Active feed forward power control loop

Details Active feed forward power control loop Active feed forward power control loop

2000-01-14

2002-07-02

Jastrzab, Jeffrey R. (Department: 3737)

Surgery: light, thermal, and electrical application

Light, thermal, and electrical application

Electrical therapeutic systems

C607S061000

Reexamination Certificate

active

06415186

ABSTRACT:

BACKGROUND OF THE INVENTION
The present invention relates to the efficient utilization of power in body tissue stimulators, and more particularly to an improved power control loop for implantable cochlear stimulator systems. Such implantable cochlear stimulator systems provide improved hearing for the hearing impaired. The power control loop serves the important function of providing power to the implanted part of the cochlear stimulator system and the efficiency of the power control loop is critical in developing the miniaturized systems of the future.
U.S. Pat. No. 4,400,590 issued Aug. 23, 1983 for “Apparatus for Multi-Channel Cochlear Implant Hearing Aid System” describes and illustrates a system for electrically stimulating predetermined locations of the auditory nerve within the cochlea of the ear, which system includes a multi-channel intra-cochlear electrode array. The electrode array comprises a plurality of exposed electrode pairs spaced along and embedded in a resilient curved base for implantation in accordance with the method of surgical implantation described in U.S. Pat. No. 3,751,605 issued Aug. 7, 1973 for “Method of Inducing Hearing.” The hearing aid system described in the '590 patent receives audio signals at a signal processor located outside the body of a hearing impaired patient. The processor converts the audio signals into analog data signals which are transmitted by a cable connection through the patient's skin to the implantable multi-channel intra-cochlear electrode array. The analog signals are applied to selected ones of the plurality of exposed electrode pairs included in the intra-cochlear electrode array to electrically stimulate predetermined locations of the auditory nerve within the cochlea of the ear where the intra-cochlear electrode array is positioned.
The cochlea stimulating system described in the '590 patent is limited in the number of channels of information, the speed of transfer of stimulating signals to the cochlea and the fidelity of the signals. Also, the cable connection through the skin of the patient to the intra-cochlear electrode array is undesired in that it interferes with the freedom of movement of the patient and represents a possible source of infection.
U.S. Pat. No. 4,532,930, issued Aug. 6, 1985 for “Cochlear Implant System For an Auditory Prosthesis” describes and illustrates a multiple electrode system which does not employ a through the skin connector. While multiple electrodes are employed to stimulate hearing, the system only operates with a single pulsatile output stimulating a single electrode channel at any given time. Such a sequential system is limited in speed of operation, and does not provide for analog operation where continuous stimulating signals, controllable in amplitude, are simultaneously applied to a number of electrode channels. Further, the system is subject to charge imbalance with misprogramming or circuit fault and inefficient use of electrical power. Moreover, once the stimulator unit is implanted there are no means for monitoring its ongoing circuit operation or power requirements so as to optimize its continued operation.
U.S. Pat. No. 4,592,359, issued Jun. 3, 1986 for “Multi-Channel Implantable Neural Stimulator” describes a cochlear implant system having 4 current sources and 4 current sinks per channel, controlled by series switches, to provide 16 different circuits for supplying 16 levels of 2 polarities to each output channel. In a pulsatile mode, the system provides for simultaneous update (amplitude control) and output to all channels. However, the system does not permit simultaneous analog update and output on all channels and the electrode pairs for each channel are not electrically isolated from all other electrode pairs whereby undesired current leakage may occur. Further, once the stimulator is implanted there are no means for monitoring its ongoing circuit operation or power requirements so as to optimize its continued operation.
U.S. Pat. No. 4,947,844, issued Aug. 14, 1990 for “Receiver/Stimulator For Hearing Prosthesis” describes and illustrates a multiple channel electrode system. The system includes an implantable receiver/stimulator connected to an implantable electrode array. Included in the implantable receiver/stimulator is a transmitter for telemetering one electrode voltage, measured during stimulation, to an external receiver for monitoring and analysis as an indicator of proper operation of the implantable stimulator. The transmitter comprises an oscillator operating at a frequency of about 1 MHz. The output of the oscillator is coupled to the implant's receiving coil. The oscillator signal, when received after transmission, is demodulated to recover the selected voltage waveforms. Unfortunately, such a telemetry system is not only limited to the monitoring of one voltage, but the simultaneous transmission of the telemetry signal and reception of the input carrier signal results in undesired modulation and possible loss of input data.
For cochlear stimulator applications, it is generally desirable to employ a cochlear stimulator that is driven by a behind-the-ear speech processor, e.g., of the type described in U.S. Pat. No. 5,824,022 issued Oct. 20, 1998 for ‘Cochlear stimulation system employing behind-the-ear speech processor with remote control.’ Behind-the-ear speech processors offer several advantages, but because of their small size are limited in the size of the battery they may carry (which in turn limits the useful life of the battery.) The small battery size results in a requirement for very low power dissipation. Although low power digital electronics have enabled digital hearing aids, this technology is only part of the answer for implantable stimulators. This is because an implantable stimulator, e.g., a cochlear stimulator, requires additional variable current to stimulate the target tissue, e.g., the auditory nerve within the cochlea, and this power must be transferred across a transcutaneous link, that is, at best, only about 50% efficient. While digital hearing aids only need to drive a transducer that uses less than one milliwatt (mW) of power, an implantable tissue stimulator may require up to 50 mW of stimulus power, which means (assuming a 50% transcutaneous link transfer efficiency) the need to transmit up to 100 mW of power to the implant device. Since power is proportional to the square of the voltage, it would thus be desirable to have a way to precisely and actively control the voltage in the implant device to track the output power requirements of the device. For example, for a cochlear stimulator, where room sound and speech levels are variant, it would be desirable to track speech and system variations and make automatic adjustments in the input power that track these variations, thereby only transmitting power to the implant device that is needed for the current conditions, thereby increasing the life of the battery.
U.S. Pat. No. 5,876,425, issued Mar. 2, 1999 for “Power Control Loop for Implantable Tissue Stimulator” describes a feedback power control loop utilizing back telemetry from the implantable device. The Implantable Cochlea Stimulator (ICS) utilizes a tank capacitor as an internal rechargable power source. The ICS monitors the voltage level of the tank capacitor and back transmits the tank capacitor voltage to the Wearable Signal Receiver and Processor (WP). Based on the back transmitted tank capacitor voltage, the WP computes the power level to be transmitted to the ICS to maintain the tank capacitor voltage within acceptable levels. While the approach taught in the '425 patent provides advantages over previous approaches that transmit power based on the peak ICS power requirement, it also results in delays in the calculation of the power requirements. The delays in the response of the power control loop result in too much power being transmitted at times, and this power is dissipated versus being stored. The requirement to continuously monitor the tank capacitor voltage requires that the impl

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