Device materials, handling, and upgradability

Implant locations and form factor

The physical form factor and volume of the pulse generator is derived by balancing the clinical therapy stimulation requirements with the technology and engineering size, weight, and power constraints of the implanted device. The form factor is governed by a number of attributes including the number of therapy electrodes, power source type and energy capacity, and the stimulation output amplitude, duty cycle, and frequency.

Significant technological advances in battery energy density, electronic circuitry miniaturization, and connector density have enabled IPG sizes to be reduced. Most IPGs for DBS and SCS are between 10 and 20 cm ³ in total volume which is dominated by the battery volume. While integrated circuits have miniaturized according to Moore's law, battery energy density has not [ ]. Some manufacturers have adopted a no-battery and no hermetic package approach such as the StimWave Freedom to reduce implanted volume, at the expense of requiring an externalized belt or patch on the overlying skin nearby to be worn during therapy delivery.

The IPG is typically implanted in a subcutaneous or submuscular pocket located below the subclavicular region, lower abdomen or upper buttocks ( Fig. 8.2 ). The size of the implanted device is primarily driven by the battery capacity and the number of header ports. More therapy electrodes require additional header ports which increase the device volume. In general, rechargeable devices or transcutaneous inductively coupled devices are smaller than nonrechargeable IPGs. Neuromodulation therapies that require less stimulation amplitude and fewer electrodes such as Dorsal Root Ganglia stimulation for chronic pain enable a smaller IPG to be realized. These smaller devices are implanted in different areas of the body and typically are implanted closer to the target therapy site. For example, StimWave Freedom promises that it can be implanted near the DRG, adjacent to the dorsal columns, or near peripheral nerves for any type of neuromodulation. Similarly, inductively powered cochlear implants are small enough for implantation in the mastoid process in the skull. Implanting the device closer to the target therapy site also has advantages in minimizing lead length. Recent RF “mid-field” charging methods provide slightly improved penetration depth for delivering energy to smaller-sized miniature implanted devices. By using 800 MHz–1.2 GHz RF antennas, the implanted receiver is reduced to an inefficient antenna which is smaller in volume than conventional inductive coils designed to operate at MHz frequencies. The dielectric constant and loss tangent of physiological saline, fat, bone, blood, and other tissues attenuate RF fields by approximately 30 dB/cm at 1 GHz. Therefore, RF powering methods have inherently poor coupling efficiency to the implanted device and require significant external transmit power levels and optimization methods such as beamforming to reduce energy losses and prevent tissue heating by exceeding the specific absorption rate (SAR) limits. Though touted as superior power transfer, mid-field powering lacks superiority for most applications—especially those requiring larger stimulation current such as DBS and SCS.

A small number of device complications around the IPG continue to occur in patients. The device shapes and form factor are designed to minimize tissue reaction and skin erosion by incorporating large-radii edges and header smoothness. The IPG thickness is especially important to minimize surgical pocket complications. Lower abdomen implant locations can tolerate approximately 12 mm thick and larger devices. In contrast, the subclavicular region requires thinner and smaller devices to prevent skin erosion. Smaller devices also provide an improved cosmesis effect for the patient to better conceal the system within subcutaneous tissues. Device form factors for SCS are shown in Fig. 8.3 [ ].

Implant considerations for recharge and telemetry

Implanting depth is an important consideration for recharge and telemetry. Implanting the device too shallow can result in skin erosion, while implanting the device too deep can result in poor recharge coupling or loss of inductive telemetry. Typical recharging depths range may be performed from 1 to 3 cm deep using conventional inductive coils placed within the titanium enclosure. Improvements in recharging have been achieved by incorporating inductive loops outside of the hermetic package (e.g., Neuros Medical Altius) to prevent the eddy current losses associated with charging through titanium. The receiving recharge coil is typically located external to the hermetic titanium package in the implanted device header to improve the efficiency of the power transmission. When the receiving coil is located external to the hermetic titanium enclosure, it is packaged in magnetically transparent materials such as polyurethane, silicone rubber, polysulfone, ceramic, glass, or biocompatible epoxy. IPG charging coils outside of the titanium enable faster charging at potentially greater depths. For midfield chargers, an RF antenna is positioned outside the titanium enclosure as well.

Other improvements have included battery chemistries which support charging at a faster C-rate without reducing the long-term battery performance or lifetime. In general, battery energy density and charge capacity has not increased every year in the same way semiconductor technology has miniaturized according to Moore's law. Looking forward, IPGs for SCS are unlikely to undergo significant miniaturization from rapid improvements in battery capacity [ ]. The header connection port volume is one potential area for size reduction, similar to pacemakers, which have adopted the newer 0.9 mm canted coil springs.

The transition to 2.4 GHz communication from the MICS band has also improved communication telemetry distance due to the higher regulated power levels in the 2.4 MHz ISM bands. Previous MICS devices required relatively close proximity due a 25 μW (−16 dbm) maximum transmit power, while 2.4 GHz connections can be performed from greater distances with transmit power as high as 0 dbm. Forward error correction methods can further improve range by 5 dB with reduced data rates. While the MICS radios previously provided the lowest energy communication available, numerous 2.4 GHz radios can be duty cycled to have a similar telemetry current consumption of approximately 1 μA quiescent current.

IPG recharge efficiency reduces and charging time increases significantly for deeper implant depths. The efficiency of recharging is impacted by whether the recharging coil is positioned in the header or inside the Titanium enclosure. Charging times can range from 0.5 to 3 h depending, with the fastest being from devices with charging coils on the outside of the hermetic enclosure. Recharging can be a significant burden for patient compliance, especially as high-frequency waveforms (1–50 kHz) consume significantly more battery life and require recharging every 1–2 days. Burst devices and tonic waveforms require far less energy and have a reduced patient recharging burden.

Device materials and biocompatibility

Materials that are implanted in the human body must be biocompatible and biostable for delivering stimulation for a duration of 5–10 years. Device manufacturers ensure that materials implanted in the body meet all required standards for implantable medical devices. International standards for biological evaluations of medical device materials include ISO 10993-1:2018 which outlines a set of comprehensive tests and protocols required for medical devices. The materials are categorized based on the duration of body contact and the type of body tissue that contact that the medical device. The standard outlines four types of body contact: noncontacting medical devices, surface-contacting devices, external communicating devices, and implanted devices.

Implanted medical devices are further defined based on the specific implantation locations of the device. These sites are categorized into two different groups: direct tissue/bone or blood contact. In addition to the site of the implant, the required tests also vary depending on the duration of the implant or contact. The standard also specifies tissue contact duration such as limited (less than 24 h), prolonged (24 h–30 days), or permanent (more than 30 days). Implanted neurostimulation devices have been evaluated by these biological tests including cytotoxicity, sensitization, irritation or intracutaneous reactivity, systemic toxicity, subacute and subchronic toxicity, genotoxicity, and hemocompatibility to ensure biocompatibility.

In addition to biocompatibility, the biostability of the materials and designs that are implanted is evaluated using other mechanical, environmental, and electrical safety standards. Biostability evaluation includes detailed mechanical, electrical, and chemical characterization of the material properties after being subjected to the human body for the defined exposure. Materials that are resistant to degradation and corrosion are key characteristics that the device manufacturers consider when selecting materials for chronic, implantable systems such as neurostimulation devices.

The materials of the implanted device that have direct tissue contact for permanent exposure periods include the device enclosure, device header and, with some designs, the recharge receiving antenna. Titanium is the most common material used for the hermetic package. It exhibits high levels of corrosion resistance, is nonmagnetic, lightweight, nontoxic and biologically compatible with human tissue and bone. Titanium also has excellent mechanical strength and durability characteristics. It is often formed into thin-walled halves that are seam-welded together to create the hermetic enclosure or by deep-drawn methods with a welded lid. Titanium has several implantable grades. Commercially pure titanium, such as Grade 1 or 2, is most commonly used. Commercially pure titanium has excellent formability and elongation which allows cold working and forming of custom device shapes and a relatively tight bend radius. Other titanium alloys, such as Grade 9 or 23, include other alloys such as aluminum or vanadium which increase the electrical resistivity of the material. This increase in electrical resistivity results in lower magnetic eddy current loss during inductive power transfer used with rechargeable systems. These alloys do not have the formability or elongation properties of commercially pure titanium which translates into larger bend radius constraints for the device form factor. However, the improved recharge performance allows the device to be implanted deeper.

Biocompatibility is considered a large risk where new materials are used. However, biocompatibility testing remains important to assess residuals associated with the manufacturing process even if all materials have known biocompatibility. Risk analysis should be performed according to ISO 10993-1:2018 to determine the risk of a particular device material or manufacturing process.

Multi-contact electrode systems provide the ability to steer current between multiple cathodes and returns using one or more electrodes on the lead. Modern IPGs can provide monopolar stimulation with the titanium hermetic package as a return common electrode. The titanium may also be connected as a low-impedance source to provide anodal intensification up to approximately 5% of the stimulation amplitude.

The electrical contacts in the device header are typically made from titanium, SS316L platinum, or platinum iridium alloy materials. Commonly used insulating materials include polyurethane, silicone rubber, tecothane, polysulfone, and biocompatible epoxy.

Lead geometries

Percutaneous SCS leads and DBS leads have cylindrical contacts to enable the lead to be inserted through a 4–5 french (1.5 mm) inner diameter needle and advanced in a nondamaging fashion. Stylets may be inserted into the bore of the lead to help steer it into position. The requirements for contact spacing and surface area vary between SCS and DBS with most leads providing between 8 and 16 cylindrical contacts for circumferential stimulation. The contacts of a percutaneous SCS lead may span from 25 to 78 mm to provide stimulation points over one to three vertebral levels. DBS leads normally have smaller contacts more tightly spaced to target small anatomical structures of the brain with finer granularity. The development of directional leads now reduces negative off-target stimulation of DBS leads.