10
Principles of Cord Activation During Spinal Cord Stimulation Gabriela Molnar, MS*; Giancarlo Barolat, MD Objectives: This study aims to review some of the basic principles of extracellular electrical stimulation used in spinal cord stimulation therapy for intractable pain. Materials and Methods: Spinal cord stimulation has been used therapeutically for more than 40 years. We present the basic principles of extracellular stimulation on which the therapy is based, describe electrode operation and current configurations, and explain the implications of these technological advances for the clinical application of spinal cord stimulation. Results: Computational studies of coupled electric field and neuron models have significantly advanced our understanding of the therapeutic effects of neurostimulation. Neurostimulation is intended to obtain maximal selectivity of desired neural elements while avoiding those resulting in side-effects. Preferential stimulation of the dorsal columns is achieved with a narrow spacing between electrodes using bipolar or tripolar electrode configurations. Stimulus parameters such as amplitude and pulse width may be used to selectively control which neuronal elements are excited during stimulation. Conclusions: A better understanding of the interaction between electric fields and the targeted neural elements may guide the selection of stimulation parameters in contemporary neurostimulators and lead to continuing advances in engineering solutions for therapeutic spinal cord stimulation. Keywords: Anode, cathode, electrode polarity, neural modeling, pulse width, spinal cord stimulation, stimulus amplitude Conflict of Interest: This study was supported by Medtronic. Giancarlo Barolat is a consultant for Medtronic and Boston Scientific, Inc. Dr. Barolat also is a consultant and equity holder for QiG Group. Gabriela Molnar is an employee at Medtronic. The planning, conduct, and conclusions of this report are those of the authors and not the company. INTRODUCTION Spinal cord stimulation (SCS) has been used in the treatment of chronic, intractable pain for more than 40 years and involves the delivery of electrical current to spinal elements including dorsal column (DC) and dorsal root (DR) fibers via electrodes placed in the dorsal epidural space. Connected to a battery-powered pulse gen- erator, those electrodes can be programmed as either cathodes (negative potential relative to a reference) or anodes (positive potential relative to a reference). When at least one cathode and one anode are programmed, they form a closed electrical circuit and generate an electric field, within the biologic tissue media, which stimulates neural elements including DC and DR fibers. Stimulation of the DC and DR fibers creates a tingling sensation called paresthe- sia, which is detected by the patient. Overlap of paresthesia with the patient’s painful area(s) has been considered necessary for clinical success (1). Paresthesia-pain overlap may be optimized by careful selection of stimulation parameters to generate an electric field that will provide maximal selectivity and stimulation of desired neural elements while avoiding stimulation of those neural elements that result in uncomfortable or painful sensations. While the exact mechanisms of SCS are not well understood, the principles under- lying stimulation of excitable tissue have been previously described (2–5). This article reviews some of those basic principles of extracel- lular electrical stimulation and discusses the selection of stimulation parameters during SCS within this context, with specific focus on electrode polarity, amplitude, and pulse width. ANATOMY RELEVANT TO SCS Sensory information from light touch and vibration stimuli is sig- naled through primary afferent axons that form DRs as they enter the spinal cord and ascend toward the brain via the DCs. These DR fibers follow a curved trajectory as they enter the spinal canal toward the spinal cord, passing through the cerebrospinal fluid (CSF) before entering the cord. In contrast to DR fibers, DC fibers are typically aligned with the long axis of the cord. The DCs are orga- nized topographically with axons from more rostral dermatomes entering the cord more laterally. In SCS, the electrodes and the targeted spinal neural elements are located at a distance from each other; this distance is primarily determined by the thickness of the dorsal CSF (dCSF). This distance varies by vertebral level and with different subject positions (e.g., lying supine vs. prone) (6,7). Both DC and DR fibers are different elements of the same type of large myelinated A-beta fiber. SCS can electrically excite these large Address correspondence to: Gabriela Molnar, MS, Medtronic Neuromodulation, 7000 Central Ave NE, MS RCE 385, Minneapolis, MN 55432, USA. Email: [email protected] * Medtronic Neuromodulation, Minneapolis, MN, USA; and Barolat Neuroscience, Denver, CO, USA For more information on author guidelines, an explanation of our peer review process, and conflict of interest informed consent policies, please go to http:// www.wiley.com/bw/submit.asp?ref=1094-7159&site=1 Neuromodulation: Technology at the Neural Interface Received: December 5, 2012 Revised: March 13, 2013 Accepted: March 13, 2013 (onlinelibrary.wiley.com) DOI: 10.1111/ner.12171 12 www.neuromodulationjournal.com Neuromodulation 2014; 17: 12–21 © 2014 International Neuromodulation Society

Principles of Cord Activation During Spinal Cord Stimulation

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Page 1: Principles of Cord Activation During Spinal Cord Stimulation

Principles of Cord Activation During SpinalCord StimulationGabriela Molnar, MS*; Giancarlo Barolat, MD†

Objectives: This study aims to review some of the basic principles of extracellular electrical stimulation used in spinal cordstimulation therapy for intractable pain.

Materials and Methods: Spinal cord stimulation has been used therapeutically for more than 40 years. We present the basicprinciples of extracellular stimulation on which the therapy is based, describe electrode operation and current configurations, andexplain the implications of these technological advances for the clinical application of spinal cord stimulation.

Results: Computational studies of coupled electric field and neuron models have significantly advanced our understanding of thetherapeutic effects of neurostimulation. Neurostimulation is intended to obtain maximal selectivity of desired neural elementswhile avoiding those resulting in side-effects. Preferential stimulation of the dorsal columns is achieved with a narrow spacingbetween electrodes using bipolar or tripolar electrode configurations. Stimulus parameters such as amplitude and pulse widthmay be used to selectively control which neuronal elements are excited during stimulation.

Conclusions: A better understanding of the interaction between electric fields and the targeted neural elements may guide theselection of stimulation parameters in contemporary neurostimulators and lead to continuing advances in engineering solutionsfor therapeutic spinal cord stimulation.

Keywords: Anode, cathode, electrode polarity, neural modeling, pulse width, spinal cord stimulation, stimulus amplitude

Conflict of Interest: This study was supported by Medtronic. Giancarlo Barolat is a consultant for Medtronic and Boston Scientific,Inc. Dr. Barolat also is a consultant and equity holder for QiG Group. Gabriela Molnar is an employee at Medtronic. The planning,conduct, and conclusions of this report are those of the authors and not the company.

INTRODUCTION

Spinal cord stimulation (SCS) has been used in the treatment ofchronic, intractable pain for more than 40 years and involves thedelivery of electrical current to spinal elements including dorsalcolumn (DC) and dorsal root (DR) fibers via electrodes placed in thedorsal epidural space. Connected to a battery-powered pulse gen-erator, those electrodes can be programmed as either cathodes(negative potential relative to a reference) or anodes (positivepotential relative to a reference). When at least one cathode and oneanode are programmed, they form a closed electrical circuit andgenerate an electric field, within the biologic tissue media, whichstimulates neural elements including DC and DR fibers. Stimulationof the DC and DR fibers creates a tingling sensation called paresthe-sia, which is detected by the patient. Overlap of paresthesia with thepatient’s painful area(s) has been considered necessary for clinicalsuccess (1). Paresthesia-pain overlap may be optimized by carefulselection of stimulation parameters to generate an electric field thatwill provide maximal selectivity and stimulation of desired neuralelements while avoiding stimulation of those neural elements thatresult in uncomfortable or painful sensations. While the exactmechanisms of SCS are not well understood, the principles under-lying stimulation of excitable tissue have been previously described(2–5). This article reviews some of those basic principles of extracel-lular electrical stimulation and discusses the selection of stimulationparameters during SCS within this context, with specific focus onelectrode polarity, amplitude, and pulse width.

ANATOMY RELEVANT TO SCS

Sensory information from light touch and vibration stimuli is sig-naled through primary afferent axons that form DRs as they enterthe spinal cord and ascend toward the brain via the DCs. These DRfibers follow a curved trajectory as they enter the spinal canaltoward the spinal cord, passing through the cerebrospinal fluid(CSF) before entering the cord. In contrast to DR fibers, DC fibers aretypically aligned with the long axis of the cord. The DCs are orga-nized topographically with axons from more rostral dermatomesentering the cord more laterally. In SCS, the electrodes and thetargeted spinal neural elements are located at a distance from eachother; this distance is primarily determined by the thickness of thedorsal CSF (dCSF). This distance varies by vertebral level and withdifferent subject positions (e.g., lying supine vs. prone) (6,7).

Both DC and DR fibers are different elements of the same type oflarge myelinated A-beta fiber. SCS can electrically excite these large

Address correspondence to: Gabriela Molnar, MS, Medtronic Neuromodulation,7000 Central Ave NE, MS RCE 385, Minneapolis, MN 55432, USA. Email:[email protected]

* Medtronic Neuromodulation, Minneapolis, MN, USA; and† Barolat Neuroscience, Denver, CO, USA

For more information on author guidelines, an explanation of our peer reviewprocess, and conflict of interest informed consent policies, please go to http://www.wiley.com/bw/submit.asp?ref=1094-7159&site=1

Neuromodulation: Technology at the Neural Interface

Received: December 5, 2012 Revised: March 13, 2013 Accepted: March 13, 2013

(onlinelibrary.wiley.com) DOI: 10.1111/ner.12171

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cutaneous afferent axons at different locations on their path towardthe brain. Neurostimulation therapy seeks to obtain maximal selec-tivity of desired neural elements (those that lead to clinical benefit)while avoiding those resulting in side-effects (8,9). The desiredneural elements are thought to be part of this large myelinatedafferent fiber system. Stimulation of the large myelinated afferentfibers at the intraspinal level can occur in four different areas: the DR,the DR entry-zone, the dorsal horn, or the DC. Electrical stimulationof this complex elicits paresthesia. Table 1 describes the clinical cor-relates of stimulating intraspinal structures. It is important to differ-entiate activation of the segmentary large myelinated afferents (DR/entry-zone/dorsal horn) from activation of the ascending long tractsin the DC. Activation of the motor fibers (ventral nerve roots, motorneurons, or corticospinal tract) is invariably perceived as intrusiveand is to be avoided. High-amplitude stimulation can evoke discom-fort or pain.

Stimulation of DC fibers generates paresthesia of several derma-tomes caudal to the level of the stimulating cathode. When thetopography of pain spans several dermatomes, the most effectivestimulation target is the DC. Stimulation of the DR should be avoidedto achieve the broadest paresthesia of the patient’s painful area(s). Inaddition to transmission of sensory information, DR fibers containmotor reflex and pain pathways. Stimulation of DR results in limitedor segmental paresthesia typically in one or two dermatomes. Nerveroots generally have a lower activation threshold than the DC due toseveral factors including their curved trajectory, the inhomogeneityand anisotropy of the conductive media they traverse, and theirorientation relative to traditional placement of the stimulation leads,which are typically aligned with the long axis of the cord (10). Further,as dCSF increases, the thresholds for DC fibers increase faster thanthe thresholds for DR fibers, leading to preferential recruitment of DRfibers when dCSF is large (9). Simultaneous activation of both DC andDR fibers will ultimately preclude successful stimulation of the DC,because stimulation of the DC will only be achieved at thresholdsthat cause motor activity, uncomfortable or painful sensations.

Selective stimulation of DC fibers rather than of DR fibers may beachieved by positioning leads to precisely control the electric fieldand selecting stimulation parameters that will be described in thefollowing sections. When the targeted area of paresthesia is the der-matome of a specific nerve root, an electrode placed on that nerveroot might be the only way to obtain concordant paresthesia. Forinstance, when addressing thoracic wall pain, stimulation of the DCwill not produce appropriate paresthesia. Stimulation of the affectednerve root(s), with laterally placed electrodes, is the only option.

PRINCIPLES OF EXTRACELLULAR STIMULATION

Action potentials are typically generated by synaptic currentsthat produce depolarization of the postsynaptic neuron and result

in an action potential conducted along the axon toward its intendedtarget to produce a desired effect, such as muscular contraction.With electrical stimulation, neurons may be artificially depolarizedby placing an electrode within the neuron to inject current (intra-cellular stimulation), or by delivering current outside the neuronthat creates a distribution of potentials within the excitable tissue(extracellular stimulation). Current is the flow of charge carried byelectrons in the wires and electrodes of a stimulatory system or byions in biologic media. Charge is transferred across a cell membranevia passive membrane properties (capacitance and resistance) andalso via active ion channels. The passage of current through excit-able tissue during extracellular stimulation, such as with SCS, gen-erates an electric field that changes the transmembrane potential ofa neuron, which is the potential difference across the membrane.That change may cause depolarization of the axon membrane, inwhich the transmembrane potential is increased and the neuron ismore likely to generate an action potential, or may cause hyperpo-larization, in which the transmembrane potential is decreasedmaking the neuron less likely to fire. Unlike physiologically gener-ated action potentials, artificial depolarization of an axon results inpropagation of the action potential in both directions (proximal anddistal). Thus, the primary effect of neurostimulation is to depolarizeor hyperpolarize the cell membranes.

In the simplest scenario of monopolar stimulation with a pointsource electrode, the extracellular potentials (Ve) generated withinthe tissue can be described by the equation:

VI

re =

4πσ(1)

where I is the current injected by the electrode into the surroundingtissue (amperes), σ is the electrical conductivity of the homoge-neous and isotropic medium (siemens per meter or S/m), and r is thedistance between the electrode and a point within the medium (4).

Formulae that describe the extracellular potentials producedfrom stimulation with nonmonopolar sources, such as dipoles, havealso been derived (4). The electric field along an axon may bedescribed by taking the first spatial derivative of the potential (Ve)along the direction of the axon, in volts per meter (V/m); however, itis the first spatial derivative of the electric field (or the second spatialderivative of the potentials) in the direction along the axon that isresponsible for the effects of extracellular stimulation.

Computational studies using models of coupled electric fieldsand neurons have helped to advance our knowledge of the effectsof neurostimulation (11). In fact, several theoretical predictions havebeen confirmed with clinical studies. This modeling is typically per-formed by a two-step process (12,13). The first step involves solvingfor the electric field within the predefined volume conductor usingmethods such as finite element analysis. This volume conductormay be simple and consists of a single compartment, such as one

Table 1. Clinical Correlates of Electrical Stimulation of Intraspinal Structures.

Anatomical structure(s) Clinical correlates Location in relation to the cathode

Dorsal root/dorsal root entry zone/dorsal horn Paresthesia Ipsilateral and at the same spine segmentDorsal columns Paresthesia Ipsilateral and caudalVentral (anterior) roots/ventral motor neurons Motor contractions Ipsilateral and at the same spine segmentCorticospinal tract Motor contractions Ipsilateral and caudalSpinothalamic tract* Sensation of warmth Contralateral and caudalAutonomic fibers (inhibition of sympathetic fibers) Sensation of warmth and vasodilation Ipsilateral and caudal

*Cannot be obtained with dorsal epidural stimulation; has only been observed during the stimulation phase of percutaneous cordotomies.

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type of tissue (also referred to as a homogeneous medium), or maybe more complex and consists of two or more compartments, eachpossessing different electrical properties (an inhomogeneousmedium). The spinal cord is typically modeled using an inhomoge-neous volume conductor to incorporate the various anatomic com-partments, such as gray matter, white matter, CSF, dura mater, andepidural fat (13). The electrical properties of each compartment maybe the same in every direction (isotropic), or may be different in oneor more directions (anisotropic). For example, white matter ismodeled as anisotropic because the electrical conductivity is higherin the direction along the nerve fibers and lower in the directiontransverse to the fibers. The electrical properties of the tissue sur-rounding the activated electrodes will impact the current pathswithin the volume conductor and will have a direct effect on thethree-dimensional spatial distribution of the electric field, as repre-sented by its potential distribution or by its current densitydistribution.

The second step of the modeling process involves the calculationof the nodal positions of a myelinated axon and the interpolation ofthe extracellular potentials at those locations to observe the effectsof the electric field on neuronal activation (10). Calculation of theneuronal response may involve the use of detailed, nonlinear,multicompartment models with parameters derived from mamma-lian myelinated axons (10,14). Both passive (resistive and capacitive)as well as active (ion channel) properties of the neuron are includedin the simulation. Computational models developed by Coburn, andlater more extensively by the Holsheimer group, have contributedto the body of theoretical knowledge surrounding SCS (11).

The electrical behavior of a myelinated nerve fiber can be repre-sented using a cable network that is described by components,including a nodal membrane resistance (Rm) in parallel with a nodalmembrane capacitance (Cm), and an intra-axonal membrane resis-tance (Ra) (Fig. 1) (12). As previously noted, the nodal membraneresistance may include both passive and active elements. ApplyingKirchhoff’s Current Law, at node n of a myelinated axon, results in anexpression describing the change in nodal transmembrane poten-tial of the fiber (Vm) to the applied extracellular potentials (Ve). Thisexpression is known as the cable equation.

CdV n

dt

V n

R

V n V n V n

R

V n

mm m

m

m m m

a

e

( ) + ( ) −−( ) + +( ) − ∗ ( )⎛

⎝⎜⎞⎠⎟

=−( )

1 1 2

1 ++ +( ) − ∗ ( )V n V n

Re e

a

1 2(2)

The second spatial derivative of the potentials along an axondrives membrane polarization (i.e., right side of Equation 2) and hasbeen termed the activating function (AF) (15).

AF nV n V n V n

Re e e

a

( ) =−( ) + +( ) − ∗ ( )1 1 2

(2A)

A positive value for the AF results in membrane depolarization atnode n, while a negative value for the AF results in membranehyperpolarization at node n. There are several limitations to the useof the AF because the term represents the input to the system (theextracellular potentials) and not the output (the change in trans-membrane potential). In addition, the time-dependent, nonlinearproperties of the axon, such as the effects of pulse width, are notconsidered; and, thus the AF would be indicative of membranepolarization at short pulse widths (16). In spite of those limitations,the AF is another tool, in addition to coupled field-neuron modeling,that has helped to elucidate some of the properties of extracellularstimulation of nerve fibers.

From the right side of Equation 2, we can also predict two recruit-ment properties of stimulation. First, neurons farther away from theelectrode will require higher amplitudes for excitation than neuronsclose to the electrode. From Equation 1, we can observe that themagnitude of the potentials decreases with increasing distancebetween the stimulating electrode and the stimulation target; simi-larly, the AF decreases with increasing distance between the stimu-lating electrode and the stimulation target. The current required fora neuron to reach threshold (Ith) is proportional to the square of thedistance of the neuron from the electrode (r) and can be describedby the current–distance relationship:

I I k rth = + ∗02 (3)

where I0 is the offset and k is the slope (3,17). The value of k for theactivation of central nervous system neurons can range from 100 to4000 μA/mm2 using 0.2-msec pulses (3).

Second, large-diameter fibers require lower currents for excita-tion than smaller diameter fibers (18). Large-diameter fibers have alarger spacing between nodes of Ranvier than small-diameterfibers; the length of the internode (the distance between successivenodes of Ranvier) is approximately 100 times the diameter of thefiber. Thus, large-diameter fibers experience increased polarizationcompared with smaller diameter fibers. Furthermore, large-diameter fibers have a lower intra-axonal membrane resistance thansmall-diameter fibers, which increases the denominator on the rightside of the cable equation (Equation 2) for large-diameter fibers.This concept is shown in Figure 2 using the AF calculated for twodifferent fiber diameters. Figure 2 illustrates that even though themaximum potential is the same, the changes induced in the trans-membrane potential as described by the AF are much larger in the10 μm-diameter fiber than in the 5 μm-diameter fiber, because thedistance between the internodes is much larger for the large-diameter fiber.

The remainder of this paper focuses on extracellular stimulationof myelinated nerve fibers because these elements are more excit-able than cell bodies. Initiation of action potentials typically occursin the axon, either at the initial segment or at a node of Ranvier, eventhough they may also be initiated in cell bodies, such as whenstimulating the motor cortex epidurally (19,20).

CATHODIC AND ANODIC STIMULATION

In addition to the electrical properties of the tissue, the spatialdistribution of the electric field is affected by the electrode polarity.

Figure 1. Electrical model of a myelinated fiber. Rm = membrane resistance,Cm = membrane capacitance, Ra = intra-axonal membrane resistance, Ve = extra-cellular potential; Vi = intracellular potential.

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Current flows from anode to cathode, depolarizes neuronal ele-ments near the cathode, and hyperpolarizes neuronal elementsnear the anode (2).

During cathodic stimulation, the negative charge of the electrodecauses charge near the electrode to redistribute, with the negativecharge collecting on the outside of the cell membrane, causing anoutward current and depolarizing the cell. Intracellular charges arealso redistributed as positive charges move from a distant portion ofthe axon to the region near the electrode. That results in an inwardcurrent and hyperpolarization at the distant portion of the axon.Stimulation under the anode results in an opposite effect on anaxon. Areas near the electrode are hyperpolarized, and areas furtherfrom the electrode are depolarized. This concept is illustrated inFigure 3 by using Equation 1 to generate the distribution of poten-tials resulting from cathodic stimulation in a uniform medium ofconductivity 0.23 S/m (similar to that of gray matter) and calculatingthe AF along a linear axon.

The two hyperpolarized regions of the axon flanking the depolar-ized area during cathodic stimulation are referred to as virtualanodes. The two depolarized regions of the axon flanking the hyper-polarized area during anodic stimulation are referred to as virtualcathodes. The distant portion of the axon that is hyperpolarizedduring cathodic stimulation or depolarized during anodic stimula-tion is smaller compared with the portion of the axon that is directlynear the electrode. Thus, while the total inward and outward cur-rents are the same, the inward current is distributed over morenodes and is of a lower magnitude than the outward current duringcathodic stimulation. Conversely during anodic stimulation, the

outward current is distributed over more nodes and is of a muchlower magnitude than the inward current. Because the outwardcurrent responsible for depolarization is of a smaller magnitudeduring anodic compared with cathodic stimulation, stimulationnear a cathode requires less current to activate the axon.

During anodic stimulation, there are two ways in which axonalexcitation can occur. First is direct depolarization at the virtual cath-odes. Activation of an axon is still possible with anodic stimulation;however, larger stimulatory currents are required, roughly three toseven times the currents required to excite a fiber using cathodicstimulation (17,21). Second is by anode break excitation that occursafter the end of a long hyperpolarizing stimulus pulse. That phe-nomenon is explained by understanding the effects of stimulationon the voltage-dependent sodium channels. Hyperpolarization ofthe membrane will remove the partial inactivation and cause theactivation gate of the sodium channels to close. At the end of thelong duration pulse, the inactivation gate will slowly close due to itsslow time constant, while the activation gate (fast time constant)will open and may lead to sodium ion influx and generation of anaction potential (22).

During cathodic stimulation, the virtual anodes may lead to thephenomenon of cathodic blocking in which the strength of thevirtual anode is large enough to prevent an action potential frompropagating along the axon (2). As with the excitation of a virtualcathode, higher currents are needed for cathodic block to occur.

The geometry of anode and cathode placement, or electrodeconfiguration, will affect the orientation of the electric fieldand therefore thresholds for nerve activation. We have previously

Figure 2. a. Potentials along a myelinated nerve fiber 5 μm (squares) and 10 μm (triangles) in diameter. b. Normalized AF along a myelinated nerve fiber 5 μm(squares) and 10 μm (triangles) in diameter. Stimulation using a point source electrode positioned 1 mm from the nerve. The peak of the activating function is largerfor the larger diameter fiber.

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examined the effects of monopolar stimulation, where the currentspreads around a cathode symmetrically and in a radial fashion fromthe electrode in a homogeneous, isotropic medium. An electrodeconfiguration is monopolar when the distance between the anodeand cathode is large, such that the current distribution resultingfrom at least one cathode and at least one anode do not interact.This may be the case when the implantable pulse generator or anelectrode located far away from the cathode is programmed as ananode. As the electrodes are moved closer together, the currentaligns more in the direction of the bipole (anode–cathode axis).Nerve fibers are more easily depolarized when aligned with thebipole; higher thresholds are required when axons are transverse tothe bipole (2).

AMPLITUDE, PULSE WIDTH, ANDTHEIR INTERACTION

The amplitude of a stimulus is the magnitude of the current orvoltage that is delivered at the electrode and is measured inmilliamps or volts, respectively. Increasing the stimulus amplitudeincreases the size of the applied electric field and increases thecharge delivered to the tissue. A larger volume of activation ofexcitable tissue may be achieved by increasing the amplitude ofstimulation.

The stimulus pulse width is the second factor in controlling thecharge delivered (Fig. 4). The pulse width is the duration of thestimulus and is measured in microseconds (μsec). Increasing pulse

width will activate a larger number of fibers and also activate lessexcitable elements such as smaller diameter fibers, which may leadto changes in perceived paresthesia. Theoretical studies have foundthat short pulse widths increase the difference of activation thresh-olds among fibers of different diameters (23,24). With lower pulsewidths, there is a greater difference in activation thresholdsbetween large and small fibers; in contrast, there would be a smallerdifference in thresholds between large and small fibers using largerpulse widths.

The amplitude and pulse width of the stimulus may be used toselectively control which neuronal elements are excited duringstimulation. These two parameters interact to determine thresholdsfor neuronal excitation; this interaction is described by thestrength–duration curve:

I IT

PWth rh

ch= +⎛⎝

⎞⎠1 (4)

where Ith is the threshold current for excitation, Irh is the rheobasecurrent, PW is the stimulus pulse width, and Tch is the chronaxie. Thecurve is plotted in Figure 4a. The rheobase current and the chro-naxie describe this curve. The rheobase current is the currentrequired for excitation at an infinitely long pulse width. The chro-naxie is the pulse width at twice the rheobase value. Neuronal ele-ments with lower chronaxie values are more excitable thanelements with higher chronaxie values. Typical values of chronaxiefor large myelinated fibers are 30–200 μsec, whereas values forcell bodies range from 1 to 10 msec (2). The strength–duration

Figure 3. Normalized activating function for cathodic (a) and anodic (b) stimulation. Stimulation using a point source electrode positioned 1 mm from the nerve.Note the hyperpolarizing (AF < 0) side lobes on the cathodic stimulation activating function, and the depolarizing (AF > 0) lobes on the anodic stimulation activatingfunction.

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relationship indicates that neurons will have lower thresholds forexcitation as the stimulus pulse width increases. The strength–duration curve will be dependent on several factors, including thedistance between the electrode and the target neurons, the polarityof the stimulus, the stimulus waveform, and the fiber diameter(18,25,26). The energy delivered is minimized when the pulse widthis approximately the chronaxie of the neuron.

The charge–duration curve can be derived from the strength–duration relationship (Equation 4) by multiplying both sides of theequation by the stimulus pulse width, where Qth is the charge atthreshold for activation:

Q I PW Tth rh ch= +( ) (5)

Figure 4b shows that the charge required for excitation increaseswith increasing pulse width. That is because accommodation occursduring a prolonged depolarizing pulse in which increased sodiuminactivation leads to lower excitability. Although shorter pulsewidths require higher amplitudes for excitation, shorter pulsewidths are generally recommended for safe stimulation in order tominimize charge injection and electrochemical reactions that occurat the electrode’s surface (27).

CLINICAL SELECTION OFSTIMULATION PARAMETERS

The principles governing neural excitation can be used to informthe selection of stimulation parameters during clinical use of SCS.The stimulation parameters of electrode polarity and geometry,amplitude, and pulse width will now be discussed in the context ofthis knowledge.

SCS Implications of Electrode Polarity and GeometryBecause nerve depolarization generally occurs near a cathode,

that electrode is typically placed at the spinal level to stimulate thecorresponding DC maximally. However, the probabilities of target-ing paresthesia to different body areas vary as a function of locationof the stimulation cathode (Fig. 5). The variation in probabilitiesobserved from Figure 5 is likely due to anatomic (e.g., cord geom-etry, DC topography, and dCSF) and lead placement differencesamong patients (28).

DC selectivity can be achieved with narrow spacing between elec-trodes using a bipolar (cathode with an adjacent anode) or guardedcathode (a cathode flanked both rostrally and caudally by an anode)configuration when the leads are oriented in the rostro-caudal direc-tion (9,29). These configurations result in an electric field that isaligned with the direction of the DC fibers rather than the DR fibersallowing more selective activation of DC. Selectivity of the DC overDR fibers is increased with narrow-spaced electrodes becausesmaller distances between the anode and cathode achieve widermediolateral recruitment of the DC. As such, Barolat et al. found thatnarrower spacing between bipolar contacts increased paresthesiacoverage (8). It has been shown via computational modeling that aguarded cathode configuration on a lead placed on the midlineprovides maximal recruitment of DC fibers (29) and that clinically,patients prefer that configuration (1). Such preferential recruitmentof DC fibers suggests that these types of electrode geometries on alead with narrow spacing between electrodes may be used to treatcomplex types of pain that require the targeting of stimulation tomultiple dermatomes. These types of leads can also allow increasedgranularity in cathode location along the lead if this finer resolutionis needed. Preferential stimulation of the DC using the narrow bipoleor guarded cathode typically comes at the expense of higher energyconsumption compared with wider spaced electrodes, with implica-tions of reduced battery longevity in primary cell devices or longerrecharge intervals in rechargeable devices (29).

The wider the electrodes are spaced on a lead, the more the fieldnear the cathode in a bipolar or guarded cathode configuration willresemble a monopolar stimulation field and the alignment of theDC fibers with the anode–cathode axis will be lost. Such electrodespacing will result in preferential recruitment of DR over DC fibers(30). Leads with a wide spacing between electrodes may be used totarget paresthesia to a larger number of areas due to the increasedspan of the electrode array, which consequently increases vertebralcoverage. Cathode selection may be done at multiple locations andmay be useful in patients with pain patterns that change with time(31). However, although wide spacing provides additional flexibilityfor selecting the location of the cathode, it comes at the expense ofreduced DC recruitment compared with the leads with narrowlyspaced electrodes. Thus, due to preferential recruitment of DR fiberswith widely spaced electrodes, these types of leads may be moreuseful for treating segmental pain.

The optimum spacing between electrodes to achieve the lowestthresholds for DC activation varies depending on the distance

Figure 4. a. Strength–duration relationship with chronaxie of 295 μsec and rheobase of 2.5 mA. b. Charge–duration relationship derived from strength–durationrelationship.

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between the electrodes and the spinal cord for a given electrodeconfiguration (9). For a given dCSF distance, the largest positive AFpeak is achieved when the largest positive peak from both theanodal and cathodal AF(s) are superimposed; this implies that thelowest DC thresholds will be achieved at this cathode–anode sepa-ration distance (5,9). Holsheimer estimated that the optimumcenter-to-center electrode spacing for a rostro-caudally orientedlead would be approximately 1.7 times dCSF, suggesting that for adCSF of 3.5 mm, the most energy-efficient spacing would beapproximately 6 mm (5).

A transverse tripole configuration (a cathode flanked laterally byanodes) can also be used to preferentially activated DC fibers andhas been investigated with computer models and in clinical settings(32–34). The lateral anodes hyperpolarize the DRs, resulting in anincrease in the amplitude required to activate them and thereforeincreasing the amplitude at which a patient would experience dis-comfort. This configuration achieves the best depth of activation ofDC fibers, but because the DC axons are aligned transversely to theanode–cathode axis, higher thresholds are required for activationcompared with the narrow bipole or guarded cathode that are ori-ented rostro-caudally and aligned with the DC fibers.

SCS Implications of AmplitudeDuring SCS, the amplitude range for a patient is generally titrated

between the perception threshold and the discomfort threshold.The perception threshold is the lowest amplitude at which pares-thesia is detected by the patient. As the amplitude is increased pastthe perception threshold, the patient will detect increased intensityand expansion of the areas of the paresthesia. Eventually, the ampli-tude will be high enough to reach the discomfort threshold, which

is the amplitude that results in intolerable stimulation such as motorside-effects, pain, or unpleasant sensations. The range between theperception threshold and the discomfort threshold is known as theusage range or therapeutic range. A typical ratio between the per-ception threshold and discomfort ranges between 1.4 and 1.7 butmay be as high as 2.8 when using a transverse tripole configurationor a narrow longitudinal bipole (33,35). A larger difference betweenthe perception threshold and the discomfort threshold is preferredin order to ensure flexibility for amplitude titration (8). Because mostpatients cannot tolerate amplitudes 40–70% higher than thethreshold for perception of paresthesia, it is unlikely that cathodicblock or anodal excitation will occur during SCS as those phenom-ena require approximately require amplitudes 3–7 times perceptionthreshold.

The amplitude of stimulation is typically delivered in a constantfashion. However, stimulation intensity depends on the patient’sposition. Specifically, the amplitude and/or energy for stimulationwhen the patient is supine is significantly lower than when stand-ing or sitting; the mean difference may range from 11% to 35%(36–39). One reason is that the spinal cord moves within the sub-arachnoid space with changes in posture, which varies the distancebetween the stimulating electrodes and the target neurons (7).Largely attributed to the effects of gravity, the spinal cord movescloser to the electrodes in a supine position and moves fartheraway from the electrodes in the prone position. The variations indCSF with the position of the patient result in variations in theamplitudes required for stimulation (6,28,32,33,40,41). Using Equa-tion 3, one can see that larger amplitudes are required to activatefibers as dCSF increases.

Computer models of SCS have been developed that include theanatomy and electrical properties (conductivity) of the modeled

Figure 5. Probability of generating paresthesia in a particular body area as a function of vertebral level of the stimulating cathode. Figure modified from HolsheimerJ, Barolat G. Neuromodulation. 1998;1 (3):129–136 (28).

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tissues, including the spinal cord gray and white matter, CSF, dura,epidural fat, and vertebral bone. Table 2 shows the electrical con-ductivity of the various tissues in the SCS model. The computermodels have shown that given a fixed location of stimulation elec-trodes, the intensity of the electric field at the level of the DCdecreases as dCSF increases when maintaining a constant ampli-tude of stimulation (42). An increase in dCSF results in an increase inthresholds for both DC and DR activation, and therefore fewerrecruitment of DC and DR fibers. Thus, constant amplitude stimula-tion generates variations in the activation of DC and DR fibers asposition (and dCSF) of the patient changes (Fig. 6a). That results inchanges in intensity and location of paresthesia, suggesting thatstimulus amplitude must vary as the distance between the elec-trodes and spinal cord varies. The reduction in amplitude in themodel between the standing and supine dCSF values is similar tothe trend in reported clinical data in which therapeutic amplitudeswhen supine are typically much lower than amplitudes when stand-ing or sitting (Fig. 6b) (36,38).

The computer model also predicts that changes in the electrodeimpedance do not vary with dCSF. The impedance for a midline leadwith a guarded cathode configuration changed less than 0.3% asdCSF was varied from 3.6 to 5.8 mm. Thus, cord movement within

the subarachnoid space has no effect on electrode impedance. Thatis because most of the impedance is determined by two factors: theelectrical conductivity and thickness of the encapsulation layer sur-rounding an electrode; and the electrical conductivity of the tissuemedium near the electrodes, including the dura and fat (6,43,44).This is illustrated in Figure 7, which shows the effects of individuallydoubling or halving the conductivity of the dura, fat, CSF, and spinalcord white matter on impedance while keeping all other conduc-tivities at the baseline values (Table 2). The impacts of electricalconductivity variations on impedance are greatest when theseoccur in tissues near the electrode; changes in tissues far away fromthe electrode have little impact on the impedance.

Lack of variation in the impedance with varying dCSF values alsois attributed to the highly conductive CSF. The conductivity of theCSF allows 80–90% of the stimulatory current to flow through theCSF; therefore, spinal cord movement within this highly conductivemedium has little impact on the impedance (6,42,45). The nonsig-nificant difference in impedance, as a function of patient position,has been confirmed in clinical studies (38,39). One implication ofthat finding is that constant amplitude stimulation systems,whether voltage- or current-controlled, are not able to automati-cally adjust for amplitude changes that occur due to changes inposition of the patient (consequent to changes in dCSF). Devicescapable of automatic adaptation of stimulus amplitudes concomi-tant with changes in position of the patient could potentiallyprovide more consistent and optimal paresthesia, and may reducedependence of the patients on their programmers (Fig. 6b). Onealgorithm that adapts the stimulation amplitude to changes inposition of the patient is programmable in the clinic in real time(39).

SCS Implications of Pulse WidthUsing a detailed computer model of SCS, with a realistic distribu-

tion of fiber sizes and density, increased stimulus pulse width

Table 2. Electrical Conductivity of Tissues Used in a Spinal Cord Stimu-lation Computer Model.

Structure Conductivity (S/m)

Vertebral bone 0.02Epidural fat 0.04Dura mater 0.03Cerebrospinal fluid (CSF) 1.7White matter 0.6 longitudinal

0.083 transverseGray matter 0.23

Figure 6. Neural activation patterns generated by varying dCSF (3.6, 4.7, and 5.8 mm) with a guarded cathode configuration and an amplitude at 60% of the usagerange using (a) constant amplitude stimulation and (b) varying amplitude stimulation. The figures to the left of the neural activation patterns quantify the dorsalcolumn recruitment area (DC area in mm2) and stimulus amplitude as a function of dCSF.

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increased recruitment of DC fibers (46). In addition, higher pulsewidths increased recruitment of medially located, smaller diameterDC fibers. Those findings were consistent with results from clinicalstudies with patients who had implanted SCS systems (47,48)—aspulse width increased, the area of stimulation-induced paresthesiaincreased, and the area of paresthesia extended toward morecaudal dermatomes. Activation of the smaller, more medial fiberswithin the DC likely corresponded to those fibers originating frommore caudal dermatomes.

The programmed pulse width varies from patient to patient butoften is within the range of 250–500 μsec (31,47,49). Studies haveindicated that the chronaxie during SCS is approximately 200–250 μsec (47,50), which corresponds with the activation of largemyelinated fibers (2). The difference between the programmedpulse width and the chronaxie suggests that patients implantedwith SCS systems may require higher pulse widths than the chro-naxie in order to obtain optimal paresthesia-pain overlap (47). Theprogrammed pulse width should be kept low to minimize thecharge injected but high enough to achieve maximal clinical benefitfor the patient.

CONCLUSION

We have reviewed the basic principles of extracellular stimulationand their importance in understanding of the effects during SCS.The electric field is influenced by the electrical properties of thetissue, the electrode placement and polarity, and the stimulationparameters. The spatial extent of activation and the types of acti-vated neural elements can be controlled by careful selection ofstimulation parameters. Better understanding of the interactionbetween the electric fields and the targeted neural elements mayguide the selection of stimulation parameters and may lead toadvances in engineering solutions for SCS.

Authorship Statements

Gabriela Molnar and Dr. Barolat coauthored the manuscript, hadcomplete access to the data needed to create this manuscript, andapproved the final manuscript.

How to Cite this Article:Molnar G., Barolat G. 2014. Principles of Cord ActivationDuring Spinal Cord Stimulation.Neuromodulation 2014; 17 (Suppl. 1): 12–21

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