122 (not visible inFIG. 12) ofmagnetic core58 fits intorecess124 ofcore cup assembly92 whileupper portion120 ofmagnetic core58 rests uponledge126 ofcore cup assembly92. Preferably,magnetic core58 is a ferrite core. Still more preferably,magnetic core58 is constructed from MN60LL high performance, low loss ferrite manufactured by Ceramic Magnetics, Inc., Fairfield, N.J.Magnetic core58 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, ofmagnetic core58 is lined with an adhesivecoated foam127 and contained incore cup assembly92.Magnetic core58 has a tendency to be brittle. Containingmagnetic core58 is core cup assembly assures that even ifmagnetic core58 has one or more fractures,magnetic core58 will still be properly positioned and continue to function.Foam127 also helps to holdmagnetic core58 together and minimize gaps between fractured segments ofmagnetic core58. Further,foam127 adds mechanical stability tomagnetic core58 helping to cushionmagnetic core58 against mechanical impacts, such as from droppingexternal antenna52 against a hard surface, and helps to prevents audible rattles which may otherwise develop from a fracturedmagnetic core58.

As shown inFIG. 13,external charging device48 can be powered either directly from internal (to charging unit50)batteries160 or indirectly fromdesktop charging device162. Desktop charging device is connectable viapower cord164 to a source of AC power, such as a standard readily available wall outlet.Desktop charging device162 can be configured as a cradle which can receive chargingunit50. Other forms of connection fromdesktop charging device162 to a power source, such as by a dedicated line cable can also be utilized.Desktop charging device162 can charge and/or rechargebatteries160 in chargingunit50, preferably by inductivecoupling using coil167 positioned indesktop charging device162 andcoil168 positioned within chargingunit50. Once charged and/or recharged,batteries160 can provide the power throughinternal circuitry168 andcable56 toexternal antenna52. Since chargingunit50 is not, in a preferred embodiment, coupled directly to the line voltage source of AC power, chargingunit50 may be used withexternal antenna52 to transfer power and/or charge implantedmedical device16 whiledesktop charging device162 is coupled to a line voltage source of AC power. The inductivecoupling using coil167 andcoil168 break the possibility of a direct connection between the line voltage source of AC power andexternal antenna52.Batteries160 also allow chargingunit50 and, hence,external charging device48, to be used in transferring power and/or charging of implantedmedical device16 while completely disconnected from either a line voltage source of AC power ordesktop charging device162. This, at least in part, allowspatient18 to be ambulatory while transferring power and/or charging implantedmedical device16.

FIG. 14 is a block diagram ofexternal charging device48 controlled bymicroprocessor212. Transmitblock214 consists of an H-bridge circuit powered from 12 volt power supply216. Transmitblock214 drivesprimary coil54 inexternal antenna52. H-bridge control signals and timing are provided conventionally bymicroprocessor212. H-bridge circuit in transmitblock214 is used to drive bothprimary coil54, used for power transfer and/or charging, and telemetry antenna218. Drive selection is done by electronicallycontrollable switch220. 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 block222 is used only during telemetry, enabled by switch224, to receive uplink signals from implantedmedical device16. Twelve volt power supply216 is a switching regulator supplying power to transmitblock214 during power transfer and/or charging as well as telemetry downlink. Nominal input voltage to 12 volt power supply216 is either 7.5 volts from lithium ion batteries226 or 10 volts from desktop charging device162 (FIG. 13).

Current measure block226 measures current to 12 volt power supply216. Current measured by current measure block226 is used in the calculation of power in along with the voltage ofbatteries160. 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 ofexternal antenna52 and/or rotatingcore cup assembly92.

Rotatingcore cup assembly92 is rotated inexternal antenna52 for better lateral alignment ofprimary coil54 andsecondary coil34. A feedback mechanism is used to determine the best rotation ofcore cup assembly92.External charging device48 can determine whether the current position of rotatingcore cup assembly92 is optimally aligned for energy transfer and/or charging.External charging device48 measures the power out ofexternal charging device48 divided by the power intoexternal charging device48. This calculation is a measure of the efficiency ofexternal charging device48. The power out is gauged by the power induced in implantablemedical device16 and is determined by multiplying the voltage ofrechargeable power source24 by the charging current in implantablemedical device16. These values are obtained by telemetry from implantedmedical device16. The power in is gauged by the power generated by chargingunit50 and is determined by multiplying the voltage of the internal voltage of chargingunit50, e.g., the voltage of a battery or batteries internal to chargingunit50, by the current drivingexternal antenna52.

The ratio of power out divided by power in can be scaled displayed topatient18, or a medical professional or other person adjusting rotatablecore cup assembly92 or positioningexternal antenna52. 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 ofexternal antenna52 and rotatablecore cup assembly92 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, afterpatient18 or other person has moved rotatablecore cup assembly92 through all of the range of positions onexternal antenna52 and can not achieve an acceptable efficiency level,patient18 or other person can removeexternal antenna52 frombracket84, realignbracket84 with bulgingarea110, reattachexternal antenna52 tobracket84 and restart the alignment and coupling efficiency process.

FIG. 15 is a flow chart illustrating an exemplary charging process usingexternal antenna52. The process starts [block126and a charging session begins [block128] with a test [block130]. The charging system performs start-up checks [block132]. 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 implantablemedical device16 is checked [block134]. 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 device48 is able to monitor [block136] charging status. Monitoring charging status can includes providing feedback to an operator to help determine the best rotational position ofcore cup assembly92.

Charge events are checked [block138]. 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 [block140]. Once charging is complete, the process terminates [block142].

As energy is transferred fromprimary coil54 ofexternal antenna52 tosecondary coil34 of implantablemedical device16, heat may also be generated in implantablemedical device16 in surrounding tissue ofpatient18. Such heat build-up in tissue ofpatient18, beyond certain limits, is undesirable and should be limited as acceptable values. Generally, it is preferable to limit the temperature ofexternal antenna52 to not more than forty-one degrees Centigrade (41° C.) and to limit the temperature of implantedmedical device16 and the skin ofpatient18 to thirty-nine degrees Centigrade (39° C.). In order to ensure that implantablemedical device16 is less than the upper limit of thirty-nine degrees Centigrade (39° C.), it is preferred that the actual temperature ofexternal antenna52 be less than thirty-nine degrees Centigrade (39° C.). In general, the temperature ofexternal antenna52 should be maintained to be less than or equal to the desired maximum temperature of implantedmedical device16. 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 chargingunit50 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 implantablemedical device16 contained withintitanium housing32 from the effects of energy transfer fromexternal charging device48 produced through inductive coupling fromprimary coil54.Magnetic shield36 is constructed of Metglas magnetic alloy 2714A (cobalt-based) manufactured by Honeywell International, Conway, S.C.Magnetic shield36 is positioned betweensecondary coil34 andhousing32 of implantablemedical device16 withsecondary coil34 facingcutaneous boundary38. Magnetic shield does not interfere with the operation ofsecondary coil34 becausemagnetic shield36 is positioned away fromprimary coil54. Also, magnetic shield does not interfere with telemetry between implantablemedical device16 and an external programmer becausemagnetic shield36 is smaller thaninternal telemetry coil44. That is,internal telemetry coil44 lies outside ofmagnetic shield36.

However, the material ofmagnetic shield36 substantially limits the electromagnetic energy induced byprimary coil54 from penetrating beyond magnetic shield. Electromagnetic waves induced byprimary coil54 that reachtitanium housing32 will tend to be absorbed bytitanium housing54 and its components and will tend to cause the temperature oftitanium housing54 to rise. As the temperature oftitanium housing54 rises, such temperature increase will be disadvantageously transferred to the surrounding tissue ofpatient18. However, any electromagnetic waves which are prevented from reachingtitanium housing32 will not cause such a temperature rise.

Thermallyconductive material62 ofexternal antenna52 is positioned to contact the skin ofpatient18 whenexternal antenna52 is placed for energy transfer, or charging, of implantedmedical device16. Thermallyconductive material62 tends to spread any heat generated at the skin surface and spread any such heat over a larger area. Thermallyconductive material62 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.

Thermallyconductive material62 is molded into the surface ofexternal antenna52 which will contact the skin surface ofpatient18 whenexternal antenna52 provides energy transfer to implantedmedical device16. Since thermallyconductive material62 should pass electromagnetic energy fromprimary coil54, thermallyconductive material62 should be constructed from a non-magnetic material. It is desirable that thermallyconductive material62 have a thermal conductivity of approximately 5.62 BTU inch/hour feet²degrees 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 10³ohm-centimeters and a surface resistivity of 10⁵ohms per square.

Energy absorptive material

62 is placed in and/or aroundprimary coil54 ofexternal antenna52 in order to absorb some of the energy generated byprimary coil54. In a preferred embodiment,energy absorptive material62 fills in otherwise empty space of rotatingcore cup assembly92. Heat generated by energy produced byprimary coil54 which is not effectively inductively coupled tosecondary coil34 will tend to cause a temperature rise in other components ofexternal antenna52 and, possibly, the skin ofpatient18. At least a portion of this temperature rise can be blocked through the use ofenergy absorptive material62.Energy absorptive material62 is chosen to absorb heat build-up in surrounding components and tend to limit further temperature increases. Preferably,energy absorptive material62 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, usingexternal antenna52.

If it is a goal to limit the temperature of the skin ofpatient18 to thirty-nine degrees Centigrade (39° C.), it is desirable to use ofenergy absorptive material62 which has a state change at or near the temperature limit. In this example, the use of anenergy absorptive material62 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 ofpatient18 to no more than the desired limit, in this example, thirty-nine degrees (39° C.).

As the temperature of surrounding components ofexternal antenna52 rise to a temperature which is just below the temperature at whichenergy absorptive material62 changes state, at least a portion of further heat energy generated byprimary coil54 and surrounding components ofexternal antenna52 will go toward providing the energy necessary forenergy absorptive material62 to change state. Asenergy absorptive material62 is in the process of changing state, its temperature is not increasing. Therefore, during the state change ofenergy absorptive material62,energy absorptive material62 is serving to at least partially limit a further rise in the temperature of components ofexternal antenna52. 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 ofpatient18,energy absorptive material62 will tend to limit the temperature components ofexternal antenna52 from reaching the temperature limit and, hence, will also tend to limit the temperature of the skin ofpatient18 from reaching the maximum desired temperature limit.

In a preferred embodiment,energy absorptive material62 is constructed from wax and, in particular, a wax which has change of state temperature of approximately the maximum temperature at whichexternal antenna52 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 betweenprimary coil54 ofexternal antenna52 and secondary coil of implantablemedical device16 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, ofexternal antenna54 is approximately nine (9) kiloHertz unloaded.

However, the inductive coupling betweenprimary coil54 ofexternal antenna52 andsecondary coil34 of implantable medical device is dependent upon the mutual inductance between the devices. The mutual inductance depends upon a number of variables.Primary coil54 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 andexternal antenna52 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_resonateequals 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 tuneexternal antenna52 to resonate at the carrier frequency of 1 megaHertz is calculated as approximately 25 nanofarads.

However, when the electrical properties ofexternal antenna52 change, either by the reflected environment or due to a physical distortion or change in the composition of theexternal antenna52, 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, ofexternal antenna52 at f_carrier. The second part is the mutual inductance, L_mutual, which is a measure of the change in current drivingexternal antenna52 and the magnetic effect, or “loading”, which the environment has onexternal antenna52. When the electrical characteristics of the environment ofexternal antenna52 change, L_selfremains constant while L_mutualvaries. The effect of a change in the overall inductance, whether that change is from L_selfor 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 fromprimary coil54 ofexternal antenna52 tosecondary coil34, 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 implantablemedical device16.

As the drive frequency, f_carrier, varies with respect to the resonant frequency, f_resonate, apparent impedance of the energy transfer system, as seen byprimary coil54, 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 drivingprimary coil54. As the impedance of the energy transfer system increases, the current drivingprimary coil54 will decreases since the voltage being applied toprimary coil54 remains relatively constant. Similarly, the current drivingprimary coil54 will increase as the impedance of the energy transfer system decreases. It can be seen then that point of maximum current drivingprimary coil54 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 drivingprimary coil54 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 drivingprimary coil54 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 implantedmedical device16 over an extended period of time or to charge a replenishable power supply within implantedmedical device16. Depending upon capacity of the replenishable power supply and the efficiency of energy transfer, chargingunit50 can be utilized for hours and typically can be used aspatient18 rests or over night aspatient18 sleeps. Further, over the extended period of time in which chargingunit50 is utilized,external antenna52 is affixed to the body ofpatient18. Aspatient18 attempts to continue a normal routine, such as by making normal movement or by sleeping, during energy transfer, it is difficult to maintainexternal antenna52 in a completely fixed position relative tosecondary coil34. Movement ofexternal antenna52 with respect tosecondary coil34 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 ofexternal antenna52, such as by fractures inmagnetic core58, for example, a change in the charge level ofrechargeable power source24 of implantablemedical device16 or a change in the power level of chargingunit50, 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 oftime charging unit50 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 implantedmedical device16 and surrounding tissue ofpatient18.

In a preferred embodiment,external charging device48 incorporatestemperature sensor87 inexternal antenna52 and control circuitry in chargingunit50 which can ensure thatexternal antenna52 does not exceed acceptable temperatures, generally a maximum of thirty-eight degrees Centigrade (38° C.).Temperature sensor87 inexternal antenna52 can be used to determine the temperature ofexternal antenna52.Temperature sensor87 can be positioned in close proximity to thermallyconductive material62 in order to obtain reasonably accurate information on the temperature of the external surface ofexternal antenna52 contactingpatient18. Preferably,temperature sensor87 is affixed to thermallyconductive material62 with a thermally conductive adhesive. Thermallyconductive material62 smoothes out any temperatures differences which otherwise might occur on the surface ofexternal antenna52 contactingpatient18.Positioning temperature sensor87 in the proximity or touching thermallyconductive material62 enables an accurate measurement of the contact temperature.

Control circuitry using the output fromtemperature sensor87 can then limit the energy transfer process in order to limit the temperature whichexternal antenna52 imparts topatient18. Astemperature sensor87 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 drivingprimary coil54, 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 bytemperature sensor87 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 ofexternal antenna52 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 utilizingtemperature sensor87.Temperature sensor87, located inexternal antenna52, is coupled toamplifier170 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 ofamplifier170 have an accuracy reflecting a temperature within one-tenth of one degree Centigrade (0.1° C.). The analog output ofamplifier170 is sent to analog-to-digital converter172 producing adigital output173 having an accuracy of one-tenth of one degree Centigrade (0.1° C.). The analog output ofamplifier170 is also sent to comparator174 which compares the analog output against a knownreference voltage176 which is set to at a predetermined level to produce apositive output178 whentemperature sensor87 reflects a temperature of thirty-eight degrees Centigrade (38° C.), the maximum temperature permitted forexternal antenna52. Control logic in chargingunit50 can then take appropriate action to limit further temperature increases such as by ceasing or limiting further energy transfer and/or charging.Temperature sensor87 is also coupled toamplifier182. Components ofamplifier182 have an accuracy reflecting a temperature within three degrees Centigrade (3° C.), much less accuracy thanamplifier170, butamplifier182 can operate over the much larger temperature range of thirty-one degrees Centigrade (31° C.) to forty-five degrees Centigrade (45° C.). The output ofamplifier182 is sent to analog-to-digital converter184 producing adigital output186 having an accuracy of three degrees Centigrade (3° C.).

Some or all of the various features of implantablemedical device16 and chargingunit50 described enable a system for transcutaneous energy transfer having a relatively high efficiency of energy transfer, especially in situations involving some latitude of maladjustment ofexternal antenna52 withsecondary coil34. High efficiency of energy transfer can enable arechargeable power source24 of implantablemedical device16 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 chargingunit50 is actually converted to chargingrechargeable power source24 instead of generating heat in implantedmedical device16 and/or surrounding tissue ofpatient18. Alternatively or in addition, high efficiency of energy transfer can result in lower temperatures being imparted to implantedmedical device16 and/or surrounding tissue ofpatient18. Alternatively or in addition, high efficiency of energy transfer can enable a greater degree of maladjustment ofexternal antenna52 withsecondary coil34 effectively resulting inpatient18 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.