Leads and Electrodes for Cardiac Implantable Electronic Devices

Leads connect the pulse generator of a cardiac implantable electronic device (CIED) to sites within the body to form an electric circuit ( Fig. 11-1 ). Leads contain five categories of basic components (electrodes, conductors, insulation materials, fixation mechanisms, and connector pieces), each with its own functional requirements and design features ( Table 11-1 ). The electrodes interact with biological tissues through body fluids and should have low polarization properties and high corrosion resistance and induce minimal fibrosis. The conductors should have low resistance against direct electric currents, high reactance against high-frequency (radiofrequency range) alternating electric currents, high tensile strength (useful during lead extraction), high resistance against metal fatigue and electrochemical corrosion, and robust joints with other lead components. The insulation materials should be resistant to mechanical and electrochemical degradation (abrasion, fatigue fracture, creep, environmental stress cracking, metal ion-induced oxidation, and hydrolysis), induce little or no thrombosis and fibrosis, and repel bacterial colonization. The fixation mechanisms should be safe and easy to deploy at implantation, secure and stable once successfully deployed, and safe and easy to disengage from biological tissues even after long-term deployment. The connector pieces are transition zones in the lead structure and need to be resistant to mechanical stresses from body movement (high frequency, low amplitude) and during CIED revisions (low frequency, high amplitude) while providing an easy, reliable, and electrically tight seal with the pulse generator for both low-voltage pacing and high-voltage defibrillation (DF) pulses. The theoretical principles behind the functional requirements and design features of lead components, as well as common failure mechanisms and the clinical follow-up needed to maximize patient safety, are discussed.

TABLE 11-1

Functional Requirements and Resulting Design Features of Lead Components

Components	Functional Requirements	Parameter Trends	Lead Design Features and Other Factors
Electrode	↓ Stimulation threshold ↑ Sensing fidelity	↑ C _dl	IrO ₂ /TiN fractal coating; ↑ geometric surface area
	↓ Stimulation threshold ↑ Sensing fidelity	↓ d _e-m	Steroid elution; ↑ geometric surface area
	↓ Interface potential ↓ Electrolysis/electrode corrosion	↑ C _dl	IrO ₂ /TiN fractal coating; charge-balanced biphasic impulse (pulse generator)
	↓ Power consumption		↓ Geometric surface area
Conductor strand	↓ Fatigue fracture/corrosion		MP35N or MP35N-LT alloy jacket
Conductor strand	↑ Electric conductivity		Silver core
Conductor coil	Magnetic resonance imaging conditionality	↑ Inductance (may ↑ resistance)	↓ Coil wire diameter; ↓ coil pitch angle; ↓ number of filars; ↑ coil diameter
	Magnetic resonance imaging conditionality	↑ Inductance (may ↑ resistance)	Inductor filter before electrode
	↑ Fixation helix deployment	↑ Torque transmission	↑ Coil wire diameter; ↑ coil pitch angle; ETFE jacket
Conductor cable	↓ Movement relative to lead body		Better lead body geometry; better implantation techniques
Insulation	↓ Mechanical degradation: abrasion; fatigue fracture; creep		Better materials; better implantation techniques
	↓ Chemical degradation: environmental stress cracking; metal ion-induced oxidation; hydrolysis		Better materials; better lead component arrangements; better implantation techniques
	Biocompatibility: ↓ fibrogenicity/thrombogenicity/bacterial colonization		Better materials; better implantation techniques
Fixation mechanism	Ease of deployment; reliability of fixation; ease of extraction; ↓ risk of cardiac perforation		Better design; better implantation techniques
Connector piece	↓ Stress concentration points; ↑ ease of pin insertion; ↓ bulk		Spiraling of cables; ↓ number of splice joints; stiffer pins (may produce abrupt transition in stiffness); DF4

C _dl , Double-layer capacitance; d _e-m , electrode-myocardium distance; DF4, quadripolar defibrillation; ETFE, ethylene tetrafluoroethylene; IrO ₂ , iridium (di)oxide; TiN, titanium nitride.

Lead Polarity

All pacing stimuli and DF pulses are capacitively coupled and require a negatively charged electrode (cathode) and a positively charged electrode (anode) and are therefore bipolar. However, whereas the cathode is typically an electrode in contact with the endocardium (for a transvenous lead) or the epimyocardium (for an epicardial lead), the anode may be located more proximally along the same lead or on another lead or may be the pulse generator case. If both the cathode and the anode are located on the same lead, the term bipolar is used to refer to the lead design. In contrast, if the lead has only one tip electrode (the cathode) and the pulse generator or another lead is used as the anode, the term unipolar is used. As discussed in Chapter 3 , cardiac pacing stimuli in CIEDs are cathodal in polarity.

Unipolar leads contain only a single conductor and a single layer of insulation coating the outside of the conductor. Thus these leads may be thinner than bipolar leads, which require two conductor cables and two layers of insulation. However, advances in bipolar lead design have enabled size reduction to occur. Early bipolar leads were often associated with inner insulation degradation, leading to a short between the two conductors and resulting in electrical noise and a rise in the apparent threshold. For this reason, unipolar leads were demonstrated to have increased reliability. With modern bipolar leads, the use of stable and reliable insulation materials has improved reliability, reducing the use and availability of unipolar leads.

Unipolar pacing leads have the advantage of being historically more reliable than bipolar pacing leads, though this advantage may be reduced by improvements in lead insulation. In addition, when a constant voltage stimulus waveform is used, the stimulation threshold is lower for unipolar than for bipolar pacing. The lower stimulation threshold for unipolar stimulation is related to the lower impedance afforded by a very large surface area anode (the pulse generator case) compared with the much smaller surface area of a ring electrode that serves as the anode during bipolar pacing. The disadvantages of unipolar stimulation include the potential for pectoralis muscle stimulation from the pulse generator (the anode) and a higher chance of cross talk between sensing channels in the atrium and ventricle due to the much larger stimulus artifact. The primary advantage of bipolar leads is the much improved sensing provided by having two closely spaced electrodes within a cardiac chamber. Because of the risk of sensing myopotentials from either the pectoralis muscles or the diaphragm, unipolar sensing is not used for determination of rate in implantable cardioverter-defibrillators (ICDs). Unipolar sensing is used for recording electrogram morphology from the distal coil of transvenous ICD leads and may offer clues to the differential diagnosis of a tachyarrhythmia. Bipolar pacing eliminates pectoralis muscle stimulation and is always preferred when the pulse generator is placed in a submuscular pocket.

Integrated Bipolar and Dedicated Bipolar Defibrillation Leads

Modern DF leads are available in two configurations—the integrated bipolar version and the dedicated bipolar version. The integrated bipolar version uses the distal (right ventricle [RV]) shock coil both as a DF electrode and as the anode for pacing and sensing. The dedicated bipolar version has a dedicated shock coil and a dedicated anode electrode for pacing and/or sensing. The merits of both have been debated often. Recent data indicate that the rate of noise and oversensing with both lead configurations is not different. The benefit of the integrated bipolar version may be its lesser complexity (fewer conductors and connections), leading to increased reliability.

Equivalent Electric Circuit of a Cardiac Implantable Electronic Device IN Vivo

An in vivo CIED is an electrochemical system with the components of a lead playing different roles in the equivalent electric circuit (see Fig. 11-1 ). The extracellular fluid (ECF) and intracellular fluid (ICF) are electrolytes containing dissolved ions. An electrical current encountering the electrode-electrolyte interface changes from the conduction of electrons in the metal components of the lead to the conduction of charged ions within the tissues and fluids of the body with a complex dynamic structure. The myocardium is also a complex electrochemical entity, both structurally and functionally.

Regardless of their actual physical forms, the components of any electric circuit can be represented as a combination of three elements: resistor, capacitor, and inductor. The impedance Z (in ohms [Ω]) of a CIED system in vivo summarizes the behavior of the equivalent electric circuit (pacing or DF) by relating the total input voltage V _total (in volts [V]) from the pulse generator to the output current I _total (in amperes [A]) via:

V_{total} = Z \cdot I_{total}

$V_{total} = Z \cdot I_{total}$

Electric Circuit Elements: Resistors, Capacitors, and Inductors

An “ideal” resistor is characterized by its resistance R (in ohms [Ω]), the ratio of the voltage V _R across to the electric current I flowing through:

V_{R} = IR

$V_{R} = IR$

By Ohm's law, the resistance R stays constant regardless of the current I.

An “ideal” capacitor is characterized by its capacitance C (in farads [F]), the ratio of the charge Q (in coulombs [C]) stored to the voltage V _C across:

C = \frac{Q}{V_{C}}

$C = \frac{Q}{V_{C}}$

The definition of capacitance C given in Equation 11-3a assumes the ratio Q/V _C stays constant as V _C changes. If this is not the case, differential capacitance is used:

C = \frac{dQ}{{dV}_{C}}

$C = \frac{dQ}{{dV}_{C}}$

An “ideal” inductor is characterized by its inductance L (in henrys [H]), the ratio of the voltage V _L across to the rate of change of the electric current I flowing through:

V_{L} = L \frac{dI}{dt}

$V_{L} = L \frac{dI}{dt}$

Inductance originates from the electromotive force (EMF) induced by change in the magnetic flux Φ through the inductor due to change in the electric current I (see below) and may not stay constant (such as when the inductor has a ferromagnetic core).

The Basic Resistor-Capacitor-Inductor Circuit

This section gives an overview of impedance, capacitance, and inductance as they relate to the flow of electric currents to illuminate the biophysics of leads and electrodes. To begin, consider an electric circuit segment consisting of a resistor, a capacitor, and an inductor in series with an electric potential V(t) applied across ( Fig. 11-2A ). If V(t) is periodic (i.e., it repeats the same pattern over time, or V(t + t ₀ ) = V(t) for some constant t ₀ ), then it can be represented as the sum of sinusoidal oscillations (i.e., a Fourier series). The same electric current I(t) flows through all three elements (Kirchhoff's current law). For simplicity and anticipating subsequent analyses, suppose I(t) is sinusoidal with frequency f (unit hertz, Hz), such that:

I (t) = I_{0} \cos ω t

$I (t) = I_{0} \cos ω t$

where I ₀ is some constant and ω = 2πf is the “angular frequency” (unit sec ⁻¹ ). By Equations 11-2 through 11-4 :

V_{R} (t) = I (t) R = I_{0} R \cos ω t

$V_{R} (t) = I (t) R = I_{0} R \cos ω t$

V_{C} (t) = \frac{Q}{C} = \frac{1}{C} \int I_{0} \cos ω t dt = \frac{I_{0}}{ω C} \sin ω t = \frac{I_{0}}{ω C} \cos (ω t - \frac{π}{2})

$V_{C} (t) = \frac{Q}{C} = \frac{1}{C} \int I_{0} \cos ω t dt = \frac{I_{0}}{ω C} \sin ω t = \frac{I_{0}}{ω C} \cos (ω t - \frac{π}{2})$

\begin{matrix} V_{L} (t) = L \frac{dI}{dt} = L \frac{d}{dt} (I_{0} \cos ω t) = - I_{0} ω L \sin ω t \\ = I_{0} ω L \cos (ω t + \frac{π}{2}) \end{matrix}

$\begin{matrix} V_{L} (t) = L \frac{dI}{dt} = L \frac{d}{dt} (I_{0} \cos ω t) = - I_{0} ω L \sin ω t \\ = I_{0} ω L \cos (ω t + \frac{π}{2}) \end{matrix}$

With respect to the electric current I(t), the voltage across the resistor V _R (t) stays in phase, the voltage across the capacitor V _C (t) lags by a quarter of a cycle (−π/2), and the voltage across the inductor V _L (t) leads by a quarter of a cycle (+π/2) (see Fig. 11-2B, C ).

Figure 11-2, Electrical Behavior of an Idealized Resistor-Capacitor-Inductor Circuit.

By Kirchhoff's voltage law, the total voltage across the three elements V(t) is given by:

\begin{matrix} V (t) = V_{R} (t) + V_{C} (t) + V_{L} (t) = I_{0} [R \cos ω t - (ω L - \frac{1}{ω C}) \sin ω t] \\ = I_{0} | Z | \cos (ω t + ϕ) \end{matrix}

$\begin{matrix} V (t) = V_{R} (t) + V_{C} (t) + V_{L} (t) = I_{0} [R \cos ω t - (ω L - \frac{1}{ω C}) \sin ω t] \\ = I_{0} | Z | \cos (ω t + ϕ) \end{matrix}$

Where:

| Z | = \sqrt{R^{2} + {(ω L - \frac{1}{ω C})}^{2}}

$| Z | = \sqrt{R^{2} + {(ω L - \frac{1}{ω C})}^{2}}$

ϕ = \tan^{- 1} \frac{ω L - 1 / ω C}{R}

$ϕ = \tan^{- 1} \frac{ω L - 1 / ω C}{R}$

V(t) is also sinusoidal but has a phase shift ϕ with respect to I(t) = I ₀ cosωt.

Equations 11-6a through 11-6c give an expression for V(t), but not in terms of I(t). To do this requires expressing V(t), I(t), and other relevant terms as complex numbers with the aid of Euler's formula e ^iθ = cosθ + isinθ, where i is the imaginary root (i ² = −1). Defining impedance Z as V(t)/I(t), then:

Z = \frac{V (t)}{I (t)} = R + i (ω L - \frac{1}{ω C})

$Z = \frac{V (t)}{I (t)} = R + i (ω L - \frac{1}{ω C})$

| Z | = \sqrt{R^{2} + {(ω L - \frac{1}{ω C})}^{2}}

$| Z | = \sqrt{R^{2} + {(ω L - \frac{1}{ω C})}^{2}}$

\arg Z = \tan^{- 1} \frac{ω L - 1 / ω C}{R}

$\arg Z = \tan^{- 1} \frac{ω L - 1 / ω C}{R}$

(See Appendix 11-1 for derivation.) Note that the expressions for |Z| in Equations 11-6b and 11-7b and for ϕ and arg Z in Equations 11-6c and 11-7c are identical.

Resistance and Reactance

If (1) the resistance of the circuit R > 0 or (2) the driving frequency ω does not match the circuit's natural frequency, , impedance Z does not vary with time from Equation 11-7a . The real part Re Z is called the resistance and is entirely due to the resistor. The imaginary part Im Z is called the reactance and has two components: X _C due to the capacitor and X _L due to the inductor:

X_{C} = - \frac{1}{ω C}

$X_{C} = - \frac{1}{ω C}$

X_{L} = ω L

$X_{L} = ω L$

Z = R + i (X_{C} + X_{L})

$Z = R + i (X_{C} + X_{L})$

Compared with resistance, reactance has a phase shift and is frequency-dependent. The signs for X _C and X _L in Equations 11-8a and 11-8b signify the phase shift of −π/2 and +π/2 with respect to the resistance R in the complex plane (see Fig. 11-2D ), corresponding to the phase shift of V _C and V _L with respect to V _R (see Fig. 11-2C ). The dependence of the amplitude of impedance |Z| on frequency can be summarized by plotting log |Z| against log ω or log f (Bode plot; see Fig 11-2E ). Capacitors are more effective in blocking low-frequency electric input (ω↓, |X _C |↑; high-pass filter), whereas inductors are more effective in blocking high-frequency electric input (ω↑,|X _L |↑; low-pass filter).

Resonance

If (1) the resistance of the circuit R = 0 and (2) the driving frequency ω matches the circuit's natural frequency, , resonance between the driving force and the electric circuit ensues. The amplitude of Z decreases with time (|Z| ∝ 1/t), which means the amplitude of current |I| increases with time (|I| ∝ t). (The phase angle of Z also changes with time and approaches 0 asymptotically. See Appendix 11-1 for details.)

In practice, the resistance of a circuit will never be truly 0, but it can be comparatively low. When the driving frequency approximates the natural frequency of the circuit, energy is not efficiently dissipated as heat in the resistor and a significant amount can be progressively stored as magnetic and electric energies in the inductor and the capacitor. When the accumulated energy exceeds the storage capacity of the circuit, an uncontrolled discharge may occur. Such a scenario can theoretically occur during magnetic resonance imaging (MRI), with potentially dangerous consequences for the patient.

Transient Currents: Damping and Time Constants

If V(t) = V ₀ constant, then:

(See Appendix 11-1 for details.)

Equations 11-9a, 11-9b, and 11-9c correspond to heavy, critical, and light damping of the transient current that flows when the total voltage V stops varying (see Fig 11-2F ). All three responses decay exponentially in amplitude over time, and the rate of decay is characterized by the time constants:

τ_{RC} = RC

$τ_{RC} = RC$

τ_{RL} = L / R

$τ_{RL} = L / R$

In general, 37% (= e ⁻¹ ) of the remaining available range is attained within one time constant period. Heavy damping is favored by a large capacitance R ² C > 4L, whereas critical and light damping is favored by a low capacitance R ² C ≤ 4L.

For heavy damping ( Equation 11-9a ), the transient electric current decreases exponentially with time with no oscillations (see Fig. 11-2F ). Significantly, what is slow to dissipate is equally slow to build up (see Fig. 11-2G, H ). Large time constants (favored by a large capacitance and a large inductance) decrease the amplitude of any unintended electric signal (“noise”) that develops within a period of time, reducing the chance of false sensing by a CIED. However, any low-amplitude electric noise also tends to persist for a longer period of time with large time constants.

For critical damping ( Equation 11-9b ), the transient current disappears most quickly (see Fig. 11-2F ), but the precise condition required (R ² C = 4L) makes it unlikely to occur fortuitously without a very deliberate attempt to tune the electric circuit in practice.

For light damping ( Equation 11-9c ), the transient current decreases exponentially with time with oscillations (see Fig. 11-2F ), but the condition required (R ² C < 4L) means the associated time constant τ _RL = L/R tends to be large. The transient current retains significant amplitude for a considerable period of time and may cause false sensing by a CIED (see Fig. 11-2F ).

Thus the chance of false sensing by a CIED is decreased by having a large capacitance in the equivalent electric circuit, which favors heavy damping over critical and light damping and reduces the amplitude of any electric noise.

Combining Impedances IN Series and IN Parallel

The equivalent impedances of two impedances Z ₁ and Z ₂ in series Z _series and in parallel Z _parallel are given by:

Z_{series} = Z_{1} + Z_{2}

$Z_{series} = Z_{1} + Z_{2}$

\frac{1}{Z_{parallel}} = \frac{1}{Z_{1}} + \frac{1}{Z_{2}} \Leftrightarrow Z_{parallel} = \frac{Z_{1} Z_{2}}{Z_{1} + Z_{2}}

$\frac{1}{Z_{parallel}} = \frac{1}{Z_{1}} + \frac{1}{Z_{2}} \Leftrightarrow Z_{parallel} = \frac{Z_{1} Z_{2}}{Z_{1} + Z_{2}}$

(Ohm's law and Kirchhoff's laws). For series connection, the equivalent impedance is dominated by the electric circuit element arrangement with the higher impedance. For parallel connection, the equivalent impedance is dominated by the electric circuit element arrangement with the lower impedance. (When more than one route exists, electricity will always take the path of the least resistance, or lowest impedance.) The equivalent impedance of more complex connections of electric circuit element arrangements can be obtained by iteration of Equations 11-11a and 11-11b .

For two capacitors of capacitance C ₁ and C ₂ , by Equations 11-8a, 11-11a, and 11-11b , their combined capacitance in series C _series and in parallel C _parallel are given by:

\begin{matrix} Z_{series} = i X_{C_{series}} = i X_{C_{1}} + i X_{C_{2}} \Rightarrow - \frac{1}{ω C_{series}} = - \frac{1}{ω} (\frac{1}{C_{1}} + \frac{1}{C_{2}}) \\ \Rightarrow \frac{1}{C_{series}} = \frac{1}{C_{1}} + \frac{1}{C_{2}} \end{matrix}

$\begin{matrix} Z_{series} = i X_{C_{series}} = i X_{C_{1}} + i X_{C_{2}} \Rightarrow - \frac{1}{ω C_{series}} = - \frac{1}{ω} (\frac{1}{C_{1}} + \frac{1}{C_{2}}) \\ \Rightarrow \frac{1}{C_{series}} = \frac{1}{C_{1}} + \frac{1}{C_{2}} \end{matrix}$

\begin{matrix} \frac{1}{Z_{parallel}} = \frac{1}{i X_{C_{parallel}}} = \frac{1}{i X_{C_{1}}} + \frac{1}{i X_{C_{2}}} \\ \Rightarrow - ω C_{parallel} = - ω (C_{1} + C_{2}) \Rightarrow C_{parallel} = C_{1} + C_{2} \end{matrix}

$\begin{matrix} \frac{1}{Z_{parallel}} = \frac{1}{i X_{C_{parallel}}} = \frac{1}{i X_{C_{1}}} + \frac{1}{i X_{C_{2}}} \\ \Rightarrow - ω C_{parallel} = - ω (C_{1} + C_{2}) \Rightarrow C_{parallel} = C_{1} + C_{2} \end{matrix}$

For two inductors of inductance L ₁ and L ₂ , by Equations 11-8b, 11-11a, and 11-11b , their combined inductance in series L _series and in parallel L _parallel are given by:

\begin{matrix} L_{series} = i X_{L_{series}} = i X_{L_{1}} + i X_{L_{2}} \Rightarrow ω L_{series} = ω (L_{1} + L_{2}) \\ \Rightarrow L_{series} = L_{1} + L_{2} \end{matrix}

$\begin{matrix} L_{series} = i X_{L_{series}} = i X_{L_{1}} + i X_{L_{2}} \Rightarrow ω L_{series} = ω (L_{1} + L_{2}) \\ \Rightarrow L_{series} = L_{1} + L_{2} \end{matrix}$

\begin{matrix} \frac{1}{Z_{parallel}} = \frac{1}{i X_{L_{parallel}}} = \frac{1}{i X_{L_{1}}} + \frac{1}{i X_{L_{2}}} \\ \Rightarrow \frac{1}{ω L_{parallel}} = \frac{1}{ω} (\frac{1}{L_{1}} + \frac{1}{L_{2}}) \Rightarrow \frac{1}{L_{parallel}} = \frac{1}{L_{1}} + \frac{1}{L_{2}} \end{matrix}

$\begin{matrix} \frac{1}{Z_{parallel}} = \frac{1}{i X_{L_{parallel}}} = \frac{1}{i X_{L_{1}}} + \frac{1}{i X_{L_{2}}} \\ \Rightarrow \frac{1}{ω L_{parallel}} = \frac{1}{ω} (\frac{1}{L_{1}} + \frac{1}{L_{2}}) \Rightarrow \frac{1}{L_{parallel}} = \frac{1}{L_{1}} + \frac{1}{L_{2}} \end{matrix}$

Impedance

Impedance is a parameter that is always measured during CIED implantation and also is used for monitoring the performance and guiding the management of a CIED system. The impedance Z measured in clinical practice is more complex than the simple ohmic relationship suggested by Equation 11-1 . The effects of resistance, capacitance, and inductance on impedance are more complex than suggested by Equations 11-8a through 11-8c , which cannot be applied if the stimulation waveform is not periodic and does not have a frequency. The testing waveform and measurement methods naturally affect the value obtained for impedance. Other than that, the interface between the lead tip and the myocardium, the geometric surface area (GSA) and electrochemical surface area (ESA) of each electrode involved in the circuit, and the structural integrity of the conductors, insulation, and electric connections all affect the impedance.

Geometrical and Electrochemical Surface Areas of Electrodes

The GSA of an electrode is its macroscopic area as estimated from its dimensions (length, width, and height). The ESA of an electrode is the microscopic area over which electrochemical processes take place and generally is estimated experimentally (e.g., with cyclic voltammetry). The ESA affects impedance indirectly through the double-layer capacitance and kinetics of charge transfer (faradaic) reactions at the electrode-electrolyte interface. The ESA is roughly proportional to the GSA. The ESA : GSA ratio is called the “roughness factor” and can be >2000 for the electrode on a CIED lead. The GSA can be used as a surrogate measure of the ESA, provided the surface structure (roughness factor) stays the same.

Measurement of Impedance IN Cardiac Implantable Electronic Devices

The waveform used for impedance measurement in the latest generations of CIEDs is a constant current pulse (specified in amperes) rather than constant voltage pulse used for stimulation (measured in volts) ( Fig. 11-3 ). When exposed to a fixed current I, the voltage V between the two electrodes in the equivalent electric circuit increases rapidly initially, then at a decreasing rate before reaching a “plateau.” The average or “plateau” voltage is divided by the “average” (if several amplitudes are used) constant current to obtain “impedance” Z = V/I. In reality, the relationship between the potential across the electrode-electrolyte interface and the current I is not linear, and the measured impedance Z varies with the measurement waveform (amplitude and duration of the constant current pulses). The waveforms and algorithms used for estimating the voltage response and impedance vary from manufacturer to manufacturer and are generally proprietary information. An implanted lead may be given different impedance values by different measurement methods. This has clinical implications, as impedance is used to monitor a lead's performance and guide its management.

Figure 11-3, Impedance Measurement in Cardiac Implantable Electronic Devices (CIEDs).

The current used for impedance measurement is usually very small (80-750 µA), so that for an “impedance” approximately 1000 Ω, the resulting voltage will still be 80-750 mV, distinctively larger than the electric signals originating from the body (approximately 10 mV) and smaller than the impulses used for cardiac stimulation (approximately 3 V for pacing and approximately 1600 V for DF). A low-amplitude, constant current waveform will also prevent unintended cardiac stimulation and patient discomfort. The constant current waveform is “charge-balanced” (i.e., as much electric charge is injected into as is removed from the electrode) to avoid unintended and undesirable electrolysis and electrode corrosion.

Lead Tip Embedment IN Myocardium

Because the electric conductivity of blood is higher than that of the myocardium, the impedance generally increases when the tip electrode becomes effectively embedded in the myocardium, by approximately 50 Ω for the right atrium (RA) and approximately 200 Ω for the RV, compared with when it is ineffectively embedded or free floating in the cardiac chamber. Alternatively, a fall in impedance by the same magnitude may indicate lead tip dislodgement, even after chronic placement ( Fig. 11-4 ).

Figure 11-4, A and B, Impedance fall due to lead dislodgement. Immediately after cardiac resynchronization therapy defibrillator (CRTD) upgrade from a dual-chamber pacemaker system implanted 7 years beforehand, the atrial lead impedance, which was previously stable, began to show a progressive downward drift associated with a gradual increase in the pacing threshold until capture could no longer be achieved. The atrial lead was considered to have dislodged and was replaced. The atrial lead impedance became stable afterward.

Multiple Small Area Electrodes

The GSA (ESA) of a stimulating cathode needs to be small in order to reduce energy consumption and lower the pacing threshold (see below). A multipolar left ventricular (LV) lead may have up to four electrodes, all designed to be capable of being the stimulating cathode in order to allow multiple pacing vector configurations. As a result, all the electrodes on a multipolar LV lead have similar (small) GSAs. In contrast, the ring electrode typically has a larger GSA than the tip electrode on a bipolar lead for use in the RA or RV. The bipolar impedance between a pair of electrodes on a LV lead deployed in the coronary sinus (CS) is thus typically higher (commonly >1600 Ω) than for a bipolar lead deployed in the RV. When a LV lead is deeply wedged inside a small-caliber side branch of the CS, both electrodes in a bipolar configuration can be heavily embedded in tissues, resulting in an extremely high impedance (>2000 Ω) that is frequently associated with either a very high pacing threshold or outright noncapture (see later).

Out-of-Range Values during Lead Implantation and Follow-Up

The causes of out-of-range impedance during lead implantation are summarized in Table 11-2 . Most of these causes are “benign” and can be easily rectified by checking the lead connection or lead redeployment. However, a high impedance may also indicate serious operative complications such as when the tip electrode perforates a cardiac chamber and enters the pericardial or even the pleural spaces.

TABLE 11-2

Causes of Out-of-Range Lead Impedance During Implantation

Impedance Value	Possible Causes
Too high	Testing clips/cables not properly connected to lead pin or programmer
	Cardiac perforation by lead tip (into the pericardial sac or lungs)
	“Wedging” of left ventricular lead tip into small caliber coronary sinus side branch (bipolar pacing)
	Pneumothorax (unipolar pacing and defibrillation)
	Lead connector pin incompletely inserted into header of pulse generator
Too low	Testing cable clips touching each other or metal surgical instruments (electric short-circuit)
	Overfixation leading to tissue damage or lead tip dislodgement from the myocardium
	Pulmonary edema (unipolar pacing and defibrillation)
	Transposition of connector pins of RV and SVC shock coils (dual-coil DF1 leads only)

DF1, Unipolar defibrillation; RV, right ventricle; SVC, superior vena cava.

In a chronically implanted lead, a high impedance (greater than the upper limit of normal) is much more frequently observed than a low impedance (less than the lower limit of normal). An abrupt rise in impedance can be either sustained or intermittent, generally signifies a disruption in the electric circuit, and is associated with electric malfunctions ( Fig. 11-5 ). Some disruptions in the electric circuit may be visualized by radiography or fluoroscopy ( Fig. 11-6 ). Electric circuit disruption associated with only intermittent impedance rise may not cause any anomaly on periodic impedance measurement and can be detected only by more sophisticated algorithms based on nonphysiologic noise. In contrast, a gradual rise in impedance reflects a change in the electrode-myocardium interface (e.g., calcification of the fibrous tissue) and can be entirely compatible with satisfactory electric function.

Figure 11-5, Causes and Patterns of High Lead Impedance Postimplantation.

Figure 11-6, Lead Electric Circuit Disruption Diagnosable From Fluoroscopy or X-Rays.

In a chronically implanted lead, a fall in impedance from the background value may indicate lead tip dislodgement (see Fig. 11-4 ) or development of an insulation breach. An insulation breach is often not associated with any detectable change in impedance, especially if it involves the outer insulation. The exposed conductor element and the surrounding blood or ECF act as two resistors arranged in parallel. The overall resistance of the arrangement is dominated by the resistor with the lower resistance ( Equation 11-11b ), which is invariably the exposed conductor element because the electric conductivity of metals is so much higher than that of blood and ECF. However, an insulation breach involving the inner insulation between conductors is associated with a significant drop in impedance if an electrical short-circuit develops between the exposed conductor element and another metal conductor (e.g., another conductor or the pulse generator can).

Electrodes

The electrode plays a pivotal role in the functioning of a CIED system, but its importance may be underappreciated because it is extremely reliable and unlikely to be the cause of any clinically significant structural failure or electric malfunction. The electrode-ECF interface is a phase boundary between solid and liquid, a transition zone between electronic and ionic current conduction, and the scene of physical and electrical interactions between artificial and biological elements (see Appendix 11-2 for brief explanations of the electromagnetic concepts relevant to the interface). In fact, the lead body is merely a transmission line connecting the pulse generator to an electrode placed at an anatomic site dictated by the electrical functions of the CIED system. Even CIEDs without leads (internal loop recorders, leadless pacemakers ) still contain electrodes. In forming an electric circuit, the lead may provide only one conductor-electrode set, with the intervening body tissues and the can of the pulse generator forming the other conductor-electrode set (unipolar sensing and pacing, single-coil DF) (see Fig. 11-1 ).

The Electrode-Electrolyte Interface: the Electric Double Layer

The electrode-electrolyte interface contains an “electric double layer,” which forms whenever a solution containing dissolved ions (an electrolyte such as blood and ECF) comes into contact with an object (which may be a gas bubble, a liquid droplet, or a solid shape such as the electrode of a lead) with which it does not mix ( Fig. 11-7 ; Appendix 11-3 ). Even when the electrode is not carrying any electric charge, certain ions adsorb onto the electrode's surface by the van der Waals force. Solvent molecules with a “dipole” (equal but opposite electric charges separated a distance apart) may also attach to the electrode's surface in a certain preferential orientation. The adsorbed ions and solvent molecule dipoles together form a compact layer of electric charge on the electrode's surface. The surface charge attracts ions of opposite polarity (which may be solvated or surrounded by solvent molecule dipoles) and repels ions of the same polarity. The aggregation of the same ionic species in a region is limited by the species' mutual lateral electrostatic repulsion and thermal motion (diffusion under concentration gradient). The balance of these counteracting forces creates a second diffuse layer of oppositely charged ions around the electrode. The compact and diffuse layers together form an electric double layer at the electrode-electrolyte interface, and the opposite charges contained in them maintain the electric neutrality of the interface.

Figure 11-7, Double Layer at the Electrode-Electrolyte Interface.

When the electrode carries an electric charge, ions of the same polarity (co-ions) or opposite polarity (counterions) and solvent molecules are still adsorbed onto its surface by the van der Waals force in the compact layer. In the diffuse layer, counterions are attracted to the electrode by electrostatic force but prevented from contacting it by the compact layer. The compact layer effectively forms a dielectric separating the two opposite electric charges on the electrode and in the diffuse layer, giving rise to a double-layer capacitance. The compact layer also forms a physical barrier between the electrode and the ion species, impeding electric charge transfer and slowing down electrochemical (reduction-oxidation) reactions.

The structure of the double layer has been deduced on the basis of both theoretical considerations and experimental observations of the electrochemical properties of the electrode-electrolyte interface. In a composite version combining features from various models, the compact layer is made of two sublayers demarcated by the inner Helmholtz plane (IHP) and the outer Helmholtz plane (OHP) (see Fig. 11-7 and Appendix 11-3 ). The IHP contains adsorbed solvent molecules and ions (including co-ions such as chloride Cl ⁻ ions for a CIED cathode in vivo) and is approximately 0.2 nm from the electrode surface. The OHP contains solvated counterions (sodium Na ⁺ ions for a CIED cathode in vivo) at their closest approach to the surface of the electrode and is approximately 0.2 nm from the IHP. The concentrations of the counterions are higher, and those of the co-ions are lower, within the diffuse layer (approximately 1-10 nm thick). Sufficiently far (approximately 30,000 nm or 0.03 mm) away from the diffuse layer, in the bulk of the solution, the concentrations of the counterions and co-ions return to their equilibrium levels.

Double-Layer Capacitance

Double-layer capacitance is a surface phenomenon and related to the potential and charge across the double layer in a complex way, rising slightly with both instead of staying constant ( Fig. 11-8 , Appendix 11-3 ). For the same electrolyte composition, the double-layer capacitance per unit area of the electrode-electrolyte interface (the ESA of the electrode) C _dl is largely fixed. For optimal functioning of a lead in a CIED system, the double-layer capacitance of an electrode should ideally be as large as possible (see below). Because the GSA of the electrode is limited by the size of the lead, and due to the need to keep the stimulation threshold low, a good way to increase the double-layer capacitance of an electrode under these constraints is by increasing the ESA per unit GSA of the electrode (i.e., the roughness factor).

Charge Transfer (Faradaic) Reactions

Charge transfer (faradaic) reactions are electrochemical reactions involving electron transfer between atomic, molecular, and ionic reactants in compulsorily coupled pairs of red uction (gain of electrons) and ox idation (loss of electrons) (redox reactions).

A large number of faradaic reactions are thermodynamically possible for an electrode in vivo. The standard reduction potential gives a measure of how thermodynamically favorable (compared with the reduction of hydrogen ions) a reaction is: the higher the potential, the more favorable the reaction (reduction) ( Table 11-3 ). The sign of the potential matters: a negative reduction potential means the opposite reaction (oxidation) is thermodynamically more favorable. By the standard reduction potential, some redox reaction pairs will occur spontaneously. If reduction and oxidation are forced to occur at two spatially segregated electrodes connected through an external circuit, energy from the redox reaction will be released as the flow of electric charges between the two electrodes, giving rise to a galvanic cell (a “battery”). Some redox reaction pairs, however, require an input of external electrical energy to occur, giving rise to an electrolytic cell.

TABLE 11-3

Selected Standard Reduction Potentials in Aqueous Solutions at 25° C Against the Normal Hydrogen Electrode for Faradaic Reactions Possible In Vivo

Data from Bard AJ, Faulkner LR: Reference tables. In: Electrochemical methods: fundamentals and applications, ed 2, New York, 2001, John Wiley, pp 808-813; and Newman J, Thomas-Alyea KE: Thermodynamics in terms of electrochemical potentials. In: Electrochemical systems, ed 3, Hoboken, NJ, 2004, John Wiley, pp 29-83.

The standard potential indicates the thermodynamic feasibility of a reaction but does not indicate the speed (or rate) at which it occurs. The rate of electrolysis is determined by many factors, including the amplitude of the applied potential (the “overpotential”), the structure of the electrodes, the concentrations of involved chemical species at the interface with the electrode and within the bulk of the solution, transport of the chemical species between the two places, and others (see Appendix 11-3 ). On the basis of electrochemical potential considerations alone, many of the chemical species within the body's ECF will undergo electrolysis for the voltages typically used for cardiac pacing (2.5 V) and DF (800 V for a transvenous ICD system) (see Table 11-3 ). Such chemical reactions are highly undesirable because the product species tend to be chemically very reactive (e.g., hydrogen peroxide) and can corrode the electrodes or lead to tissue injury. Some product species can even be gaseous at body temperature and may form microbubbles within the bloodstream. Electrolysis is generally not a significant problem with cardiac stimulation in vivo, because the pulse duration is too limited to saturate the double-layer capacitance at the electrode-electrolyte interface ( Box 11-1 ). Charge-balanced current waveforms remove any excess electric charge that has accumulated on the electrodes and reverse (at least partially) any chemical reactions that may have occurred ( Fig. 11-9 ). An electric potential can also be applied to an electrode (even when the CIED is not stimulating) to prevent thermodynamically favorable faradaic reactions from occurring spontaneously. These processes together protect the body and the electrodes from harmful electrolytic reactions in vivo. Electrolytic reactions may become more significant if the potential stays in the same direction (i.e., direct current) and the pulse duration is longer (e.g., in the second range) with electric impulses used for ventricular fibrillation induction, but such stimulation is performed only infrequently and no clinically significant adverse effects have been reported.

Box 11-1

Charging and Discharging of Lead Electrodes' Double-Layer Capacitance

Assuming the double-layer capacitance C _dl is 0.2 µF⋅mm ⁻² for a lead electrode in vivo, the total capacitance C for a pacing electrode with an electrochemical surface area (ESA) of 5.7 mm ² is 1.14 µF and for a shock coil with an ESA of 660 mm ² is 132 µF. If the resistance R for pacing is 800 Ω, the corresponding time constant τ _RC = RC is 0.912 msec for the low-voltage pacing circuit, significantly longer than the typical pulse width of a pacing impulse (0.4 msec). Similarly, if the resistance R for defibrillation (DF) is 60 Ω, the corresponding time constant τ _RC is 7.9 msec for high-voltage DF circuit, significantly longer than the typical pulse width of a DF impulse (4 msec). This means the double-layer capacitance at the electrode-electrolyte interface is not saturated during either pacing or defibrillation and hence able to absorb most of the electric charges delivered to the electrodes, diverting them away from faradaic charge transfer reactions that may corrode the electrodes or produce noxious chemical species.

Figure 11-9, A-C, Charge-balanced current waveform used in cardiac stimulation. I a , Anodic current; I c , cathodic current; t a , anodic phase period; t c , cathodic phase period; t ip , interphase dwell. Cardiac stimulation is assumed to be at an output of 2.5 V at 0.4-msec pulse width with 800-Ω impedance. The charge-balanced current waveform removes electric charge accumulated on the electrodes after stimulation, accelerating dissipation of interface polarization due to ion concentration gradients in the double layer and reversing (at least partially) any faradaic redox reactions. The impulses used for cardiac pacing are “constant voltage” instead of “constant current,” as the pulse width is typically much shorter than the time constant for the corresponding resistor-capacitor circuit and hence the voltage from the output capacitor is effectively “constant” throughout the duration of the impulse. Provided the “impedance” (see Fig. 11-3 ) stays constant, the pacing current is also constant. Electronic circuitry within the pulse generator is used to generate the anodic current and make sure the biphasic waveform is “charge-balanced.”

The impedance related to charge transfer processes (see Fig. 11-1 ) comes from redox reaction kinetics on the electrode surface and mass transport (by diffusion) of reactants within the solution to the electrode surface (the Warburg impedance). Because charge transfer does not play a major role in the functioning of the CIED and leads in vivo, these concepts are not discussed further.

Pseudocapacitance

Pseudocapacitance is an alternative way of electric charge (energy) storage alongside double-layer capacitance in the presence of a suitable electrode-electrolyte interface. Pseudocapacitance operates by (1) surface faradaic redox reactions that change only the valence state but not the chemical bonds or physical phases of the atomic particles constituting the electrode and the adsorbed desolvated ions from the electrolyte or (2) intercalation and/or electrosorption of the desolvated ions in atomic-sized spaces or pores in the electrodes. These processes are rapid and highly reversible, and the voltage of the surface redox reaction is proportional to the state of charge of the electrode-electrolyte interface. In these senses, pseudocapacitance behaves more similarly to a conventional capacitor than to a battery because the voltage is largely independent of the state of charge. Pseudocapacitance can be many times the value of double-layer capacitance.

Of the two materials most commonly used for coating lead electrodes, iridium oxide (IrO ₂ ) has much more pseudocapacitance than titanium nitride (TiN).

Polarization and Depolarization of the Electrode-Electrolyte Interface

The term polarization has different meanings in electromagnetism, electrochemistry, and other contexts. For CIEDs, interface polarization and depolarization are taken to mean charging and discharging of the double-layer capacitance C _dl between the electrode and the ECF. A capacitor functions electrically as an electric charge energy store and a high-pass filter. Interface (de)polarization allows transient electric currents to flow without inducing electrolysis and electrode corrosion (charge transfer reactions) and is essential for the functioning of the electrode in the equivalent electric circuit of a CIED in vivo (see Fig. 11-1 ). Interface (de)polarization interferes with cardiac sensing (see Table 11-1 ). In general, interface (de)polarization involving a high polarization charge but a low polarization potential is desirable, which means a high double-layer capacitance C _dl .

Electric Consequences of High Double-Layer Capacitance

A high C _dl reduces the interface potential, leaving a higher portion of the interelectrode voltage across the electrolyte to be experienced by the cell membrane (see Appendix 11-4 ), so that the stimulation threshold becomes lower ( Fig. 11-10A ). A lower interface potential (see Fig. 11-10B, C ) facilitates the detection of any evoked myocardial response (allowing the operation of automatic capture management algorithms) and is less likely to be sensed by the CIED (reducing the risk of inappropriate administration or withholding of device therapy). However, a high C _dl also means the interface potential will linger longer (see Fig. 11-2F ), but an opposite charge-balanced electric impulse will accelerate its dissipation (see Fig. 11-9 ). A high C _dl reduces the impedance and attenuation of low-frequency signals ( Equation 11-8a ) (see Figs. 11-2E and 11-10D ), allowing high-fidelity sensing of intrinsic myocardial signals (see Fig. 11-10E, F ). A high C _dl absorbs the majority (ideally the entirety) of the stimulation current, preventing electrochemically possible faradaic electrolysis reactions (see Appendix 11-3 ) that may produce noxious chemicals or corrode the electrode in vivo. A high C _dl allows the GSA of the electrode to be smaller, thereby reducing the stimulation threshold (see below).

Figure 11-10, Electric Effects of Increased Double-Layer Capacitance.

Mechanisms of Increasing Double-Layer Capacitance

Double-layer capacitance can be increased by increasing either the ESA (compared with the GSA; i.e., the roughness factor) or the surface capacitance density (capacitance per unit surface area). Increasing the ESA while maintaining the same GSA is achieved by either everted (“fractal” coating; roughly, repetitive self-superimposition of the same fundamental geometric shape on an ever-decreasing scale) ( Fig. 11-11A ) or inverted (pores, channels, indentations, and cracks) ( Fig. 11-12A -D) “folding” of the surface. The two processes produce opposite sides of the same surface (see Fig. 11-11B ). Inverted surface folding suffers from the inherent problem that access to the extra ESA is fixed by the GSA and requires active convection venting of the electrolyte solution (ECF) for the ESA to become effective (see Fig. 11-11B ). In contrast, everted surface folding allows access to the ESA, so that the ESA is effective even in the absence of active convection of the electrolyte solution (ECF).

Figure 11-11, Fractal Coating and Everted and Inverted Surface Folding.

Figure 11-12, Inverted and Everted Folding of Electrode Surface.

The surface capacitance density can be increased through pseudocapacitance of the material (such as IrO ₂ and platinum) coating or forming the electrode.

{Ir}^{3 +} \to {Ir}^{4 +} + e^{-}

${Ir}^{3 +} \to {Ir}^{4 +} + e^{-}$

Ir {(OH)}_{n} \to {IrO}_{x} {(OH)}_{n - x} + x H^{+} + x e^{-}

$Ir {(OH)}_{n} \to {IrO}_{x} {(OH)}_{n - x} + x H^{+} + x e^{-}$

Pt + H_{2} O \to PtO + 2 H^{+} + 2 e^{-}

$Pt + H_{2} O \to PtO + 2 H^{+} + 2 e^{-}$

PtO + 2 H^{+} + 2 e^{-} \to PtH + {OH}^{-}

$PtO + 2 H^{+} + 2 e^{-} \to PtH + {OH}^{-}$

where n and x are positive integers. The reactions involving iridium occur within the coating without phase change of the reactants. The reactions involving platinum occur in a monolayer of platinum atoms on the surface of the electrode.

Fractal Coating

“Fractal” coating is realized in practice by a variety of manufacturing processes (sputtering, arc vapor deposition) that deposit layer after layer of granules of the same material (see Fig. 11-12E-H ) on the surface of an electrode. Because it is very difficult to control the size of the granules, the end product may be not truly a “fractal” but just a highly textured or porous surface.

In the context of CIED in vivo, if the GSA of an electrode is 1.3 mm ² , assuming a double-layer capacitance density of approximately 20 µF/cm ² (range, 10-40 µF/cm ² ) (see Appendix 11-3 ), the total double-layer capacitance will be 0.26 µF. Fractal coating can increases the ESA with respect to the GSA by more than 2000 times. Assuming fractal coating increases the ESA by up to 1000 times the GSA, the total double-layer capacitance will be 260 µF. In comparison, the capacitance of a tantalum electrolytic capacitor used in the high-voltage circuit of the ICDs ranges from 0.001 to 1000 µF.

The importance of fractal coating for enhancing CIED capabilities is illustrated by the subcutaneous ICD. The can of the transvenous ICD forms one of the DF electrodes. Even though the bulk of the can is typically made of titanium, its surface is covered by a thin layer of titanium dioxide, which is only slightly roughened at a microscopic level ( Fig. 11-13A ). The transvenous ICD may be used as an electrode in some auxiliary sensing channels; however, it is typically not used as the sensing channel for cardiac rhythm detection, because postshock polarization may interfere with sensing. In contrast, interface polarization around the textured can of a subcutaneous ICD typically results in much less artifact, and postshock sensing is not impaired (see Fig. 11-10B ). Fractal coating of the subcutaneous ICD can with TiN increases the ESA (see Fig. 11-13B ) and double-layer capacitance enough so that the interface polarization does not compromise accurate cardiac rhythm detection.

$Figure 11-13, Fractally and Nonfractally Coated Implantable Cardioverter Defibrillator (ICD) Pulse Generator Cans.$

The complex surface structure of fractal coating may encourage tissue ingrowth, helping to fix the lead electrode in position with respect to the surrounding myocardium during cardiac movement, which in turn reduces inflammation, fibrosis, and myocardial damage, independent of steroid elution. The electrode-myocardium distance may thus be shorter, reducing the stimulation threshold and increasing sensing fidelity.

Geometric Surface Area

The electric consequences of decreasing the GSA of electrodes for a lead can be illustrated by a simplified version of the equivalent electric circuit ( Figs. 11-14 and 11-15A ). In general, reducing the GSA of the stimulating cathode concentrates the voltage drop across the circuit to the myocardium, reducing the capture threshold (see Appendix 11-4 ) and energy consumption at the same time. These desirable functional effects can be accentuated by increasing the GSA of the nonstimulating anode, even when it is situated a considerable distance away (see later). A high impedance does not automatically improve the performance of a lead. For example, if the high lead impedance is achieved by increasing the resistance of the conductor elements, the electric performance of the lead will be worse in terms of capture threshold and energy consumption. The lead performance is optimized by accentuating the disparity in impedance between the electrode-myocardium interface and other components in the electric circuit. (In fact, highly conductive materials such as silver are intentionally used to lower the resistance of the conductor elements in leads.)

Figure 11-14, Distribution of Impedance, Pacing Threshold, and Energy Consumption.

Figure 11-15, Electrodes With Different Surface Areas and Interelectrode Distances.

Even though reducing the GSA of the electrode lowers the stimulation threshold and power consumption, it cannot be reduced indefinitely, as double-layer capacitance C _dl decreases with the GSA (even with ESA enhancement technologies) and interface polarization begins to significantly interfere with cardiac stimulation and sensing. Too small a GSA will also reduce the statistical probability of achieving a short electrode-myocardium distance and hence may increase the stimulation threshold (see Appendix 11-4 ).

An electrode with a large GSA (ESA) can be deliberately chosen to be the anode to reduce the chances of unintended anodal myocardial capture (see Fig. 11-14 ). The RV shock coil and the pulse generator can are routinely used as the anode, with the CS lead providing the cathode for LV stimulation in cardiac resynchronization therapy (CRT). Some older CS leads had different GSAs (ESAs) for the electrodes (notably the Medtronic 4194 CS lead, with distal tip electrode ESA 5.8 mm ² and proximal coil electrode ESA 38 mm ² ) (see Fig. 11-15B ). However, the quadripolar CS leads from St. Jude Medical (Quartet) and Boston Scientific (Acuity X4) have three proximal electrodes with identical GSAs (ESAs). In the case of the Boston Scientific Acuity X4 lead, the three proximal electrodes can be programmed as an anode or cathode. The distal electrodes for the St. Jude Medical and Boston Scientific leads have smaller GSAs (ESAs) than the proximal three electrodes and as such are not used as an anode. Medtronic's Attain Performa uses four electrodes with all of the same GSAs (ESAs), and all four can be used interchangeably as anode or cathode.

Most tip and ring electrodes are fractally coated with either IrO ₂ or TiN, but the shock coil electrodes for DF leads are typically made of platinum. The shock coil ESA is approximately 660 mm ² , whereas the tip and ring electrode ESA is approximately 5.7 mm ² . The large GSA of the shock coil compensates for the lack of fractal coating, and interface polarization is not an issue even if the RV shock coil is part of the ventricular sensing circuit (i.e., an integrated bipolar DF lead ). Despite being an anode with a large GSA and usually a considerably farther distance away from the myocardium than the tip electrode, anodal capture of the myocardium from the RV shock coil can still occasionally occur.

Interelectrode Distance

The interelectrode distance has been altered in different lead models in attempts to achieve specific performance attributes, such as reduction of far-field ventricular sensing for an atrial lead (e.g., the Optisense atrial lead, St. Jude Medical, St. Paul, MN; tip-ring electrode separation 1.1 mm instead of the typical 10-12 mm; see Fig 11-15C ) or phrenic nerve stimulation (e.g., the Attain Performa quadripolar CS lead, Medtronic, Minneapolis, MN; the interelectrode distance is 1.3 mm for the middle two electrodes instead of 21 mm for the other electrode pairs; see Fig. 11-15D ). However, the impact of these design features on the intended attributes may be limited, as alternative designs in other lead models have comparable performance in clinical use. The interelectrode distance may also be deliberately manipulated to lower pacing threshold and energy consumption (see Fig. 11-14A -C).

For ICD leads, electrode location has a significant impact on sensing and DF performance. Clinical studies of early transvenous ICD leads (Endotak 60 series, Boston Scientific) demonstrated that with a 6-mm tip-to-shock coil spacing, transient diminution in R-wave amplitude could occur, leading to a failure to redetect VF after a failed first shock. Conversely, in experimental studies, if the tip to coil spacing was over 20 mm, DF thresholds would be elevated.

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