Abstract:
A fixed frequency external power source having an external coil is inductively coupled with an implanted coil of an implanted medical device. The implant device has an electronic impedance transformer as part of its load circuit. Such electronic impedance transformer sets a proper voltage and current ratio (impedance) so that the coil set, i.e., the external coil and the implanted coil, are loaded with an optimal load. Such optimal loading, in turn, significantly minimizes any mismatch loss from the inductive link between the external coil and the implant coil, and allows wide ranges in the voltage and load resistance and coil separation, while at the same time maintaining an optimal load condition. The impedance transformer is especially applicable to fully implantable cochlear stimulation systems wherein, during one mode of operation, a relatively large power level must be transferred for charging the implanted power storage element, e.g., a rechargeable battery, but wherein another mode of operation, the implant is operated and powered from an external unit and a relatively small power level is transferred to the implant device. The ratio of these power levels may be large, e.g., about 30 to 1, and unless the coil set, i.e., the external coil and implanted coil, are altered between these different load conditions, a huge mismatch loss may occur, which mismatch greatly reduces the power transfer efficiency. The impedance transformer of the invention minimizes such a mismatch loss.

Description:
The present application claims the benefit of U.S. Provisional Application Serial No. 60/239,288 filed Oct. 11, 2000, which application is incorporated herein by reference. 
    
    
     BACKGROUND OF THE INVENTION 
     The present invention relates to an implantable electrical device, e.g., an implantable medical device such as an implantable cochlear stimulation system, which receives its operating power and/or which receives recharging power from an external (non-implanted) power source. 
     Implantable electrical devices are used for many purposes. A common type of implantable device is a tissue stimulator. A tissue stimulator includes one or more electrodes in contact with desired tissue. An electrical stimulation current is generated by the stimulator and applied to the tissue through the electrode(s). 
     In order for an implanted device to perform its intended function, e.g., to generate an electrical stimulation current, it needs a power source. Some implanted devices, e.g., cardiac pacemakers, employ a high capacity battery that has sufficient power stored therein to provide operating power for the device for several years. Other implanted devices, e.g., a cochlear stimulation system, do not use an implanted power source but rather receive a continuous stream of power from an external source through an rf (radio frequency) or inductive link. Yet other implanted devices include a rechargeable power source, e.g., a rechargeable battery, that must be regularly recharged, e.g., once a day, or 2-3 times per week, from an external source in order for the implanted device to operate. The present invention is intended for use with the latter two types of implanted devices, e.g., those that receive a continuous stream of operating power from an external source, and/or those that must receive power at regular intervals in order to recharge an implantable power source. 
     Power is typically coupled to an implanted device through inductive coupling. Inductive coupling advantageously avoids the use of wires that must pass through or penetrate the skin. With inductive coupling, an external coil receives an ac power signal. An implanted coil connected to, or forming part of, the implantable device, is placed in close proximity to the external coil so that magnetic flux generated by the ac power signal in the external coil induces an ac power signal in the second coil, much like the primary winding of a transformer couples energy to a secondary winding of the transformer, even though the two windings are not directly connected to each other. When inductively coupling power to an implanted device in this manner, an optimum power transfer condition exists only when there is a good impedance match between the implant device and the external device. While impedance matching schemes can and have been used in the external device, such matching schemes are only effective for a given distance between the external coil and the implant coil, and for a given load attached to the implant device. 
     Unfortunately, neither the load associated with the implant device nor the separation distance between the external coil and the implant coil are constants. Each of these parameters are, in practice, variables, that may vary, e.g., from 3-to-15 mm for the separation distance, and 20 to 300 ohms for the load. As a result, optimum power transfer between the external device and implant device is rarely achieved. Thus, a less than optimum power transfer condition exists and much of the energy sent to the external coil is lost. What is needed, therefore, is a way to assure that optimum power transfer conditions exist between the external coil and implant device at the time a power transfer is made. 
     For many implant devices, optimum power transfer has heretofore generally not been a serious concern inasmuch as the external device (which has generally comprised a relatively large device that is worn or carried by the patient) has been viewed as having a potentially infinite power source (through recharging and/or replacing its battery). Unfortunately, however, transferring large amounts of power without concern for how much power is lost is not only inefficient, but may create regulatory problems. That is, most regulatory agencies stipulate the power levels that may be used with an implant device. 
     Further, new generation external devices are being made smaller and smaller to accommodate the needs and desires of the user. For example, a behind-the-ear (BTE) external device may be used with an implantable cochlear stimulator (ICS). Such a BTE external device is about the same size as a conventional behind-the-ear hearing aid. Such smaller devices, as a practical manner, do not have a potentially infinite power source, but must be powered using a small button battery, or equivalent. Such a small battery must provide power for both the external unit and the implant unit, and achieving an efficient power transfer is a key element in assuring a long battery life. 
     It is known in the art, see, e.g., U.S. Pat. No. 4,654,880, to include the external coil and implant coil (as coupled to each other based on a given separation distance and load) in the oscillator circuit that sets the frequency of the signal that is coupled between the external coil and implant coil. Such circuit is somewhat self-compensating because as the transfer efficiency starts to go down (e.g., because the separation distance changes, or because the load changes) the frequency of the signal used to couple energy into the implant coil automatically changes in a direction that tends to retune the coupled coils so that the energy transfer becomes more efficient. 
     It is also known in the art, see, e.g., U.S. Pat. No. 5,179,511, to use a self-regulating Class E amplifier, combined with current feedback, to better control the frequency of the coupling signal so as to achieve a more optimum energy transfer. 
     Disadvantageously, changing the frequency of the signal coupled into the implant circuit may also create regulatory problems. That is, regulatory agencies are typically very strict about the frequencies of signals that are allowed to be transmitted, even if only transmitted over short distances. 
     One technique known in the art for optimally transferring power is through the use of a DC-to-DC converter. Disadvantageously, stability problems may arise when using a DC-to-DC converter. More particularly, switching regulators, a common form of DC-to-DC converters, are prone to “bistability”, as discussed in the article: “Source resistance: the efficiency killer in DC-DC converter circuits”, which article is attached hereto as Appendix A and is incorporated herein by reference. 
     In view of the above, it is evident that what is needed is a transmission scheme for use with a medical implant device that optimally transfers power to the implant device from an external device at a fixed frequency, i.e., that transfers power into the implant device from the external device with minimum power loss. 
     SUMMARY OF THE INVENTION 
     The present invention addresses the above and other needs by providing a fixed frequency external power source that is inductively coupled with an implanted device. Unlike prior art implanted devices, however, the implant device of the present invention utilizes an electronic impedance transformer as part of the load circuit in the implant device. Such electronic impedance transformer stabilizes, or makes constant, the load resistance. While the impedance seen looking into the external coil is still very much a function of the coil separation, and hence may not be optimal (this impedance follows a parabolic shaped loss curve, well known in the art, as a function of coil separation distance), it is now possible, with an adjustable stabilized load resistance (made possible by the impedance transformer of the present invention) for a smart external device to measure the impedance seen looking into the external coil (which impedance includes both the coil separation loss and the stabilized load resistance made possible by the invention) and vary the internal impedance transformer to achieve an overall better power transfer. Hence, the invention makes possible the proper voltage and current ratio (resistance) to exist, so that the coil set, i.e., the external coil and the implanted coil, are loaded with the “best available” load under the circumstances. Such best possible load, in turn, minimizes mismatch losses from the inductive link between the external coil and the implant coil, and allows wide ranges in the voltage and load resistance and coil separation, while at the same time maintains a best possible load condition. 
     The present invention is especially applicable to fully implantable cochlear stimulation systems. A representative fully implantable cochlear stimulation system is disclosed, e.g., in U.S. Pat. No. 6,067,474 and/or in U.S. patent application Ser. No. 09/404,966, filed Sep. 24, 1999, which patent and patent application are incorporated herein by reference. In a fully implantable system (FIS), the FIS preferably operates using power from an implanted power source, such as a rechargeable battery, which power source must be periodically recharged by transferring large amounts of power to the implant device. However, the FIS must also be able to operate, from time to time or in the event of a battery or other failure, using an external behind-the-ear (BTE) unit, or other external unit, which requires a power transfer at much lower power levels than are needed for recharging. That is, in the FIS, during one mode of operation, a relatively large power level must be transferred for charging the implanted power storage element, e.g., a rechargeable battery. However, in another mode of operation, the implant is operated and powered from a BTE unit, or other external unit, during which mode a relatively small power level is transferred to the implant device. The ratio of these power levels may be, e.g., about 30 to 1. Unless the coil set, i.e., the external coil and implanted coil, are altered between these different load conditions, a mismatch loss on the order of 14dB may occur, which mismatch may reduce the transfer efficiency from about 70% to about 3%! The present invention advantageously eliminates such a mismatch loss. 
     In accordance with one aspect of the invention, a time-varying impedance transformer is utilized to make the mismatch loss constant. The control of the mismatch is determined by the load impedance, once all other components are fixed. However, because an implant device of the type with which the present invention is used may require a range of output voltages, and output currents, the effective load resistance is not equivalent to a single load resistance, but rather varies as a function of time dependent upon the required circuit operation. The time-varying impedance transformer provided in the implant device by the invention thus operates to stabilize (make constant insofar as possible) the ratio of output voltage and output current as seen by the coil set, thereby rendering the mismatch loss constant, even though the individual output voltages and output currents do vary. 
     In accordance with another aspect of the invention, a switching regulator circuit is employed as the time varying impedance transformer. Advantageously, a switching regulator circuit provides for the efficient transfer of electrical power from one voltage level to another. A switching regulator operates as a DC-to-DC impedance transformer. That is, at its input, the switching regulator consumes the required current at the source voltage level, and transforms the current to a new level at a different output voltage. Since energy is neither created nor destroyed, the switching regulator functions as a power transformer, with some loss occurring (as determined by the converter efficiency). Hence, in accordance with the present invention, a switching regulator included as part of the implant circuitry is controlled in an appropriate manner so that the resulting impedance transforming property of the switching regulator reduces mismatch loss variations. 
     It is thus a feature of the present invention to provide an implantable medical device, e.g., an implantable cochlear stimulator or other implantable neural stimulator, that employs a switching regulator circuit as part of the implanted circuitry. The switching regulator is controlled, as energy is inductively coupled into the implant circuitry through a coil set that includes an external coil and an implanted coil, to operate as a varying impedance transformer. More particularly, the impedance is varied so as to minimize mismatch losses as seen at the power source, thereby improving the power transfer efficiency into the implant device. 
     It is a further feature of the invention to provide an implantable time-varying impedance transformer wherein there are no circuit value or wiring changes needed to handle varying output load impedance. Rather, all control is entirely electronic. 
     One advantage of the invention is that the frequency of the carrier signal (the signal applied to the external coil) is fixed, thereby avoiding regulatory or other problems incident to using variable frequency carrier signals. 
     An additional advantage of the invention is that the effective DC load resistance of the implanted circuitry (output voltage divided by output current) is transformed to effect an AC circuit mismatch loss, so that the lowest insertion loss of the coil set (i.e., the highest power transfer efficiency between the external coil and implanted coil) may be utilized. 
     Still another advantage of the invention is that the voltage transform ratio of the coil set is, within certain practical constraints, relatively independent of the voltage at the output load. 
     Another advantage of the invention, when used in combination with a smart external power source that can regularly measure the impedance as seen looking into the external coil and communicate this measured impedance to the implant device, is that the impedance transforming process that occurs in the implant device, acting upon the measured impedance information obtained from the smart external device, may also be used to compensate for variations in transfer efficiency that occur due to coil separation. 
    
    
     BRIEF DESCRIPTION OF THE DRAWINGS 
     The above and other aspects, features and advantages of the present invention will be more apparent from the following more particular description thereof, presented in conjunction with the following drawings and appendix wherein: 
     FIG. 1 is block diagram illustrating the use of an external device to inductively couple power into an implant device; 
     FIG. 2 is a block diagram illustrating the use of a stabilization boost regulator as part of the implanted circuits in accordance with the present invention; 
     FIG. 3 shows a functional block diagram of the stabilization boost regulator shown in FIG. 2; 
     FIG. 4 illustrates a functional block diagram of the impedance sensing circuit shown in FIG. 3; 
     FIG. 5 depicts a simplified functional block diagram of the pulse width (PW) and frequency control circuit shown in FIG. 3; 
     FIG. 6 is a timing diagram that illustrates the operation of the stabilization boost regulator shown in FIG. 2 for two different power modes, a low power mode and a high power mode; and 
     FIG. 7 illustrates an unconventional extended control loop that is used with the present invention in order to prevent bistability problems, which extended control loop includes both a fold-over loop, and an under-voltage loop. 
    
    
     Appendix A is a copy of the article entitled “Source resistance: the efficiency killer in DC-DC converter circuits”, previously referenced. 
     Corresponding reference characters indicate corresponding components throughout the several views of the drawings. The present invention addresses the above and other needs by 
     DETAILED DESCRIPTION OF THE INVENTION 
     The following description is of the best mode presently contemplated for carrying out the invention. This description is not to be taken in a limiting sense, but is made merely for the purpose of describing the general principles of the invention. The scope of the invention should be determined with reference to the claims. 
     Turning first to FIG. 1, there is shown a block diagram of a power transfer system  10  that includes an external power transfer system  12  and an implanted device  14 . The external power transfer system  12 , also referred to as the external system  12 , includes a power source  20 , an external device or circuit  22 , and an external coil  24 . As used herein, the word “external” means not-implanted, e.g., on the outside of the skin  30 , of a patient or user. 
     The implanted device  14 , as seen in FIG. 1, includes an implanted coil  32 , an implanted circuit(s)  34 , and a load  36 . As used herein, the term “implanted” means implantable subcutaneously, i.e., placed under the skin  30  of the user or patient. For some embodiments of the invention, it is contemplated that the power transfer system  12  may itself be implanted, wholly or partially, the invention for such embodiment thus being directed to the transmission of power between two implantable devices. 
     For purposes of the present invention, the function performed by the implanted or implantable device  14  is not important. The implant device  14  may perform any desired function, e.g., tissue stimulation, sensing and monitoring physiological parameters, injecting medication into the blood or tissue of the patient, and the like. A preferred application for the invention is for use with an implantable cochlear stimulator (ICS), which ICS provides stimulation to the auditory nerve fibers in the cochlea of the patient as a function of sounds sensed external to the patient. A representative ICS system is illustrated in U.S. Pat. No. 5,603,726, incorporated herein by reference. 
     Regardless of the type of function performed by the implant device  14 , it must receive operating power from the external system  12 . Typically, power is transferred into the implant device  14  via inductive coupling. That is, an ac power signal, generated by the external device  12  is applied to the external coil  24 . This ac power signal induces a corresponding ac power signal in the implanted coil  32  whenever the external coil  24  and the implant coil  32  are sufficiently close to each other so as to permit the alternating magnetic field created by passage of the ac power signal in the external coil  24  to pass through the implanted coil  32 . Such magnetic coupling of two coils is commonly referred to as inductive coupling. 
     The power coupled from the external coil  24  to the implanted coil  32  is represented in FIG. 1 by the straight arrow  38 . The magnitude of the coupled power  38  is predominately a function of the distance between the two coils  24  and  32  as well as the impedance match between the implant device  14  and the external device  12 . 
     The distance between the two coils  24  and  32  is referred to herein as the “implant distance”, and is represented in FIG. 1 as the distance “d”. 
     The input impedance of the implanted device  14  is represented in FIG. 1 by the symbol Z L . The value of Z L  is determined in large part by the value of the load  36  attached to the implanted device  34 , as well as the components used to make the implanted circuit(s)  34 . The value of the load  36  varies significantly from patient to patient, and over time for the same patient, depending upon the operation of the implanted device  34 . Thus, the input impedance Z L  of the implant device  14  is not a constant, as has often been assumed in the past, but is a variable that may vary over time and from patient to patient by as much as a factor of 30 or more. 
     The output impedance of the external system  12  is represented in FIG. 1 by the symbol Z 0 . The value of Z 0  is determined in large part by the circuit components from which the external device  22  is made. 
     It is an object of the present invention to provide, as part of the implanted circuits  34 , an impedance transformer that automatically adjusts the input impedance Z L  of the implanted system  14  so as to provide the lowest power insertion loss of the coil set. Here, the “coil set” refers to the external coil  24  and the implanted coil  32 . The power transfer through the coil set follows the well known relationships of the “double-tuned” network that has been used in radio circuits since its earliest times. In a weakly coupled coil set (coupling coefficient less than 0.2), the wires of the coils contribute to the losses of the power transfer. The parameters that effect the power transfer are: coil geometry (diameters and spacings), operating frequency, numbers of turns in the coils, coil resistance, and source and load resistance. Generally, the coils are of different diameters, and have different numbers of turns. The impedance levels Z 0  and Z L , using the relationships of the double-tuned network are therefore generally different. However, for any given coil set, operated at a given spacing and frequency, there is an optimum Z 0  and Z L  that provide the lowest insertion loss. These values may be calculated from the parameters listed above. Under such minimum insertion loss condition, a maximum amount of power  38  may advantageously be coupled into the implant device. 
     Turning next to FIG. 2, there is shown a block diagram illustrating the use of a stabilization boost regulator  40  as part of an implanted device  14  in accordance with the present invention. The implanted device  14  shown in FIG. 2 comprises an implantable cochlear stimulator (ICS), although this is only exemplary. The implanted device  14  includes an implanted coil  32 , as described in FIG.  1 . The implanted coil  32  interfaces with an external pocket head piece (PHP)  25 , driven by an external pocket speech processor (PSP)  23 . 
     The PSP  23  includes a power source  20 , e.g., a battery, that provides operating power to a boost regulator circuit  21 . The boost regulator circuit  21 , in turn, provides operating power to a suitable transmitter circuit  26 , Tx, that drives the PHP circuitry  27  and external coil  24 , with a suitable modulated ac signal that is coupled into the ICS circuitry  14  through the implanted coil  32 , as described above in connection with FIG.  1 . 
     It should be noted that there is much circuitry included in a typical PSP  23 , PHP  25  and ICS  14  that is not shown in FIG.  2 . Such additional circuitry relates to how audible signals are sensed and converted to appropriate control signals for directing the ICS to stimulate cochlear tissue through a suitable electrode implanted in the cochlea, thereby allowing the user to experience the sensation of hearing through direct electrical stimulation of his/her auditory nerve. Such additional circuitry is not relevant to the present invention, and is therefore not disclosed. (The interested reader can refer, e.g., to U.S. Pat. Nos. 3,751,605; 4,207,441; 4,408,608; 4,428,377; 4,532,930 and 5,603,726, which patents are incorporated herein by reference, for a detailed description of the circuitry associated with and the operation of various types of cochlear implant systems.) Rather, the present invention focuses on the manner in which the implant system  14 , which may be any type of implant system, e.g., a cochlear implant system, may more efficiently receive and process power received from an external source through the use of an implanted time-varying impedance transformer that monitors the load coupled to the implant device, and makes adjustments, as required, to present a more or less constant load to the external circuitry that couples power into the implant device. 
     Thus, in operation, as shown in FIG. 2, the ICS  14  receives operating power from the PSP  23  through the PHP  25 , and more particularly through the coil set that includes the external coil  24  and the implanted coil  32 . Such power, when received at the implanted coil  32  is rectified using rectifier circuit  35 . The rectified power is then available to power the ICS circuitry, including any rechargeable battery that may be included as part of the ICS, or an implanted speech processor, e.g., through a charging port  37 , as well as a stabilization boost regulator circuit  40 . The stabilization regulator circuit  40  provides power for the circuits of the ICS, which circuits are represented in FIG. 2 by the load  36 . For an ICS, the load  36  includes, in addition to most of the signal processing circuitry that receives and processes commands from the speech processor, an electrode array that has been implanted in the cochlea of the user. It is this stabilization regular circuit  40 , for the particular ICS embodiment shown in FIG. 2, that functions as a time-varying impedance transformer in accordance with the principles of the present invention. 
     In the implementation of the invention shown in FIG. 2, the various efficiencies, η, for each stage of the system, are represented. By way of example, the efficiency of the external boost regulator  21  is shown as being less than 0.085 (η&lt;0.85); the efficiency of the external coil  24  is 0.5&lt;η coil &lt;0.7; the efficiency of the rectifier  35  is η rectifier =1/(1+2×V diod /V 0 ); and the efficiency of the stabilization boost regulator  40  is η Boost &lt;0.7. A preferred transfer efficiency through the entire system should approach 70%, although transfer efficiencies much less than 70%, e.g., 35-50%, may still represent a significant improvement over what has been achieved in the past. Proper operation of the circuitry shown in FIG. 2 maintains a proper voltage/current ratio at its input as the output load (V 0 , R 0 , l 0 , P 0 ) varies. 
     Turning next to FIG. 3, a functional block diagram of one embodiment of a stabilization boost regulator circuit  40  is illustrated. It is to be emphasized that the block diagram of FIG. 3, as well as the other block diagrams included in the figures, is functional in nature. Those of skill in the art, given the functional descriptions presented herein, will be able to fashion suitable circuitry, whether dedicated analog, digital, or combinations of analog/digital circuitry, and/or whether implemented using state-diagram driven, or firmware/software controlled circuits, to carry out the indicated functions. 
     The stabilization boost regulator circuit  40  shown in FIG. 3 functions as a switching regulator circuit. As such, it may include, but does not have to include, a step up circuit  42  that selectively multiplies, or steps up, the voltage V IN  of the “Power In” signal by a prescribed amount, as needed in order to allow the output voltage, V 0 , to be within a desired range. That is, the output of the step up circuit  42 , when used, will typically be a voltage having a value of nV IN , where n is an appropriate multiplication factor. Such voltage step up circuits are known in the art. 
     In one specific embodiment, a single-ended primary inductance converter (SEPIC) circuit was employed as the “step up circuit  42 ” that operated continuously between step-up and step-down at its output. In practice, any suitable switching circuit may be used, including step up, step down, or buck/boost circuits. The essential element is that the switching mechanism resemble an electronically-controlled transformer. 
     Following step up circuit  42 , if used, is a switch  42  that switches the signal nV IN  in accordance with a desired duty cycle. That is, the switch  42  is turned ON for a desired portion of a time period T, during which ON time the power signal is passed through the switch  43  to the next functional element in the circuit. When the switch is turned OFF, then the power signal is not allowed to pass through the switch  43 . The result, at the output of the switch  43 , is a pulsed waveform  41  having a series of pulses with a pulse width PW (representing the ON time of the switch  43 ) repeating every T seconds. The frequency, f, of the pulsed waveform is thus 1/T. The duty cycle of the waveform, expressed as a percentage, is 
     
       
         Duty Cycle=( PW )/( T )×100%. 
       
     
     The duty cycle may thus vary from 0% (switch  43  OFF all the time), to 50% (switch  43  ON half of the time and OFF half of the time), to 100% (switch  43  ON all the time), or any other value between 0% and 100%. As the pulsed signal is applied to a suitable filter  44 , e.g., a low pass filter, it is converted to a dc level, or output voltage V 0 , the amplitude of which varies as a function of the duty cycle. Thus, control of the duty cycle of the switched waveform nV IN  controls the amplitude of the output voltage V 0 . 
     Still referring to FIG. 3, the output of the filter circuit  44  is coupled to the load, R L , through a sensing resistor R s , or equivalent current-sensing element. When a sensing resistor R S  is used, it will typically be a very small value, e.g., 1 ohm or less. Such element is used to provide a measure of the output current I 0  that is flowing into the load R L . Any suitable element that senses dc current flow, such as a dc current probe, may be used as the sensing element R S . 
     For the embodiment shown in FIG. 3, the voltage developed across the sensing element R S  provides a measure of the current I 0  flowing therethrough. This value of I 0 , coupled with the output voltage V 0 , is coupled to an impedance sensing circuit  45 . The impedance sensing circuit  45  determines the output impedance based on the measured values of I 0  and V 0  (impedance, Z, is equal to V 0 /I 0 ), and produces a voltage signal V Z  as a function thereof. The signal V Z  is then applied to a suitable pulse-width and frequency (PW and Freq.) control circuit  46  that, in turn, generates the appropriate control signals for controlling the step up circuit  42  and switch  43  so that a desired impedance transformation takes place. 
     Next, with respect to FIG. 4, there is shown one embodiment of a representative functional block diagram of the impedance sensing circuit  45  shown in FIG.  3 . Such impedance sensing circuit  45  includes a differential amplifier  47  that monitors the voltage across the sense resistor R S  and generates an output voltage, VI, as a function thereof. The output voltage V 0  and the current-sense voltage V I , are then applied to a division circuit  48  where the output voltage V 0  is divided by the current-sense voltage V I , to produce the impedance voltage V Z . The impedance voltage V Z  is then applied to the PW and Freq. Control circuit  46  (FIG. 3) so that appropriate adjustments are made to the PW and duty cycle in order to automatically bring about needed impedance transformation whenever a significant change in the power (V 0  and I 0 ) delivered to the load R L  is sensed. In this manner, i.e., with appropriate impedance transformations automatically occurring within the implanted device  14 , the impedance seen by the external source  23  remains relatively constant, thus maintaining the efficiency with which power is coupled into the implanted device  14 . 
     It should also be noted that the impedance transforming process made possible by the invention may also be used to compensate for variations in transfer efficiency that occur due to coil separation, i.e., the separation between the external coil  24  and the implanted coil  32 . 
     Turning next to FIG. 5, a simplified functional block diagram of one embodiment of the pulse width (PW) and frequency control circuit  46  is shown. As seen in FIG. 5, a first differential amplifier  51  compares the impedance voltage V Z  to a first reference voltage V R1 , producing an output voltage V ΔZ  representative of changes in the impedance voltage compared to the first reference voltage V R1 . That is, the output voltage V ΔZ  provides an indication of changes in the output impedance Z 0 =V 0 /I 0 . The voltage V ΔZ  drives a voltage controlled oscillator (VCO)  54 . The VCO  54  generates a pulsed waveform  54   a  having fixed narrow pulses. This waveform  54   a  is applied to a pulse width (PW) control circuit  55 , which converts the narrow pulse widths of the waveform  54   a  to a waveform  55   a  having wider pulse widths that vary as a function of a PW control signal  52   a . The PW control signal  52   a  is generated by a second differential amplifier  52 , or comparator circuit, that compares the output voltage V ΔZ (which represents changes in output impedance) to a second reference signal V R2 , and causes the pulse width to change as a function of the difference between V ΔZ  and V R2 . In this manner, changes in the output impedance signal V Z  relative to a first threshold V R1  affect the frequency of the switch control signal  55   a , and changes in the signal V ΔZ  relative to a second threshold V R2  cause the pulse width (PW) of the switch control signal  55   a  to vary. As the pulse width and frequency of the control signal  55   a  vary, an impedance transformation takes place that makes the coupling of power into the implant device  14  more efficient. 
     In the event that the output voltage V 0  changes a significant amount due to large changes that occur in the power delivered to the load of the implant device, e.g., should changes be needed in the impedance transformation beyond those possible through just frequency and pulse width control, then the multiplication factor n associated with the step-up control circuit  42  (FIG. 3) is adjusted accordingly. Control of the multiplication factor n is triggered by control signal  53   a , generated by a third differential amplifier  53 , or comparator circuit, that compares the change in impedance voltage V ΔZ  to a third reference signal V R3 . 
     Thus, under a low power operating mode, e.g., a mode where the power (V 0 , I 0 ) delivered to the load R L  is relatively low, a typical switch control signal  55   a  may have an amplitude V, a pulse width PW 1 , and a period T 1  that is as shown in FIG. 6, waveform (A). In such mode, the duty cycle of the stabilization boost regulator  40  remains relatively low. Under a high power operating mode, e.g., a mode where the power (V 0 , I 0 ) delivered to the load R L  is relative high, a typical switch control signal  55   a  may have an amplitude nV, a pulse width PW 2 , and a period T 2  that is as shown in FIG. 6, waveform (B). In such mode, the duty cycle of the stabilization boost regulator  40  may be relatively high. In either operating mode, however, the impedance of the implant device  14 , as seen by the external charging device  12 , may remain relatively constant, thereby promoting the efficient coupling of power into the implant device. 
     It should be noted that the changes in frequency that occur in the control signal  55   a  (e.g., as the period varies from T 1  to T 2 , FIG. 6) are not related to the frequency of the rf signal that is inductively coupled form the external coil  24  to the implanted coil  32 . Such rf frequency of the coupling signal may remain at a constant frequency, e.g., 49 MHz, thereby simplifying the external circuitry and minimizing regulatory requirements associated with variable frequency transmissions. 
     Other embodiments of the invention may also be used. Further, as those of skill in the art will recognize, the use of a switching regulator circuit may easily lead to instability problems. That is, switching regulators are prone to “bistability”, as discussed in article: “Source resistance: the efficiency killer in DC-DC converter circuits”, previously referenced. Such article discusses, intra alia, the common design criteria and design problems associated with DC-DC converters. 
     In order to avoid the bistability problems that have plagued prior art designs, an unconventional extended control loop may be employed. Such extended control loop prevents bistable runaway, and allows full operation of the circuitry (FIG. 3 et seq.) described herein. Basically, such extended control loop includes a fold-over loop, and an under-voltage loop, which in combination prevent the switching regulator from running away, and tell the RF source to increase or decrease its output level. A block diagram of this approach is shown in FIG.  7 . 
     As seen in FIG. 7, the system includes external components  58  and an implanted unit  60 , e.g., an ICS. The external components are typically realized in a BTE unit coupled to a pocket head piece, or PHP, and include an RF signal source coupled to a transmitter  72  through a modulator  74 . The transmitter  72  drives the external coil  24 . A back telemetry (BT) receiver  75 , also included in the external components  58 , receives back telemetry signals from the implanted unit  60 , which signals are inputted to a processor  77 . The processor  77 , in turn, processes the signals in an appropriate manner so as to use the information contained therein to modulate the RF power signal that is sent to the transmitter  72 . 
     The implanted unit  60  includes the implanted coil  32 . Signals received through the implanted coil  32  are rectified and filtered using the rectifier and filter circuits  62  to create an input voltage V IN . The input voltage V IN , as well as an input current, I IN , are applied to a switching regulator  64 . An output voltage (V O ) setpoint register  65  provides a reference voltage, or signal, to the switching regulator  64  that defines the desired output voltage, V O . The switching regulator thus generates the output voltage V O  at a certain power level, P O , into the ICS load R L . 
     A voltage control loop, comprising a buffer amplifier driving a resistor R 1  provides a feedback (FB) signal to the switching regulator. A difference amplifier  67  compares a reference voltage from the switching regulator  64  with the buffered output voltage (output of buffered amplifier  66 ) to generate an error voltage, V O  Error, on signal line  68 . This signal is compared with an Rin Error signal, generated as discussed below, and then integrated through the use of an integration circuit  69  to create a Power Error signal. The Power Error signal is applied to a back telemetry transmitter circuit (BT TX)  70 , from where it is transmitted to the back telemetry receiver  75  in the external BTE/PHP unit  58 . The external unit  58  uses the received Power Error signal to adjust the level of the input power transmitted to the ICS. 
     The V IN  and I IN  signals that are applied to the switching regulator  64  are also applied to a voltage/current divider circuit  71 . The ratio of the V IN  and I IN  signals (i.e., the input voltage V IN  divided by the input current I IN ) provides a measure of the input impedance, or input resistance, R IN , of the implanted ICS unit  60 . Such measure of input resistance R IN  is compared with an optimum resistance R OPT  by a summer circuit  73 , which summer circuit  73  subtracts R OPT  from R IN  to arrive at an R IN  Error signal. As indicated above, the R IN  Error signal is compared with the V O  Error signal, in a summer circuit  75 , to create the Power Error signal that is sent by back telemetry to the external BTE unit  58 . 
     As further shown in FIG. 7, the R IN  Error signal is integrated, by integrator circuit  76  and then applied through diode D 1  to the feedback (FB) signal applied to the switching regulator. The output voltage V O  portion of the FB signal is derived from a voltage control loop made up of the buffer amplifier  66  and resistor R 1 . The integrated R IN  Error signal portion of the FB signal is derived from a foldback loop made up of the V/I divider circuit  71 , the summer  73 , the integrator circuit  76  and the diode D 1 . Advantageously, use of a voltage control loop and a foldback loop as shown in FIG. 7 may be used to avoid bistability problems in the operation of the switching regulator circuit, when needed. 
     From the above, it is thus seen that the present invention provides an implantable medical device, e.g., an implantable cochlear stimulator or other implantable neural stimulator, that employs a switching regulator circuit as part of the implanted circuitry. Advantageously, the switching regulator is controlled, as energy is inductively coupled into the implant circuitry through a coil set that includes an external coil and an implanted coil, to operate as a varying impedance transformer. More particularly, the impedance is varied so as to make mismatch losses as seen at the power source appear substantially constant. In turn, a constant mismatch loss allows an optimum power transfer efficiency (minimum power insertion loss) to occur. 
     It is further seen that the invention provides an implantable time-varying impedance transformer wherein there are no circuit value or wiring changes needed to handle varying output load impedance. Rather, all control is entirely electronic and preferably automatic. 
     It is also seen that the invention may operate using a carrier signal (the signal applied to the external coil) having a fixed frequency, thereby avoiding regulatory or other problems incident to using variable frequency carrier signals. 
     It is additionally seen that an advantage of the invention is that the effective DC load resistance of the implanted circuitry (output voltage divided by output current) is transformed to effect an AC circuit mismatch loss, so that the lowest insertion loss of the coil set (i.e., the highest power transfer efficiency between the external coil and implanted coil) may be utilized. 
     While the invention herein disclosed has been described by means of specific embodiments and applications thereof, numerous modifications and variations could be made thereto by those skilled in the art without departing from the scope of the invention set forth in the claims.