Patent Application
20160279427
The present invention relates generally to an external charger used to inductively charge a rechargeable battery within an implantable medical device such as a neurostimulator.
Implantable stimulation devices generate and deliver electrical stimuli to nerves and tissues for the therapy of various biological disorders, such as pacemakers to treat cardiac arrhythmia, defibrillators to treat cardiac fibrillation, cochlear stimulators to treat deafness, retinal stimulators to treat blindness, muscle stimulators to produce coordinated limb movement, spinal cord stimulators to treat chronic pain, cortical and deep brain stimulators to treat motor and psychological disorders, occipital nerve stimulators to treat migraine headaches, and other neural stimulators to treat urinary incontinence, sleep apnea, shoulder sublaxation, etc. The present invention may find applicability in all such applications and in other implantable medical device systems, although the description that follows will generally focus on the use of the invention in a Bion™ microstimulator device system of the type disclosed in U.S. Patent Application Publication 2010/0268309.
Microstimulator devices typically comprise a small generally-cylindrical housing which carries electrodes for producing a desired stimulation current. Devices of this type are implanted proximate to the target tissue to allow the stimulation current to stimulate the target tissue to provide therapy for a wide variety of conditions and disorders. A microstimulator usually includes or carries stimulating electrodes intended to contact the patient's tissue, but may also have electrodes coupled to the body of the device via a lead or leads. A microstimulator may have two or more electrodes. Microstimulators benefit from simplicity. Because of their small size, the microstimulator can be directly implanted at a site requiring patient therapy.
The battery 145 supplies power to the various components within the microstimulator 100, such the electrical circuitry 144 and the coil 147. The battery 145 also provides power for therapeutic stimulation current sourced or sunk from the electrodes 142. The power source 145 may be a primary battery, a rechargeable battery, a capacitor, or any other suitable power source. Systems and methods for charging a rechargeable battery 145 will be described further below.
The coil 147 is configured to receive and/or emit a magnetic field that is used to communicate with, or receive power from, one or more external devices that support the implanted microstimulator 100, examples of which will be described below. Such communication and/or power transfer may be transcutaneous as is well known.
The programmable memory 146 is used at least in part for storing one or more sets of data, including electrical stimulation parameters that are safe and efficacious for a particular medical condition and/or for a particular patient. Electrical stimulation parameters control various parameters of the stimulation current applied to a target tissue including, but not limited to, the frequency, pulse width, amplitude, burst pattern (e.g., burst on time and burst off time), duty cycle or burst repeat interval, ramp on time and ramp off time of the stimulation current, etc.
The illustrated microstimulator 100 includes electrodes 142-1 and 142-2 on the exterior of the capsule 202. The electrodes 142 may be disposed at either end of the capsule 202 as illustrated, or placed along the length of the capsule. There may also be more than two electrodes arranged in an array along the length of the capsule. One of the electrodes 142 may be designated as a stimulating electrode, with the other acting as an indifferent electrode (reference node) used to complete a stimulation circuit, producing monopolar stimulation. Or, one electrode may act as a cathode while the other acts as an anode, producing bipolar stimulation. Electrodes 142 may alternatively be located at the ends of short, flexible leads. The use of such leads permits, among other things, electrical stimulation to be directed to targeted tissue(s) a short distance from the surgical fixation of the bulk of the device 100.
The electrical circuitry 144 produces the electrical stimulation pulses that are delivered to the target nerve via the electrodes 142. The electrical circuitry 144 may include one or more microprocessors or microcontrollers configured to decode stimulation parameters from memory 146 and generate the corresponding stimulation pulses. The electrical circuitry 144 will generally also include other circuitry such as the current source circuitry, the transmission and receiver circuitry coupled to coil 147, electrode output capacitors, etc.
The external surfaces of the microstimulator 100 are preferably composed of biocompatible materials. For example, the capsule 202 may be made of glass, ceramic, metal, or any other material that provides a hermetic package that excludes water but permits passage of the magnetic fields used to transmit data and/or power. The electrodes 142 may be made of a noble or refractory metal or compound, such as platinum, iridium, tantalum, titanium, titanium nitride, niobium or alloys of any of these, to avoid corrosion or electrolysis which could damage the surrounding tissues and the device.
The microstimulator 100 may also include one or more infusion outlets 201, which facilitate the infusion of one or more drugs into the target tissue. Alternatively, catheters may be coupled to the infusion outlets 201 to deliver the drug therapy to target tissue some distance from the body of the microstimulator 100. If the microstimulator 100 is configured to provide a drug stimulation using infusion outlets 201, the microstimulator 100 may also include a pump 149 that is configured to store and dispense the one or more drugs.
Turning to
An external charger 151 provides power used to recharge the battery 145 (
It is generally desirable to charge the battery 145 as quickly as possible to minimize inconvenience to the patient. One way to decrease charging time is to increase the strength of the magnetic charging field by increasing the excitation current in the coil 157 of the external charger. Increasing the charging field will increase the current/voltage induced in the coil 147 of the microstimulator 100, which increases the battery charging current, Ibat. However, the strength of the magnetic charging field can only be increased so far before implant heating becomes a concern. One skilled in the art will understand that implant heating is an inevitable side effect of charging using magnetic fields. Heating can result from several different sources, such as eddy currents in conductive portions of the implant, or heating of the various components in the power circuitry 160. Implant heating is a serious safety concern; if an implant exceeds a given safe temperature (e.g., 41° C.), the tissue surrounding the implant may be aggravated or damaged.
The art has recognized that heating can be controlled by controlling the intensity of the magnetic charging field produced at the external charger 151. For example, the current flowing through charging coil 157 can be reduced to reduce the temperature of the implant during charging. The art has also recognized that heating can be regulated by duty cycling the charging field, i.e., by turning the charging field at the external charger 151 on and off.
While changing the intensity or duty cycling of the magnetic charging field produced by the external charger 151 can be an effective means of controlling implant temperature, the inventors have realized that such approaches do not adequately address important issues. First, known prior approaches do not address whether the magnetic charging field intensity, duty cycle, or both, should be modified as a means of temperature control. Moreover, such prior techniques are not understood to consider efficient charging of the implant battery 145. Thus, one can change the intensity and/or duty cycle of the magnetic charging field to arrive at suitable temperature control, but the particular parameters chosen may provide a charging power to the battery that is unnecessarily low, which would prolong charging. Prolonged charging is inefficient, because that patient must wait an inordinate amount of time to fully charge the battery 145 in his or her implant. Understandably, patients do not desire charging to take any longer than necessary.
Finding optimal charging conditions (intensity, duty cycle) thus remains unknown with such prior art techniques, and this disclosure presents a technique to combat this problem, and to make charging more efficient from both a time and implant heating perspective.
An improved external charger for a battery in an implantable medical device (implant), and technique for charging the battery using such improved external charger, is disclosed. In one example, simulation data is used to model the power dissipation of the charging circuitry in the implant at varying levels of implant power. A power dissipation limit is chosen to constrain the charging circuitry from producing an inordinate amount of heat to the tissue surrounding the implant, and duty cycles are determined for the various levels of input intensities to ensure that the power limit is not exceeded. A maximum simulated average battery current determines the optimal (i.e., quickest) battery charging current, and at least an optimal value for a parameter indicative of that current, for example, the voltage across the battery charging circuitry, is determined and stored in the external charger. During charging, the actual value for that parameter is reported from the implant to the external charger, which in turn adjusts the intensity and/or duty cycle of the magnetic charging field consistent with the simulation to ensure that charging is as fast as possible, while still not exceeding the power dissipation limit. As a result, charging is optimized to be as fast as possible, while still safe from a tissue heating perspective.
Prior to discussing the disclosed technique, reference is made to the microstimulator power circuitry 160 of
Various components in the power circuitry 160 within the implant will draw power during the reception of a magnetic charging field from the external charger 151. In particular, the coil 147, its associated tank capacitor 162, the rectification circuitry (diode) 164, charging circuitry 170, and the battery 145 itself will all dissipate power in the form of heat. (Capacitor 166 will draw a comparatively negligible amount of power, and thus is not further discussed). The sum total of the powers dissipated by each of these components must be considered when understanding how the tissue surrounding the implant 100 will heat up during a charging session. For example, animal studies show that for a particular multiple-electrode microstimulator device, a radiated power of 32 mW will raise the temperature of the tissue surrounding the implant by approximately 4° C., while a total radiated power of 25.6 mW will raise the temperature by 3.2° C. Of course, these values are only exemplary, and could vary; future values could be determined that are more accurate, safer, etc. In any event, such animal studies have correlated power dissipation to tissue heating for a given implant.
An aspect of the disclosed technique seeks to keep the total dissipated power at or below a limit to ensure that the patient's tissue will not overheat. Because a 4° C. rise in tissue temperature is generally accepted as safe for a patient, one example of the technique labors to keep the total power dissipated from the power circuitry 160 at or below 32 mW. Of course, different limits could be chosen, such as the 25.6 mW/3.2° C. limit discussed above.
The inventors have noticed through simulations that power dissipation from the various components in the power circuitry 160 is complex and non-linear in nature. One such simulation 200 is illustrated in
Simulation 200 shows the effect of varying the intensity (e.g., current) in the external controller's charging coil 157 (Iprim(rms)) on the various components in the power circuitry 160 of the implant 100, with each successive row representing an increasing value for Iprim(rms). Because the simulation 200 results will vary depending on how full or depleted the implant battery 145 is at a given moment, the depicted simulation assumes a battery with a particular voltage of Vbat=3.1 V. Although not depicted, other simulations 200 at other battery voltages (e.g., 3.3V, 3.7V, 4.1V, etc.) may also be generated to provide accurate simulation results as battery capacity starts to fill during charging. For example, if the battery 145 has a full capacity of Vbat=4.1V, then simulations 200 may be generated for Vbat=3.1 V, 3.3V, 3.7V, and 4.1V to cover a range of expected battery capacity. However, if the various parameters within simulation 200 do not vary appreciably with Vbat, then the generation of additional simulations 200 for different battery capacities may not be necessary. A simulation program useful in generating a simulation 200 is Mentor Graphics Design Architect.
The simulation 200 assumes a particular coupling factor between the primary coil 157 in the external charger 151 and the secondary coil 147 in the implant 100, which coupling factor is modeled taking into account factors affecting such coupling, such as coil inductances, coil alignment, the distance and permittivity of any materials (e.g., tissue, air) between the coils, etc. In the depicted simulation, a coupling factor k=0.017 was chosen to conservatively simulate a worst case alignment between the charging coils 157 and 147. In any event, the coupling factor ultimately results in a simulated induced current in charging coil 147 in the implant (Isec(rms)), a current in the associated tank capacitor 162 (Icap(rms)), a voltage across the coil 147 (Vcoil(rms)), a DC voltage produced by the rectifier circuit (diode) 164 (Vna), a battery charging current (Ibat), a battery voltage (Vbat) resulting from the input of the battery charging current, which battery voltage takes into account the internal resistance of the battery 145. Of course, relevant parameters for the various components in the power circuitry 160 (resistances, capacitance, inductances, coupling factor, etc.) are input into the simulation program to allow it to generate the simulation results.
Of particular interest in simulation 200 is the voltage across the charging circuitry 170, Vnab, which represents the difference between Vna and Vbat. Because the charging circuitry 170 is in line with the battery charging current, Ibat, any voltage build up across the charging circuitry comprises undesired heat generation. Unfortunately, modeling shows that the amount of heat dissipation from the charging circuitry 170 increases essentially exponentially as the battery charging current increases. This is shown in
From the various simulated voltages and currents in
A review of the Ptotal parameter in simulation 200 illustrates a tissue heating concern for the designer. As discussed earlier, an acceptable level of total power dissipated by the power circuitry 160 should not exceed the 32 mW power dissipation limit in one example—a temperature known by experimentation to increase surrounding tissue by 4° C. However, all but the top three rows in
One solution to keep the total power at or below 32 mW is to duty cycle the power at the external charger 151, which computed duty cycle is shown in
The results of such duty cycling are shown in
As will be seen further below, it is advantageous to telemeter data (e.g., Vnab, Vbat) back to the external charger 151 to allow charging to be iteratively optimized in real time. As can be seen in
In the fourth row of the simulation 200 (Iprim(rms)=800 mA), the simulated total power dissipated was 38.6 mW, above the 32 mW limit. Therefore, duty cycling is imposed as a heat control measure, in addition to the desire for an off period to allow for back telemetry. Such duty cycling equals 82.9% ( 32/38.6) to ensure a total dissipated power of not more than 32 mW. The fifth row is similarly processed to determine a duty cycle of 61.2%, and its effects on Iprim(rms) and Ibat are shown.
Additionally shown to the right in
Note from
The disclosed technique seeks to maintain charging at this optimal average battery current. To so maintain Ibat(avg)(opt) during charging, it is useful to monitor a parameter indicative of the battery charging current, Ibat. One convenient parameter comprises Vnab, i.e., the voltage that builds across the charging circuitry 170, although other parameters indicative of the battery charging current could also be used (e.g., Vna). The Vnab parameter is easily measured in the implant, and as noted earlier represents wasted heat.
Prior to discussing use of the technique in an actual charging session, steps to this point in the process are summarized in
Also shown in
Traditionally, such back telemetry from the implant to the external charger is used to transmit the capacity of the battery 145 during charging (Vbat), which informs the external charger 151 when the battery is full and that charging can cease. Battery capacity is similarly reported in disclosed system, but additionally, the Vnab value measured at the implant 100 is also transmitted. Reporting of Vnab to the external charger 151 can take place at any suitable interval during charging, such as once every 100 seconds or so. The more frequently Vnab is reported, the more frequently charging can be optimized during the charging session.
With the basic structure of the external charger 151 understood, attention can now focus on how charger 151 operates during an actual charging session, which basic steps are shown in
Periodically during charging, for example, perhaps every 100 seconds, the battery voltage (Vbat) and the voltage across the charging circuitry (Vnab) are measured at the implant 100, and telemetered to the external charger. Again, such telemetry can comprise RF or LSK telemetry performed during the telemetry window (TW) or off periods in the duty cycle. How often to communicate, just like the time used for communication during the telemetry window (TW), may also be determined by the length of the needed communication between implant and charger. Increasing the frequency of communication will reduce temperature ripple in the implant 100.
Once Vnab is reported, the microcontroller 300 consults memory 302 to see if Vnab is optimal, i.e., if Vnab=Vnab(opt) for the reported Vbat. If not, intensity of the magnetic charging field is changed. For example, and referring to memory 300 in
At the same time, the duty cycle of the magnetic charging field would also be changed to match the Vnab being reported. Modifying the duty cycle to match Vnab is important to ensure proper compliance with the power dissipation limit. For example, and referring again to
As shown in
To this point in the disclosure, it has been assumed that there is a single optimal Vnab value, Vnab(opt). However, Vnab(opt) can also represent a range of acceptable Vnab values. For example, the simulation 200 in
It should be understood that various parameters (e.g., Vnab(opt); a DC corresponding to a particular Vnab) can be interpolated or extrapolated from the simulation 200, and are therefore not necessarily constrained to actual values appearing in the simulation. However, such interpolation was not shown to keep discussion of the technique simple.
Many of the parameters determined herein (e.g., Vnab(opt)) result from the simulation 200, which simulation provides a convenient expedient for understanding the external charger/implant system. However, not all implementations of the technique will require the use of a simulation. Instead, empirical data, experimental models, direct analytical tools, or values chosen by other means, could be used depending upon consideration of factors deemed important by the designer.
The disclosed technique limits the total power dissipated by the implant. However, the technique can be constrained to control heating at only a portion of the implant. For example, in larger implants or implants with low heat conductivity, the technique can be employed to limit the local heating at any section of the implant. In such an application, the technique can use a parameter (perhaps different from Vnab) indicative of heating to that section, and limiting heating of that particular section to tolerable limits. Thus, this modification to the technique would only consider power dissipated as heat in the relevant section of the device.
Vnab is used in this disclosure as the measure indicative of excess power dissipation. However, other parameters from the implant indicative of incoming power and which can be used to control that power can also be used, such total power delivered to the battery, ripple of the coil voltage, ripple of the rectified voltage, on time of the rectifying circuit, duty cycle of the rectifying circuit, etc. Of course, these parameters could be measured or inferred in the implant in different ways.
Even though the technique describes the periodic measurement of parameters in the implant during a charging session, and periodic adjustment of the magnetic charging field, “periodic” should not be understood as necessarily taking such actions at set intervals. Instead, “periodic” should be understood as taking a plurality of such actions over time, even if not at set intervals.
While the inventions disclosed have 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 literal and equivalent scope of the inventions set forth in the claims.
This is a continuation of U.S. patent application Ser. No. 14/292,169, filed May 30, 2014 (U.S. Pat. No. 9,358,399), which is a continuation of U.S. patent application Ser. No. 12/575,733, filed Oct. 8, 2009 (U.S. Pat. No. 8,744,592). Priority is claimed to these applications, and they are incorporated herein by reference in their entireties.
| Number | Date | Country | |
|---|---|---|---|
| Parent | 14292169 | May 2014 | US |
| Child | 15172856 | US | |
| Parent | 12575733 | Oct 2009 | US |
| Child | 14292169 | US |
