JP2018519968A - Implantable pulse generator for generating spinal stimulation signals for the human body - Google Patents

Abstract

An implantable pulse generator (IPG) that generates a spinal cord stimulation signal for a human body includes a timing generator and a high-frequency generator. The timing generator generates a timing signal representing a multi-channel stimulus signal. The high frequency generator determines whether to modulate the timing signal and, if the determination is affirmative, modulates the timing signal at the burst frequency according to the stored burst parameter. Thus, IPG provides the ability to generate both low and high frequency stimulation signals in different channels according to user programming.

Description

Detailed Description of the Invention

CROSS REFERENCE TO RELATED APPLICATIONS This application is filed on March 14, 2014, claiming priority from US Provisional Application No. 61 / 792,654, entitled “SPINAL CORD STIMULATOR SYSTEM,” filed March 15, 2013. Is a continuation-in-part of US patent application Ser. No. 14 / 213,186, which is hereby incorporated by reference in its entirety.
[Technical field]

The present disclosure relates to a stimulation device that uses electrical pulses in a medical environment, and more particularly applies an electrical stimulation signal to the spinal cord to control pain.
[Background technology]

  A spinal cord stimulator (SCS) is used to apply a pulsed electrical signal to the spinal cord to control chronic pain. Spinal cord stimulation consists of, in its simplest form, a stimulation electrode embedded in the epidural space, an implantable pulse generator embedded in the lower abdominal region or buttocks, a conductive wire connecting the electrode to the electrical pulse generator, An electric pulse generator remote control device and an electric pulse generator charger are provided. Spinal cord stimulation has a pronounced analgesic effect and is now mainly used for the treatment of post-lumbar pain syndrome, complex regional pain syndrome, and cancer pain due to ischemia.

  Neural stimulation pain electrotherapy began in 1965 shortly after Melzack and Wall advocated a gate control theory. This theory suggested that both nerves that transmit painful peripheral stimuli and nerves that transmit tactile and vibrational sensations terminate at the dorsal horn (gate) of the spinal cord. It was hypothesized that input to the dorsal horn of the spinal cord could be manipulated to “close the gate” to the nerve. As an application of the gate control theory, Sheary et al. Embedded the first spinal cord stimulation device directly in the posterior spinal column for the treatment of chronic pain in 1971.

  Spinal cord stimulation does not remove pain. The electrical impulse from the stimulator invalidates the pain information so that the patient does not feel intense pain. In essence, the stimulator masks pain. Trial embedding is performed before implanting the permanent stimulator. The physician first implants a trial stimulator through the skin (percutaneously) to perform the stimulation as a trial. Since the percutaneous trial stimulator tends to move away from its original position, it is considered temporary. If the trial is successful, the physician can then implant a permanent stimulator. The permanent stimulator is implanted under the skin of the abdomen and the lead is inserted under the skin and sent subcutaneously into the spinal canal. This arrangement of the stimulator in the abdomen is a more stable and effective position. The lead consists of an array of electrodes and can be percutaneous or paddle type. The percutaneous electrode is more easily inserted than the paddle type and is inserted through an incision on the spinal cord and a laminectomy.

  There are several problems that exist with currently available implantable pulse generators that limit the full benefit of posterior spinal column stimulation in terms of effectiveness and patient user ease of use.

  One problem is that the circuit in the current generator consumes too much power. This requires frequent charging and is very inconvenient for the patient. Another problem is that current generators are limited in generating different stimulation patterns simultaneously to treat different parts of the body at the same time. Thus, if a patient has varying degrees of pain at different parts of the body, it is difficult if not impossible to treat all pain areas emotionally.

Therefore, it would be desirable to provide a system and method for generating stimulation patterns that solves the problems discussed above.
[Summary of Invention]

  According to one aspect of the present invention, an implantable pulse generator (IPG) that generates a spinal cord stimulation signal for the human body is provided, which does not involve any intervention from the processor that controls the overall operation of the IPG. Without it, it has a programmable signal generator that can generate a signal based on stored signal parameters. While the signal generator is generating a signal, the processor can be in a standby mode to substantially save battery power.

  According to another aspect of the present invention, an implantable pulse generator (IPG) is provided that generates a spinal cord stimulation signal for a human body, wherein each channel is associated with at least two electrodes and is specific to the associated electrode. A control register for storing stimulation signal parameters for the stimulation channel in a state capable of representing a plurality of stimulation signal patterns. The arbiter continuously receives the timing signal representing the stimulation signal pattern and selects one of the multiple channels as the dynamic treatment channel in order to avoid two channels being activated simultaneously. The arbiter provides the flexibility to program different pulse parameters for multiple stimulation channels without the possibility of overloading the power supply that generates the stimulation signal pattern.

  In accordance with another aspect of the present invention, an implantable pulse generator (IPG) for generating a spinal cord stimulation signal for the human body is provided, which includes a timing generator and a high frequency generator. The timing generator generates a timing signal representing a multi-channel stimulus signal. The high frequency generator determines whether to modulate the timing signal and, if the determination is affirmative, modulates the timing signal at the burst frequency according to the stored burst parameter. Thus, IPG provides the ability to generate both low and high frequency stimulation signals in different channels according to user programming.

  In accordance with another aspect of the present invention, an implantable pulse generator (IPG) for generating a spinal cord stimulation signal for the human body is provided, which includes a timing generator and a high frequency generator. The timing generator generates a timing signal representing a multi-channel stimulus signal. The high frequency generator determines whether to modulate the timing signal and, if the determination is affirmative, modulates the timing signal at the burst frequency according to the stored burst parameter. The high frequency generator can also independently control the pulse frequency of each channel according to the stored parameters. Thus, IPG provides the ability to generate both low and high frequency stimulation signals in different channels according to user programming to provide maximum flexibility in therapy.

According to another aspect of the present invention, an implantable pulse generator (IPG) for generating a spinal cord stimulation signal for a human body, comprising an electrode driver for each electrode, and a stimulation signal for each stimulation pattern channel In order to allow independent amplitude control, an implantable pulse generator (IPG) is provided that outputs an output current corresponding to a conditioned signal for transmission to the associated electrode.
[Brief description of drawings]

FIG. 1 depicts various components that can be included in a spinal cord stimulation system, according to embodiments, during trial implantation and permanent implantation.
FIG. 2 depicts an exploded view of an implantable pulse generator (IPG) assembly, according to an embodiment.
FIG. 3 depicts a feedthrough assembly of an implantable pulse generator (IPG) assembly, according to an embodiment.
FIG. 4 depicts a lead contact system of an implantable pulse generator (IPG) assembly, according to an embodiment.
FIG. 5 depicts a lead contact assembly of an implantable pulse generator (IPG) assembly, according to an embodiment.
FIG. 6 depicts a head unit assembly of an implantable pulse generator (IPG) assembly, according to an embodiment.
FIG. 7 depicts an RF antenna of an implantable pulse generator (IPG) assembly, according to an embodiment.
FIG. 8 depicts a percutaneous lead according to an embodiment.
FIG. 9 depicts a paddle-type conductor according to an embodiment.
FIG. 10 depicts a conductor extension, according to an embodiment.
FIG. 11 depicts a conductor splitter, according to an embodiment.
FIG. 12 depicts a sleeve anchor, according to an embodiment.
FIG. 13 depicts a mechanical locking anchor, according to an embodiment.
FIG. 14 illustrates communication via a wireless dongle with a tablet / clinician programmer and a smartphone / portable / patient programmer during trial and / or permanent implantation, according to an embodiment.
FIG. 15 depicts a chewy needle, according to an embodiment.
FIG. 16 depicts a stylet, according to an embodiment.
FIG. 17 depicts a passing elevator according to an embodiment.
FIG. 18 depicts a tunneling device according to an embodiment.
FIG. 19 depicts a torque wrench according to an embodiment.
FIG. 20 is a functional block diagram of several components in an implantable pulse generator, according to an embodiment.
FIG. 21 is a functional block diagram of the signal generator of FIG.
FIG. 22 is a functional block diagram of the high frequency generator of FIG.
FIG. 23 is a functional block diagram of the electrode driver of FIG.
FIG. 24 is a functional illustration of two of the current drivers of FIG.
FIG. 24 illustrates an exemplary electrode waveform for an active channel, according to an embodiment of the present invention.
FIG. 25 shows the classification of electrodes for different channels according to an embodiment of the present invention.
FIG. 26 illustrates an exemplary electrode waveform showing an asymmetric pulse pattern between positive and negative pulses, according to an embodiment of the present invention.
FIG. 27 illustrates an exemplary electrode waveform for two arbitrated stimulation channels according to an embodiment of the present invention.
[Mode for Carrying Out the Invention]

Implantable pulse generator (IPG)
FIG. 1 illustrates various components that can be included in an SCS system for trial and permanent installation periods. The spinal cord stimulator (SCS) 100 is an implantable device used to deliver electrical pulse therapy to the spinal cord to treat chronic pain. The implantable component of the system consists of an implantable pulse generator (IPG) 102 and a number of stimulation electrodes 130. The IPG 102 is implanted subcutaneously at a depth of only 30 mm in an area that is not painful for the patient, while the stimulation electrode 130 is implanted directly into the epidural space. The electrodes 130 are wired to the IPG 102 via leads 140, 141 that hold the stimulation pulses isolated from each other to deliver the correct therapy to each individual electrode 130.

  The therapy delivered consists of electrical pulses with a controlled current amplitude in the range from +12.7 mA to 12.7 mA (current range 0-25.4 mA). These pulses can be programmed for both a length of 10 μS to 2000 μS and a frequency of 0.5 Hz to 1200 Hz. At any moment, the sum of the currents supplied from the anode electrode 130 must be equal to the sum of the currents attenuated by the cathode electrode 130. Furthermore, each individual pulse is biphasic, meaning that another pulse of opposite amplitude is generated after the set off state period when the initial pulse is completed. The electrodes 130 can be categorized into stimulation sets to deliver pulses over a wider area or to target a specific area, but the sum of the currents generated at any one given time is Cannot exceed 20 mA. The user can also program different stimulation sets (up to 8) with different parameters to target different areas with different therapies.

  FIG. 2 depicts an exploded view of the IPG 102. The IPG 102 consists of two main dynamic components 104, 106, a battery 108, an antenna 110, several support circuits, and a number of output capacitors 112. The first of the main active components is the microcontroller 104 transceiver 104. It is responsible for receiving, decrypting and executing both commands and requests from external remotes. If necessary, it passes those commands or requests to the second main component, ASIC 106. The ASIC 106 receives digital data from the microcontroller 104, performs all signal processing, and generates signals necessary for stimulation. These signals are then passed to the stimulation electrode 130 in the epidural space.

  The ASIC 106 includes a digital section and an analog section. The digital section is divided into multiple sections, including timing generators, arbitration controllers, pulse burst adjusters, and electrode logic. The analog section receives incoming pulses from the digital section and amplifies them to deliver the correct therapy. There are also a number of digital register storage elements utilized by the respective sections, both digital and analog.

  All of the digital elements in ASIC 106 are made up of a standard subset of digital logic, including logic gates, timers, counters, registers, comparators, flip-flops, and decoders. These elements are ideal for processing stimulation pulses, since all of them can function very fast, that is, much faster than the required pulse width. All devices function at one single voltage, typically 5.0, 3.3, 2.5, or 1.8 volts.

  The timing generator is the base of each of the stimulus sets. It generates actual rising and falling edge triggers for each phase of the biphasic pulse. It accomplishes this by taking the incoming clock sent from the microcontroller 104 and sending it to the counter. For the purposes of this discussion, assume that the counter counts these clock rising edges indefinitely. The output of the counter is sent to six different comparators. The other input of the comparator is connected to a specific register that is programmed by the microcontroller 104. If the grand total is equal to the value stored in the register, the comparator asserts a positive signal.

  The first comparator is connected to the set signal of the SR flip-flop. The SR flip-flop remains positive until the reset signal is asserted, to which the second comparator is connected. The output of the SR flip-flop is the first phase of the biphasic pulse. Its rising and falling edges are values stored in registers and programmed by the microcontroller 104. The third and fourth comparators and registers operate in exactly the same way to produce the second phase of the biphasic pulse using the second SR flip-flop.

  The fifth comparator is connected to the reset of the final SR flip-flop in the timing generator. This flip-flop is set by the first comparator, which is the rising edge of the first pulse. The reset is then triggered by the value of the programmed microprocessor in a register connected to the comparator. This allows a “hold-off” period after the falling edge of the second pulse. The output of this third SR flip-flop can be thought of as an envelope of a biphasic pulse that indicates when this particular timing generator is valid.

  The final comparator of the system is again connected to a register that stores the frequency value from the microprocessor. Basically, when the total number reaches this value, it triggers a comparator that is fed back to the counter, resets this value to zero, and begins the entire pulse generation period again. The ASIC 106 may include a number of these timing generators so that each can control a range from two to all of the electrodes 130 connected to the IPG 102 at the same time. However, if there are multiple timing generators and multiple channels are effectively programmed, a mechanism is needed to suppress the second channel from turning on when another is already enabled. The

  The next circuit block included in the IPG 102 is an arbitration unit. The arbiter works by examining each of the timing generator envelope signals to make sure that only one of them can be valid at the same time. If the second attempts to start, the arbitration unit suppresses the signal.

  The arbiter accomplishes this by drawing each of the channel envelope signals into a rising edge detection circuit. When one is triggered, it is sent to the set pin of the SR flip-flop. The outputs of this SR flip-flop are sent to all other rising edge detectors to suppress them from triggering. The channel envelope signal is also sent to the falling edge detector which is then sent to reset the SR flip-flop. The output of the SR flip-flop is then connected to a switch that combines all of its outputs and switches on / off a particular biphasic pulse train for that channel. As a result, the output of this circuit element is one biphasic pulse train, and is a signal indicating which timing generator the specific pulse train has come out of. In principle, the circuit looks for a valid channel. If it finds one, it suppresses everything else until the channel becomes inactive.

  The next section of the circuit is then very similar to the timing generator to create a fast burst pulse train, combined with the stimulation pulse train to create a bursty biphasic pulse train as needed. work.

  It accomplishes this by taking the incoming clock sent from the microcontroller 104 and sending it to the counter. The counter can count these rising clock edges indefinitely. The counter is only valid during the period of one phase of the biphasic signal and begins counting as soon as a rising edge is detected. The output of the counter, along with the microcontroller programmed register, is sent to the comparator, whose output is connected to a reset pin on the counter. As a result, this counter will only count up to the programmed value and reset. This program value is the burst frequency.

  The output of the comparator is then sent to an edge detection circuit and then to a flip-flop that couples it to the actual stimulus pulse train to produce a single phase burst stimulus pulse. The entire circuit is replicated for the second phase of the signal resulting in the desired bursted biphasic pulse train. The stimulus signal is immediately handed over to the electrode logic stage.

  The electrode logic circuit conditions and directs the biphasic signal to the analog section of the ASIC 106. Currently, the biphasic signal contains all of the timing information, but does not contain any needed amplitude information. The input signal includes a biphasic pulse train and another signal indicating which timing generator the current valid train is from. Each electrode logic cell has a register for each timing generator that stores the amplitude value of this particular electrode 130 for that timing generator. The electrode logic cell uses a designation signal to determine which register to derive the amplitude value from, for example, when the third timing generator is passed through the arbitration circuit, the electrode logic circuit is in the third register. Will read the value from

  When a value is pulled from a register, it passes through a series of logic gates. The gate first determines that electrode 130 is valid. If not, no further action is taken and the analog section of the electrode output is not activated, thus saving valuable battery 108 power. Next, a determination is made if the particular electrode 130 is an anode or a cathode. If the electrode is considered an anode, the electrode logic circuit passes amplitude information and a biphasic signal to a digital-to-analog converter (DAC) in the analog section of the ASIC 106. If the electrode is considered a cathode, the electrode logic passes amplitude information and a biphasic signal to the negative current DAC in the analog section of the ASIC 106. The electrode logic must make those decisions for each phase of the biphasic signal each time the electrode 130 switches between the anode and the cathode.

  The analog elements in the ASIC 106 are specially designed to produce the desired signal. The basis of analog IC design is a field effect transistor (FET), and the type of high current multiple output design required in SCS100 means that the bulk of silicon in ASIC 106 is dedicated to the analog section. To do.

  The signal from the electrode output is sent to the respective current DAC when that particular electrode 130 is to be enabled. Each electrode 130 has a positive and negative current DAC triggered by an electrode logic circuit, both of which are not active at the same time. The job of each current DAC, when activated, is to take a digital value representing the stimulation current amplitude and generate an analog representation of this value that is sent to the output stage. This circuit forms a half of the barrier between the digital and analog sections of the ASIC 106.

  The digital section of the ASIC 106 is built on technology that only allows a small voltage to exist. When moving to the analog section, the output of the current DAC (which is a low level analog signal) must be amplified to a high voltage for use in the analog section. The circuit that performs this task is called a power level shifter. Since this circuit is made under two different manufacturing technologies and requires a high precision analog circuit made under a digital base, it can be difficult to implement.

  As the voltage is converted for use in the analog portion of the ASIC 106, the voltage is passed to the output current stage. There are two current sources for each electrode output. One will supply positive current and the other will reduce negative current, but both will never be active at the same time. The current source itself is composed of analog elements that are similar to Howland current sources. There is an input stage and an amplification stage with feedback through the sensing component to maintain a constant current. The input stage takes an analog voltage value from the power level shifter and generates an output pulse designated for the amplifier. The amplifier then produces a pulse that is a constant voltage but a variable voltage. The power supply can source or sink up to 12.7 mA at 0.1 mA resolution for loads up to 1.2 k ohms. This moves to a 15 volt range that will vary depending on the load to keep the current constant.

The microcontroller 104 to the ASIC 106 interface is designed to be as simple as possible with a minimum bus “chatter” to conserve battery 108 life. The ASIC 106 can be a collection of registers programmed via an I ² C or SPI bus. Since the ASIC 106 governs total power management, there will also be a power good (PG) line between the two chips 104, 106 to inform the microcontroller 104 if it is safe to power up. The ASIC 106 also needs to use a pin for the microcontroller 104 to generate a hardware interrupt in case of an emergency in the ASIC 106. The final connection is a time reference for all of the stimulation networks. The ASIC 106 requires two clocks, one clock is sent directly from the microcontroller 104 clock output, for its internal digital circuitry, another clock needs to be integrated by the microcontroller 104 This is to base all stimuli sent to the ASIC 106. All commands and requests to the ASIC 106 will be made on the I ² C or SPI bus and will involve reading register addresses or writing to registers. Even if the ASIC 106 generates a hardware interrupt, it will be the responsibility of the microcontroller 104 to poll the ASIC 104 and determine the cause of the interrupt.

  The radio interface is based on FCC's MedRadio standard, operating in the 402-405 MHz range, utilizing up to 10 channels for telemetry. The protocol implemented is chosen to minimize transmission and maximize battery 108 life. All processing is done to the remote user / programmer so that only the transmitted data will be used by the microcontroller 104 to the ASIC 106 bus. That is, all of the wireless packets will contain not only the register address but also the necessary overhead information, the data to store in the register, and the command byte that instructs the microcontroller 104 how to handle the data. The overhead section of the wireless protocol will include a synchronization bit, a start byte, an address that is synchronized with the IPG 102 serial number, and a CRC byte to ensure proper transmission. The packet length is kept as short as possible to maintain the battery 108 life. Because the IPG 102 cannot wait for the entire time of the packet due to the battery 108 life, it makes a round with a duty cycle shorter than 0.05% of that time. This time value can be kept as short as the data packet is shortened. User commands required to run the system are executed by the entire system using the flow.

  The IPG 102 uses an embedded grade lithium ion battery 108 using zero voltage 215 mAHr technology. The voltage of battery 108 at full capacity is 4.1V, which supplies current only until it is consumed to 3.3V, which is considered 100% discharge. The remaining capacity of the battery 108 can be estimated at any time by measuring the voltage across the terminals. The maximum charging rate is 107.5 mA. The regulated constant current low voltage (CCCV) type can be applied to faster charging of the battery 108.

  The internal secondary coil 109 is composed of a maximum of 30 turns of 30 AWG copper magnet wires. The coil ID, OD, and thickness are 30, 32, and 2 mm, respectively. Inductance L2 is measured to be 58 uH and an 80 nF capacitor is connected to it to create a series resonant tank at a frequency of 74 kHz. In inductive charging technology, two types of rectifiers, either bridge full-wave rectifiers or voltage doubler full-wave rectifiers, are believed to convert inductive AC to usable DC. In order to obtain higher voltages, double voltage full wave rectifiers are used in this application. The rectifier is built with a fast Schottky diode to improve its function at high frequencies on the order of 100 kHz. Zener diodes and 5V voltage regulators are also used for regulation. This circuit would be able to supply 100 mA of current to the power management IC that induces AC voltage, rectifies to DC, regulates to 5V, and charges the internal battery 108 through CCCV regulation.

  The adjusted 5V 100 mA output from the resonant tank is sent to, for example, a power management integrated circuit (PMIC) MCP 73843. This particular chip was specially designed by the microchip to charge the lithium ion battery 108 to 4.1V by CCCV regulation. The fast charge current can be adjusted by replacing the resistor and is set to a 96 mA threshold current in the example circuit. The chip charges the battery 108 to 4.1 V as long as the received current exceeds 96 mA. However, if the supply current drops below 96 mA, it stops charging the battery 108 until the supply is again higher than 96. For various practical reasons, if the distance between the coils increases, the internal secondary coil 109 will receive a current less than the regulated value, and instead of slowly charging the battery 108, it will receive more than 96mA. Suspend charging completely. Those skilled in the art will appreciate that other power management chips can be used and the power management chip is not limited to the PMIC MCP 734432 chip.

  All functions of the IPG 102 are externally controlled using a portable remote control specially designed for this device. If the remote control device is lost or damaged, it is desirable for additional controls to operate the IPG 102 in addition to the remote control device. For this purpose, a Hall effect based magnet switch was incorporated to either turn the IPG 102 on or off using an external piece of magnet. The magnet switch functions as a main control for the IPG 102 for turning on or off. A sufficiently strong south pole turns on the output and a sufficiently strong north pole turns off the output. The output is latched so that the switch continues to hold its state even after the magnet has been removed from its surroundings.

  The IPG 102 is an active medical implant that generates electrical signals that stimulate the spinal cord. The signal is conveyed through a stimulation lead 140 that plugs directly into the IPG 102. The IPG 102 communicates via the RF wireless antenna 110 to recharge wirelessly through the induction coil 109 and change the stimulation parameters. The IPG 102 can be fixed to the fascia by being embedded up to 3 cm below the epidermis and passing two sutures through the holes in the epoxy header 114. Lead 140 is electrically connected to IPG 102 through lead contact system 116, a cylindrical spring-based contact system with an internal contact silicone seal. Conductor 140 is secured to IPG 102 with a set screw 117 that operates within locking housing 118. Set screw compression on the stationary contact of lead 140 can be governed by a disposable torque wrench. Wireless recharging is accomplished by aligning the external induction coil on the charger with the internal induction coil 109 in the IPG 102. The RF antenna in the remote dongle 200 communicates with the RF antenna 110 in the epoxy resin header 114 of the IPG 102. FIG. 2 shows an exploded view of the IPG 102 assembly.

  The IPG 102 is an assembly of a sealed titanium (6Al-4V) housing 120 that houses a battery 108, circuitry 104, 106, and a charging coil 109. The IPG 102 further includes an epoxy resin header 114 (see FIG. 6) that houses the lead contact assembly 116, the locking housing 118, and the RF antenna 110 (see FIGS. 6 and 7). The internal electronics are connected to components in the epoxy resin head through hermetic feedthrough 122, as shown in FIG. The feedthrough 122 is a titanium (6Al-4V) flange having an alumina window and a gold ring. Within the alumina window, 34 interfaces internally with direct solder to the circuit board, and externally interfaces with a series of laser welded platinum iridium wires to the antenna 110 and conductor contact 126. Platinum iridium (90-10) pins. The IPG 102 interfaces with 32 electrical contacts 126 arranged in four rows of 8 contacts 126. Thirty-two of the pins 124 of the feedthrough 122 are interfaced with the contact 126, while two are interfaced with the antenna 110, one to the ground plane and the other to the antenna 110 feed.

  4 and 5 depict the lead contact system 115 and assembly 116, respectively. Conductor contact 126 consists of MP35N housing 128 with platinum iridium 90-10 spring 129. Each contact 126 is separated by a silicon seal 127. At the proximal end of each stack of eight contacts 126 is a titanium (6Al-4V) cap 125 that serves as a stop for the lead 140. At the distal end, there is a titanium (6Al-4V) set screw 119 and a block 118 for securing the lead. At the lead entry point, there is a silicone tube 123 that provides strain relief as the lead 140 exits the head unit 114, and over the set screw 119, another silicone tube 131 with a small inner tube allows a torque wrench to enter. Allows but does not allow the set screw 119 to retract. In addition to contact 126 and antenna 110, header 114 also includes a radiopaque titanium (6Al-4V) tag 132 that allows for identification of the device under fluoroscopy. The overmold of the header 114 is Epotek 301, a two part, biocompatible epoxy. 4, 5, 6, and 7 depict views of lead contact system 115, lead contact assembly 116, head unit assembly 114, and RF antenna 110, respectively.

  Inside the titanium (6Al-4V) container 120 are a circuit board 105, a battery 108, a charging coil 109, and an internal plastic support frame. The circuit board 105 can be a multilayer FR-4 board with copper traces and solder mask coating. The non-solder mask area of the plate can be electroless nickel immersion gold. The embedded battery 108, all surface mount components, ASIC 106, microcontroller 104, charging coil 109, and feedthrough 122 will be soldered to the circuit board 105. A plastic frame made of either polycarbonate or ABS will provide a sliding fit between the network 105 and the container 120 to maintain the position of the battery 108 and prevent movement. The charging coil 109 is a wound coated copper.

Conductor As depicted in FIG. 8, the percutaneous stimulation lead 140 is a fully implantable electromedical accessory device for use with the implantable SCS 100. The primary function of the lead is to transmit electrical signals from the IPG 102 to the targeted stimulation area on the spinal cord. Percutaneous stimulation lead 140 provides peripheral stimulation. The percutaneous stimulation lead 140 provides a strong, flexible, and biocompatible electrical connection between the IPG 102 and the stimulation region. The lead 140 is surgically implanted through a subarachnoid puncture needle or epidural needle and is driven through the spinal canal using a progressive probe that penetrates the center of the lead 140. The lead 140 is mechanically secured to the patient using either anchors or sutures that penetrate tissue and are tied around the body of the lead 140. The lead 140 is secured to the proximal end with a set screw 119 on the IPG 102 that applies radial pressure to the blank contact on the distal end of the proximal contact.

  Percutaneous stimulation lead 140 is comprised of a combination of implantable materials. The stimulation electrode 130 at the distal end and the electrical contact at the proximal end are made of 90-10 platinum iridium alloy. This alloy is utilized for its biocompatibility and electrical conductivity. The electrode 130 is geometrically cylindrical. The polymer of the lead 140 is a polyurethane, chosen for its biocompatibility, flexibility, and high lubricity to reduce friction while penetrating tissue. The polyurethane tubing has a multi-lumen cross section with one central lumen 142 and eight external lumens 144. The central cavity 142 functions as a tube for containing the progressive probe during implantation, while the external cavity 144 is electrically and mechanically between the wires 146 that transmits stimulation from the proximal contact to the distal electrode 130. Bring about separation. These wires 146 are bundles of MP35N yarn with a 28% silver core. The wires 146 are individually coated with ethylene tetrafluoroethylene (ETFE) to provide an additional non-conductive barrier. Wire 146 is laser welded to contact and electrode 130 to create an electrical connection between the respective contacts on the proximal and distal ends. The lead 140 employs a platinum iridium plug 148 molded into the distal tip of the central cavity 142 to prevent the advancing probe from puncturing the distal tip of the lead 140. Conductor 140 is available in a variety of four and eight electrode 130 configurations. These leads 140 have 4 and 8 proximal contacts (+1 fixed contact), respectively. The configuration depends on the number of electrodes 130, the distance between the electrodes 130, the total length of the electrodes 130, and the total length of all the conductive wires 140.

  As depicted in FIG. 9, paddle stimulation lead 141 is a fully implantable electromedical accessory device for use with implantable SCS 100. The main function of the paddle lead 141 is to transmit electrical signals from the IPG 102 to the targeted stimulation area on the spinal cord. The paddle lead 141 provides unidirectional stimulation across the two-dimensional array of electrodes 130 and allows greater accuracy when targeting the stimulation zone. The paddle stimulation lead 141 provides a strong, flexible, and biocompatible electrical connection between the IPG 102 and the stimulation region. The lead 141 is surgically implanted through a small incision, usually in conjunction with a laminectomy or laminectomy, and positioned using forceps or a small surgical tool. Lead 141 is mechanically secured to the patient using either anchors or sutures that penetrate tissue and are tied around the body of lead 141. The lead 141 is secured to the proximal end with a set screw on the IPG 102 that applies radial pressure to the stationary contact on the distal end of the proximal contact.

  The paddle stimulation lead 141 is made of a combination of implantable materials. The stimulation electrode 130 at the distal end and the electrical contact at the proximal end are made of a 90-10 platinum iridium alloy that is utilized for its biocompatibility and electrical conductivity. The polymer of lead 141 is a polyurethane chosen for its biocompatibility, flexibility, and high lubricity to reduce friction while penetrating tissue. The polyurethane tubing has a multi-lumen cross section with one central lumen 142 and eight external lumens 144. The central cavity 142 functions as a tube for containing the progressive probe during implantation, while the external cavity 144 is electrically and mechanically between the wires 146 that transmits stimulation from the proximal contact to the distal electrode 130. Bring about separation. These wires 146 are bundles of MP35N yarn with a 28% silver core. The wires 146 are individually coated with ethylene tetrafluoroethylene (ETFE) to provide an additional non-conductive barrier. At the distal tip of the paddle lead 141 there is a two dimensional array of flat rectangular electrodes 130 formed into a flat silicon body 149. In an embodiment, one side of the rectangular electrode 130 is exposed to provide a unidirectional stimulus. Wire 146 is laser welded to contact and electrode 130 to create an electrical connection between the respective contacts on the proximal and distal ends. A polyester mesh 147 that adds stability to the molded body 149 while improving aesthetics by covering the wire 146 routing is also molded into the distal silicon paddle. The number of individual 8-contact conductors 141 used for each paddle 141 is governed by the number of electrodes 130. The electrode 130 for each paddle 141 range from 8 to 32 divides between one proximal lead 141 end and four proximal lead 141 ends. Each proximal lead 141 has 8 contacts (+1 fixed contact). The configuration depends on the number of electrodes 130, the spacing between the electrodes 130, the total length of the electrodes, and the total length of all the conducting wires.

  As depicted in FIG. 10, the lead extension 150 is a fully implantable electromedical accessory device for use with the implantable SCS 100 and either the percutaneous lead 140 or the paddle 141 lead. is there. The primary function of the lead extension 150 is to increase the overall length of the leads 140, 141 by transmitting electrical signals from the IPG 102 to the proximal ends of the stimulation leads 140, 141. This extends the full range of conductors 140, 141 if the provided conductors 140, 141 are not long enough. Conductor extension 150 provides a strong, flexible, and biocompatible electrical connection between IPG 102 and the proximal end of stimulation leads 140, 141. The extension 150 can be mechanically secured to the patient using either anchors or sutures that penetrate tissue and are tied around the body of the extension 150. The extension 150 is secured to the proximal end with a set screw 119 on the IPG 102 that applies radial pressure to a stationary contact on the distal end of the proximal contact of the extension 150. The stimulation leads 140, 141 use set screws 152 inside the molded tip of the extension 150 to apply radial pressure to the stationary contact at the proximal end of the stimulation leads 140, 141, It is fixed to the expansion unit 150 in the same manner.

  The conductor extension 150 is made of a combination of embedded materials. At the distal tip of the extension 150, there is a 1 × 8 array of implantable electrical contacts 154, each consisting of an MP35 housing 128 and a 90-10 platinum iridium spring. A silicon seal 127 separates each of the housings 128. At the proximal end of the contact, there is a titanium (6Al4V) cap that serves as a conductor stop, and at the distal tip is a titanium (6Al4V) block and a set screw 152 for securing the conductor. The electrical contact at the proximal end consists of a 90-10 platinum iridium alloy utilized for its biocompatibility and electrical conductivity. The polymer 156 of the lead 150 is a polyurethane selected for its biocompatibility, flexibility, and high lubricity to reduce friction while penetrating tissue. Polyurethane tubing 158 has a multi-lumen cross section with one central lumen 142 and eight external lumens 144. The central cavity 142 functions as a tube for containing the progressive probe during implantation, while the external cavity 144 is electrically and mechanically between the wires 146 that transmits stimulation from the proximal contact to the distal electrode. Bring separation. These wires 146 are bundles of MP35N yarn with a 28% silver core. The wires 146 are individually coated with ethylene tetrafluoroethylene (ETFE) to provide an additional non-conductive barrier. Each conductor extension 150 has eight proximal cylindrical contacts (+1 fixed contact).

  As depicted in FIG. 11, the lead splitter 160 is a fully implantable electromedical accessory device used in conjunction with the SCS 100 and usually a set of 4-contact percutaneous leads 140. The main function of the conductor splitter 160 is to divide one conductor 140 of eight contacts into a set of four contact conductors 140. The splitter 160 transmits electrical signals from the IPG 102 to the proximal ends of the two 4-contact percutaneous stimulation leads 140. This allows the surgeon to reach the stimulation area more by increasing the number of stimulation leads 140 available. Conductor splitter 160 provides a strong, flexible and biocompatible electrical connection between IPG 102 and the proximal end of stimulation lead 140. Splitter 160 may be mechanically secured to the patient using either anchors or sutures that penetrate tissue and are tied around the body of splitter 160. The splitter 160 is secured to the proximal end with a set screw 119 on the IPG 102 that applies radial pressure to a stationary contact on the distal end of the proximal contact of the splitter 160. The stimulation leads 140 are similar using a set of set screws inside the shaped tip of the splitter 160 to apply radial pressure to a stationary contact at the proximal end of each stimulation lead 140. It is fixed to the splitter 160 by this method.

  Conductor splitter 160 comprises a combination of embedded materials. At the distal tip of splitter 160, there is a 2 × 4 array of implantable electrical contacts 162, each contact 162 consisting of an MP35 housing 128 and a 90-10 platinum iridium spring. A silicon seal 127 separates each of the housings 128. At the proximal end of each row of contacts 162 there is a titanium (6Al4V) cap that serves as a lead stop, and at the distal tip there is a titanium (6Al4V) block and a lead screw set screw. The electrical contact at the proximal end of the splitter 160 consists of a 90-10 platinum iridium alloy that is utilized for its biocompatibility and electrical conductivity. The polymer 164 of the lead 160 is a polyurethane chosen for its biocompatibility, flexibility, and high lubricity to reduce friction while penetrating tissue. Polyurethane tubing 166 has a multi-lumen cross section with one central lumen 142 and eight external lumens 144. The central cavity 142 functions as a tube for containing the progressive probe during implantation, while the external cavity 144 is electrically and mechanically between the wires 146 that transmits stimulation from the proximal contact to the distal electrode 130. Bring about separation. These wires 146 are bundles of MP35N yarn with a 28% silver core. The wires 146 are individually coated with ethylene tetrafluoroethylene (ETFE) to provide an additional non-conductive barrier. Each conductor splitter 160 has two rows of eight proximal contacts (+1 fixed contacts) and four contacts 162 at the distal end.

Anchor As depicted in FIGS. 12 and 13, the lead anchor 170 is a fully implantable electromedical accessory used in conjunction with both the percutaneous stimulation lead 140 and the paddle stimulation lead 141. The primary function of the lead anchor 170 is to prevent movement of the distal tip of the lead 140, 141 by mechanically securing the lead 140, 141 to the tissue. There are currently two types of anchors 170 depicted in FIG. 12, one sleeve 171, and a locking mechanism 172 depicted in FIG. 13, each having a slightly different interface. For one sleeve type anchor 171, the conductors 140, 141 pass through the central through hole 174 of the anchor 171, and then the suture is turned around the outside of the anchor 171 and the conductor in the anchor 171. 140, 141 are fastened to fix. The anchor 171 can then be sutured to the fascia. The locking anchor 172 uses a set screw 176 for locking purposes and a bi-directional disposable torque wrench for locking and disengaging. Tactile and audible feedback is provided for both locking and disengaging.

  Both anchors 171, 172 can be molded from embedded silicon, but the locking anchor 172 uses an internal titanium assembly for locking. The three-part mechanism includes a housing 175, a locking screw 176, and an occlusion locking screw 177 to prevent the locking screw from moving backward. All three components can be titanium (6Al4V). The two-way torque wrench can have a plastic body and a stainless steel hex shaft.

Wireless dongle The wireless dongle 200 is capable of communicating between a trial generator 107 or IPG 102 and a smartphone / portable device 202 or tablet 204, as shown in FIG. Is a hardware connection to. During the trial or permanent implantation phase, the wireless dongle 200 is connected to the tablet 204 through connection pins specific to the tablet 204, and the clinician programmer's software on the tablet 204 is used to control the stimulation parameters. Commands from the clinician programmer software are communicated to the wireless dongle 200 which is then transmitted from the wireless dongle 200 using the RF signal to the trial generator 107 or IPG 102. Once the parameters on the clinician programmer have been set, the parameters can be stored on the tablet 204 and communicated to the patient programmer's software on the smartphone / portable device 202. The wireless dongle 200 includes an antenna, a microcontroller (having the same specifications as the IPG 102 and the trial generator 107), and a pin connector that connects to the smartphone / portable device 202 and the tablet 204.

Charger IPG 102 has a rechargeable lithium ion battery 108 that powers its activities. The external induction charger 210 (FIG. 1) recharges the attached battery 108 inside the IPG 102 wirelessly. The charger 210 is packed in a housing and includes a printed circuit board (PCB) including a rechargeable battery, a primary coil of wires, and an electronic device. In operation, the charger 210 generates a magnetic field and causes a voltage in the secondary coil 109 of the IPG 102. The induced voltage is then rectified and used to charge the battery 108 inside the IPG 102. In order to maximize the coupling between the coils, both the inner and outer coils are combined with a capacitor to resonate them at a specific shared frequency. The coil that functions as the inductor L forms an LC resonant tank. The charger uses a class E amplifier topology to generate an alternating current in the primary coil around the resonant frequency. The features of the charger 210 are not limited to the following,
● Charging the IPG 102 wirelessly ● Charging to a maximum depth of 30 mm ● The integrated alignment sensor provides high power transfer efficiency, indicating alignment between the charger and the IPG 102 ● The alignment sensor is Includes small and portable to provide audible and visual feedback to the user.

  The cylindrical lithium ion battery protection type is used as a battery for the charger 210. Class E power amplifier topologies are a commonly used type of amplifier for inductive chargers, especially for implantable electromedical devices. Because of the relatively high theoretical efficiency of class E power amplifiers, it is often used for devices that require high efficiency power transfer. A 0.1 ohm high wattage resistor is used in series to sense current through this circuit.

The primary coil L1 is made of 60 turns of 44 AWG litz linear 100 / 44-100 yarn each. The litz wire solves the skin effect problem and keeps its impedance low at high frequencies. The inductance of this coil was initially set at 181 uH, but lining it with a ferrite plate increases the inductance to 229.7 uH. The attached ferrite plate concentrates the generated magnetic field toward the implant. Such a mechanism helps the secondary coil to receive more magnetic field, which helps to induce high power.
When the switch is on, the resonance is at the following frequency:

When the switch is off, the frequency moves to the next

In continuous operation, the resonant frequency will be in the following range.

In order to approximate the on and off resonant frequencies, a relatively large value of C1 can be chosen by a simple criterion as follows.
C1 = nC2 and a value of n = 4 was used in the above example, but in most cases 3 <n <10.

The voltage in these class E amplifiers typically rises to about 300 VAC. The selected capacitor must be able to withstand these high voltages, sustain high currents, and further maintain a low equivalent series resistance (ESR). Higher ESR results in unnecessary power loss in the form of heat. The circuit is connected to the battery through an inductor that functions as a choke. The choke helps to smooth the supply to the circuit. The N-channel MOSFET functions as a switch in this class E power amplifier. FET having a low on-resistance and high drain current _{I d} is preferable.

  In summary, the circuit can recharge the battery 108 of the IPG 102 from 0% to 100% in about 2 hours and 45 minutes with a distance between the coils of 29 mm. The primary coil and class E amplifier draw 0.866A of DC current to accomplish this task. In order to improve the efficiency of the circuit, feedback closed loop control is implemented to reduce losses. Loss is minimal when the MOSFET is switched on and when the drain side voltage approaches zero.

  If the output meets the criteria, the controller retrieves and verifies the output from the operational amplifier and then triggers the driver to turn on the MOSFET in the next cycle. The controller can use a delay timer, an OR gate, and a 555 timer in a monostable configuration to condition the signal for the driver. If the device is switched on, the circuit will not function immediately, so that there is no active feedback loop. When the circuit begins to function, feedback becomes active. In order to provide an active feedback loop, an initial external trigger is applied to jump start the system.

Alignment Sensor The efficiency of power transfer between the external charger 210 and the internal IPG 102 will be maximized only when the charger 210 and the IPG 102 are correctly aligned. An alignment sensor is provided to ensure correct alignment as part of the external circuit and is based on the principle of firing impedance. When the external coil is brought close to the internal coil, the impedance of both circuits changes. Sensing is based on measuring the reflected impedance and checking if it exceeds a threshold. The beeper provides audible feedback to the patient and the LED provides visual feedback.

  As the circuit impedance changes, so does the current passing through it. A high power 0.1 ohm resistor can be used in a series of circuits to monitor the change in current. The voltage drop across the resistor is amplified by a factor of 40 and then compared to a fixed threshold using an operational amplifier voltage comparator. The output is sent to the timer chip, which in turn activates the beeper and LED so as to provide feedback to the user.

  The circuit can sense alignment up to a distance of about 30 mm. Current fluctuations in the circuit depend on many factors, not just the reflected impedance, and the circuit is sensitive to other parameters of the circuit as well. To reduce sensitivity to other parameters, one option is to eliminate the interference of all other factors and improve the functionality of the reflective impedance sensor, which is the limited available for the network. It is very difficult to carry out in a space. Another option is to use a dedicated sensor chip to measure the reflected impedance.

  The second design uses sensors designed for proximity detectors or metal detectors for alignment sensing. Tips designed to detect metal bodies due to the effect of eddy currents on the HF loss of the coil can be used for this application. TDE0160 is an example of such a chip.

  The external charger is designed to work at 75-80 kHz, while the proximity sensor is designed for 1 MHz. The sensor circuit is designed to be compatible with the rest of the exterior and is fine tuned to detect the internal IPG 102 from a distance of about 30 mm.

Programmer The clinician programmer is an application installed on the tablet 204. The clinician programmer is used by the clinician to set stimulation parameters on the trial generator 107 or IPG 102 during trial and permanent implantation in the operating room. A clinician programmer can save multiple settings for multiple patients and can be used to adjust stimulation parameters outside the operating room. The clinician programmer can change the stimulation parameters through the RF wireless dongle 200 when the trial generator 107 or IPG 102 implanted in the patient is within the RF range. In addition, the clinician programmer can set or change stimulation parameters on the trial generator 107 and / or IPG 102 via the Internet when both the tablet 204 and the patient programmer on the smartphone / portable device 202 access the Internet. You can also

  The patient programmer is an application that is installed on the smartphone / portable device 202. The patient programmer is used by the patient to set stimulation parameters in the trial generator 107 or IPG 102 after trial and permanent implantation outside the operating room. The clinician programmer can save multiple settings for multiple patients and when the clinician programmer tablet 204 and the patient programmer smartphone / portable device 202 are within a wireless range such as Bluetooth® with each other, the wireless Can be communicated to the patient programmer. If the clinician programmer tablet 204 and the patient programmer smartphone / portable device 202 are out of radio range of each other, data may be communicated over the internet where both devices 202, 204 have wireless access such as Wi-Fi. The patient programmer can change the stimulation parameters on the trial generator 107 or IPG 102 through the RF wireless dongle 200 when the trial generator 107 or IPG implanted in the patient is within the RF range.

The Chewy needle 240 as depicted in FIG. 15 is used in conjunction with a saline-filled syringe for placement of a resistance disappearance needle and a percutaneous stimulation lead 140 during placement of the lead 140 in the spinal canal. Used. The Chuchy epidural needle 240 is slowly inserted into the spinal canal using a resistance loss technique to measure the depth of the needle 240. When inserted to the proper depth, the percutaneous stimulation lead 140 is passed through the needle 240 and between the vertebrae.

  The epidural needle 240 is a non-coring 14G stainless steel subarachnoid puncture needle 240 that can be used in lengths of 5 inches (127 mm) and 6 inches (152.4). The distal tip 242 of the needle 240 is slightly curved to direct the stimulation lead 140 into the spinal canal. Proximal end 246 is a standard luer lock connection 248.

Probe The probe 250 as depicted in FIG. 16 is used to insert the tip of the percutaneous stimulation lead 140 into the desired stimulation zone by adding rigidity and operability. The wire 252 of the probe 250 passes through the central lumen 142 of the percutaneous lead 140 and stops at the distal tip of the lead 140 with an occlusion plug. The tip of the probe 250 has both a straight tip and a curved tip. A small handle 254 is used at the proximal end of the probe 250 to rotate the probe 250 within the central cavity 142 to assist in insertion. This handle 254 is detachable and allows the anchor 170 to pass over the conductor 140 while the probe 250 is still in place. The wire 252 of the probe 250 is a PTFE-coated stainless steel wire, and the handle 254 is plastic.

Passing Elevator A passing elevator 260 as depicted in FIG. 17 removes tissue in the spinal canal and helps the surgeon adjust the size of the lead relative to the anatomy. Used before. The passing elevator 260 includes a paddle-shaped tip 262 that is flexible to remove obstructions from the spinal canal. The flexible tip is attached to the surgical handle 264.

  The passing elevator 260 is an integral disposable plastic appliance made of a flexible high strength material having high lubricity. Flexibility allows the instrument to easily adapt to the angle of the spinal canal, and lubricity allows the instrument to easily penetrate tissue.

Tunneling Instrument The tunneling instrument 270 as depicted in FIG. 18 is used to create a subcutaneous conduit for passing the stimulation lead 140 from the entry point in the spinal canal to the IPG implantation site. Tunneling instrument 270 is a long skewer shaped instrument having a ringlet handle 272 at the proximal end 274. Instrument 270 is covered by a plastic sheath 276 having a tapered tip 278 that allows instrument 270 to easily penetrate tissue. When the IPG 102 implantation zone is bridged to the lead 140 entry point in the spinal canal, the inner core 275 is removed and the sheath 276 remains. A lead 140 can then be penetrated through the sheath 276 to the IPG 102 implantation site. Tunneling device 270 is often bent to assist in traveling through tissue.

  The tunneling device 270 is made of a 304 stainless steel core having a fluorinated ethylene propylene (FEP) sheath 276. 304 stainless steel is used due to its strength and ductility while bent, and sheath 276 is used due to its strength and ductility.

Torque Wrench In order to tighten the internal set screw 119 and create a radial force against the frictional contact of the stimulation leads 140, 141 and prevent the leads 140, 141 from coming off, a torque wrench 280 as shown in FIG. It is used in combination with the conductor extension 150 and the conductor splitter 160. The torque wrench 280 is also used to lock and disengage the anchor 170. Torque wrench 280 is a small, disposable medical instrument used in any SCS100 case. Torque wrench 280 provides the surgeon with audible and tactile feedback that leads 140, 141 are fixed to IPG 102, extension 150, or splitter 160, or anchor 170 is in the locked or disengaged position.

  The torque wrench 280 is a 0.9 mm stainless steel hexagonal shaft 282 assembled with a plastic body 284. The torque rating of the wrench 280 is bi-directional, primarily to provide feedback that the anchor 170 has been locked or disengaged. The torque rating allows for a firm fixation of the set screws 119, 152 to the stimulation leads 140, 141 without overtightening.

Trial patch Trial patches are used in conjunction with the pulse generator 107 under trial to provide a clean ergonomic protective cover at the entry point of the stimulation leads 140, 141 in the spinal canal. The patch also covers and includes the trial generator 107. The patch is a large adhesive bandage that is applied to the patient after surgery during the trial phase. The patch completely covers the wires 140, 141 and the generator 107 and is anchored to the patient with an antimicrobial adhesive.

  The patch is a watertight 150 mm × 250 mm antibacterial adhesive patch. The watertight patch allows the patient to take a shower during the trial phase and the antibacterial adhesive reduces the risk of infection. The patch is made of polyethylene, silicone, urethane, acrylate, and rayon.

Magnetic Switch A magnetic switch is a coin-sized magnet that, when placed near the IPG 102, can switch the IPG 102 on or off. The direction in which the magnet faces the IPG 102 determines whether the magnetic switch switches the IPG on or off.

  As shown in FIG. 20, the implantable pulse generator (IPG) 102 includes an RF transceiver module 103, a processor such as the microcontroller 104, and a programmable signal generator such as the ASIC 106. The transceiver module 103 manages wireless communication between the microcontroller 104 and an external remote control (eg, a dongle 200 connected to either the smartphone / mobile 202 or tablet 204).

  In the embodiment shown in FIG. 20, the treatment control module 14 stored in the flash memory 13 of the microcontroller 104 has any of the IPGs 102 except for a stand-alone charger (not shown) that charges the battery 108. It is executed by the microcontroller to control the operation of the components and circuits. Specifically, the treatment control module 14 handles the programming of the RF transceiver module 103 and the signal generator 106, among other functions. Communication between the microcontroller 104, the RF transceiver module 103, and the signal generator 106 is performed over a bus 16, such as a serial peripheral interface (SPI) bus.

  One exemplary microcontroller 104 may be the MSP430F5328 from Texas Instruments, Dallas, Texas, because it has very low power usage, large amounts of memory, integrated peripherals, and small physical size.

  As shown in more detail in FIG. 21, the signal generator 106 includes a memory (control register 18), a timing generator 20, an arbitration circuit 22, a high frequency generator 24, and an electrode driver, all coupled together. 26. All of the components in FIG. 21 access and provide signal parameters stored in the control register 18 through the register bus 28.

  One of the many new features of the IPG 102 is that the control register 18 in the signal generator 106 has all of the signal parameters required to drive the electrodes E1-E32, independent of the microcontroller 104. Having enough memory to store. As a result, the microcontroller 104 programs all of the pulse parameters in the control register 18 and the treatment control module 14 sets the stimulus signal to the signal generator by setting the stimulus enabled pin STIM-EM to logic high. When instructed to generate, the microcontroller 104 may be placed in a standby mode. In the illustrated embodiment where the microcontroller is MSP430F5328, the treatment control module 14 places the microcontroller in LPM3 standby mode. In the LPM3 mode, the master clock (main clock) that drives the instruction execution unit of the microcontroller 104 is stopped, and basically the microcontroller is stopped to save battery power.

  The microcontroller can be activated from standby mode by an interrupt signal IRQ that can be transmitted by the transceiver module 103 when it receives the appropriate instructions from the remote control device.

  As previously mentioned, in IPG 102, electrodes E1-E32 may be grouped into stimulation sets (stimulation channels). Each stimulation channel corresponds to one specific stimulation signal / pattern applied to the associated electrode. In the illustrated embodiment, the IPG 102 can accommodate up to 16 channels (ch1-ch16). Each electrode can belong to one or more channels, up to a maximum number of channels, and each channel can be associated with at least two electrodes, up to a maximum of 32 electrodes. Thus, one electrode can belong to all 16 electrodes.

  For example, as shown in FIG. 25, channel 1 includes electrodes E1 and E2, channel 2 includes E2 and E3, while channel 3 includes E2-E4 and E6-E8. Therefore, the electrode E2 belongs to channels 1, 2, and 3, and the electrode E3 belongs to channels 2 and 3, whereas the electrodes E4 and E5 to E8 belong only to channel 3, and the electrode E1 only belongs to channel 1. Belonging to. In FIG. 25, electrode E5 is not used and therefore does not belong to any channel. Data relating the electrodes to a particular stimulation channel is stored in the control register 18.

  As will be described more fully later herein, for each channel, the IPG 102 has a first phase pulse (pulse 1) of biphasic pulses (see, eg, FIG. 26) for all of the stimulation channels and The amplitude, frequency and duration of both second phase pulses (pulse 2) can be programmed. As seen in FIG. 26, the waveforms of the biphasic pulses including pulse 1 and pulse 2 of the active channel (ch_act) are combined into a signal pulse pattern applied to electrodes belonging to the active channel such as E1 and E2. Become. FIG. 26 shows variations in pulse width and amplitude. As can be seen, the width of pulse 1 is wider than that of pulse 2, but the amplitude for pulse 1 as scaled by pulse scaler 38 (see, for example, the positive pulse of E1) is pulse 2 (See, for example, the negative pulse at E1).

  As an example that the pulse parameters for pulse 1 are independent of the parameters for pulse 2 for maximum flexibility in dealing with pain, and with reference to FIG. 25, the electrodes E1 and E2 The current path between (channel 1) affects the nerve path to the left leg, while the current path between electrodes E2 and E3 (channel 2) affects the nerve path to the right leg To do. If the patient complains of pain with the right leg rather than the left leg, the physician programming IPG 102 may associate electrodes E1 and E2 to channel 1 with a stimulation pulse pattern having a frequency of 100 Hz and electrode E2 to channel 2 And E3 may be associated with a pulse pattern having a frequency of 1000 Hz and a higher current amplitude than channel 1. In this way, a smaller current is applied to the left leg channel 1 and a larger current is applied to the right leg channel 2. The flexibility of the IPG 102 advantageously provides the right amount of current to each affected area of the patient.

  The control register 18 includes a standard read / write register 18 that can be accessed by the SPI bus 16 and the register bus 28. The control register 18 is configured as an 8-bit register array, each having a unique address. The control register 18 is programmed by the microcontroller 104 so that the signal generator 106 stores all the pulse parameters that are necessary to generate the stimulation channel pattern without any intervention from the microcontroller. The Pulse parameters include stimulation channel timing settings, current (pulse amplitude) scaler settings, calibration data, and electrode group parameters.

  For each channel, the control register 18 has the rising and falling edges of the channel itself, the rising and falling edges of each of the two pulses (pulse 1 and pulse 2), the period of the biphasic pulse, the active channel period (channel envelope). Line), and current scaling (pulse amplitude) values for both pulses (pulse 1 and pulse 2). For each channel, control register 18 also stores burst frequency data (as described later herein), such as the burst period for both pulses (pulse 1 and pulse 2). For each channel, the control register 18 also stores data regarding which electrodes E1-E32 belong to that channel. For each channel, and for each channel within that channel, the control register 18 provides source / sink data for both pulses of the biphasic pulse (pulse 1 and pulse 2), ie, each electrode is pulse 1 and It also remembers whether it will source current or sink current during pulse 2. All parameters are defined in relation to the origin and are in microseconds.

  The timing generator 20 is based on the pulse timing parameters stored in the control register 18 and includes pulse 1, pulse 2, and channel pulse “ch” (channel envelope waveform) for all 16 channels. Generate a signal. When all 16 channels are programmed by the clinician, the timing generator 20 generates channel 1, pulse 2, and channel pulse data ch for all 16 channels.

  As an example, the pulse 1 and pulse 2 waveforms of FIG. 26 show the output waveform from the timing generator 20 for a typical biphasic pulse for a particular channel. The channel envelope data ch_act, along with the beginning and end of the active part of each channel, also defines the channel period, which can be defined as the time between two adjacent rising edges of the channel. Pulse 1 and pulse 2 define the beginning and end of each of the two phases of the biphasic pulse.

  Due to the flexibility of the IPG 102, more than one channel may be active at any time when multiple channel stimulation patterns are programmed. The arbiter 22 is designed to solve the overlapping channel (channel contention) problem by ensuring that only one channel is active at any given time. In particular, the circuit in the arbitration unit 22 is designed using two rules. The first rule is that when an active channel is selected (ie, a channel that is currently in progress), all other channels that are going to become active are suppressed and discarded. The second rule is that when two or more channels are about to become active on the simultaneous rising edges of ch, the active channel will be determined based on a predetermined channel priority. .

  In the illustrated embodiment, the arbitrator 22 is programmed so that the lowest numbered channel will be given priority and the remaining simultaneous channels will be discarded. Since IPG 102 has 16 channels (ch1 to ch16), channel 1 has the highest priority while channel 16 has the lowest priority.

  The output of the arbitration unit 22 includes pulse timing signals p1_act and p2_act having the same waveform as that of the active channel pulse 1 and pulse 2. The arbiter 22 is an electrode driver, as well as the channel envelope of the active channel (eg, ch_act as shown in FIG. 26) for use by the high frequency generator 24 and as described later herein. 26 also outputs the channel number (ch_code) of the active channel that will be used by H.26. In the illustrated embodiment, the channel number is a 4-bit code that identifies the number of the active channel. For example, “0001” represents channel 2 while “1111” represents channel 16.

  FIG. 27 provides an example of channel arbitration by the arbitration unit 22. In FIG. 27, channel number 1 has three electrodes E1 to E3, and channel number 2 has two electrodes E1 and E2. Since the channel periods for the two channels are different, they will sometimes overlap. In the figure, channel 1 is active (ch1_active) when channel 2 is about to become active. At that point, the arbiter will execute the first rule, suppress channel 2 and prevent it from becoming active. Therefore, only the pulse 1 / pulse 2 signal for driving the electrode for channel 1 (active channel) is output by the arbitration unit 22. The pulse 1 / pulse 2 signal and the channel envelope signal for channel 2 (ch2_active) are indicated by dotted lines indicating that they are suppressed by the arbitration unit 22.

  The high frequency generator 24 receives the waveforms of p1_act and p2_act from the adjustment unit 22, and determines whether to modulate the received signal based on the parameters stored in the control register 18. If the determination is negative, the high frequency generator 24 passes the unchanged received pulse signal to the electrode driver 26.

  However, if the decision is affirmative, the high frequency generator 24 modulates the received signal to the burst frequency programmed in the control register 18. The burst frequency is higher than the frequency of the received signals p1_act and p2_act.

  The electrode driver 26 receives the output (p1, p2, and ch_code) of the high frequency generator 24, amplifies the received signal according to the pulse amplitude parameter stored in the control register 18, and is applied through the electrodes E1 to E32. Output the final stimulus pattern for each channel. As described above, the burst pulse parameters stored in the control register 18 are such that an asymmetric pulse shape can be generated by the electrode driver 26, with positive and negative pulses having different frequency values. 1 and pulse 2 have different frequency values.

  FIG. 22 is a more detailed functional block diagram of the high-frequency generator of FIG. The high frequency generator 24 includes a burst generator 30 and a burst multiplexer 32.

  The burst multiplexer 32 receives the burst parameter stored in the control register 18 for all channels, selects the burst parameter associated with the active channel, and outputs the selected burst parameter to the burst generator 30. Specifically, for burst options, there are pulse 1 / pulse 2 burst register pairs for each channel, and in the illustrated embodiment, there are a total of 32 registers for 16 possible channels. The burst multiplexer 32 is a vector MUX that selects the pulse 1 / pulse 2 register pair corresponding to the active channel number. The select line 34 to the burst multiplexer 32 is the active channel number (ch_code) from the arbitration unit 22 that identifies the active channel at any given time. The selected burst parameter for pulse 1 / pulse 2 is transmitted to the burst generator 30.

  Within burst generator 30 there are two individual burst generator circuits, one for pulse 1 and one for pulse 2. The pulse 1 and pulse 2 signals belonging to the active channel are passed from the arbitration unit 22 to the burst generator circuit. If the pulse 1 / pulse 2 burst parameter data stored in the control register 18 is zero (and therefore the selected burst parameter to the burst generator 30 is also zero), no burst occurs, in this case The pulse 1 / pulse 2 signal from the arbitration unit 22 is transmitted to the electrode driver 26 without being changed. When the pulse 1 / pulse 2 burst register stored in the control register 18 is programmed, the pulse 1 / pulse 2 duration is based on the burst parameter selected from the burst multiplexer 32 and the corresponding programmed burst signal. Will be replaced (it will be modulated into a burst signal).

  As an example, FIG. 27 shows that pulse 1 for channel 1 is programmed with high frequency modulation, while pulse 2 for the channel is not programmed. Specifically, one pulse 1 pulse is modulated (replaced) by a burst pulse with a frequency 5 higher. Therefore, the frequency of the new modulation pulse 1 is five times that of the original pulse 1 signal.

  FIG. 23 is a more detailed functional block diagram of the electrode driver 26 of FIG. The electrode driver 26 includes a pulse amplitude multiplexer 36, a pulse scaler 38, and a current driver 40.

  The pulse amplitude multiplexer 36 receives the amplitude parameter stored in the control register 18 for all channels, selects the amplitude parameter associated with the active channel, and outputs the selected amplitude parameter to the pulse scaler 38.

  Specifically, 512 bytes (16 × 32 bytes—32 bytes for each channel) are stored in the control register 18 to store the pulse amplitude data. Each of the 16 channels is associated with 32 bytes, each byte representing pulse amplitude information for each pulse 1 and pulse 2 of the 32 electrodes E1-E32. 7 bits from 1 byte are used to store the amplitude information for pulse 1 and pulse 2 and the remaining bits (MSB) are the polarity of the pulse at the relevant electrode as described later in this specification. Define

  Similar to burst multiplexer 32, pulse amplitude multiplexer 36 is a vector MUX that selects 32 bytes of the amplitude parameter for pulse 1 / pulse 2 corresponding to the active channel number. The select line 34 to the amplitude multiplexer 36 is the active channel number (ch_code) from the arbiter 22 that identifies the active channel at any given time. The selected amplitude parameters of pulse 1 / pulse 2 for all 32 electrodes E1-E32 are transmitted to the pulse scaler 38. The pulse scaler 38 outputs the amplitude magnification for all electrodes of the active channel. Therefore, the pulse scaler 38 includes 32 identical scalers corresponding to the 32 electrodes E1 to E32. In the illustrated embodiment, each scaler includes a D / A converter that converts the digital amplitude values into corresponding analog values I1-I32.

  As mentioned above, the signal generator 106 supports asymmetric pulse amplitude features, which means that the amplitude for pulse 1 and pulse 2 can be different. The amplitude scaling data is stored in the control register 18. In the embodiment shown, 4 bits, 2 bits for pulse 1 and 2 bits for pulse 2, are used to specify the scaling factor for pulse 1 and pulse 2 for each electrode. Further, the signal generator 106 can accommodate asymmetric pulse width variations between pulse 1 and pulse 2.

  The pulse scaler 38 adjusts the amplitude parameter for pulse 1 and pulse 2 by the associated magnification stored in the associated 2 bits. In the illustrated embodiment, the pulse scaler 38 shifts the contents of the selected amplitude parameter (7-bit data from the amplitude multiplexer 36) to the right by the number stored at the corresponding 2-bit magnification. To perform the scaling function. Since there are four possibilities in a 2-bit number (0, 1, 2, or 3), the amplitude can be reduced to 1/2, 1/4, or 1/8.

  For example, assuming that the active channel is channel 1, the amplitude parameter for channel 1 electrode E1 in control register 18 is binary "1111111" (decimal 127), while the associated 2-bit magnification is In pulse 1, the binary number is “11”, and in pulse 2, the binary number is “01”. Equation 127 represents a current of 12.7 mA. In pulse 1, the 7-bit amplitude content is shifted to the right by 3 bits (binary "11") for pulse 1 and 1 bit (binary "01") for pulse 2. Therefore, the pulse scaler 38 pulses the amplitude of 127 to 15 (binary “1111”) in pulse 1 corresponding to 1.5 mA current for pulse 1 and 6.3 mA current for pulse 2. 2 is reduced to 63 (binary number “111111”).

  In the illustrated embodiment, the scale factor parameter is stored with the unused bit burst parameter and passed from the burst multiplexer 32 to the pulse scaler 38. However, a dedicated memory can be assigned to the control register 18. Note that the stimulus programming software needs to ensure that at any given moment, the algebraic sum of all electrode currents is zero and the dc average of the current per cycle at each electrode is also zero. .

  Another function performed by the pulse scaler 38 is to convert the scaled amplitude digital value to an analog output current that is linearly proportional to the value of the digital value. The analog output current for each electrode is supplied to the current driver 40. In the embodiment shown, the analog output current corresponds to 1/20 of the actual current supplied to the associated electrode.

  The current driver 40 amplifies the analog output current from the pulse scaler 38, and switches the amplified current to the appropriate electrodes E1 to E32 based on the pulse parameters stored in the control register 18. Since there are 32 electrodes, there are 32 electrode drivers 40 in the illustrated embodiment. In the illustrated embodiment, each current driver 40 is a current driver that amplifies the input signal from the pulse scaler 38 by a factor of 20.

  FIG. 24 is a functional diagram of two of the current drivers of FIG. Each electrode driver 40 can be a current source capable of sinking or sourcing current whose amplitude is based on a control current signal coming from a pulse scaler 38. In the illustrated embodiment, the current source includes a pair of NMOS current source 42 and PMOS current source 44 coupled in series between the voltage supply and ground. Each of the PMOS and NMOS current sources can be implemented as a current mirror as is well known. NMOS current source 44 sinks current from the associated electrode to ground, while PMOS current source 44 sources current from the positive voltage supply to the associated electrode. Each electrode driver 40 includes switches 46 and 48 connected in series between a voltage supply and ground to source or sink current.

  The control register 18 stores pulse parameters relating to whether a particular electrode will source or sink current. In the illustrated embodiment, bit 7 (MSB) of each byte of the pulse amplitude parameter supplied to burst multiplexer 32 indicates that a particular current source 40 will sink current or source current. Used to specify what will be done. If the bit is zero, switch 46 will be turned on during pulse 1, while switch 48 will be turned off, and switch 46 will be turned off during pulse 2. On the other hand, the switch 48 will be turned on. If the bit is set (ie, the bit is “1”), during pulse 1, switch 46 will be turned off while switch 48 will be turned on and the pulse During 2, switch 46 will be turned on while switch 48 will be turned off.

  FIG. 24 shows current paths for two electrodes Ej and Ek when bit 7 is set when Ej = 0 and bit 7 is set when Ek = 1. During pulse 1, Ej switch 46 and Ek switch 48 are turned on, creating a current path 50. In that case, the Ej PMOS current source 42 will source current, while the Ek NMOS current source 44 will sink current to ground. On the other hand, during the pulse 2, the switch 46 for Ek and the switch 48 for Ej are turned on, and a current path 52 is created. In that case, the PMOS current source 42 for Ek will source current, while the NMOS current source 44 for Ej will sink current to ground.

  The pulse shape at electrode Ej will appear similar to the E1 waveform as shown in FIG. 26, while the pulse shape at electrode Ek will appear similar to the E2 waveform.

  The specific embodiments described above represent only a portion of the methods of practicing the present invention. Many other embodiments are possible within the spirit of the invention. Accordingly, the scope of the invention is not limited to the above specification, but instead is indicated by the appended claims, along with the full scope of their corresponding patents.

FIG. 6 depicts various components that can be included in a spinal cord stimulation system, according to embodiments, during trial implantation and permanent implantation. FIG. 3 depicts an exploded view of an implantable pulse generator (IPG) assembly, according to an embodiment. 1 depicts a feedthrough assembly of an implantable pulse generator (IPG) assembly, according to an embodiment. 1 depicts a lead contact system of an implantable pulse generator (IPG) assembly, according to an embodiment. 1 depicts a lead contact assembly of an implantable pulse generator (IPG) assembly, according to an embodiment. 1 depicts a head unit assembly of an implantable pulse generator (IPG) assembly, according to an embodiment. 1 depicts an RF antenna of an implantable pulse generator (IPG) assembly, according to an embodiment. 1 depicts a percutaneous lead according to an embodiment. 1 depicts a paddle-type lead according to an embodiment. Figure 8 depicts a conductor extension, according to an embodiment. 1 depicts a conductor splitter, according to an embodiment. Figure 3 depicts a sleeve anchor, according to an embodiment. 1 depicts a mechanical locking anchor, according to an embodiment. FIG. 4 illustrates communication via a wireless dongle with a tablet / clinician programmer and a smartphone / portable / patient programmer during trial and / or permanent implantation, according to an embodiment. 1 depicts a chuhe needle, according to an embodiment. Figure 8 depicts a stylet, according to an embodiment. 1 depicts a passing elevator, according to an embodiment. 1 depicts a tunneling device, according to an embodiment. 1 depicts a torque wrench, according to an embodiment. 2 is a functional block diagram of several components in an implantable pulse generator, according to an embodiment. FIG. It is a functional block diagram of the signal generator of FIG. It is a functional block diagram of the high frequency generator of FIG. It is a functional block diagram of the electrode driver of FIG. FIG. 24 is a functional illustration of two of the current drivers of FIG. Fig. 4 illustrates an exemplary electrode waveform for an active channel, according to an embodiment of the present invention. Fig. 4 shows the classification of electrodes for different channels according to an embodiment of the invention. FIG. 4 illustrates an exemplary electrode waveform showing an asymmetric pulse pattern between positive and negative pulses, according to an embodiment of the present invention. FIG. FIG. 4 illustrates an exemplary electrode waveform for two arbitrated stimulation channels according to an embodiment of the present invention.

Claims (18)

An implantable pulse generator (IPG) for generating a spinal cord stimulation signal for a human body,
A control register that stores stimulation signal parameters for a plurality of channels, wherein each channel can be associated with at least two electrodes of the plurality of electrodes, each channel stimulating the associated electrode A signal pattern, wherein the signal parameter includes a burst parameter;
A timing generator adapted to generate a timing signal representing a stimulation signal for the plurality of channels according to the stored signal parameters;
Adapted to determine whether to modulate the timing signal received from the timing generator and whether to modulate the received timing signal at a burst frequency according to the stored burst parameters; An implantable pulse generator (IPG) comprising: a high frequency generator;   The IPG of claim 1, wherein when the high frequency generator determines that the received timing signal is not modulated, the high frequency generator outputs the received timing signal whose frequency is not changed. The high-frequency generator is
Receiving the burst parameter stored in the control register;
Selecting the burst parameter associated with an active channel;
The IPG of claim 1, comprising a burst multiplexer that outputs the selected burst parameter.   4. The high frequency generator includes a burst generator that receives the selected burst parameter from the burst multiplexer and modulates the received timing signal in accordance with the received burst parameter. IPG. A processor having memory;
A treatment control module stored in the memory and executable by the processor, the treatment control module transmitting the stimulation signal parameters to the control register and placing the processor in a standby mode Adapted to generate the timing signal according to the stored signal parameters without intervention from the processor while the processor is in the standby mode. The IPG of claim 1, further comprising a treatment control module adapted.   The IPG of claim 5, further comprising a transceiver module adapted to wirelessly communicate with a remote control and to transmit an interrupt signal to the processor to wake up the processor from the standby mode. .   The IPG according to claim 1, further comprising: an arbitration unit that continuously receives the generated timing signal and selects one of the plurality of channels as an active treatment channel.   The IPG according to claim 7, wherein when the one channel is selected as the active treatment channel, the arbitrator is adapted to suppress all other channels.   When the generated timing signal indicates that two or more channels are about to be active at the same time when no active channel is selected, the arbitration unit is configured to The IPG of claim 7, wherein the IPG is adapted to select the one channel as the active treatment channel based on a predetermined priority of the channel. A method of generating a spinal cord stimulation signal for a human body,
Storing stimulation signal parameters for a plurality of channels in a control register, wherein each channel can be associated with at least two of the plurality of electrodes, each channel being associated with the associated electrode A stimulation signal pattern of the signal, wherein the signal parameter includes a burst parameter;
Generating a timing signal representing a stimulation signal for the plurality of channels according to the stored signal parameters;
Determining whether to modulate the timing signal received from the timing generator;
Modulating the received timing signal at a burst frequency according to the stored burst parameter if it is determined that the timing signal is to be modulated.   The method of claim 10, wherein if it is determined that the timing signal should not be modulated, the received timing signal is output without changing frequency.   Further comprising selecting the burst parameter associated with an active channel of the plurality of channels, wherein the modulating step comprises modulating the received timing signal according to the selected burst parameter. 11. The method of claim 10, comprising. Transmitting the stimulus signal parameter to the control register by a processor;
11. The method of claim 10, further comprising: placing the processor in standby mode, wherein the timing signal is generated without intervention from the processor while the processor is in the standby mode. .   The method of claim 13, further comprising receiving an interrupt signal from a transceiver module by the processor to wake the processor from the standby mode.   The method of claim 10, further comprising: continuously receiving the generated timing signal; and selecting one of the plurality of channels as an active treatment channel.   16. The method of claim 15, further comprising suppressing all other channels when the one channel is selected as the active treatment channel.   A predetermined priority of the two or more channels when the generated timing signal indicates that two or more channels are about to be active at the same time when no active channel is selected; The method of claim 15, further comprising selecting the one channel as the active treatment channel based on degree.   The method of claim 10, wherein the signal parameter includes an identification of all electrodes belonging to each channel, and each electrode can be associated with more than one channel.