Abstract:
In a cochlear implant system, the implantable stimulator includes a monitor which monitors parameters associated with the stimulation signals and/or the power stored in an energy storage element which stores energy transmitted from the processor. This parameter or parameters is/are analyzed and one or more feedback signals are generated and transmitted back to the processor. The processor uses the feedback signal to insure that power is transmitted to the stimulator optimally and that the stimulation signals are compliant.

Description:
BACKGROUND OF THE INVENTION 
     A. Field of Invention 
     This invention pertains to an optimization circuit in a cochlear implant system and more particularly to a circuit which monitors one or more parameters within the implant such as the internal power supply level and the compliance of the stimulation signals applied by the implant. If an undesirable condition is indicated by these parameters, the circuit generates control signals to correct the condition by adjusting the coupling between the internal and external components of the system. 
     B. Description of the Prior Art 
     Certain patients suffer from a hearing disability in the inner ear which cannot be satisfactorily assisted by normal hearing aids. However, if the aural nerve is intact, the patient may have some aural functions restored with a cochlear implant system. A typical cochlear implant system presently available includes an external component or processor and an internal component often called the implanted stimulator. The external component includes a microphone for receiving ambient sounds and converting them into electrical signals, a processor for processing said electrical signals into encoded signals and a transmitter transmitting said encoded signals to the internal component. 
     The internal component includes a receiver receiving the encoded signals, a decoder for decoding said signals into stimulation signals and an electrode array including both intracochlear electrodes extending into the patient&#39;s cochlear and optionally one or more extra-cochlear electrodes. The stimulation signals are applied in the form of pulses having durations and waveshapes determined by the processor. 
     Because the internal component of the cochlear implant system is relatively small, it is not normally provided with its own permanent power supply. Instead, the internal component is energized transcutaneously by RF signals received from the external component with the use of two inductively coupled coils, one provided in the external component and the other being provided within the internal component. The external component sends data to the internal component, by first encoding the data into the RF signals and then transmitting it across the transcutaneous link. The internal component decodes the data from the received RF signals and also stores the received RF energy in a capacitor to power its electronics. In order to achieve efficient power transfer across the transcutaneous link, both coils are tuned to resonate, at or close to the operating frequency or the transmitter and are held in axial alignment with the aid of a magnetic coupling. 
     The amount of energy being transferred to the internal component depends mainly on the amount of inductive coupling between the two coils as well as the resonance frequency of the respective coils. The former is dependent on the thickness of the tissue separating the two coils, which thickness varies over the patient population. Hence, for identical cochlear implant systems the efficiency of energy transfer varies from one patient to another. 
     The required amount of energy varies with the patient, (due to the electrode-tissue interface impedance being patient specific) the system programming, and the sound environment. Therefore, every cochlear implant system must be designed so that adequate power is delivered to the internal component for all patients under all conditions. Hence, there is an excess energy transfer across the link for patients with relatively smaller separation between the coils, or a low electrode-tissue interface impedance, resulting in a shorter battery life, than optimally desired. 
     Attempts have been made by others to resolve this problem but they have not been entirely satisfactory. For example, U.S. Pat. No. 5,603,726 discloses a multichannel cochlear implant system in which the implantable section generates signals to a wearable processor indicative of the status of the implantable section, such as its power level and stimulation voltages. The information is used by the wearable processor to modify the characteristics of the signals transmitted. More particularly, the implantable section has an internal power supply capable of producing several outputs having different nominal DC levels. Additionally, the implantable section is also capable of providing unipolar or bipolar stimulation pulses between various intercochlear electrodes as well as an indifferent electrode. A telemetry transmitter is used to send data to the wearable processor, the data being indicative of the voltage levels of the power supply outputs, the amplitudes of the stimulation signals and other parameters. The wearable processor uses the power level signals to adjust the amplitude (and therefore the power) of the RF signals transmitted to the implantable section. However, this approach is disadvantageous because it requires an RF transmitter having a variable programmable amplitude, and utilizes a fixed tuning of the transmit coil, therefore making no attempt to modulate the voltage on the tank capacitors to track the voltage required to maintain system compliance. Obviously such a transmitter is expensive to make and more complex then a standard RF transmitter having a preset amplitude. Moreover, sending information from the implantable section about the amplitude of the stimulation pulses after these pulses have already been applied is ineffective because, if one of these pulses is out of compliance, the external section can do nothing about it, except crank up the power to insure that future pulses are compliant. However, merely cranking the power, without any further intelligence wastes energy. 
     Commonly assigned application Ser. No. 09/244,345 filed Feb. 4, 1999 entitled HIGH COMPLIANCE OUTPUT STAGE FOR A TISSUE STIMULATOR, incorporated herein by reference, describes a cochlear implant system wherein the generation of stimulation pulses is monitored, (i.e. the compliance of the stimulation generation circuit) and a voltage multiplier is used if necessary to ensure that the stimulation pulses are or the desired intensity. This application essentially deals with a system of improving the internal power supply in order to eliminate stimulation pulses, and as such, there is no provision in this application for transmission of data back to the external section. 
     OBJECTIVES AND SUMMARY OF THE INVENTION 
     In view of the above disadvantages of the prior art, it is an objective of the present invention to provide a power control circuit for a cochlear implant which is constructed and arranged to automatically and dynamically optimize the power transferred to the internal component based on one or more preselected criteria by adjusting an inductive coupling therebetween. 
     A further objective is to provide a power control circuit for a cochlear implant which is constructed and arranged to automatically and dynamically regulate the inductive coupling with the internal component thereof to insure that power is not wasted, thereby increasing the life of the external component battery. 
     Other objectives and advantages of the invention shall become apparent from the following description. 
     Briefly, a cochlear implant system constructed in accordance with this invention includes an external speech processor and an implantable stimulator having electronic circuitry, the two components being coupled to each other inductively by respective coils. Each coil is part of a tank circuit. The external speech processor transmits RF signals through the coupling. The implantable stimulator uses these signals for two purposes. First, the energy of the signals is stored in a storage element such as a capacitor and used to power the electronic circuitry. Second, the signals are decoded and used to derive the stimulation signals applied to the aural nerve. 
     In one embodiment of the invention, a parameter indicative of the voltage of the storage element is monitored and sent back to the speech processor via a secondary channel. The external speech processor then adjusts the frequency of its tank circuit to regulate the power transferred to the internal component to optimize it. 
     Additionally, or alternatively, the compliance of the stimulation signals is monitored and used as a feedback signal to control the frequency of the tank circuit to optimize power transfer to the internal component. This adjustment can be done either based on statistical basis, or in response to an individual and specific out of compliance condition. 
    
    
     BRIEF DESCRIPTION OF THE DRAWINGS 
     FIG. 1 shows a schematic diagram of a cochlear system constructed in accordance with the present invention; 
     FIG. 2 shows a schematic diagram of the external component of the cochlear system of FIG. 1; 
     FIG. 3 shows a schematic diagram of the internal component of the cochlear system of FIG. 1; 
     FIGS. 4A,  4 B and  4 C show the power control signals transmitted from the internal to the external components respectively to indicate the power level induced within the internal component; 
     FIGS. 5A and 5B show flow charts for the operation of internal and external components of FIGS. 1-3, respectively; and 
     FIG. 6 shows two sets of typical biphasic stimulation signals defined by the speech processor; 
     FIG. 7 shows the current pulses required to produce the stimulation pulses of FIG. 6; and 
     FIG. 8 shows the corresponding waveforms across the current source. 
    
    
     DETAILED DESCRIPTION OF THE INVENTION 
     Referring first to FIG. 1, a cochlear implant system  10  constructed in accordance with this invention includes an external component  12  and an internal component  14 . The external component includes a speech processor  12 A and is associated with a microphone  16  for sensing ambient sounds and generating corresponding electrical signals. These signals are sent to the speech processor  12 A which processes the signals and generates corresponding encoded signals. The encoded signals are provided to a transmitter (including a transmit coil  20 ) for transmission to the internal component  14 . 
     The internal component  14  (which may also be referred to as an implantable stimulator) receives the power and data via a receive coil  22 . The RF power signal is stored by a power supply  24  (See FIG. 3) which provides power for the internal component  14 . The data signals control the operation of the internal component  14  so as to generate the required stimulation pulses which are applied to the auditory nerve of the patient via an electrode array  28 . 
     The structure of the external speech processor  12 A is shown in more detail in FIG.  2 . First, the audio signals received from microphone  16  are fed to a signal processor  30 . This signal processor  30  maps the audio signals into excitation signals in accordance with one or more mapping algorithms stored in a map memory  31 . These excitation signals are encoded by a digital data encoder  34 . The encoder data is combined with an RF signal in the data and power transmitter  36 , and passed to the transmit coil  20  via a tuneable tank circuit  38 . 
     In accordance with the present invention, encoded telemetry data is received back from the internal component  14  via coil  20 , and is decoded by telemetry decoder  52 . The decoder telemetry data is passed to the tuning adjuster and power controller  40 , which uses the telemetry data to generate a tuning adjustment signal. The tuneable tank circuit  38  adjusts the tuning of the transmit coil  20  according to the tuning adjustment signal as described in more detail below. This can be achieved, or example, by using an electrically controlled variable capacitor in conjunction with a series tuning capacitor, or by any of various similar means known to the art. Power to the whole system  10  is provided by a power supply  50  which typically includes a battery. 
     Referring now to FIG. 3, the internal component  14  includes a housing (not shown) which is hermetically sealed. The component  14  also includes a receiver tank circuit  32  having the receive coil  22  and a capacitor  66 . Signals received through this tank circuit are fed to a power supply  24  generating an output voltage Vdd. The power supply is represented in FIG. 3 by a diode  68  charging a capacitor  70 . The power supply  24  uses the energy of the received RF signals to charge up the capacitor  70 . 
     The RF signals are also fed to a data decoder  60 . The data decoder  60  derives from the RF signal the digital excitation signals generated by the data encoder  34  and generates corresponding stimulation control signals. These signals are fed to a programmable current source  62  and a switching control circuit  64 . These two circuits cooperate in response to the signals from data decoder  60  to apply the cochlear stimulation signals to predetermined electrodes or electrode array  28  in a known manner which is beyond the scope of this invention. 
     Implant  14  further includes a compliance monitor  66  which generates an output that is fed to a telemetry encoder  8  as discussed more fully below; and a power supply monitor  82  which is used to monitor the voltage Vdd generated by power supply  24  and which provides a voltage condition signal to telemetry encoder  80 . 
     The compliance monitor  66  and power supply monitor  82  each sense certain specific functions of the internal component and transmit them to the telemetry encoder  80 . The telemetry encoder  80  then transmits this information to the telemetry decoder  52 . The data is decoded and used to adjust the power transmit between the coils, if necessary. 
     An exemplary mode of operation indicating the voltage monitoring made is now described in conjunction with FIGS. 4A, B and C and  5 A and  5 B. At predetermined intervals, for example, every 100 ms, or alternatively after every stimulation pulse, the telemetry encoder  80  generates a first pulse F. (Step  100 ). This pulse may have a duration of about 1 ms. This pulse F indicates to the external speech processor  12 A that the implantable stimulator  14  is sending data. 
     Next, the power supply monitor  82  compares the power supply output voltage Vdd to a threshold value Vt and sends the result to the telemetry encoder  80 . More specifically, starting with step  102 , the power supply monitor  82  first determines if Vdd&gt;Vt. If it is, then in step  104 , a parameter pw (pulse width) is set to a predetermined value A, of for example, 2 ms by the telemetry encoder  80 . 
     If in step  102  Vdd is not larger than Vt then in step  106  a check is performed to determine if Vdd is approximately equal to Vt. If it is, then in step  108  parameter pw is set to zero. Tf it is not then, Vdd must be smaller than Vt and in step  110  the parameter pw is set to a predetermined value B of, for example, 1 ms. 
     Next, in step  112  a pulse D is generated having a pulse width A or B, or no pulse is generated, depending on the outcome of the decisions  102  and  106 . The pulse D (if present) is generated a period T after pulse F. T may be about 1 ms. The results of this step are seen in FIGS. 4A,  4 B,  4 C. 
     For FIG. 4A it has been determined that Vdd&gt;Vt, and hence pulse D with a pulse width A is sent about 1 ms after pulse F. 
     In FIG. 4B, Vdd has been found to be about equal to Vt and hence no pulse D is present. 
     In FIG. 4C, Vdd is found to be smaller that Vt and hence pulse D having a pulse width B is sent about 1 ms after period F, pulse width B being generally shorter than pulse width A. For example, pulse width A may be 2 ms and pulse width B may be about 1 ms. 
     Pulse F and, if present, pulse D are then sent to the tank circuit  32 . As a result, a corresponding signal appears on the transmit coil  20 , which is then decoded by the telemetry decoder  52 . 
     The operation of the telemetry decoder  52  is now described in conjunction with FIG.  5 B. Starting with step  120 , a pulse F is first detected which indicates that the power supply monitor  82  is sending information about the status of the power supply  24 . Next in step  122  a check is made to determine if a pulse D is present following pulse F. If this pulse is not detected, then in step  130  the previous operations are continued with no change. 
     If in step  122 , a pulse D is detected then in step  124  a determination is made as to whether this pulse D has a pulse width A or a pulse width B. A telemetry pulse D having a relatively long pulse width, in a range corresponding to the pulse width A (for example if pulse D exceeds 1.5 ms), indicates that the implant supply voltage is high (i.e. Vdd&gt;Vt). In step  126 , the tuning adjuster and power controller  40  therefore adjusts the tunable tank circuit  38  to reduce the power transferred to the implant. A preferred method to accomplish this effect is to reduce the resonance frequency of the tank circuit. 
     If the telemetry pulse is less than 1.5 ms, (indicating a pulse width B and that the power supply Vdd&lt;Vt) then in step  128  the tuning adjuster and power controller  40  adjusts the tunable tank circuit  38  to increase the transferred power. 
     The tunable tank circuit  38  is adjusted by the tuning adjuster and power controller  40  via means of a tuning capacitor (not shown) which is preferably a voltage dependent capacitor. It should be appreciated that the tunable tank circuit  38  could also be tuned by other known means as would be understood by one skilled in the art. 
     Similarly, the above mentioned operation may be performed in respect of the compliance monitor signal, as described in more detail below. 
     Briefly, referring to FIG. 3; under the control of commands from data decoder  60 , the programmable current source  62  generates current pulses which are applied to the electrodes by switching control circuit  64 . FIG. 6 depicts two typical stimulation current waveforms  70  and  73  which may be requested by the signal processor  30 . It can be seen that each waveform is biphasic, consisting of two current pulses of equal amplitude and opposite polarity. Thus, lower amplitude biphasic current waveform  70  consists of positive and negative pulses  71  and  72  respectively, and higher amplitude current waveform  73  consists of positive and negative pulses  74  and  75 . 
     Next, FIG. 7 depicts the corresponding current waveforms that must be generated by the programmable current source  62  to produce the desired stimulation current waveforms  70  and  73 . That is, the programmable current source  62  must generate two lower amplitude square waves  76  and  77  to generate stimulus pulses  71  and  72  respectively, and two larger amplitude square waves  78  and  79  to generate the stimulus pulses  74  and  75 . Pulses  77  and  79  are reversed by the switching control circuit  64 . However, if the current pulses  78  and  79  exceed the capability of the power supply  24 , an out of compliance condition occurs. This problem is resolved in the present invention as follows. 
     Referring to FIG. 8 the voltage waveform  80  represents the voltage Vn at the output of the programmable current source  62 . It can be seen from the shape of the voltage waveform  80  that the load contains a capacitive component. The level Vc marks the minimum voltage across the programmable current source  62  at which compliance with the desired current waveforms of FIG. 7 can be maintained. The voltage Vca is a little higher than Vc as shown and is selected to provide a safety margin. As seen in FIG. 8, pulse  83  required to generate pulses  78  and  74  of FIGS. 7 and 6 respectively, starts off at a level above Vca but decreases linearly toward a minimum value (P) which is substantially below level Vc and therefore is not attainable. When this pulse reaches Vca (at point  85 ), the compliance monitor  66  generates a compliance monitor signal indicating an out of compliance condition. The signal is encoded by the telemetry encoder  80  and transmitted to the external processor. The signal may be the same signal as when VDD drops below VT as discussed above, or it may be a different signal, as would be appreciated by one skilled in the art. In response, the tuning adjuster and power controller commands the tunable tank circuit  38  to increase the voltage transmitted to the internal section. 
     The adjustment of the link tuning or RF power generated can be performed for every instance of a compliance monitor signal being received from the implant and may be maintained at a high level for a predetermined time, after which the RF power can be dropped to a previous level. 
     Alternatively, the frequency of the compliance monitor signal may be monitored by the tuning adjuster and power controller  40 . The link tuning or RF power generated could then be adjusted to maintain a desired ratio of compliance monitor signals to stimulation signals. For example, the link tuning or RF power generated could be adjusted to keep the ratio of compliance monitor signals to stimulation pulses to a desired target of for example 5%, i.e. For this purpose, the tuning adjuster and power controller  40  includes a counter which counts every instance of non-compliance. After a predetermined number of stimulation pulses, for example a thousand, the counter is checked to determine the number of non-compliant instances. If the counter shows a number over the desired target (i.e. 50 for a 5% target) then the tuning adjuster and power controller  40  adjusts the tank circuit  38  to increase its power level. On the other hand for a number of non-compliant instances below the target, the power level is increased. Of course, this determination could also be made within the implant by the compliance monitor itself, as would be evident to one skilled in the art. 
     Obviously numerous modifications can be made to the invention without departing from its scope as defined in the appended claims.