The Technology: The Anatomy of a Spinal Cord and Nerve Root Stimulator: The Lead and the Power Source

Blackwell Publishing IncMalden, USAPMEPain Medicine1526-2375American Academy of Pain Medicine20067S1S27S34 Original ArticleThe TechnologyBradley PAIN MEDICINE Volume 7 Number S1 2006 The Technology: The Anatomy of a Spinal Cord and Nerve Root Stimulator: The Lead and the Power Source Kerry Bradley, MS Advanced Bionics Corporation, A Boston Scientific Company, Valencia, California, USA ABSTRACT ABSTRACT Introduction S Spinal cord stimulation (SCS) has been used for over 30 years to treat a variety of pain conditions. Success rates in SCS have improved due to more judicious patient selection, standardization of techniques, understanding of technical and clinical goals, and, quite importantly, advances in the stimulation technology. This article discusses characteristics of the stimulating leads and the power source and how they contribute to optimizing SCS therapy. Key Words. Spinal Cord Stimulation; Lead; Electrode; Rechargeable pinal cord stimulation (SCS) has been used for over 30 years to treat a variety of pain conditions. The most successful results have been in the treatment of neuropathic pain, often with a vascular component [1]. In SCS, an array of stimulating metal contacts is placed in the dorsal epidural space. These contacts, connected to a pulse generator, are then programmed in contact combinations of anodes and cathodes, with the intent of generating an electric field, which stimulates the axons of dorsal root (DR) and dorsal column (DC) fibers in the spinal column. Stimulation of these fibers results in inhibition of activity in the lateral spinothalamic tract (a known pain pathway), as well as increased activity in descending antinociceptive pathways [2]. Clinically, SCS creates an epiphenomenon called paresthesia, which has been described by patients as a tingling sensation in their skin. In the application of SCS, a crucial technical result that is correlated to a successful clinical outcome is the overlap of the painful areas with paresthesia [3]. In order to optimize this result, the SCS system, like any neurostimulation system, must provide adequate neural selectivity; that is, maximal control over stimulation of the targeted nerves while avoiding stimulation of undesired neurons. Technology has accelerated in Reprint requests to: Kerry Bradley, MS, 25129 Rye Canyon Loop, Valencia, CA 91355, USA. Tel: 661-362-1903; Fax: 661-362-1508; E-mail: kerrybradley@earthlink.net. the last decade to continually improve this capability. This article will discuss two categories of technology that enable the improvement in stimulation selectivity, the lead and the power source. Stimulation Selectivity In common electrical stimulation applications, such as cardiac pacing and defibrillation, there are usually no more than 1 2 stimulating contacts near the tissue to be stimulated, e.g., a bipolar pair of contacts placed in the right ventricular apex for bradycardia pacing. Also, the physiologic result of the stimulation is usually binary. For example, in a cardiac pacer, enough charge needs to be delivered to the endocardial stimulation site to trigger a heart contraction, and no more [4]. In neurostimulation applications, however, the nerves surrounding the contacts often have multiple functions (e.g., sensory, motor, autonomic), complex geometries (e.g., DRs, DCs), and differing thresholds [5]. The intent of neurostimulation is to activate only those nerves that create the therapeutic benefit. This benefit is proportional to the number of targeted nerves stimulated and inversely related to the number of undesired nerves stimulated. Thus, in neurostimulation, recruitment is a graded function and should be confined to the targeted nerves. In sum, neurostimulation differs greatly from other electrical stimulation in that it requires a high degree of stimulation selectivity. This requirement is the main driving force behind the technological devel- American Academy of Pain Medicine 1526-2375/06/$15.00/S27 S27 S34

S28 Bradley opments in lead and stimulator power source design. The Lead Two types of lead are presently in use for SCS: percutaneous and paddle. The main differences between the two types have been described in the literature [6 8]. Briefly, percutaneous leads are catheter-like, being cylindrical and flexible, and these mechanical features allow them to be introduced into the spinal column via a spinal needle (Touhy, Husted, or the like). The contacts on the leads are cylindrical, most commonly 3 mm in length, and made of platinum iridium alloys. The lead itself consists of an isodiametric polyurethane body containing wires that connect the intraspinal contacts to a proximal connector, for connection to the stimulator electronics [6]. Paddle leads, in contrast, differ in that their distal ends are relatively flat and wide, with round or rectangular plate contacts placed on one side of the flat portion, and are constructed of a flexible silicone. Given their shape, most paddle leads must be placed via surgical procedure, such as laminotomy or laminectomy [8]. The proximal portion of the paddle is typically similar to the percutaneous leads, terminating in one or two long isodiametric bodies that provide connection to the stimulating electronics. From a stimulation selectivity standpoint, the key differences between lead types can be broken down into several categories: number of contacts, intercontact center-to-center spacing, contact length, contact width, contact shape, intraspinal lead shape, and for paddle designs with multiple columns, the relative orientation of the contacts among columns [9,10]. The specifics of these features will be described for both percutaneous and paddle-type leads. Percutaneous Leads All modern percutaneous leads have multiple contacts, ranging from 4 to 8 (see Figure 1). Percutaneous leads with multiple contacts have been shown to provide statistically improved long-term outcomes [3], primarily due to their ability to Figure 1 Percutaneous leads of 4 mm and 9 mm centerto-center spacing, respectively. Figure 2 Effects of center-to-center spacing on stimulation field. Widely spaced contacts on left have greater rostrocaudal and lateral stimulation field, where narrowly spaced contacts on right have more confined stimulation field, laterally and rostrocaudally [15] (reprinted with permission of the authors). allow for postoperative reprogramming to capture stimulation targets following micro- or macromigration. For leads placed at or near the physiological midline, mathematical modeling has shown that the center-to-center spacing between active contacts is a key factor in determining the fiber types to be stimulated. Lead models where the anode(s) and cathode were separated by a centerto-center spacing of 4 mm have been shown to be mathematically optimal for preferentially stimulating DC fibers over DR fibers [11,12]. Clinically, leads that use center-to-center anode cathode separations of <10 mm have been shown to be superior at capturing challenging pain targets, such as covering the low back with paresthesia [13,14]. The value of narrow center-to-center spacing is that it confines the stimulation field along the axis of the lead, as well as increases the peak of the stimulation field beneath the cathode (see

The Technology Figure 2). Typically, the DC fibers are the primary targets in SCS, as the DC contains fibers of passage from all cutaneous regions caudal to the stimulating contacts, so stimulating DC fibers means maximizing the stimulation paresthesia coverage at any segmental level. Also, the DC contains predominantly touch and vibration sensory fibers, which are least likely to generate functional deficits (motor or autonomic side effects) when stimulated maximally. In contrast, primary activation of DR fibers would generate paresthesia in only a few dermatomes at the segmental location of the stimulating contacts [16,17]. While this may be a specific stimulation goal for a given patient, i.e., where the SCS device functions as a dorsal root stimulator, in general, broad regions of paresthesia should be more likely to maintain paresthesia pain overlap if the pain happens to migrate to different body locations, as in complex regional pain syndrome [2]. Additionally, DRs carry a wide range of sensory fibers (pain as well as touch and vibration) and make more direct synaptic reflex links with the musculoskeletal system. Maximal stimulation of DR fibers may then result in pain as well as motor deficits. Thus, it may be clinically advantageous to use leads with narrow center-tocenter spacing as this should result in more selective activation of DC fibers. The cost of using narrow combinations is an increase in stimulation energy [11], but such selectivity may mean the difference between a successful and failed technical result [14]. Another key aspect of narrow intercontact spacing is the increased targeting resolution enabled by multiple contacts in a confined region. In keeping with the goal of stimulation selectivity, a prominent factor in the success of an SCS implant is the placement of the stimulating contact(s) as close as possible to the desired neural targets to optimize paresthesia pain overlap [18]. This is due to the nature of electric fields: the activation field strength rapidly falls off (inverse cube root) with distance from the contact to the nerve [19], so placement of the contact will determine which nerves are stimulated first. If first-stimulated nerves are not the targeted ones, an increase in stimulation energy (amplitude or pulse width) is necessary to reach the neural targets. This can mean sensations in undesired locations and, possibly, motor activation prior to target capture [17]. Indeed, this is the reason why significant physician effort is expended during intraoperative placement of the lead [20]. Having a large number of contacts within a confined space increases the likelihood of S29 having a stimulating contact close to the neural target. Additionally, slight migration of the lead may result in a shift of the activating field and loss of the desired paresthesia. If a large number of narrowly spaced contacts are employed on a lead, slight migrations may be accounted for by reprogramming [3]. Paddle Leads Modern paddle leads also have multiple contacts, ranging from 2 to 16, arranged in single or multiple columns (see Figure 3). Paddle leads have been shown to have superior technical results when compared with percutaneous leads [21]. The primary reason for the improved stimulation appears to be mechanical. The flat, relatively rigid shape of the paddle lead tends to compress the pliable dorsal dural sac. This reduces the distance between the contact and the DCs, and thus increasing the likelihood of primary DC fiber activation [22]. Nonetheless, percutaneous leads can still achieve technical results similar to paddles. In a prospective, randomized comparison, paddle leads provided statistically superior paresthesia pain overlap when programmed with simple bipolar combinations. However, when the contact combinations were exhaustively tested and the optimal contact combination was programmed both for percutaneous lead and for the paddle, the statistical difference disappeared [21]. This result indicates that stimulation selectivity is a result of several factors: number of contacts, contact and lead geometry, and electric field shaping via the anode cathode combination [23]. Additionally, in comparisons of percutaneous and paddle leads, the radial nature of the percutaneous contact allows current flow in a radial direction. This has been shown to increase the likelihood of dorsal current flow that can result in uncomfortable axial sensations via stimulation of sensory fibers in the ligamentum flavum at therapeutic current levels. On paddle leads with ventrally directed contacts, this result was not observed [22]. Although the nature of percutaneous lead implant makes it difficult to control the dorsoventral position of the contacts, mathematical modeling has shown that percutaneous contacts, when placed directly on the dura mater, achieve near unidirectional current flow toward the spinal cord, similar to that of paddle leads (see Figure 4) [23]. Indeed, those same models have shown that placement of a single percutaneous lead on midline can achieve slightly superior DC

S30 Figure 3 Typical two-column paddle leads. penetration when compared with paddles with wide (4 mm) contacts [23]. The relative orientation of contacts on paddle leads with multiple columns of contacts is also of importance. Mathematical models suggest that superior DC penetration is achieved when the active contacts are arranged on the lead in perfect parallel, that is, with contacts aligned rather than staggered [10]. When contacts are staggered, the cathodic fields do not optimally superimpose and, if two columns of narrow bipoles or guarded cathodes are programmed, greater anode-cathode cancellation can occur. Novel paddle designs that seek to maximize DC penetration have been developed and clinically evaluated. The most notable of these was the transverse tripole design [24]. Inspired by results from mathematical modeling, which revealed that DR fibers would be typically activated prior to DC fibers in nominal SCS, the transverse tripole lead sought to reverse this recruitment order by judicious contact placement on the paddle. The transverse tripole consisted of two long lateral contacts to the left and right sides of the approximately 10- Bradley mm-wide paddle, with two smaller contacts placed longitudinally in the center of the paddle, between the anodes. The intent of this design was to program the lateral contacts as anodes proximal to the DR fibers. The anodic field would then hyperpolarize the roots and allow for the center contacts, programmed as cathodes, to achieve selective activation of the DC fibers and deeper penetration into the DCs. Clinically, the system was evaluated in a multicenter study, which showed that very selective paresthesia could be generated in many dermatomes caudal to the segmental level of the paddle [25]. The cost of such selectivity, again, was increased stimulation energy [26]. Single Versus Dual Leads In the first two decades of SCS, a single lead was placed as close to physiological midline as possible in each patient. This was due primarily to limitations in technology, as only a single lead with a few contacts could be driven by early stimulators [27]. In recent years, with the advent of stimulators which can deliver stimulation to two leads, there has been a trend toward the implant and use of dual, parallel leads over a specific spinal segment. There have been several reasons for doing so: 1) placement of two leads, one to each side of the physiological midline, can provide greater control over the laterality of the paresthesia pattern for patients with bilateral pain, a common condition [7]; 2) it is known that percutaneous leads tend to migrate slightly in the acute phase of implant and having two leads, with a larger number of contacts at the targeted spinal segment, allows for maximum programming flexibility in order to recapture the neural target [28]; and 3) mathematical modeling has shown that, for certain anatomic and geometric conditions, two leads straddling the physiological midline can create stimulation fields that superimpose and achieve good penetration in the DCs [10]. All of these reasons have certain A B Figure 4 Effects of dorsoventral position of percutaneous lead in epidural space. (A) Intimate dural contact yields minimal dorsal current flow. (B) If contact is located more dorsally in epidural space, current flow becomes more radial and may activate neurons in posterior regions [23] (reprinted with permission of the authors).

The Technology merit. Nonetheless, the placement of a single lead on or very near the physiological midline has been shown clinically and in mathematical modeling to be the optimal stimulation construct, as it achieved the best DC activation at the lowest stimulation energy [10,21]. The choice of number of leads and their placement in the spinal cord depend upon the physician s technique and his/her goals for each patient, as well as the patient s pain pattern. The Power Source As multiple stimulating contacts on narrowly spaced leads have been shown to improve clinical outcomes, and narrowly spaced contact combinations require higher stimulation energies, it becomes apparent that significant stimulation power will be required to drive such systems. Additionally, the growing use of dual-lead implants also requires an increase in stimulation energy. These new technologies and techniques entail special requirements for the stimulation power source, both from a battery standpoint and from the method for generating the stimulating field, i.e., the manner of delivering the stimulation pulse. Neurostimulators presently are powered using one of three different power sources: radiofrequency coupling (RF), primary cell, and rechargeable cell. RF stimulators were historically the first to be used for SCS, in the 1970s and 1980s. The first primary cell devices, employing lithiumthionyl-chloride batteries derived from pacemaker technology, appeared in 1980 [27]. These two choices were the only ones available for approximately two decades, and the clinical tradeoffs between them were relatively clear: primary cell implantable pulse generators (IPG) are more convenient for the patient, as the entire system is fully implantable and could be used during most activities of daily living. Primary cell IPGs, however, have limited life (average of 4 years, although sometimes less than 1 year), and are relatively large compared with similar bradycardia pacemaker devices [29]. The limited battery life also implies that the stimulation power may be compromised [1]. In contrast, RF units do not suffer from stimulation power constraints, as their power source is external, and has no specified end of life. RF units can drive multiple stimulation channels, up to 16 contacts, and be programmed with relatively high stimulation amplitudes, pulse width, and rate values [30]. However, RF units are very inconvenient for patients, as they require an S31 antenna to be worn over the implant during use of the device. This means that patients cannot sleep, shower, or swim with the equipment and, if the patient has sensitive skin, the taping of the antenna might result in chronic irritation [6]. In the last two decades, a revolution in portable consumer electronics took place, largely enabled by the development and use of lithium-ion batteries. Li-ion technology has many features, such as high energy density in a small size, memory-less charge retention behavior, and safer design [31], which make it possible to build a rechargeable implantable-grade medical device. In 2004, the first rechargeable cell SCS IPG became commercially available, which eliminated many of the tradeoffs between primary cell IPGs and RF devices. With the introduction of Li-ion rechargeable batteries in fully implantable neurostimulators, the stimulation power available in an IPG is no longer limited by concerns about short device life. Presently available rechargeable IPGs have lifetimes up to 9 years, and device size is now generally correlated to recharge interval rather than time-to-replacement. Recharge interval is dependent upon several parameters: stimulation energy, depth of discharge, useable battery capacity, state of battery degradation, depth of discharge, etc. In the simplest sense, the more stimulation energy a patient uses (e.g., high-amplitude settings and/or round-the-clock SCS usage), the more frequently a patient will recharge. Similarly, a larger battery capacity will extend the time between recharges, but will also require more time spent recharging by the patient to fill a depleted battery. Additionally, battery degradation over time and the depth of battery discharge must be understood. Li-ion cells have a fixed number of charge discharge cycles and the battery degrades slightly with each cycle. Over time, as the battery is cycled, the battery holds less charge. Also, in certain Li-ion technologies, if the battery is allowed to reach full discharge, chemical changes to the battery can take place that further degrade the battery s capacity. From the patient s point of view, device implant depth and the interface with the IPGs recharging equipment can be important [32]. In general, rechargeable neurostimulators require new awareness and responsibilities, but have also enabled the implementation of newer stimulation technologies [33]. One new feature made possible by rechargeable cell IPGs is the use of a current-controlled stimulation architecture. When controlling the stimu-

S32 lating current, there are several methods. The standard current-controlled system uses a single source for all contacts. In contact combinations where there is a single anode and cathode, the current to each contact is precisely controlled. However, if there is more than one anode or cathode in the contact combination, the current at each contact is no longer controlled; the impedance of each contact will determine where the current flows [34]. This impedance can change over time, as well as change differentially among the contacts [35]. This can affect the stimulation outcome, as a large impedance on a few contacts may so restrict the flow of current that the stimulation field is drastically altered, to the detriment of the therapy [36]. This challenge can be overcome in current-controlled systems that dedicate a current-source to each contact, such as that found in multiple independent constant-current (MICC) stimulation architecture. In such systems, the current at each contact is controlled, allowing for shaping of the stimulation field with minimum impedance sensitivity to directing current flow and designed to provide improved neural selectivity. To better understand the theoretical reasons for using MICC stimulation, a simple cascade of relationships between the nerve and the contact is instructive (see Figure 5). Since the nerve is stimulated when the transmembrane potential is driven to threshold [37], then controlling the A nerve fires when the transmembrane potential is driven to threshold The transmembrane potential is driven by the spatial gradient of the Electric Field at the nerve The Electric Field is proportional to the current density at the nerve Cathode Anode + Bradley transmembrane potential of the targeted nerves is a key factor to success. When stimulating extracellularly, the transmembrane potential is driven by the second spatial derivative of the electric field along the nerve [38]. This electric field is directly proportional to the current density, which in turn is determined by the current that flows along the nerve [39]. Thus, direct control of the current along the nerve is key to controlling the stimulation of the nerve [40]. If a MICC system is used, then the current flowing at the targeted nerve, often a few millimeters from the stimulating contacts, should be precisely controlled when using many contacts and should achieve the desired neural selectivity. Also, with rechargeable-cell systems, the capability of much larger stimulation pulse parameter outputs are possible, such as rate and pulse width. Studies have suggested that very high rates of stimulation in SCS may be useful for recapture of relief in the presence of breakthrough pain [30]. Additionally, relatively long pulse width values can be programmed, and these may result in more recruitment of smaller DC fibers [41], and thus achieve wider regions of paresthesia coverage. Conclusion Advances in lead technology, such as the increase in the number of contacts, with a concomitant decrease in the distance between them, have + Depolarized + Hyperpolarized + All geometry being equal, the current density is a direct function of the CURRENT at the nerve Therefore, CURRENT at the nerve is a key determinant of neurostimulation Figure 5 Cascade of relationships between nerve stimulation and electric field generation.

The Technology enabled better dorsal DC selectivity. Also, dual percutaneous leads of multiple contacts and paddle leads with multiple columns of contacts are now commonly used to create complex stimulation fields to achieve better neural targeting. With these newer lead technologies and techniques, the power source for neurostimulators has also advanced to use rechargeable lithium-ion cells. Use of rechargeable technology not only results in extended stimulator lifetime, but also enables higher stimulation power, such that current-controlled stimulation can be employed to control stimulation fields and further improve neural selectivity. Editors Note: At the time of this publication, there is a lack of prospective controlled clinical trials demonstrating improved efficacy or clinical outcomes of one manufacturer system (lead or generator design) over another. References 1 Barolat G. Spinal cord stimulation for persistent pain management. In: Gildenberg PL, Tasker RR, Franklin PO, eds. Textbook of Stereotactic and Functional Neurosurgery. New York: McGraw- Hill; 1998:1519 38. 2 Linderoth B, Foreman RD. Physiology of spinal cord stimulation: Review and update. Neuromodulation 1999;2:150 64. 3 North RB, Ewend MG, Lawton MT, Piantadosi S. Spinal cord stimulation for chronic, intractable pain: Superiority of multi-channel devices. Pain 1991;44:119 30. 4 Hayes DL, Furman S. Cardiac pacing: How it started, where we are, where we are going. PACE 2004;27(5):693 704. 5 Barolat G. Epidural spinal cord stimulation: Anatomical and electrical properties of the intraspinal structures relevant to spinal cord stimulation and clinical correlations. Neuromodulation 1998;1(2):63 71. 6 Barolat G, North RB. Spinal cord stimulation equipment and implantation techniques. In: Burchiel K, ed. Surgical Management of Pain. New York: Thieme Medical Publishers, Inc.; 2002:535 48. 7 Oakley JC. Spinal cord stimulation: Patient selection, technique, and outcomes. Neurosurg Clin N Am 2003;14:365 80. 8 Barolat G. Experience with 509 plate electrodes implanted epidurally from C1 to L1. Stereotact Funct Neurosurg 1993;61:60 79. 9 Holsheimer J, Struijk JJ, Tas NR. Effects of electrode geometry and combination on nerve fibre selectivity in spinal cord stimulation. Med Biol Eng Comput 1995;33:676 82. S33 10 Holsheimer J, Wesselink WA. Effect of anode cathode configuration on paresthesia coverage in spinal cord stimulation. Neurosurgery 1997;41:654 60. 11 Holsheimer J, Wesselink WA. Optimum electrode geometry for spinal cord stimulation: The narrow bipole and tripole. Med Biol Eng Comput 1997;35:493 7. 12 Manola L, Holsheimer J, Veltink PH. Technical performance of percutaneous leads for spinal cord stimulation: A modeling study. Neuromodulation 2005;8(2):88 99. 13 Law JD. Spinal stimulation: Statistical superiority of monophasic stimulation of narrowly separated, longitudinal bipoles having rostral cathodes. Appl Neurophysiol 1983;46:129 37. 14 Law JD. Targeting a spinal stimulator to treat the failed back surgery syndrome. Appl Neurophysiol 1987;50:437 8. 15 Alo KM, Holsheimer J. New trends in neuromodulation for the management of neuropathic pain. Neurosurgery 2002;50(4):690 704. 16 Dimitrijevic MR, Faganel J, Sharkey PC, Sherwood AM. Study of sensation and muscle twitch responses to spinal cord stimulation. Int Rehabil Med 1980;2:76 81. 17 Hunter JP, Ashby P. Segmental effects of epidural spinal cord stimulation in humans. J Physiol 1994;474(3):407 19. 18 Oakley JC, Prager JP. Spinal cord stimulation mechanisms of action. Spine 2002;27(22):2574 83. 19 Struijk JJ, Holsheimer J, van Veen BK, Boom HBK. Epidural spinal cord stimulation: Calculation of field potentials with special reference to dorsal column nerve fibers. IEEE Trans Biomed Eng 1991;38(1):104 10. 20 Lazorthes Y, Verdie J-C. Technical evolution and long-term results of chronic spinal cord stimulation. In: Lazorthes Y, Upton ARM, eds. Neurostimulation: An Overview. Mt Kisco, New York: Futura Publishing Co., Inc.; 1985:67 86. 21 North RB, Kidd DH, Olin JC, Sieracki JM. Spinal cord stimulation electrode design: Prospective, randomized controlled trial comparing percutaneous and laminectomy electrodes Part I. Technical outcomes. Neurosurgery 2002;51:381 90. 22 North RB, Lanning A, Hessels R, Cutchis PN. Spinal cord stimulation with percutaneous and plate electrodes: Side effects and quantitative comparisons. Neurosurg Focus 1990;2:1 5. 23 Manola L, Holsheimer J. Technical performance of percutaneous and laminectomy leads analyzed by modeling. Neuromodulation 2004;7(4):231 41. 24 Holsheimer J. Effectiveness of spinal cord stimulation in the management of chronic pain: Analysis of technical drawbacks and solutions. Neurosurgery 1997;40(5):990 9. 25 Oakley JC, et al. Transverse tripolar spinal cord stimulation: Results of an international multicenter study. Neuromodulation 2006 (in press).

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